Synthetic hydrogel composite

Abstract

The cellulose-reinforced hydrogel is a cellulose nanofiber network structure and an interstitial hydrogel portion within the interstitial region of the cellulose nanofiber network structure, the interstitial hydrogel portion may include an interstitial hydrogel portion containing polyvinyl alcohol (PVA) and having a crystallinity of 20% or more of the hydrogel component.

Description

Technical Field

[0001] Claims of Priority

[0001] This patent application claims the priority of U.S. Patent Application No. 17 / 845,881, entitled "SYNTHETIC HYDROGEL COMPOSITE," filed on June 21, 2022, and U.S. Provisional Patent Application No. 63 / 338,439, entitled "SYNTHETIC HYDROGEL COMPOSITE," filed on May 4, 2022, the entire contents of each of which are incorporated herein by reference. Incorporation by Reference

[0002] All publications and patent applications mentioned in this specification are hereby incorporated by reference in their entirety to the same extent as if each individual publication or patent application were specifically and individually indicated to be incorporated by reference.

Background Art

[0002]

[0003] Each year, approximately 900,000 people in the United States suffer from damage to the articular cartilage that lines the ends of the bones. Osteoarthritis (OA) is an example of a degenerative joint disease that is a common cause of disability. Articular cartilage lesions most commonly occur in the knee and can cause debilitating pain. Cartilage has no blood vessels and has a very limited ability to heal. One approach to treating cartilage lesions is an attempt to regrow the cartilage using techniques such as microfracture or autologous chondrocyte implantation. Unfortunately, these methods have a high failure rate, long rehabilitation times, and show decreased effectiveness in elderly patients. Transplantation of fresh osteochondral allografts can accelerate recovery by eliminating the need to regrow the cartilage. Unfortunately, the limited supply of fresh allografts restricts the number of these procedures. Failure of these treatment strategies typically results in a more invasive total knee arthroplasty. While total knee arthroplasty can be successful in elderly patients, it may not be suitable for younger patients who are at risk of implant damage during their lifetime and may require a second invasive surgery. Therefore, there is a clear need to treat cartilage lesions with a low failure rate, enable rapid recovery, and minimize invasive treatment options that are widely available.

[0003]

[0004] In consideration of the need for less invasive alternatives to total knee arthroplasty for the treatment of OA, efforts have continued to replace damaged cartilage with devices made from traditional orthopedic materials such as cobalt-chromium alloys or ultra-high molecular weight polyethylene. However, these materials have a much higher coefficient of friction (COF) than cartilage and can cause unacceptable levels of wear on the opposing cartilage surfaces. In addition, these materials are much harder than cartilage and can therefore cause abnormal stress distributions in the joint, potentially causing damage to the surrounding cartilage.

[0004]

[0005] Hydrogels, hydrogel networks swollen with water, are promising synthetic materials for cartilage replacement because the hydrogels can be made to have mechanical and tribological properties similar to those of natural cartilage. However, it is necessary to improve the physical properties of the hydrogels to withstand wear and fracture that the implants may encounter. For example, the hydrogels need to have an elastic modulus, coefficient of friction, and wear resistance similar to those of cartilage, but at the higher end of the strength range reported for cartilage. Described herein are methods, hydrogel compositions, and devices (e.g., implants) that can address these needs.

SUMMARY OF THE INVENTION

MEANS FOR SOLVING THE PROBLEM

[0005]

[0006] The present disclosure generally relates to an artificial cartilage material in an implant suitable for cartilage repair, including a hydrogel composite, and a method for attaching the hydrogel composite to the surface of the implant.

[0006]

[0007] Described herein is a hydrogel material for use as artificial cartilage in an implant. The hydrogel can be infused into a nanofiber material (e.g., a nanofiber network) and bonded to the surface of an implant such as a porous base. The composite hydrogel has physical properties such as strength, elastic modulus, and wear resistance, and coefficient of friction (COF) that approximate or exceed those of healthy cartilage bonded to bone. The method includes a strengthening process that increases the crystallinity of the hydrogel and decreases the water content, thereby improving its mechanical properties for implementation as a cartilage replacement. As described herein, the strengthening of the hydrogel can include one or more of the steps of drying, annealing, and rehydrating to affect the crystal structure of the hydrogel. The method can further include the steps of fixing a nanofiber material to the surface of the implant, infiltrating the hydrogel into the nanofiber network, and annealing the hydrogel.

[0007]

[0008] To mimic cartilage, a method of generating synthetic cartilage by infiltrating a hydrogel into a nanofiber network structure is described in International Patent Application No. PCT / US2021 / 040031, the entire content of which is incorporated herein by reference. The methods described herein can be used to form hydrogels that have an elastic modulus, coefficient of friction, and wear resistance similar to cartilage, while matching or exceeding the higher end of the strength range of cartilage.

[0008]

[0009] Described herein are hydrogels that can be intermeshed with nanofiber network structures, such as cellulose nanofiber network structures, for mimicking or replacing cartilage, and methods of making and using the hydrogels. The incorporated hydrogel can have a crystalline structure that imparts high tensile strength and / or compressive strength to the hydrogel. In some examples, a reinforcing hydrogel for use in the implants described herein can include a cross-linked cellulose nanofiber network structure; and a hydrogel injected into the interstitial regions of the cross-linked cellulose nanofiber network structure, where the hydrogel has a crystallinity of 20% or more. In some examples, the hydrogel includes polyvinyl alcohol (PVA). In any of these examples, the hydrogel may exclude (or substantially exclude) PAMPS. The hydrogel can be PVA that is annealed as described herein and is greater than 90% PVA (e.g., greater than 92%, greater than 93%, greater than 94%, greater than 95%, greater than 96%, greater than 97%, greater than 98%, greater than 99%, etc.).

[0009]

[0010] As demonstrated herein, crystallites formed during annealing strengthen other amorphous polymeric hydrogels by acting as cross-links that redistribute the applied stress and prevent crack propagation. Crystallites also increase the solid content and strength of the hydrogel by reducing the amount of water taken up by the PVA after annealing.

[0010]

[0011] Described herein is an implant comprising an implant body and a cross-linked cellulose nanofiber network structure bonded to the porous surface of the implant body by cement; and a cellulose-reinforced hydrogel material comprising a hydrogel impregnated in the cross-linked cellulose nanofiber network structure, wherein the hydrogel has a crystallinity of 20% or more. The implant body may comprise a porous surface. For example, the implant body may be a titanium body having a porous surface facing the bone and a non-porous surface facing the hydrogel.

[0011]

[0012] The hydrogels described herein may have a water content that contributes to imparting high tensile strength and / or compressive strength to the hydrogel. In some examples, the hydrogel may have at least 20 weight percent (wt%) water and may have a tensile strength exceeding that of cartilage, for example, exceeding 40 megapascals (MPa).

[0012]

[0013] The composition of the interstitial hydrogel may be selected to maximize crystallinity. For example, some hydrogel polymers and / or polymer mixtures have been found to interfere with crystal formation and thereby reduce the tensile and compressive strengths of the composite hydrogel.

[0013]

[0014] Generally, a hydrogel can be composed of one or more polymers that contribute to the formation of a crystal structure. In some examples, the hydrogel may include polyvinyl alcohol (PVA). In some cases, the hydrogel may contain only one type of polymer. In some modified forms, the hydrogel is polyvinyl alcohol (PVA), poly(2-acrylamido-2-methyl-1-propanesulfonic acid sodium salt) (PAMPS), poly(N,N'-dimethylacrylamide) (PDMAAm), a copolymer of 1-vinylimidazole and methacrylic acid, an amphiphilic triblock copolymer, a polyampholyte hydrogel, a PVA-tannic acid hydrogel, a poly(N-acryloyl)glycine amide hydrogel, a polyacrylic acid-acrylamide-C18 hydrogel, guanine-boric acid reinforced PDMAAm, a polyelectrolyte hydrogel, a poly(acrylonitrile-co-1-vinylimidazole) hydrogel (e.g., a mineralized poly(acrylonitrile-co-1-vinylimidazole) hydrogel), a polyacrylic acid-Fe3+-chitosan hydrogel, a poly(methacrylic acid) gel, a graphene oxide / xonotlite reinforced polyacrylamide (PAAm) gel, a poly(stearyl methacrylate)-polyacrylic acid gel, an annealed PVA-polyacrylic acid hydrogel, a supramolecular hydrogel from a multiurea-bonded segmented copolymer, a polyacrylonitrile-PAAm hydrogel, a microsilica reinforced DMA gel, an agar-polyhydroxyethyl methacrylate gel, a polyfacryloyloethyl trimethylammonium chloride hydrogel, a poly(3-(methylacryloylamino)propyl)trimethylammonium chloride hydrogel, a poly(sodium p-styrenesulfonate) hydrogel, a polyethylene glycol diacrylate hydrogel, and a polyethylene glycol hydrogel. In some cases, it may be beneficial to exclude PAMPS (e.g., having no PAMPS, having less than 0.1%, less than 0.5%, less than 1% of PAMPS, etc.).

[0014]

[0015] The nanofiber network structure may include a cellulose nanofiber network structure. The nanofiber network structure may include a crosslinked cellulose nanofiber network structure. In some examples, the nanofiber network structure includes bacterial cellulose (BC). Additionally or alternatively, the nanofiber network structure may include at least one of electrospun polymer nanofibers, polyvinyl alcohol (PVA) nanofibers, aramid nanofibers, aramid-PVA nanofibers, wet-spun silk protein nanofibers, chemically crosslinked cellulose nanofibers, and polycaprolactone (PCL) fibers.

[0015]

[0016] The cellulose-reinforced hydrogel may include a cellulose nanofiber network structure; and a hydrogel impregnated in the cellulose nanofiber network structure, and the hydrogel has a crystallinity of 20% or more. The cellulose-reinforced hydrogel material may have a tensile strength of 40 MPa or more. The cellulose-reinforced hydrogel material may have a compressive strength of 59 MPa or more.

[0016]

[0017] Described herein is a cellulose-reinforced hydrogel comprising a water content of at least 20 wt% and a compressive strength of more than 59 MPa. The cellulose-reinforced hydrogel may include a hydrogel comprising bacterial cellulose and / or polyvinyl alcohol (PVA).

[0017]

[0018] Described herein is a method for forming a cellulose-reinforced hydrogel, comprising the steps of infiltrating a hydrogel into a cellulose nanofiber network to form a cellulose-reinforced hydrogel, and annealing the hydrogel to increase the crystalline content of the hydrogel. The step of annealing the hydrogel may include heating the cellulose-reinforced hydrogel. The step of annealing the hydrogel may include heating the cellulose-reinforced hydrogel to reduce the water content of the hydrogel. In some examples, the cellulose-reinforced hydrogel may be heated to a temperature in the range of 90 to 140 °C. The step of annealing the hydrogel may include rehydrating the hydrogel. The step of rehydrating the hydrogel may include increasing the water content of the hydrogel to at least 20 wt%. The method may further include the step of removing excess hydrogel from the surface of the cellulose nanofiber network. The step of removing excess hydrogel may include removing the excess hydrogel manually or by shaping the cellulose-reinforced hydrogel.

[0018]

[0019] Described herein is an implant knee joint resurfacing device, comprising an upper support surface including a cellulose nanofiber network and a cellulose-reinforced hydrogel comprising a hydrogel impregnated in the cellulose nanofiber network, the cellulose-reinforced hydrogel having a crystallinity of 20% or more.

[0019]

[0020] A hydrogel (e.g., a cellulose-reinforced hydrogel) can be attached to a metal base with a shear strength of more than 0.2 MPa.

[0021] When used in partial knee joint resurfacing, the implant can be configured to abrade the opposing cartilage surface to an extent that is not significantly greater than the extent to which the cartilage abrades itself. The upper support surface of the implant may have a coefficient of friction (COF) that is not statistically different from the coefficient of friction of cartilage.

[0020]

[0022] The implants described in this specification can be configured as medical implants and can include tissue engagement portions (e.g., bone engagement portions such as rods, screens, nails, etc.). The first surface of the implant to which the nanofiber network can be fixed can be porous. For example, the first surface can be more than 40% porous to a depth of 1 mm or more.

[0021]

[0023] The nanofiber network can be fixed to the implant (e.g., the porous surface of the implant) by any suitable method. For example, the nanofiber network may be fixed to the implant by a cement such as α-TCP cement. In some examples, the cement can be a resin-based (dental) cement containing silicate or other types of fillers in an organic resin matrix, such as zinc oxide eugenol, glass ionomer, calcium silicate, polycarboxylate cement, zinc phosphate, acrylate or methacrylate resin cement (e.g., a methacrylate cement such as "RelyX (trademark) Unicem 2 Self-Adhesive Resin Cement" or "RelyX (trademark) Ultimate Adhesive Resin Cement"), and one or more of resin-modified glass ionomer cements. The cement may include (but is not limited to) an adhesive such as phosphoserine (PPS). In some variants, the cement may include reinforcing particles such as stainless steel particles (e.g., stainless steel powder, SSP).

