Title of Invention	HYDROXYPHENYL CROSS-LINKED MACROMOLECULAR NETWORK AND APPLICATIONS THEREOF
Abstract	A synthetic, implantable tissue matrix material comprising a macromolecular network comprising wherein R1 and R2 each comprises a structure selected from the group consisting of a) polycarboxylate molecules that have been substituted at CO2H sites thereon with a hydroxyphenyl compound at a substitution rate less than 10 percent based on a total number of CO2H sites on the polycarboxylate molecules, b) polyamines, that have been substituted at primary amine sites thereon with a hydroxyphenyl compound, and c) copolymers thereof, and wherein R1 and R2 can be the same or different structures.

Title of Invention

HYDROXYPHENYL CROSS-LINKED MACROMOLECULAR NETWORK AND APPLICATIONS THEREOF

Abstract

A synthetic, implantable tissue matrix material comprising a macromolecular network comprising wherein R1 and R2 each comprises a structure selected from the group consisting of a) polycarboxylate molecules that have been substituted at CO2H sites thereon with a hydroxyphenyl compound at a substitution rate less than 10 percent based on a total number of CO2H sites on the polycarboxylate molecules, b) polyamines, that have been substituted at primary amine sites thereon with a hydroxyphenyl compound, and c) copolymers thereof, and wherein R1 and R2 can be the same or different structures.

Full Text	HYDROXYPHENYL CROSS-LINKED MACROMOLECULAR NETWORK AND APPLICATIONS THEREOF [0001] This application claims priority from U.S. provisional patent application Serial No. 60/586,585 filed July 9, 2004, the contents of both of which are incorporated herein by reference in their entirety. BACKGROUND OF THE INVENTION [0002] Articular cartilage performs an essential function in healthy joints. It is responsible for absorbing and dissipating impact and frictional loads in order to divert these loads away from bones, to protect the bones from damage. Cartilage performs this function by transferring the loading force to a fluid phase within a three-dimensional network of aggrecan molecules, themselves constrained (described in the next paragraph) within the joint space. Aggrecan molecules have up to 100 chondroitin sulfate (CS) chains attached to a core protein, with each chondroitin sulfate chain possessing multiple negatively charged sulfate groups along their length. The effect of all these sulfate groups is to cause each of the chondroitin sulfate chains in a single aggrecan molecule to repel one another, (resulting in the aggrecan molecule having the maximum possible volume at rest), and also to cause adjacent aggrecan molecules in a cartilage aggregate to repel one another. [0003] In healthy cartilage, aggrecan molecules are attached to long hyaluronan chains, which are in turn constrained in large cartilage aggregates within the joint space by an extracellular collagen fibril matrix. Thus, even though adjacent chondroitin sulfate chains in each aggrecan molecule (and adjacent aggrecan molecules attached to the same or a different hyaluronan chain) repel one another, they are nonetheless constrained within the collagen matrix. See Fig. 1 depicting normal, healthy cartilage. Because the chondroitin sulfate chains are so repulsive, the hyaluronan-aggrecan network (or macromolecular network) expands as much as possible within the constraints of the collagen matrix to achieve the lowest possible energy state at rest; i.e. to allow the maximum possible spacing between adjacent negatively charged sulfate groups. As a result, network molecules are highly resistant to being shifted or displaced in order to avoid approaching an adjacent network molecule. These large cartilage aggregates are trapped at one fifth their free solution volume within a meshwork of collagen fibers, which resist any further swelling. Cartilage aggregates with their high negative charge density bind large solvent domains, and contribute to cartilage's ability to absorb loads and resist deformation. Upon compression, the distance between the fixed-negative charge groups on the proteoglycans decreases, which increases the charge-to-charge repulsive forces as well as the concentration of free-floating positive counterions (such as Ca2+ and Na+). Both effects contribute to the viscoelastic nature of cartilage and its ability to resist deformation and absorb compressive loads, further described below. [0004] Within the macromolecular network are water molecules which provide a substantially continuous fluid phase. The macromolecular network diverts impact and frictional loads away from bones by transferring them to the continuous fluid (water) phase as follows. As a joint undergoes a load, the force is absorbed first by the macromolecular network, where it acts on and tends to deform or compress the network. The force sets up pressure gradients in the fluid phase in order to induce fluid flow to accommodate network deformation or compression resulting from the load. But the fluid cannot negotiate the tight macromolecular network, packed with the repulsive chondroitin sulfate chains, sufficiently to ac: mmodate a bulk flow of water without shifting or displacing the network molecules. Hence, individual water molecules may diffuse within the network, but the bulk fluid phase is substantially constrained from flowing through the network except at a much slowed rate due to the resistance to displacement of network molecules. Because the water molecules cannot flow readily despite the pressure gradients, the energy from the impact or frictional load is transferred to and absorbed by the fluid phase where it contributes to compressing the liquid water until the water can be sufficiently displaced to accommodate the network conformation and the pressure gradients have subsided. The overall result is that cartilage absorbs the potentially harmful load, thereby diverting it from bone. [0005] Through this elegant mechanism, normal cartilage is capable of absorbing significant loads by transferring the bulk of the loading force to a fluid phase constrained within a macromolecular network. This arrangement has yet to be adequately duplicated via artificial or synthetic means in the prior art. Consequently, there is no adequate remedy for cartilage degenerative disorders, such as arthritic disorders, where the aggrecan molecules become separated from their hyaluronan chains and are digested or otherwise carried out from the cartilage aggregates. [0006] Osteoarthritis and rheumatoid arthritis affect an estimated 20.7 and 2.1 million Americans, respectively. Osteoarthritis alone is responsible for roughly 7 million physician visits a year. For severe disabling arthritis, current treatment involves total joint replacement with on average 168,000 total hip replacements and 267,000 total knee replacements performed per year in the U.S. alone. Defects in articular cartilage present a complicated treatment problem because of the limited capacity of chondrocytes to repair cartilage. Treatment strategies to date have focused on the use of autologous chondrocytes expanded in culture or the recruitment of mesenchymal stem cells in vivo by chemotactic or mitogenic agents. The intent of these strategies is to increase and/or activate the chondrocyte population so as to resynthesize a normal, healthy articular cartilage surface. One major difficulty associated with these strategies is the inability to maintain these agents at the site of the defect. Hyaluronan has been proposed as a candidate for the development of biomaterials for local delivery of chondrocytes or bioactive agents because of its unique properties, including excellent biocompatibility, degradability, and Theological and physiochemical properties. However, it has been unknown whether chondrocytes suspended in a tissue engineered hyaluronan matrix would be able to synthesize a new cartilage matrix with mechanical properties comparable to normal, healthy articular cartilage. This is because conventional biomaterials made from hyaluronan are formed through chemistries that are incompatible with maintaining cell viability. Chondrocytes must be introduced to the matrices after matrix formation with variable and normally poor results. [0007] Accordingly, there is a need in the art for an artificial or synthetic matrix that can effectively divert a loading force from bones in an effective manner. Preferably, such a matrix can be provided in situ or in vivo to repair or replace articular cartilage during an orthopedic surgical procedure. Most preferably, the artificial or synthetic matrix can be provided to an in situ or in vivo target site as a liquid or a plurality of liquids, and can set up in place to provide a substantially seamless integration with existing cartilaginous and/or bony tissue in a patient. [0008] It also is desirable to provide an artificial or synthetic matrix that can be used or adapted to synthesize a variety of replacement tissues. SUMMARY OF THE INVENTION [0009] A synthetic, implantable tissue matrix material is provided including a macromolecular network that includes the following structure wherein R1 and R2 each comprises a structure selected from the group consisting of polycarboxylates, polyamines, polyhydroxyphenyl molecules, and copolymers thereof, and wherein R1 and R2 can be the same or different structures. [0010] A variety of synthetic, implantable tissue materials also are provided which include or are composed of the tissue matrix material mentioned in the preceding paragraph, including a synthetic, implantable cartilage material; a synthetic, implantable vocal cord material; a synthetic, implantable vitreous material; a synthetic, implantable soft tissue material; and a synthetic, implantable mitral valve material. BRIEF DESCRIPTION OF THE DRAWINGS [0011] Fig. 1 is a schematic diagram of normal, healthy human cartilage. [0012] Fig. 2 is a schematic diagram of a dihydroxyphenyl cross-linked macromolecular network according to the invention. [0013] Fig. 3 is a structural formula of a hyaluronan molecule. [0014] Figs. 4a-4c are graphs showing comparative results for mechanical testing in a confined compression test (equilibrium stress versus applied strain) of T-HA (Fig. 4a), T- Aggrecan (Fig. 4b) and 50% T-HA/50% T-Aggrecan composite (Fig. 4c) hydrogels according to the invention versus published results for articular cartilage plugs (Example 3). The relationship between glycosaminoglycan (GAG) concentration and material compressive strength is shown in Fig. 4d. [0015] Fig. 5 is a graph showing comparative data of glucose utilization for chondrocytes embedded in T-HA hydrogels (1.7% and 4.7% T-HA) compared to cultured on tissue culture plastic (control). [0016] Fig. 6 is a series of four photographs illustrating a surgical procedure to implant a T-HA hydrogel into articular cartilage defects according to an aspect of the invention described in Example 6. [0017] Fig. 7 is a series of two photographs showing the T-HA hydrogel implants one month after implantation into the medial trochlar facet of a Yucatan pig as described in Example 6, as well as the opposing (articulating) patella surface. [0018] Fig. 8 is a series of photographs illustrating the histological results of control side (unfilled) and experimental side (TB-HA hydrogel filled) canine vocal cords, 3 months post-operatively, following a vocal cord repair procedure using a T-HA hydrogel as a synthetic vocal cord material as described in Example 7. [0019] Fig. 9 is a series of photographs illustrating the histological results of surgically augmented vocal cords in a rabbit model using a T-HA hydrogel as a synthetic vocal cord material, also as described in Example 7. [0020] Fig. 10 is a series of photographs of control (unoperated) and experimental (surgically replaced) eyes one month post-operative, following a vitreous replacement procedure using T-HA hydrogel as a synthetic vitreous material as described in Example 8. [0021] Fig. 11 shows comparative electroretinogram (ERG) results recorded for both control and vitreous replaced eyes in response to flashes of light in a rabbit model as described in Example 8. [0022] Fig. 12 is a series of electron micrographs of the retina from four quadrants of control (unoperated) and experimental (surgically replaced) eyes one month post-operative, following a vitreous replacement procedure using T-HA hydrogel as a synthetic vitreous material as described in Example 8. [0023] Fig. 13 is a series of photographs showing representative results of histological results for a 100 mg/ml T-HA hydrogel plug implanted subcutaneously into an immunocompetent rat at one month post-operatively as described in Example 9. [0024] Fig. 14 is a photograph of a cadaveric canine heart used to specify T-HA hydrogel materials for mitral valve repair as described in Example 10. DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS OF THE INVENTION [0025] As used herein, the term polycarboxylate means a molecule, structure or species having a chain length of at least two functional groups or units, wherein at least two such groups or units of the chain are or comprise carboxylic acid groups that are sterically accessible to a nucleophilic substitution reaction as described herein. Also as used herein, the term polyamine means a molecule, structure or species having a chain length of at least two functional groups or units, wherein at least two such groups or units of the chain are or comprise primary amine groups that are available for a nucleophilic substitution reaction. Also as used herein, a polyhydroxyphenyl molecule means a molecule having a chain length of at least two functional groups or units, wherein at least two such groups or units of the chain are or comprise hydroxyphenyl groups that can be linked to another hydroxyphenyl group via a C-C bond. Also as used herein, a hydrogel is a material that is prepared comprising a macromolecular network that is used or useful in tissue replacement or engineering applications, e.g. as artificial cartilage, as a material to coat surgical instruments to prevent tissue irritation, or to provide a semi-permeable membrane such as for use in an artificial kidney, etc. [0026] The invention includes a novel structure of a macromolecular network that has been formed by linking hydroxyphenyl groups attached to adjacent long chain macromolecules, resulting in effectively cross-linking the macromolecules to provide a large network. The basic cross-linking structure of the network is shown below where R1 and R2 are each long chain macromolecules. R1 and R2 can be the same molecule or different molecules, but it will be understood that to provide a suitable network, R1 and R2 will be different molecules for at least a portion of the dihydroxyphenyl linkages in a network according to the invention. It is not necessary, though it is preferred, that R1 and R2 are the same species of molecule. [0027] By providing a plurality of these dihydroxyphenyl linkages between adjacent macromolecules, a network of dihydroxyphenyl cross-linked macromolecules is provided as shown schematically in Fig. 2. In the figure, the macromolecules are represented schematically by cylindrical strands, each preferably having at least two hydroxyphenyl groups attached along its length. It is noted that not every hydroxyphenyl group must be linked to another hydroxyphenyl group. [0028] Briefly, the disclosed invention involves covalent coupling of hydroxyphenyl containing compounds, including but not limited to tyramine, through their primary amine (or carboxyl) groups to carboxyl (or primary amine) groups on various polymeric scaffold materials, including but not limited to hyaluronan or chondroitin sulfate (e.g. in the form of aggrecan), via a carbodiimide-mediated reaction. After isolation and purification of the hydroxyphenyl-substituted polymeric scaffolds, the hydroxyphenyl residues are selectively cross-linked by horseradish peroxidase (HRP) in the presence of very dilute hydrogen peroxide to form hydrogels. As will become apparent, the hydrogels made as described herein are or can be used as a fully implantable, non-immunogenic synthetic tissue matrix material that can be implanted into the body for a variety of purposes as will be described. As used herein, 'implantable' refers both to surgical implantation of a hydrogel as through a surgical incision, and to provision of the hydrogel within the body via injection, e.g. using a syringe. Whether surgically implanted or injected, the implantable hydrogels can be provided within the body already cross-linked (ex vivo cross-linking) or otherwise it can be cross- linked in situ at the site of implantation within the body as will be further described. [0029] The first step in providing the macromolecular network is to prepare or provide the long-chain macromolecules having periodic hydroxyphenyl groups attached. In one embodiment, the macromolecules are polyhydroxyphenyl molecules which already have multiple or periodic hydroxyphenyl groups, such as polyphenols. Suitable polyphenols include polyamino acids (e.g. polytyrosine), epigallocatechin (EGC), and epigallocatechin gallate (EGCG) isolated from green tea, less preferably other polyphenols. [0030] In a further embodiment, the hydroxyphenyl groups can be added to the macromolecules periodically or randomly along their length via a chemical reaction. A preferred method of adding hydroxyphenyl groups to the macromolecules is to utilize a carbodiimide-mediated substitution reaction pathway to provide an amide bond between a primary amine having a hydroxyphenyl group and a carboxylic acid group attached to the macromolecules. In this method, the long-chain macromolecule preferably is a polycarboxylate molecule, having periodic carboxylic acid groups along its length. The hydroxyphenyl groups are provided as part of smaller molecules having primary amine groups that can be attached to the carboxyl carbon atoms of a carboxylic acid group on the long-chain macrqmolecules via the carbodiimide pathway. The reaction proceeds as follows: where: Structure A is a carbodiimide; Structure B is a polycarboxylate (though only one CO2H group is shown); Structure C is the product of Reaction A and is an activated O-acylisourea; Structure D is a primary amine having a hydroxyphenyl group; Structure E is a hydroxyphenyl-substituted polycarboxylate; and Structure F is an acylurea byproduct; wherein individual Rs can be individually selected, the same or different from one another, to be a straight chain or branched alkane or acyl group, or any other structure that does not interfere with the carbodiimide reaction pathway to provide the amide bond between the NH2 and CO2H groups as shown in Structure E above. [0031] In the above-illustrated pathway, Reaction A represents a carbodiimide activation of the carboxyl group to provide an activated O-acylisourea intermediate. The electropositive carbon atom of this intermediate is receptive to nucleophilic attack by the lone pair of electrons on a nitrogen atom of an adjacent primary amine molecule having an attached hydroxyphenyl group. The products of this nucleophilic substitution reaction (Reaction B) are a hydroxyphenyl-substituted polycarboxylate and an acylurea byproduct which can be dialyzed out to provide a substantially pure hydroxyphenyl-substituted polycarboxylate product. [0032] Certain side-reactions are possible in the above-described carbodiimide reaction pathway chemistry and should be considered by the person having ordinary skill in the art. First, the carbodiimide can react with nucleophiles other than the carboxylate oxygen atom of the polycarboxylate molecule required to form the desired O-acylisourea (reaction A). Such nucleophiles may include the amine and/or hydroxyphenyl groups of Structure D illustrated above. In particular, there are three potential side-reactions for Reaction A which can reduce the effective concentration of the carbodiimide and the primary amine having the hydroxyphenyl group (Structures A and D), and potentially lead to the creation of undesired adducts on the polycarboxylate (Structure B): [0033] The product of an amine reaction with the carbodiimide (Reaction C) will not have a free amine group effectively reducing the amount of tyramine available for reaction with the O-acylisourea. This reaction also reduces the amount of carbodiimide available for formation of the desired O-acylisourea. The products of the hydroxyphenyl reaction (Reaction D) are not UV absorbent, which will make their detection by UV-spectroscopy in the final hydroxyphenyl-substituted polycarboxylate product (explained below) more difficult. However, because these products still contain free amine groups, they can form amide bonds with the polycarboxylate molecule via Reaction B. This can give rise to two unproductive hyaluronan-substituted structures, neither of which can participate in the peroxidase cross-linking reaction in the second step (described below) of preparing the cross- linked network due to the absence of an extractable phenolic hydroxyl hydrogen atom needed to generate the free radical (also explained below). Finally, the carbodiimide can react non- productively with water (Reaction E) to produce the same acylurea shown above as a byproduct of Reaction B, but with none of Structure E, the desired product. [0034] Once the desired O-acylisourea product has been formed in Reaction A, there is again the possibility for certain additional side-reactions: [0035] The O-acylisourea (Structure C) can be hydrolyzed as shown in Reaction F releasing the original unmodified polycarboxylate (Structure B) and the acylurea of the carbodiimide (Structure F). This is an unproductive reaction similar to reaction E, which reduces the effective concentration of the carbodiimide. The O-acylisourea, can also undergo an intramolecular rearrangement (Reaction G) to form two unreactive N-acylureas. These structures form unproductive adducts on the carboxylate molecule which cannot contribute to the peroxidase catalyzed cross-linking reaction (step 2 discussed below) for preparing a network according to the invention. The O-acylisourea can also react (Reaction H) with a second carboxyl group on either the same or a different polycarboxylate molecule to form an acid anhydride. This molecule can then react with Structure D to form the desired amide and regenerate the second carboxyl group. Thus there are two potential side-reactions for the O- acylisourea, which can reduce the effective concentration of the carbodiimide (Reactions F and G), and potentially lead to creation of undesired adducts on the polycarboxylate molecule. [0036] Negative effects of these side reactions can be addressed through conventional techniques without undue experimentation. [0037] Alternatively to the pathway shown above where the macromolecule (Structure B) is a polycarboxylate, the macromolecule can be a polyamine having multiple or periodic amine groups along its length, wherein the hydroxyphenyl groups then are provided as part of smaller carboxylic acid molecules. Suitable polyamines include: polyhexosamines such as chitosan (polyglucosamine); polyamino acids such as polylysine; polydeoxyribonucleotides such as poly (dA) (polydeoxyadenylic acid), poly(dC) (polydeoxycytidylic acid), and poly(dG) (polydeoxyguanylic acid); and polyribonucleotides such as poly(A) (polyadenylic acid), poly(C) (polycytidylic acid), and poly(G) (polyguanylic acid). The carbodiimide-mediated reaction pathway proceeds exactly as explained above to form the amide bond between the amine group and carboxylic acid group except that, as will be understood by a person having ordinary skill in the art, the resulting product will be hydroxyphenyl-substituted polyamine instead of a polycarboxylate. Other peptides and/or proteins also can be used as the macromolecules in the present invention, either which have hydroxyphenyl groups disposed along their length, or to which hydroxyphenyl groups can be provided via a substitution reaction as described herein. For example, in addition to the peptides already disclosed herein, polyarginine can be used as the macromolecule. [0038] When substituting onto a polycarboxylate molecule, suitable hydroxyphenyl- containing compounds for use in the present invention include those having a free primary amine that can be used to modify scaffold materials having multiple or periodic CO2H groups, including tyrosine (2-amino-3-(4-hydroxyphenyl) proprionic acid) and tyramine (tyrosamine or 2-(4-hydroxyphenyl) ethylamine). When substituting onto a polyamine, suitable hydroxyphenyl-containing compounds include those having a free CO2H group that can be used to modify scaffold materials having multiple or periodic primary NH2 groups, including tyrosine, 3-(4-hydroxyphenyl) propionic acid and 4-hydroxyphenylacetic acid. [0039] The second step in preparing a cross-linked macromolecular network according to the invention is to link the resulting macromolecules, now having one or more hydroxyphenyl groups attached, via a dihydroxyphenyl linking structure. In this step hydroxyphenyl groups attached to different macromolecules are linked via the reaction mechanism shown below using a peroxide reagent in the presence of a peroxidase: [0040] (It is noted that some dihydroxyphenyl linking may occur between different hydroxyphenyl groups attached to the same molecule as well). Peroxidase in the presence of a dilute peroxide (preferably H2O2) is able to extract the phenolic hydroxyl hydrogen atom from hydroxyphenyl containing compounds (such as tyramine) leaving the phenolic hydroxyl oxygen with a single unshared electron, an extremely reactive free radical. The free radical isomerizes to one of the two equivalent ortho-position carbons and then two such structures dimerize to form a covalent bond effectively cross-linking the structures, which after enolizing generates a dihydroxyphenyl dimer (a dihydroxyphenyl linkage such as dityramine linkage as described below). [0041] For clarity, only a single dihydroxyphenyl linking reaction is shown above, but it will be understood that several or multiple such linkages will be produced when macromolecules having attached hydroxyphenyl groups are subjected to the reaction conditions (peroxide and peroxidase). Hydrogen peroxide is indicated in the above mechanism, but other suitable peroxides can be used. Also, the peroxidase preferably is horseradish peroxidase (HRP). Alternatively, any other suitable enzyme (or other agent) can be used that is capable of generating free-radicals for cross-linking scaffold materials containing hydroxyphenyl groups, preferably under ordinary metabolic conditions as described below. [0042] We have shown that the interaction of horseradish peroxidase (Type II) and hydrogen peroxide (H2O2) is suitable for the production of cross-linked macromolecular networks. The mechanism comprises four distinct steps: (a) binding of peroxide to the heme- Fe(IH) complex of the peroxidase to form an unstable peroxide complex, "Compound I"; (b) oxidation of the iron to generate a ferryl species with a pi-cation radical in the heme porphyrin ring, "Compound II"; (c) reduction of Compound II by one substrate (i.e. hydroxyphenyl or water) molecule to produce a product (i.e. hydroxyphenyl or superoxide) radical and another ferryl species, "Compound III"; (d) reduction of Compound HI by a second substrate (i.e. hydroxyphenyl or water) molecule to release a second product (i.e. hydroxyphenyl or superoxide) radical and regenerate the native enzyme. Thus the peroxidase enzyme can either form hydroxyphenyl radicals required for cross-linking through interaction of hydroxyphenyl groups at the enzyme active site to directly create the desired radicals or through first generation of superoxide radicals, which then diffuse from the enzyme and interact with hydroxyphenyl groups to generate the desired radicals. Other compounds that have the potential to produce the same effect include any porphyrin containing compound (i.e. Photofrin below), which includes the peroxidase family, hemoproteins, or the structurally related chlorin compounds. [0043] A number of other free radical initiators can be used to crosslink the hydroxyphenyl modified macromolecules described herein. A majority are based on the formation or inclusion of reactive oxygen species (ROS) such as, but not limited to, molecules of hydrogen peroxide, ions of hypochlorite, radicals like the hydroxyl radical, and the superoxide anion which is both ion and radical. Additional reactive molecules such as reactive nitrogen species or reactive sulfur species, or those free radical species involved in synthetic polymerization have the potential to be used for hydroxyphenyl cross-linking. [0044] ROS are commonly produced in nature through the use of enzymes, and substrates. Additional enzymatic systems which have the potential to be used in the cross- linking process, as a result of production of superoxide radicals, include, but are not limited to xanthine-xanthine oxidase and NADPH-NADPH oxidase. [0045] Another class of ROS free radical initiators that can be used involves the use of metallic cations. One example is based on the Fenton reaction, which takes place between hydrogen peroxide and a bivalent cation, such as Fe2+. This process generates powerful free radicals when the catalyst reacts with hydrogen peroxide. The principal chemical reaction associated with Fenton's reaction is shown below: where, Fe2+ = ferrous ion, Fe3+ = ferric ion, OH. = hydroxyl radicals [0046] In addition to the initiation reaction described above that produces hydroxyl radicals, the Fenton's process can also produce superoxide radicals and hydroperoxide anions by additional chain propagation reactions described below. The perhydroxyl radical is known to be a weaker reductant compared to superoxide radical and hydroperoxide anions. where O2. = superoxide radical anion, HO2 = hydroperoxide anion, HO2' = perhydroxyl radical. [0047] We have demonstrated the ability for this reaction to crosslink tyramine substituted hyaluronan in the laboratory using ferrous sulfate in conjunction with hydrogen peroxide. Compounds which include, but are not limited to, bivalent cations of copper, chromium, vanadium and cobalt can be used in a similar manner. It is to be noted that while the hydroxyl free radical can be used to form a dityramine crosslink, it has also been shown to cleave HA chains, and thus may ultimately be unsuitable for ideal hydrogel formation. Additional molecules or methods which can generate ROS include: • rubidium or cesium ions in the presence of oxygen to form superoxide radicals; • trivalent cations, which with hydrogen peroxide form free radicals and bivalent cations as shown below, which can subsequently follow the reactions involved in the Fenton process. • the cytotoxic and antitumor therapy Photofrin, which upon illumination with laser light at a wavelength of 630 nm causes propagation of a radical generating reaction that produces superoxide and hydroxyl radicals. In the absence of light, but the presence of hydrogen peroxide, the porphorin ring in Photofrin should operate by the same reaction as for the peroxidase enzyme above. • UV light and hydrogen peroxide to form hydroxyl and superoxide free radicals. • the persulfate family in combination with TEMED. [0048] As noted above, one alternative method for generating such free-radicals is to use Photofrin as an alternative, non-enzymatic, light-activated cross-linking agent to cross- link the macromolecular network described herein, e.g. tyramine-substituted hyaluronan to form tyramine cross-linked hyaluronan hydrogels. Photofrin®, which is known in the art, generates free radicals which could initiate the cross-linking reaction as described herein in a manner similar to the peroxidase-H2O2 mechanism described above. Photofrin® is a porfimer sodium manufactured in powder or cake form by Wyeth-Ayerst Lederle Parenterals, Inc. [0049] The dihydroxyphenyl cross-linked macromolecular network is superior to conventional cartilage or other tissue replacement or substitution methods and products, at least with respect to the ability to carry out an in situ cross-linking procedure, because the preferred cross-linking reaction is enzyme driven (peroxidase). This means the cross-linking reaction is carried out under ordinary in vivo or metabolic conditions of temperature such as 35-39°C (e.g. about 37°C), pH range of 6-7 (e.g. about 6.5), reagents etc. (A peroxide, such as hydrogen peroxide, is the only required reagent for the cross-linking reaction). In addition, Photofrin already is used in in vivo applications, e.g. ablative treatment of Barrett's esophagus, and the iron-based cross-linking mechanism also can be optimized for in vivo performance. Thus, the cross-linking reaction can be performed in vivo, to provide a cross- linked hydrogel at a surgical situs, such as an orthopedic surgical situs, to promote maximum seamless integration between the hydrogel and native tissue such as bony and cartilaginous tissue. Integration of the new hydrogel scaffold with native cartilage matrix may occur immediately as the hydroxyphenyl-substituted macromolecular scaffold quickly penetrates into the existing cartilage matrix prior to cross-linking, and cross-links not only with other hydroxyphenyl-substituted macromolecular scaffold material but potentially with tyrosine residues of resident proteins in the existing cartilage matrix. This would eliminate a typical problem found with pre-formed matrix plugs, which is their poor integration into the native cartilage tissue. The ability to cross-link the hydrogel directly on the articular surface eliminates the need to surgically enlarge a defect to fit a pre-cast plug, as is necessary for hydrogels whose chemistries are toxic to or otherwise prohibit their formation inside the patient. It should be noted that most cartilage damage as a result of arthritis presents as a variable thinning of the articular surface, not holes of defined shape. [0050] For the peroxidase mechanism, because the cross-linking reaction requires both the peroxide and a peroxidase (preferably horseradish peroxidase), solutions containing all but one of these components can be prepared for convenient application to a surgical site. For example, a solution comprising a tyramine - (or other hydroxyphenyl containing species) substituted polycarboxylate (such as tyramine-substituted hyaluronan, etc.) and the peroxidase can be prepared, with a second solution prepared containing the peroxide. Alternatively, the peroxide and the peroxidase can be swapped between the first and second solutions, the importan thing being that the peroxide and peroxidase are kept separate (i.e. in separate solutions) until the cross-linking reaction is to be carried out. Then, the first solution is applied, (e.g. to an in vivo surgical situs), and the second solution is applied or sprayed over the first, in vivo, to cause in situ cross-linking of the tyramine residues. The cross linking reaction occurs in vivo. Other combinations will be evident from the present disclosure which are within the skill of a person of ordinary skill in the art. [0051 ] Furthermore, because the cross-linking reaction occurs under ordinary metabolic conditions, additional living cells, such as chondrocytes, progenitor cells, stem cells, etc., can be provided directly to a medium containing the non-cross-linked hydroxyphenyl-substituted polycarboxylates or polyamines (or polyphenols), i.e. to the first or second solution from the preceding paragraph, wherein the cell-rich medium is applied with the macromolecules to the site in vivo, and the molecules are subsequently cross-linked via addition of peroxidase and peroxide. The result is a cross-linked macromolecular network containing the desired cells dispersed within it. Such a cell-enriched network is not possible in conventional tissue replacement matrices due to the harsh conditions of temperature and pH under which they are prepared. Further, as described below in Example 5, it has been demonstrated that the cells provided to the network as described above remain viable even after cross-linking of tyramine-substituted hyaluronan (also described below). [0052] In a preferred embodiment particularly suitable for preparing synthetic cartilage as well as other synthetic or artificial tissues, the macromolecule used to produce the network is hyaluronan or hyaluronic acid (HA), and the hydroxyphenyl group is supplied in the form of tyramine. Hyaluronan (HA) is a ubiquitous molecule, which is most concentrated in specialized tissues such as cartilage, vocal cords, vitreous, synovial fluid, umbilical cord, and dermis. In these tissues, its function is manifold, influencing tissue viscosity, shock absorption, wound healing, and space filling. HA has been shown to influence many processes within the extracellular matrix (ECM) in native tissues where it is present including matrix assembly, cell proliferation, cell migration and embryonic/tissue development. [0053] HA is composed of repeating pairs of glucuronic acid (glcA) and N- acetylglucosamine (glcNAc) residues linked by a B 1,3 glycosidic bond as shown in Fig. 3. The glucuronic acid residue is particularly pertinent to the production of a macromolecular network as described herein as this sugar provides an available carboxyl group periodically along the repeat disaccharide structure of HA that is useful for hydroxyphenyl, i.e. tyramine, substitution. For each hyaluronan chain, this simple disaccharide is repeated up to 10,000 times or greater resulting in macromolecule that can have a molecular weight on the order 10 million daltons (10 megadaltons). Adjacent disaccharide units of HA are linked by a B1.4 glycosidic bond, also seen in Fig. 3. Each glcA residue has a carboxylic acid group (CO2H) attached to the number 5 carbon atom of the glucose ring. Under biological conditions, HA is a negatively charged, randomly coiled polymer filling a volume more than 1,000 times greater than would be expected based on molecular weight and composition alone. As noted above, the strong negative charges attract cations and water, which allow HA to assume the form of a strongly hydrated gel in vivo, giving it a unique viscoelastic and shock-absorbing property. HA represents a readily available and desirable scaffolding material for tissue engineering applications as it is non-immunogenic, non-toxic and non-inflammatory. Also as a naturally occurring extracellular matrix (ECM) molecule it offers the advantages of being recognized by cell receptors, of interacting with other ECM molecules, and of being metabolized by normal physiological pathways. [0054] Tyramine is a phenolic molecule having an ethyl amine group attached para to the OH group on the benzene ring. When these species are used, the mechanism for tyramine substitution onto the singly bound oxygen atom of a CO2H group on HA proceeds via the carbodiimide-mediated reaction mechanism described above as illustrated immediately below. The preferred carbodiimide species is l-ethyl-3-(3- dimethylaminopropyl)carbodiimide (EDC) as shown. where: Structure A is EDC; Structure B is hyaluronan (though only one CO2H group is shown); Structure C is the product of Reaction A and is l-ethyI-3-(3-dimethylaminopropyI) isourea; Structure D is tyramine; Structure E is tyramine-substituted hyaluronan; and Structure F is l-ethyl-3-(3-dimethylaminopropyl) urea (EDU). [0055] In the above pathway, a negatively charged oxygen atom of the carboxyl group of the hyaluronan molecule attacks, via a nucleophilic reaction mechanism, the electron-deficient diimide carbon atom on the carbodiimide molecule (EDC) to form the activated O-acylisourea (Reaction A). The result is that the carbon atom of the HA carboxylate group becomes sufficiently electron deficient to be susceptible to nucleophilic attack by the unshared pair of electrons on the amine group of a tyramine molecule (Reaction B). Reaction A is preferably catalyzed by a suitable catalyst that will result in the formation of an active ester during Reaction A, thus permitting the reaction to be carried out at substantially neutral pH (e.g. pH=6.5). Suitable catalysts include N-hydroxysuccinimide (NHS), less preferably 1-hydroxybenzotriazole (HOBt) or N-hydroxysulfosuccinimide (NHSS), less preferably another suitable catalyst or combinations thereof effective to enhance the carbodiimide reaction by formation of an active ester in order to minimize the unproductive hydrolysis of carbodiimides at higher pHs. Less preferably other carbodiimides besides EDC can be used, including l-cyclohexyI-3-[2-(4- methylmorpholino)ethyl]carbodiimide (CMC), and dicyclohexylcarbodiimide (DCC). [0056] The result of Reaction A above is O-acylisourea-substituted hyaluronan; essentially the EDC molecule has been temporarily substituted onto the carboxylic acid group of a glcA residue from the HA molecule, making the carbon atom of the carboxylic acid group slightly positively charged. The electron pair from the terminal amine group of a tyramine molecule is then substituted onto the carbon atom via a nucleophilic substitution reaction as explained in the preceding paragraph (Reaction B). The result of Reaction B is the tyramine-substituted HA molecule (T-HA) and acylurea, a byproduct. It will be understood that Reactions A and B will result in a plurality of tyramine substitutions on the periodic glcA residues of HA molecules; a single substitution has been shown here for brevity and clarity. [0057] After formation of T-HA, a plurality of T-HA molecules are reacted via peroxide and peroxidase enzyme to cross-link T-HA molecules as previously described and illustrated above. That is, the hydroxyphenyl groups on the tyramine residues now attached to HA molecules react with peroxide (preferably H2O2) in the presence of a peroxidase to remove the phenolic hydrogen atom resulting in a tyramine free radical, with the unpaired electron associated with the phenolic oxygen atom. This free radical species isomerizes or resonates, resulting in a resonance structure (or free radical isomer) with the unpaired electron now associated with an ortho carbon atom on the phenolic ring. In this position, the unpaired electron quickly reacts with a similarly situated unpaired electron on another tyramine free radical to form a covalent bond therebetween. The result is a free-radical driven dimerization reaction between different tyramine free radical residues attached to different glcAs of the same or different HA molecules. This dimerized species further enolizes to restore the now-linked tyramine residues, resulting in a dityramine linkage structure. It will be understood that a plurality of reactions as herein described will occur between adjacent tyramine residues, resulting in a cross-linked macromolecular network of T-HA molecules having the following cross-linking structure: [0058] The cross-linked T-HA network can be provided with aggrecan molecules in a conventional manner, e.g. via link proteins, to provide a cross-linked T-HA network having aggrecan molecules attached to the HA chains. Thus, a network similar to that found in a normal cartilage aggregate can be provided, with the dityramine bonds holding the network together thereby constraining the contained aggrecan network, instead of collagen fibrils as in normal cartilage. similarly non-displaceable chondroitin sulfate network (and concomitant water- impermeability) as in normal cartilage, but without the extracellular collagen fibril matrix or the HA chains found in normal cartilage. In fact, by directly cross-linking chondroitin sulfate molecules, (instead of their core HA molecules as in the previously described embodiment), the repulsive force between adjacent chondroitin sulfate molecules may be strengthened, resulting in even stronger fluid flow resistance compared to normal cartilage. This may result in greater loading force absorption and dissipation capacity than normal cartilage because the interstitial fluid phase is even more constrained from flowing. In this embodiment, where chondroitin sulfate molecules are directly cross-linked, certain cartilage degenerative conditions are entirely circumvented; e.g. conditions where the core protein to which chondroitin sulfate molecules are ordinarily bonded in normal cartilage becomes cleaved between the HA binding domain (G1) and the second globular domain (G2) thus allowing the chondroitin sulfate rich region to diffuse out from the cartilage aggregate. In this embodiment, because the chondroitin sulfate molecules are directly cross-linked to one another, unassociated with an aggrecan or other proteoglycan molecule, they cannot be cleaved or carried away as in normal cartilage. [0062] Nonetheless, a tyramine cross-linked T-HA network (having an HA backbone chain with attached aggrecan molecules, which in turn include chondroitin sulfate chains) may be preferred because of the high availability of HA. This may be beneficial in the case of cartilage replacement or repair using the present invention, because the body's normal metabolic pathway for generating cartilage may be able to build directly onto an implanted tyramine cross-linked T-HA network as will be described. [0063] One further particular application where a cross-linked network according to the invention will have substantial utility is in the production of an artificial kidney. The kidney filters blood by two mechanisms: one is by size exclusion and the second is by charge exclusion. MEMS devices have been designed for use in artificial kidney devices, which contain precisely defined micropores that can effectively mimic only the size exclusion characteristics of the kidney. In a healthy kidney, the charge exclusion related filtration is the result of heparan sulfate proteoglycans present in a basement membrane, which separates two distinct cell types important for other kidney related functions. To mimic this charge barrier in the MEMS engineered artificial kidney, hydrogels can be prepared composed of either heparan sulfate or heparin that are cross-linked via dihydroxyphenyl (dityramine) links as described herein and provided within the pores of the MEMS device. This heparin/heparan sulfate hydrogel can then be sandwiched between two hyaluronan derived hydrogels (e.g. T- HA described above) as described herein, and containing one of each of the cell types normally found in a normally functioning kidney. The central heparin/heparan sulfate hydrogel provides the charge exclusion properties for the device. The outer two hyaluronan hydrogel layers provide protection from the immune system and fouling by normal cellular and molecular debris. Inclusion of the two cell types on opposite sides of the filtration barrier provides a cellular component in its normal physiologic orientation. [0064] In another promising application, the hydrogels herein described can be applied in developing an artificial pancreas. A problem in development of an artificial pancreas is the short half life of MEMS engineered glucose sensors due to fouling of the detector electrode in vivo. Coating of the surface of these detectors with a hyaluronan hydrogel (e.g. T-HA) as described herein would permit diffusion of the small molecular weight glucose molecules that they are designed to detect while providing protection from the immune system and fouling by normal cellular and molecular debris. [0065] In summary, it will be evident from the foregoing that macromolecules useful as scaffold materials for formation of hydrogels include but are not limited to polycarboxylates (containing free carboxylate groups), polyamines (containing free primary amine groups), polyphenols (containing free hydroxyphenyl groups) and their copolymers, examples of which have been described above. When polyphenols are used, the first step in preparing the network described above can be omitted because polyphenols already contain multiple or periodic hydroxyphenyl groups. Otherwise, both polycarboxylates and polyamines must have hydroxyphenyl groups added or substituted along their length, preferably via the above-described carbodiimide reaction pathway. The second step in preparing the network is to carry out an enzyme driven dimerization reaction between two hydroxyphenyl groups attached to adjacent macromolecules (whether polycarboxylates, polyamines or polyphenols) in order to provide a cross-linked structure. This step is carried out using a peroxide reagent (preferably hydrogen peroxide) in the presence of a suitable enzyme (preferably HRP) under metabolic conditions of temperature and pH. [0066] In the case of the preferred dityramine cross-linked T-HA network, in the first step the carboxyl groups on high molecular weight hyaluronan (HA) are substituted with tyramine which introduces reactive hydroxyphenyl groups into the HA molecule. This tyramine substitution reaction preferably is mediated by the carbodiimide, l-ethy]-3-(3- dimetbylaminopropyl)carbodiimide (EDC) with the degree of tyramine substitution on HA controlled by the molar ratios and absolute concentrations of tyramine, EDC and HA used in the reaction mix. Excess reagents such as unused tyramine and EDC are subsequently removed by dialysis, allowing isolation and recovery of high molecular weight tyramine- substituted HA (T-HA). The percent tyramine substitution within each T-HA preparation is easily calculated by measuring: 1) the concentration of tyramine present in the preparation, which is quantitated spectrophotometrically based on the unique UV-absorbance properties of tyramine at 275 ran (see Example 2 below); and 2) the concentration of total carboxyl groups in the HA preparation, which is quantitated spectrophotometrically by a standard hexuronic acid assay. By this technique, T-HA preparations which contain a percent tyramine substitution of only 4 - 6% have been routinely synthesized experimentally. At this level of tyramine substitution, the vast majority (preferably at least 60, 70, 80, 90, or 95, percent) of the HA molecule remains chemically unaltered, and therefore biologically functional. From this formulation of T-HA (i.e. 4 - 6% tyramine substitution) a wide range of biomaterials with a wide range of physical properties can be produced by simply varying the concentration of the T-HA used in the second step of the process. [0067] In the cross-linking reaction, solutions of T-HA are cross-linked to form hydrogels through an enzyme (peroxidase) driven reaction, which catalyzes the formation of a covalent bond between two tyramine adducts on adjacent HA molecules, producing a single dityramine cross-link. The formation of multiple, e.g. hundreds, of these dityramine cross- links per HA molecule result in formation of a stable 3-dimentional scaffold or hydrogel. Addition of very dilute peroxide (preferably H2O2) is required to initiate the cross-linking reaction as it is the peroxide, not -HA, that is the actual substrate for the peroxidase enzyme. The products of the reaction of the peroxidase enzyme on peroxide are free radicals that are preferentially taken up by the hydroxyphenyl rings of tyramine resulting in the formation of the dityramine cross-links. The dityramine linked structures are fluorescent blue (see Example 2), a property which is used to both image the hydrogc ■ and to quantify the degree of cross-linking within the hydrogels. Since the cross-linking reaction is enzyme driven, the hydrogels can be formed under physiologic conditions, and therefore can be formed in the presence of included cells or bioactive agents, or directly adjacent to living tissue while maintaining cell and tissue viability. [0068] The resulting hydrogels are optically clear with a wide range of physical properties depending on the initial T-HA concentration. For example, hydrogels formed from T-HA solutions of 6.25, 12.5,25, 50 and 100 mg/ml T-HA have been shown experimentally to have physical properties (rigidity, rheology and texture) of a jelly, a gelatin, a dough, a resilient rubber-like composition (similar to a rubber ball), and a cartilage-like material respectively- see Example 3. These materials have potential applications in a wide range of clinical settings including tissue engineering of both orthopedic (i.e. cartilage, bone, tendon, meniscus, intervertebral disk, etc.) and non-orthopaedic (kidney, liver, pancreas, etc.) tissues, gene and drug delivery, coating of non-biological devices for in vivo implantation (i.e. glucose sensors, artificial hearts, etc.), wound repair, biosensor design, and vocal cord reconstruction. [0069] Advantageous properties of the hydrogels described herein include the ability to: 1) provide easy characterization and quality control; 2) integrate with existing tissue matrices; 3) directly incorporate into newly formed matrices; 4) directly include cells and bioactive factors; 5) maintain biocompatibility, 6) control bioresorption; 7) cast easily into complicated anatomical shapes (see Example 4 below); and 8) exhibit the mechanical properties of native tissues such as articular cartilage. [0070] Current biologically-based surgical procedures for cartilage repair include autologous chondrocyte implantation, drilling, abrasion chondroplasty, microfracture, and mosaic arthroplasty. All these procedures treat only focal articular cartilage injuries, and not cartilage denuded joint surfaces such as seen in severe osteoarthritis and rheumatoid arthritis. Also, they use either cartilage tissue plugs or expanded chondrocytes harvested from the patient to fill cartilage defects. These tissues or chondrocytes are expected to fill the defect by synthesizing entirely de novo material, such as newly synthesized hyaline cartilage, that has integrated with existing cartilage matrices and has the biomechanical properties of normal cartilage. However, such procedures all promote the formation of a reparative tissue (fibrocartilage) rather than true hyaline cartilage with further mechanical damage to fibrocartilage thought to predispose the joint to osteoarthritis. Furthermore, the availability of endogenous cartilage as a repair material is quite limited with its acquisition presenting its own risks and morbidity to the patient. As evident from the foregoing discussion and as will become further apparent based on the following Examples, the synthetic macromolecular networks and resulting hydrogels disclosed herein present practical materials for promising new therapies in patients suffering from cartilage degenerative diseases. The materials are entirely synthesized from commercially available ex vivo reagents and so involve no morbidity to the patient which conventionally would be required to harvest endogenous material. In addition, the hydrogel (particularly T-HA) can be implanted as an effective cartilage substitute in cartilage denuded joints as a direct intervention for patients suffering from cartilage-degenerative diseases because they can be synthesized so as to emulate the behavior of normal, healthy cartilage. [0071] Rather than relying on synthetic or natural materials or on chondrocytes to produce de novo an implantable, synthetic cartilage-like extracellular matrix (ECM), the present inventors initially focused on purifying the molecules that give cartilage its form and structural characteristics, and then minimally modifying these molecules to make a material resistant to biological degradation. While chondrocytes still may be relied on for maintenance of the synthetic ECM provided by the macromolecular (e.g. T-HA) network post-implantation (e.g. chondrocytes can be embedded into the hydrogel materials as described above), they are not relied on for de novo synthesis. Instead, the basic structure of the synthetic materials described here is modified by cross-linking via a dihydroxyphenyl, preferably dityramine linkage chemistry, to ensure its survival. On further development and experimentation, as will be seen in the following Examples it was discovered mat hydrogels can be made from such materials having a wide array of viscoelastic and other physical properties that can be tuned by appropriate and judicious selection of reagent concentrations and cross-linking conditions to approximate or simulate the properties of other native tissues for which it is or may be desirable to provide a synthetic implantable substitute. [0072] As the Examples below demonstrate, the present hydrogels can be prepared having widely varying properties that are suitable for any number of synthetic tissue implantation or augmentation, as well as other clinical applications. As already described, the present materials can be used to repair cartilage defects produced as a result of either injury or disease. Defects due to injury that can be so repaired can be sports- or accident- related, and may involve only the superficial cartilage layer, or may include the underlying subchondral bone. Defects due to disease which can be repaired using the compositions described herein include those resulting from osteoarthritis and rheumatoid arthritis. Whether from injury or disease, such defects may be in either mature or growth plate cartilage. Formulations for hydrogels for synthetic growth plate cartilage may require the inclusion of unsubstituted scaffold material in order to allow for controlled bioresorption of the biomaterial during growth. [0073] Another potential clinical application for treatment of damaged or arthritic joints is as a replacement for synovial fluid. Conventionally referred to as viscosupplementation therapy, this currently involves injection of a solution of uncross- linked HA into a damaged or arthritic joint, which provides sustained pain relief for weeks even though the HA is cleared from the joint in 1-2 days. Use of the T-HA hydrogels described herein should provide an extended benefit due to their longer i?i vivo persistence compared to uncross-linked HA. [0074] Another field where the hydrogels described herein will be useful is the repair, reconstruction or augmentation of cartilaginous as well as soft tissues of the head and neck. The availability of biomaterials for soft tissue augmentation and head and neck reconstruction has remained a fundamental challenge in the field of plastic and reconstructive surgery. Significant research and investment has been undertaken for the development of a material with appropriate biological compatibility and life span. The outcomes of this research have not been promising. When placed in immunocompetent animals the structural integrity of currently proposed materials has been shown to fail as the framework is absorbed. Furthermore, though conventional synthetic materials offer excellent lifespan, they present certain unavoidable pitfalls. For example, silicones have been fraught with concerns of safety and long-term immune related effects. Synthetic polymers PTFE (gortex) and silastic offer less tissue reactivity but do not offer tissue integration and can represent long term risks of foreign body infections and extrusion. The materials described in this application will be useful to prepare a synthetic soft-tissue scaffold material for the augmentation or repair of soft-tissue defects of the head and neck. In particular, a cross-linked tyramine-substituted hyaluronan (T-HA) hydrogel, which is non-inflammatory, non-immunogenic, and which can be prepared having the appropriate degree of viscoelasticity (see Examples below), could be used as an effective implantable scaffold material. In addition, the unique ability of the preferred enzyme-driven cross-linking chemistry to maintain cell viability permits inclusion of cells such as chondrocytes directly into the hydrogels during formation which can be performed in situ at a defect site. Thus, the need to sculpt or mold an anatomically compatible graft shape to fit a particular defect site is eliminated. [0075] The dityramine cross-linked T-HA network described above has particular utility for producing artificial or synthetic cartilage. The present hydrogel materials can be used, for example, as a novel, biocompatible and biocompliant material to prepare cartilage implants which are frequently used in reconstructive procedures of the head and neck to repair cartilaginous or bony defects secondary to trauma or congenital abnormalities. Applications specific to the ear include otoplasty and auricular reconstruction, which are often undertaken to repair cartilaginous defects due to trauma, neoplasm (i.e., squamous cell carcinoma, basal cell carcinoma, and melanoma), and congenital defects such as microtia. Applications specific to the nose include cosmetic and reconstructive procedures of the nose and nasal septum. Dorsal hump augmentation, tip, shield and spreader grafts are frequently used in cosmetic rhinoplasty. Nasal reconstruction following trauma, neoplasm, autoimmune diseases such as Wegeners granulomatosis, or congenital defects require cartilage for repair. Septal perforations are difficult to manage and often fail treatment Cartilage grafts would be ideal for these applications, as autologous or donor cartilage is often unavailable. Applications specific to the throat include laryngotracheal reconstruction, which in children usually requires harvesting costal cartilage, which is not without morbidity. Auricular and septal cartilage is often inadequate for this application. Synthetic cartilaginous materials prepared from hydrogels disclosed herein can be synthesized to suit each of the foregoing applications, based on tuning parameters of hydrogel synthesis such as reagent concentration, substitution and cross-Unking rates, etc., as evident from the below Examples. Laryngotracheal reconstruction is usually performed for airway narrowing due to subglottic or tracheal stenosis. The etiology may be traumatic (i.e., intubation trauma, or tracheotomy) or idiopathic. Other possibilities include chin and cheek augmentation, and use in ectropion repair of the lower eyelid, in addition to numerous craniofacial applications. It should be noted that these applications may not need cartilage with the exacting mechanical properties of articular cartilage. Inclusion of a cell population or bioactive agents may also be desirable. [0076] The hydrogel materials described herein also can be used for repair and narrowing of the nasal cavity, normally following overly aggressive surgical resection, to prevent the chronic pooling of fluid in the nasal passages that leads to infection and encrustation. Another promising application is in laryngotracheal reconstruction in both children and adults, as a result of laryngotracheal injury due for example to intubation during a surgical procedure such as cardiovascular surgery. Damaged tracheal cartilage at the anterior and posterior portion of the tracheal ring can be replaced with pre-cast hydrogel bile ducts of the liver. The bile ducts form a branching network within the liver and exit the liver via two main branches that are combined into the common bile duct which drains the liver and gallbladder of bile into the duodenum. The bile ducts are very narrow in diameter, measuring only up to 2 nun normally at their largest most distal portions, and yet they must normally drain liters of bile every day from the liver into the duodenum. Any blockage of these ducts can result in a serious condition known as jaundice, which allows many toxins and especially hemoglobin breakdown products to accumulate in the body. PSC is a scarring or stricturing disease of the bile ducts within the liver and in the extrahepatic bile ducts described above that connect the liver to the small intestine. The bile duct strictures of PSC may be treated or prevented with the present hydrogels. for treatment of chronic pancreatitis. Chronic pancreatitis is a chronic inflammatory disease of the pancreas that may be complicated by scars or strictures of the pancreatic ducts. These strictures block the drainage of pancreatic juice, which normally must exit the pancreas through a system of ducts or drainage conduits into the small intestine. The pancreatic juice contains many digestive enzymes and other elements important to normal digestion and nutrient absorption. Blockage or narrowing of the pancreatic ducts by chronic pancreatitis can results in severe complications in which the pancreas autodigests and forms life-threatening abdominal infections and or abscesses. The pancreatic strictures of chronic pancreatitis may be treated or prevented with the present hydrogels. for treatment of gallstone-induced bile duct and pancreatic duct strictures. Gallstones are a very common disorder, a principal complication of which is the formation of bile duct and pancreatic duct strictures, which may be treated or prevented with the hydrogels. for treatment of ischemic bowel disease. The intestines are prone to the formation of scars or strictures when their blood supply is compromised. Compromised blood flow is called ischemia, and can be caused by many pathologies, including cardiovascular disease, atherosclerosis, hypotension, hypovolemia, renal or hepatic disease-induced hypoalbuminemia, vasculitis, drug-induced disease, and many others. The end stage result of all of these etiologies can result in intestinal strictures that block off the bowel and prevent its normal function. The present hydrogels may be used to treat or prevent ischemic bowel strictures, for treatment of radiation-induced intestinal strictures. Radiation therapy for cancer is associated with numerous morbidities, important among which is intestinal stricture formation. The present hydrogels may be used to treat or prevent radiation-induced intestinal strictures. [0079] In addition to making synthetic tissues, the hydrogels disclosed here also can be used to provide a coating for non-biological structures or devices to be used in surgery or otherwise for in vivo implantation, such as surgical instruments, or ceramic or metal prostheses. Such a coating would provide a barrier between the non-biologic device material and living tissue. The role of hydrogels as a barrier for non-biologic devices includes, but is not limited to: 1) prevention of absorption of macromolecules and/or cells on the surfaces of non-biologic devices, which can lead to protein fouling or thrombosis at the device surface; 2) presentation of a non-toxic, non-inflammatory, non-immunogenic, biologically compatible surface for devices made from otherwise non-biologically compatible materials; 3) compatibility with device function such as diffusion of glucose for a glucose sensor, transmission of mechanical force for a pressure sensor, or endothelization of a vascular graft or stent; 4) enhancement of device function, such as providing a charge barrier to an existing size barrier in a MEMS based artificial nephron; 5) incorporation into non-biologic devices of a viable cell population entrapped within an aqueous, physiologically compatible environment; and 6) inclusion of drugs or bioactive factors such as growth factors, anti-viral agents, antibiotics, or adhesion molecules designed to encourage vascularization, epithelization or endothelization of the device. [0080] Based on the foregoing, the hydrogels of the present invention may be used to provide a non-allergenic coating for a variety of implantable devices including an implantable glucose sensor for management of diabetes. In addition, the hydrogels may be used to provide: a charge barrier for the development of MEMS-based artificial nephrons; an aqueous, physiologically compatible environment in which embedded kidney cells such as podocytes can be incorporated into a MEMS-based artificial nephron design; and a coating for implantable MEMS devices designed for a variety of purposes including, but not limited to, drug delivery, mechanical sensing, and as a bio-detection system. [0081] The disclosed hydrogels, and particularly a hyaluronan-based hydrogel, also may be covalently attached to silicon-based devices, e.g. through first covalent attachment of the primary amine of tyramine to the silicon surface to provide a hydroxyphenyl coated surface chemistry. This may use the same chemistry used to bind DNA that has been modified with a free amine to silicon surfaces. The HA-based hydrogel then is covalently coupled to the hydroxyphenyl coated surface by the same peroxidase driven chemistry used in its preferred cross-linking mode described above. [0082] The hydrogels also can be used for coating non-biologic cardiovascular devices such as catheters, stents and vascular grafts. These would include devices made from materials conventionally not used because of their biological incompatibility, but which have superior design characteristics to those devices currently in use. Bioactive factors could be incorporated into the hydrogels to promote endothelization or epithelization of the hydrogel, and thus of the implanted device. [0083] A particularly promising application mentioned above is in the design and implementation of an implantable artificial glucose sensor for the treatment and management of diabetes. Effective glycemic control requires frequent measurement of blood glucose levels, which currently requires a pin prick (or "finger stick") to obtain a blood sample. There is tremendous clinical interest in a reliable, cost-effective method of blood glucose measurement and in preventing hypoglycemia, which is the cause of most severe life- threatening events. From a technological standpoint, microsensors have been very successful over the last decade in a wide variety of applications. The successful development of a biocompatible long term implantable glucose sensor would significantly impact routine monitoring of glucose levels by diabetic individuals and play a major contributory role in the further development of a bioartificial pancreas. [0084] A design of a sensor for use during cardiovascular surgery has been published, Clark IX!, Lyons C, "Electrode system for continuous monitoring in cardiovascular surgery," Annals of New York Academy of Science, 102:29-45 (1962).. Subsequently, efforts have been directed toward developing and testing an implantable device that could mimic the native glucose/insulin control system. Besides the obvious advantage of serving as part of a bioartificial pancreas, such a system could be coupled with telemetry hardware and thereby give the patient advance warning of hypoglycemia. [0085] Prior work on implantable glucose sensors generally follows one of two approaches. The first involves placing sensors into blood vessels such as the vena cava or the carotid artery. The second involves placing sensors subcutaneously. These sensors may involve a microdialysis probe or more commonly, an amperometric enzymatic-based transducer. It is believed the risk of thrombosis and hematogenous spread of infection mitigate against the long term use of intravascular sensors. While the exact relationship between blood and subcutaneous glucose concentrations is still being investigated, recent work suggests that mass transfer modeling methods can significantly improve the estimates of blood glucose levels that are based on subcutaneous data. Furthermore, there are significant advantages associated with subcutaneous sensors: clinical safety, ease of insertion and removal, ease of coupling these sensors to a telemetry system and cost. There is substantial evidence that subcutaneous placement of a glucose sensor will work and will lead to much longer life of the sensor than if it were to contact blood directly. [0086] However, a major problem in the design of any continuous glucose sensor for clinical use remains the long-term drift of the sensor, usually caused by fouling of the electrode when exposed to human tissue or the gradual loss of enzyme activity. The introduction of various membranes to act as a glucose or a hydrogen peroxide barrier has, in general, improved sensor performance but it has not resulted in long term stability. The much heralded membrane for this purpose, Nafion, rapidly deteriorates when implanted in the body. Introduction of an implant into subcutaneous tissue elicits both acute and chrome inflammatory responses. Together these result in a complexly orchestrated growth of new tissue which ultimately envelops the implant with a foreign body capsule (FBC). In the short term, it is likely that inflammatory cells metabolize glucose and thereby cause artifacts in the glucose readings. When discussing the problems with long-term use of subcutaneous sensors, experts maintain that the diminished response in vivo can be ascribed to the protein or cellular coating around the sensor which interferes with the mass transport of glucose. If suitable covering membranes for the sensor could be provided to exclude interfering substances or control coating or encapsulation with proteins and cells, the excellent performance in vitro may be matched in vivo. The use of the HA-based hydrogels described herein as a coating agent to both minimize the FBC and keep it away from the sensor membrane should prove a useful solution. [0087] The purpose is to control the tissue response to an implantable glucose sensor using a HA-coating on the sensor membrane. A sheath of HA-based hydrogel will give the sensor membrane "breathing space" by preventing proteins and cells from clogging the diffusion of glucose and oxygen into the sensor. Prior experience has indicated that HA and its derivatives are extremely biocompatible and as a consequence are used in situations where the host tissue response needs to be minimized (e.g., in eye implantation surgery). Thus, sensor performance should be enhanced in the long term when HA-based hydrogels are cast around sensor membranes as it relates to the development of an implantable glucose sensor with the long term perspective that such a sensor should result in improved blood glucose monitoring and ultimately improved quality of life for the diabetic population. In addition, the novel cross-linking structure of the HA-based hydrogels herein disclosed will ensure long-term maintenance of such a coating which will provide significant longevity to a subcutaneously implanted glucose sensor. [0088] Still another promising application is in the production of a bioartificial kidney for the treatment of end-stage renal disease (ESRD). The only current treatment options for ESRD patients are renal replacement therapy (all forms of dialysis) and transplantation. Transplantation is limited by the shortage of donor organs, and is complicated by the necessary and expensive life-long use of immunosuppressive drugs. Alternatively, although dialysis can prolong the life of ESRD patients, average life expectancy on dialysis is reduced by 50%, and the remaining quality of life is far from ideal. Repeated vascular access and handling of the patient's blood leads to frequent and sometimes life threatening infections. [0089] The functional unit of the kidney is the nephron. The nephron begins with a filtering structure, the glomerulus, which is of a tuft of capillaries surrounded by epithelial cells (podocytes) and supported by mesenchymal cells (the mesangium). The glomerulus is connected directly to the tubule of the nephron, a long tube lined with a single layer epithelium of polarized cells. The tubule cells function to salvage fluid, electrolytes and nutrients from the filtrate (by both intracellular transport and pericellular movement) concentrating the filtrate into urine. All nephrons connect into the collecting system, a network of epithelial-lined tubes, which has some additional reabsorptive properties, but primarily functions to direct the urine to the bladder. The filtration unit of the nephron, the glomerulus, consists of the endothelial cell of the capillary arteriolar wall, the podocyte surrounding the exterior of the capillary, and the glomerular basement membrane (GBM) sandwiched between the two cell types. The glomerular capillaries are some of the smallest vascular beds in the body, and the glomerular endothelial cells are specialized for their function by being fenestrated to allow direct contact of the blood plasma to the filtration barrier. Although these fenestrated endothelia do restrict the movement of leukocytes and very large molecules into the filtrate, the permselectivity of the filtration barrier is defined by the podocyte and the GBM. [0090] The GBM is a classic basement membrane structure composed of the prototypic molecules: type IV collagen (a3, a4, a5 heterotrimers), laminin (Laminin-11, a5, P2, yl heterotrimers), HS proteoglycans (perlecan and agrin) and nidogen (nidogen-1 and -2); as well as several additional ECM molecules including, collagen V, fibronectin, a CS proteoglycan (bamacan) and several small leucine-rich proteoglycans (biglycan, decorin, podocan). The GBM is synthesized by both the endothelial cell and the podocyte. Each cell produces a complete basement membrane which subsequently fuses during development to form one, double thickness basement membrane. The GBM has important functions in providing the appropriate microenvironment and substrata for the podocytes and endothelial cells. Without a normal GBM, both cell types lose their typical morphology and cellular differentiation characteristics, which subsequently destroys glomerulus function. The GBM also functions in filtration by restricting the movement of water and has some contribution to the size and charge selectivity, however, the majority of permselectivity is dictated by the podocyte. [0091] The podocyte is a highly specialized epithelial cell and has unique function in the glomerulus. The podocyte extends lamellipodia that wrap around the capillaries, branching into very fine interdigitations with other podocytes. On cross section, these interdigitated cellular extensions are called foot processes (FP) and the spaces between the FPs, where filtration occurs, are called slits. The podocyte synthesizes a macromolecular structure that spans the slit, the slit diaphragm (SD), which forms a bridge between two adjacent FPs. The molecular composition and structure of the SD is not fully understood. The SD appears to be a modified adherens junction containing additional podocyte-specific proteins, the most notable being nephrin. Nephrin extends from the plasma membrane of one FP and forms a homodimeric interaction with another nephrin molecule extending from the adjacent FP, creating a zipper-like structure when viewed in cross section by electron microscopy. How the SD and nephrin function as a permselective barrier is not known, but is currently a very active area of research. [0092] Biological microelectromechanical systems (bioMEMS) are a promising area of exploration for development of next generation bioartificial kidneys. Drug delivery systems, immunoisolators, and capillary networks, as well as precise control of cell differentiation and growth have been demonstrated for bioMEMS. The kidney is the first organ for which chronic substitutive therapy has been accepted, and application of the bioMEMS toolkit to treatment of ESRD is both evolutionary in the technology and revolutionary in the end product. Silicon micromachining technology has evolved such that structures with feature sizes on the order of 1 - 100 nanometers can be reliably produced in quantity. These dimensions are on the order of those for the glomerular slit diaphragm. The facility with which standard silicon bulk and surface micromachining technology permits microfluidic control, patterned deposition of cells and extracellular matrix proteins, and immunoisolation of cells, lends itself to tissue engineering of artificial organs. The engineering of nanoscale semiconductor filtration membranes could permit independent control and investigation of charge-size selectivity with the potential to lead to the tissue engineering of a bioartificial glomerulus and eventually a complete nephronal unit. [0093] One of the first components in the miniaturization of a bioartificial kidney is development of a nanofabricated hemofiltration membrane (NHM) from bioMEMS components. The NHM is intended to serve the hemofiltration function of the glomerulus in the nephron-like devices of a bioartificial kidney. NHM arrays can be fabricated using standard silicon micromachining techniques containing slit pores of approximately the dimensions of the glomerular slit diaphragm, and conducted experiments to demonstrate its size barrier characteristics similar to those of the glomerular basement membrane. The chemistries and hydrogels described in this patent application can be used to provide two additional and necessary components to the filtration characteristics of the NHM that are required for glomerular function. The first is a charge barrier component similar to those of the glomerular basement membrane. This would be provided by application of a layer of heparan sulfate (HS) based hydrogel. HS is a type of GAG similar to HA and CS. The second addition is inclusion of the podocytes, which are responsible for the majority of the filtration function of the glomerulus through the slit diaphragm. The podocytes would be applied to the surface of the HS-based hydrogel layer in a HA-based hydrogel layer, which would also serve to provide a layer of biocompatibility. The presence of the HS layer should facilitate proper matrix-cell interactions and stimulate the deposition of an appropriate basement membrane. [0094] The hydrogels described herein, including but not limited to tyramine-based hyaluronan hydrogels, also can be used as research and clinical reagents. One promising application is controlled or extended release drug delivery. In this application, the drug can be trapped within a sphere or other suitable shape of hydrogel material composed of a central spherical or other shaped core of hydrogel formulated at a relatively high macromolecular concentration (and thus lowest porosity), onto which concentric spherical layers of hydrogel are coated, each successively coated layer being formulated of a progressively lower macromolecular concentration (and thus higher porosity). Release of the drug is then controlled by the rate of hydrogel degradation if so engineered, binding of the drug to the hydrogel scaffold and diffusion of the drug through the scaffold pores. The hydrogel sphere then is implanted into a patient at an appropriate location to effect extended release of the drug. [0095] Targeted drug deliver also can be achieved through an affinity-based strategy based on designed affinity of drug laden hydrogel particles to specific tissue and cell types. To this end, the hydrogels can be used as an affinity-based medium for the selective binding and thus purification of specific cell populations through incorporation of targeted cell binding molecules within the hydrogels during or prior to cross-linking. Once a select cell population is bound to the hydrogel affinity-based medium they could be released for further investigation, or directly entrapped while bound to the hydrogel affinity-based medium into other formulations of the hydrogels for other tissue engineering or clinical applications. [0096] Such an affinity-based medium also can be used for the selective binding and purification of hyaluronan binding proteins. As the entire medium can be made solely of hyaluronan with no other support material background binding should be quite low. By using other materials as the scaffold material (such as aggrecan) other affinity-based media can be prepared for purification of molecules that selectively bind to those scaffold materials. [0097] Such an affinity-based medium also can be used for selective binding and purification of specific macromolecules or cell populations through incorporation of protein A within the hydrogels during cross-linking. Antibodies specific to the macromolecule or cell population of interest can then be used to coat the protein A infused hydrogels with the antibodies optimally oriented with their antigen binding (Fab) arms directed outward and their constant (Fc) domain bound to the protein A. Once a select cell population or macromolecule is bound to the protein A hydrogel, it could be released for further investigation, or directly entrapped while bound to the protein A hydrogel into other formulations of the hydrogel for other tissue engineering or clinical applications. Alternatively antibody could be directly incorporated into the hydrogels. [0098] The disclosed hyaluronan-based hydrogel materials also have utility as a diagnostic for the presence of hyaluronidases which can be predictive of the metastatic potential of certain cancers; e.g. by coating of a biopsy slide with hyaluronan hydrogel and measurement of the extent and localization of the loss of intrinsic fluorescence of the hydrogel material due to its dityramine cross-links as the hydrogel is digested by endogenous hyaluronidases. By using other materials as the scaffold material (such as aggrecan) other degradative enzymes could be detected such as metaloproteinases. [0099] Further aspects of the invention will be understood in conjunction with one or more of the following examples, which are provided by way of illustration. EXAMPLES EXAMPLE 1 [0100] Experimental quantities of tyramine-substituted hyaluronan hydrogels having dityramine cross-links according to the invention have been prepared as follows. HA is dissolved at 1 mg/ml based on hexuronic acid in 250 mM 2-(N-morpholino)ethanesulfonic acid (MES), 150 mM NaCl, 75 mM NaOH, pH 6.5 containing a 10-fold molar excess of tyramine relative to the molar concentration of HA carboxyl groups. Tyramine substitution onto the carboxyl groups is then initiated by the addition of a 10-fold molar excess of EDC relative to the molar concentration of the HA carboxyl groups. A l/10th molar ratio of N- hydroxysuccinimide (NHS) relative to the molar amount of EDC is added to the reactions to assist the EDC catalyzed amidation reaction by formation of active esters. Reactions are carried out at room temperature for 24 hours, after which the macromolecular fraction is recovered from unreacted small molecular weight reactants such as tyramine, EDC, NHS, and MES by exhaustive dialysis versus 150 mM NaCl and then ultrapure water followed by lyophilization. After lyophilization, the tyramine-substituted HA (T-HA) product is dissolved to working concentrations of between 5 and 100 mg/ml in PBS (which is a buffer compatible with cell suspension, in vivo tissue contact, and the cross-linking reaction) to provide various concentration preparations depending on the desired rigidity of the final hydrogel. Alternatively, the solvent can be any other suitable solvent besides PBS that will not substantially negatively impact the enzyme activity and that will not interfere with cross- linking reaction via selective uptake of free radicals generated by the enzyme. Suitable alternative solvents include water, conventional biological tissue culture media, and cell freezing solution (generally composed of about 90% blood serum and about 10% dimethyl sulfoxide). Prior to suspension of cells (see Example 5) or contact with tissues in vivo, the T- HA should be filtered through a 0.2 urn filter. Next, tyrarnine-tyrarnine linking is carried out by adding 10 U/ml of type II horseradish peroxidase (HRP) to each T-HA preparation. Cross-linking is initiated by the addition of a small volume (1-5 u.1) of a dilute hydrogen peroxide solution (0.012%-0.00012% final concentration) to yield the final hydrogel with desired rigidity. For preparation of larger quantities or volumes of a desired hydrogel, quantities of reagents provided in this paragraph could be scaled up appropriately by a person of ordinary skill in the art. EXAMPLE 2 [0101] An experiment was conducted to determine the degree of tyramine substitution (and consequent dityramine cross-linking) for a T-HA macromolecular network according to the invention. Initially, three formulations of (uncrosslinked) tyramine-substituted hyaluronan (T-HA) were prepared as described above, designated OX, IX or 10X. The OX formulation was prepared using no EDC (i.e. containing no carbodiimide), meaning there was no carbodiimide present to mediate the reaction for creating an amide bond between the NH2 group on tyramine and a CO2H group on the HA molecules. Thus, the OX formulation can be considered a control. The IX formulation contained a 1:1 stoichiometric ratio of EDC based on the quantity of CO2H groups present on the HA molecules in the reaction mixture. The 10X formulation contained a 10:1 stoichiometric ratio (or 10-fold excess) of EDC based on the quantity of CO2H groups present on the HA molecules in the reaction mixture. In all three formulations, a stoichiometric excess of tyramine was provided relative to the quantity of CO2H groups on HA. In all three formulations (OX, IX and 10X) the reactants and the appropriate amount of EDC for the formulation were combined in a vial and agitated to facilitate the tyramine-substitution reaction. All three formulations were allowed to react for 24 hours at room temperature, after which the vial contents were dialyzed to remove unreacted tyramine molecules, EDC and acylurea (EDU) byproducts of the reaction. These molecules were easily separated from HA and any formed T-HA molecules through dialysis due to the relatively small size of tyramine, EDC and EDU compared to macromolecular HA. Once unreacted tyramine and EDC were removed, the remaining contents for each formulation were analyzed to determine the rate of tyramine substitution relative to the total number of available CO2H sites present on HA molecules. [0102] Tyramine exhibits a UV absorbance peak at 275 nm, making the degree of tyramine substitution easily detectible against a tyramine calibration curve. Based on UV- spectroscopic analysis of the above three T-HA formulations, it was discovered that the HA- tyramine substitution reaction carried out with no EDC present (formulation OX) resulted in substantially zero tyramine substitution onto the HA molecules. This confirmed the importance of using a carbodiimide reaction pathway in the tyramine substitution reaction. However, the tyramine absorption in the T-HA formulation prepared using a 1:1 EDCrCOjH stoichiometric ratio in the tyramine substitution reaction (formulation IX) resulted in a tyramine substitution rate of about 1.7% relative to all available CO2H groups on the HA chains. The 10X formulation (10:1 EDC:CO2H ratio) resulted in about a 4.7% substitution rate. [0103] Subsequently, hydrogen peroxide and horseradish peroxidase (HRP) were added to each of the three dialyzed HA/T-HA formulations (OX, IX and 10X) at 5 mg/mL and the resulting formulations were allowed to react to completion. After reaction in the presence of peroxide and HRP, it was observed that the OX formulation remained entirely liquid, having a strong meniscus; no gel formation was observed, confirming the fact that no or substantially no tyramine substitution had occurred when no EDC was used in the tyramine substitution reaction. For the IX formulation, only a very weak meniscus was observed and the contents of the vial had gelled, confirming that both tyramine substitution and cross-linking had occurred. For the 10X formulation, a relatively rigid gel had formed, and in fact had shrunk relative to the initial volume of fluid in the container, leaving a quantity of liquid (having a meniscus) on top. The gel prepared from the 10X formulation (having a 4.7% tyramine substitution rate) was much firmer and more rigid than that from the 1X formulation having a 1.7% tyramine substitution rate. [0104] The dityramine structure exhibits a blue fluorescence on exposure to UV light. The products of each of the above formulations were exposed to UV light to detect the presence of dityramine cross-links. As expected, both the IX and 10X hydrogels exhibited blue fluorescence (the 10X hydrogel fluorescence being more intense than that of the IX hydrogel), while the OX formulation exhibited no blue fluorescence at all. This confirmed the presence of dityramine cross-links in both hydrogels, and that the occurrence of dityramine in the more rigid hydrogel (10X) was greater than in the less rigid hydrogel (IX). [0106] The overall result was that the importance of the carbodiimide-mediated reaction pathway was demonstrated, and it was confirmed that the relative rigidity of a hydrogel formed from a cross-linked T-HA network is proportional to the degree of dityramine cross-linking, which is in turn proportional to the degree of tyramine-substitution - onto HA. It was quite a surprising and unexpected result that even a 1.7% tyramine- substitution rate (and subsequent cross-linking rate to form dityramine links) provided a suitably firm T-HA gel (or hydrogel). A 4.7% substitution (and cross-linking) rate resulted in even a firmer T-HA gel. Also surprising was that a ten-fold stoichiometric excess of carbodiimide (EDC) relative to the quantity of carboxylic acid groups present in the reaction mixture (formulation 10X) resulted in only about a 4.5-4.7% tyramine substitution rate, yet stable and cohesive tyramine cross-linked T-HA networks were nonetheless achieved. [0107] This means that the majority of the carboxylic acid groups on the HA molecules are unsubstituted and not tyramine cross-linked, essentially remaining the same as in the native HA molecule, yet the resulting network is a cohesive and stable hydrogel. Therefore, when used as a cartilage substitute in vivo, because a majority of the HA molecules in the invented T-HA network or gel are essentially unaltered compared to HA in normal cartilage, it is believed that the body's native metabolic pathways (aided or unaided by cells provided within the T-HA network) may recognize the invented network as native biologic material, and will be able to carry out ordinary synthesis and metabolism functions with respect thereto. In addition, it is noted that HA is a highly ubiquitous material in the body, and is non-immunogenic in humans. As a result, the cross-linked macromolecular network, comprised a majority of unaltered native HA, will have substantial application in a wide variety of tissue engineering applications where it is desirable or necessary to provide synthetic tissue in a human body. This represents a significant advance over the state of the art. Therefore, quite surprisingly, a high degree of tyramine substitution, e.g. greater than about 10-20%, maybe undesirable; the above described experiments demonstrated that such high degrees of substitution are unnecessary to provide a suitable T-HA network. Preferably, a dihydroxyphenyl (e.g. dityramine) cross-linked polycarboxylate (e.g. HA) network has a hydroxyphenyl (tyramine) substitution rate of less than 50, preferably less than 40, preferably less than 30, preferably less than 20, preferably less than 15, preferably less than 10, preferably less than 9, preferably less than 8, preferably less than 7, preferably less than 6, preferably less than 5, percent based on the total quantity of CO2H groups present on the polycarboxylate (HA) molecules. EXAMPLE 3 [0108] Conventionally, it has been believed that natural cartilage exhibits its viscoelastic properties and its ability to resist deformation and absorb compressive loads principally as a result of the repulsive forces between negatively charged SO42- groups on adjacent chondroitin sulfate chains present in the aggrecan matrix. An experiment was performed to determine the efficacy of various macromolecular networks within the scope of the invention to resist deformation and absorb compression compared to natural cartilage. In particular, three such networks were prepared, respectively, composed of the following: 1) dityramine cross-linked HA molecules (T-HA); 2) dityramine cross-linked chondroitin sulfate molecules in the form of aggrecan (T-Aggrecan); and 3) a composite material composed of 50% T-HA and 50% T-Aggrecan. Formulations of uncross-linked T-HA and T- Aggrecan were prepared and purified as in Example 1, each having a tyramine substitution rate of about 5%. From these T-HA and T-Aggrecan formulations, five different concentrations of the T-HA alone, T-Aggrecan alone and a 50:50 mixture of T-HA and T- Aggrecan were prepared: Concentration 1: 6.25 mg total T-GAG / mL water Concentration 2: 12.5 mg total T-GAG / mL water Concentration 3: 25 mg total T-GAG / mL water Concentration 4: 50 mg total T-GAG / mL water Concentration 5: 100 mg total T-GAG / mL water. [0109] The notation T-GAG is used herein to embrace both T-HA and T-Aggrecan. Though aggrecan technically is not a glycosaminoglycan (GAG), for purposes of this example T-GAG nonetheless is defined to embrace both T-HA and T-Aggrecan hydrogels. Each of the above preparations was then reacted in the presence of hydrogen peroxide and horseradish peroxidase, also as in Example 1, to form dityramine cross-links between the T- GAG molecules and provide respectively Hydrogels 1,2, 3,4 and 5 for each of the three material compositions. Each of the fifteen hydrogels (five concentrations for each of the three material compositions) was found to be a stable and substantially coherent material with the physical properties of each hydrogel varying relative to the concentration of T-GAG in the preparation from which it was made. For example, qualitatively T-HA Concentration 1 resulted in T-HA Hydrogel 1 having rigidity and rheological properties comparable to that of Vaseline or jelly; the hydrogel was stable and coherent yet could be caused to flow or spread on application of an external force, e.g. from a spatula or other conventional tool. T-HA Hydrogel 1 exhibited excellent adhesive properties making it an ideal candidate for a nonallergenic coating material for surgical instruments during surgery, e.g. ophthalmologic surgery. T-HA Hydrogel 2 was more rigid than T-HA Hydrogel 1 due to the greater concentration of T-HA in the preparation from which it was made, and the consequent predicted decrease in intramolecular cross-linking and increase in intermolecular cross- linking associated with increased T-HA concentration. T-HA Hydrogel 2 exhibited rheological and rigidity properties characteristic of gelatins, with a degree of viscoelastic reboundability on external loading. On greater loading, T-HA Hydrogel 2 was found to break up into smaller pieces instead of flowing, also characteristic of a gelatinous material. T-HA Hydrogel 3 had the properties and consistency of a dough or