[0022]

[0024] The cement can extend at least 5 microns (e.g., 6 microns or more, 7 microns or more, 8 microns or more, 10 microns or more, 15 microns or more, 20 microns or more, etc.) into the nanofiber network from the first surface. The cement may not bind to the hydrogel in some cases. In some examples, cement fixation can be completed before impregnating with the hydrogel (the cement hardens or dries).

[0023]

[0025] Cement can bind to the nanofiber network structure but does not bind directly to the hydrogel. This can be the result of a method of forming a network structure-reinforced hydrogel in which the nanofiber network structure (e.g., a cellulose nanofiber network structure) is first fixed to the implant body (e.g., cement fixation) before impregnating the hydrogel. Since the cement can be cured on the nanofiber network structure, it does not bind directly to the hydrogel.

[0024]

[0026] Other adhesives may include surgical adhesives such as cyanoacrylate, gelling / resorcinol / formaldehyde (GRF), and / or fibrin.

[0027] The implant can be formed from any suitable biocompatible material. For example, the surface of the implant body may be titanium. The surface of the implant body may be one or more of a stainless steel alloy, a titanium alloy, a Co-Cr alloy, tantalum, gold, niobium, bone, aluminum oxide, zirconium oxide, hydroxyapatite, tricalcium phosphate, sodium calcium phosphosilicate, poly(methyl methacrylate), polyetheretherketone, polyethylene, polyamide, polyurethane, or polytetrafluoroethylene.

[0025]

[0028] As described, the attachment surface to which the nanofiber network structure is fixed can be porous. Alternatively, the attachment surface may be non-porous. For example, the attachment surface can be 20% or more (such as 30% or more, 40% or more, 50% or more, 60% or more, 70% or more, etc.) porous to a depth of 0.5 mm or more (such as 0.6 mm, 0.7 mm, 0.8 mm, 0.9 mm, 1 mm or more, etc.). As used herein, the percentage by which a surface is porous (e.g., the percent porosity of the surface) may refer to the percentage of the surface within the depth that forms open spaces within the surface that are absent. These open spaces can be referred to as pores, and some of them can be connected to each other (e.g., fluidly connected). The attachment surface can be configured in contact with another surface (e.g., a bone surface, etc.) and may be subjected to load-bearing surfaces, and thus may generally be referred to herein as an upper support surface.

[0026]

[0029] In any of these devices (e.g., devices, systems including implants), at least a portion of the nanofiber network structure can be mineralized. For example, at least a portion such as a region near the interface with the surface can be mineralized with hydroxyapatite. Mineralization can extend at least 5 microns (e.g., at least 7 microns, at least 8 microns, at least 9 microns, at least 10 microns, at least 15 microns, at least 20 microns, etc.) from the surface into the nanofiber network structure.

[0027]

[0030] Generally, the nanofiber network structure can be connected to the upper support surface of the implant. The cross-linked cellulose nanofiber network structure can be attached to the upper load surface by clamps and / or adhesives. For example, the nanofiber network structure can be bonded to the upper load surface by cement; in some examples, the cement does not bond to the hydrogel; the cement bonds only to the nanofiber network structure. Alternatively, in some examples, the nanofiber network structure can be connected to the implant, such that the nanofiber network structure is fixed on the upper support surface without using a chemical adhesive such as epoxy. Instead, the nanofiber network structure can be fixed on the upper support surface by clamps. For example, the clamp can fix the nanofiber network structure (e.g., a sheet of one or more BCs) on the upper support surface at the periphery of the upper support surface. Thus, generally, the use of an adhesive (such as epoxy) is optional.

[0028]

[0031] Any suitable implant may be used. The surface of the implant (e.g., the upper support surface, which may equivalently be simply called the support surface) may be an area where at least a nanofiber network structure is adhered thereon, and may be made of titanium, stainless steel, etc., and the support surface (e.g., the upper support surface) may be convex, flat, concave, or a mixture thereof. For example, the surface of the implant body may include one or more of a stainless steel alloy, a titanium alloy, a Co-Cr alloy, tantalum, gold, niobium, bone, aluminum oxide, zirconium oxide, hydroxyapatite, tricalcium phosphate, sodium calcium phosphosilicate, poly(methyl methacrylate), polyetheretherketone, polyethylene, polyamide, polyurethane, or polytetrafluoroethylene.

[0029]

[0032] Also described herein are methods of making and / or using these implants. For example, described herein is a method of attaching a hydrogel to a surface such that the hydrogel is fixed to the surface with a shear strength of greater than 1 MPa. Any of these methods may include a step of infiltrating the hydrogel into the cellulose nanofiber network structure to form a cellulose-reinforced hydrogel, and a step of annealing the hydrogel to increase the crystalline substances of the hydrogel. For example, the step of annealing the hydrogel may include a step of heating the cellulose-reinforced hydrogel. In some examples, the step of annealing the hydrogel may include a step of heating the cellulose-reinforced hydrogel to reduce the water content of the hydrogel. For example, the cellulose-reinforced hydrogel may be heated to a temperature in the range of 90 to 140 °C. In some cases, the step of annealing the hydrogel may include a step of rehydrating the hydrogel. The step of rehydrating the hydrogel may include a step of increasing the water content of the hydrogel to at least 20 wt%. The method may also include a step of removing excess hydrogel from the surface of the cellulose nanofiber network structure. The step of removing excess hydrogel may include a step of manually removing the excess hydrogel or removing the excess hydrogel by shaping the cellulose-reinforced hydrogel.

[0030]

[0033] In some examples, the outer surface of the hydrogel can be formed to be smooth (e.g., having a roughness of less than 30 microns). For example, the methods described herein can include mechanically polishing the outer surface of the hydrogel to a roughness of less than 30 microns. In some cases, the outer surface can be smoothly formed by molding, including using a smooth mold to mold the heated polymer. For example, the step of infiltrating the hydrogel into the nanofiber network structure may include a step of molding the hydrogel such that the outer surface of the hydrogel has a roughness of less than 30 microns. Molding the outer surface can also enable the manufacturer to form the outer surface into any desired shape. For example, the shape can be concave, convex, saddle-shaped, etc. Any desired shape (and smoothness) can be formed, for example, by molding and / or polishing.

[0031]

[0034] In any of these methods, the step of fixing the (e.g., dry) nanofiber network structure can include clamping and / or cementing the lyophilized nanofiber network structure. As described above, any of these devices and methods may use a dry nanofiber network structure including a cellulose nanofiber network structure. The dry nanofiber network structure may include at least one of electrospun polymer nanofibers, polyvinyl alcohol (PVA) nanofibers, aramid nanofibers, aramid-PVA nanofibers, wet-spun silk protein nanofibers, chemically cross-linked cellulose nanofibers, or polycaprolactone (PCL) fibers.

[0032]

[0035] Any of the methods described herein can include a step of rehydrating the nanofiber network structure. After fixing the nanofiber network structure to the implant surface, it includes a step of rehydrating it.

[0033]

[0036] Any of these methods can include a step of mineralizing at least a portion of the nanofiber network structure adjacent to the surface.

[0037] Described herein is an implant for knee joint surface reconstruction or partial knee joint surface reconstruction. For example, the upper support surface of the implant may include a hydrogel having a water content of at least 20 wt%, and the hydrogel is attached to the metal base with a shear strength exceeding 0.2 MPa.

[0034]

[0038] Any of the methods described herein may include a step of mechanically polishing the outer surface of a hydrogel (e.g., a cellulose-reinforced hydrogel) to a roughness of less than 50 microns (e.g., less than 50 microns, less than 40 microns, less than 30 microns, less than 25 microns, less than 20 microns, less than 15 microns, less than 10 microns, etc.). The step of mechanically polishing may include a step of scraping off the hydrogel adhering to the surface with fine sandpaper or the like as described herein.

[0035]

[0039] Any of these methods may include a step of rehydrating the nanofiber network structure. The nanofiber network structure may be rehydrated before impregnating with the hydrogel, or the impregnation may rehydrate the nanofiber network structure.

[0036]

[0040] For example, described herein is an implant comprising an implant body having a top hearing surface, an anchoring base (which may extend from the back of the top hearing surface), a cellulose-reinforced hydrogel, a cross-linked cellulose nanofiber network structure fixed on the top hearing surface of the implant body; and a interstitial hydrogel portion within the interstitial region of the cross-linked cellulose nanofiber network structure, the interstitial hydrogel portion having a crystallinity of 20% or more, the implant comprising the cellulose-reinforced hydrogel. The interstitial hydrogel portion may be polyvinyl alcohol (PVA). The cellulose-reinforced hydrogel may contain at least 20% by weight of water. The cellulose-reinforced hydrogel may have a tensile strength of more than 40 MPa. The cross-linked cellulose nanofiber network structure may be chemically cross-linked. The cross-linked cellulose nanofiber network structure may include bacterial cellulose (BC). The cellulose-reinforced hydrogel may have a compressive strength of more than 59 MPa. The cross-linked cellulose nanofiber network structure may be fixed on the top hearing surface by a clamp. In some examples, the cross-linked cellulose nanofiber network structure includes one or more sheets of bacterial cellulose (BC) held on the top hearing surface by a clamp fixed to a lip or rim of the top hearing surface. The clamp may be used to fix the cross-linked cellulose nanofiber network structure without the need for epoxy. Alternatively, any of these implants may include an adhesive.

[0037]

[0041] Also described herein is a method of forming an implant having a cellulose-reinforced hydrogel, the method comprising: attaching a cross-linked cellulose nanofiber network to an upper support surface of the implant; infiltrating a hydrogel component into the interstitial region of the cross-linked cellulose nanofiber network to form a cellulose-reinforced hydrogel; and annealing the cellulose-reinforced hydrogel such that the crystalline material of the hydrogel component has a crystallinity of 20% or more. The hydrogel component may include polyvinyl alcohol (PVA). The step of annealing the cellulose-reinforced hydrogel may include heating the cellulose-reinforced hydrogel. For example, the step of annealing the cellulose-reinforced hydrogel may include heating the cellulose-reinforced hydrogel to reduce the water content of the cellulose-reinforced hydrogel. In some examples, the cellulose-reinforced hydrogel is heated to a temperature in the range of 90 to 140 °C. The step of annealing the cellulose-reinforced hydrogel may include rehydrating the cellulose-reinforced hydrogel. The step of rehydrating the cellulose-reinforced hydrogel may include increasing the water content of the cellulose-reinforced hydrogel to at least 20 wt%. Any of these methods may include removing excess hydrogel component from the surface of the cross-linked cellulose nanofiber network. For example, the excess hydrogel component may be removed manually or by shaping the cellulose-reinforced hydrogel. In any of these examples, the cross-linked cellulose nanofiber network may include bacterial cellulose (BC). In some examples, the step of attaching the cross-linked cellulose nanofiber network to the upper support surface includes clamping the cross-linked cellulose nanofibers at the periphery of the upper support surface.

[0038]

[0042] Also described herein is an implant for knee joint surface reconstruction, comprising a cellulose nanofiber network structure and a hydrogel component impregnated in the cellulose nanofiber network structure, the hydrogel component having a crystallinity of 20% or more, and an upper support surface comprising a cellulose-reinforced hydrogel. The hydrogel component may include polyvinyl alcohol (PVA). The cellulose-reinforced hydrogel may contain at least 20% by weight of water. The cellulose-reinforced hydrogel may have a tensile strength of more than 40 MPa. The cellulose-reinforced hydrogel may be attached to the metal base of the upper support surface with a shear strength of more than 0.2 MPa. The upper support surface may have a coefficient of friction that is not statistically greater than the coefficient of friction (COF) of cartilage.

[0039]

[0043] Generally, the methods and apparatuses described herein may use any of the methods, apparatuses, and compositions described in International Patent Application No. PCT / US2021 / 040031, entitled "NANOFIBER REINFORCEMENT OF ATTACHED HYDROGELS", filed on July 1, 2021, the entire contents of which are incorporated herein by reference.

[0040]

[0044] All methods and apparatuses described herein may be used in any combination to achieve the benefits contemplated herein and described herein.

[0045] A better understanding of the features and advantages of the methods and apparatuses described herein will be obtained by reference to the following detailed description, which illustrates exemplary embodiments, and the accompanying drawings.

Brief Description of the Drawings

[0041]

Figure 1A

[0046] Figure of an exemplary process for attaching a hydrogel to a porous substrate by the Nanofiber-Enhanced STicking (NEST) method. In this example, a nanofiber sheet (e.g., bacterial cellulose) is attached to a surface (e.g., a porous substrate such as porous titanium) with an adhesive (e.g., α-TCP cement), and then the hydrogel component is infiltrated into the nanofiber sheet.

Figure 1B

[0047] Figure showing an example of a hydrogel bonded to a titanium plug.

Figure 1C

[0048] Figure showing a SEM image of the surface of an exemplary freeze-dried bacterial cellulose sheet.

Figure 2A

[0049] Figure 2A shows various hydrogel samples before and after annealing and rehydration. It illustrates the effect on the morphological structure of different hydrogel compositions.

Figure 2B

Figure 2C

Figure 2D

Figure 3A

[0050] Figures 3A - 3E are graphs showing the effect of annealing on the mechanical properties of various hydrogel compositions. Figure 3A is a graph illustrating the tensile stress - strain curve for PVA hydrogels annealed at different temperatures.

Figure 3B

Figure 3C

Figure 3D

Figure 3E

Figure 4A

[0051] Figures 4A to 4E are graphs illustrating the effect of annealing on the mechanical properties of BC-PVA hydrogel. Figure 4A is a graph showing the tensile stress-strain curve of BC-PVA hydrogel annealed at different temperatures.

Figure 4B

Figure 4C

Figure 4D

Figure 4E

Figure 5A

[0052] Figures 5A and 5B are graphs showing the effect of PAMPS on annealed BC-PVA hydrogels. Figure 5A is a graph showing the tensile strength, tensile modulus, and solid content weight fraction of BC-PVA-PAMPS hydrogels prepared using solutions containing different concentrations of AMPS monomers, and the BC-PVA samples were annealed prior to the penetration of AMPS. Figure 5B is a graph showing the compressive strength and modulus of the BC-PVA-PAMPS hydrogels.

Figure 5B

[0052] Figures 5A and 5B are graphs showing the effect of PAMPS on annealed BC-PVA hydrogels. Figure 5A is a graph showing the tensile strength, tensile modulus, and solid content weight fraction of BC-PVA-PAMPS hydrogels prepared using solutions containing different concentrations of AMPS monomers, and the BC-PVA samples were annealed prior to the penetration of AMPS. Figure 5B is a graph showing the compressive strength and modulus of the BC-PVA-PAMPS hydrogels.

Figure 6A

[0053] Figures 6A to 6D are diagrams showing various aspects of measuring the wear and coefficient of friction (COF) of various hydrogel compositions and comparing such measured values with cartilage. Figure 6A is a diagram schematically showing a method of measuring the wear and COF of a hydrogel against cartilage.

Figure 6B

Figure 6C

Figure 6D

Figure 7A

[0054] Figures 7A-7C are diagrams showing various aspects of measuring the wear of various hydrogel compositions and comparing such measured values with cartilage. Figure 7A is a diagram schematically showing a method of measuring the wear of cartilage against a hydrogel.

Figure 7B

Figure 7C

Figure 8A

[0055] Figures 8A-8D are diagrams illustrating various aspects of the shear strength tests of various hydrogel compositions and cartilage. Figure 8A is a diagram showing the results of a shear test of porcine cartilage and hydrogel fixed to a metal pin using an adhesive and a shape memory alloy clamp.

Figure 8B

Figure 8C

Figure 8D

Figure 9

[0056] Figure 9A is a front view of an annealed BC-PVA hydrogel incorporated into an implant for partial knee joint surface reconstruction before being subjected to mechanical stress. Figure 9B is a side view of an annealed BC-PVA hydrogel incorporated into an implant for partial knee joint surface reconstruction before being subjected to mechanical stress. Figure 9C is a front view of an annealed BC-PVA hydrogel incorporated into an implant for partial knee joint surface reconstruction after being subjected to mechanical stress. Figure 9D is a side view of an annealed BC-PVA hydrogel incorporated into an implant for partial knee joint surface reconstruction after being subjected to mechanical stress. Figure 9E is a front view of an annealed BC-PVA hydrogel incorporated into an implant for partial knee joint surface reconstruction after being subjected to mechanical stress. Figure 9F is a side view of an annealed BC-PVA hydrogel incorporated into an implant for partial knee joint surface reconstruction after being subjected to mechanical stress.

Figure 10A

[0057] Figure 10A schematically shows an example of an implant comprising a hydrogel attached (e.g., forming a surface) as described herein.

Figure 10B

Figure 11

[0058] It is a diagram showing an exemplary method of forming and attaching a hydrogel surface.

Figure 12A

[0059] Figure 12A is an image showing an example of a method of attaching a hydrogel to a metal plug, including using a clamp (e.g., a shape memory alloy clamp).

Figure 12B

[0060] FIG. showing an example of a jig that can be used to align and form the materials described herein as described herein (e.g., aligning BCs including rods, cut BCs, and ring clamps). FIG. 12B shows a perspective view of the jig.

Figure 12C

Figure 12D

[0061] FIG. showing an image of an exemplary sheet of bacterial cellulose (BC) cut (e.g., having legs or crenellations) for wrapping the end of a support surface (e.g., a metal rod, head, etc.) from above.

Figure 12E

[0062] FIG. showing the dimensions of diameter D, length L, and width W of an example of the sheet.

Figure 13

[0063] FIG. showing an example of a test jig that can be used for testing the shear of cartilage from bone and / or the shear of hydrogel material from a test rod.

Figure 14

[0064] FIG. 14A shows an example of a process for attaching a BC-PVA-PAMPS hydrogel to a titanium implant for the treatment of osteochondral defects.

[0065] FIGS. 14B and 14C show views of the repair of the implant.

Figure 15

[0066] Figures 15A to 15D are diagrams showing hydrogel samples before and after annealing and rehydration. All samples were annealed at 90 °C for 25 hours and rehydrated in PBS solution at 23 °C for 24 hours. Figure 15A is a diagram showing a sample of BC without PVA. Figure 15B is a diagram showing a sample of BC annealed in a 10 wt.% PVA solution. Figure 15C is a diagram showing a sample of 40 wt.% PVA. Figure 15D is a diagram showing a BC sample infiltrated with 40 wt.% PVA at 120 °C for 24 hours in a hot water cylinder before annealing and rehydration.

Figure 16

[0067] Figure 16A is a diagram showing the FTIR spectra of BC-PVA, annealed BC-PVA and annealed PVA hydrogels. It shows the increase in hydrogen bonding that occurs during annealing.

[0068] Figure 16B is a diagram showing an enlarged region highlighting the shift of the hydroxyl peak.

Figure 17A

[0069] Figure 17A is a diagram showing the DSC thermogram of a frozen-thawed and annealed PVA hydrogel.

Figure 17B

Figure 17C

Figure 18

[0070] It is a diagram showing the DSC thermogram of an annealed BC-PVA hydrogel sample. It shows how the peaks were integrated.

Mode for Carrying Out the Invention

[0042]

[0071] Disclosed herein is a hydrogel composition for long-term cartilage repair. The hydrogel has a crystal structure that imparts to the hydrogel a tensile strength and a compressive strength equal to or exceeding that of cartilage. The hydrogel is incorporated into a nanofiber network structure (e.g., cellulose) to increase wear characteristics and / or facilitate adhesion to the implant body. The hydrogel has been found to withstand the high compressive and shear stresses associated with knee joint movement and is thus suitable for implementation in knee implants. The hydrogel can be characterized by one or more traits and properties such as crystal structure, tensile strength, compressive strength, water content, coefficient of friction (COF), and / or other traits and properties.

[0043]

[0072] The method of forming a hydrogel implant described herein can be used to produce a hydrogel-coated orthopedic implant having a surface that mimics or improves the mechanical and / or tribological properties of cartilage. A method of forming an implant having a hydrogel comprising a bacterial cellulose (BC) network structure infused with polyvinyl alcohol (PVA) and sodium poly(2-acrylamido-2-methyl-1-propanesulfonate) (PAMPS), referred to as BC-PVA-PAMPS hydrogel, has been previously described in International Patent Application No. PCT / US2021 / 040031, the entire content of which is incorporated herein by reference. In the preparation of BC-PVA-PAMPS hydrogel, a freeze-thaw method is used to gelate the PVA-water mixture after penetration into BC. This freeze-thaw gelation process was used to increase the strength of the PVA hydrogel and prevent the degradation of PVA in the subsequent PAMPS penetration step. The increase in strength upon freezing and thawing of PVA is due to the crystallization and phase separation of PVA chains.

[0044]

[0073] The methods described herein improve these existing methods by enhancing the mechanical properties of the reinforced hydrogel, thereby improving the performance of the implant even when subjected to high impact and shear forces. For example, the tensile and compressive strengths of the reinforced hydrogel can increase up to and beyond the tensile and compressive strengths of cartilage. The method includes a mechanical strengthening process that increases the crystallinity of PVA and reduces the water content with respect to the freeze-thaw process. The mechanical strengthening process may include steps of drying, annealing, and / or rehydrating the hydrogel. When implemented in the hydrogel, the crystal structure within the hydrogel can increase even though they are interlocked with each other within the fibrous network structure. In addition, the hydrogel can be substantially bubble-free and crack-free after undergoing the crystal reconfiguration process.

[0045]

[0074] Figures 1A - 1C show an example of an apparatus in which a hydrogel is bonded to an implant surface as described herein. The hydrogel can be connected to the implant surface by first attaching a layer of nanofiber material, such as cellulose (e.g., bacterial cellulose), to the implant base using an adhesive (e.g., cement). The nanofiber material may be dried (e.g., before attaching to the implant base). The attachment surface of the implant base may be porous, for example, to enhance adhesion. Next, the hydrogel component can be infiltrated into the nanofiber layer. In this way, the nanofiber portion can penetrate into the porous bacterial cellulose network structure and be fixed with an adhesive (e.g., cement) that can be fixed to the surface, generating a bond that interlocks with each other without water interference. When the hydrogel component penetrates into the nanofiber network structure, the reinforced hydrogel is processed (e.g., annealed), and the mechanical properties of the reinforced hydrogel can be enhanced by changing the crystal structure of the hydrogel component.

[0046]

[0075] For example, in FIG. 1A, the nanofiber portion is bacterial cellulose (BC) 101 and is applied to the surface of the prepared implant (shown here as a titanium base with pores) 103. Cement (e.g., any suitable medical or dental grade cement can be used) is applied to fix the dry bacterial cellulose to the implant surface. Thereafter, the hydrogel component can penetrate the nanofiber portion, resulting in a complete hydrogel 107 attached to the base 103 via the bacterial cellulose 101. Next, the reinforced hydrogel 107 undergoes a crystal reconstitution process to enhance its mechanical properties.

[0047]

[0076] FIG. 1B shows an example of a titanium implant (e.g., a plug) with a cellulose-reinforced hydrogel attached as described herein. In this example, the nanofiber portion of the hydrogel (e.g., BC) is bonded to the porous surface of the implant via an adhesive, and the hydrogel is linked to the nanofiber portion. Any suitable adhesive (e.g., cement) can be used to adhere the nanofiber portion of the hydrogel to the implant surface. In some variations, the cement is α-tricalcium phosphate (α-TCP), a hydroxyapatite-forming cement that can be used for hydrogel attachment due to its biocompatibility, osteoconductivity, and shear strength, which can exceed those of cyanoacrylate. In some cases, α-TCP associates with phosphoserine (PPS) to facilitate adhesion. In some cases, hydroxyapatite is reinforced with stainless steel powder (SSP) (e.g., having an average particle size of 150 μm) to prevent crack propagation. As described in more detail below, in some examples, no adhesive is used and the nanofiber portion is fixed to the support surface by mechanical means (such as clamps).

[0048]

[0077] As described herein, the nanofiber portion (e.g., BC) can be dried (e.g., freeze-dried) to increase adhesion to the nanofibers. FIG. 1C is a scanning electron microscope (SEM) image of the surface of an exemplary freeze-dried sheet of BC, showing that it consists of many nanoscale fibers that provide a large surface area for adhesion using an adhesive. In some examples, multiple freeze-thaw cycles can be performed to increase the tensile strength (when the hydrogel is injected therein) and / or increase the shear strength of the adhesion of the reinforcing hydrogel to the implant base.

[0049]

[0078] As described above, existing techniques for generating cartilage-equivalent hydrogels involve infiltrating polyvinyl alcohol (PVA) and poly(2-acrylamido-2-methyl-1-propanesulfonic acid sodium salt) (PAMPS) into a bacterial cellulose (BC) nanofiber network. This hydrogel exhibited a tensile strength of 22.6 MPa and a compressive strength of 20 MPa. In comparison, the reported ranges of tensile and compressive strength for human cartilage are 8.1 - 40 MPa and 14 - 59 MPa, respectively. Thus, there is room to improve the strength of the hydrogel such that it has a similar elastic modulus, coefficient of friction, and wear resistance to cartilage, while being at the higher end of the strength ranges reported for cartilage or exceeding cartilage in strength. Considering the higher tensile strength of annealed PVA compared to freeze-thawed PVA, tests were conducted to determine whether changing from a freeze-thaw to an annealing process could improve the mechanical strength of the BC-PVA-PAMPS hydrogel while maintaining appropriate control over hydrogel shape and defect content. Considering that the tensile strength of the BC-PVA-PAMPS hydrogel (22.6 MPa) was already similar to that of the PVA hydrogel produced by annealing (20 MPa), it was not clear that changing to the annealing process would result in further improvement in mechanical strength for the BC-reinforced hydrogel. Additionally, the presence of BC or PAMPS potentially interfered with the crystallization of PVA that occurs during the annealing process, thereby preventing the improvement in mechanical strength that results from crystallization. Also, it was not clear whether it would be possible to obtain high-quality, bubble-free, crack-free samples after annealing of BC-reinforced PVA. Obtaining samples that are as defect-free as possible may be necessary to maximize the mechanical strength of the hydrogel. Finally, it was not clear whether the lower water content of the annealed hydrogel would result in too high a COF and opposing surface wear.