malleable paste, also not flowing on application of an external loading force. This material also exhibited substantially greater viscoelastic properties compared to T-HA Hydrogels 1 and 2. T-HA Hydrogel 4 was a highly rigid and coherent gel that strongly resisted breaking up on application of an external loading force. T-HA Hydrogel 4 was a highly resilient rubber-like composition that actually generated substantial springing force upon sudden compression (e.g. dropping onto the floor). This ability of T-HA Hydrogel 4 to generate such a springing force in response to a sudden compression may make this material ideal for certain joint replacement/repair applications where the joint undergoes repeated and periodic compressional loading (e.g. the ankle joint). In addition to the properties described for T-HA Hydrogel 4, T-HA Hydrogel 5 had cartilage- like properties with both the appearance of articular cartilage and the feel of cartilage upon cutting with a surgical blade. [0110] Confined compression tests were performed to quantitatively determine the compressive mechanical properties of the fifteen different hydrogels. A custom built polycarbonate confining chamber, and porous polypropylene filter platen (20 µm pores, 20% porosity) were used to perform the confined compression testing. Five cylindrical plugs (7.1 mm in diameter, approximately 3 mm in thickness) at each hydrogel concentration for each of the three material compositions were made using the confining chamber and the freeze- thaw technique described in Example 4 below. The following testing protocol was followed for a series of stress relaxation tests in confined compression. All testing was performed using an Instron 5543 machine under computer control, which recorded the time- displacement-load data at a frequency of 10 Hz. A ±5 N or ±50 N load cell (Sensotec) was used to monitor load throughout each test. A step of 30 µm (30 µm /sec), representing 1 % strain, was applied until the sample reached equilibrium. This was defined as a relaxation rate that slowed to less than 10 mN min-1, at which time the next step was automatically started, until 20 cycles (representing approximately 20% strain) were completed. The thickness of each sample tested in confined compression was determined mechanically, by measuring the displacement at which the compressive response initiated relative to the bottom of the chamber as measured with the Instron 5543 machine. The measured thickness was used to calculate the strain percentage for each step. [0111] The compressive mechanical properties of the fifteen hydrogels were determined as described in the preceding paragraph. Load data was normalized by sample cross-sectional area (39.6 mm2) to compute stress. The equilibrium stress was plotted against the applied strain for each material formulation. The aggregate modulus at each step was defined as the equilibrium stress divided by the applied strain. For each material, the aggregate modulus was defined as the slope of the equilibrium stress-strain data in the most linear range. Figs. 4a, 4b and 4c display the equilibrium compression behavior for the five concentrations of T-HA, T-Aggrecan and 50:50 T-HA/T-Aggrecan composite hydrogels, respectively. All fifteen hydrogels were testable in confined compression, and demonstrated characteristic stress relaxation responses typical of biphasic materials (such as cartilage). The aggregate moduli for the 6.25 mg/ml and 12.5 mg/ml T-GAG hydrogels were 1-2 orders of magnitude lower than articular cartilage. The 25 mg/ml T-GAG hydrogels, as well as the 50 mg/ml T-aggrecan hydrogel, displayed aggregate moduli on the order of, but at least 30% lower than that of articular cartilage. All the 100 mg/ml T-GAG hydrogels, as well as the 50 mg/ml T-HA and the all the composite hydrogels, displayed aggregate moduli, equal to or exceeding reported literature values for articular cartilage. These data demonstrate the ability to characterize hydrogels using standard mechanical assays, and to generate hydrogels with similar mechanical properties to a wide variety of tissues including that of articular cartilage, using a variety of glycosaminoglycans as the hydrogel scaffold material. [0112] The aggregate moduli for the five concentrations of the T-HA, T-Aggrecan and composite materials composed of 50% T-HA and 50% T-aggrecan are summarized below in Table 1. [0113] Fig. 4d shows the measured aggregate modulus as a function of concentration for the T-HA, T-Aggrecan and composite hydrogels. As the concentration of the T-HA hydrogels increases, a plateau is reached for the aggregate modulus while the T-Aggrecan hydrogels display a linear relationship. Interestingly, the composite hydrogels show a relationship indicative of an exponential increase in compressive properties as concentration increases. This indicates that the moduli of other hydrogel materials can be predicted by further exploring and modeling these relationships. [0114] Based on the above experiments it was surprisingly and unexpectedly discovered that a dityramine cross-linked GAG network (HA or aggrecan) will produce a coherent hydrogel material whose rigidity and other physical (rheological) properties can be tuned by varying the T-GAG concentration prior to cross-linking the tyramine groups to suit a particular application. The coherence and elastic properties of these hydrogels was observed even absent any (or substantially any) SO42- groups in the network to supply the charge-to-charge repulsive forces to generate the material's compression resistance and elasticity. This was a highly surprising and unexpected result with substantial positive consequences in tissue engineering applications. Hyaluronan is a highly ubiquitous and non- immunogenic molecule found in humans. Therefore, hydrogels comprised of dityramine cross-linked hyaluronan networks can be used to provide suitable tissue replacement materials that can be implanted within a human body, whose rigidity can be tuned based on the application as evidenced by this example. As these materials can or will be composed of predominantly unaltered hyaluronan which is non-immunogenic, the hydrogels should result in zero or substantially zero immune response. This is an important advantage over many conventional tissue engineered 'materials whose formation chemistries prevent their application in vivo due to harsh reaction conditions or reagents, and whose final chemical structures are more likely to induce an immune response. EXAMPLE 4 [0115] A number of methods of preparing hydrogels such as those described in Example 3 have been developed to cast or form the hydrogel into a predetermined three- dimensional shape. This is important for myriad tissue engineering applications where it is necessary to provide artificial tissue material to fill a native tissue defect or void within a patient. [0116] A first method is to employ an in situ forming technique where the hydrogel is formed in place, i.e. in position and in the shape of its final application and structure. The in situ formation method has been carried out experimentally as follows. Tyramine-substituted hyaluronan (T-HA) was prepared via the carbodiimide-mediated pathway described herein. Following dialyzation to remove unreacted tyramine, EDC, NHS, etc., and dissolution at the desired concentration in PBS (see Example 1 above), a small quantity of horseradish peroxidase enzyme was added to the T-HA liquid preparation to form a first solution. This first solution was provided into a laboratory container (to simulate an in vivo situs) having a specific interior geometry. Subsequently, a second solution was prepared containing very dilute hydrogen peroxide (0.012%-0.00012% final concentration). A small volume of this second solution relative to the first solution was then injected into the container already containing the first solution to initiate the dityramine cross-linking reaction to yield the hydrogel. Hydrogels prepared by this technique have been prepared having varying rigidity and rheological properties as described above in Example 3, and conformed well to the interior surface contour of the container in which they were formed. Because the principal reagents (H2O2, hyaluronan and peroxidase) are either nonallergenic or diffusible molecules, and because the cross-linking reaction proceeds under metabolic conditions of temperature and pH, this technique can be performed in vivo at a surgical situs in a patient as a surgical procedure to produce a defect-conforming hydrogel. This method is particularly attractive for reconstructive facial surgery in which the uncross-linked T-HA preparation (with peroxidase) can be injected and manipulated subcutaneously by the surgeon to produce the desired facial contours and then the hydrogel subsequently cross-linked by injection of a it freezes into^a solid ice form conforming to the shape and contour of the inner mold surface. However, the silicone mold, having a glass transition temperature below -80°C, remains soft and malleable and the solid ice form of the first solution is easily removed. Because the first solution expands as it freezes, suitable mechanical hardware should be used to ensure the silicone mold does not deform or expand as the solution freezes. Preferably, port holes are provided in the mold to allow for expansion and discharge of the first solution as it expands during the freezing process. [0119] Once the solid ice form of the first solution has been demolded, minute defects or flaws in the three-dimensional structure can be repaired by carving with a suitable tool, and more of the liquid first solution can be added to fill surface voids, which liquid instantly freezes on contact with the solid ice form. Also, the ice form can be placed back on the dry ice surface if desired to ensure uniform temperature and freezing of any added first solution material. Once the three-dimensional shape of the ice form has been perfected, it is immersed in a liquid peroxide solution to initiate thawing of the frozen water and dityramine cross-linking from the outside-in. This is possible do to the rapid kinetics of the cross-linking reaction. Cross-linking is determined to be complete once the last remaining frozen water has melted at the center of the forming hydrogel form, which can be easily observed because the forming hydrogel is substantially clear. [0120] Very successful experiments have been performed according to this freeze- thaw technique to produce a solid hydrogel in the shape of a human ear. Other structures that could be formed by this method, such as intervertebral discs, meniscus, etc. will be evident to those skilled in the art. It should be noted in this freeze-thaw technique, the threshold glass transition temperature of-80°C for the mold material is selected to correspond roughly with the surface temperature of solid CO2 (dry ice), to ensure the mold material does not become brittle when the first solution is frozen to produce the solid ice form. However, if another cooling material, other than CO2 is used, then the threshold glass transition temperature for suitable mold materials may be adjusted accordingly. [0121] For the three methods of hydrogel formation described above, the first solution contained both the peroxidase and T-HA, while the second solution contained the peroxide. While it may be possible to switch the peroxidase and peroxide in the first and second solutions respectively, it is less preferred to provide the peroxide in the first solution with the T-HA. This is because once the peroxide, peroxidase and T-HA are combined, the T-HA rapidly begins to form a cross-linked macromolecular network. If the peroxidase (which is a macromolecular molecule) is not already uniformly distributed with the T-HA it may be unable or substantially hindered from diffusing through the pore structure of the forming hydrogel to facilitate uniform cross-linking throughout the entire T-HA/peroxide solution. The result could be non-uniform and/or incomplete cross-linking of the T-HA and a non- uniform hydrogel. Conversely, the relatively small peroxide molecule (hydrogen peroxide is only one oxygen atom larger than water) can diffuse through the hydrogel pore structure with relative ease, resulting in a uniform hydrogel structure. [0122] In addition, the macromolecular size of the peroxidase allows it to be similarly retained as the T-HA within porous molds that are only porous to small molecular weight peroxides which easily and uniformly diffuse through both the molds and newly forming macromolecular networks (i.e. hydrogels). For these reasons it is preferred to start with the peroxidase uniformly distributed with the T-HA in the first solution, and to provide the peroxide separately in the second solution. [0123] A fourth method is an alternating sprayed or brushed layering technique. The first solution is prepared as described above and contains both the peroxidase and T-HA. However, the second solution not only contains the peroxide as described above, but also T- HA at the same concentration as in the first solution. Then a thin layer of the first solution is applied at the desired location (in situ) followed by an overlying thin layer of the second solution. This procedure is repeated such that alternating layers of the first and second solutions are successively applied until the defect or application situs has been completed. The very thin alternating layers of the first and second solutions promote virtually complete dityramine cross-linking ensuring a highly coherent final hydrogel having the desired rheological properties based on the initial T-HA concentration of the two solutions. The thin nature of the layers is desirable to ensure that free radicals produced by the peroxidase in the first solution layers are able to penetrate completely adjacent second solution layers and complete cross-linking independent of peroxidase diffusion into the second solution layer (see above). T-HA is included in both solutions to ensure uniform T-HA concentration throughout the final hydrogel. This technique has been performed in laboratory bench experiments and has provided contour-conforming and volume-filling coherent hydrogels. This technique is highly applicable where it is desired to provide a thin, but variable layer of tyramine cross-linked HA, such as on the surface of a denuded osteoarthritic joint in which little if any native healthy cartilage remains in the patient at the implant site. [0124] All four of the above techniques have been described with respect to dityramine cross-linked hyaluronan, however it will be understood that other combinations within the scope of the present invention (other dihydroxyphenyl cross-linked macromolecules, such as polycarboxylates, polyamines, polyhydroxyphenyl molecules and copolymers thereof) can be molded via the above techniques. EXAMPLE 5 [0125] Rat chondrocytes were embedded in (cross-linked) T-HA hydrogels to measure their ability to survive the cross-linking reaction. Isolated chondrocytes were suspended in the 1.7% and 4.7% T-HA hydrogels described in Example 2 by providing these live cells to the first solution to be co-dispersed with the T-HA and peroxidase, followed by introduction of the peroxide-containing second solution to initiate dityramine cross-linking. The chondrocyte-embedded 1.7% and 4.7% T-HA hydrogels exhibited uniformly distributed chondrocytes with the optical clarity of the gels allowing visualization throughout the gel. Glucose utilization was used as an indicator of cell viability after cross-linking to form the hydrogels as chondrocytes are voracious with respect to glucose consumption, depleting the medium of glucose in less than 24 hours. The results showed that chondrocytes embedded in T-HA hydrogels showed essentially the same glucose consumption profile over 24 hours as the same chondrocytes cultured in monolayer (Fig. 5). This continued for up to 7 days indicating that the cells were alive and metabolically active. Medium glucose was measured by standard hexokinase assay. [0126] Fluorescent images of frozen sections of T-HA hydrogels containing both chondrocytes and cartilage tissue were also generated. HA samples from both the hydrogel scaffold and cartilage matrix were visualized by fluorescent staining with biotinylated HA binding protein (b-HABP) reagent while cell nuclei were visualized with standard DAPI stain. The b-HABP reagent is prepared from purified cartilage aggrecan (the G1 domain only) and link protein, and recognizes and irreversibly binds to stretches of native HA equivalent to those normally bound by aggrecan and link protein in cartilage. The results showed a more intense staining of the T-HA hydrogel with b-HABP than the cartilage as the hyaluronan in the tissue is already occupied by native aggrecan and link protein. No visible distinction could be seen between the T-HA scaffold of the hydrogel and the matrix of suspended cartilage tissue suggesting seamless integration. These results demonstrated the feasibility of maintaining the viability of chondrocytes during the hydrogel cross-linking reactions, and the ability of the hydrogel to integrate seamlessly into existing cartilage matrix, both of which are advantageous for application to cartilage repair. The results also demonstrated that sufficient stretches of the T-HA remain chemically unaltered, and available for binding by newly synthesized aggrecan and link protein in situ. The results also demonstrated that oxygen, carbon dioxide, glucose and insulin are diffusable through T-HA hydrogels according to the invention at a rate that is not limiting to chondrocyte metabolism, which is important not only to the development of cartilage substitutes but to other applications such as glucose sensor design and development of an artificial kidney. [0127] In order to include cells such as chondrocytes in hydrogels molded into intricate anatomical shapes using the freeze/thaw technique described in Example 4, it is desirable that the enzyme driven cross-linking reaction proceed in the presence of standard cell freezing solutions such as those containing 10% dimethylsulfoxide (DMSO)/90% fetal bovine serum (FBS). This has been demonstrated in the laboratory for all of the T-HA hydrogel formulations described in Example 3. The ability to directly incorporate a solution containing 90% FBS also demonstrates the ability to include bioactive factors such as growth factors, hormones and factors controlling cell differentiation, as these are normal components ofFBS. EXAMPLE 6 [0128] An experiment was conducted whereby a T-HA hydrogel as described hereinabove was implanted into Yucatan minipigs in order to repair articular cartilage defects. Following is a description of that experiment, including the experimental methods and results obtained, after a brief discussion of the background for this application. Background Tissue Description [0129] Articular Cartilage Structure and Function - As discussed above, articular cartilage is the resilient load-bearing tissue that forms the articulating surfaces of diarthrodial joints. It absorbs mechanical shock and deflects or spreads applied load over greater surface area of subchondral bone. It consists primarily of a large extracellular matrix (ECM) with a sparse population of highly specialized cells (chondrocytes) distributed throughout the tissue. The primary components of the ECM are water, cartilage aggregates and type II collagen. Cartilage aggregates are composed of hyaluronan (HA), aggrecan (the large cartilage-specific proteoglycan), and link protein (LP), a small glycoprotein. Aggrecan contains a central core protein to which is attached ~ 100 chondroitin sulfate (CS) chains. The core protein has three globular domains with the N-terminal globular 1 (G1) domain having binding sites for both HA and LP. LP has sequence homology to the G1 domain of aggrecan, and contains binding sites for both HA and the G1 domain of aggrecan. Each cartilage aggregate is composed of a single HA chain, to which are attached hundreds of aggrecan/LP duplexes. These large cartilage aggregates are trapped at one fifth of their free solution volume within a tight meshwork of type II collagen fibers, which resist further swelling. This molecular architecture contributes to the tissues mechanical properties and function as described below. [0130] Swelling Pressure - The HA and CS chains in cartilage aggregates contain repeating carboxyl and/or sulfate groups. In solution, these groups become ionized (COO' and SO3"), and in the physiologic environment they require positive counter ions such as Na+ to maintain overall electroneutrality. These free-floating ions within the interstitial water are present at a higher concentration than that found in the surrounding fluids (i.e. synovial fluid) giving rise to an osmotic pressure (Donnan pressure). In cartilage, ions are prevented from flowing out of the tissue along the concentration gradient by the fixed nature of their negative counter ions (i.e. the COO" and SO3" groups on the HA and/or CS chains), and the need to maintain electroneutrality. Water flow into the tissue to equilibrate the concentration gradient is resisted by the inextensible nature of the collagen meshwork preventing further swelling. [0131] Alternatively, tight cartilage aggregate packing causes the fixed-negative charge groups to be spaced only 10 to 15 angstroms apart, resulting in strong charge-to- charge repulsive forces (electrorepulsive forces). As with the Donnan effect, the tendency to swell to lessen these repulsive forces is resisted by the inextensible nature of the collagen meshwork. When compressed, the distances between charge groups decrease, thus increasing the charge-to-charge repulsive forces and increasing the free-floating positive counter ions concentration. Thus both the Donnan and electrorepulsion effects are intensified by compression. Both effects contribute to the swelling pressure of articular cartilage and its ability to resist deformation and absorb compressive loads. [0132] Stress Shielding Effect - Articular cartilage is often described as a viscoelastic, biphasic material, composed of a solid phase (cartilage aggregates, collagen, etc.) and a fluid phase (water and dissolved ions). The macromolecular architecture of the ECM of articular cartilage functions to deflect applied forces during loading from the wear susceptible solid phase of the tissue to the wear resistant fluid phase or water. This stress shielding occurs due to the elegant design of the cartilage ECM which produces a material with very low permeability creating a drag during interstitial fluid flow. Interstitial fluid pressure is generated during compressive loading, and during dynamic loading, is the primary force responsible for supporting the applied load with matrix compression a minor factor. During compression, the porosity is reduced further, which increases the already high frictional drag forces. The load support is gradually transferred from the fluid phase (as the fluid pressure dissipates) to the solid phase. Typically, for normal cartilage, this equilibration process takes 2.5 to 6.0 hours to achieve. Thus, load support through fluid pressurization predominates within the tissue. Need for Synthetic Material [0133] Increased water content and decreased proteoglycan content are the most apparent early changes in osteoarthritic cartilage. These changes reflect an increase in tissue permeability. Increased permeability diminishes the fluid pressurization mechanism of load support in cartilage (stress shielding), requiring the collagen-aggrecan solid matrix to bear more load, which may be an important contributing factor in the development and progression of cartilage degeneration. Bioartificial cartilage substitutes that do not mimic the low permeability of normal, healthy articular cartilage may be predisposed to degeneration by a similar mechanism. [0134] One of the most difficult challenges facing orthopedic surgeons is treating patients who have suffered focal cartilage lesions, but are too young or too active for a total joint replacement. These localized cartilage defects can be very debilitating. Restoring these localized areas without a total joint replacement would be a preferred approach with significant benefits including reduced surgical requirements, shorter recovery times, lower cost, and slowing or arresting the further degradation of the load bearing surface. [0135] This example demonstrates the application of a tyramine-substituted HA (T- HA) hydrogel for repair of this type of localized cartilage defect. Experimental Description Design of Extracellular Matrix Material having Desired Properties [0136] Natural articular cartilage has the elastic as well as physical and chemical properties described above, which impart its unique ability to absorb mechanical loads and to deflect impact loads away from the subchondral bone. To produce a suitable synthetic cartilage material made from a T-HA hydrogel as disclosed herein, it was important to design the macromolecular network for the hydrogel so as to emulate those properties as nearly as possible through judicious selection of reagent concentrations, cross-linking conditions, incorporated living cells as well as other molecules, etc. [0137] The results from confined compression testing of T-HA hydrogels (see Example 3 above) provided an initial basis for the synthesis of an appropriate synthetic T-HA material having properties matched to those measured for normal articular cartilage. They also illustrate the spectrum of material properties that can be manufactured from a single formulation of T-HA. Based on those data, a T-HA hydrogel composed of, inter alia a macromolecular network of dityramine cross-linked hyaluronan molecules was selected based on the following criteria in order to produce a synthetic implantable cartilage material. [0138] A material composed solely of HA was chosen because hydrogel compositions with the compressive properties of cartilage could be formed using HA alone as the scaffold material (Example 3), while avoiding possible host response to the protein component of aggrecan if it were used as a scaffold material. Reaction conditions (tyramine/EDC ratio) were chosen to produce a percent tyramine substitution of 5% (Example 2) as this provided sufficient cross-linking to produce a material with the compressive properties of cartilage (Example 3) while maintaining the majority of the native HA structure. HA was substituted with ~5% tyramine, as described in Example 1, except that the HA was dissolved at 5 mg/ml rather than 1 mg/ml to conserve reagents. The absolute concentration of all other reagents remained the same so that the tyramine and EDC were at a 2-fold rather than 10-fold molar excess based on the molar concentration of HA carboxyl groups. A concentration of 125 mg/ml of HA in sterile saline was chosen as this concentration in saline had a compressive aggregate modulus most closely resembling that of articular cartilage (saline data not shown). It was also the concentration deemed most appropriate based on the experience of our clinician collaborator. Peroxidase was added at 10 U/ml prior to application in an in situ cross-linking protocol as described below. The in situ cross-linking protocol was chosen as it provided the best opportunity for integration with surrounding cartilage matrix. It also allowed easy and complete filling of the surgically produced cartilage defect without the need to know or measure the defect's exact dimensions. Pre-cast (in vitro cross-linked) plugs would either require exact dimensions or sculpting of pre-formed shapes to fit the defect. No cells or bioactive factors where added in this experiment as this experiment was intended to evaluate the hydrogel material independent of complicating factors derived from inclusion of cells or bioactive factors. However, cells or bioactive factors could be included as described hereinabove to produce desired effects. Surgical Procedure [0139] Pre-Operative - After arrival in the biological resource unit, Yucatan minipigs (~7-8 months of age, -30-35 kg) were maintained for a minimum of 7 days to ensure full acclimatization. After pre-medication with Ketamine (20 mg/kg I.M.) as anesthesia and Ambipen (40,000 U/kg I.M.) as prophylaxis antibiotic, the animal's rear legs were shaved and painted in Betadyne as simultaneous bilateral surgery of both knees was performed. A general anesthesia was maintained by inhalation with Isoflurane (1-2.5% volume) in O2 following intubation. Thiopental was used as needed (to effect 25 mg/ml I. V.). During surgery, the animal was monitored for heart rate, respiration rate, body temperature, etc. [0140] Opening - A longitudinal midline skin incision was made and carried down sharply through the pre-patellar bursa. Electrocautery was used for hemostasis. The lateral border of the patella was identified and a lateral para-patellar arthrotomy was performed. The lateral retinaculum and musculature were tagged with #1 vicryl suture. The patella was dislocated medially to expose the femoral trochlea. [0141] Cartilage repair - As seen in Fig. 6, two circular full thickness chondral defects (-4.5 mm in diameter, panel B of Fig. 6) were created in the medial trochlear facet of the femoral chondyle (panel A of Fig. 6) using an Acufex 4.5 mm mosaiaplasty chisel and a sharp, curved currette taking care as much as possible not to disrupt the osteochondral plate. The defects were filled with in situ cross-linked T-HA hydrogel (125 mg/ml in sterile saline) as follows to produce a hydrogel implant having the composition described above in order to reproduce the in vitro measured compressive properties of natural cartilage as described above. Initially, each defect was rinsed with 0.01 cc of 0.6% hydrogen peroxide, and then immediately blotted dry with sterile gauze. Subsequently, a plug containing 0.15 cc of uncross-linked hydrogel paste (panel C of Fig. 6) having the composition and prepared as described above was inserted into and used to fill each defect with the surgeon smoothing the surface of the hydrogel implant with fingertips to match the contour of the articular surface. A sterile piece of filter paper (Whatman 50) soaked in 0.6% hydrogen peroxide was pressed against the surface of the hydrogel implants for five minutes to cross-link the hydrogel in the defect. During the 5 minutes, the filter paper was rubbed back and forth across the implant surface to prevent integration with the filter paper, and to effectively polish the implant surface. After 5 minutes the filter paper was removed, excess hydrogel trimmed from the site, and then -0.01 cc (1 drop) of 0.6% hydrogen peroxide added to the surface of each implant plug (panel D of Fig. 6). The Patella was reduced anatomically over the femoral trochlea. The Patella was dislocated and reduced again to ensure secure primary stability of the hydrogel. [0142] Closing - The joint was irrigated with sterile saline. The wounds were closed in layers with vicryl sutures. Specifically, the arthrotomy was closed with interrupted #1 vicryl suture, the subcutaneous tissue was closed with interrupted 2-0 vicryl suture and the skin layers were closed with interrupted 3-0 vicryl suture. No restriction of movement was required after surgery. [0143] Post-Operative - The animal returned to full weight bearing immediately following surgery. Analgesia was provided by Buprenorphine (0.02 mg/kg I.M.) for 24 hours and a Fentanyl Patch (50 mcg/hr) for 3 post-operative days. Post-operative prophylactic antibiotic in the form of cephalexin 500 mg twice per day was given for 7 days. The animals were kept in a conventional animal run. [0144] Post-Implantation Data - At one month post-implantation, the animal was euthanized with overdose of the barbiturate, Beuthanasia D Special (1 ml/10 kg B.W., I.V.) under general anesthesia. After euthanasia, the entire knee joint was carefully dissected, macroscopically evaluated and photo documented. As seen in Fig. 7, macroscopic inspection of the knees at 1 month revealed no significant effusion and no evidence of inflammatory reaction. The lesions were partially filled with a white material (the implanted T-HA hydrogel as well as other factors or cells which may have migrated into the hydrogel post- operatively) and the surrounding articular cartilage and opposing articular surface (patella) were normal in appearance except for a slight abrasion appearing on the opposing articular surface as seen in panel B of Fig. 7. It is not evident this abrasion was the result of rubbing against the implants, particularly given its location on the Patella in a position which does not appear as though it would have abraded against the implants during normal articulation of the joint. [0145] The results indicate no apparent negative effect on joint health as a result of the hydrogen peroxide or peroxidase reaction used for in situ cross-linking of the hydrogel, and demonstrated the utility of the hydrogels disclosed herein, comprising a dityramine cross- linked hyaluronan macromolecular network, as a synthetic implantable extracellular matrix for use as a synthetic in vivo cartilage replacement or implant material. EXAMPLE 7 [0146] An experiment was conducted whereby a T-HA hydrogel as described hereinabove was implanted into canine and rabbit models in order to repair vocal cord defects as well as to augment vocal cords. Following is a description of that experiment, including the experimental methods and results obtained, after, a brief discussion of the background for this application. Background Tissue Description [0147] The vocal cords are complex, multilayered structures under very fine neuromuscular control. The overlying mucosa is composed of a non-keratinized, stratified squamous epithelium, with a multilayered, lamina propria deep to the epithelium. Underlying the lamina propria is a muscular layer consisting of the thyroarytenoid muscle which inserts into the thyroid cartilage anteriorly and the vocal process of the arytenoid cartilage posteriorly. The thyroarytenoid muscle can stiffen or relax, altering the tension on the lamina propria and thereby altering the vibratory dynamics of the epithelium, which produces the finely coordinated vibrations responsible for high quality speech production. [0148] The biomechanics of human voice production have been attributed to the action of certain biological macromolecules naturally found within the extracellular matrix (ECM) of the lamina propria. Hyaluronan (HA) is a ubiquitous molecule, which is most concentrated in specialized tissues such as the vocal cords, synovial fluid, umbilical cord, dermis, and cartilage. In these tissues, its function is manifold, influencing tissue viscosity, shock absorption, wound healing, and space filling. [0149] The unique structure of HA elucidates its multiple functions. It consists of D- glucuronic acid and N-acety]glucosamine arranged in repeating disaccharide chains, containing as many as 30,000 repeating disaccharide units with a mass of more than 10 megadaltons. As HA is a polysaccharide instead of a protein, it is non-antigenie. Under biological conditions, it is a negatively charged, randomly coiled polymer, filling a volume more than 1,000 times greater than expected based on molecular weight and composition alone. The strong negative charges attract cations and water, allowing it to assume the form of a strongly hydrated gel, and giving HA its unique viscoelastic and shock-absorbing property. [0150] Vocal cord viscoelasticity is essential to high quality voice production, as it directly affects the initiation and maintenance of phonation and the regulation of vocal cord fundamental frequency. HA in the human glottis is concentrated in the lamina propria and its importance has been quantified by comparing the biomechanical properties of cadaveric vocal cords with and without HA. Treatment of the vocal cords with hyaluronidase led to a 35% average reduction in vocal cord stiffness and a 70% mean reduction in high frequency vocal fold viscosity, thus illustrating the significance of HA in these tissues. Need for a Synthetic Material [0151] Vocal Cord Repair - Defects in the vocal cords have a dramatic effect on vocal production. Arising either de novo, or resulting from surgical intervention, heterogeneous masses within the vocal cords disrupt the finely coordinated vibrations responsible for high quality speech production. Patients with de novo lesions usually present early in the disease course, due to persistent hoarseness. When presenting with an early stage malignant process (T1-T2 in the Tumor, Node, Metastasis staging system), patients undergo rapid treatment consisting of either external beam radiation therapy or endoscopic surgery. Such patients are counseled that poor post-treatment voice quality is an expected side effect of effective tumor eradication. When presenting with a presumed benign process, patients are faced with a conundrum, for the surgical treatment often produces speech quality as poor as that caused by the lesion itself. Unfortunately, current standard laryngeal operative technique cannot provide for effective removal of either benign or malignant lesions without causing poor vocal outcomes secondary to vocal fold scarring. This is due to the mechanism of wound healing in the unique anatomy of the larynx. The superficial, vibratory surfaces of the vocal cords become tethered to the deeper layers by the post-treatment scar, preventing physiologic phonatory oscillation. [0152] HA in the human glottis is concentrated in the lamina propria, a histological layer separating the vocalis muscle from the overlying epithelium. The lamina propria allows the epithelium to vibrate over the taut vocalis muscle, like waves propagating over a pond. This "mucosal wave" is the sine qua non of effective speech production. In the presence of benign or malignant lesions of the vocal cords, the mucosal wave is disrupted. Even in the normal process of healing, scar bands and disorganized collagen "tether" the superficial mucosa to the deeper layers of the vocal cords, disrupting the normal mucosal wave and impairing vocal production. The shock absorbing nature of HA allows it to act as a tissue damper, protecting the mucosal surfaces from the oscillatory trauma experienced during phonation. HA also appears to facilitate wound repair by minimizing fibrosis and scarring, thereby protecting the vocal cord from the permanent damage resulting from trauma. [0153] The development of a technique, permitting the restoration of a fully vibratory phonatory surface on vocal cords undergoing laser or cold surgical treatment, would enable a large population of patients with both benign and malignant processes to undergo treatment of their tumors with the expectation of unprecedented post-operative speech outcomes. [0154] Vocal Cord Augmentation - A variety of disorders and diseases adversely affect glottic function, vocal quality and the ability to communicate. Approximately 7 million people in the United States suffer from dysphonia or voice impairment, and those affected by vocal cord paresis/paralysis are a significant subset of this population. It is estimated that 1-4% of all cardiac and thyroid surgeries in the United States result in vocal cord paresis or complete paralysis due to inadvertent vagus nerve or recurrent laryngeal nerve injury during surgery. [0155] Another condition affecting vocal cord function is unilateral vocal fold paralysis (UVP). In UVP the problem is malposition of an insensate vocal cord. While medialization results immediately following nerve injury due to opposing tensions of the laryngeal adductors and abductors, the paralyzed vocal cord rapidly lateralizes to a paramedian position. The arytenoid cartilage prolapses into the larynx following recurrent laryngeal nerve injury, resulting in a change in vertical height of the vocal cord, as well as decreased dynamic tension often resulting in vocal fold bowing. Atrophy with resultant shortening and bowing of the vocal cord occurs later as the thyroarytenoid muscle atrophies due to a lack of neural stimulation. As a result of atrophy and lateralization, the contralateral vocal cord cannot fully contact the paralyzed cord, leading to manifestations of UVP. [0156] Such manifestations include breathy hoarseness, a weak cough, an inability to valsalva (protect the airway), and difficulty swallowing; complications include aspiration (solids and liquids) and recurrent pneumonia. This can result in a life-threatening condition due to increased incidence of recurrent pulmonary infections. [0157] As only one functional vocal cord is required for normal voice production, successful treatment consists of "medialization" of the paralyzed vocal fold, thereby enabling it to contact the contralateral mobile vocal fold. This normalizes voice production and prevents aspiration minimizing the risk of aspiration pneumonia. Vocal fold paralysis is currently treated in two ways: open trans-cervical approaches or trans-oral endoscopic vocal cord injection, also known as injection laryngoplasty therapy (ILT). Ishiki-type I thyroplasty is the most commonly performed trans-cervical approach, where a "window" is created in the thyroid cartilage to allow placement of a silastic implant into the body of the atrophic, paralyzed vocal fold, in effect pushing it into a more medial position. This procedure has a permanent effect, although complications include implant migration, extrusion, or infection. [0158] In ILT, the paralyzed vocal cord is medialized by the endoscopic injection of an exogenous substance. A wide variety of synthetic and biologic materials are currently available as an injectant for treatment of UVP, including: gelfoam, hydroxyapatite, autologous fat or facia, acellular cadaveric dermis (Cymetra®), collagen or Teflon®/Gortex®. Unfortunately, all have proven to be less than ideal with none fulfilling all desired criteria for the ideal material for long-term vocal cord augmentation. Such limitations create the need for either re-injection or over-injection to account for the projected loss of volume. [0159] A biocompatible, injectable material such as a T-HA hydrogel as disclosed herein can be designed to mimic the Theological properties of the natural vocal cord tissue and persist indefinitely in vivo without migration. The design, chemistry, and material properties of a T-HA hydrogel as described herein can be tuned to produce an injectable bioimplant that is uniquely suited to otolaryngology treatments such as ILT through judicious selection of component concentrations and cross-linking methodologies as described hereinabove. [0160] A suitable biocompatible, high-longevity synthetic material also is desirable to treat pre-existing sulcus or scarring that can develop due to trauma or develop spontaneously with aging. Such a material also could be used advantageously in place of saline as a diagnostic and surgical aid prior to vocal cord surgery. Conventionally, saline is injected within the HA matrix of the vocal cord lamina propria between a lesion to be surgically removed and the underlying ligament. This is done to: a) determine if t: lesion involves the underlying ligament; and b) to make the surgery easier by increasing the distance between the lesion and the ligament (cold instruments), or providing a heat sink (laser). Ligament involvement complicates the surgery with penetration of the ligament to be avoided if possible. This procedure could benefit from the incorporation of hepatocyte growth factors in the hydrogel, which is used to increase HA production in the lamina propria J decrease collagen production associated with scarring. Experimental Description Design of Extracellular Matrix Material having Desired Properties [0161] An ideal synthetic matrix or biomaterial for vocal cord augmentation will have the following characteristics: 1) biocompatible, so there is no unfavorable immunologic response; 2) easily injectable to allow a surgeon to control the exact amount and location of injection through a small needle; 3) readily available with minimal preparation for optimal time efficiency arid potential application to the outpatient office setting; 4) possess the same or similar biomechanical properties to the vocal fold component being augmented to cause minimal alteration in the natural function of the augmented structure; 5) resistant to resorption or migration, so that the initial augmentation result is maintained; and 6) easily removable in the event of revision surgery. [0162] The T-HA hydrogels disclosed herein meet all six of these criteria, most important being that its biomechanical properties can be tuned through judicious selection of reactant/synthesis parameters and GAG (e.g. HA) concentration to produce the necessary macromolecular network for producing a hydrogel having desired viscoelastic and biomechanical properties. Specifically, results of both in vitro and in vivo preliminary studies have allowed favorable comparisons to be drawn against the above criteria. First, transplantation of T-HA hydrogels into various animal species including rat, rabbit, dog and pig all have demonstrated little to no host immune response. Second, uncross-linked T-HA hydrogels at the concentrations required for vocal cord augmentation at the lamina propria or muscle level easily pass through a 21 gauge needle. Third, medical-grade HA is readily available and has been used for years in FDA-approved formulations such as Healon and Restylane. Furthermore, uncross-linked T-HA hydrogel and hydrogen peroxide (cross- linking agent) solutions can be readily pre-made, in off-the-shelf formulations that require no preparation by the surgeon. Fourth, the T-HA hydrogels disclosed herein can be formulated to match the mechanical properties of the various tissues of the glottis including the lamina propria, thyroarytenoid muscle, and thyroid or arytenoid cartilages. Fifth, the unique cross- links and non-protein nature of the T-HA hydrogels have demonstrated resistance to resorption in in vivo experiments. This implies that the initial augmentation result will be maintained. Finally, formation of a solid continuous implant through the novel in situ cross- linking protocol made possible through the non-immunogenic enzyme-driven cross-linking architecture described herein should prevent migration and allow for easy location and removal of the hydrogel implant should revision surgery be required. [0163] With respect to the fourth point above, the results from confined compression testing of T-HA hydrogels (see Example 3 above) provided an initial basis for the synthesis of an appropriate synthetic T-HA material having properties matched to those of normal vocal cord tissue. Based on those data, a T-HA hydrogel composed of, inter alia a macromolecular network of dityramine cross-linked hyaluronan molecules was selected based on the following criteria in order to produce a synthetic implantable vocal cord repair or augmentation material. [0164] Choices of scaffold material (HA alone), percent tyramine substitution (-5%), protocol for tyramine substitution of HA (modified from Example 1), and the non- incorporation of cells and bioactive factors were as described in Example 6. A concentration range of between 2.5 and 10 mg/ml of T-HA hydrogel in sterile saline most closely matched the Theological and vibratory properties of the vocal cord lamina propria. The 2.5 mg/ml concentration of T-HA hydrogel was deemed most appropriate based on the extensive clinical experience of our clinician collaborators with both the vocal cord tissue and other injectable materials used for vocal cord repair and augmentation. In vitro cross-linked hydrogel was used rather than an in situ cross-linkm? protocol based on the experience of our clinician collaborators with other injectable materials for vocal cord repair and augmentation. In vitro cross-linking was as described in Example 1. Based on the results of the rabbit and canine experiments described below a preferred embodiment for vocal cord augmentation using the disclosed hydrogel material as envisioned by the inventors would use the in situ cross-linking protocol and a T-HA concentration to more closely match injection into the deeper muscle layers of the vocal cord. Surgical Procedure [0165] Pre-Operative — After arrival in the biological resource unit, mongrel dogs or New Zealand white rabbits (depending on the experiment, described below) were maintained for 7 days for full acclimatization. After pre-medication and general anesthesia per IACUC- approved protocols, each animal was intubated and maintained in stage III surgical anesthesia. The animal was positioned supine. After grasping and superiorly retracting the tongue, a Dedo laryngoscope was placed transorally providing good exposure of the larynx. The tip of the laryngoscope was positioned several centimeters proximal to the superior surface of the true vocal cords. The laryngoscope was then suspended. A rigid videostroboscopic telescope was positioned above the true vocal cords, permitting complete inspection and imaging of the larynx. [0166] Vocal Cord Repair -For vocal cord repair, lateral-based micro flaps were raised in both vocal cords of dogs, and then soft tissue defects were created equivalent to 50% of the vocal cord mass, including lamina propria and underlying muscle. One side underwent soft tissue reconstruction (filling) with the T-HA hydrogel, while the contralateral side served as an unrepaired (unfilled) control. The microflap was then redraped over the hydrogel such that the epithelium was completely continuous. This study used 2.5 mg/ml ex vivo cross-linked T-HA hydrogel in saline (5% tyramine substituted) prepared as described above to approximate the rheological properties of the lamina propria. After surgery, the dogs were weaned from the anesthesia and transferred to the recovery room. The animals received analgesia for 1 to 2 days per IACUC-approved protocols. [0167] Injection Laryngoplasty Therapy (Vocal Cord Augmentation) - After anesthesia, a 27 gauge laryngeal needle was used to inject approximately 0.25 ml of 2.5 mg/ml ex vivo cross-linked T-HA hydrogel in saline (5% tyramine substituted) at the left anterior and posterior membranous vocal cord of rabbits. The injections were made in the superficial layer of the lamina propria. [0168] Based on the results from the above experiments with dogs (vocal cord repair) and rabbits (ILT), the following preferred embodiment for vocal cord augmentation using the disclosed hydrogel material is envisioned by the inventors. For ILT, the lateralized and atrophied vocal cord are injected with 50 mg/ml of the uncross-linked T-HA hydrogel in saline (5% tyramine substituted) with peroxidase at the level of the thyroarytenoid muscle. Preferably one bolus of hydrogel would be used to obtain the desired medialization, but no more than two. A 21 gauge laryngeal needle is used to inject the hydrogel. Cross-linking would be initiated by injection of a small volume of dilute hydrogen peroxide through a 27 gauge needle into the center of the implanted bolus of hydrogel using the 21 gauge needle, which would not have been withdrawn, for purpose of orientation. Cross-linking of the hydrogel into a solid implant would be achieved within minutes and verified by feel. After surgery, the animals would be weaned from the anesthesia and transferred to the recovery room. The animals would receive analgesia for 1 to 2 days per IACUC-approved protocols. At the time of euthanasia, the vocal cords of each animal would be carefully dissected, macroscopically evaluated and photo documented. Post-Implantation Data [0169] Vocal Cord Repair - At the time of euthanasia, the vocal cords of the dogs were carefully dissected, macroscopically evaluated and photo documented. The degree of wound healing was assessed histologically by a consulting pathologist with specific attention paid to inflammatory infiltrates, HA staining (normal matrix production), hydrogel staining and collagen staining (scarring). Fig. 8 shows representative histological results of control side (unfilled) and experimental side (T-HA hydrogel filled) vocal cords stained with alcian blue for one of the dogs three months following surgery. Gross observation indicated a more normal appearance and vibratory properties for the T-HA hydrogel-treated vocal cord compared to untreated controls. The histological results indicated significant scarring in the untreated control vocal cord along the wound track as indicated by a lack of deposition of GAG (i.e., HA) and increased collagen deposition when compared to the T-HA-filled wound track of the experimentally repaired vocal cord. [0170] Only small foci of T-HA hydrogel could be found in the experimental vocal cord at 12 weeks, which show a minimal foreign body response with a layer of surrounding mast cells observed. This may indicate degradation of the T-HA hydrogel with concomitant deposition of normal HA-containing tissue matrix. However, given the low concentration (2.5 mg/ml) and thus the very fluid nature of the hydrogel used in this study, it is more likely that much of the hydrogel was lost from the wound site prior to closure of the site as the epithelial microflap knitted to the opposed underlying tissue. Hydrogel between the microflap and opposed underlying tissue is actually predicted to inhibit the knitting process contributing to hydrogel loss from the wound site. Thus, the positive wound healing effect seen is believed due to only a thin layer of hydrogel retained at the wound site rather than the volume-filling bolus of hydrogel initially implanted. These results indicate the ability of the hydrogel to prevent scarring and match the rheologic properties of the lamina propria. [0171] Injection Laryngoplasty Therapy (Vocal Cord Augmentation) - At the time of euthanasia, the vocal cords of the rabbits were carefully dissected, macroscopically evaluated and photo documented. Histological evaluation by a consulting pathologist was used to assess the inflammatory response and retention of the hydrogel. Representative histological results with Alcian Blue staining to detect the hydrogel and hematoxylin & eosin (H&E) for general morphology are shown in Fig. 9 for one of the rabbits that underwent the augmentation procedure at two weeks post-operatively. As seen in the figure, pockets of T- HA hydrogel could be found in the injected vocal cords at 2 weeks, which show a minimal foreign body response with a layer of surrounding mast cells observed. EXAMPLE 8 [0172] An experiment was conducted whereby a T-HA hydrogel as described hereinabove was implanted into a rabbit model in order to fill the vitreous cavity of the eye in order to prevent or treat vitreo-retinal diseases such as retinal detachment. Following is a description of that experiment, including the experimental methods and results obtained, after a brief discussion of the background for this application. Background Tissue Description and the Need for a Synthetic Material [0173] Vitreo-retinal diseases, such as retinal detachment, diabetic retinopathy and others, are among the most common causes of blindness. The vitreous cavity of the eye normally is filled with a gel like substance. In retinal detachment surgery, the vitreous is surgically removed (a procedure called "vitrectomy"), the retina is re-attached against the back wall of the eye, and a replacement substance is injected into the vitreous cavity. Vitreous substitutes are used for a number of different purposes in the vitreous cavity of the eye. These include (1) achieving a long term tamponade after retinal re-attachment surgery to keep the retina apposed to the wall of the eye; (2) in intra-operative procedures such as unfolding of retinal tears, the removal of subretinal fluid and the flotation and removal of dislocated intraocular lens components; (3) for developing a sustained release system that could maintain therapeutic drug levels in the posterior segment of the eye over long periods of time. [0174] A number of different compounds are used as vitreous substitutes after retinal re-attachment surgery. These compounds have physical properties that permit successful retinal re-attachment but fail in other important surgical goals. Gases injected into the eye provide short term retinal tamponade but re-absorb quickly and cause significant optical distortion while they are in the eye. Perfluorocarbon liquids are effective intra-operative tools for flattening the detached retina but cause unacceptable toxicity when left in the eye for prolonged periods of time. Silicon oil is used as a medium term retinal tamponade but also carries a risk of toxicity and causes significant optical distortion. [0175] Alternatively, substitute vitreous compounds are desirable for use as safe, long term or time-released drug delivery vehicles in the eye. Many chronic inflammatory and infectious conditions of the eye, such as sarcoidosis, idiopathic posterior uveitis and cytomegalovirus retinitis, necessitate intraocular injections of medication. Repeat intra- ocular injections pose risks such as bleeding, retinal detachment and infection. A stable, non- toxic vehicle is needed for sustained intravitreal drug delivery. [0176] Hyaluronan (HA) is an acellular substance that is an essential component of natural vitreous in humans and other mammals. Formulations of hyaluronan are already in use in some ophthalmic surgical procedures. For example, sodium hyaluronate is the most commonly used viscoelastic surgical device for anterior segment and cataract surgery. Unfortunately, sodium hyaluronate and other previously tested hyaluronan substitutes are dissolved relatively quickly in human tissues. These substances have not proven effective in vitreous surgery because of their failure to provide long-term retinal tamponade. Experimental Description Design of Extracellular Matrix Material having Desired Properties [0177] The most common need for a vitreous replacement is during retinal detachment surgery. A significant challenge is maintaining the retina flat against the wall of the eye for a prolonged period of time post-operatively. An ideal vitreous substitute should be optically clear to allow maximum visual rehabilitation during the recovery period. Finally, retinal detachments that occur inferiorly in the eye pose a particular challenge. For the retina to remain flat post-operatively, the vitreous replacement must be directly apposed to the area of the retinal tear. To tamponade inferior breaks the patient must often lie in a face down position for weeks after the surgery. None of the vitreous substitutes in use today satisfy all of the current clinical needs. [0178] There is a need for a non-toxic, optically clear, vitreous substitute that will result in improved surgical results and post-operative visual rehabilitation in patients undergoing retinal detachment surgery. [0179] Hydrogels made from a tyramine-substituted and cross-linked hyaluronan (T- HA) macromolecular network as disclosed herein present an ideal choice for a synthetic vitreous material. Specifically, the novel enzyme-driven cross-linking chemistry described above for cross-linking the tyramine-substituted hyaluronan macromolecules using a peroxidase and H2O2 allows the resulting hydrogels to be cross-linked ex vivo, and to remain stable in animal tissues. For example, studies in rats have demonstrated that this material does not degrade over several months when injected subcutaneously (see Example 9). At low concentrations the hydrogels are optically clear, easily injected through a syringe or a vitrectomy port and have a specific gravity higher than water. These physical properties make T-HA gels an ideal substrate for vitreous replacement. [0180] Based in part on confined compression testing data reported in Example 3 above, it was possible to design a T-HA hydrogel composed of, inter alia a macromolecular network of dityramine cross-linked hyaluronan molecules, having elastic and other physical properties matched to natural vitreous material in order to produce a synthetic implantable vitreous substitute. [0181] Choices of scaffold material (HA alone), percent tyramine substitution (-5%), protocol for tyramine substitution of HA (modified from Example 1), and exclusion of cells and bioactive factors were as described in Example 6. A concentration range of between 2.5 and 10 mg/ml of T-HA hydrogel in sterile saline most closely matched the rheologic, optical (clarity, refractive index) and gravimetric (density) properties of the vitreous of the eye. The 10 mg/ml concentration of T-HA hydrogel was deemed most appropriate based on the extensive clinical experience of our clinician collaborators. In vitro cross-linked hydrogel was used rather than an in situ cross-linking protocol based on the experience of our clinician collaborators with the potential sensitivity of the thin, highly-specialized layer of cells in the retina. In vitro cross-linking was as described in Example 1. An insoluble steroid was added to the cross-linked T-HA to allow visualization by the surgeons during the operative procedure because the hydrogel material itself was optically transparent and colorless. Surgical Procedure [0182] Rabbits underwent unilateral vitrectomy surgery (left eye only) using standard vitreoretinal surgical techniques with replacement of the natural vitreous of the eye with the T-HA hydrogel described above in order to evaluate the hydrogel material as a vitreous substitute. Following general anesthesia (ketamine: 50 mg/kg; xylazine: 5 mg/kg), the rabbit was prepped and draped in a sterile fashion. The left eye was dilated with mydriacyl and phenylephrine. Two drops of topical Ciloxan was instilled over the eye before and after the case. Topical proparacaine drops were instilled. Under an operating microscope, a 270° conjunctival peritomy was performed using wescott scissors. An infusion port was created 2.5 mm posterior to the limbus and the infusion cannula was secured to the sclera using 7-0 vicryl suture. A lens ring was sutured to the sclera using 7-0 vicryl suture. A 30° prism vitrectomy lens was placed on the lens ring. A second port was created and the vitrectomy instrument was inserted into the vitreous cavity. A complete core and peripheral vitrectomy was performed. At this point one of the ports was sutured closed with a 7-0 vicryl suture. The BSS bottle was lowered to patient level and -1.2 cc of a mixture of 3 mg/ml preservative free triamcinolone acetonide (steroid) and the T-HA hydrogel (10 mg/ml, 5% tyramine substitution) in BSS was injected into the vitreous cavity through an 18 gauge syringe. The steroid is the same as usually administered after vitreoretinal surgery with its milky appearance allowing visualization of the otherwise optically transparent hydrogel material. As the milky solution was injected it was directly visualized filling the vitreous cavity. The infusion was stopped when at 50% fill or when the T-HA material was seen backing up through the irrigation canula (100% fill). After filling the vitreous cavity 7-0 vicryl suture was used to close the remaining ports. 8-0 vicryl suture was used to close the conjunctiva. Topical bacitracin ointment was placed on the eye after closing. A topical antibiotic/bacitracin ointment was applied to the eye bid x 1 week, and the rabbit was placed in a recovery cage. After the rabbit had regained sternal recumbancy, it was returned to its home cage. Post-Implantation Data [0183] At 1 month post-implantation, rabbits were anesthetized with ketamine (50 mg/kg; 10 mg/kg/hr thereafter) and xylazine (5 mg/kg; 0.5 mg/kg/hr thereafter), the pupils dilated with eye drops (1% tropicamide; 2.5% phenylephrine), and the corneal surface anesthetized with an eye drop (0.5% proparacaine). After full pupil dilation, the status of the retina and eye were examined by indirect ophthalmoscopy followed by fundus photography. In addition, intraocular pressure (IOP) was measured using a Tonopen, a device that is used clinically and which makes minimal contact with the corneal surface. FinaUy, the rabbit was placed on a heating pad in darkness for 1 hour, and electroretinograms (ERus) recorded for both control and vitreous replaced eyes in response to flashes of light. ERG electrodes consist of a corneal contact lens and two platinum 0.5 inch Grass needle electrodes, placed in the cheek and trunk. While still under anesthesia, the rabbit was euthanized by using an intravenous dose of Beuthanasia D Special (1 ml/5 kg). Both control and vitreous replaced eyes were then enucleated and fixed in 10% buffered formalin for 24 hours, for histological evaluation. [0184] The results at 1 month post-implantation, indicated minimal post-operative inflammation of the surgically treated eye with normal IOP. By one week, a cataract had formed in the vitreous replaced eye relative to the un-operated eye and to BSS control operated eyes creating a limited view of the posterior segment of the experimental eye. Gross observation of the sectioned eyes showed the cataract in the anterior segment of the vitreous replaced eye (Fig. 10). The remainder of the anterior segment as well as the entire posterior segment of the eye looked similar by gross observation (Fig. 10). Hydrogel was recovered from the experimental eye at 1 month post-implantation, and was a clear gel-like substance similar to its pre-injection form (Fig. 10). The ERG for the vitreous replaced eye was normal compared to the un-operated control eye, and indicated that the retinal cells were alive and remained functional (Fig. 11). Finally electron micrographs from the four quadrants of the retina show normal morphology for the retina from the vitreous replaced eye compared to the normal un-operated eye (Fig. 12). These data indicate the T-HA hydrogel can be used as a vitreous substitute without causing infection or inflammation of the eye, and without damaging the retina, and illustrate a method by which the T-HA hydrogel can be used as a retinal tamponade for reattachment of a detached retina. Related Ophthalmologic Applications [0185] In addition to the foregoing retinal tamponade application following retinal reattachment surgery, the T-HA hydrogel described in this example also could be used for the following related applications: " as a vitreous replacement for intra-operative procedures such as unfolding of retinal tears, the removal of subretinal fluid and floatation and removal of dislocated intraocular lens components. ■ as a vitreous replacement incorporating a sustained release drug delivery system to maintain therapeutic drug levels (steroids, antibiotics, anti-viral drugs, etc.) in the posterior segment of the eye over long periods of time to treat chronic inflammatory and infectious conditions of the eye such as sarcoidosis, idiopathic posterior uveitis and cytomegalovirus retinitis. ■ for anterior segment surgery including as a substitute for plastic polymer inserts in corneal refractive surgery. These inserts, implanted surgically in the cornea, are used to change the shape of the cornea and correct mild myopia. The optical clarity and biocompatibility of the T-HA hydrogel with human tissues make it well suited to this application. " as a substitute for partial or full thickness corneal grafting procedures, e.g. for anterior segment surgery, necessitated as a result of corneal scarring from infection, keratoconus, or other causes. The optical and physical properties of the hydrogels make them compatible with use as corneal tissue substitutes. ■ as a viscoelastic device during anterior segment and cataract surgery. At low concentrations, hydrogels can maintain anterior chamber shape and pressure while allowing the surgeon to clearly visualize ocular structures. ■ for oculoplastic surgery including subcutaneous injection to smooth wrinkles in the face. ■ for oculoplastic surgery as an ocular implant in patients undergoing enucleation or exoneration surgery. The hydrogel formed in the dimensions of a human eye can be used as an implant to fill the orbit and improve cosmetic appearance of the individual after globe removal. ■ to coat MEMS devices for use in vitreo-retinal surgery. • to expand the utility of laser vision correction surgery (LASIK) to include those cases where volume needs to be added to the cornea rather than removed to correct vision. Corrective laser surgery would be used to produce the exact dimensions required for optimal visual outcome following implantation of an intentionally oversized plug of the hydrogel. ■ as a replacement for the typical gases, perfluorocarbon liquids and silicon oils normally used as tamponades in eye surgery. These applications include but are not limited to the following: giant retinal tears, proliferative vitreoretinopathy (PVR), large breaks with "fish-mouth" phenomenon, posterior breaks or macular holes, the restoration of intraocular volume after drainage of subretinal fluid, total retinal detachment with multiple breaks and large meridional folds, retinal detachment caused by ocular trauma or complicated by PVR or associated with choroidal coloboma, dislocated lenses, suprachoroidal and submacular hemorrhage, rhegmatogenous retinal detachments without PVR, severe proliferative diabetic retinopathy, chronic uveitis with profound hypotony, and infectious retinitis. EXAMPLE 9 [0186] An experiment was conducted whereby plugs of T-HA hydrogels as described hereinabove were implanted subcutaneously into immunocompetent rats in order to investigate and demonstrate their in vivo persistence and longevity as well as to measure any host immune response. As described in detail above, hydrogels comprising a cross-linked macromolecular network (such as a tyramine-substituted and cross-linked hyaluronan network) can be prepared having a range of physical and viscoelastic properties. These materials, for example, can be tuned to emulate natural soft tissue and could be used for repair or augmentation of soft tissue defects, as in plastic surgery or reconstructive surgery. In particular, as described in detail in Examples 3 and 4 above, the viscoelasticity, rigidity and other physical properties of the material can be tuned across a wide range to emulate like properties of a wide variety of native soft tissues, and the material can be cast or formed into a variety of complex anatomical shapes which would make it ideal for casting replacement or reconstructed tissue components; e.g., in the shape of an ear or of a nose for facial reconstruction. [0187] While it already was clear from the noted examples that these materials could be cast into appropriate shapes and could be given appropriate physical properties, the present experiment demonstrates the feasibility of using the T-HA hydrogels as synthetic tissue matrix or replacement materials in vivo. Following is a description of the experiment, including the experimental methods and results obtained, after a brief discussion of the background for this application. Background Tissue Description and the Need for a Synthetic Material [0188] The availability of biomaterials for soft tissue augmentation and head and neck reconstruction has remained a fundamental challenge in the field of plastic and reconstructive surgery. Significant research and investment has been undertaken for the development of a material with appropriate biological compatibility and life span. The present focus in tissue engineering has been directed at attempts toward fibroblast and chondrocyte cultures as a method of creating endogenous cartilage and collagen bearing structures useful for implantation. The archetypal standard of this avenue of research has been the nude mouse with a neo-cartilage ear on its back. This is based on the concept of chondrocyte culture on a poly-lactic or poly-glycolic acid framework. The presumption is that the chondrocytes can produce the extracellular matrix (ECM) for the production of cartilage and create a new functional biological filling agent with complete compatibility. The outcomes of this research have not been promising in regards to their clinical application. When placed in immunocompetent animals the structural integrity of the neo-cartilage has been shown to fail as the framework is absorbed. Fundamentally, while chondrocytes can successfully be cultured and propagated they apparently cannot be made to produce cartilage on a framework prior to its hydrolysis by the host defense mechanisms. [0189] Conventionally, clinicians have been limited by the use of xenogenic materials such as bovine collagen and unmodified hyaluronan (HA) as well as synthetic materials such as silicone, silastic and hydroxyapatite. The synthetic materials are prone to foreign body reactions and infection while the biological substrates are prone to breakdown over time. In addition, synthetic PTFE (gortex) polymers and silastic offer less tissue reactivity but do not offer tissue integration and also can represent long term risks of foreign body infections and extrusion. [0190] Instead of a tissue engineering model where chondrocytes are required to produce a cartilage ECM, the hydrogels disclosed herein are or can be based on the same materials that provide cartilage its functionality and feel (HA). In the present invention, hyaluronan is used directly as the substrate for the creation of a stable tissue engineered material to replace natural HA-rich soft tissues. In essence, the HA-based hydrogels used in herein incorporate the same material that gives cartilage its form and structural characteristics, but it is modified (tyramine-substituted and cross-linked) to make the material resistant to biological degradation. Thus, an ideal synthetic extracellular matrix material suitable for in vivo implantation and longevity is achieved. Experimental Description Design of Extracellular Matrix Material having Desired Properties [0191] To provide a synthetic soft-tissue and cartilage substitute for use in head and neck reconstruction based on the disclosed HA materials the following points were considered: 1) optimization of enzyme-selective cross-linked hydrogels using hyaluronan as the scaffold material; and 2) application of T-HA hydrogels as cartilage substitutes and soft- tissue fillers. These include characterization of the effect of the cross-linked hydrogels in vivo. [0192] Again, based in part on confined compression testing data reported in Example 3 above, it was possible to design T-HA hydrogels composed of, inter alia a macromolecular network of dityramine cross-linked hyaluronan molecules, having elastic and other physical properties matched to natural soft tissues which were suitable for in vivo rat implantation to determine their immunogenic and longevity characteristics. [0193] Choices of scaffold material (HA alone), percent tyramine substitution (-5%), protocol for tyramine substitution of HA (modified from Example 1), and exclusion of cells and bioactive factors were as described in Example 6. A concentration range of between 6.25 and 100 mg/ml of T-HA hydrogel encompassed the wide spectrum of physical properties required of a material for facial reconstruction. Therefore the same five concentration used in Example 3 were deemed appropriate for testing in a subcutaneous rat model based on the extensive clinical experience of our clinician collaborators. In vitro cross-linked hydrogels were used so as to produce hydrogels of defined shape for analysis of shape retention, a property deemed important by our clinician collaborators. In vitro cross-linking was as described in Example 1. Surgical Procedure [0194] T-HA hydro gel plugs of defined shape, mass and volume (7 mm in diameter and 3 mm in thickness) and defined mechanical properties based on HA concentration were surgically implanted subcutaneously in the backs of immunocompetent rats to allow evaluation of their in vivo persistence and host immune response based on previously published protocols for the evaluation of collagen and other HA-based hydrogels. After induction of anesthesia with intraperitoneal injection of ketamine (100 mg/kg) and xylazine (5 mg/kg), the rat received a single intramuscular injection of 60,000 units of procaine penicillin for infection prophylaxis. A 1cm stab incision was made with a #11 surgical blade in the lower lumbar region of the rat. A 14 g needle was used as a trocar to dissect in the subcutaneous plane to create a pocket. Three preformed hydrogel plugs (~7.1 mm diameter x 3 mm thick) of one of the HA concentrations to be tested were inserted into the surgical pocket. A single absorbable stitch (3-0 Chromic) was placed to re-approximate the skin edge. At 1- week, 1 month, 3 months, and 6 months post-implantation, rats were sacrificed by CO2 asphyxiation and the T-HA hydrogel plugs with surrounding tissue excised and stored in formalin at 4°C until time for histological evaluation. Post-Implantation Data [0195] Hydrogel compositions tested included plugs made from concentrations of 6.25, 12.5, 25, 50, and 100 mg/ml of HA; generating hydrogel plugs with a wide spectrum of physical properties ranging from that of gel to a paste to a rubber-like material (see Example 3). Implanted T-HA hydrogel plugs were collected at 1 week, 1 month, 3 months, and 6 months post-implantation. Excised plugs were evaluated for their in vivo persistence and host immune response. [0196] Fig. 13 shows representative results of histological staining with H&E, alcian blue, MC/giemsa, Movats, Reticular, and Trichrome stains for the 100 mg/ml TB-HA hydrogel plug from the 1-month time point. Clearly defined in Fig. 13 are the surface hair follicles, the superficial muscle layer, the hydrogel plug, and the thin fibrous capsule surrounding the hydrogel plug as a result of a minimal foreign body response. An artifact exists as a result of hydrogel shrinkage from the paraffin embedding process, which can be avoided through the use of frozen sections. The results show very little immune response with only a thin layer of mast cells surrounding the plug, and no evidence for host cell infiltration into the plug. When measured, the volume of the void left by the plug during histological processing is 3 mm (the original plug thickness) indicating little to no biodegradation or deformation of the hydrogel matrix. Staining indicated that the plugs had little protein, such as collagen or elastin, deposited within them and remained primarily composed of HA hydrogel. These results indicate that the hydrogel plugs over a broad range of five concentrations persisted through 6 months with little evidence of degradation, host immune response, and cellular infiltration providing a wide range of injectable materials for use in soft tissue reconstruction. EXAMPLE 10 [0197] It will be apparent from the foregoing discussion and the Examples that hydrogels described herein composed of a cross-linked (in situ or ex vivo) macromolecular network of hyaluronan molecules cross-linked via a suitable dihydroxyphenyl cross-linking chemistry as herein described are suitable as a synthetic, implantable extracellular matrix tissue material for a variety of tissue engineering and repair applications. A particular such application for which the disclosed hydrogel materials will have particular utility is in the repair or augmentation of the mitral valve in a heart. [0198] The mitral valve is one of the most complex connective tissue structures in the body. It consists of two leaflets and numerous chordae tendineae. These chordae have a highly aligned collagenous core and a thin outer sheath of elastic fibers and endothelial cells. Both leaflets are laminated tissues containing a heavily collagenous layer on the ventricular side, a predominantly elastic layer on the atrial side, and an inner spongiosa layer containing abundant proteoglycans (PGs) and hyaluronan (HA). The relative thicknesses of these layers vary between the two leaflets and also within each leaflet from its attachment edge to its free edge. [0199] The variability of the different leaflet layers, and hence the structural constituents within the mitral valve, are determined by the specific functional roles of the leaflets and chordae. The closed valve maintains a balance of tensile and compressive loads, in which the chordae and the flat central region of the anterior leaflet are in tension, whereas the free edge of the anterior leaflet and most of the posterior leaflet are in appositional compression. Accordingly, the most collagenous components of the mitral apparatus are the chordae and the portion of the anterior leaflet between the annulus and the upper appositional border. In the posterior leaflet and in the free edge of the anterior leaflet, the collagenenous layer is relatively thinner, whereas the PG rich spongiosa is substantially thicker. The wide diversity of glycosaminoglycans (GAGs) and their parent PGs exert considerable yet variable control over the physical properties of the extracellular matrix. [0200] Functional mitral regurgitation (MR) refers to the regurgitation that occurs with a structurally normal valve as a consequence of left ventricular (LV) dysfunction, and as a result, almost half the patients with LV dysfunction have at least moderate MR. Functional MR plays a pivotal role in the pathophysiology of congestive heart failure (CHF), a major cause of cardiac morbidity and mortality. Several studies have shown that the presence of functional MR in patients with CHF is associated with poor outcomes. Although this observation could suggest that MR is merely an indicator of CHF severity, it is also increasingly apparent that the development of the MR hastens the progression of CHF. The precise mechanism of functional MR remains controversial and can relate to mitral annular dilatation in the septal-lateral (S-L) axis or tethering of the leaflets secondary to progressive ventricular remodeling. MR leads to greater volume overload of the LV with progressive annular dilatation and increased MR, creating a "vicious cycle" which exacerbates the problem. MR is commonly considered to be one of the initiators of CHF, as well as an ongoing impetus of the progression of the disease. [0201] Surgical annuloplasty is a widely used method for mitral valve repair and can provide long-term benefit. However, the surgical procedure requires access to and manipulation of the valve annulus via atriotomy. In addition, the procedure requires the patient to be placed on cardiopulmonary bypass (CPB). The prolonged CPB time has been suggested as a cause of not only postoperative LV dysfunction but also main organ dysfunction. The use of heparin during CPB results in an increased risk of bleeding complications. The increased morbidity and mortality profile leads many care providers directly to non-treatment options of MR in the earlier stage heart failure patients. [0202] Recently, several minimally invasive methods of mitral valve repair have been developed. For example, several investigators have reported the preliminary methodology of off-pump mitral valve repair procedures through a thoracic incision. Others have reported new devices that can be inserted percutaneously into the coronary sinus and great cardiac vein to reduce the S-L dimension of the mitral annulus. There are the possibilities of adverse effects, however, such as obstruction or disturbance of the coronary circulation, by chronic placement of the device in the coronary sinus, [0203] Surgical therapy for functional MR, including mitral valve repair with an annuloplasty ring and replacement with an artificial valve has been limited in patients with severe CHF by relatively high operative mortality rates due to the effects of CPB. Therefore, there is a need for a minimally invasive procedure that will not compromise coronary circulation and will allow for reduction of the S-L dimension of the mitral annulus to reduce functional MR as well as other forms of MR. Myxomatous changes in mitral valve tissues can lead to leaflet prolapse and mitral regurgitation. [0204] The T-HA hydrogel materials disclosed herein could be adapted for this purpose; i.e. mitral annular remodeling resulting from a nonabsorbable substance injection (namely a T-HA hydrogel material designed to have the necessary viscoelastic and other physical properties) into the posterior mitral annulus using an epicardial approach. This procedure would enable the S-L dimension to be efficiently reduced, thus reducing MR, without employing CPB or implanting a device into the coronary sinus. This mitral annular remodeling procedure could be modified to allow for percutaneously injection of the substance through the coronary sinus. [0205] This application for a T-HA hydrogel as disclosed herein would enable a nonabsorbable substance to be percutaneously injected through the coronary sinus for severe CHF patients with functional MR who are unable to receive conventional mitral valve surgery. This minimally invasive approach would obviate the need for CPB and sternotomy, as well as diminish the risk of major side effects from conventional surgical therapy such as postoperative LV dysfunction and resulting poor organ perfusion. In addition, such a procedure would provide patients with mild to moderate CHF with an option for early restoration of mitral valve competence, arresting the initiation and progression of devastating heart failure. [0206] In particular, a hydrogel composed of a cross-linked macromolecular network as described herein, particularly of HA, could be designed based on the principles as elucidated in Example 3 above, in order to produce a hydrogel material having all of the following characteristics which would be considered desirable for this application: ■ Injectable and nonabsorbable; ■ Low-grade inflammatory reaction; ■ Low evidence of foreign body migration; ■ Ease of collagen encapsulation which contributes to the prevention of migration; ■ Not especially malleable nor especially rigid. [0207] A protocol has been established for the injection of a T-HA hydrogel material to augment the mitral valve of a beating heart. That protocol is described as follows. [0208] Injection Procedure - Two dimensional epicardial echocardiography (2D EE) and transesophageal echocardiography (2D TEE) are performed to evaluate LV end-diastolic and end-systolic volumes (EDV and ESV), stroke volume (SV), ejection fraction (EF), the S- L dimension of mitral annulus, and the degree of MR. Hemodynamic data such as LVP, LAP, the central venous pressure (CVP), the pulmonary arterial pressure (PAP), the pulmonary capillary wedge pressure (PCWP), CO, LAD flow, and LCX flow should be collected. LV end-diastolic and end-systolic pressure-volume relations can be obtained by transient TVC occlusion using an occlusion catheter to assess LV contractility and compliance (baseline before injection). [0209] A commercially available cardiac stabilizer used for ofT-pump coronary bypass grafting can be used to stabilize the target region. Under 2D TEE guidance, the uncross-linked T-HA hydrogel composition (50 mg/ml in sterile saline) is injected into the posterior mitral annulus from the outside of the heart while the heart is beating. During the injection, 2D can be used to assess the position of the tip of the needle for the injection, the range occupied by the substance in the posterior annulus. Once an appropriate fill and repositioning of the mitral valve has been accomplished, cross-linking is initiated by injection of 0.2 cc of 0.6% hydrogen peroxide. After completion of cross-linking of the hydrogel implant, data including hemodynamics, coronary flow, LV pressure-volume loops (LV P-V loops), 2D EE, and 2D TEE should be collected (data after injection). [0210] The foregoing injection methodology was developed and evaluated using cadaveric dog and pig hearts as models. Fig. 14 shows a cadaveric dog heart in which a T- HA hydrogel material was injected and cross-linked in situ via the foregoing injection methodology. To produce the T-HA hydrogel material for this experiment, the scaffold material (HA alone), percent tyramine substitution (~5%), protocol for tyramine substitution of HA (modified from Example I) and non-inclusion of cells and bioactive factors were as described in Example 6. Concentrations of 25, 50 and 100 mg/ml of T-HA hydrogel in saline most closely matched those of cardiac (heart) tissue required for mitral valve closure. The 50 mg/ml T-HA hydrogel was deemed most appropriate for a mitral annular remodeling procedure by percutaneously injection based on the extensive clinical experience of our clinician collaborators. Peroxidase would be added at 10 U/ml prior to application in an in situ cross-linking protocol as the one described below. The in situ cros linking protocol is preferred as it would allow the uncross-linked T-HA to pass through an appropriately sized needle while the cross-linked hydrogel may not. In addition, the in situ cross-linking protocol allows the surgeon first to properly position (close) the mitral valves by injection of the uncross-linked hydrogel and then cross-link the hydrogel into a solid implant only after visual confirmation that the mitral valve had been properly repositioned. [0211] In Fig. 14, the implanted hydrogel was bisected post-'^Dlantation to demonstrate its solid, viscoelastic character following in situ cross- ng, as well as to evaluate its placement in the heart. In particular, these models were used to 1) evaluate the appropriate concentration of hydrogel required to both mimic the consistency of cardiac muscle after cross-linking yet pass through the injection port prior to cross-linking; 2) demonstrate reproducibility for the in vivo cross-linking protocols for complete cross-linking in vivo at the required volumes (~2 ml); and development suitable injection techniques. All of these goals were met, establishing confidence that the injection procedure as well as a suitable T-HA hydrogel can precisely accommodate anatomical constraint on the mitral annulus. [0212] Although the above-described embodiments constitute the preferred embodiments, it will be understood that various changes or modifications can be made thereto without departing from the spirit and the scope of the present invention as set forth in the appended claims. WE CLAIM: 1. A synthetic, implantable tissue matrix material comprising a macromolecular network comprising wherein R1 and R2 each comprises a structure selected from the group consisting of a) polycarboxylate molecules that have been substituted at CO2H sites thereon with a hydroxyphenyl compound at a substitution rate less than 10 percent based on a total number of CO2H sites on the polycarboxylate molecules, b) polyamines, that have been substituted at primary amine sites thereon with a hydroxyphenyl compound, and c) copolymers thereof, and wherein R1 and R2 can be the same or different structures. 2. The synthetic, implantable tissue matrix material as claimed in claim 1, wherein R1 is a polycarboxylate. 3. The synthetic, implantable tissue matrix material as claimed in claim 1, wherein R1 is a polyamine. 4. The synthetic, implantable tissue matrix material as claimed in claim 1, wherein R1 comprises a structure selected from the group consisting of glycosaminoglycans. 5. The synthetic, implantable tissue matrix material as claimed in claim 1, wherein R1 comprises hyaluronan. 6. The synthetic, implantable tissue matrix material as claimed in claim 1, wherein R1 comprises chondroitin sulfate. 7. The synthetic, implantable tissue matrix material as claimed in claim 6, said chondroitin sulfate being in the form of aggrecan. 8. The synthetic, implantable cartilage material comprising the synthetic, implantable tissue matrix material as claimed in claim 1. 9. The synthetic, implantable vocal cord material comprising the synthetic, implantable tissue matrix material as claimed in claim 1. 10. The synthetic, implantable vitreous material comprising the synthetic, implantable tissue matrix material as claimed in claim 1. 11. The synthetic, implantable soft tissue material comprising the synthetic, implantable tissue matrix material as claimed in claim 1. 12. The synthetic, implantable mitral valve material comprising the synthetic, implantable tissue matrix material as claimed in claim 1. 13. The synthetic, implantable tissue matrix material as claimed in claim 1, said network comprising said polycarboxylate molecules, wherein at least one dihydroxyphenyl linkage is formed between two hydroxyphenyl groups attached respectively to adjacent polycarboxylate molecules. 14. The synthetic, implantable tissue matrix material as claimed in claim 1, said network comprising said polyamine molecules, wherein at least one dihydroxyphenyl linkage is formed between two hydroxyphenyl groups attached respectively to adjacent polyamine molecules. 15. The synthetic, implantable tissue matrix material as claimed in claim 13, said polycarboxylate molecules having a hydroxyphenyl compound substitution rate less than 9 percent based on the total number of CO2H sites present on said polycarboxylate molecules. 16. The synthetic, implantable tissue matrix material comprising a macromolecular network comprising a plurality of tyramine-substituted hyaluronan molecules, at least two adjacent hyaluronan molecules being linked by a dityramine linkage. 17. The synthetic, implantable cartilage material comprising the synthetic, implantable tissue matrix material as claimed in claim 16. 18. The synthetic, implantable vocal cord material comprising the synthetic, implantable tissue matrix material as claimed in claim 16. 19. The synthetic, implantable vitreous material comprising the synthetic, implantable tissue matrix material as claimed in claim 16. 20. The synthetic, implantable soft tissue material comprising the synthetic, implantable tissue matrix material as claimed in claim 16. 21. The synthetic, implantable mitral valve material comprising the synthetic, implantable tissue matrix material as claimed in claim 16. 22. The synthetic, implantable tissue matrix material as claimed in claim 16, having a tyramine substitution rate on said hyaluronan molecules of 1.7%-10% based on the total number of CO2H sites present on said hyaluronan molecules. 23. The synthetic, implantable tissue matrix material as claimed in claim 16, having a tyramine substitution rate on said hyaluronan molecules of 1.7%- 5% based on the total number of CO2H sites present on said hyaluronan molecules. 24. The synthetic, implantable tissue matrix material as claimed in claim 1, wherein R1 and R2 each comprise hydroxyphenyl-substituted hyaluronan. 25. The synthetic, implantable tissue matrix material as claimed in claim 24, wherein hyaluronan molecules for each of R1 and R2 have aggrecan attached thereto. 26. The synthetic, implantable tissue matrix material as claimed in claim 24, said hydroxyphenyl-substituted hyaluronan having a hydroxyphenyl substitution rate less than 9 percent. 27. The synthetic, implantable tissue matrix material as claimed in claim 24, said hydroxyphenyl-substituted hyaluronan having a hydroxyphenyl substitution rate less than 5 percent. 28. The synthetic, implantable tissue matrix material as claimed in claim 24, said hydroxyphenyl-substituted hyaluronan being tyramine-substituted hyaluronan. 29. The synthetic, implantable tissue matrix material as claimed in claim 28, said tyramine-substituted hyaluronan having a tyramine substitution rate less than 9 percent. 30. The synthetic, implantable tissue matrix material as claimed in claim 28, said tyramine-substitution hyaluronan having a tyramine substitution rate less than 5 percent. 31. The synthetic, implantable tissue matrix material as claimed in claim 1, wherein R1 and R2 each comprise hydroxyphenyl-substituted chondroitin sulfate. 32. The synthetic, implantable tissue matrix material as claimed in claim 31, wherein chondroitin sulfate molecules for each of R1 and R2 are in the form of aggrecan. 33. The synthetic, implantable tissue matrix material as claimed in claim 31, said hydroxyphenyl-substituted chondroitin sulfate being tyramine-substituted chondroitin sulfate. 34. The synthetic, implantable tissue matrix material as claimed in claim 33, said tyramine-substituted chondroitin sulfate having a tyramine substitution rate less than 9 percent. ABSTRACT Title: Hydroxypenyl cross-linked macromolecular network and applications thereof. A synthetic, implantable tissue matrix material comprising a macromolecular network comprising wherein R1 and R2 each comprises a structure selected from the group consisting of a) polycarboxylate molecules that have been substituted at CO2H sites thereon with a hydroxyphenyl compound at a substitution rate less than 10 percent based on a total number of CO2H sites on the polycarboxylate molecules, b) polyamines, that have been substituted at primary amine sites thereon with a hydroxyphenyl compound, and c) copolymers thereof, and wherein R1 and R2 can be the same or different structures.