[0050]

[0079] As demonstrated herein, reinforcement using annealed PVA BC resulted in a 3.2-fold improvement in tensile strength (from 15.6 to 50.5 MPa) and a 1.7-fold increase in compressive strength (from 56.7 to 95.4 MPa). The more highly crystallized BC-PVA hydrogel obtained from annealing is the first hydrogel with tensile and compressive strengths exceeding those of cartilage. Reinforcement of PVA with BC can essentially remove the deformations and air bubbles that can occur during annealing without reinforcement. When tested against cartilage, annealed BC-PVA wore the opposing cartilage surface to the same extent as cartilage and had a wear resistance three times greater than that of cartilage. The COF of BC-PVA against cartilage was equivalent to the COF of cartilage against cartilage. In contrast to the results using freeze-thawed BC-PVA, addition of PAMPS to annealed BC-PVA decreased the tensile strength of the hydrogel due to loss of crystallized PVA and increased water content. The improvement in tensile strength of annealed BC-PVA was 68% greater than the shear strength of cartilage against bone, enabling its attachment to the metal base. Due to its high strength, high wear resistance, and low COF, annealed BC-PVA is an excellent material for replacement of damaged cartilage.

[0051]

[0080] For various hydrogel compositions, the tests and measurements performed and described herein demonstrate how a particular hydrogel composition that has undergone one or more strengthening processes, such as an annealing process, can increase the strength of the hydrogel to the upper limit or beyond that of cartilage while other characteristics, such as the coefficient of friction, reach levels similar to those of cartilage.

[0052]

[0081] One possible drawback of annealing a hydrogel is that the hydrogel can develop bubbles and cracks, especially when the sample thickness increases or the water content increases. Figures 2A - 2D show various hydrogel samples before and after annealing and rehydration, illustrating the effect on the morphological structure of different hydrogel compositions. Specifically, these samples illustrate the effect of the hydrogel composition on the shape of the samples after drying, annealing, and rehydration. All samples were annealed at 90 °C for 25 hours and rehydrated in a phosphate buffered saline (PBS) solution at 23 °C for 24 hours. Figure 2A shows a sample of BC (without PVA), demonstrating how the BC sample wrinkled and folded at the edges after annealing and rehydration. Figure 2B shows a sample of BC annealed in a 10 wt% PVA solution. As shown, the BC annealed in a 10 wt% PVA solution also demonstrates substantial deformation. The PVA layer formed on top of the BC after annealing contains numerous bubbles and easily peels off from the BC film. The bubbles may be the result of water evaporation. Figure 2D shows 40 wt% PVA samples before and after annealing, demonstrating that such hydrogels also formed numerous bubbles and deformed during the annealing process.

[0053]

[0082] Figure 2E shows a BC sample that was infiltrated with 40 wt% PVA at 120 °C for 24 hours in a hot water bomb before annealing and rehydration. As shown, the reinforcement of 40 wt% PVA with BC enabled the hydrogel to either hardly deform or not deform at all after annealing and retain its shape. The lack of such deformation may be attributed to the higher solid content and tensile modulus of BC-reinforced PVA. That is, the nanoscale network structure of the BC layer is thought to suppress the formation of large bubbles seen in the 40 wt% PVA sample (Figure 2C). A comparison of Figure 2B and Figure 2D shows that the method of infiltrating a high concentration of PVA into BC in a hot water bomb and subsequently removing the excess PVA from the BC surface results in a more uniform hydrogel than when the BC sample is placed in a more dilute PVA solution that is concentrated via drying. These results demonstrate that, unlike in the case of BC alone, PVA alone, or a combination of BC and 10 wt% PVA, BC infiltrated with 40 wt% PVA can retain its shape and remain relatively free of bubbles and other defects after annealing.

[0054]

[0083] Figures 3A - 3E illustrate the effect of annealing on the mechanical properties of various hydrogel compositions and analyze the effect of annealing on PVA hydrogels as a reference point. PVA was fully hydrolyzed and had a molecular weight of 145,000 g mol-1. A 40 wt.% PVA solution was dried at 90 °C for 24 hours, annealed at 90 °C, 120 °C, or 140 °C, and then placed in 0.15 M PBS solution for 24 hours for rehydration. PVA samples that had undergone freeze - thaw cycles were tested for comparison. Figures 3A and 3B show that annealing the hydrogel dramatically increased the tensile and compressive strengths for samples that had undergone freeze - thaw cycles. Figures 3C and 3D show that annealing increased the tensile strength by 60 - fold (from 0.26 to 15.6 MPa) and the compressive strength by 9 - fold (from 14.8 to 140.8 MPa) for the freeze - thaw process. The increase in annealing temperature from 90 °C to 140 °C resulted in an increase in tensile strength and elastic modulus. The increase in strength and elastic modulus was due to an increase in the crystallinity and solid content of the hydrogel after annealing. Figure 2E demonstrated that the crystallinity and solid content of the annealed PVA hydrogel were much greater than those of the freeze - thawed PVA hydrogel. For example, a PVA hydrogel prepared via the freeze - thaw process had a total solid content of 0.09 and a PVA crystallinity of 0.21, while a PVA hydrogel prepared via annealing at 90 °C had a total solid content of 0.42 and a PVA crystallinity of 0.58. The crystallites formed during annealing redistribute the applied stress and act as tough cross - links that prevent crack propagation, strengthening other amorphous polymer hydrogels. The crystallites also increase the solid content and strength of the hydrogel by reducing the amount of water taken up by PVA when immersed in 0.15 M PBS after annealing.

[0055]

[0084] Figures 4A - 4E illustrate the effect of annealing on the BC - PVA hydrogel. Referring to Figures 3A - 3E, the same annealing process as described above for the PVA hydrogel was applied to the BC - PVA hydrogel. Similar to the PVA hydrogel, the BC - PVA hydrogel was dried at 90 °C for 24 hours, annealed at 90 °C, 120 °C or 140 °C, and then placed in 0.15 M PBS solution for 24 hours for rehydration. Figures 4A and 4C show that the tensile strength of the annealed BC - PVA hydrogel reached 50.4 MPa, a 4.6 - fold increase compared to BC - PVA after freeze - thaw cycles, and a 3.2 - fold increase compared to annealed PVA not reinforced with BC. Figures 4B and 4D show that the compressive strength increased from 55.32 MPa to 95.35 MPa after annealing. Similar to the PVA hydrogel, this dramatic increase in strength can be attributed to the increase in crystallinity and solid content after annealing. Figure 3E shows that the crystallinity of the BC - PVA hydrogel increased from 0.07 after freeze - thaw cycles to 0.4 after annealing. The solid weight fraction of the BC - PVA hydrogel increased from 0.11 after freeze - thaw cycles to 0.53 after annealing. These results indicate that PVA can still form crystallites within the nanofiber BC network structure, and these crystallites increase the solid content and strength of the hydrogel.

[0056]

[0085] Figures 5A and 5B illustrate the effect of PAMPS on annealed BC-PVA hydrogels. As described above, it has been previously found that the incorporation of PAMPS into BC-PVA hydrogels made by freeze-thaw cycles resulted in an increase in the tensile and compressive strengths of the hydrogels. Therefore, PAMPS was incorporated into annealed BC-PVA hydrogels to determine the effect of PAMPS addition. As shown in Figure 5A, the addition of PAMPS to annealed BC-PVA hydrogels resulted in a decrease in the solids content from 0.53 to 0.37 relative to BC-PVA alone. Differential scanning calorimetry (DSC) thermograms showed that the peak from the melt-crystalline PVA disappeared after the addition of 10 wt% PAMPS, indicating that the addition of PAMPS disrupted the PVA crystallites formed during the annealing process. The decrease in solids content and loss of crystallinity upon addition of PAMPS led to a decrease in the tensile strength of the hydrogel (from 48.9 MPa to 20.8 MPa), a decrease in the tensile modulus (from 444.8 MPa to 150.5 MPa), and a decrease in the compressive strength (from 98.1 MPa to 56.0 MPa in Figure 5B). The increase in water content and loss of strength of the hydrogel are thought to be due to the fact that PAMPS is a negatively charged polymer and this negative charge creates an osmotic pressure that causes the hydrogel to swell with water.

[0057]

[0086] Figures 6A - 6D illustrate various aspects of measuring the wear and coefficient of friction (COF) of different hydrogel compositions and comparing such measurements to cartilage. Figure 6A is a schematic diagram of an example of a method for performing a wear test. To test the wear of the hydrogel in fetal bovine serum (FBS), the pin-on-disk configuration described in Example 11 below was used. Porcine cartilage plugs were placed under a pressure of 1 MPa and rotated at 319 revolutions per minute (maximum linear velocity was 100 mm s -1It was rotated 106 times against the hydrogel surface at a speed of (). The wear resistance of potential substitutes for cartilage exceeds that of cartilage, ensures durability, and minimizes the generation of wear powder that may potentially cause harmful biological reactions. The wear resistance of BC-PVA-PAMPS hydrogel is equivalent to that of cartilage and has been shown to be superior to PVA or PVA-PAMPS when tested against a stainless steel pin. These hydrogels were prepared by applying freeze-thaw cycles to crystallize PVA. Figures 6B to 6D compare the wear resistance of PVA-based hydrogels (PVA, BC-PVA, and BC-PVA-PAMPS) dried and annealed at 90 °C with the wear resistance of porcine cartilage against porcine cartilage plugs in FBS.

[0058]

[0087] Figure 6B shows a cross-sectional Micro-CT image of the hydrogel obtained at the center of the wear track to measure the maximum wear depth. Figure 6C compares the wear depth of the hydrogel with that of the cartilage. The wear depth of the BC-PVA hydrogel with 0% AMPS was 70.1 μm. The addition of 20% AMPS reduced the average wear depth to 65.9 μm, but the difference between the AMPS samples at 0% and 20% was not statistically significant. This comparison illustrates that the negative charge and higher water content resulting from incorporating PAMPS into the annealed BC-PVA hydrogel do not significantly improve wear resistance. Both of these values were one-third of the wear depth of the cartilage sample, which was 227.8 μm. The wear depth of annealed and rehydrated PVA was 301.0 μm, four times the size of any BC-PVA sample. These results indicate that the presence of BC in the hydrogel can dramatically improve the wear resistance of the annealed PVA hydrogel to be superior to that of cartilage.

[0059]

[0088] As shown in Figure 6D, the COF was recorded during the wear test. The cartilage maintained a constant COF of 0.020 during the test. The COF of BC-PVA decreased from 0.040 to 0.021 during the test. The BC-PVA hydrogel with 20% AMPS had a similar COF as the case without AMPS. In contrast, the COF of PVA increased dramatically from 0.033 to 0.135 during the test. Previous studies have similarly demonstrated that the COF of PVA against cartilage increases over time, while the COF of cartilage against cartilage remains constant. The increase in COF regarding the PVA-cartilage interface was attributed to the damaged PVA moving onto the cartilage surface, which in turn decreased the ability of the cartilage surface to maintain the lubricating water layer. Incorporation of BC into PVA clearly inhibits damage to the hydrogel and enables it to maintain a low coefficient of friction similar to that of cartilage during the wear test. The presence of AMPS in the hydrogel is not considered necessary to maintain a low COF and high wear resistance.

[0060]

[0089] The material used for cartilage replacement on one side of the joint, i.e., on the femoral condyle, should not cause wear of the cartilage on the opposing surface, i.e., on the tibial plateau. Traditional orthopedic materials such as cobalt-chrome and ultra-high molecular weight polyethylene are known to damage the opposing cartilage surface to a greater extent than hydrogels due to their higher COF and hardness of traditional orthopedic materials. To evaluate the wear on cartilage caused by BC-PVA and BC-PVA-PAMPS hydrogels, hydrogel plugs were generated for the wear test (see Example 7 below). The hydrogel plugs were pressed against the cartilage samples at a pressure of 1 MPa (see Figure 7A) and rotated 106 times at a speed of 319 rotations per minute (the maximum linear velocity at the circumference of the pin was 100 mm / s). -1 and rotated 106 times at a speed of 319 rotations per minute (the maximum linear velocity at the circumference of the pin was 100 mm / s).

[0061]

[0090] FIG. 7B shows a cross-sectional Micro-CT image of a cartilage sample obtained at the center of the wear scar to measure the maximum wear depth. FIG. 7C compares the wear depths of cartilage caused by the hydrogel or cartilage. The wear caused by BC-PVA on cartilage (247 ± 16 μm) was not significantly different from the wear caused by cartilage on cartilage (228 ± 12 μm). The addition of PAMPS to BC-PVA reduced the wear on the opposing cartilage surface to 81 ± 27 μm, which was significantly lower than the wear of cartilage against cartilage.

[0062]

[0091] FIGS. 8A - 8D show various aspects of the shear strength tests of different hydrogel compositions and cartilage. A hypothesis was made that an increase in the tensile strength of the hydrogel would also increase the shear strength. For use as a cartilage replacement material, the synthetic hydrogel must have the same shear strength as the joint between cartilage and bone and be fixed to the defect site. One way to achieve this is to have a hydrogel that adheres directly to bone or cartilage with sufficient strength. Alternatively, the hydrogel can be attached to a metal base such as titanium that has the ability to integrate with bone. As described above, the ability to attach a BC-PVA-PAMPS hydrogel to a metal base with a shear strength equivalent to that of the cartilage-bone interface is described in PCT / US2021 / 040031. Previous tests have shown that the strength of hydrogel attachment is limited by the tensile force required to break the hydrogel that curves over the end of the metal base. Therefore, an increase in the tensile strength of the hydrogel should in turn increase the shear strength with which the hydrogel is attached to the metal base.

[0063]

[0092] The arrangements used for the shear tests are described in Examples 7 and 12 below. FIGS. 8A - 8D show the results of shear tests of plugs of porcine cartilage against bone extracted from porcine knees, tests of BC - PVA - PAMPS hydrogels made by previous freeze - thaw processes, and tests of BC - PVA hydrogels annealed at 90° C. and then rehydrated. Both hydrogels are attached to stainless - steel rods with a combination of RelyX Ultimate cement and shape - memory alloy rings. The shear strength of BC - PVA of 1.98 is significantly greater than the shear strength of porcine cartilage (the p - value from one - way ANOVA is less than 0.05). The mean value of the shear strength for BC - PVA is also 40% greater than that of BC - PVA - PAMPS, but due to measurement error, the difference in these values is not statistically significant. A comparison of the samples after failure shows that porcine cartilage was completely sheared from the underlying bone, while both BC - PVA - PAMPS (made by the freeze - thaw process) and BC - PVA (made by annealing at 90° C. followed by rehydration) were fractured on one side of the cylindrical sample but remained attached. These results show that the shear strength of adhesion for annealed BC - PVA is greater than the shear strength of porcine cartilage.

[0064]

[0093] The above - mentioned are the tensile strength, compressive strength, and shear strength of BC - reinforced hydrogels attached to metal pins having a diameter of 5.2 mm. This size is convenient for testing, but such a diameter is too small to serve as an implant for partial knee - surface reconstruction. In addition, the sample lacked the curvature necessary to mimic the natural curvature of the femoral condyle. FIGS. 9A - 9F show the application of annealed BC - PVA hydrogels to implants for partial knee - surface reconstruction, demonstrating the ability of the hydrogels to adhere to a metal base having the size and shape of a representative implant for partial knee - surface reconstruction.

[0065]

[0094] Figures 9A and 9B show the implant before mechanical testing. The implant sample has a diameter of 20 mm and a radius of curvature of 20 mm. The 20 mm implant diameter is a typical size used for osteochondral allografts, and the 20 mm radius of curvature is within the typical range of curvature for the femoral condyle. The peak force applied to the knee during jogging was measured to be 5551 N for a body weight of 100 kg. The tibiofemoral contact area was measured to be 1500 mm 2 at 3100 N. Based on the assumption that the contact area does not increase for higher forces, the peak stress applied to the knee contact area during jogging is 3.7 MPa. Figures 9C and 9D show the implant after being subjected to a compressive stress of 16 MPa, which is 4.3 times the magnitude of the peak physiological force applied to the femoral condyle. After this test, there were no signs of cracking or damage on the hydrogel surface. This test represents that implants produced using annealed BC-PVA hydrogel can withstand the compressive forces in the knee without cracking. The peak anterior shear force in the knee for walking was measured to be 294 N, which is 30% of body weight, or for an individual of 100 kg. This is the highest shear force measured in the knee for any of the investigated daily activities. The tibiofemoral contact area was measured to be 1500 mm 2 at 3100 N, which is approximately equivalent to the peak vertical force during walking for an individual with a body weight of 100 kg. Since the peak vertical force coincides with the peak shear force, using this tibiofemoral contact area, the peak shear stress experienced by the cartilage during walking can be calculated to be 0.2 MPa (294 N ÷ 1500 mm 2 ). The shear test on the implant represents that no failure occurred until a stress of 0.9 MPa was applied. Since the implant can withstand a shear load that is 4.5 times the magnitude of the load experienced by the cartilage in the knee, this result shows that the annealed BC-PVA hydrogel and the attachment method have sufficient strength for the production of implants for partial knee joint surface reconstruction.

[0066]

[0095] Implants for partial knee resurfacing can be relatively large and can be curved to mimic the natural curvature of the femoral condyles. FIGS. 14A - 14C show images of an implant with a diameter of 20 mm and a radius of curvature of 20 mm. A 20 - mm implant diameter is typical of those used for osteochondral allografts, and a 20 - mm radius of curvature is within the typical range of curvature for the femoral condyles. In this example, five pieces of BC were cut into an octagonal shape with eight legs to allow for folding back at the implant end at BC. To improve integration with the bone, a commercially available 0.25 - mm - thick coating of pure titanium was applied by a plasma spraying process under the implant's stem and base. FIGS. 14B and 14C show an example of how such an implant is used for knee resurfacing. FIG. 14B shows an example of a cartilage defect. The surgery involves drilling holes in the defect site that are complementary to the shape of the hydrogel - capped implant. Next, the hydrogel - capped implant is pushed into the holes to replace the damaged cartilage.

[0067]

[0096] The methods and apparatus described herein are used to reinforce annealed PVA hydrogels with BC, and for the first time can provide hydrogels having a compressive strength and a tensile strength greater than cartilage. Annealing increased the tensile strength of BC-PVA fivefold and the compressive strength 1.8-fold with respect to the freeze-thaw process due to the greater crystallization and lower water content achieved by annealing. The reinforcement of PVA with BC reduced the wear of the hydrogel to one quarter with respect to PVA alone and to one third with respect to cartilage. The annealed BC-PVA hydrogel caused a minimal opposing surface wear, similar to that caused by the cartilage itself. The attachment of BC to a metal plug via an adhesive and / or clamp, followed by the penetration and annealing of PVA, enabled the attachment of the BC-PVA hydrogel to the metal support with a shear strength greater than the attachment of the cartilage to bone. These advancements in hydrogel strength and attachment can enable the production of implants with hydrogel surfaces and titanium supports that can enable durable joint surface reconstruction in the arthroplasty of damaged cartilage.

[0068]

[0097] As used herein, an implant can have a structure suitable for implanting into the body. In some (non-limiting) examples, the implant can have a shape that allows it to be implanted into bone with a hydrogel attached to the outer-facing surface. For example, FIGS. 10A and 10B show examples of implants with a hydrogel attached as described herein. In FIG. 10A, the implant includes a base 1001 (e.g., a titanium base) having an elongated pin shape that can be, for example, 2 mm × 7 mm (tapering from about 3 mm at the end to about 1.5 mm). The base can include one or more channels, openings, passageways, etc. for bone growth. The implant can also include an upper portion 1005 that can be curved (e.g., single-curved or double-curved). For example, the surface can be curved with a radius of curvature of about 17 mm (single-curved) or about 19 mm × 12 mm (double-curved). In FIG. 10A, the upper portion has a diameter 1007 of approximately 7 mm. The outer surface of the implant can have a thickness 1009 of approximately 1 mm or more and can be about 70% or more porous. The hydrogel can be attached to the upper surface. The hydrogel in this example is a triple-network structure hydrogel of BC-PVA-PAMPS, where BC is cemented to the porous upper portion, while PVA-PAMPS is impregnated within BC. FIG. 10B shows an implant similar to that shown in FIG. 10A, with the hydrogel attached (e.g., via cementing the nanofiber portion of the hydrogel to the porous surface of the implant as shown). The implant in FIG. 10B is titanium.

[0069]

[0098] As described above, any of these implant surfaces can include a porous structure. The porosity of the implant surface can be, for example, between 10% porous and 90% porous, such as between 30% porous and 90% porous, between 55% porous and 95% porous, between 65% porous and 85% porous, etc. The depth of the pores can also vary. For example, the surface can be porous to a depth between 0.1 mm and 5 mm, between 0.2 mm and 3 mm, between 0.5 mm and 2 mm (e.g., 0.2 mm or more, 0.3 mm or more, 0.5 mm or more, 0.75 mm or more, 1 mm or more, 1.5 mm or more, etc.).

[0070]

[0099] As described above, any suitable nanofiber network structure including, but not limited to, nanofiber bacterial cellulose may be used. Other nanofiber network structures may include electrospun polymer nanofibers such as polyvinyl alcohol (PVA) nanofibers, aramid nanofibers (e.g., aramid-PVA nanofibers), wet-spun silk protein nanofibers, chemically crosslinked cellulose nanofibers, or polycaprolactone fibers (e.g., 3D woven PCL fibers). Additionally, any suitable double network structure hydrogel including, but not limited to, PVA and PAMPS may be used. For example, other hydrogel-forming polymers may be poly(N,N'-dimethylacrylamide) (PDMAAm), a copolymer of 1-vinylimidazole and methacrylic acid, a double network structure hydrogel based on an amphiphilic triblock copolymer, a polyampholyte hydrogel, a PVA-tannic acid hydrogel, a poly(N-acryloyl)glycinamide hydrogel, a polyacrylic acid-acrylamide-C18 hydrogel, guanine-boric acid reinforced PDMAAm, a polyelectrolyte hydrogel, a poly(acrylonitrile-co-l-vinylimidazole) hydrogel (e.g., a mineralized poly(acrylonitrile-co-l-vinylimidazole) hydrogel), a polyacrylic acid-Fe3+-chitosan hydrogel, a poly(methacrylic acid) gel, a graphene oxide / xonotlite reinforced polyacrylamide (PAAm) gel, a poly(stearyl methacrylate)-polyacrylic acid gel, an annealed PVA-polyacrylic acid hydrogel, a supramolecular hydrogel from a multiurea-bonded segmented copolymer, a polyacrylonitrile-PAAm hydrogel, a microsilica reinforced DMA gel, an agar-polyhydroxyethyl methacrylate gel, a polyfacryloyloethyltrimethylammonium chloride hydrogel, a poly(3-(methylacryloylamino)propyl)trimethylammonium chloride hydrogel, a poly(sodium p-styrenesulfonate) hydrogel, a polyethylene glycol diacrylate hydrogel, a polyethylene glycol hydrogel, or a hydrogel composed of a combination of these polymers.

[0071]

[0100] The implants described herein can be formed from any suitable material including, but not limited to, titanium and stainless steel. For example, a hydrogel can be attached to an implant surface (e.g., a base including a porous base) formed from a stainless steel alloy, other titanium alloys, Co-Cr alloys, tantalum, gold, niobium, bone, Al oxide, Zr oxide, hydroxyapatite, tricalcium phosphate, sodium calcium phosphosilicate (bioactive glass), poly(methyl methacrylate), polyetheretherketone, polyethylene, polyamide, polyurethane, polytetrafluoroethylene, or other materials used in implant fabrication, as described herein.

[0072]

[0101] Any of the implants described herein can include a hydrogel having a surface that is substantially smooth and / or is shaped in a predetermined three-dimensional configuration such as, but not limited to, concave, convex, saddle-shaped, etc. For example, any of these devices (e.g., implants) can have a surface roughness of less than 30 microns. In some cases, the surface can be smoothly formed by molding. In some cases, the surface can be smoothly formed by polishing or sanding. For example, when an additional hydrogel component forms a network structure (e.g., a nanofiber-reinforced network structure), the hydrogel coating can optionally be finished by polishing; specifically, the surface can be sanded to a roughness of less than 30 microns. The polishing can be performed by sanding (e.g., using a fine-grit sanding surface such as 600, 400, 320 grid, etc.).

[0073]

[0102] Figure 11 is a flowchart showing an exemplary method of fabricating the device described herein. The surface to which the hydrogel is to be attached, for example, the implant surface, may optionally be prepared, 1101. For example, the surface may be made porous. In some examples, the porosity may be at least 0.5 mm deep (e.g., 1 mm or more deep). The porosity may be described as percent porosity (e.g., porous between 10% and 90%, between 20% and 90%, greater than 30%, greater than 40%, greater than 50%, etc.).

[0074]

[0103] Next, the nanofiber portion may be prepared for attachment to the surface, 1103. For example, the nanofiber portion is dried (e.g., freeze-dried). The nanofiber portion may be dry or substantially dry and applied to the attachment surface, 1105. Next, the nanofiber portion may be fixed to the surface, 1107. In some variations, an adhesive (e.g., cement) may be applied to the surface before the nanofiber portion is applied, and / or the adhesive may be applied to the nanofiber portion on the surface. In some variations, the adhesive may be applied to the nanofiber portion before attaching it to the surface. In some examples, no adhesive is used at all.

[0075]

[0104] If used, the adhesive is applied and may be dried at a drying temperature (e.g., room temperature, 30 degrees, etc.) (e.g., for a predetermined time, e.g., 1 hour, 2 hours, 3 hours, 6 hours, 12 hours, 24 hours, etc.). Once dried, the nanofiber portion cemented to the surface may optionally be rehydrated, for example, by the addition of an aqueous solution, 1109.