Full Text

HYDROXYPHENYL CROSS-LINKED MACROMOLECULAR NETWORK AND
APPLICATIONS THEREOF
[0001] This application claims priority from U.S. provisional patent application Serial
No. 60/586,585 filed July 9, 2004, the contents of both of which are incorporated herein by
reference in their entirety.
BACKGROUND OF THE INVENTION
[0002] Articular cartilage performs an essential function in healthy joints. It is
responsible for absorbing and dissipating impact and frictional loads in order to divert these
loads away from bones, to protect the bones from damage. Cartilage performs this function
by transferring the loading force to a fluid phase within a three-dimensional network of
aggrecan molecules, themselves constrained (described in the next paragraph) within the joint
space. Aggrecan molecules have up to 100 chondroitin sulfate (CS) chains attached to a core
protein, with each chondroitin sulfate chain possessing multiple negatively charged sulfate
groups along their length. The effect of all these sulfate groups is to cause each of the
chondroitin sulfate chains in a single aggrecan molecule to repel one another, (resulting in the
aggrecan molecule having the maximum possible volume at rest), and also to cause adjacent
aggrecan molecules in a cartilage aggregate to repel one another.
[0003] In healthy cartilage, aggrecan molecules are attached to long hyaluronan
chains, which are in turn constrained in large cartilage aggregates within the joint space by an
extracellular collagen fibril matrix. Thus, even though adjacent chondroitin sulfate chains in
each aggrecan molecule (and adjacent aggrecan molecules attached to the same or a different
hyaluronan chain) repel one another, they are nonetheless constrained within the collagen
matrix. See Fig. 1 depicting normal, healthy cartilage. Because the chondroitin sulfate
chains are so repulsive, the hyaluronan-aggrecan network (or macromolecular network)
expands as much as possible within the constraints of the collagen matrix to achieve the
lowest possible energy state at rest; i.e. to allow the maximum possible spacing between
adjacent negatively charged sulfate groups. As a result, network molecules are highly
resistant to being shifted or displaced in order to avoid approaching an adjacent network
molecule. These large cartilage aggregates are trapped at one fifth their free solution volume
within a meshwork of collagen fibers, which resist any further swelling. Cartilage aggregates
with their high negative charge density bind large solvent domains, and contribute to
cartilage's ability to absorb loads and resist deformation. Upon compression, the distance

between the fixed-negative charge groups on the proteoglycans decreases, which increases
the charge-to-charge repulsive forces as well as the concentration of free-floating positive
counterions (such as Ca2+ and Na+). Both effects contribute to the viscoelastic nature of
cartilage and its ability to resist deformation and absorb compressive loads, further described
below.
[0004] Within the macromolecular network are water molecules which provide a
substantially continuous fluid phase. The macromolecular network diverts impact and
frictional loads away from bones by transferring them to the continuous fluid (water) phase as
follows. As a joint undergoes a load, the force is absorbed first by the macromolecular
network, where it acts on and tends to deform or compress the network. The force sets up
pressure gradients in the fluid phase in order to induce fluid flow to accommodate network
deformation or compression resulting from the load. But the fluid cannot negotiate the tight
macromolecular network, packed with the repulsive chondroitin sulfate chains, sufficiently to
ac: mmodate a bulk flow of water without shifting or displacing the network molecules.
Hence, individual water molecules may diffuse within the network, but the bulk fluid phase is
substantially constrained from flowing through the network except at a much slowed rate due
to the resistance to displacement of network molecules. Because the water molecules cannot
flow readily despite the pressure gradients, the energy from the impact or frictional load is
transferred to and absorbed by the fluid phase where it contributes to compressing the liquid
water until the water can be sufficiently displaced to accommodate the network conformation
and the pressure gradients have subsided. The overall result is that cartilage absorbs the
potentially harmful load, thereby diverting it from bone.
[0005] Through this elegant mechanism, normal cartilage is capable of absorbing
significant loads by transferring the bulk of the loading force to a fluid phase constrained
within a macromolecular network. This arrangement has yet to be adequately duplicated via
artificial or synthetic means in the prior art. Consequently, there is no adequate remedy for
cartilage degenerative disorders, such as arthritic disorders, where the aggrecan molecules
become separated from their hyaluronan chains and are digested or otherwise carried out
from the cartilage aggregates.
[0006] Osteoarthritis and rheumatoid arthritis affect an estimated 20.7 and 2.1 million
Americans, respectively. Osteoarthritis alone is responsible for roughly 7 million physician
visits a year. For severe disabling arthritis, current treatment involves total joint replacement

with on average 168,000 total hip replacements and 267,000 total knee replacements
performed per year in the U.S. alone. Defects in articular cartilage present a complicated
treatment problem because of the limited capacity of chondrocytes to repair cartilage.
Treatment strategies to date have focused on the use of autologous chondrocytes expanded in
culture or the recruitment of mesenchymal stem cells in vivo by chemotactic or mitogenic
agents. The intent of these strategies is to increase and/or activate the chondrocyte
population so as to resynthesize a normal, healthy articular cartilage surface. One major
difficulty associated with these strategies is the inability to maintain these agents at the site of
the defect. Hyaluronan has been proposed as a candidate for the development of
biomaterials for local delivery of chondrocytes or bioactive agents because of its unique
properties, including excellent biocompatibility, degradability, and Theological and
physiochemical properties. However, it has been unknown whether chondrocytes suspended
in a tissue engineered hyaluronan matrix would be able to synthesize a new cartilage matrix
with mechanical properties comparable to normal, healthy articular cartilage. This is because
conventional biomaterials made from hyaluronan are formed through chemistries that are
incompatible with maintaining cell viability. Chondrocytes must be introduced to the
matrices after matrix formation with variable and normally poor results.
[0007] Accordingly, there is a need in the art for an artificial or synthetic matrix that
can effectively divert a loading force from bones in an effective manner. Preferably, such a
matrix can be provided in situ or in vivo to repair or replace articular cartilage during an
orthopedic surgical procedure. Most preferably, the artificial or synthetic matrix can be
provided to an in situ or in vivo target site as a liquid or a plurality of liquids, and can set up
in place to provide a substantially seamless integration with existing cartilaginous and/or
bony tissue in a patient.
[0008] It also is desirable to provide an artificial or synthetic matrix that can be used
or adapted to synthesize a variety of replacement tissues.

SUMMARY OF THE INVENTION
[0009] A synthetic, implantable tissue matrix material is provided including a
macromolecular network that includes the following structure

wherein R1 and R2 each comprises a structure selected from the group consisting of
polycarboxylates, polyamines, polyhydroxyphenyl molecules, and copolymers thereof, and
wherein R1 and R2 can be the same or different structures.
[0010] A variety of synthetic, implantable tissue materials also are provided which
include or are composed of the tissue matrix material mentioned in the preceding paragraph,
including a synthetic, implantable cartilage material; a synthetic, implantable vocal cord
material; a synthetic, implantable vitreous material; a synthetic, implantable soft tissue
material; and a synthetic, implantable mitral valve material.
BRIEF DESCRIPTION OF THE DRAWINGS
[0011] Fig. 1 is a schematic diagram of normal, healthy human cartilage.
[0012] Fig. 2 is a schematic diagram of a dihydroxyphenyl cross-linked
macromolecular network according to the invention.
[0013] Fig. 3 is a structural formula of a hyaluronan molecule.
[0014] Figs. 4a-4c are graphs showing comparative results for mechanical testing in a
confined compression test (equilibrium stress versus applied strain) of T-HA (Fig. 4a), T-
Aggrecan (Fig. 4b) and 50% T-HA/50% T-Aggrecan composite (Fig. 4c) hydrogels
according to the invention versus published results for articular cartilage plugs (Example 3).
The relationship between glycosaminoglycan (GAG) concentration and material compressive
strength is shown in Fig. 4d.
[0015] Fig. 5 is a graph showing comparative data of glucose utilization for
chondrocytes embedded in T-HA hydrogels (1.7% and 4.7% T-HA) compared to cultured on
tissue culture plastic (control).

[0016] Fig. 6 is a series of four photographs illustrating a surgical procedure to
implant a T-HA hydrogel into articular cartilage defects according to an aspect of the
invention described in Example 6.
[0017] Fig. 7 is a series of two photographs showing the T-HA hydrogel implants one
month after implantation into the medial trochlar facet of a Yucatan pig as described in
Example 6, as well as the opposing (articulating) patella surface.
[0018] Fig. 8 is a series of photographs illustrating the histological results of control
side (unfilled) and experimental side (TB-HA hydrogel filled) canine vocal cords, 3 months
post-operatively, following a vocal cord repair procedure using a T-HA hydrogel as a
synthetic vocal cord material as described in Example 7.
[0019] Fig. 9 is a series of photographs illustrating the histological results of
surgically augmented vocal cords in a rabbit model using a T-HA hydrogel as a synthetic
vocal cord material, also as described in Example 7.
[0020] Fig. 10 is a series of photographs of control (unoperated) and experimental
(surgically replaced) eyes one month post-operative, following a vitreous replacement
procedure using T-HA hydrogel as a synthetic vitreous material as described in Example 8.
[0021] Fig. 11 shows comparative electroretinogram (ERG) results recorded for both
control and vitreous replaced eyes in response to flashes of light in a rabbit model as
described in Example 8.
[0022] Fig. 12 is a series of electron micrographs of the retina from four quadrants of
control (unoperated) and experimental (surgically replaced) eyes one month post-operative,
following a vitreous replacement procedure using T-HA hydrogel as a synthetic vitreous
material as described in Example 8.
[0023] Fig. 13 is a series of photographs showing representative results of
histological results for a 100 mg/ml T-HA hydrogel plug implanted subcutaneously into an
immunocompetent rat at one month post-operatively as described in Example 9.
[0024] Fig. 14 is a photograph of a cadaveric canine heart used to specify T-HA
hydrogel materials for mitral valve repair as described in Example 10.

DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS OF THE INVENTION
[0025] As used herein, the term polycarboxylate means a molecule, structure or
species having a chain length of at least two functional groups or units, wherein at least two
such groups or units of the chain are or comprise carboxylic acid groups that are sterically
accessible to a nucleophilic substitution reaction as described herein. Also as used herein, the
term polyamine means a molecule, structure or species having a chain length of at least two
functional groups or units, wherein at least two such groups or units of the chain are or
comprise primary amine groups that are available for a nucleophilic substitution reaction.
Also as used herein, a polyhydroxyphenyl molecule means a molecule having a chain length
of at least two functional groups or units, wherein at least two such groups or units of the
chain are or comprise hydroxyphenyl groups that can be linked to another hydroxyphenyl
group via a C-C bond. Also as used herein, a hydrogel is a material that is prepared
comprising a macromolecular network that is used or useful in tissue replacement or
engineering applications, e.g. as artificial cartilage, as a material to coat surgical instruments
to prevent tissue irritation, or to provide a semi-permeable membrane such as for use in an
artificial kidney, etc.
[0026] The invention includes a novel structure of a macromolecular network that has
been formed by linking hydroxyphenyl groups attached to adjacent long chain
macromolecules, resulting in effectively cross-linking the macromolecules to provide a large
network. The basic cross-linking structure of the network is shown below

where R1 and R2 are each long chain macromolecules. R1 and R2 can be the same molecule
or different molecules, but it will be understood that to provide a suitable network, R1 and R2
will be different molecules for at least a portion of the dihydroxyphenyl linkages in a network
according to the invention. It is not necessary, though it is preferred, that R1 and R2 are the
same species of molecule.
[0027] By providing a plurality of these dihydroxyphenyl linkages between adjacent
macromolecules, a network of dihydroxyphenyl cross-linked macromolecules is provided as

shown schematically in Fig. 2. In the figure, the macromolecules are represented
schematically by cylindrical strands, each preferably having at least two hydroxyphenyl
groups attached along its length. It is noted that not every hydroxyphenyl group must be
linked to another hydroxyphenyl group.
[0028] Briefly, the disclosed invention involves covalent coupling of hydroxyphenyl
containing compounds, including but not limited to tyramine, through their primary amine (or
carboxyl) groups to carboxyl (or primary amine) groups on various polymeric scaffold
materials, including but not limited to hyaluronan or chondroitin sulfate (e.g. in the form of
aggrecan), via a carbodiimide-mediated reaction. After isolation and purification of the
hydroxyphenyl-substituted polymeric scaffolds, the hydroxyphenyl residues are selectively
cross-linked by horseradish peroxidase (HRP) in the presence of very dilute hydrogen
peroxide to form hydrogels. As will become apparent, the hydrogels made as described
herein are or can be used as a fully implantable, non-immunogenic synthetic tissue matrix
material that can be implanted into the body for a variety of purposes as will be described.
As used herein, 'implantable' refers both to surgical implantation of a hydrogel as through a
surgical incision, and to provision of the hydrogel within the body via injection, e.g. using a
syringe. Whether surgically implanted or injected, the implantable hydrogels can be provided
within the body already cross-linked (ex vivo cross-linking) or otherwise it can be cross-
linked in situ at the site of implantation within the body as will be further described.
[0029] The first step in providing the macromolecular network is to prepare or
provide the long-chain macromolecules having periodic hydroxyphenyl groups attached. In
one embodiment, the macromolecules are polyhydroxyphenyl molecules which already have
multiple or periodic hydroxyphenyl groups, such as polyphenols. Suitable polyphenols
include polyamino acids (e.g. polytyrosine), epigallocatechin (EGC), and epigallocatechin
gallate (EGCG) isolated from green tea, less preferably other polyphenols.
[0030] In a further embodiment, the hydroxyphenyl groups can be added to the
macromolecules periodically or randomly along their length via a chemical reaction. A
preferred method of adding hydroxyphenyl groups to the macromolecules is to utilize a
carbodiimide-mediated substitution reaction pathway to provide an amide bond between a
primary amine having a hydroxyphenyl group and a carboxylic acid group attached to the
macromolecules. In this method, the long-chain macromolecule preferably is a
polycarboxylate molecule, having periodic carboxylic acid groups along its length. The

hydroxyphenyl groups are provided as part of smaller molecules having primary amine
groups that can be attached to the carboxyl carbon atoms of a carboxylic acid group on the
long-chain macrqmolecules via the carbodiimide pathway. The reaction proceeds as follows:

where:
Structure A is a carbodiimide;
Structure B is a polycarboxylate (though only one CO2H group is shown);
Structure C is the product of Reaction A and is an activated O-acylisourea;
Structure D is a primary amine having a hydroxyphenyl group;
Structure E is a hydroxyphenyl-substituted polycarboxylate; and
Structure F is an acylurea byproduct;
wherein individual Rs can be individually selected, the same or different from one
another, to be a straight chain or branched alkane or acyl group, or any other structure that
does not interfere with the carbodiimide reaction pathway to provide the amide bond between
the NH2 and CO2H groups as shown in Structure E above.

[0031] In the above-illustrated pathway, Reaction A represents a carbodiimide
activation of the carboxyl group to provide an activated O-acylisourea intermediate. The
electropositive carbon atom of this intermediate is receptive to nucleophilic attack by the lone
pair of electrons on a nitrogen atom of an adjacent primary amine molecule having an
attached hydroxyphenyl group. The products of this nucleophilic substitution reaction
(Reaction B) are a hydroxyphenyl-substituted polycarboxylate and an acylurea byproduct
which can be dialyzed out to provide a substantially pure hydroxyphenyl-substituted
polycarboxylate product.
[0032] Certain side-reactions are possible in the above-described carbodiimide
reaction pathway chemistry and should be considered by the person having ordinary skill in
the art. First, the carbodiimide can react with nucleophiles other than the carboxylate oxygen
atom of the polycarboxylate molecule required to form the desired O-acylisourea (reaction
A). Such nucleophiles may include the amine and/or hydroxyphenyl groups of Structure D
illustrated above. In particular, there are three potential side-reactions for Reaction A which
can reduce the effective concentration of the carbodiimide and the primary amine having the
hydroxyphenyl group (Structures A and D), and potentially lead to the creation of undesired
adducts on the polycarboxylate (Structure B):

[0033] The product of an amine reaction with the carbodiimide (Reaction C) will not
have a free amine group effectively reducing the amount of tyramine available for reaction
with the O-acylisourea. This reaction also reduces the amount of carbodiimide available for
formation of the desired O-acylisourea. The products of the hydroxyphenyl reaction
(Reaction D) are not UV absorbent, which will make their detection by UV-spectroscopy in
the final hydroxyphenyl-substituted polycarboxylate product (explained below) more
difficult. However, because these products still contain free amine groups, they can form
amide bonds with the polycarboxylate molecule via Reaction B. This can give rise to two
unproductive hyaluronan-substituted structures, neither of which can participate in the
peroxidase cross-linking reaction in the second step (described below) of preparing the cross-
linked network due to the absence of an extractable phenolic hydroxyl hydrogen atom needed
to generate the free radical (also explained below). Finally, the carbodiimide can react non-

productively with water (Reaction E) to produce the same acylurea shown above as a
byproduct of Reaction B, but with none of Structure E, the desired product.
[0034] Once the desired O-acylisourea product has been formed in Reaction A, there
is again the possibility for certain additional side-reactions:

[0035] The O-acylisourea (Structure C) can be hydrolyzed as shown in Reaction F
releasing the original unmodified polycarboxylate (Structure B) and the acylurea of the

carbodiimide (Structure F). This is an unproductive reaction similar to reaction E, which
reduces the effective concentration of the carbodiimide. The O-acylisourea, can also undergo
an intramolecular rearrangement (Reaction G) to form two unreactive N-acylureas. These
structures form unproductive adducts on the carboxylate molecule which cannot contribute to
the peroxidase catalyzed cross-linking reaction (step 2 discussed below) for preparing a
network according to the invention. The O-acylisourea can also react (Reaction H) with a
second carboxyl group on either the same or a different polycarboxylate molecule to form an
acid anhydride. This molecule can then react with Structure D to form the desired amide and
regenerate the second carboxyl group. Thus there are two potential side-reactions for the O-
acylisourea, which can reduce the effective concentration of the carbodiimide (Reactions F
and G), and potentially lead to creation of undesired adducts on the polycarboxylate
molecule.
[0036] Negative effects of these side reactions can be addressed through conventional
techniques without undue experimentation.
[0037] Alternatively to the pathway shown above where the macromolecule
(Structure B) is a polycarboxylate, the macromolecule can be a polyamine having multiple or
periodic amine groups along its length, wherein the hydroxyphenyl groups then are provided
as part of smaller carboxylic acid molecules. Suitable polyamines include: polyhexosamines
such as chitosan (polyglucosamine); polyamino acids such as polylysine;
polydeoxyribonucleotides such as poly (dA) (polydeoxyadenylic acid), poly(dC)
(polydeoxycytidylic acid), and poly(dG) (polydeoxyguanylic acid); and polyribonucleotides
such as poly(A) (polyadenylic acid), poly(C) (polycytidylic acid), and poly(G) (polyguanylic
acid). The carbodiimide-mediated reaction pathway proceeds exactly as explained above to
form the amide bond between the amine group and carboxylic acid group except that, as will
be understood by a person having ordinary skill in the art, the resulting product will be
hydroxyphenyl-substituted polyamine instead of a polycarboxylate. Other peptides and/or
proteins also can be used as the macromolecules in the present invention, either which have
hydroxyphenyl groups disposed along their length, or to which hydroxyphenyl groups can be
provided via a substitution reaction as described herein. For example, in addition to the
peptides already disclosed herein, polyarginine can be used as the macromolecule.
[0038] When substituting onto a polycarboxylate molecule, suitable hydroxyphenyl-
containing compounds for use in the present invention include those having a free primary

amine that can be used to modify scaffold materials having multiple or periodic CO2H
groups, including tyrosine (2-amino-3-(4-hydroxyphenyl) proprionic acid) and tyramine
(tyrosamine or 2-(4-hydroxyphenyl) ethylamine). When substituting onto a polyamine,
suitable hydroxyphenyl-containing compounds include those having a free CO2H group that
can be used to modify scaffold materials having multiple or periodic primary NH2 groups,
including tyrosine, 3-(4-hydroxyphenyl) propionic acid and 4-hydroxyphenylacetic acid.
[0039] The second step in preparing a cross-linked macromolecular network
according to the invention is to link the resulting macromolecules, now having one or more
hydroxyphenyl groups attached, via a dihydroxyphenyl linking structure. In this step
hydroxyphenyl groups attached to different macromolecules are linked via the reaction
mechanism shown below using a peroxide reagent in the presence of a peroxidase:

[0040] (It is noted that some dihydroxyphenyl linking may occur between different
hydroxyphenyl groups attached to the same molecule as well). Peroxidase in the presence of
a dilute peroxide (preferably H2O2) is able to extract the phenolic hydroxyl hydrogen atom
from hydroxyphenyl containing compounds (such as tyramine) leaving the phenolic hydroxyl
oxygen with a single unshared electron, an extremely reactive free radical. The free radical
isomerizes to one of the two equivalent ortho-position carbons and then two such structures
dimerize to form a covalent bond effectively cross-linking the structures, which after
enolizing generates a dihydroxyphenyl dimer (a dihydroxyphenyl linkage such as dityramine
linkage as described below).
[0041] For clarity, only a single dihydroxyphenyl linking reaction is shown above,
but it will be understood that several or multiple such linkages will be produced when
macromolecules having attached hydroxyphenyl groups are subjected to the reaction
conditions (peroxide and peroxidase). Hydrogen peroxide is indicated in the above
mechanism, but other suitable peroxides can be used. Also, the peroxidase preferably is
horseradish peroxidase (HRP). Alternatively, any other suitable enzyme (or other agent) can
be used that is capable of generating free-radicals for cross-linking scaffold materials
containing hydroxyphenyl groups, preferably under ordinary metabolic conditions as
described below.
[0042] We have shown that the interaction of horseradish peroxidase (Type II) and
hydrogen peroxide (H2O2) is suitable for the production of cross-linked macromolecular
networks. The mechanism comprises four distinct steps: (a) binding of peroxide to the heme-
Fe(IH) complex of the peroxidase to form an unstable peroxide complex, "Compound I"; (b)
oxidation of the iron to generate a ferryl species with a pi-cation radical in the heme
porphyrin ring, "Compound II"; (c) reduction of Compound II by one substrate (i.e.
hydroxyphenyl or water) molecule to produce a product (i.e. hydroxyphenyl or superoxide)
radical and another ferryl species, "Compound III"; (d) reduction of Compound HI by a
second substrate (i.e. hydroxyphenyl or water) molecule to release a second product (i.e.
hydroxyphenyl or superoxide) radical and regenerate the native enzyme. Thus the peroxidase
enzyme can either form hydroxyphenyl radicals required for cross-linking through interaction
of hydroxyphenyl groups at the enzyme active site to directly create the desired radicals or
through first generation of superoxide radicals, which then diffuse from the enzyme and
interact with hydroxyphenyl groups to generate the desired radicals. Other compounds that

have the potential to produce the same effect include any porphyrin containing compound
(i.e. Photofrin below), which includes the peroxidase family, hemoproteins, or the structurally
related chlorin compounds.
[0043] A number of other free radical initiators can be used to crosslink the
hydroxyphenyl modified macromolecules described herein. A majority are based on the
formation or inclusion of reactive oxygen species (ROS) such as, but not limited to,
molecules of hydrogen peroxide, ions of hypochlorite, radicals like the hydroxyl radical, and
the superoxide anion which is both ion and radical. Additional reactive molecules such as
reactive nitrogen species or reactive sulfur species, or those free radical species involved in
synthetic polymerization have the potential to be used for hydroxyphenyl cross-linking.
[0044] ROS are commonly produced in nature through the use of enzymes, and
substrates. Additional enzymatic systems which have the potential to be used in the cross-
linking process, as a result of production of superoxide radicals, include, but are not limited
to xanthine-xanthine oxidase and NADPH-NADPH oxidase.
[0045] Another class of ROS free radical initiators that can be used involves the use
of metallic cations. One example is based on the Fenton reaction, which takes place between
hydrogen peroxide and a bivalent cation, such as Fe2+. This process generates powerful free
radicals when the catalyst reacts with hydrogen peroxide. The principal chemical reaction
associated with Fenton's reaction is shown below:

where, Fe2+ = ferrous ion, Fe3+ = ferric ion, OH. = hydroxyl radicals
[0046] In addition to the initiation reaction described above that produces hydroxyl
radicals, the Fenton's process can also produce superoxide radicals and hydroperoxide anions
by additional chain propagation reactions described below. The perhydroxyl radical is known
to be a weaker reductant compared to superoxide radical and hydroperoxide anions.

where O2. = superoxide radical anion, HO2 = hydroperoxide anion, HO2' = perhydroxyl
radical.

[0047] We have demonstrated the ability for this reaction to crosslink tyramine
substituted hyaluronan in the laboratory using ferrous sulfate in conjunction with hydrogen
peroxide. Compounds which include, but are not limited to, bivalent cations of copper,
chromium, vanadium and cobalt can be used in a similar manner. It is to be noted that while
the hydroxyl free radical can be used to form a dityramine crosslink, it has also been shown
to cleave HA chains, and thus may ultimately be unsuitable for ideal hydrogel formation.
Additional molecules or methods which can generate ROS include:
• rubidium or cesium ions in the presence of oxygen to form superoxide radicals;
• trivalent cations, which with hydrogen peroxide form free radicals and bivalent
cations as shown below, which can subsequently follow the reactions involved in
the Fenton process.

• the cytotoxic and antitumor therapy Photofrin, which upon illumination with laser
light at a wavelength of 630 nm causes propagation of a radical generating
reaction that produces superoxide and hydroxyl radicals. In the absence of light,
but the presence of hydrogen peroxide, the porphorin ring in Photofrin should
operate by the same reaction as for the peroxidase enzyme above.
• UV light and hydrogen peroxide to form hydroxyl and superoxide free radicals.
• the persulfate family in combination with TEMED.
[0048] As noted above, one alternative method for generating such free-radicals is to
use Photofrin as an alternative, non-enzymatic, light-activated cross-linking agent to cross-
link the macromolecular network described herein, e.g. tyramine-substituted hyaluronan to
form tyramine cross-linked hyaluronan hydrogels. Photofrin®, which is known in the art,
generates free radicals which could initiate the cross-linking reaction as described herein in a
manner similar to the peroxidase-H2O2 mechanism described above. Photofrin® is a
porfimer sodium manufactured in powder or cake form by Wyeth-Ayerst Lederle Parenterals,
Inc.
[0049] The dihydroxyphenyl cross-linked macromolecular network is superior to
conventional cartilage or other tissue replacement or substitution methods and products, at
least with respect to the ability to carry out an in situ cross-linking procedure, because the
preferred cross-linking reaction is enzyme driven (peroxidase). This means the cross-linking

reaction is carried out under ordinary in vivo or metabolic conditions of temperature such as
35-39°C (e.g. about 37°C), pH range of 6-7 (e.g. about 6.5), reagents etc. (A peroxide, such
as hydrogen peroxide, is the only required reagent for the cross-linking reaction). In addition,
Photofrin already is used in in vivo applications, e.g. ablative treatment of Barrett's
esophagus, and the iron-based cross-linking mechanism also can be optimized for in vivo
performance. Thus, the cross-linking reaction can be performed in vivo, to provide a cross-
linked hydrogel at a surgical situs, such as an orthopedic surgical situs, to promote maximum
seamless integration between the hydrogel and native tissue such as bony and cartilaginous
tissue. Integration of the new hydrogel scaffold with native cartilage matrix may occur
immediately as the hydroxyphenyl-substituted macromolecular scaffold quickly penetrates
into the existing cartilage matrix prior to cross-linking, and cross-links not only with other
hydroxyphenyl-substituted macromolecular scaffold material but potentially with tyrosine
residues of resident proteins in the existing cartilage matrix. This would eliminate a typical
problem found with pre-formed matrix plugs, which is their poor integration into the native
cartilage tissue. The ability to cross-link the hydrogel directly on the articular surface
eliminates the need to surgically enlarge a defect to fit a pre-cast plug, as is necessary for
hydrogels whose chemistries are toxic to or otherwise prohibit their formation inside the
patient. It should be noted that most cartilage damage as a result of arthritis presents as a
variable thinning of the articular surface, not holes of defined shape.
[0050] For the peroxidase mechanism, because the cross-linking reaction requires
both the peroxide and a peroxidase (preferably horseradish peroxidase), solutions containing
all but one of these components can be prepared for convenient application to a surgical site.
For example, a solution comprising a tyramine - (or other hydroxyphenyl containing species)
substituted polycarboxylate (such as tyramine-substituted hyaluronan, etc.) and the
peroxidase can be prepared, with a second solution prepared containing the peroxide.
Alternatively, the peroxide and the peroxidase can be swapped between the first and second
solutions, the importan thing being that the peroxide and peroxidase are kept separate (i.e. in
separate solutions) until the cross-linking reaction is to be carried out. Then, the first solution
is applied, (e.g. to an in vivo surgical situs), and the second solution is applied or sprayed
over the first, in vivo, to cause in situ cross-linking of the tyramine residues. The cross
linking reaction occurs in vivo. Other combinations will be evident from the present
disclosure which are within the skill of a person of ordinary skill in the art.

[0051 ] Furthermore, because the cross-linking reaction occurs under ordinary
metabolic conditions, additional living cells, such as chondrocytes, progenitor cells, stem
cells, etc., can be provided directly to a medium containing the non-cross-linked
hydroxyphenyl-substituted polycarboxylates or polyamines (or polyphenols), i.e. to the first
or second solution from the preceding paragraph, wherein the cell-rich medium is applied
with the macromolecules to the site in vivo, and the molecules are subsequently cross-linked
via addition of peroxidase and peroxide. The result is a cross-linked macromolecular
network containing the desired cells dispersed within it. Such a cell-enriched network is not
possible in conventional tissue replacement matrices due to the harsh conditions of
temperature and pH under which they are prepared. Further, as described below in Example
5, it has been demonstrated that the cells provided to the network as described above remain
viable even after cross-linking of tyramine-substituted hyaluronan (also described below).
[0052] In a preferred embodiment particularly suitable for preparing synthetic
cartilage as well as other synthetic or artificial tissues, the macromolecule used to produce the
network is hyaluronan or hyaluronic acid (HA), and the hydroxyphenyl group is supplied in
the form of tyramine. Hyaluronan (HA) is a ubiquitous molecule, which is most concentrated
in specialized tissues such as cartilage, vocal cords, vitreous, synovial fluid, umbilical cord,
and dermis. In these tissues, its function is manifold, influencing tissue viscosity, shock
absorption, wound healing, and space filling. HA has been shown to influence many
processes within the extracellular matrix (ECM) in native tissues where it is present including
matrix assembly, cell proliferation, cell migration and embryonic/tissue development.
[0053] HA is composed of repeating pairs of glucuronic acid (glcA) and N-
acetylglucosamine (glcNAc) residues linked by a B 1,3 glycosidic bond as shown in Fig. 3.
The glucuronic acid residue is particularly pertinent to the production of a macromolecular
network as described herein as this sugar provides an available carboxyl group periodically
along the repeat disaccharide structure of HA that is useful for hydroxyphenyl, i.e. tyramine,
substitution. For each hyaluronan chain, this simple disaccharide is repeated up to 10,000
times or greater resulting in macromolecule that can have a molecular weight on the order 10
million daltons (10 megadaltons). Adjacent disaccharide units of HA are linked by a B1.4
glycosidic bond, also seen in Fig. 3. Each glcA residue has a carboxylic acid group (CO2H)
attached to the number 5 carbon atom of the glucose ring. Under biological conditions, HA is
a negatively charged, randomly coiled polymer filling a volume more than 1,000 times

greater than would be expected based on molecular weight and composition alone. As noted
above, the strong negative charges attract cations and water, which allow HA to assume the
form of a strongly hydrated gel in vivo, giving it a unique viscoelastic and shock-absorbing
property. HA represents a readily available and desirable scaffolding material for tissue
engineering applications as it is non-immunogenic, non-toxic and non-inflammatory. Also as
a naturally occurring extracellular matrix (ECM) molecule it offers the advantages of being
recognized by cell receptors, of interacting with other ECM molecules, and of being
metabolized by normal physiological pathways.
[0054] Tyramine is a phenolic molecule having an ethyl amine group attached para to
the OH group on the benzene ring. When these species are used, the mechanism for tyramine
substitution onto the singly bound oxygen atom of a CO2H group on HA proceeds via the
carbodiimide-mediated reaction mechanism described above as illustrated immediately
below. The preferred carbodiimide species is l-ethyl-3-(3-
dimethylaminopropyl)carbodiimide (EDC) as shown.

where:
Structure A is EDC;
Structure B is hyaluronan (though only one CO2H group is shown);
Structure C is the product of Reaction A and is l-ethyI-3-(3-dimethylaminopropyI)
isourea;
Structure D is tyramine;
Structure E is tyramine-substituted hyaluronan; and
Structure F is l-ethyl-3-(3-dimethylaminopropyl) urea (EDU).
[0055] In the above pathway, a negatively charged oxygen atom of the carboxyl
group of the hyaluronan molecule attacks, via a nucleophilic reaction mechanism, the
electron-deficient diimide carbon atom on the carbodiimide molecule (EDC) to form the
activated O-acylisourea (Reaction A). The result is that the carbon atom of the HA
carboxylate group becomes sufficiently electron deficient to be susceptible to nucleophilic
attack by the unshared pair of electrons on the amine group of a tyramine molecule (Reaction
B). Reaction A is preferably catalyzed by a suitable catalyst that will result in the formation
of an active ester during Reaction A, thus permitting the reaction to be carried out at
substantially neutral pH (e.g. pH=6.5). Suitable catalysts include N-hydroxysuccinimide
(NHS), less preferably 1-hydroxybenzotriazole (HOBt) or N-hydroxysulfosuccinimide
(NHSS), less preferably another suitable catalyst or combinations thereof effective to enhance
the carbodiimide reaction by formation of an active ester in order to minimize the
unproductive hydrolysis of carbodiimides at higher pHs. Less preferably other carbodiimides
besides EDC can be used, including l-cyclohexyI-3-[2-(4-
methylmorpholino)ethyl]carbodiimide (CMC), and dicyclohexylcarbodiimide (DCC).
[0056] The result of Reaction A above is O-acylisourea-substituted hyaluronan;
essentially the EDC molecule has been temporarily substituted onto the carboxylic acid group
of a glcA residue from the HA molecule, making the carbon atom of the carboxylic acid
group slightly positively charged. The electron pair from the terminal amine group of a
tyramine molecule is then substituted onto the carbon atom via a nucleophilic substitution
reaction as explained in the preceding paragraph (Reaction B). The result of Reaction B is
the tyramine-substituted HA molecule (T-HA) and acylurea, a byproduct. It will be
understood that Reactions A and B will result in a plurality of tyramine substitutions on the

periodic glcA residues of HA molecules; a single substitution has been shown here for
brevity and clarity.
[0057] After formation of T-HA, a plurality of T-HA molecules are reacted via
peroxide and peroxidase enzyme to cross-link T-HA molecules as previously described and
illustrated above. That is, the hydroxyphenyl groups on the tyramine residues now attached
to HA molecules react with peroxide (preferably H2O2) in the presence of a peroxidase to
remove the phenolic hydrogen atom resulting in a tyramine free radical, with the unpaired
electron associated with the phenolic oxygen atom. This free radical species isomerizes or
resonates, resulting in a resonance structure (or free radical isomer) with the unpaired
electron now associated with an ortho carbon atom on the phenolic ring. In this position, the
unpaired electron quickly reacts with a similarly situated unpaired electron on another
tyramine free radical to form a covalent bond therebetween. The result is a free-radical
driven dimerization reaction between different tyramine free radical residues attached to
different glcAs of the same or different HA molecules. This dimerized species further
enolizes to restore the now-linked tyramine residues, resulting in a dityramine linkage
structure. It will be understood that a plurality of reactions as herein described will occur
between adjacent tyramine residues, resulting in a cross-linked macromolecular network of
T-HA molecules having the following cross-linking structure:

[0058] The cross-linked T-HA network can be provided with aggrecan molecules in a
conventional manner, e.g. via link proteins, to provide a cross-linked T-HA network having
aggrecan molecules attached to the HA chains. Thus, a network similar to that found in a
normal cartilage aggregate can be provided, with the dityramine bonds holding the network
together thereby constraining the contained aggrecan network, instead of collagen fibrils as in
normal cartilage.

similarly non-displaceable chondroitin sulfate network (and concomitant water-
impermeability) as in normal cartilage, but without the extracellular collagen fibril matrix or
the HA chains found in normal cartilage. In fact, by directly cross-linking chondroitin sulfate
molecules, (instead of their core HA molecules as in the previously described embodiment),
the repulsive force between adjacent chondroitin sulfate molecules may be strengthened,
resulting in even stronger fluid flow resistance compared to normal cartilage. This may result
in greater loading force absorption and dissipation capacity than normal cartilage because the
interstitial fluid phase is even more constrained from flowing. In this embodiment, where
chondroitin sulfate molecules are directly cross-linked, certain cartilage degenerative
conditions are entirely circumvented; e.g. conditions where the core protein to which
chondroitin sulfate molecules are ordinarily bonded in normal cartilage becomes cleaved
between the HA binding domain (G1) and the second globular domain (G2) thus allowing the
chondroitin sulfate rich region to diffuse out from the cartilage aggregate. In this
embodiment, because the chondroitin sulfate molecules are directly cross-linked to one
another, unassociated with an aggrecan or other proteoglycan molecule, they cannot be
cleaved or carried away as in normal cartilage.
[0062] Nonetheless, a tyramine cross-linked T-HA network (having an HA backbone
chain with attached aggrecan molecules, which in turn include chondroitin sulfate chains)
may be preferred because of the high availability of HA. This may be beneficial in the case
of cartilage replacement or repair using the present invention, because the body's normal
metabolic pathway for generating cartilage may be able to build directly onto an implanted
tyramine cross-linked T-HA network as will be described.
[0063] One further particular application where a cross-linked network according to
the invention will have substantial utility is in the production of an artificial kidney. The
kidney filters blood by two mechanisms: one is by size exclusion and the second is by charge
exclusion. MEMS devices have been designed for use in artificial kidney devices, which
contain precisely defined micropores that can effectively mimic only the size exclusion
characteristics of the kidney. In a healthy kidney, the charge exclusion related filtration is the
result of heparan sulfate proteoglycans present in a basement membrane, which separates two
distinct cell types important for other kidney related functions. To mimic this charge barrier
in the MEMS engineered artificial kidney, hydrogels can be prepared composed of either
heparan sulfate or heparin that are cross-linked via dihydroxyphenyl (dityramine) links as

described herein and provided within the pores of the MEMS device. This heparin/heparan
sulfate hydrogel can then be sandwiched between two hyaluronan derived hydrogels (e.g. T-
HA described above) as described herein, and containing one of each of the cell types
normally found in a normally functioning kidney. The central heparin/heparan sulfate
hydrogel provides the charge exclusion properties for the device. The outer two hyaluronan
hydrogel layers provide protection from the immune system and fouling by normal cellular
and molecular debris. Inclusion of the two cell types on opposite sides of the filtration barrier
provides a cellular component in its normal physiologic orientation.
[0064] In another promising application, the hydrogels herein described can be
applied in developing an artificial pancreas. A problem in development of an artificial
pancreas is the short half life of MEMS engineered glucose sensors due to fouling of the
detector electrode in vivo. Coating of the surface of these detectors with a hyaluronan
hydrogel (e.g. T-HA) as described herein would permit diffusion of the small molecular
weight glucose molecules that they are designed to detect while providing protection from the
immune system and fouling by normal cellular and molecular debris.
[0065] In summary, it will be evident from the foregoing that macromolecules useful
as scaffold materials for formation of hydrogels include but are not limited to
polycarboxylates (containing free carboxylate groups), polyamines (containing free primary
amine groups), polyphenols (containing free hydroxyphenyl groups) and their copolymers,
examples of which have been described above. When polyphenols are used, the first step in
preparing the network described above can be omitted because polyphenols already contain
multiple or periodic hydroxyphenyl groups. Otherwise, both polycarboxylates and
polyamines must have hydroxyphenyl groups added or substituted along their length,
preferably via the above-described carbodiimide reaction pathway. The second step in
preparing the network is to carry out an enzyme driven dimerization reaction between two
hydroxyphenyl groups attached to adjacent macromolecules (whether polycarboxylates,
polyamines or polyphenols) in order to provide a cross-linked structure. This step is carried
out using a peroxide reagent (preferably hydrogen peroxide) in the presence of a suitable
enzyme (preferably HRP) under metabolic conditions of temperature and pH.
[0066] In the case of the preferred dityramine cross-linked T-HA network, in the first
step the carboxyl groups on high molecular weight hyaluronan (HA) are substituted with
tyramine which introduces reactive hydroxyphenyl groups into the HA molecule. This

tyramine substitution reaction preferably is mediated by the carbodiimide, l-ethy]-3-(3-
dimetbylaminopropyl)carbodiimide (EDC) with the degree of tyramine substitution on HA
controlled by the molar ratios and absolute concentrations of tyramine, EDC and HA used in
the reaction mix. Excess reagents such as unused tyramine and EDC are subsequently
removed by dialysis, allowing isolation and recovery of high molecular weight tyramine-
substituted HA (T-HA). The percent tyramine substitution within each T-HA preparation is
easily calculated by measuring: 1) the concentration of tyramine present in the preparation,
which is quantitated spectrophotometrically based on the unique UV-absorbance properties of
tyramine at 275 ran (see Example 2 below); and 2) the concentration of total carboxyl groups
in the HA preparation, which is quantitated spectrophotometrically by a standard hexuronic
acid assay. By this technique, T-HA preparations which contain a percent tyramine
substitution of only 4 - 6% have been routinely synthesized experimentally. At this level of
tyramine substitution, the vast majority (preferably at least 60, 70, 80, 90, or 95, percent) of
the HA molecule remains chemically unaltered, and therefore biologically functional. From
this formulation of T-HA (i.e. 4 - 6% tyramine substitution) a wide range of biomaterials with
a wide range of physical properties can be produced by simply varying the concentration of
the T-HA used in the second step of the process.
[0067] In the cross-linking reaction, solutions of T-HA are cross-linked to form
hydrogels through an enzyme (peroxidase) driven reaction, which catalyzes the formation of
a covalent bond between two tyramine adducts on adjacent HA molecules, producing a single
dityramine cross-link. The formation of multiple, e.g. hundreds, of these dityramine cross-
links per HA molecule result in formation of a stable 3-dimentional scaffold or hydrogel.
Addition of very dilute peroxide (preferably H2O2) is required to initiate the cross-linking
reaction as it is the peroxide, not -HA, that is the actual substrate for the peroxidase enzyme.
The products of the reaction of the peroxidase enzyme on peroxide are free radicals that are
preferentially taken up by the hydroxyphenyl rings of tyramine resulting in the formation of
the dityramine cross-links. The dityramine linked structures are fluorescent blue (see
Example 2), a property which is used to both image the hydrogc ■ and to quantify the degree
of cross-linking within the hydrogels. Since the cross-linking reaction is enzyme driven, the
hydrogels can be formed under physiologic conditions, and therefore can be formed in the
presence of included cells or bioactive agents, or directly adjacent to living tissue while
maintaining cell and tissue viability.