[0076]

[0105] The nanofiber portion may then be penetrated by other components of the hydrogel, which begins to impregnate within the nanofiber portion fixed to the surface, 1111. The other components may include one or more polymer components that can form a hydrogel and can crystallize during a subsequent annealing process. In some examples, the polymer component includes polyvinyl alcohol (PVA). In some examples, the polymer component includes only PVA.

[0077]

[0106] When the polymer hydrogel component is injected into the nanofiber portion, a mechanical strengthening process can be implemented to strengthen the hydrogel. The mechanical strengthening process can include drying, annealing, and rehydrating the hydrogel. The drying and / or annealing step can include heating the hydrogel to a predetermined temperature (e.g., in the range of 90°C to 140°C), followed by rehydrating (e.g., in a PBS solution). The resulting hydrogel can have an increased crystal structure. For example, an interstitial polymer hydrogel (e.g., PVA) can have a crystallinity of at least 20%. In some examples, the tensile strength of the resulting hydrogel is at least 40 MPa. In some examples, the compressive strength of the resulting hydrogel is at least 59 MPa. Additionally, the hydrogel can have a water content of about 20 wt% or more. Once the mechanical strengthening process is complete, the hydrogel surface may optionally be polished, 1115.

[0078]

[0107] In some examples, the devices described herein can form part of a surgical implant for treating defects such as osteochondral defects. For example, the surgical implant can include a surface covered with a hydrogel; this surface can act as an interface between one or more other body regions, including hard tissues such as bone and cartilage. Repair of cartilage lesions using a hydrogel can benefit from the long-term retention of the hydrogel at the defect site. Adhesion of the hydrogel to the base (substrate) of the bone, which enables integration with the bone, can allow for long-term retention of the hydrogel, but current methods of forming a bond with the hydrogel have a shear strength of less than one-tenth of the osteochondral junction. The devices and methods described herein can include bonding the hydrogel to a surface (e.g., a base) with a shear strength several times greater than the shear strength achieved heretofore.

[0079]

[0108] Figures 12A - 12C show an overview of an example of how a hydrogel can be attached to a metal base (e.g., the upper support surface). In this example, as shown in the example of Figure 12D, the lyophilized BC sheet was cut into an octagonal shape with 8 protrusions (e.g., "legs") that can be bent over the end of the implant. This cutting can remove excess BC that would otherwise fold onto the cylindrical side surface if not cut. Next, the cut BC piece was placed in a jig that facilitates centering and alignment of the BC piece, the metal rod, and the ring clamp. The metal rod was pushed down through the jig so that the ring pressed the BC piece onto the metal rod. This process of pressing the ring onto the BC and the rod could also be done manually. As shown in Figures 12B - 12C, the use of alignment features can help keep the small pieces consistently centered during assembly. Next, the sample can be clamped (starting at a temperature of 50°C) by heating it to 90°C in, for example, a furnace and initiating the clamp in a shape - memory alloy material pre - installed as described herein. Next, this component was heated with PVA in a hot - water cylinder at 120°C for 24 hours to allow the polymer to penetrate into the BC. Next, as described herein, this component was dried, annealed, and re - hydrated.

[0080]

[0109] The following are example methods for preparing and testing various hydrogel samples described herein.

Example

[0081]

[0110] Example 1: Fabrication of BC-PVA-PAMPS Hydrogel

[0111] The BC sheet was compressed to a thickness of 0.5 mm and placed in a hot water reactor together with a mixture of polyvinyl alcohol (PVA) (40 wt.%) and deionized water (60 wt.%). The hot water reactor was sealed and heated at 120 °C for 24 hours to allow PVA to diffuse into the pores of BC, forming a BC-PVA hydrogel. The BC-PVA hydrogel was taken out of the reactor at high temperatures (e.g., above 85 °C). The residual PVA solution was removed by rubbing the surface of the BC-PVA sample with a metal spatula. The sample was frozen at -78 °C for 30 minutes and thawed at room temperature to physically crosslink the PVA network. Next, the BC-PVA hydrogel was immersed in a solution of sodium 2-acrylamido-2-methylpropanesulfonate (AMPS) (30 wt.%), N,N’-methylenebisacrylamide (MBAA) (60 mM), 2-hydroxy-4’-(2-hydroxyethoxy)-2-methylpropiophenone (I2959) (50 mM) and potassium persulfate (KPS) (0.5 mg mL-1) for 24 hours. The hydrogel was cured on one side with a UV transilluminator for 15 minutes each and further cured in an oven at 60 °C for 8 hours to ensure it was flat and fully cured. The resulting BC-PVA-PAMPS hydrogel was stored in phosphate buffered saline (PBS) for at least 24 hours before further characterization.

[0082]

[0112] Example 2: Fabrication of Annealed BC

[0113] The BC sheet was compressed to a thickness of 0.5 mm. Next, the BC sheet was placed in a furnace at 90 °C for 24 hours before annealing at 90 °C for an additional 1 hour. The resulting annealed BC was cut into the desired shape and stored in 0.15 M PBS for at least 24 hours.

[0083]

[0114] Example 3: Fabrication of PVA Hydrogel

[0115] To fabricate PVA hydrogels, a slurry of PVA (40 wt.%) and DI water (60 wt.%) was mixed in a metal baking dish (diameter: 203.2 mm) and heated in an autoclave sterilizer at 120 °C for 20 minutes. To fabricate annealed PVA hydrogels, the obtained hydrogels were dried in an oven at 90 °C for 24 hours before further annealing at 90 °C, 120 °C or 140 °C for 1 hour. To fabricate freeze-thawed PVA hydrogels, the autoclaved hydrogels were frozen at -80 °C for 30 minutes and thawed at 23 °C for 30 minutes. Before testing, the obtained PVA hydrogels were cut into the desired shape and stored in 0.15 M PBS for at least 24 hours.

[0084]

[0116] Example 4: Fabrication of Annealed BC-40wt.%PVA Hydrogel

[0117] The BC sheet was compressed to a thickness of 0.5 mm and placed in a hot water reactor together with a mixture of PVA (40 wt.%) and DI water (60 wt.%). The hot water reactor was sealed and heated at 120 °C for 24 hours to diffuse PVA into the voids of BC and form BC-PVA hydrogels. The BC-PVA hydrogels were taken out of the reactor at high temperatures (e.g., 85 °C or higher). Note that the hot water reactor is pressurized with hot steam and there is a risk of burns. When opening the reactor, personal protective equipment including a laboratory coat, heat-resistant gloves and a full-coverage face shield must be used. The residual PVA solution was removed by rubbing the surface of the BC-PVA sample with a metal spatula. The samples were dried in an oven at 90 °C for 24 hours before further annealing at 90 °C, 120 °C or 140 °C for 1 hour. Before testing, the obtained annealed BC-PVA hydrogels were cut into the desired shape and stored in 0.15 M PBS for at least 24 hours.

[0085]

[0118] Example 5: Fabrication of Annealed BC-10wt.%PVA Hydrogel

[0119] The BC sheet was compressed to a thickness of 0.5 mm and placed in a baking dish (15.6 cm × 8.6 cm × 4.2 cm). Approximately 30 mL of a 10 wt.% PVA solution was added to the baking dish. The baking dish was placed in an oven at 90 °C for 24 hours and annealed at 90 °C for an additional 1 hour. The obtained annealed BC-PVA hydrogel was cut into the desired shape and stored in 0.15 M PBS for at least 24 hours.

[0086]

[0120] Example 6: Fabrication of Annealed BC-PVA-PAMPS Hydrogel

[0121] The BC sheet was compressed to a thickness of 0.5 mm and placed in a hot water reactor together with a mixture of PVA (40 wt.%) and DI water (60 wt.%). The hot water reactor was sealed and heated at 120 °C for 24 hours to diffuse PVA into the voids of BC and form a BC-PVA hydrogel. The BC-PVA hydrogel was taken out of the reactor at high temperature (above 85 °C). Note that the hot water reactor is pressurized with hot steam and there is a risk of burns. When opening the reactor, personal protective equipment including a lab coat, heat-resistant gloves and a full-coverage face shield must be used. The residual PVA solution was removed by rubbing the surface of the BC-PVA sample with a metal spatula. The sample was dried in an oven at 90 °C for 24 hours before annealing at 90 °C, 120 °C or 140 °C for an additional 1 hour. Next, the annealed BC-PVA hydrogel was immersed in a solution of AMPS (30 wt.%), MBAA (60 mM), I2959 (50 mM) and KPS (0.5 mg mL-1) for 24 hours. The hydrogel was cured on one side for 15 minutes each with an ultraviolet transilluminator (VWR International) and further cured in an oven at 60 °C for 8 hours to ensure it was flat and fully cured. The obtained annealed BC-PVA-PAMPS hydrogel was stored in PBS for at least 24 hours before further characterization.

[0087]

[0122] Example 7: Fabrication of Hydrogel on Stainless Steel Pins

[0123] The preparation of all hydrogel samples was started by cutting freeze-dried BC. The BC was cut into an octagonal shape with a diameter D and eight legs with a leg length L and a width W = 0.383D (see, for example, FIGS. 2A-2D). The samples were labeled BC-D-L after cutting. Eight star-shaped pieces (BC-D-L) were generated by MATLAB and loaded into Adobe Illustrator. In Adobe Illustrator, the stroke of the shape was changed to 0.0001 pt to ensure accurate cutting. The file was sent to a laser cutter (Epilog Fusion M2) using the print function, and the laser cutter was selected as the printer. Vector processing was used at 100% speed, 20% output, and 100% frequency. To cut the BC, a clean metal plate was placed on the bed of the laser cutter, and the freeze-dried BC was placed on that metal plate. Another metal plate was placed at the end of the BC to ensure that the BC stayed in place. The focus was adjusted, and the shape was cut by the machine. After cutting, the BC was collected and stored in a Petri dish for later use.

[0088]

[0124] To prepare the shear test specimens, six pieces of BC were adhered to stainless steel rods using one layer of cement and clamps. An overview of the assembly method is shown in FIGS. 12A - 12C and 13. The stainless steel test rods were machined to have an upper section with a diameter of 5.2 mm and a height of 2 mm, and a bottom section with a diameter of 6.75 mm and a height of 13 mm. Three pieces of BC - 6.5 - 2.25 and three pieces of BC - 6.5 - 2 were placed in an alignment jig. As an optional step, an adhesive (e.g., Scotchbond Universal Adhesive) may be applied to the layer of BC in contact with the rod and the upper surface of the rod. If used, the adhesive was allowed to cure for 20 seconds before further air - blowing for 5 seconds. In one example, next, approximately 0.15 g of RelyX Ultimate cement was applied to the same surface that was coated with Scotchbond Universal Adhesive. The rod was pushed into the BC layer and then into a shape - memory alloy ring clamp. The cement was cured for 1 hour. The specimen was heated in an oven at 175 °C for 10 minutes to shrink the clamp. Next, the specimen was immersed in DI water in a centrifuge tube for 1 hour before later use. Next, the specimens with BC on top were subjected to a specific hydrogel fabrication process.

[0089]

[0125] Compression test specimens were fabricated without using cement, using stainless steel rods and clamps. Wear test specimens were fabricated without using cement, using stainless steel rods with a diameter of 5.7 mm and a height of 38 mm, three pieces of BC - 6.5 - 2, and shape - memory alloy rings.

[0090]

[0126] Example 8: Uniaxial Tension Test and Compression Test

[0127] The monotonic tensile test was performed in the load frames of Instron 1321 (Instron, Norwood, MA, USA) and TestResources 830 (TestResources, Shakopee, MN, USA) at 0.25 mm per second -1It was run at a speed of

[0091]

[0128] The compression properties of all samples were measured using an axial torsion system (TestResources 830LE63). Cylindrical PVA samples were cut out from the hydrogel film with a steel punch having a diameter of 4 mm. The BC-PVA samples were attached to metal pins for compression testing to obtain sufficiently thick samples. The dimensions of the samples were measured with calipers before the test. The compression properties were measured at a strain rate of 0.05 seconds -1 . The ultimate compression strength was obtained as the maximum stress measured before fracture. The compression modulus was derived as the slope at a stress of 0.4 MPa of the stress-strain curve for comparison with previous studies on human cartilage.

[0092]

[0129] Example 9: Differential Scanning Calorimetry

[0130] Differential scanning calorimetry (DSC) was performed on the hydrogels to determine the crystallinity of PVA. The test was completed on a TA TGA550. In a typical experiment, approximately 5 mg of the hydrogel sample was placed in an aluminum pan under a nitrogen gas flow and heated from 25 °C to 300 °C at a rate of 10 °C / min. Typical thermograms for the PVA, BC-PVA, and BC-PVA-PAMPS hydrogels are shown in FIGS. 4A - 4D.

[0093]

[0131] The calculation of how much PVA was crystallized, i.e., the degree of crystallization, was adopted from Hassan et al. 22. After obtaining the DSC thermogram, the area under the melting peak over the range of 140 - 220 °C (shown in Figures 5A and 5B) was integrated to obtain a value with the unit of J·°C·s-1·g-1. This numerical value was then divided by the heating rate (0.17 °C·s -1 ) to obtain ΔH (J·g -1 ). Next, the crystallinity of PVA was calculated by dividing the ΔH for the sample by the heat quantity ΔH e = 138.6 J / g required to melt the 100% crystalline PVA sample and the weight fraction w PVA of PVA in the sample.

[0094]

Equation

[0095]

[0132] Here, χ PVA is the crystallinity of PVA.

[0133] Example 10: Measurement of Solid Weight Fraction

[0134] The weight of approximately 1 g of the hydrated hydrogel was measured before drying at 90 °C for 24 hours. Next, the weight of the dehydrated sample was measured. The weight after dehydration was divided by the weight before dehydration to determine the solid weight fraction of the hydrogel sample.

[0096]

[0135] Example 11: Wear Test

[0136] The abrasion resistance of the hydrogel and porcine cartilage samples was determined in a pin-on-disk configuration as shown in FIGS. 6A - 6D. The pin-on-disk method was used with an Anton Paar Rheometer (MR302) and a tribological accessory (SCF7). Cartilage samples were harvested from porcine femurs using an osteochondral autograft transfer system (Arthrex). The femurs were purchased from a local grocery store and frozen at -78°C before sample collection. Hydrogel samples were polished with #600, #800, #1000, #1200, #1500, #2000, #2500, and #3000 sandpapers and made smooth before testing. Hydrogel pins were fabricated by using the method described in Example 7 above. Disks of hydrogel or porcine cartilage with a 12.7 mm diameter were adhered to the sample holder with a cyanoacrylate adhesive (Gorilla Glue Company). The test parameters were as follows: 1,000,000 rotations; angular velocity: 319 rotations per minute (maximum linear velocity: 100 mm s-1); vertical force: 28.26 N (pressure: 1 MPa). A pressure of 1 MPa was applied to each sample for 5 minutes before the test was started. The test was conducted in FBS. FBS is often used during wear testing to mimic the lubrication provided by synovial fluid.

[0097]

[0137] After the wear test, the samples were rehydrated in FBS for 24 hours, and the gel was recovered from the applied pressure before measuring the wear depth with a high-resolution X-ray computed tomography (Micro-CT) scanner (Nikon XTH 225 St). 3D models of the reconstructed Micro-CT images were rendered with Avizo 9 Lite. To measure the wear depth, slices of the 3D model were obtained at the center of the wear track. The wear depth was measured from the image of the central slice using ImageJ.

[0098]

[0138] To calculate the COF, the total frictional force (F) was determined from the torque (T) and the radius (R) of the pin in the pin-on-disk configuration:

[0099]

Equation

[0100]

[0139] Next, the COF can be calculated by

[0101]

Number

[0102] Here, F

[0140] is the vertical force (28.26 N). The linear velocity (v) is N calculated by v = ωR v = ωR where ω is the angular velocity of the pin.

[0103]

[0141] In the formula, ω was the angular velocity of the pin.

[0142] The above results are particularly unexpectedly surprising. As described in this specification, annealed BC-PVA provides a material that maintains a low coefficient of friction (0.21), similar to cartilage (0.2), over more than 1 million wear cycles. In contrast, annealed PVA alone has a COF that increases from 0.033 to 0.135 over 1 million cycles. Particularly surprisingly, adding BC to PVA resulted in a long-term COF that is one-sixth that of PVA alone. This lower COF made possible by BC has significant implications for preventing wear of the opposing cartilage surface. Similarly, it is surprising that adding BC to annealed PVA reduces the wear depth to less than one-third that of annealed PVA alone. This greater wear resistance made possible by BC can also have significant implications for the long-term durability of the implant.

[0104]

[0143] Example 12: Shear Test

[0144] The shear tests were performed on an 830LE63-axis torsion testing machine equipped with a 45.36 kg (100 pound (lb)) load cell. Each test was performed in a customized shear test jig (see Figure 13). To shear the hydrogel sample on cartilage or metal, the sample was fixed in a cylindrical hole on the left side of the jig. The hole diameter was 6 mm for the but plug and 7 mm for the hydrogel sample. A spacer was added directly under the sample to accurately align the shear plane at the cartilage-bone or hydrogel-metal interface. The right side of the jig was machined to have a complementary semi-cylindrical shape and was used to extrude the hydrogel or cartilage from its substrate. The diameter of the right semi-cylindrical shape matched the diameter of the left semi-cylindrical shape (either 6 mm or 7 mm). To minimize cracking and peeling, rubber was deployed between the sample and the right shear jig and pressure was applied during the shear test. A crosshead displacement rate of 2 mm min-1 was used for all measurements.

[0105] Supplementary Materials

[0145] As described, in some of these examples, bacterial cellulose (BC) was purchased from Gia Gia Nguyen Co., Ltd. Poly(vinyl alcohol) (PVA) (fully hydrolyzed, molecular weight: 145,000 g mol -1 )), N,N’-methylenebisacrylamide (MBAA, 97.0%), 2-hydroxy-4’-(2-hydroxyethoxy)-2-methylpropiophenone (I2959), potassium persulfate (KPS) and sodium 2-acrylamido-2-methylpropanesulfonate (AMPS, 50 wt.% solution in water) were purchased from Sigma Aldrich. Phosphate buffered saline (PBS) was purchased from VWR International. Fetal bovine serum (FBS, from Canada, typically collected from cows 12 - 24 months old) was purchased from Corning. Shape memory alloy ring clamps were purchased from Intrinsic Devices.

[0106]

[0146] In one example, the BC sheet was compressed to a thickness of 0.5 mm and placed in a hot water reactor together with a mixture of PVA (40 wt.%) and DI water (60 wt.%). The hot water reactor was sealed and heated at 135 °C for 24 hours to diffuse PVA into the voids of the BC and form a BC-PVA hydrogel. The BC-PVA hydrogel was removed from the reactor at high temperature (above 85 °C). Note that the hot water reactor is pressurized with hot steam and there is a risk of burns. When opening the reactor, personal protective equipment including a laboratory coat, heat-resistant gloves and a full-coverage face shield must be used. The residual PVA solution was removed by rubbing the surface of the BC-PVA sample with a metal spatula. The sample was frozen at -78 °C for 30 minutes and thawed at room temperature to physically crosslink the PVA network. Next, the BC-PVA hydrogel was immersed in a solution of AMPS (30 wt.%), MBAA (60 mM), I2959 (50 mM) and KPS (0.5 mg mL-1) for 24 hours. The hydrogel was cured on one side with an ultraviolet transilluminator (VWR International) for 15 minutes each and further cured in an oven at 60 °C for 8 hours to ensure it was flat and fully cured. The resulting BC-PVA-PAMPS hydrogel was stored in PBS for at least 24 hours before further characterization.

[0107]

[0147] As described, the BC sheet was compressed, for example, to a thickness of about 0.5 mm. In some examples, the BC sheet was then placed in an oven at 90 °C for 24 hours before annealing at 90 °C for an additional 1 hour. The resulting annealed BC was cut into the desired shape and stored in PBS (0.15 M) for at least 24 hours.

[0108]

[0148] A PVA hydrogel can be fabricated using any suitable method. For example, to fabricate a PVA hydrogel, a slurry of PVA (40 wt.%) and DI water (60 wt.%) was mixed in a metal baking dish (diameter: 203.2 mm) and heated at 120 °C for 20 minutes in an autoclave sterilizer. To produce an annealed PVA hydrogel, the resulting hydrogel was dried in an oven at 90 °C for 24 hours before being further annealed at 90 °C, 120 °C or 140 °C for 1 hour. To produce a freeze - thawed PVA hydrogel, the autoclaved hydrogel was frozen at - 80 °C for 30 minutes and thawed at 23 °C for 30 minutes. Before testing, the resulting PVA hydrogel was cut into the desired shape and stored in PBS (0.15 M) for at least 24 hours.

[0109]

[0149] The annealed BC - 40 wt.% PVA hydrogel may be fabricated by any suitable technique. For example, the BC sheet may be compressed to a thickness of 0.5 mm and placed in a hot water reactor together with a mixture of PVA (40 wt.%) and DI water (60 wt.%). The hot water reactor was sealed and heated at 135 °C for 24 hours to diffuse PVA into the pores of BC and form a BC - PVA hydrogel. The BC - PVA hydrogel was taken out of the reactor at a high temperature (above 85 °C). The remaining PVA solution was removed by rubbing the surface of the BC - PVA sample with a metal spatula. The sample was dried in an oven at 90 °C for 24 hours before being further annealed at 90 °C, 120 °C or 140 °C for 1 hour. Before testing, the resulting annealed BC - PVA hydrogel was cut into the desired shape and stored in PBS (0.15 M) for at least 24 hours.

[0110]

[0150] In one example, an annealed BC-10wt.% PVA hydrogel was fabricated by compressing the BC sheet to a thickness of about 0.5 mm and placing it in a baking dish (15.6 cm × 8.6 cm × 4.2 cm). Approximately 30 mL of a 10wt.% PVA solution was added to the baking dish. The baking dish was placed in an oven at 90 °C for 24 hours and annealed at 90 °C for an additional 1 hour. The resulting annealed BC-PVA hydrogel was cut into the desired shape and stored in PBS (0.15 M) for at least 24 hours.

[0111]

[0151] An annealed BC-PVA-PAMPS hydrogel was fabricated by compressing the BC sheet to 0.5 mm thickness and placing it in a hot water reactor together with a mixture of PVA (40 wt.%) and DI water (60 wt.%). The hot water reactor was sealed and heated at 120 °C for 24 hours to diffuse PVA into the voids of BC and form a BC-PVA hydrogel. The BC-PVA hydrogel was taken out of the reactor at high temperature (> 85 °C). Note that the hot water reactor is pressurized with hot steam and there is a risk of burns. When opening the reactor, personal protective equipment including a lab coat, heat-resistant gloves and a full-coverage face shield must be used. The residual PVA solution was removed by rubbing the surface of the BC-PVA sample with a metal spatula. The sample was dried in an oven at 90 °C for 24 hours before further annealing at 90 °C, 120 °C or 140 °C for 1 hour. Next, the annealed BC-PVA hydrogel was immersed in a solution of AMPS (30 wt.%), MBAA (60 mM), I2959 (50 mM) and KPS (0.5 mg mL-1) for 24 hours. The hydrogel was cured on one side for 15 minutes each with an ultraviolet transilluminator (VWR International) and further cured in an oven at 60 °C for 8 hours to ensure it was flat and fully cured. The resulting annealed BC-PVA-PAMPS hydrogel was stored in PBS (0.15 M) for at least 24 hours before further characterization.

[0112] Fabrication of Hydrogel on Stainless Steel Pins

[0152] In one example, the preparation of the hydrogel sample on the stainless-steel pins was started by cutting the lyophilized BC. The BC was cut into an octagonal shape with eight legs of diameter D, leg length L, and width W = 0.383D. The sample was labeled BC-D-L after cutting. Eight star-shaped pieces (BC-D-L) were generated by MATLAB and loaded into Adobe Illustrator. In Adobe Illustrator, the stroke of the shape was changed to 0.0001 pt to ensure accurate cutting. The file was sent to a laser cutter (Epilog Fusion M2) using the print function, and the laser cutter was selected as the printer. Vector processing was used at 100% speed, 20% output, and 100% frequency. To cut the BC, a clean metal plate was placed on the bed of the laser cutter, and the lyophilized BC was placed on that metal plate. Another metal plate was placed at the end of the BC to prevent the BC from moving. The focus was adjusted, and the shape was cut by the machine. After cutting, the BC was collected and stored in a Petri dish for later use.

[0113]

[0153] To prepare the shear test specimens, six pieces of BC were adhered to stainless steel rods using one layer of cement and clamps. The stainless steel test rods were machined to have an upper section with a diameter of 5.2 mm and a height of 2 mm, and a bottom section with a diameter of 6.75 mm and a height of 13 mm. Three pieces of BC-6.5-2.25 and three pieces of BC-6.5-2 were placed in an alignment jig. Scotchbond Universal adhesive was applied to the layer of BC in contact with the rod and the upper surface of the rod. The adhesive was allowed to cure for 20 seconds before blowing air for an additional 5 seconds. Next, approximately 0.15 g of RelyX Ultimate cement was applied to the same surface coated with Scotchbond Universal adhesive. The rod was pushed into the BC layer and then into a shape memory alloy ring clamp. The cement was cured for 1 hour. The specimens were heated in an oven at 175 °C for 10 minutes to shrink the clamps. Next, the specimens were immersed in DI water in centrifuge tubes for 1 hour before later use. Next, the specimens with BC on top were subjected to a specific hydrogel fabrication process.

[0114]

[0154] The compression test specimens were fabricated without using cement, using stainless steel rods and clamps. The wear test specimens were fabricated without using cement, using stainless steel rods with a diameter of 5.7 mm and a height of 38 mm, three pieces of BC-6.5-2, and shape memory alloy rings.