[0068] The resulting hydrogels are optically clear with a wide range of physical
properties depending on the initial T-HA concentration. For example, hydrogels formed from
T-HA solutions of 6.25, 12.5,25, 50 and 100 mg/ml T-HA have been shown experimentally
to have physical properties (rigidity, rheology and texture) of a jelly, a gelatin, a dough, a
resilient rubber-like composition (similar to a rubber ball), and a cartilage-like material
respectively- see Example 3. These materials have potential applications in a wide range of
clinical settings including tissue engineering of both orthopedic (i.e. cartilage, bone, tendon,
meniscus, intervertebral disk, etc.) and non-orthopaedic (kidney, liver, pancreas, etc.) tissues,
gene and drug delivery, coating of non-biological devices for in vivo implantation (i.e.
glucose sensors, artificial hearts, etc.), wound repair, biosensor design, and vocal cord
reconstruction.
[0069] Advantageous properties of the hydrogels described herein include the ability
to: 1) provide easy characterization and quality control; 2) integrate with existing tissue
matrices; 3) directly incorporate into newly formed matrices; 4) directly include cells and
bioactive factors; 5) maintain biocompatibility, 6) control bioresorption; 7) cast easily into
complicated anatomical shapes (see Example 4 below); and 8) exhibit the mechanical
properties of native tissues such as articular cartilage.
[0070] Current biologically-based surgical procedures for cartilage repair include
autologous chondrocyte implantation, drilling, abrasion chondroplasty, microfracture, and
mosaic arthroplasty. All these procedures treat only focal articular cartilage injuries, and not
cartilage denuded joint surfaces such as seen in severe osteoarthritis and rheumatoid arthritis.
Also, they use either cartilage tissue plugs or expanded chondrocytes harvested from the
patient to fill cartilage defects. These tissues or chondrocytes are expected to fill the defect
by synthesizing entirely de novo material, such as newly synthesized hyaline cartilage, that
has integrated with existing cartilage matrices and has the biomechanical properties of normal
cartilage. However, such procedures all promote the formation of a reparative tissue
(fibrocartilage) rather than true hyaline cartilage with further mechanical damage to
fibrocartilage thought to predispose the joint to osteoarthritis. Furthermore, the availability
of endogenous cartilage as a repair material is quite limited with its acquisition presenting its
own risks and morbidity to the patient. As evident from the foregoing discussion and as will
become further apparent based on the following Examples, the synthetic macromolecular
networks and resulting hydrogels disclosed herein present practical materials for promising

new therapies in patients suffering from cartilage degenerative diseases. The materials are
entirely synthesized from commercially available ex vivo reagents and so involve no
morbidity to the patient which conventionally would be required to harvest endogenous
material. In addition, the hydrogel (particularly T-HA) can be implanted as an effective
cartilage substitute in cartilage denuded joints as a direct intervention for patients suffering
from cartilage-degenerative diseases because they can be synthesized so as to emulate the
behavior of normal, healthy cartilage.
[0071] Rather than relying on synthetic or natural materials or on chondrocytes to
produce de novo an implantable, synthetic cartilage-like extracellular matrix (ECM), the
present inventors initially focused on purifying the molecules that give cartilage its form and
structural characteristics, and then minimally modifying these molecules to make a material
resistant to biological degradation. While chondrocytes still may be relied on for
maintenance of the synthetic ECM provided by the macromolecular (e.g. T-HA) network
post-implantation (e.g. chondrocytes can be embedded into the hydrogel materials as
described above), they are not relied on for de novo synthesis. Instead, the basic structure of
the synthetic materials described here is modified by cross-linking via a dihydroxyphenyl,
preferably dityramine linkage chemistry, to ensure its survival. On further development and
experimentation, as will be seen in the following Examples it was discovered mat hydrogels
can be made from such materials having a wide array of viscoelastic and other physical
properties that can be tuned by appropriate and judicious selection of reagent concentrations
and cross-linking conditions to approximate or simulate the properties of other native tissues
for which it is or may be desirable to provide a synthetic implantable substitute.
[0072] As the Examples below demonstrate, the present hydrogels can be prepared
having widely varying properties that are suitable for any number of synthetic tissue
implantation or augmentation, as well as other clinical applications. As already described,
the present materials can be used to repair cartilage defects produced as a result of either
injury or disease. Defects due to injury that can be so repaired can be sports- or accident-
related, and may involve only the superficial cartilage layer, or may include the underlying
subchondral bone. Defects due to disease which can be repaired using the compositions
described herein include those resulting from osteoarthritis and rheumatoid arthritis. Whether
from injury or disease, such defects may be in either mature or growth plate cartilage.
Formulations for hydrogels for synthetic growth plate cartilage may require the inclusion of

unsubstituted scaffold material in order to allow for controlled bioresorption of the
biomaterial during growth.
[0073] Another potential clinical application for treatment of damaged or arthritic
joints is as a replacement for synovial fluid. Conventionally referred to as
viscosupplementation therapy, this currently involves injection of a solution of uncross-
linked HA into a damaged or arthritic joint, which provides sustained pain relief for weeks
even though the HA is cleared from the joint in 1-2 days. Use of the T-HA hydrogels
described herein should provide an extended benefit due to their longer i?i vivo persistence
compared to uncross-linked HA.
[0074] Another field where the hydrogels described herein will be useful is the repair,
reconstruction or augmentation of cartilaginous as well as soft tissues of the head and neck.
The availability of biomaterials for soft tissue augmentation and head and neck reconstruction
has remained a fundamental challenge in the field of plastic and reconstructive surgery.
Significant research and investment has been undertaken for the development of a material
with appropriate biological compatibility and life span. The outcomes of this research have
not been promising. When placed in immunocompetent animals the structural integrity of
currently proposed materials has been shown to fail as the framework is absorbed.
Furthermore, though conventional synthetic materials offer excellent lifespan, they present
certain unavoidable pitfalls. For example, silicones have been fraught with concerns of
safety and long-term immune related effects. Synthetic polymers PTFE (gortex) and silastic
offer less tissue reactivity but do not offer tissue integration and can represent long term risks
of foreign body infections and extrusion. The materials described in this application will be
useful to prepare a synthetic soft-tissue scaffold material for the augmentation or repair of
soft-tissue defects of the head and neck. In particular, a cross-linked tyramine-substituted
hyaluronan (T-HA) hydrogel, which is non-inflammatory, non-immunogenic, and which can
be prepared having the appropriate degree of viscoelasticity (see Examples below), could be
used as an effective implantable scaffold material. In addition, the unique ability of the
preferred enzyme-driven cross-linking chemistry to maintain cell viability permits inclusion
of cells such as chondrocytes directly into the hydrogels during formation which can be
performed in situ at a defect site. Thus, the need to sculpt or mold an anatomically
compatible graft shape to fit a particular defect site is eliminated.

[0075] The dityramine cross-linked T-HA network described above has particular
utility for producing artificial or synthetic cartilage. The present hydrogel materials can be
used, for example, as a novel, biocompatible and biocompliant material to prepare cartilage
implants which are frequently used in reconstructive procedures of the head and neck to
repair cartilaginous or bony defects secondary to trauma or congenital abnormalities.
Applications specific to the ear include otoplasty and auricular reconstruction, which are
often undertaken to repair cartilaginous defects due to trauma, neoplasm (i.e., squamous cell
carcinoma, basal cell carcinoma, and melanoma), and congenital defects such as microtia.
Applications specific to the nose include cosmetic and reconstructive procedures of the nose
and nasal septum. Dorsal hump augmentation, tip, shield and spreader grafts are frequently
used in cosmetic rhinoplasty. Nasal reconstruction following trauma, neoplasm, autoimmune
diseases such as Wegeners granulomatosis, or congenital defects require cartilage for repair.
Septal perforations are difficult to manage and often fail treatment Cartilage grafts would be
ideal for these applications, as autologous or donor cartilage is often unavailable.
Applications specific to the throat include laryngotracheal reconstruction, which in children
usually requires harvesting costal cartilage, which is not without morbidity. Auricular and
septal cartilage is often inadequate for this application. Synthetic cartilaginous materials
prepared from hydrogels disclosed herein can be synthesized to suit each of the foregoing
applications, based on tuning parameters of hydrogel synthesis such as reagent concentration,
substitution and cross-Unking rates, etc., as evident from the below Examples.
Laryngotracheal reconstruction is usually performed for airway narrowing due to subglottic
or tracheal stenosis. The etiology may be traumatic (i.e., intubation trauma, or tracheotomy)
or idiopathic. Other possibilities include chin and cheek augmentation, and use in ectropion
repair of the lower eyelid, in addition to numerous craniofacial applications. It should be
noted that these applications may not need cartilage with the exacting mechanical properties
of articular cartilage. Inclusion of a cell population or bioactive agents may also be desirable.
[0076] The hydrogel materials described herein also can be used for repair and
narrowing of the nasal cavity, normally following overly aggressive surgical resection, to
prevent the chronic pooling of fluid in the nasal passages that leads to infection and
encrustation. Another promising application is in laryngotracheal reconstruction in both
children and adults, as a result of laryngotracheal injury due for example to intubation during
a surgical procedure such as cardiovascular surgery. Damaged tracheal cartilage at the
anterior and posterior portion of the tracheal ring can be replaced with pre-cast hydrogel

bile ducts of the liver. The bile ducts form a branching network within the liver
and exit the liver via two main branches that are combined into the common bile
duct which drains the liver and gallbladder of bile into the duodenum. The bile
ducts are very narrow in diameter, measuring only up to 2 nun normally at their
largest most distal portions, and yet they must normally drain liters of bile every
day from the liver into the duodenum. Any blockage of these ducts can result in a
serious condition known as jaundice, which allows many toxins and especially
hemoglobin breakdown products to accumulate in the body. PSC is a scarring or
stricturing disease of the bile ducts within the liver and in the extrahepatic bile
ducts described above that connect the liver to the small intestine. The bile duct
strictures of PSC may be treated or prevented with the present hydrogels.
for treatment of chronic pancreatitis. Chronic pancreatitis is a chronic
inflammatory disease of the pancreas that may be complicated by scars or
strictures of the pancreatic ducts. These strictures block the drainage of pancreatic
juice, which normally must exit the pancreas through a system of ducts or
drainage conduits into the small intestine. The pancreatic juice contains many
digestive enzymes and other elements important to normal digestion and nutrient
absorption. Blockage or narrowing of the pancreatic ducts by chronic pancreatitis
can results in severe complications in which the pancreas autodigests and forms
life-threatening abdominal infections and or abscesses. The pancreatic strictures
of chronic pancreatitis may be treated or prevented with the present hydrogels.
for treatment of gallstone-induced bile duct and pancreatic duct strictures.
Gallstones are a very common disorder, a principal complication of which is the
formation of bile duct and pancreatic duct strictures, which may be treated or
prevented with the hydrogels.
for treatment of ischemic bowel disease. The intestines are prone to the formation
of scars or strictures when their blood supply is compromised. Compromised
blood flow is called ischemia, and can be caused by many pathologies, including
cardiovascular disease, atherosclerosis, hypotension, hypovolemia, renal or
hepatic disease-induced hypoalbuminemia, vasculitis, drug-induced disease, and
many others. The end stage result of all of these etiologies can result in intestinal
strictures that block off the bowel and prevent its normal function. The present
hydrogels may be used to treat or prevent ischemic bowel strictures,
for treatment of radiation-induced intestinal strictures. Radiation therapy for

cancer is associated with numerous morbidities, important among which is
intestinal stricture formation. The present hydrogels may be used to treat or
prevent radiation-induced intestinal strictures.
[0079] In addition to making synthetic tissues, the hydrogels disclosed here also can
be used to provide a coating for non-biological structures or devices to be used in surgery or
otherwise for in vivo implantation, such as surgical instruments, or ceramic or metal
prostheses. Such a coating would provide a barrier between the non-biologic device material
and living tissue. The role of hydrogels as a barrier for non-biologic devices includes, but is
not limited to: 1) prevention of absorption of macromolecules and/or cells on the surfaces of
non-biologic devices, which can lead to protein fouling or thrombosis at the device surface;
2) presentation of a non-toxic, non-inflammatory, non-immunogenic, biologically compatible
surface for devices made from otherwise non-biologically compatible materials; 3)
compatibility with device function such as diffusion of glucose for a glucose sensor,
transmission of mechanical force for a pressure sensor, or endothelization of a vascular graft
or stent; 4) enhancement of device function, such as providing a charge barrier to an existing
size barrier in a MEMS based artificial nephron; 5) incorporation into non-biologic devices of
a viable cell population entrapped within an aqueous, physiologically compatible
environment; and 6) inclusion of drugs or bioactive factors such as growth factors, anti-viral
agents, antibiotics, or adhesion molecules designed to encourage vascularization,
epithelization or endothelization of the device.
[0080] Based on the foregoing, the hydrogels of the present invention may be used to
provide a non-allergenic coating for a variety of implantable devices including an implantable
glucose sensor for management of diabetes. In addition, the hydrogels may be used to
provide: a charge barrier for the development of MEMS-based artificial nephrons; an
aqueous, physiologically compatible environment in which embedded kidney cells such as
podocytes can be incorporated into a MEMS-based artificial nephron design; and a coating
for implantable MEMS devices designed for a variety of purposes including, but not limited
to, drug delivery, mechanical sensing, and as a bio-detection system.
[0081] The disclosed hydrogels, and particularly a hyaluronan-based hydrogel, also
may be covalently attached to silicon-based devices, e.g. through first covalent attachment of
the primary amine of tyramine to the silicon surface to provide a hydroxyphenyl coated
surface chemistry. This may use the same chemistry used to bind DNA that has been

modified with a free amine to silicon surfaces. The HA-based hydrogel then is covalently
coupled to the hydroxyphenyl coated surface by the same peroxidase driven chemistry used
in its preferred cross-linking mode described above.
[0082] The hydrogels also can be used for coating non-biologic cardiovascular
devices such as catheters, stents and vascular grafts. These would include devices made from
materials conventionally not used because of their biological incompatibility, but which have
superior design characteristics to those devices currently in use. Bioactive factors could be
incorporated into the hydrogels to promote endothelization or epithelization of the hydrogel,
and thus of the implanted device.
[0083] A particularly promising application mentioned above is in the design and
implementation of an implantable artificial glucose sensor for the treatment and management
of diabetes. Effective glycemic control requires frequent measurement of blood glucose
levels, which currently requires a pin prick (or "finger stick") to obtain a blood sample.
There is tremendous clinical interest in a reliable, cost-effective method of blood glucose
measurement and in preventing hypoglycemia, which is the cause of most severe life-
threatening events. From a technological standpoint, microsensors have been very successful
over the last decade in a wide variety of applications. The successful development of a
biocompatible long term implantable glucose sensor would significantly impact routine
monitoring of glucose levels by diabetic individuals and play a major contributory role in the
further development of a bioartificial pancreas.
[0084] A design of a sensor for use during cardiovascular surgery has been published,
Clark IX!, Lyons C, "Electrode system for continuous monitoring in cardiovascular surgery,"
Annals of New York Academy of Science, 102:29-45 (1962).. Subsequently, efforts have
been directed toward developing and testing an implantable device that could mimic the
native glucose/insulin control system. Besides the obvious advantage of serving as part of a
bioartificial pancreas, such a system could be coupled with telemetry hardware and thereby
give the patient advance warning of hypoglycemia.
[0085] Prior work on implantable glucose sensors generally follows one of two
approaches. The first involves placing sensors into blood vessels such as the vena cava or the
carotid artery. The second involves placing sensors subcutaneously. These sensors may
involve a microdialysis probe or more commonly, an amperometric enzymatic-based

transducer. It is believed the risk of thrombosis and hematogenous spread of infection
mitigate against the long term use of intravascular sensors. While the exact relationship
between blood and subcutaneous glucose concentrations is still being investigated, recent
work suggests that mass transfer modeling methods can significantly improve the estimates
of blood glucose levels that are based on subcutaneous data. Furthermore, there are
significant advantages associated with subcutaneous sensors: clinical safety, ease of insertion
and removal, ease of coupling these sensors to a telemetry system and cost. There is
substantial evidence that subcutaneous placement of a glucose sensor will work and will lead
to much longer life of the sensor than if it were to contact blood directly.
[0086] However, a major problem in the design of any continuous glucose sensor for
clinical use remains the long-term drift of the sensor, usually caused by fouling of the
electrode when exposed to human tissue or the gradual loss of enzyme activity. The
introduction of various membranes to act as a glucose or a hydrogen peroxide barrier has, in
general, improved sensor performance but it has not resulted in long term stability. The
much heralded membrane for this purpose, Nafion, rapidly deteriorates when implanted in the
body. Introduction of an implant into subcutaneous tissue elicits both acute and chrome
inflammatory responses. Together these result in a complexly orchestrated growth of new
tissue which ultimately envelops the implant with a foreign body capsule (FBC). In the short
term, it is likely that inflammatory cells metabolize glucose and thereby cause artifacts in the
glucose readings. When discussing the problems with long-term use of subcutaneous
sensors, experts maintain that the diminished response in vivo can be ascribed to the protein
or cellular coating around the sensor which interferes with the mass transport of glucose. If
suitable covering membranes for the sensor could be provided to exclude interfering
substances or control coating or encapsulation with proteins and cells, the excellent
performance in vitro may be matched in vivo. The use of the HA-based hydrogels described
herein as a coating agent to both minimize the FBC and keep it away from the sensor
membrane should prove a useful solution.
[0087] The purpose is to control the tissue response to an implantable glucose sensor
using a HA-coating on the sensor membrane. A sheath of HA-based hydrogel will give the
sensor membrane "breathing space" by preventing proteins and cells from clogging the
diffusion of glucose and oxygen into the sensor. Prior experience has indicated that HA and
its derivatives are extremely biocompatible and as a consequence are used in situations where

the host tissue response needs to be minimized (e.g., in eye implantation surgery). Thus,
sensor performance should be enhanced in the long term when HA-based hydrogels are cast
around sensor membranes as it relates to the development of an implantable glucose sensor
with the long term perspective that such a sensor should result in improved blood glucose
monitoring and ultimately improved quality of life for the diabetic population. In addition,
the novel cross-linking structure of the HA-based hydrogels herein disclosed will ensure
long-term maintenance of such a coating which will provide significant longevity to a
subcutaneously implanted glucose sensor.
[0088] Still another promising application is in the production of a bioartificial kidney
for the treatment of end-stage renal disease (ESRD). The only current treatment options for
ESRD patients are renal replacement therapy (all forms of dialysis) and transplantation.
Transplantation is limited by the shortage of donor organs, and is complicated by the
necessary and expensive life-long use of immunosuppressive drugs. Alternatively, although
dialysis can prolong the life of ESRD patients, average life expectancy on dialysis is reduced
by 50%, and the remaining quality of life is far from ideal. Repeated vascular access and
handling of the patient's blood leads to frequent and sometimes life threatening infections.
[0089] The functional unit of the kidney is the nephron. The nephron begins with a
filtering structure, the glomerulus, which is of a tuft of capillaries surrounded by epithelial
cells (podocytes) and supported by mesenchymal cells (the mesangium). The glomerulus is
connected directly to the tubule of the nephron, a long tube lined with a single layer
epithelium of polarized cells. The tubule cells function to salvage fluid, electrolytes and
nutrients from the filtrate (by both intracellular transport and pericellular movement)
concentrating the filtrate into urine. All nephrons connect into the collecting system, a
network of epithelial-lined tubes, which has some additional reabsorptive properties, but
primarily functions to direct the urine to the bladder. The filtration unit of the nephron, the
glomerulus, consists of the endothelial cell of the capillary arteriolar wall, the podocyte
surrounding the exterior of the capillary, and the glomerular basement membrane (GBM)
sandwiched between the two cell types. The glomerular capillaries are some of the smallest
vascular beds in the body, and the glomerular endothelial cells are specialized for their
function by being fenestrated to allow direct contact of the blood plasma to the filtration
barrier. Although these fenestrated endothelia do restrict the movement of leukocytes and
very large molecules into the filtrate, the permselectivity of the filtration barrier is defined by

the podocyte and the GBM.
[0090] The GBM is a classic basement membrane structure composed of the
prototypic molecules: type IV collagen (a3, a4, a5 heterotrimers), laminin (Laminin-11, a5,
P2, yl heterotrimers), HS proteoglycans (perlecan and agrin) and nidogen (nidogen-1 and -2);
as well as several additional ECM molecules including, collagen V, fibronectin, a CS
proteoglycan (bamacan) and several small leucine-rich proteoglycans (biglycan, decorin,
podocan). The GBM is synthesized by both the endothelial cell and the podocyte. Each cell
produces a complete basement membrane which subsequently fuses during development to
form one, double thickness basement membrane. The GBM has important functions in
providing the appropriate microenvironment and substrata for the podocytes and endothelial
cells. Without a normal GBM, both cell types lose their typical morphology and cellular
differentiation characteristics, which subsequently destroys glomerulus function. The GBM
also functions in filtration by restricting the movement of water and has some contribution to
the size and charge selectivity, however, the majority of permselectivity is dictated by the
podocyte.
[0091] The podocyte is a highly specialized epithelial cell and has unique function in
the glomerulus. The podocyte extends lamellipodia that wrap around the capillaries,
branching into very fine interdigitations with other podocytes. On cross section, these
interdigitated cellular extensions are called foot processes (FP) and the spaces between the
FPs, where filtration occurs, are called slits. The podocyte synthesizes a macromolecular
structure that spans the slit, the slit diaphragm (SD), which forms a bridge between two
adjacent FPs. The molecular composition and structure of the SD is not fully understood.
The SD appears to be a modified adherens junction containing additional podocyte-specific
proteins, the most notable being nephrin. Nephrin extends from the plasma membrane of one
FP and forms a homodimeric interaction with another nephrin molecule extending from the
adjacent FP, creating a zipper-like structure when viewed in cross section by electron
microscopy. How the SD and nephrin function as a permselective barrier is not known, but is
currently a very active area of research.
[0092] Biological microelectromechanical systems (bioMEMS) are a promising area
of exploration for development of next generation bioartificial kidneys. Drug delivery
systems, immunoisolators, and capillary networks, as well as precise control of cell
differentiation and growth have been demonstrated for bioMEMS. The kidney is the first

organ for which chronic substitutive therapy has been accepted, and application of the
bioMEMS toolkit to treatment of ESRD is both evolutionary in the technology and
revolutionary in the end product. Silicon micromachining technology has evolved such that
structures with feature sizes on the order of 1 - 100 nanometers can be reliably produced in
quantity. These dimensions are on the order of those for the glomerular slit diaphragm. The
facility with which standard silicon bulk and surface micromachining technology permits
microfluidic control, patterned deposition of cells and extracellular matrix proteins, and
immunoisolation of cells, lends itself to tissue engineering of artificial organs. The
engineering of nanoscale semiconductor filtration membranes could permit independent
control and investigation of charge-size selectivity with the potential to lead to the tissue
engineering of a bioartificial glomerulus and eventually a complete nephronal unit.
[0093] One of the first components in the miniaturization of a bioartificial kidney is
development of a nanofabricated hemofiltration membrane (NHM) from bioMEMS
components. The NHM is intended to serve the hemofiltration function of the glomerulus in
the nephron-like devices of a bioartificial kidney. NHM arrays can be fabricated using
standard silicon micromachining techniques containing slit pores of approximately the
dimensions of the glomerular slit diaphragm, and conducted experiments to demonstrate its
size barrier characteristics similar to those of the glomerular basement membrane. The
chemistries and hydrogels described in this patent application can be used to provide two
additional and necessary components to the filtration characteristics of the NHM that are
required for glomerular function. The first is a charge barrier component similar to those of
the glomerular basement membrane. This would be provided by application of a layer of
heparan sulfate (HS) based hydrogel. HS is a type of GAG similar to HA and CS. The
second addition is inclusion of the podocytes, which are responsible for the majority of the
filtration function of the glomerulus through the slit diaphragm. The podocytes would be
applied to the surface of the HS-based hydrogel layer in a HA-based hydrogel layer, which
would also serve to provide a layer of biocompatibility. The presence of the HS layer should
facilitate proper matrix-cell interactions and stimulate the deposition of an appropriate
basement membrane.
[0094] The hydrogels described herein, including but not limited to tyramine-based
hyaluronan hydrogels, also can be used as research and clinical reagents. One promising
application is controlled or extended release drug delivery. In this application, the drug can

be trapped within a sphere or other suitable shape of hydrogel material composed of a central
spherical or other shaped core of hydrogel formulated at a relatively high macromolecular
concentration (and thus lowest porosity), onto which concentric spherical layers of hydrogel
are coated, each successively coated layer being formulated of a progressively lower
macromolecular concentration (and thus higher porosity). Release of the drug is then
controlled by the rate of hydrogel degradation if so engineered, binding of the drug to the
hydrogel scaffold and diffusion of the drug through the scaffold pores. The hydrogel sphere
then is implanted into a patient at an appropriate location to effect extended release of the
drug.
[0095] Targeted drug deliver also can be achieved through an affinity-based strategy
based on designed affinity of drug laden hydrogel particles to specific tissue and cell types.
To this end, the hydrogels can be used as an affinity-based medium for the selective binding
and thus purification of specific cell populations through incorporation of targeted cell
binding molecules within the hydrogels during or prior to cross-linking. Once a select cell
population is bound to the hydrogel affinity-based medium they could be released for further
investigation, or directly entrapped while bound to the hydrogel affinity-based medium into
other formulations of the hydrogels for other tissue engineering or clinical applications.
[0096] Such an affinity-based medium also can be used for the selective binding and
purification of hyaluronan binding proteins. As the entire medium can be made solely of
hyaluronan with no other support material background binding should be quite low. By using
other materials as the scaffold material (such as aggrecan) other affinity-based media can be
prepared for purification of molecules that selectively bind to those scaffold materials.
[0097] Such an affinity-based medium also can be used for selective binding and
purification of specific macromolecules or cell populations through incorporation of protein
A within the hydrogels during cross-linking. Antibodies specific to the macromolecule or
cell population of interest can then be used to coat the protein A infused hydrogels with the
antibodies optimally oriented with their antigen binding (Fab) arms directed outward and their
constant (Fc) domain bound to the protein A. Once a select cell population or macromolecule
is bound to the protein A hydrogel, it could be released for further investigation, or directly
entrapped while bound to the protein A hydrogel into other formulations of the hydrogel for
other tissue engineering or clinical applications. Alternatively antibody could be directly
incorporated into the hydrogels.

[0098] The disclosed hyaluronan-based hydrogel materials also have utility as a
diagnostic for the presence of hyaluronidases which can be predictive of the metastatic
potential of certain cancers; e.g. by coating of a biopsy slide with hyaluronan hydrogel and
measurement of the extent and localization of the loss of intrinsic fluorescence of the
hydrogel material due to its dityramine cross-links as the hydrogel is digested by endogenous
hyaluronidases. By using other materials as the scaffold material (such as aggrecan) other
degradative enzymes could be detected such as metaloproteinases.
[0099] Further aspects of the invention will be understood in conjunction with one or
more of the following examples, which are provided by way of illustration.
EXAMPLES
EXAMPLE 1
[0100] Experimental quantities of tyramine-substituted hyaluronan hydrogels having
dityramine cross-links according to the invention have been prepared as follows. HA is
dissolved at 1 mg/ml based on hexuronic acid in 250 mM 2-(N-morpholino)ethanesulfonic
acid (MES), 150 mM NaCl, 75 mM NaOH, pH 6.5 containing a 10-fold molar excess of
tyramine relative to the molar concentration of HA carboxyl groups. Tyramine substitution
onto the carboxyl groups is then initiated by the addition of a 10-fold molar excess of EDC
relative to the molar concentration of the HA carboxyl groups. A l/10th molar ratio of N-
hydroxysuccinimide (NHS) relative to the molar amount of EDC is added to the reactions to
assist the EDC catalyzed amidation reaction by formation of active esters. Reactions are
carried out at room temperature for 24 hours, after which the macromolecular fraction is
recovered from unreacted small molecular weight reactants such as tyramine, EDC, NHS,
and MES by exhaustive dialysis versus 150 mM NaCl and then ultrapure water followed by
lyophilization. After lyophilization, the tyramine-substituted HA (T-HA) product is
dissolved to working concentrations of between 5 and 100 mg/ml in PBS (which is a buffer
compatible with cell suspension, in vivo tissue contact, and the cross-linking reaction) to
provide various concentration preparations depending on the desired rigidity of the final
hydrogel. Alternatively, the solvent can be any other suitable solvent besides PBS that will
not substantially negatively impact the enzyme activity and that will not interfere with cross-
linking reaction via selective uptake of free radicals generated by the enzyme. Suitable
alternative solvents include water, conventional biological tissue culture media, and cell
freezing solution (generally composed of about 90% blood serum and about 10% dimethyl

sulfoxide). Prior to suspension of cells (see Example 5) or contact with tissues in vivo, the T-
HA should be filtered through a 0.2 urn filter. Next, tyrarnine-tyrarnine linking is carried out
by adding 10 U/ml of type II horseradish peroxidase (HRP) to each T-HA preparation.
Cross-linking is initiated by the addition of a small volume (1-5 u.1) of a dilute hydrogen
peroxide solution (0.012%-0.00012% final concentration) to yield the final hydrogel with
desired rigidity. For preparation of larger quantities or volumes of a desired hydrogel,
quantities of reagents provided in this paragraph could be scaled up appropriately by a person
of ordinary skill in the art.
EXAMPLE 2
[0101] An experiment was conducted to determine the degree of tyramine substitution
(and consequent dityramine cross-linking) for a T-HA macromolecular network according to
the invention. Initially, three formulations of (uncrosslinked) tyramine-substituted
hyaluronan (T-HA) were prepared as described above, designated OX, IX or 10X. The OX
formulation was prepared using no EDC (i.e. containing no carbodiimide), meaning there was
no carbodiimide present to mediate the reaction for creating an amide bond between the NH2
group on tyramine and a CO2H group on the HA molecules. Thus, the OX formulation can be
considered a control. The IX formulation contained a 1:1 stoichiometric ratio of EDC based
on the quantity of CO2H groups present on the HA molecules in the reaction mixture. The
10X formulation contained a 10:1 stoichiometric ratio (or 10-fold excess) of EDC based on
the quantity of CO2H groups present on the HA molecules in the reaction mixture. In all
three formulations, a stoichiometric excess of tyramine was provided relative to the quantity
of CO2H groups on HA. In all three formulations (OX, IX and 10X) the reactants and the
appropriate amount of EDC for the formulation were combined in a vial and agitated to
facilitate the tyramine-substitution reaction. All three formulations were allowed to react for
24 hours at room temperature, after which the vial contents were dialyzed to remove
unreacted tyramine molecules, EDC and acylurea (EDU) byproducts of the reaction. These
molecules were easily separated from HA and any formed T-HA molecules through dialysis
due to the relatively small size of tyramine, EDC and EDU compared to macromolecular HA.
Once unreacted tyramine and EDC were removed, the remaining contents for each
formulation were analyzed to determine the rate of tyramine substitution relative to the total
number of available CO2H sites present on HA molecules.
[0102] Tyramine exhibits a UV absorbance peak at 275 nm, making the degree of

tyramine substitution easily detectible against a tyramine calibration curve. Based on UV-
spectroscopic analysis of the above three T-HA formulations, it was discovered that the HA-
tyramine substitution reaction carried out with no EDC present (formulation OX) resulted in
substantially zero tyramine substitution onto the HA molecules. This confirmed the
importance of using a carbodiimide reaction pathway in the tyramine substitution reaction.
However, the tyramine absorption in the T-HA formulation prepared using a 1:1 EDCrCOjH
stoichiometric ratio in the tyramine substitution reaction (formulation IX) resulted in a
tyramine substitution rate of about 1.7% relative to all available CO2H groups on the HA
chains. The 10X formulation (10:1 EDC:CO2H ratio) resulted in about a 4.7% substitution
rate.
[0103] Subsequently, hydrogen peroxide and horseradish peroxidase (HRP) were
added to each of the three dialyzed HA/T-HA formulations (OX, IX and 10X) at 5 mg/mL
and the resulting formulations were allowed to react to completion. After reaction in the
presence of peroxide and HRP, it was observed that the OX formulation remained entirely
liquid, having a strong meniscus; no gel formation was observed, confirming the fact that no
or substantially no tyramine substitution had occurred when no EDC was used in the
tyramine substitution reaction. For the IX formulation, only a very weak meniscus was
observed and the contents of the vial had gelled, confirming that both tyramine substitution
and cross-linking had occurred. For the 10X formulation, a relatively rigid gel had formed,
and in fact had shrunk relative to the initial volume of fluid in the container, leaving a
quantity of liquid (having a meniscus) on top. The gel prepared from the 10X formulation
(having a 4.7% tyramine substitution rate) was much firmer and more rigid than that from the
1X formulation having a 1.7% tyramine substitution rate.
[0104] The dityramine structure exhibits a blue fluorescence on exposure to UV light.
The products of each of the above formulations were exposed to UV light to detect the
presence of dityramine cross-links. As expected, both the IX and 10X hydrogels exhibited
blue fluorescence (the 10X hydrogel fluorescence being more intense than that of the IX
hydrogel), while the OX formulation exhibited no blue fluorescence at all. This confirmed the
presence of dityramine cross-links in both hydrogels, and that the occurrence of dityramine in
the more rigid hydrogel (10X) was greater than in the less rigid hydrogel (IX).
[0106] The overall result was that the importance of the carbodiimide-mediated
reaction pathway was demonstrated, and it was confirmed that the relative rigidity of a

hydrogel formed from a cross-linked T-HA network is proportional to the degree of
dityramine cross-linking, which is in turn proportional to the degree of tyramine-substitution -
onto HA. It was quite a surprising and unexpected result that even a 1.7% tyramine-
substitution rate (and subsequent cross-linking rate to form dityramine links) provided a
suitably firm T-HA gel (or hydrogel). A 4.7% substitution (and cross-linking) rate resulted in
even a firmer T-HA gel. Also surprising was that a ten-fold stoichiometric excess of
carbodiimide (EDC) relative to the quantity of carboxylic acid groups present in the reaction
mixture (formulation 10X) resulted in only about a 4.5-4.7% tyramine substitution rate, yet
stable and cohesive tyramine cross-linked T-HA networks were nonetheless achieved.
[0107] This means that the majority of the carboxylic acid groups on the HA
molecules are unsubstituted and not tyramine cross-linked, essentially remaining the same as
in the native HA molecule, yet the resulting network is a cohesive and stable hydrogel.
Therefore, when used as a cartilage substitute in vivo, because a majority of the HA
molecules in the invented T-HA network or gel are essentially unaltered compared to HA in
normal cartilage, it is believed that the body's native metabolic pathways (aided or unaided
by cells provided within the T-HA network) may recognize the invented network as native
biologic material, and will be able to carry out ordinary synthesis and metabolism functions
with respect thereto. In addition, it is noted that HA is a highly ubiquitous material in the
body, and is non-immunogenic in humans. As a result, the cross-linked macromolecular
network, comprised a majority of unaltered native HA, will have substantial application in a
wide variety of tissue engineering applications where it is desirable or necessary to provide
synthetic tissue in a human body. This represents a significant advance over the state of the
art. Therefore, quite surprisingly, a high degree of tyramine substitution, e.g. greater than about 10-20%, maybe undesirable; the above described experiments demonstrated that such
high degrees of substitution are unnecessary to provide a suitable T-HA network. Preferably,
a dihydroxyphenyl (e.g. dityramine) cross-linked polycarboxylate (e.g. HA) network has a
hydroxyphenyl (tyramine) substitution rate of less than 50, preferably less than 40, preferably
less than 30, preferably less than 20, preferably less than 15, preferably less than 10, preferably less than 9, preferably less than 8, preferably less than 7, preferably less than 6,
preferably less than 5, percent based on the total quantity of CO2H groups present on the
polycarboxylate (HA) molecules.

EXAMPLE 3
[0108] Conventionally, it has been believed that natural cartilage exhibits its
viscoelastic properties and its ability to resist deformation and absorb compressive loads
principally as a result of the repulsive forces between negatively charged SO42- groups on
adjacent chondroitin sulfate chains present in the aggrecan matrix. An experiment was
performed to determine the efficacy of various macromolecular networks within the scope of
the invention to resist deformation and absorb compression compared to natural cartilage. In
particular, three such networks were prepared, respectively, composed of the following: 1)
dityramine cross-linked HA molecules (T-HA); 2) dityramine cross-linked chondroitin
sulfate molecules in the form of aggrecan (T-Aggrecan); and 3) a composite material
composed of 50% T-HA and 50% T-Aggrecan. Formulations of uncross-linked T-HA and T-
Aggrecan were prepared and purified as in Example 1, each having a tyramine substitution
rate of about 5%. From these T-HA and T-Aggrecan formulations, five different
concentrations of the T-HA alone, T-Aggrecan alone and a 50:50 mixture of T-HA and T-
Aggrecan were prepared:
Concentration 1: 6.25 mg total T-GAG / mL water
Concentration 2: 12.5 mg total T-GAG / mL water
Concentration 3: 25 mg total T-GAG / mL water
Concentration 4: 50 mg total T-GAG / mL water
Concentration 5: 100 mg total T-GAG / mL water.
[0109] The notation T-GAG is used herein to embrace both T-HA and T-Aggrecan.
Though aggrecan technically is not a glycosaminoglycan (GAG), for purposes of this
example T-GAG nonetheless is defined to embrace both T-HA and T-Aggrecan hydrogels.
Each of the above preparations was then reacted in the presence of hydrogen peroxide and
horseradish peroxidase, also as in Example 1, to form dityramine cross-links between the T-
GAG molecules and provide respectively Hydrogels 1,2, 3,4 and 5 for each of the three
material compositions. Each of the fifteen hydrogels (five concentrations for each of the
three material compositions) was found to be a stable and substantially coherent material with
the physical properties of each hydrogel varying relative to the concentration of T-GAG in
the preparation from which it was made. For example, qualitatively T-HA Concentration 1
resulted in T-HA Hydrogel 1 having rigidity and rheological properties comparable to that of
Vaseline or jelly; the hydrogel was stable and coherent yet could be caused to flow or spread
on application of an external force, e.g. from a spatula or other conventional tool. T-HA

Hydrogel 1 exhibited excellent adhesive properties making it an ideal candidate for a
nonallergenic coating material for surgical instruments during surgery, e.g. ophthalmologic
surgery. T-HA Hydrogel 2 was more rigid than T-HA Hydrogel 1 due to the greater
concentration of T-HA in the preparation from which it was made, and the consequent
predicted decrease in intramolecular cross-linking and increase in intermolecular cross-
linking associated with increased T-HA concentration. T-HA Hydrogel 2 exhibited
rheological and rigidity properties characteristic of gelatins, with a degree of viscoelastic
reboundability on external loading. On greater loading, T-HA Hydrogel 2 was found to break
up into smaller pieces instead of flowing, also characteristic of a gelatinous material. T-HA
Hydrogel 3 had the properties and consistency of a dough or malleable paste, also not flowing
on application of an external loading force. This material also exhibited substantially greater
viscoelastic properties compared to T-HA Hydrogels 1 and 2. T-HA Hydrogel 4 was a highly
rigid and coherent gel that strongly resisted breaking up on application of an external loading
force. T-HA Hydrogel 4 was a highly resilient rubber-like composition that actually
generated substantial springing force upon sudden compression (e.g. dropping onto the floor).
This ability of T-HA Hydrogel 4 to generate such a springing force in response to a sudden
compression may make this material ideal for certain joint replacement/repair applications
where the joint undergoes repeated and periodic compressional loading (e.g. the ankle joint).
In addition to the properties described for T-HA Hydrogel 4, T-HA Hydrogel 5 had cartilage-
like properties with both the appearance of articular cartilage and the feel of cartilage upon
cutting with a surgical blade.
[0110] Confined compression tests were performed to quantitatively determine the
compressive mechanical properties of the fifteen different hydrogels. A custom built
polycarbonate confining chamber, and porous polypropylene filter platen (20 µm pores, 20%
porosity) were used to perform the confined compression testing. Five cylindrical plugs (7.1
mm in diameter, approximately 3 mm in thickness) at each hydrogel concentration for each
of the three material compositions were made using the confining chamber and the freeze-
thaw technique described in Example 4 below. The following testing protocol was followed
for a series of stress relaxation tests in confined compression. All testing was performed
using an Instron 5543 machine under computer control, which recorded the time-
displacement-load data at a frequency of 10 Hz. A ±5 N or ±50 N load cell (Sensotec) was
used to monitor load throughout each test. A step of 30 µm (30 µm /sec), representing 1 %
strain, was applied until the sample reached equilibrium. This was defined as a relaxation

rate that slowed to less than 10 mN min-1, at which time the next step was automatically
started, until 20 cycles (representing approximately 20% strain) were completed. The
thickness of each sample tested in confined compression was determined mechanically, by
measuring the displacement at which the compressive response initiated relative to the
bottom of the chamber as measured with the Instron 5543 machine. The measured thickness
was used to calculate the strain percentage for each step.
[0111] The compressive mechanical properties of the fifteen hydrogels were
determined as described in the preceding paragraph. Load data was normalized by sample
cross-sectional area (39.6 mm2) to compute stress. The equilibrium stress was plotted against
the applied strain for each material formulation. The aggregate modulus at each step was
defined as the equilibrium stress divided by the applied strain. For each material, the
aggregate modulus was defined as the slope of the equilibrium stress-strain data in the most
linear range. Figs. 4a, 4b and 4c display the equilibrium compression behavior for the five
concentrations of T-HA, T-Aggrecan and 50:50 T-HA/T-Aggrecan composite hydrogels,
respectively. All fifteen hydrogels were testable in confined compression, and demonstrated
characteristic stress relaxation responses typical of biphasic materials (such as cartilage). The
aggregate moduli for the 6.25 mg/ml and 12.5 mg/ml T-GAG hydrogels were 1-2 orders of
magnitude lower than articular cartilage. The 25 mg/ml T-GAG hydrogels, as well as the 50
mg/ml T-aggrecan hydrogel, displayed aggregate moduli on the order of, but at least 30%
lower than that of articular cartilage. All the 100 mg/ml T-GAG hydrogels, as well as the 50
mg/ml T-HA and the all the composite hydrogels, displayed aggregate moduli, equal to or
exceeding reported literature values for articular cartilage. These data demonstrate the ability
to characterize hydrogels using standard mechanical assays, and to generate hydrogels with
similar mechanical properties to a wide variety of tissues including that of articular cartilage,
using a variety of glycosaminoglycans as the hydrogel scaffold material.
[0112] The aggregate moduli for the five concentrations of the T-HA, T-Aggrecan
and composite materials composed of 50% T-HA and 50% T-aggrecan are summarized
below in Table 1.

[0113] Fig. 4d shows the measured aggregate modulus as a function of concentration
for the T-HA, T-Aggrecan and composite hydrogels. As the concentration of the T-HA
hydrogels increases, a plateau is reached for the aggregate modulus while the T-Aggrecan
hydrogels display a linear relationship. Interestingly, the composite hydrogels show a
relationship indicative of an exponential increase in compressive properties as concentration
increases. This indicates that the moduli of other hydrogel materials can be predicted by
further exploring and modeling these relationships.
[0114] Based on the above experiments it was surprisingly and unexpectedly
discovered that a dityramine cross-linked GAG network (HA or aggrecan) will produce a
coherent hydrogel material whose rigidity and other physical (rheological) properties can be
tuned by varying the T-GAG concentration prior to cross-linking the tyramine groups to suit
a particular application. The coherence and elastic properties of these hydrogels was
observed even absent any (or substantially any) SO42- groups in the network to supply the
charge-to-charge repulsive forces to generate the material's compression resistance and
elasticity. This was a highly surprising and unexpected result with substantial positive
consequences in tissue engineering applications. Hyaluronan is a highly ubiquitous and non-
immunogenic molecule found in humans. Therefore, hydrogels comprised of dityramine
cross-linked hyaluronan networks can be used to provide suitable tissue replacement
materials that can be implanted within a human body, whose rigidity can be tuned based on
the application as evidenced by this example. As these materials can or will be composed of
predominantly unaltered hyaluronan which is non-immunogenic, the hydrogels should result

in zero or substantially zero immune response. This is an important advantage over many
conventional tissue engineered 'materials whose formation chemistries prevent their
application in vivo due to harsh reaction conditions or reagents, and whose final chemical
structures are more likely to induce an immune response.
EXAMPLE 4
[0115] A number of methods of preparing hydrogels such as those described in
Example 3 have been developed to cast or form the hydrogel into a predetermined three-
dimensional shape. This is important for myriad tissue engineering applications where it is
necessary to provide artificial tissue material to fill a native tissue defect or void within a
patient.
[0116] A first method is to employ an in situ forming technique where the hydrogel is
formed in place, i.e. in position and in the shape of its final application and structure. The in
situ formation method has been carried out experimentally as follows. Tyramine-substituted
hyaluronan (T-HA) was prepared via the carbodiimide-mediated pathway described herein.
Following dialyzation to remove unreacted tyramine, EDC, NHS, etc., and dissolution at the
desired concentration in PBS (see Example 1 above), a small quantity of horseradish
peroxidase enzyme was added to the T-HA liquid preparation to form a first solution. This
first solution was provided into a laboratory container (to simulate an in vivo situs) having a
specific interior geometry. Subsequently, a second solution was prepared containing very
dilute hydrogen peroxide (0.012%-0.00012% final concentration). A small volume of this
second solution relative to the first solution was then injected into the container already
containing the first solution to initiate the dityramine cross-linking reaction to yield the
hydrogel. Hydrogels prepared by this technique have been prepared having varying rigidity
and rheological properties as described above in Example 3, and conformed well to the
interior surface contour of the container in which they were formed. Because the principal
reagents (H2O2, hyaluronan and peroxidase) are either nonallergenic or diffusible molecules,
and because the cross-linking reaction proceeds under metabolic conditions of temperature
and pH, this technique can be performed in vivo at a surgical situs in a patient as a surgical
procedure to produce a defect-conforming hydrogel. This method is particularly attractive
for reconstructive facial surgery in which the uncross-linked T-HA preparation (with
peroxidase) can be injected and manipulated subcutaneously by the surgeon to produce the
desired facial contours and then the hydrogel subsequently cross-linked by injection of a

it freezes into^a solid ice form conforming to the shape and contour of the inner mold surface.
However, the silicone mold, having a glass transition temperature below -80°C, remains soft
and malleable and the solid ice form of the first solution is easily removed. Because the first
solution expands as it freezes, suitable mechanical hardware should be used to ensure the
silicone mold does not deform or expand as the solution freezes. Preferably, port holes are
provided in the mold to allow for expansion and discharge of the first solution as it expands
during the freezing process.
[0119] Once the solid ice form of the first solution has been demolded, minute defects
or flaws in the three-dimensional structure can be repaired by carving with a suitable tool,
and more of the liquid first solution can be added to fill surface voids, which liquid instantly
freezes on contact with the solid ice form. Also, the ice form can be placed back on the dry
ice surface if desired to ensure uniform temperature and freezing of any added first solution
material. Once the three-dimensional shape of the ice form has been perfected, it is
immersed in a liquid peroxide solution to initiate thawing of the frozen water and dityramine
cross-linking from the outside-in. This is possible do to the rapid kinetics of the cross-linking
reaction. Cross-linking is determined to be complete once the last remaining frozen water has
melted at the center of the forming hydrogel form, which can be easily observed because the
forming hydrogel is substantially clear.
[0120] Very successful experiments have been performed according to this freeze-
thaw technique to produce a solid hydrogel in the shape of a human ear. Other structures that
could be formed by this method, such as intervertebral discs, meniscus, etc. will be evident to
those skilled in the art. It should be noted in this freeze-thaw technique, the threshold glass
transition temperature of-80°C for the mold material is selected to correspond roughly with
the surface temperature of solid CO2 (dry ice), to ensure the mold material does not become
brittle when the first solution is frozen to produce the solid ice form. However, if another
cooling material, other than CO2 is used, then the threshold glass transition temperature for
suitable mold materials may be adjusted accordingly.
[0121] For the three methods of hydrogel formation described above, the first solution
contained both the peroxidase and T-HA, while the second solution contained the peroxide.
While it may be possible to switch the peroxidase and peroxide in the first and second
solutions respectively, it is less preferred to provide the peroxide in the first solution with the
T-HA. This is because once the peroxide, peroxidase and T-HA are combined, the T-HA

rapidly begins to form a cross-linked macromolecular network. If the peroxidase (which is a
macromolecular molecule) is not already uniformly distributed with the T-HA it may be
unable or substantially hindered from diffusing through the pore structure of the forming
hydrogel to facilitate uniform cross-linking throughout the entire T-HA/peroxide solution.
The result could be non-uniform and/or incomplete cross-linking of the T-HA and a non-
uniform hydrogel. Conversely, the relatively small peroxide molecule (hydrogen peroxide is
only one oxygen atom larger than water) can diffuse through the hydrogel pore structure with
relative ease, resulting in a uniform hydrogel structure.
[0122] In addition, the macromolecular size of the peroxidase allows it to be similarly
retained as the T-HA within porous molds that are only porous to small molecular weight
peroxides which easily and uniformly diffuse through both the molds and newly forming
macromolecular networks (i.e. hydrogels). For these reasons it is preferred to start with the
peroxidase uniformly distributed with the T-HA in the first solution, and to provide the
peroxide separately in the second solution.
[0123] A fourth method is an alternating sprayed or brushed layering technique. The
first solution is prepared as described above and contains both the peroxidase and T-HA.
However, the second solution not only contains the peroxide as described above, but also T-
HA at the same concentration as in the first solution. Then a thin layer of the first solution is
applied at the desired location (in situ) followed by an overlying thin layer of the second
solution. This procedure is repeated such that alternating layers of the first and second
solutions are successively applied until the defect or application situs has been completed.
The very thin alternating layers of the first and second solutions promote virtually complete
dityramine cross-linking ensuring a highly coherent final hydrogel having the desired
rheological properties based on the initial T-HA concentration of the two solutions. The thin
nature of the layers is desirable to ensure that free radicals produced by the peroxidase in the
first solution layers are able to penetrate completely adjacent second solution layers and
complete cross-linking independent of peroxidase diffusion into the second solution layer
(see above). T-HA is included in both solutions to ensure uniform T-HA concentration
throughout the final hydrogel. This technique has been performed in laboratory bench
experiments and has provided contour-conforming and volume-filling coherent hydrogels.
This technique is highly applicable where it is desired to provide a thin, but variable layer of
tyramine cross-linked HA, such as on the surface of a denuded osteoarthritic joint in which

little if any native healthy cartilage remains in the patient at the implant site.
[0124] All four of the above techniques have been described with respect to
dityramine cross-linked hyaluronan, however it will be understood that other combinations
within the scope of the present invention (other dihydroxyphenyl cross-linked
macromolecules, such as polycarboxylates, polyamines, polyhydroxyphenyl molecules and
copolymers thereof) can be molded via the above techniques.
EXAMPLE 5
[0125] Rat chondrocytes were embedded in (cross-linked) T-HA hydrogels to
measure their ability to survive the cross-linking reaction. Isolated chondrocytes were
suspended in the 1.7% and 4.7% T-HA hydrogels described in Example 2 by providing these
live cells to the first solution to be co-dispersed with the T-HA and peroxidase, followed by
introduction of the peroxide-containing second solution to initiate dityramine cross-linking.
The chondrocyte-embedded 1.7% and 4.7% T-HA hydrogels exhibited uniformly distributed
chondrocytes with the optical clarity of the gels allowing visualization throughout the gel.
Glucose utilization was used as an indicator of cell viability after cross-linking to form the
hydrogels as chondrocytes are voracious with respect to glucose consumption, depleting the
medium of glucose in less than 24 hours. The results showed that chondrocytes embedded in
T-HA hydrogels showed essentially the same glucose consumption profile over 24 hours as
the same chondrocytes cultured in monolayer (Fig. 5). This continued for up to 7 days
indicating that the cells were alive and metabolically active. Medium glucose was measured
by standard hexokinase assay.
[0126] Fluorescent images of frozen sections of T-HA hydrogels containing both
chondrocytes and cartilage tissue were also generated. HA samples from both the hydrogel
scaffold and cartilage matrix were visualized by fluorescent staining with biotinylated HA
binding protein (b-HABP) reagent while cell nuclei were visualized with standard DAPI
stain. The b-HABP reagent is prepared from purified cartilage aggrecan (the G1 domain
only) and link protein, and recognizes and irreversibly binds to stretches of native HA
equivalent to those normally bound by aggrecan and link protein in cartilage. The results
showed a more intense staining of the T-HA hydrogel with b-HABP than the cartilage as the
hyaluronan in the tissue is already occupied by native aggrecan and link protein. No visible
distinction could be seen between the T-HA scaffold of the hydrogel and the matrix of
suspended cartilage tissue suggesting seamless integration. These results demonstrated the

feasibility of maintaining the viability of chondrocytes during the hydrogel cross-linking
reactions, and the ability of the hydrogel to integrate seamlessly into existing cartilage matrix,
both of which are advantageous for application to cartilage repair. The results also
demonstrated that sufficient stretches of the T-HA remain chemically unaltered, and available
for binding by newly synthesized aggrecan and link protein in situ. The results also
demonstrated that oxygen, carbon dioxide, glucose and insulin are diffusable through T-HA
hydrogels according to the invention at a rate that is not limiting to chondrocyte metabolism,
which is important not only to the development of cartilage substitutes but to other
applications such as glucose sensor design and development of an artificial kidney.
[0127] In order to include cells such as chondrocytes in hydrogels molded into
intricate anatomical shapes using the freeze/thaw technique described in Example 4, it is
desirable that the enzyme driven cross-linking reaction proceed in the presence of standard
cell freezing solutions such as those containing 10% dimethylsulfoxide (DMSO)/90% fetal
bovine serum (FBS). This has been demonstrated in the laboratory for all of the T-HA
hydrogel formulations described in Example 3. The ability to directly incorporate a solution
containing 90% FBS also demonstrates the ability to include bioactive factors such as growth
factors, hormones and factors controlling cell differentiation, as these are normal components
ofFBS.
EXAMPLE 6
[0128] An experiment was conducted whereby a T-HA hydrogel as described
hereinabove was implanted into Yucatan minipigs in order to repair articular cartilage
defects. Following is a description of that experiment, including the experimental methods
and results obtained, after a brief discussion of the background for this application.
Background
Tissue Description
[0129] Articular Cartilage Structure and Function - As discussed above, articular
cartilage is the resilient load-bearing tissue that forms the articulating surfaces of diarthrodial
joints. It absorbs mechanical shock and deflects or spreads applied load over greater surface
area of subchondral bone. It consists primarily of a large extracellular matrix (ECM) with a
sparse population of highly specialized cells (chondrocytes) distributed throughout the tissue.
The primary components of the ECM are water, cartilage aggregates and type II collagen.

Cartilage aggregates are composed of hyaluronan (HA), aggrecan (the large cartilage-specific
proteoglycan), and link protein (LP), a small glycoprotein. Aggrecan contains a central core
protein to which is attached ~ 100 chondroitin sulfate (CS) chains. The core protein has three
globular domains with the N-terminal globular 1 (G1) domain having binding sites for both
HA and LP. LP has sequence homology to the G1 domain of aggrecan, and contains binding
sites for both HA and the G1 domain of aggrecan. Each cartilage aggregate is composed of a
single HA chain, to which are attached hundreds of aggrecan/LP duplexes. These large
cartilage aggregates are trapped at one fifth of their free solution volume within a tight
meshwork of type II collagen fibers, which resist further swelling. This molecular
architecture contributes to the tissues mechanical properties and function as described below.
[0130] Swelling Pressure - The HA and CS chains in cartilage aggregates contain
repeating carboxyl and/or sulfate groups. In solution, these groups become ionized (COO'
and SO3"), and in the physiologic environment they require positive counter ions such as Na+
to maintain overall electroneutrality. These free-floating ions within the interstitial water are
present at a higher concentration than that found in the surrounding fluids (i.e. synovial fluid)
giving rise to an osmotic pressure (Donnan pressure). In cartilage, ions are prevented from
flowing out of the tissue along the concentration gradient by the fixed nature of their negative
counter ions (i.e. the COO" and SO3" groups on the HA and/or CS chains), and the need to
maintain electroneutrality. Water flow into the tissue to equilibrate the concentration
gradient is resisted by the inextensible nature of the collagen meshwork preventing further
swelling.
[0131] Alternatively, tight cartilage aggregate packing causes the fixed-negative
charge groups to be spaced only 10 to 15 angstroms apart, resulting in strong charge-to-
charge repulsive forces (electrorepulsive forces). As with the Donnan effect, the tendency to
swell to lessen these repulsive forces is resisted by the inextensible nature of the collagen
meshwork. When compressed, the distances between charge groups decrease, thus increasing
the charge-to-charge repulsive forces and increasing the free-floating positive counter ions
concentration. Thus both the Donnan and electrorepulsion effects are intensified by
compression. Both effects contribute to the swelling pressure of articular cartilage and its
ability to resist deformation and absorb compressive loads.

[0132] Stress Shielding Effect - Articular cartilage is often described as a viscoelastic,
biphasic material, composed of a solid phase (cartilage aggregates, collagen, etc.) and a fluid
phase (water and dissolved ions). The macromolecular architecture of the ECM of articular
cartilage functions to deflect applied forces during loading from the wear susceptible solid
phase of the tissue to the wear resistant fluid phase or water. This stress shielding occurs due
to the elegant design of the cartilage ECM which produces a material with very low
permeability creating a drag during interstitial fluid flow. Interstitial fluid pressure is
generated during compressive loading, and during dynamic loading, is the primary force
responsible for supporting the applied load with matrix compression a minor factor. During
compression, the porosity is reduced further, which increases the already high frictional drag
forces. The load support is gradually transferred from the fluid phase (as the fluid pressure
dissipates) to the solid phase. Typically, for normal cartilage, this equilibration process takes
2.5 to 6.0 hours to achieve. Thus, load support through fluid pressurization predominates
within the tissue.
Need for Synthetic Material
[0133] Increased water content and decreased proteoglycan content are the most
apparent early changes in osteoarthritic cartilage. These changes reflect an increase in tissue
permeability. Increased permeability diminishes the fluid pressurization mechanism of load
support in cartilage (stress shielding), requiring the collagen-aggrecan solid matrix to bear
more load, which may be an important contributing factor in the development and
progression of cartilage degeneration. Bioartificial cartilage substitutes that do not mimic the
low permeability of normal, healthy articular cartilage may be predisposed to degeneration by
a similar mechanism.
[0134] One of the most difficult challenges facing orthopedic surgeons is treating
patients who have suffered focal cartilage lesions, but are too young or too active for a total
joint replacement. These localized cartilage defects can be very debilitating. Restoring these
localized areas without a total joint replacement would be a preferred approach with
significant benefits including reduced surgical requirements, shorter recovery times, lower
cost, and slowing or arresting the further degradation of the load bearing surface.
[0135] This example demonstrates the application of a tyramine-substituted HA (T-
HA) hydrogel for repair of this type of localized cartilage defect.

Experimental Description
Design of Extracellular Matrix Material having Desired Properties
[0136] Natural articular cartilage has the elastic as well as physical and chemical
properties described above, which impart its unique ability to absorb mechanical loads and to
deflect impact loads away from the subchondral bone. To produce a suitable synthetic
cartilage material made from a T-HA hydrogel as disclosed herein, it was important to design
the macromolecular network for the hydrogel so as to emulate those properties as nearly as
possible through judicious selection of reagent concentrations, cross-linking conditions,
incorporated living cells as well as other molecules, etc.
[0137] The results from confined compression testing of T-HA hydrogels (see
Example 3 above) provided an initial basis for the synthesis of an appropriate synthetic T-HA
material having properties matched to those measured for normal articular cartilage. They
also illustrate the spectrum of material properties that can be manufactured from a single
formulation of T-HA. Based on those data, a T-HA hydrogel composed of, inter alia a
macromolecular network of dityramine cross-linked hyaluronan molecules was selected
based on the following criteria in order to produce a synthetic implantable cartilage material.
[0138] A material composed solely of HA was chosen because hydrogel compositions
with the compressive properties of cartilage could be formed using HA alone as the scaffold
material (Example 3), while avoiding possible host response to the protein component of
aggrecan if it were used as a scaffold material. Reaction conditions (tyramine/EDC ratio)
were chosen to produce a percent tyramine substitution of 5% (Example 2) as this provided
sufficient cross-linking to produce a material with the compressive properties of cartilage
(Example 3) while maintaining the majority of the native HA structure. HA was substituted
with ~5% tyramine, as described in Example 1, except that the HA was dissolved at 5 mg/ml
rather than 1 mg/ml to conserve reagents. The absolute concentration of all other reagents
remained the same so that the tyramine and EDC were at a 2-fold rather than 10-fold molar
excess based on the molar concentration of HA carboxyl groups. A concentration of 125
mg/ml of HA in sterile saline was chosen as this concentration in saline had a compressive
aggregate modulus most closely resembling that of articular cartilage (saline data not shown).
It was also the concentration deemed most appropriate based on the experience of our
clinician collaborator. Peroxidase was added at 10 U/ml prior to application in an in situ

cross-linking protocol as described below. The in situ cross-linking protocol was chosen as it
provided the best opportunity for integration with surrounding cartilage matrix. It also
allowed easy and complete filling of the surgically produced cartilage defect without the need
to know or measure the defect's exact dimensions. Pre-cast (in vitro cross-linked) plugs
would either require exact dimensions or sculpting of pre-formed shapes to fit the defect. No
cells or bioactive factors where added in this experiment as this experiment was intended to
evaluate the hydrogel material independent of complicating factors derived from inclusion of
cells or bioactive factors. However, cells or bioactive factors could be included as described
hereinabove to produce desired effects.
Surgical Procedure
[0139] Pre-Operative - After arrival in the biological resource unit, Yucatan
minipigs (~7-8 months of age, -30-35 kg) were maintained for a minimum of 7 days to
ensure full acclimatization. After pre-medication with Ketamine (20 mg/kg I.M.) as
anesthesia and Ambipen (40,000 U/kg I.M.) as prophylaxis antibiotic, the animal's rear legs
were shaved and painted in Betadyne as simultaneous bilateral surgery of both knees was
performed. A general anesthesia was maintained by inhalation with Isoflurane (1-2.5%
volume) in O2 following intubation. Thiopental was used as needed (to effect 25 mg/ml
I. V.). During surgery, the animal was monitored for heart rate, respiration rate, body
temperature, etc.
[0140] Opening - A longitudinal midline skin incision was made and carried down
sharply through the pre-patellar bursa. Electrocautery was used for hemostasis. The lateral
border of the patella was identified and a lateral para-patellar arthrotomy was performed.
The lateral retinaculum and musculature were tagged with #1 vicryl suture. The patella was
dislocated medially to expose the femoral trochlea.
[0141] Cartilage repair - As seen in Fig. 6, two circular full thickness chondral
defects (-4.5 mm in diameter, panel B of Fig. 6) were created in the medial trochlear facet of
the femoral chondyle (panel A of Fig. 6) using an Acufex 4.5 mm mosaiaplasty chisel and a
sharp, curved currette taking care as much as possible not to disrupt the osteochondral plate.
The defects were filled with in situ cross-linked T-HA hydrogel (125 mg/ml in sterile saline)
as follows to produce a hydrogel implant having the composition described above in order to
reproduce the in vitro measured compressive properties of natural cartilage as described

above. Initially, each defect was rinsed with 0.01 cc of 0.6% hydrogen peroxide, and then
immediately blotted dry with sterile gauze. Subsequently, a plug containing 0.15 cc of
uncross-linked hydrogel paste (panel C of Fig. 6) having the composition and prepared as
described above was inserted into and used to fill each defect with the surgeon smoothing the
surface of the hydrogel implant with fingertips to match the contour of the articular surface.
A sterile piece of filter paper (Whatman 50) soaked in 0.6% hydrogen peroxide was pressed
against the surface of the hydrogel implants for five minutes to cross-link the hydrogel in the
defect. During the 5 minutes, the filter paper was rubbed back and forth across the implant
surface to prevent integration with the filter paper, and to effectively polish the implant
surface. After 5 minutes the filter paper was removed, excess hydrogel trimmed from the
site, and then -0.01 cc (1 drop) of 0.6% hydrogen peroxide added to the surface of each
implant plug (panel D of Fig. 6). The Patella was reduced anatomically over the femoral
trochlea. The Patella was dislocated and reduced again to ensure secure primary stability of
the hydrogel.
[0142] Closing - The joint was irrigated with sterile saline. The wounds were closed
in layers with vicryl sutures. Specifically, the arthrotomy was closed with interrupted #1
vicryl suture, the subcutaneous tissue was closed with interrupted 2-0 vicryl suture and the
skin layers were closed with interrupted 3-0 vicryl suture. No restriction of movement was
required after surgery.
[0143] Post-Operative - The animal returned to full weight bearing immediately
following surgery. Analgesia was provided by Buprenorphine (0.02 mg/kg I.M.) for 24 hours
and a Fentanyl Patch (50 mcg/hr) for 3 post-operative days. Post-operative prophylactic
antibiotic in the form of cephalexin 500 mg twice per day was given for 7 days. The animals
were kept in a conventional animal run.
[0144] Post-Implantation Data - At one month post-implantation, the animal was
euthanized with overdose of the barbiturate, Beuthanasia D Special (1 ml/10 kg B.W., I.V.)
under general anesthesia. After euthanasia, the entire knee joint was carefully dissected,
macroscopically evaluated and photo documented. As seen in Fig. 7, macroscopic inspection
of the knees at 1 month revealed no significant effusion and no evidence of inflammatory
reaction. The lesions were partially filled with a white material (the implanted T-HA
hydrogel as well as other factors or cells which may have migrated into the hydrogel post-

operatively) and the surrounding articular cartilage and opposing articular surface (patella)
were normal in appearance except for a slight abrasion appearing on the opposing articular
surface as seen in panel B of Fig. 7. It is not evident this abrasion was the result of rubbing
against the implants, particularly given its location on the Patella in a position which does not
appear as though it would have abraded against the implants during normal articulation of the
joint.
[0145] The results indicate no apparent negative effect on joint health as a result of
the hydrogen peroxide or peroxidase reaction used for in situ cross-linking of the hydrogel,
and demonstrated the utility of the hydrogels disclosed herein, comprising a dityramine cross-
linked hyaluronan macromolecular network, as a synthetic implantable extracellular matrix
for use as a synthetic in vivo cartilage replacement or implant material.
EXAMPLE 7
[0146] An experiment was conducted whereby a T-HA hydrogel as described
hereinabove was implanted into canine and rabbit models in order to repair vocal cord defects
as well as to augment vocal cords. Following is a description of that experiment, including
the experimental methods and results obtained, after, a brief discussion of the background for
this application.
Background
Tissue Description
[0147] The vocal cords are complex, multilayered structures under very fine
neuromuscular control. The overlying mucosa is composed of a non-keratinized, stratified
squamous epithelium, with a multilayered, lamina propria deep to the epithelium.
Underlying the lamina propria is a muscular layer consisting of the thyroarytenoid muscle
which inserts into the thyroid cartilage anteriorly and the vocal process of the arytenoid
cartilage posteriorly. The thyroarytenoid muscle can stiffen or relax, altering the tension on
the lamina propria and thereby altering the vibratory dynamics of the epithelium, which
produces the finely coordinated vibrations responsible for high quality speech production.
[0148] The biomechanics of human voice production have been attributed to the
action of certain biological macromolecules naturally found within the extracellular matrix
(ECM) of the lamina propria. Hyaluronan (HA) is a ubiquitous molecule, which is most

concentrated in specialized tissues such as the vocal cords, synovial fluid, umbilical cord,
dermis, and cartilage. In these tissues, its function is manifold, influencing tissue viscosity,
shock absorption, wound healing, and space filling.
[0149] The unique structure of HA elucidates its multiple functions. It consists of D-
glucuronic acid and N-acety]glucosamine arranged in repeating disaccharide chains,
containing as many as 30,000 repeating disaccharide units with a mass of more than 10
megadaltons. As HA is a polysaccharide instead of a protein, it is non-antigenie. Under
biological conditions, it is a negatively charged, randomly coiled polymer, filling a volume
more than 1,000 times greater than expected based on molecular weight and composition
alone. The strong negative charges attract cations and water, allowing it to assume the form
of a strongly hydrated gel, and giving HA its unique viscoelastic and shock-absorbing
property.
[0150] Vocal cord viscoelasticity is essential to high quality voice production, as it
directly affects the initiation and maintenance of phonation and the regulation of vocal cord
fundamental frequency. HA in the human glottis is concentrated in the lamina propria and its
importance has been quantified by comparing the biomechanical properties of cadaveric
vocal cords with and without HA. Treatment of the vocal cords with hyaluronidase led to a
35% average reduction in vocal cord stiffness and a 70% mean reduction in high frequency
vocal fold viscosity, thus illustrating the significance of HA in these tissues.
Need for a Synthetic Material
[0151] Vocal Cord Repair - Defects in the vocal cords have a dramatic effect on vocal
production. Arising either de novo, or resulting from surgical intervention, heterogeneous
masses within the vocal cords disrupt the finely coordinated vibrations responsible for high
quality speech production. Patients with de novo lesions usually present early in the disease
course, due to persistent hoarseness. When presenting with an early stage malignant process
(T1-T2 in the Tumor, Node, Metastasis staging system), patients undergo rapid treatment
consisting of either external beam radiation therapy or endoscopic surgery. Such patients are
counseled that poor post-treatment voice quality is an expected side effect of effective tumor
eradication. When presenting with a presumed benign process, patients are faced with a
conundrum, for the surgical treatment often produces speech quality as poor as that caused by
the lesion itself. Unfortunately, current standard laryngeal operative technique cannot
provide for effective removal of either benign or malignant lesions without causing poor

vocal outcomes secondary to vocal fold scarring. This is due to the mechanism of wound
healing in the unique anatomy of the larynx. The superficial, vibratory surfaces of the vocal
cords become tethered to the deeper layers by the post-treatment scar, preventing physiologic
phonatory oscillation.
[0152] HA in the human glottis is concentrated in the lamina propria, a histological
layer separating the vocalis muscle from the overlying epithelium. The lamina propria allows
the epithelium to vibrate over the taut vocalis muscle, like waves propagating over a pond.
This "mucosal wave" is the sine qua non of effective speech production. In the presence of
benign or malignant lesions of the vocal cords, the mucosal wave is disrupted. Even in the
normal process of healing, scar bands and disorganized collagen "tether" the superficial
mucosa to the deeper layers of the vocal cords, disrupting the normal mucosal wave and
impairing vocal production. The shock absorbing nature of HA allows it to act as a tissue
damper, protecting the mucosal surfaces from the oscillatory trauma experienced during
phonation. HA also appears to facilitate wound repair by minimizing fibrosis and scarring,
thereby protecting the vocal cord from the permanent damage resulting from trauma.
[0153] The development of a technique, permitting the restoration of a fully vibratory
phonatory surface on vocal cords undergoing laser or cold surgical treatment, would enable a
large population of patients with both benign and malignant processes to undergo treatment
of their tumors with the expectation of unprecedented post-operative speech outcomes.
[0154] Vocal Cord Augmentation - A variety of disorders and diseases adversely
affect glottic function, vocal quality and the ability to communicate. Approximately 7
million people in the United States suffer from dysphonia or voice impairment, and those
affected by vocal cord paresis/paralysis are a significant subset of this population. It is
estimated that 1-4% of all cardiac and thyroid surgeries in the United States result in vocal
cord paresis or complete paralysis due to inadvertent vagus nerve or recurrent laryngeal nerve
injury during surgery.
[0155] Another condition affecting vocal cord function is unilateral vocal fold
paralysis (UVP). In UVP the problem is malposition of an insensate vocal cord. While
medialization results immediately following nerve injury due to opposing tensions of the
laryngeal adductors and abductors, the paralyzed vocal cord rapidly lateralizes to a
paramedian position. The arytenoid cartilage prolapses into the larynx following recurrent
laryngeal nerve injury, resulting in a change in vertical height of the vocal cord, as well as

decreased dynamic tension often resulting in vocal fold bowing. Atrophy with resultant
shortening and bowing of the vocal cord occurs later as the thyroarytenoid muscle atrophies
due to a lack of neural stimulation. As a result of atrophy and lateralization, the contralateral
vocal cord cannot fully contact the paralyzed cord, leading to manifestations of UVP.
[0156] Such manifestations include breathy hoarseness, a weak cough, an inability to
valsalva (protect the airway), and difficulty swallowing; complications include aspiration
(solids and liquids) and recurrent pneumonia. This can result in a life-threatening condition
due to increased incidence of recurrent pulmonary infections.
[0157] As only one functional vocal cord is required for normal voice production,
successful treatment consists of "medialization" of the paralyzed vocal fold, thereby enabling
it to contact the contralateral mobile vocal fold. This normalizes voice production and
prevents aspiration minimizing the risk of aspiration pneumonia. Vocal fold paralysis is
currently treated in two ways: open trans-cervical approaches or trans-oral endoscopic vocal
cord injection, also known as injection laryngoplasty therapy (ILT). Ishiki-type I thyroplasty
is the most commonly performed trans-cervical approach, where a "window" is created in the
thyroid cartilage to allow placement of a silastic implant into the body of the atrophic,
paralyzed vocal fold, in effect pushing it into a more medial position. This procedure has a
permanent effect, although complications include implant migration, extrusion, or infection.
[0158] In ILT, the paralyzed vocal cord is medialized by the endoscopic injection of
an exogenous substance. A wide variety of synthetic and biologic materials are currently
available as an injectant for treatment of UVP, including: gelfoam, hydroxyapatite,
autologous fat or facia, acellular cadaveric dermis (Cymetra®), collagen or
Teflon®/Gortex®. Unfortunately, all have proven to be less than ideal with none fulfilling
all desired criteria for the ideal material for long-term vocal cord augmentation. Such
limitations create the need for either re-injection or over-injection to account for the projected
loss of volume.
[0159] A biocompatible, injectable material such as a T-HA hydrogel as disclosed
herein can be designed to mimic the Theological properties of the natural vocal cord tissue
and persist indefinitely in vivo without migration. The design, chemistry, and material
properties of a T-HA hydrogel as described herein can be tuned to produce an injectable

bioimplant that is uniquely suited to otolaryngology treatments such as ILT through judicious
selection of component concentrations and cross-linking methodologies as described
hereinabove.
[0160] A suitable biocompatible, high-longevity synthetic material also is desirable to
treat pre-existing sulcus or scarring that can develop due to trauma or develop spontaneously
with aging. Such a material also could be used advantageously in place of saline as a
diagnostic and surgical aid prior to vocal cord surgery. Conventionally, saline is injected
within the HA matrix of the vocal cord lamina propria between a lesion to be surgically
removed and the underlying ligament. This is done to: a) determine if t: lesion involves the
underlying ligament; and b) to make the surgery easier by increasing the distance between the
lesion and the ligament (cold instruments), or providing a heat sink (laser). Ligament
involvement complicates the surgery with penetration of the ligament to be avoided if
possible. This procedure could benefit from the incorporation of hepatocyte growth factors in
the hydrogel, which is used to increase HA production in the lamina propria J decrease
collagen production associated with scarring.
Experimental Description
Design of Extracellular Matrix Material having Desired Properties
[0161] An ideal synthetic matrix or biomaterial for vocal cord augmentation will have
the following characteristics: 1) biocompatible, so there is no unfavorable immunologic
response; 2) easily injectable to allow a surgeon to control the exact amount and location of
injection through a small needle; 3) readily available with minimal preparation for optimal
time efficiency arid potential application to the outpatient office setting; 4) possess the same
or similar biomechanical properties to the vocal fold component being augmented to cause
minimal alteration in the natural function of the augmented structure; 5) resistant to
resorption or migration, so that the initial augmentation result is maintained; and 6) easily
removable in the event of revision surgery.
[0162] The T-HA hydrogels disclosed herein meet all six of these criteria, most
important being that its biomechanical properties can be tuned through judicious selection of
reactant/synthesis parameters and GAG (e.g. HA) concentration to produce the necessary
macromolecular network for producing a hydrogel having desired viscoelastic and
biomechanical properties. Specifically, results of both in vitro and in vivo preliminary studies

have allowed favorable comparisons to be drawn against the above criteria. First,
transplantation of T-HA hydrogels into various animal species including rat, rabbit, dog and
pig all have demonstrated little to no host immune response. Second, uncross-linked T-HA
hydrogels at the concentrations required for vocal cord augmentation at the lamina propria or
muscle level easily pass through a 21 gauge needle. Third, medical-grade HA is readily
available and has been used for years in FDA-approved formulations such as Healon and
Restylane. Furthermore, uncross-linked T-HA hydrogel and hydrogen peroxide (cross-
linking agent) solutions can be readily pre-made, in off-the-shelf formulations that require no
preparation by the surgeon. Fourth, the T-HA hydrogels disclosed herein can be formulated
to match the mechanical properties of the various tissues of the glottis including the lamina
propria, thyroarytenoid muscle, and thyroid or arytenoid cartilages. Fifth, the unique cross-
links and non-protein nature of the T-HA hydrogels have demonstrated resistance to
resorption in in vivo experiments. This implies that the initial augmentation result will be
maintained. Finally, formation of a solid continuous implant through the novel in situ cross-
linking protocol made possible through the non-immunogenic enzyme-driven cross-linking
architecture described herein should prevent migration and allow for easy location and
removal of the hydrogel implant should revision surgery be required.
[0163] With respect to the fourth point above, the results from confined compression
testing of T-HA hydrogels (see Example 3 above) provided an initial basis for the synthesis
of an appropriate synthetic T-HA material having properties matched to those of normal
vocal cord tissue. Based on those data, a T-HA hydrogel composed of, inter alia a
macromolecular network of dityramine cross-linked hyaluronan molecules was selected
based on the following criteria in order to produce a synthetic implantable vocal cord repair
or augmentation material.
[0164] Choices of scaffold material (HA alone), percent tyramine substitution (-5%),
protocol for tyramine substitution of HA (modified from Example 1), and the non-
incorporation of cells and bioactive factors were as described in Example 6. A concentration
range of between 2.5 and 10 mg/ml of T-HA hydrogel in sterile saline most closely matched
the Theological and vibratory properties of the vocal cord lamina propria. The 2.5 mg/ml
concentration of T-HA hydrogel was deemed most appropriate based on the extensive clinical
experience of our clinician collaborators with both the vocal cord tissue and other injectable
materials used for vocal cord repair and augmentation. In vitro cross-linked hydrogel was

used rather than an in situ cross-linkm? protocol based on the experience of our clinician
collaborators with other injectable materials for vocal cord repair and augmentation. In vitro
cross-linking was as described in Example 1. Based on the results of the rabbit and canine
experiments described below a preferred embodiment for vocal cord augmentation using the
disclosed hydrogel material as envisioned by the inventors would use the in situ cross-linking
protocol and a T-HA concentration to more closely match injection into the deeper muscle
layers of the vocal cord.
Surgical Procedure
[0165] Pre-Operative — After arrival in the biological resource unit, mongrel dogs or
New Zealand white rabbits (depending on the experiment, described below) were maintained
for 7 days for full acclimatization. After pre-medication and general anesthesia per IACUC-
approved protocols, each animal was intubated and maintained in stage III surgical
anesthesia. The animal was positioned supine. After grasping and superiorly retracting the
tongue, a Dedo laryngoscope was placed transorally providing good exposure of the larynx.
The tip of the laryngoscope was positioned several centimeters proximal to the superior
surface of the true vocal cords. The laryngoscope was then suspended. A rigid
videostroboscopic telescope was positioned above the true vocal cords, permitting complete
inspection and imaging of the larynx.
[0166] Vocal Cord Repair -For vocal cord repair, lateral-based micro flaps were
raised in both vocal cords of dogs, and then soft tissue defects were created equivalent to
50% of the vocal cord mass, including lamina propria and underlying muscle. One side
underwent soft tissue reconstruction (filling) with the T-HA hydrogel, while the contralateral
side served as an unrepaired (unfilled) control. The microflap was then redraped over the
hydrogel such that the epithelium was completely continuous. This study used 2.5 mg/ml ex
vivo cross-linked T-HA hydrogel in saline (5% tyramine substituted) prepared as described
above to approximate the rheological properties of the lamina propria. After surgery, the
dogs were weaned from the anesthesia and transferred to the recovery room. The animals
received analgesia for 1 to 2 days per IACUC-approved protocols.
[0167] Injection Laryngoplasty Therapy (Vocal Cord Augmentation) - After
anesthesia, a 27 gauge laryngeal needle was used to inject approximately 0.25 ml of 2.5
mg/ml ex vivo cross-linked T-HA hydrogel in saline (5% tyramine substituted) at the left

anterior and posterior membranous vocal cord of rabbits. The injections were made in the
superficial layer of the lamina propria.
[0168] Based on the results from the above experiments with dogs (vocal cord repair)
and rabbits (ILT), the following preferred embodiment for vocal cord augmentation using the
disclosed hydrogel material is envisioned by the inventors. For ILT, the lateralized and
atrophied vocal cord are injected with 50 mg/ml of the uncross-linked T-HA hydrogel in
saline (5% tyramine substituted) with peroxidase at the level of the thyroarytenoid muscle.
Preferably one bolus of hydrogel would be used to obtain the desired medialization, but no
more than two. A 21 gauge laryngeal needle is used to inject the hydrogel. Cross-linking
would be initiated by injection of a small volume of dilute hydrogen peroxide through a 27
gauge needle into the center of the implanted bolus of hydrogel using the 21 gauge needle,
which would not have been withdrawn, for purpose of orientation. Cross-linking of the
hydrogel into a solid implant would be achieved within minutes and verified by feel. After
surgery, the animals would be weaned from the anesthesia and transferred to the recovery
room. The animals would receive analgesia for 1 to 2 days per IACUC-approved protocols.
At the time of euthanasia, the vocal cords of each animal would be carefully dissected,
macroscopically evaluated and photo documented.
Post-Implantation Data
[0169] Vocal Cord Repair - At the time of euthanasia, the vocal cords of the dogs
were carefully dissected, macroscopically evaluated and photo documented. The degree of
wound healing was assessed histologically by a consulting pathologist with specific attention
paid to inflammatory infiltrates, HA staining (normal matrix production), hydrogel staining
and collagen staining (scarring). Fig. 8 shows representative histological results of control
side (unfilled) and experimental side (T-HA hydrogel filled) vocal cords stained with alcian
blue for one of the dogs three months following surgery. Gross observation indicated a more
normal appearance and vibratory properties for the T-HA hydrogel-treated vocal cord
compared to untreated controls. The histological results indicated significant scarring in the
untreated control vocal cord along the wound track as indicated by a lack of deposition of
GAG (i.e., HA) and increased collagen deposition when compared to the T-HA-filled wound
track of the experimentally repaired vocal cord.

[0170] Only small foci of T-HA hydrogel could be found in the experimental vocal
cord at 12 weeks, which show a minimal foreign body response with a layer of surrounding
mast cells observed. This may indicate degradation of the T-HA hydrogel with concomitant
deposition of normal HA-containing tissue matrix. However, given the low concentration
(2.5 mg/ml) and thus the very fluid nature of the hydrogel used in this study, it is more likely
that much of the hydrogel was lost from the wound site prior to closure of the site as the
epithelial microflap knitted to the opposed underlying tissue. Hydrogel between the
microflap and opposed underlying tissue is actually predicted to inhibit the knitting process
contributing to hydrogel loss from the wound site. Thus, the positive wound healing effect
seen is believed due to only a thin layer of hydrogel retained at the wound site rather than the
volume-filling bolus of hydrogel initially implanted. These results indicate the ability of the
hydrogel to prevent scarring and match the rheologic properties of the lamina propria.
[0171] Injection Laryngoplasty Therapy (Vocal Cord Augmentation) - At the time of
euthanasia, the vocal cords of the rabbits were carefully dissected, macroscopically evaluated
and photo documented. Histological evaluation by a consulting pathologist was used to
assess the inflammatory response and retention of the hydrogel. Representative histological
results with Alcian Blue staining to detect the hydrogel and hematoxylin & eosin (H&E) for
general morphology are shown in Fig. 9 for one of the rabbits that underwent the
augmentation procedure at two weeks post-operatively. As seen in the figure, pockets of T-
HA hydrogel could be found in the injected vocal cords at 2 weeks, which show a minimal
foreign body response with a layer of surrounding mast cells observed.
EXAMPLE 8
[0172] An experiment was conducted whereby a T-HA hydrogel as described
hereinabove was implanted into a rabbit model in order to fill the vitreous cavity of the eye in
order to prevent or treat vitreo-retinal diseases such as retinal detachment. Following is a
description of that experiment, including the experimental methods and results obtained, after
a brief discussion of the background for this application.
Background
Tissue Description and the Need for a Synthetic Material
[0173] Vitreo-retinal diseases, such as retinal detachment, diabetic retinopathy and
others, are among the most common causes of blindness. The vitreous cavity of the eye

normally is filled with a gel like substance. In retinal detachment surgery, the vitreous is
surgically removed (a procedure called "vitrectomy"), the retina is re-attached against the
back wall of the eye, and a replacement substance is injected into the vitreous cavity.
Vitreous substitutes are used for a number of different purposes in the vitreous cavity of the
eye. These include (1) achieving a long term tamponade after retinal re-attachment surgery to
keep the retina apposed to the wall of the eye; (2) in intra-operative procedures such as
unfolding of retinal tears, the removal of subretinal fluid and the flotation and removal of
dislocated intraocular lens components; (3) for developing a sustained release system that
could maintain therapeutic drug levels in the posterior segment of the eye over long periods
of time.
[0174] A number of different compounds are used as vitreous substitutes after retinal
re-attachment surgery. These compounds have physical properties that permit successful
retinal re-attachment but fail in other important surgical goals. Gases injected into the eye
provide short term retinal tamponade but re-absorb quickly and cause significant optical
distortion while they are in the eye. Perfluorocarbon liquids are effective intra-operative
tools for flattening the detached retina but cause unacceptable toxicity when left in the eye for
prolonged periods of time. Silicon oil is used as a medium term retinal tamponade but also
carries a risk of toxicity and causes significant optical distortion.
[0175] Alternatively, substitute vitreous compounds are desirable for use as safe, long
term or time-released drug delivery vehicles in the eye. Many chronic inflammatory and
infectious conditions of the eye, such as sarcoidosis, idiopathic posterior uveitis and
cytomegalovirus retinitis, necessitate intraocular injections of medication. Repeat intra-
ocular injections pose risks such as bleeding, retinal detachment and infection. A stable, non-
toxic vehicle is needed for sustained intravitreal drug delivery.
[0176] Hyaluronan (HA) is an acellular substance that is an essential component of
natural vitreous in humans and other mammals. Formulations of hyaluronan are already in
use in some ophthalmic surgical procedures. For example, sodium hyaluronate is the most
commonly used viscoelastic surgical device for anterior segment and cataract surgery.
Unfortunately, sodium hyaluronate and other previously tested hyaluronan substitutes are
dissolved relatively quickly in human tissues. These substances have not proven effective in
vitreous surgery because of their failure to provide long-term retinal tamponade.

Experimental Description
Design of Extracellular Matrix Material having Desired Properties
[0177] The most common need for a vitreous replacement is during retinal
detachment surgery. A significant challenge is maintaining the retina flat against the wall of
the eye for a prolonged period of time post-operatively. An ideal vitreous substitute should
be optically clear to allow maximum visual rehabilitation during the recovery period.
Finally, retinal detachments that occur inferiorly in the eye pose a particular challenge. For
the retina to remain flat post-operatively, the vitreous replacement must be directly apposed
to the area of the retinal tear. To tamponade inferior breaks the patient must often lie in a
face down position for weeks after the surgery. None of the vitreous substitutes in use today
satisfy all of the current clinical needs.
[0178] There is a need for a non-toxic, optically clear, vitreous substitute that will
result in improved surgical results and post-operative visual rehabilitation in patients
undergoing retinal detachment surgery.
[0179] Hydrogels made from a tyramine-substituted and cross-linked hyaluronan (T-
HA) macromolecular network as disclosed herein present an ideal choice for a synthetic
vitreous material. Specifically, the novel enzyme-driven cross-linking chemistry described
above for cross-linking the tyramine-substituted hyaluronan macromolecules using a
peroxidase and H2O2 allows the resulting hydrogels to be cross-linked ex vivo, and to remain
stable in animal tissues. For example, studies in rats have demonstrated that this material
does not degrade over several months when injected subcutaneously (see Example 9). At low
concentrations the hydrogels are optically clear, easily injected through a syringe or a
vitrectomy port and have a specific gravity higher than water. These physical properties
make T-HA gels an ideal substrate for vitreous replacement.
[0180] Based in part on confined compression testing data reported in Example 3
above, it was possible to design a T-HA hydrogel composed of, inter alia a macromolecular
network of dityramine cross-linked hyaluronan molecules, having elastic and other physical
properties matched to natural vitreous material in order to produce a synthetic implantable
vitreous substitute.

[0181] Choices of scaffold material (HA alone), percent tyramine substitution (-5%),
protocol for tyramine substitution of HA (modified from Example 1), and exclusion of cells
and bioactive factors were as described in Example 6. A concentration range of between 2.5
and 10 mg/ml of T-HA hydrogel in sterile saline most closely matched the rheologic, optical
(clarity, refractive index) and gravimetric (density) properties of the vitreous of the eye. The
10 mg/ml concentration of T-HA hydrogel was deemed most appropriate based on the
extensive clinical experience of our clinician collaborators. In vitro cross-linked hydrogel
was used rather than an in situ cross-linking protocol based on the experience of our clinician
collaborators with the potential sensitivity of the thin, highly-specialized layer of cells in the
retina. In vitro cross-linking was as described in Example 1. An insoluble steroid was added
to the cross-linked T-HA to allow visualization by the surgeons during the operative
procedure because the hydrogel material itself was optically transparent and colorless.
Surgical Procedure
[0182] Rabbits underwent unilateral vitrectomy surgery (left eye only) using standard
vitreoretinal surgical techniques with replacement of the natural vitreous of the eye with the
T-HA hydrogel described above in order to evaluate the hydrogel material as a vitreous
substitute. Following general anesthesia (ketamine: 50 mg/kg; xylazine: 5 mg/kg), the rabbit
was prepped and draped in a sterile fashion. The left eye was dilated with mydriacyl and
phenylephrine. Two drops of topical Ciloxan was instilled over the eye before and after the
case. Topical proparacaine drops were instilled. Under an operating microscope, a 270°
conjunctival peritomy was performed using wescott scissors. An infusion port was created
2.5 mm posterior to the limbus and the infusion cannula was secured to the sclera using 7-0
vicryl suture. A lens ring was sutured to the sclera using 7-0 vicryl suture. A 30° prism
vitrectomy lens was placed on the lens ring. A second port was created and the vitrectomy
instrument was inserted into the vitreous cavity. A complete core and peripheral vitrectomy
was performed. At this point one of the ports was sutured closed with a 7-0 vicryl suture.
The BSS bottle was lowered to patient level and -1.2 cc of a mixture of 3 mg/ml preservative
free triamcinolone acetonide (steroid) and the T-HA hydrogel (10 mg/ml, 5% tyramine
substitution) in BSS was injected into the vitreous cavity through an 18 gauge syringe. The
steroid is the same as usually administered after vitreoretinal surgery with its milky
appearance allowing visualization of the otherwise optically transparent hydrogel material.
As the milky solution was injected it was directly visualized filling the vitreous cavity. The
infusion was stopped when at 50% fill or when the T-HA material was seen backing up

through the irrigation canula (100% fill). After filling the vitreous cavity 7-0 vicryl suture
was used to close the remaining ports. 8-0 vicryl suture was used to close the conjunctiva.
Topical bacitracin ointment was placed on the eye after closing. A topical
antibiotic/bacitracin ointment was applied to the eye bid x 1 week, and the rabbit was placed
in a recovery cage. After the rabbit had regained sternal recumbancy, it was returned to its
home cage.
Post-Implantation Data
[0183] At 1 month post-implantation, rabbits were anesthetized with ketamine (50
mg/kg; 10 mg/kg/hr thereafter) and xylazine (5 mg/kg; 0.5 mg/kg/hr thereafter), the pupils
dilated with eye drops (1% tropicamide; 2.5% phenylephrine), and the corneal surface
anesthetized with an eye drop (0.5% proparacaine). After full pupil dilation, the status of the
retina and eye were examined by indirect ophthalmoscopy followed by fundus photography.
In addition, intraocular pressure (IOP) was measured using a Tonopen, a device that is used
clinically and which makes minimal contact with the corneal surface. FinaUy, the rabbit was
placed on a heating pad in darkness for 1 hour, and electroretinograms (ERus) recorded for
both control and vitreous replaced eyes in response to flashes of light. ERG electrodes
consist of a corneal contact lens and two platinum 0.5 inch Grass needle electrodes, placed in
the cheek and trunk. While still under anesthesia, the rabbit was euthanized by using an
intravenous dose of Beuthanasia D Special (1 ml/5 kg). Both control and vitreous replaced
eyes were then enucleated and fixed in 10% buffered formalin for 24 hours, for histological
evaluation.
[0184] The results at 1 month post-implantation, indicated minimal post-operative
inflammation of the surgically treated eye with normal IOP. By one week, a cataract had
formed in the vitreous replaced eye relative to the un-operated eye and to BSS control
operated eyes creating a limited view of the posterior segment of the experimental eye. Gross
observation of the sectioned eyes showed the cataract in the anterior segment of the vitreous
replaced eye (Fig. 10). The remainder of the anterior segment as well as the entire posterior
segment of the eye looked similar by gross observation (Fig. 10). Hydrogel was recovered
from the experimental eye at 1 month post-implantation, and was a clear gel-like substance
similar to its pre-injection form (Fig. 10). The ERG for the vitreous replaced eye was normal
compared to the un-operated control eye, and indicated that the retinal cells were alive and
remained functional (Fig. 11). Finally electron micrographs from the four quadrants of the

retina show normal morphology for the retina from the vitreous replaced eye compared to the
normal un-operated eye (Fig. 12). These data indicate the T-HA hydrogel can be used as a
vitreous substitute without causing infection or inflammation of the eye, and without
damaging the retina, and illustrate a method by which the T-HA hydrogel can be used as a
retinal tamponade for reattachment of a detached retina.
Related Ophthalmologic Applications
[0185] In addition to the foregoing retinal tamponade application following retinal
reattachment surgery, the T-HA hydrogel described in this example also could be used for the
following related applications:
" as a vitreous replacement for intra-operative procedures such as unfolding of
retinal tears, the removal of subretinal fluid and floatation and removal of
dislocated intraocular lens components.
■ as a vitreous replacement incorporating a sustained release drug delivery system
to maintain therapeutic drug levels (steroids, antibiotics, anti-viral drugs, etc.) in
the posterior segment of the eye over long periods of time to treat chronic
inflammatory and infectious conditions of the eye such as sarcoidosis, idiopathic
posterior uveitis and cytomegalovirus retinitis.
■ for anterior segment surgery including as a substitute for plastic polymer inserts in
corneal refractive surgery. These inserts, implanted surgically in the cornea, are
used to change the shape of the cornea and correct mild myopia. The optical
clarity and biocompatibility of the T-HA hydrogel with human tissues make it
well suited to this application.
" as a substitute for partial or full thickness corneal grafting procedures, e.g. for
anterior segment surgery, necessitated as a result of corneal scarring from
infection, keratoconus, or other causes. The optical and physical properties of the
hydrogels make them compatible with use as corneal tissue substitutes.
■ as a viscoelastic device during anterior segment and cataract surgery. At low
concentrations, hydrogels can maintain anterior chamber shape and pressure while
allowing the surgeon to clearly visualize ocular structures.
■ for oculoplastic surgery including subcutaneous injection to smooth wrinkles in
the face.
■ for oculoplastic surgery as an ocular implant in patients undergoing enucleation or
exoneration surgery. The hydrogel formed in the dimensions of a human eye can

be used as an implant to fill the orbit and improve cosmetic appearance of the
individual after globe removal.
■ to coat MEMS devices for use in vitreo-retinal surgery.
• to expand the utility of laser vision correction surgery (LASIK) to include those
cases where volume needs to be added to the cornea rather than removed to
correct vision. Corrective laser surgery would be used to produce the exact
dimensions required for optimal visual outcome following implantation of an
intentionally oversized plug of the hydrogel.
■ as a replacement for the typical gases, perfluorocarbon liquids and silicon oils
normally used as tamponades in eye surgery. These applications include but are
not limited to the following: giant retinal tears, proliferative vitreoretinopathy
(PVR), large breaks with "fish-mouth" phenomenon, posterior breaks or macular
holes, the restoration of intraocular volume after drainage of subretinal fluid, total
retinal detachment with multiple breaks and large meridional folds, retinal
detachment caused by ocular trauma or complicated by PVR or associated with
choroidal coloboma, dislocated lenses, suprachoroidal and submacular
hemorrhage, rhegmatogenous retinal detachments without PVR, severe
proliferative diabetic retinopathy, chronic uveitis with profound hypotony, and
infectious retinitis.
EXAMPLE 9
[0186] An experiment was conducted whereby plugs of T-HA hydrogels as described
hereinabove were implanted subcutaneously into immunocompetent rats in order to
investigate and demonstrate their in vivo persistence and longevity as well as to measure any
host immune response. As described in detail above, hydrogels comprising a cross-linked
macromolecular network (such as a tyramine-substituted and cross-linked hyaluronan
network) can be prepared having a range of physical and viscoelastic properties. These
materials, for example, can be tuned to emulate natural soft tissue and could be used for
repair or augmentation of soft tissue defects, as in plastic surgery or reconstructive surgery.
In particular, as described in detail in Examples 3 and 4 above, the viscoelasticity, rigidity
and other physical properties of the material can be tuned across a wide range to emulate like
properties of a wide variety of native soft tissues, and the material can be cast or formed into
a variety of complex anatomical shapes which would make it ideal for casting replacement or
reconstructed tissue components; e.g., in the shape of an ear or of a nose for facial

reconstruction.
[0187] While it already was clear from the noted examples that these materials could
be cast into appropriate shapes and could be given appropriate physical properties, the present
experiment demonstrates the feasibility of using the T-HA hydrogels as synthetic tissue
matrix or replacement materials in vivo. Following is a description of the experiment,
including the experimental methods and results obtained, after a brief discussion of the
background for this application.
Background
Tissue Description and the Need for a Synthetic Material
[0188] The availability of biomaterials for soft tissue augmentation and head and
neck reconstruction has remained a fundamental challenge in the field of plastic and
reconstructive surgery. Significant research and investment has been undertaken for the
development of a material with appropriate biological compatibility and life span. The
present focus in tissue engineering has been directed at attempts toward fibroblast and
chondrocyte cultures as a method of creating endogenous cartilage and collagen bearing
structures useful for implantation. The archetypal standard of this avenue of research has
been the nude mouse with a neo-cartilage ear on its back. This is based on the concept of
chondrocyte culture on a poly-lactic or poly-glycolic acid framework. The presumption is
that the chondrocytes can produce the extracellular matrix (ECM) for the production of
cartilage and create a new functional biological filling agent with complete compatibility.
The outcomes of this research have not been promising in regards to their clinical application.
When placed in immunocompetent animals the structural integrity of the neo-cartilage has
been shown to fail as the framework is absorbed. Fundamentally, while chondrocytes can
successfully be cultured and propagated they apparently cannot be made to produce cartilage
on a framework prior to its hydrolysis by the host defense mechanisms.
[0189] Conventionally, clinicians have been limited by the use of xenogenic materials
such as bovine collagen and unmodified hyaluronan (HA) as well as synthetic materials such
as silicone, silastic and hydroxyapatite. The synthetic materials are prone to foreign body
reactions and infection while the biological substrates are prone to breakdown over time. In
addition, synthetic PTFE (gortex) polymers and silastic offer less tissue reactivity but do not
offer tissue integration and also can represent long term risks of foreign body infections and
extrusion.

[0190] Instead of a tissue engineering model where chondrocytes are required to
produce a cartilage ECM, the hydrogels disclosed herein are or can be based on the same
materials that provide cartilage its functionality and feel (HA). In the present invention,
hyaluronan is used directly as the substrate for the creation of a stable tissue engineered
material to replace natural HA-rich soft tissues. In essence, the HA-based hydrogels used in
herein incorporate the same material that gives cartilage its form and structural
characteristics, but it is modified (tyramine-substituted and cross-linked) to make the material
resistant to biological degradation. Thus, an ideal synthetic extracellular matrix material
suitable for in vivo implantation and longevity is achieved.
Experimental Description
Design of Extracellular Matrix Material having Desired Properties
[0191] To provide a synthetic soft-tissue and cartilage substitute for use in head and
neck reconstruction based on the disclosed HA materials the following points were
considered: 1) optimization of enzyme-selective cross-linked hydrogels using hyaluronan as
the scaffold material; and 2) application of T-HA hydrogels as cartilage substitutes and soft-
tissue fillers. These include characterization of the effect of the cross-linked hydrogels in
vivo.
[0192] Again, based in part on confined compression testing data reported in Example
3 above, it was possible to design T-HA hydrogels composed of, inter alia a macromolecular
network of dityramine cross-linked hyaluronan molecules, having elastic and other physical
properties matched to natural soft tissues which were suitable for in vivo rat implantation to
determine their immunogenic and longevity characteristics.
[0193] Choices of scaffold material (HA alone), percent tyramine substitution (-5%),
protocol for tyramine substitution of HA (modified from Example 1), and exclusion of cells
and bioactive factors were as described in Example 6. A concentration range of between 6.25
and 100 mg/ml of T-HA hydrogel encompassed the wide spectrum of physical properties
required of a material for facial reconstruction. Therefore the same five concentration used in
Example 3 were deemed appropriate for testing in a subcutaneous rat model based on the
extensive clinical experience of our clinician collaborators. In vitro cross-linked hydrogels
were used so as to produce hydrogels of defined shape for analysis of shape retention, a
property deemed important by our clinician collaborators. In vitro cross-linking was as

described in Example 1.
Surgical Procedure
[0194] T-HA hydro gel plugs of defined shape, mass and volume (7 mm in diameter
and 3 mm in thickness) and defined mechanical properties based on HA concentration were
surgically implanted subcutaneously in the backs of immunocompetent rats to allow
evaluation of their in vivo persistence and host immune response based on previously
published protocols for the evaluation of collagen and other HA-based hydrogels. After
induction of anesthesia with intraperitoneal injection of ketamine (100 mg/kg) and xylazine
(5 mg/kg), the rat received a single intramuscular injection of 60,000 units of procaine
penicillin for infection prophylaxis. A 1cm stab incision was made with a #11 surgical blade
in the lower lumbar region of the rat. A 14 g needle was used as a trocar to dissect in the
subcutaneous plane to create a pocket. Three preformed hydrogel plugs (~7.1 mm diameter x
3 mm thick) of one of the HA concentrations to be tested were inserted into the surgical
pocket. A single absorbable stitch (3-0 Chromic) was placed to re-approximate the skin edge.
At 1- week, 1 month, 3 months, and 6 months post-implantation, rats were sacrificed by CO2
asphyxiation and the T-HA hydrogel plugs with surrounding tissue excised and stored in
formalin at 4°C until time for histological evaluation.
Post-Implantation Data
[0195] Hydrogel compositions tested included plugs made from concentrations of
6.25, 12.5, 25, 50, and 100 mg/ml of HA; generating hydrogel plugs with a wide spectrum of
physical properties ranging from that of gel to a paste to a rubber-like material (see Example
3). Implanted T-HA hydrogel plugs were collected at 1 week, 1 month, 3 months, and 6
months post-implantation. Excised plugs were evaluated for their in vivo persistence and host
immune response.
[0196] Fig. 13 shows representative results of histological staining with H&E, alcian
blue, MC/giemsa, Movats, Reticular, and Trichrome stains for the 100 mg/ml TB-HA
hydrogel plug from the 1-month time point. Clearly defined in Fig. 13 are the surface hair
follicles, the superficial muscle layer, the hydrogel plug, and the thin fibrous capsule
surrounding the hydrogel plug as a result of a minimal foreign body response. An artifact
exists as a result of hydrogel shrinkage from the paraffin embedding process, which can be
avoided through the use of frozen sections. The results show very little immune response
with only a thin layer of mast cells surrounding the plug, and no evidence for host cell

infiltration into the plug. When measured, the volume of the void left by the plug during
histological processing is 3 mm (the original plug thickness) indicating little to no
biodegradation or deformation of the hydrogel matrix. Staining indicated that the plugs had
little protein, such as collagen or elastin, deposited within them and remained primarily
composed of HA hydrogel. These results indicate that the hydrogel plugs over a broad range
of five concentrations persisted through 6 months with little evidence of degradation, host
immune response, and cellular infiltration providing a wide range of injectable materials for
use in soft tissue reconstruction.
EXAMPLE 10
[0197] It will be apparent from the foregoing discussion and the Examples that
hydrogels described herein composed of a cross-linked (in situ or ex vivo) macromolecular
network of hyaluronan molecules cross-linked via a suitable dihydroxyphenyl cross-linking
chemistry as herein described are suitable as a synthetic, implantable extracellular matrix
tissue material for a variety of tissue engineering and repair applications. A particular such
application for which the disclosed hydrogel materials will have particular utility is in the
repair or augmentation of the mitral valve in a heart.
[0198] The mitral valve is one of the most complex connective tissue structures in the
body. It consists of two leaflets and numerous chordae tendineae. These chordae have a
highly aligned collagenous core and a thin outer sheath of elastic fibers and endothelial cells.
Both leaflets are laminated tissues containing a heavily collagenous layer on the ventricular
side, a predominantly elastic layer on the atrial side, and an inner spongiosa layer containing
abundant proteoglycans (PGs) and hyaluronan (HA). The relative thicknesses of these layers
vary between the two leaflets and also within each leaflet from its attachment edge to its free
edge.
[0199] The variability of the different leaflet layers, and hence the structural
constituents within the mitral valve, are determined by the specific functional roles of the
leaflets and chordae. The closed valve maintains a balance of tensile and compressive loads,
in which the chordae and the flat central region of the anterior leaflet are in tension, whereas
the free edge of the anterior leaflet and most of the posterior leaflet are in appositional
compression. Accordingly, the most collagenous components of the mitral apparatus are the
chordae and the portion of the anterior leaflet between the annulus and the upper appositional

border. In the posterior leaflet and in the free edge of the anterior leaflet, the collagenenous
layer is relatively thinner, whereas the PG rich spongiosa is substantially thicker. The wide
diversity of glycosaminoglycans (GAGs) and their parent PGs exert considerable yet variable
control over the physical properties of the extracellular matrix.
[0200] Functional mitral regurgitation (MR) refers to the regurgitation that occurs
with a structurally normal valve as a consequence of left ventricular (LV) dysfunction, and as
a result, almost half the patients with LV dysfunction have at least moderate MR. Functional
MR plays a pivotal role in the pathophysiology of congestive heart failure (CHF), a major
cause of cardiac morbidity and mortality. Several studies have shown that the presence of
functional MR in patients with CHF is associated with poor outcomes. Although this
observation could suggest that MR is merely an indicator of CHF severity, it is also
increasingly apparent that the development of the MR hastens the progression of CHF. The
precise mechanism of functional MR remains controversial and can relate to mitral annular
dilatation in the septal-lateral (S-L) axis or tethering of the leaflets secondary to progressive
ventricular remodeling. MR leads to greater volume overload of the LV with progressive
annular dilatation and increased MR, creating a "vicious cycle" which exacerbates the
problem. MR is commonly considered to be one of the initiators of CHF, as well as an
ongoing impetus of the progression of the disease.
[0201] Surgical annuloplasty is a widely used method for mitral valve repair and can
provide long-term benefit. However, the surgical procedure requires access to and
manipulation of the valve annulus via atriotomy. In addition, the procedure requires the
patient to be placed on cardiopulmonary bypass (CPB). The prolonged CPB time has been
suggested as a cause of not only postoperative LV dysfunction but also main organ
dysfunction. The use of heparin during CPB results in an increased risk of bleeding
complications. The increased morbidity and mortality profile leads many care providers
directly to non-treatment options of MR in the earlier stage heart failure patients.
[0202] Recently, several minimally invasive methods of mitral valve repair have been
developed. For example, several investigators have reported the preliminary methodology of
off-pump mitral valve repair procedures through a thoracic incision. Others have reported
new devices that can be inserted percutaneously into the coronary sinus and great cardiac
vein to reduce the S-L dimension of the mitral annulus. There are the possibilities of adverse

effects, however, such as obstruction or disturbance of the coronary circulation, by chronic
placement of the device in the coronary sinus,

[0203] Surgical therapy for functional MR, including mitral valve repair with an
annuloplasty ring and replacement with an artificial valve has been limited in patients with
severe CHF by relatively high operative mortality rates due to the effects of CPB. Therefore,
there is a need for a minimally invasive procedure that will not compromise coronary
circulation and will allow for reduction of the S-L dimension of the mitral annulus to reduce
functional MR as well as other forms of MR. Myxomatous changes in mitral valve tissues
can lead to leaflet prolapse and mitral regurgitation.
[0204] The T-HA hydrogel materials disclosed herein could be adapted for this
purpose; i.e. mitral annular remodeling resulting from a nonabsorbable substance injection
(namely a T-HA hydrogel material designed to have the necessary viscoelastic and other
physical properties) into the posterior mitral annulus using an epicardial approach. This
procedure would enable the S-L dimension to be efficiently reduced, thus reducing MR,
without employing CPB or implanting a device into the coronary sinus. This mitral annular
remodeling procedure could be modified to allow for percutaneously injection of the
substance through the coronary sinus.
[0205] This application for a T-HA hydrogel as disclosed herein would enable a
nonabsorbable substance to be percutaneously injected through the coronary sinus for severe
CHF patients with functional MR who are unable to receive conventional mitral valve
surgery. This minimally invasive approach would obviate the need for CPB and sternotomy,
as well as diminish the risk of major side effects from conventional surgical therapy such as
postoperative LV dysfunction and resulting poor organ perfusion. In addition, such a
procedure would provide patients with mild to moderate CHF with an option for early
restoration of mitral valve competence, arresting the initiation and progression of devastating
heart failure.
[0206] In particular, a hydrogel composed of a cross-linked macromolecular network
as described herein, particularly of HA, could be designed based on the principles as
elucidated in Example 3 above, in order to produce a hydrogel material having all of the
following characteristics which would be considered desirable for this application:

■ Injectable and nonabsorbable;
■ Low-grade inflammatory reaction;
■ Low evidence of foreign body migration;
■ Ease of collagen encapsulation which contributes to the prevention of migration;
■ Not especially malleable nor especially rigid.
[0207] A protocol has been established for the injection of a T-HA hydrogel material
to augment the mitral valve of a beating heart. That protocol is described as follows.
[0208] Injection Procedure - Two dimensional epicardial echocardiography (2D EE)
and transesophageal echocardiography (2D TEE) are performed to evaluate LV end-diastolic
and end-systolic volumes (EDV and ESV), stroke volume (SV), ejection fraction (EF), the S-
L dimension of mitral annulus, and the degree of MR. Hemodynamic data such as LVP,
LAP, the central venous pressure (CVP), the pulmonary arterial pressure (PAP), the
pulmonary capillary wedge pressure (PCWP), CO, LAD flow, and LCX flow should be
collected. LV end-diastolic and end-systolic pressure-volume relations can be obtained by
transient TVC occlusion using an occlusion catheter to assess LV contractility and compliance
(baseline before injection).
[0209] A commercially available cardiac stabilizer used for ofT-pump coronary
bypass grafting can be used to stabilize the target region. Under 2D TEE guidance, the
uncross-linked T-HA hydrogel composition (50 mg/ml in sterile saline) is injected into the
posterior mitral annulus from the outside of the heart while the heart is beating. During the
injection, 2D can be used to assess the position of the tip of the needle for the injection, the
range occupied by the substance in the posterior annulus. Once an appropriate fill and
repositioning of the mitral valve has been accomplished, cross-linking is initiated by injection
of 0.2 cc of 0.6% hydrogen peroxide. After completion of cross-linking of the hydrogel
implant, data including hemodynamics, coronary flow, LV pressure-volume loops (LV P-V
loops), 2D EE, and 2D TEE should be collected (data after injection).
[0210] The foregoing injection methodology was developed and evaluated using
cadaveric dog and pig hearts as models. Fig. 14 shows a cadaveric dog heart in which a T-
HA hydrogel material was injected and cross-linked in situ via the foregoing injection

methodology. To produce the T-HA hydrogel material for this experiment, the scaffold
material (HA alone), percent tyramine substitution (~5%), protocol for tyramine substitution
of HA (modified from Example I) and non-inclusion of cells and bioactive factors were as
described in Example 6. Concentrations of 25, 50 and 100 mg/ml of T-HA hydrogel in saline
most closely matched those of cardiac (heart) tissue required for mitral valve closure. The 50
mg/ml T-HA hydrogel was deemed most appropriate for a mitral annular remodeling
procedure by percutaneously injection based on the extensive clinical experience of our
clinician collaborators. Peroxidase would be added at 10 U/ml prior to application in an in
situ cross-linking protocol as the one described below. The in situ cros linking protocol is
preferred as it would allow the uncross-linked T-HA to pass through an appropriately sized
needle while the cross-linked hydrogel may not. In addition, the in situ cross-linking protocol
allows the surgeon first to properly position (close) the mitral valves by injection of the
uncross-linked hydrogel and then cross-link the hydrogel into a solid implant only after visual
confirmation that the mitral valve had been properly repositioned.
[0211] In Fig. 14, the implanted hydrogel was bisected post-'^Dlantation to
demonstrate its solid, viscoelastic character following in situ cross- ng, as well as to
evaluate its placement in the heart. In particular, these models were used to 1) evaluate the
appropriate concentration of hydrogel required to both mimic the consistency of cardiac
muscle after cross-linking yet pass through the injection port prior to cross-linking; 2)
demonstrate reproducibility for the in vivo cross-linking protocols for complete cross-linking
in vivo at the required volumes (~2 ml); and development suitable injection techniques. All
of these goals were met, establishing confidence that the injection procedure as well as a
suitable T-HA hydrogel can precisely accommodate anatomical constraint on the mitral
annulus.
[0212] Although the above-described embodiments constitute the preferred
embodiments, it will be understood that various changes or modifications can be made
thereto without departing from the spirit and the scope of the present invention as set forth in
the appended claims.

WE CLAIM:
1. A synthetic, implantable tissue matrix material comprising a macromolecular network
comprising

wherein R1 and R2 each comprises a structure selected from the group consisting of
a) polycarboxylate molecules that have been substituted at CO2H sites thereon with a
hydroxyphenyl compound at a substitution rate less than 10 percent based on a total
number of CO2H sites on the polycarboxylate molecules,
b) polyamines, that have been substituted at primary amine sites thereon with a
hydroxyphenyl compound, and
c) copolymers thereof, and wherein R1 and R2 can be the same or different structures.
2. The synthetic, implantable tissue matrix material as claimed in claim 1, wherein R1 is a
polycarboxylate.

3. The synthetic, implantable tissue matrix material as claimed in claim 1, wherein R1 is a
polyamine.
4. The synthetic, implantable tissue matrix material as claimed in claim 1, wherein R1
comprises a structure selected from the group consisting of glycosaminoglycans.
5. The synthetic, implantable tissue matrix material as claimed in claim 1, wherein R1
comprises hyaluronan.
6. The synthetic, implantable tissue matrix material as claimed in claim 1, wherein R1
comprises chondroitin sulfate.
7. The synthetic, implantable tissue matrix material as claimed in claim 6, said
chondroitin sulfate being in the form of aggrecan.
8. The synthetic, implantable cartilage material comprising the synthetic, implantable
tissue matrix material as claimed in claim 1.
9. The synthetic, implantable vocal cord material comprising the synthetic, implantable
tissue matrix material as claimed in claim 1.
10. The synthetic, implantable vitreous material comprising the synthetic, implantable
tissue matrix material as claimed in claim 1.

11. The synthetic, implantable soft tissue material comprising the synthetic, implantable
tissue matrix material as claimed in claim 1.
12. The synthetic, implantable mitral valve material comprising the synthetic,
implantable tissue matrix material as claimed in claim 1.
13. The synthetic, implantable tissue matrix material as claimed in claim 1, said network
comprising said polycarboxylate molecules, wherein at least one dihydroxyphenyl
linkage is formed between two hydroxyphenyl groups attached respectively to adjacent
polycarboxylate molecules.
14. The synthetic, implantable tissue matrix material as claimed in claim 1, said network
comprising said polyamine molecules, wherein at least one dihydroxyphenyl linkage is
formed between two hydroxyphenyl groups attached respectively to adjacent polyamine
molecules.

15. The synthetic, implantable tissue matrix material as claimed in claim 13, said
polycarboxylate molecules having a hydroxyphenyl compound substitution rate less than
9 percent based on the total number of CO2H sites present on said polycarboxylate
molecules.
16. The synthetic, implantable tissue matrix material comprising a macromolecular
network comprising a plurality of tyramine-substituted hyaluronan molecules, at least
two adjacent hyaluronan molecules being linked by a dityramine linkage.
17. The synthetic, implantable cartilage material comprising the synthetic, implantable
tissue matrix material as claimed in claim 16.
18. The synthetic, implantable vocal cord material comprising the synthetic, implantable
tissue matrix material as claimed in claim 16.
19. The synthetic, implantable vitreous material comprising the synthetic, implantable
tissue matrix material as claimed in claim 16.
20. The synthetic, implantable soft tissue material comprising the synthetic, implantable
tissue matrix material as claimed in claim 16.
21. The synthetic, implantable mitral valve material comprising the synthetic,
implantable tissue matrix material as claimed in claim 16.

22. The synthetic, implantable tissue matrix material as claimed in claim 16, having a
tyramine substitution rate on said hyaluronan molecules of 1.7%-10% based on the total
number of CO2H sites present on said hyaluronan molecules.
23. The synthetic, implantable tissue matrix material as claimed in claim 16, having a
tyramine substitution rate on said hyaluronan molecules of 1.7%- 5% based on the total
number of CO2H sites present on said hyaluronan molecules.
24. The synthetic, implantable tissue matrix material as claimed in claim 1, wherein R1
and R2 each comprise hydroxyphenyl-substituted hyaluronan.
25. The synthetic, implantable tissue matrix material as claimed in claim 24, wherein
hyaluronan molecules for each of R1 and R2 have aggrecan attached thereto.
26. The synthetic, implantable tissue matrix material as claimed in claim 24, said
hydroxyphenyl-substituted hyaluronan having a hydroxyphenyl substitution rate less than
9 percent.
27. The synthetic, implantable tissue matrix material as claimed in claim 24, said
hydroxyphenyl-substituted hyaluronan having a hydroxyphenyl substitution rate less than
5 percent.

28. The synthetic, implantable tissue matrix material as claimed in claim 24, said
hydroxyphenyl-substituted hyaluronan being tyramine-substituted hyaluronan.
29. The synthetic, implantable tissue matrix material as claimed in claim 28, said
tyramine-substituted hyaluronan having a tyramine substitution rate less than 9 percent.
30. The synthetic, implantable tissue matrix material as claimed in claim 28, said
tyramine-substitution hyaluronan having a tyramine substitution rate less than 5 percent.
31. The synthetic, implantable tissue matrix material as claimed in claim 1, wherein R1
and R2 each comprise hydroxyphenyl-substituted chondroitin sulfate.
32. The synthetic, implantable tissue matrix material as claimed in claim 31, wherein
chondroitin sulfate molecules for each of R1 and R2 are in the form of aggrecan.
33. The synthetic, implantable tissue matrix material as claimed in claim 31, said
hydroxyphenyl-substituted chondroitin sulfate being tyramine-substituted chondroitin
sulfate.
34. The synthetic, implantable tissue matrix material as claimed in claim 33, said
tyramine-substituted chondroitin sulfate having a tyramine substitution rate less than 9
percent.

ABSTRACT

Title: Hydroxypenyl cross-linked macromolecular network and applications thereof.
A synthetic, implantable tissue matrix material comprising a macromolecular network
comprising

wherein R1 and R2 each comprises a structure selected from the group consisting of a)
polycarboxylate molecules that have been substituted at CO2H sites thereon with a
hydroxyphenyl compound at a substitution rate less than 10 percent based on a total
number of CO2H sites on the polycarboxylate molecules, b) polyamines, that have been
substituted at primary amine sites thereon with a hydroxyphenyl compound, and c)
copolymers thereof, and wherein R1 and R2 can be the same or different structures.

HYDROXYPHENYL CROSS-LINKED MACROMOLECULAR NETWORK AND APPLICATIONS THEREOF

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