[0115]

[0155] The monotonic tensile test was performed on an Instron 1321 load frame (Instron, Norwood, MA, USA) and a Test Resources 830LE63 axial torsion testing machine (TestResources, Shakopee, MN, USA) at 0.25 mm per second -1It was executed at the speed of . The completed hydrogel was cut into a V shape of ASTM D638-14 type with a test titanium punch (see, for example, FIGS. 15A to 15D). The dimensions of the sample were measured with calipers before the test. The ultimate tensile strength (UTS) was the maximum stress measured before fracture in the case of the BC-PVA sample, or the maximum compressive stress at 80% strain for the PVA sample. The tensile modulus was obtained as the slope at a stress of 1 MPa of the stress-strain curve for comparison with previous studies on human cartilage.

[0116]

[0156] The compressive properties of all samples were measured using a Test Resources 830LE63-axis torsional testing machine. A cylindrical sample of PVA was cut out from the film of the hydrogel sample with a steel punch having a diameter of 4 mm. The BC-PVA sample was attached to a metal pin for compression testing in order to obtain a sufficiently thick sample. The dimensions of the sample were measured with calipers before the test. The compressive properties were measured at a strain rate of 0.05 seconds -1 . When the material broke before a strain of 0.8, the compressive strength was obtained as the stress at a strain of 0.8, or the stress at fracture. The compressive modulus was derived as the slope at a stress of 0.4 MPa of the stress-strain curve for comparison with previous studies on human cartilage.

[0117]

[0157] Differential scanning calorimetry (DSC) was performed on the hydrogel sample to determine the crystallinity of PVA. The test was completed with a TA Instruments TGA550. In a typical test, approximately 5 mg of the hydrogel sample was placed in an aluminum pan and heated from 25 °C to 300 °C under a nitrogen gas flow at a scan rate of 10 °C / min. Typical thermograms for PVA, BC-PVA and BC-PVA-PAMPS hydrogels are shown in FIGS. 17A to 17C.

[0118]

[0158] Calculations were performed regarding how much PVA was crystallized, that is, the degree of crystallization. After obtaining the DSC thermogram, the area under the melting peak in the range of 140 to 220 °C (shown in FIG. 18) was integrated to obtain J·°C·second -1· g -1A value with the unit of was obtained. This numerical value was then divided by the heating rate (0.17 °C·s -1 ), and ΔH (J·g -1 ) was obtained. Next, the crystallinity (X PVA ) of PVA was calculated by dividing the ΔH for the sample by the amount of heat required to melt a 100% crystalline PVA sample (ΔHc = 138.6 J / g) and the weight fraction, w PVA of PVA in the sample:

[0119]

Number

[0120]

[0159] The weight of approximately 1 g of hydrated hydrogel was measured before drying at 90 °C for 24 hours. Next, the weight of the dehydrated sample was measured. The weight after dehydration was divided by the weight before dehydration to determine the solid weight fraction of the hydrogel sample.

[0121]

[0160] Fourier transform infrared (FTIR) spectroscopy was performed on the hydrogel sample to analyze the changes in the bonds after annealing. The hydrogel sample was cut into 1 cm × 1 cm squares before the test. The test was completed with a Thermo Scientific Nicolet iS50 FT-IR. In a typical test, the sample was held under the detector, the number of scans was set to 32, the resolution was set to 4 (0.482 cm -1 ), and the format was set to % transmittance. Typical FTIR spectra are shown in FIGS. 16A - 16B.

[0122]

[0161] The wear resistance of the hydrogel and porcine cartilage samples was determined in a pin-on-disk configuration. The pin-on-disk method was used with an Anton Paar Rheometer (MR302) and a tribological accessory (SCF7). The cartilage samples were harvested from porcine femurs using an osteochondral autograft transfer system (Arthrex). The femurs were purchased from a local grocery store and frozen at -78 °C before sample collection. The hydrogel samples were polished with #600, #800, #1000, #1200, #1500, #2000, #2500, and #3000 sandpapers and smoothed before testing. The hydrogel pins were fabricated by using the method described in Section 2.8. Disks of hydrogel or porcine cartilage with a diameter of 12.7 mm were adhered to the sample holder with a cyanoacrylate adhesive (Gorilla Glue Company). The test parameters were as follows: 1,000,000 rotations; angular velocity: 319 rotations per minute (maximum linear velocity: 100 mm s-1); vertical force: 28.26 N (pressure: 1 MPa). Before starting the test, a pressure of 1 MPa was applied to each sample for 5 minutes. The test was carried out in FBS. FBS is often used during wear testing to mimic the lubrication provided by synovial fluid.

[0123]

[0162] After the wear test, the samples were rehydrated in FBS for 24 hours and the gel was recovered from the applied pressure before measuring the wear depth with a high-resolution X-ray computed tomography (Micro-CT) scanner (Nikon XTH 225 St). 3D models of the reconstructed Micro-CT images were rendered with Avizo 9 Lite. To measure the wear depth, slices of the 3D model were obtained at the center of the wear track. The wear depth was measured from the image of the central slice using ImageJ.

[0124]

[0163] To calculate the COF, the total frictional force (F) was determined from the torque (T) and the radius (R) of the pin in the pin-on-disk configuration:

[0125]

Equation

[0126]

[0164] Next, the COF can be calculated by

[0127]

Number

[0128] which can be calculated by

[0165] Here, F N is the vertical force (28.26 N). The linear velocity (v) was calculated by v = ωR where ω was the angular velocity of the pin.

[0129]

[0166] In the formula, ω was the angular velocity of the pin.

[0167] The shear test was performed on a Test Resources 830LE63 axial torsion testing machine equipped with a 100 lb. load cell. Each test was performed in a customized shear test fixture (see Figure 13). To shear cartilage from bone or hydrogel from a metal sample, the sample was fixed in the cylindrical hole on the left side of the fixture. A spacer was added directly below the sample to accurately align the shear plane at the cartilage-bone or hydrogel-metal interface. The right side of the fixture was machined to have a complementary semi-cylindrical shape and was used to extrude the hydrogel or cartilage from its substrate. To minimize cracking and delamination, a rubber spacer was placed between the sample and the right shear fixture, and pressure was applied during the shear test. A crosshead displacement rate of 2 mm min-1 was used for all measurements.

[0130]

[0168] A human-sized osteochondral implant with a diameter of 20 mm was fabricated. The upper surface of the implant had a radius of curvature of 20 mm to match the natural curvature of the femoral condyle. This implant was fabricated with 5 layers of BC. To improve integration with bone, a commercially available 0.25-mm-thick coating of pure titanium was applied to the stem and under the base of the implant by a plasma spraying process.

[0131]

Table 1

[0132]

[0169] The above table shows the mechanical properties of annealed BC-PVA, annealed BC-PVA-PAMPS, annealed PVA hydrogels, cartilage, and previously reported hydrogels. This table compares annealed BC-PVA with the properties of freeze-thawed hydrogels and other previously reported hydrogels and shows significant improvements in properties.

[0133]

[0170] Abbreviations used in this table: BC: bacterial cellulose; PVA: poly(vinyl alcohol); PAMPS: poly(2-acrylamido-2-methyl-1-propanesulfonic acid sodium salt); PAAm: polyacrylamide; CNC: cellulose nanocrystal, PA: phenyl acrylate; HA: hydroxyapatite, HACC: 2-hydroxypropyltrimethylammonium chloride chitosan, PCL: polycaprolactone, CS: chitosan. For clarity, the references in this table report four measurement criteria, namely, strength and elastic modulus in tension and compression, and are limited to publications having a tensile strength and a compressive strength higher than 3 MPa.

[0134]

[0171] Any of the methods (including user interfaces) described herein may be implemented as software, hardware, or firmware, and may be described as a non-transitory computer-readable storage medium storing a series of instructions executable by a processor (e.g., a computer, tablet, smartphone, etc.), which, when executed by the processor, causes the processor to perform any of the steps including, but not limited to, display, communication with a user, analysis, change of parameters (including timing, frequency, intensity, etc.), decision-making, warning, etc.

[0135]

[0172] All combinations of the foregoing concepts and the additional concepts discussed in detail below (provided such concepts are not mutually inconsistent) are contemplated as being part of the subject matter of the invention disclosed herein and may be used to achieve the advantages described herein.

[0136]

[0173] As used herein, when a feature or element is described as being "on" another feature or element, it may be directly on the other feature or element or intervening features or elements may also be present. In contrast, when a feature or element is described as being "directly on" another feature or element, no intervening features or elements are present. When a feature or element is described as being "coupled", "attached" or "connected" to another feature or element, it may be directly coupled, attached or connected to the other feature or element or intervening features or elements may also be present. In contrast, when a feature or element is described as being "directly coupled", "directly attached" or "directly connected" to another feature or element, no intervening features or elements are present. Features or elements described or shown with respect to one embodiment may be applicable to other embodiments. It will also be understood by those skilled in the art that reference to a structure or feature "adjacent" to another feature may have portions that overlap or are underlying the adjacent feature.

[0137]

[0174] The terms used in this specification are for the sole purpose of describing particular embodiments and are not intended to limit the present invention. For example, as used in this specification, the singular forms "a", "an", and "the" are intended to include the plural forms as well, unless the context clearly dictates otherwise. The terms "comprises" and / or "comprising", as used in this specification, specify the presence of the stated features, steps, operations, elements, and / or components, but do not preclude the presence or addition of one or more other features, steps, operations, elements, components, and / or groups thereof. As used in this specification, the term "and / or" includes any and all combinations of one or more of the associated listed items and may be abbreviated as " / ".

[0138]

[0175] Spatially relative terms such as "under", "below", "lower", "over", "upper", etc. may be used in this specification to facilitate description of the relationship of one element or feature to another (s) element or feature(s) as illustrated in the figures. It will be understood that spatially relative terms are intended to encompass different orientations of the device in use or operation in addition to the orientation shown in the figures. For example, if the device in the figures is turned upside down, an element described as "under" or "beneath" another element or feature would then be positioned "over" the other element or feature. Thus, the exemplary term "under" can encompass both an orientation of above and below. The device may be otherwise oriented (rotated 90 degrees or at other orientations), and the spatially relative descriptions used herein are to be construed accordingly. Similarly, terms such as "above", "below", "vertical", "horizontal", etc. are used in this specification for illustrative purposes only, unless otherwise specifically stated.

[0139]

[0176] The terms "first" and "second" may be used in this specification to describe various features / elements (including steps), but these features / elements should not be limited by these terms unless the context indicates otherwise. These terms may be used to distinguish one feature / element from another. Thus, without departing from the teachings of the present invention, the first feature / element discussed below may be referred to as the second feature / element, and similarly, the second feature / element discussed below may be referred to as the first feature / element.

[0140]

[0177] Throughout this specification and the following claims, unless the context requires otherwise, the word "comprise", and variations such as "comprises" and "comprising", are meant to imply that various components can be used jointly in methods and aspects (e.g., compositions and devices including devices and methods). For example, it will be understood that the term "comprising" means the inclusion of any recited element or step, but not the exclusion of any other element or step.

[0141]

[0178] In general, all of the apparatuses and methods described herein should be understood to be inclusive, but all or secondary components and / or steps may alternatively be exclusive, and various components, steps, sub-components or sub-steps may be expressed as "consisting of" or alternatively "consisting essentially of".

[0142]

[0179] As used in this specification and the claims, including when used in the examples, unless otherwise specified, all numbers should be read as being preceded by the term "about" or "approximately", even if the term is not specifically indicated. The terms "about" or "approximately" can be used when describing a magnitude and / or a position, and indicate that the recited value and / or position are within a reasonably expected range of the value and / or position. For example, a numerical value can have a value that is + / −0.1% of the recited value (or range of values), + / −1% of the recited value (or range of values), + / −2% of the recited value (or range of values), + / −5% of the recited value (or range of values), + / −10% of the recited value (or range of values), etc. Any numerical value given herein should be understood to include approximately that value or about that value, unless the context indicates otherwise. For example, if the value "10" is disclosed, "about 10" is also disclosed. Any numerical range recited herein is intended to include all sub-ranges subsumed therein. It is also understood that when a value is disclosed, the possible ranges "less than that value", "greater than or equal to that value", and between that value are also disclosed, as would be appropriately understood by one of ordinary skill in the art. For example, if the value "X" is disclosed, "less than X" as well as "greater than or equal to X" (where X is a numerical value here) are also disclosed. Throughout this application, data is provided in a number of different formats, and it is understood that this data represents ranges with respect to endpoints and starting points, as well as any combination of data points. For example, if a particular data point "10" and a particular data point "15" are disclosed, values greater than 10 and 15, greater than or equal to 10 and 15, less than 10 and 15, less than or equal to 10 and 15, and equal to 10 and 15 are also considered to be disclosed, as well as between 10 and 15. It is also understood that each unit between two particular units is also disclosed. For example, if 10 and 15 are disclosed, 11, 12, 13, and 14 are also disclosed.

[0143]

[0180] Although various exemplary embodiments have been described above, many modifications may be made to the various embodiments without departing from the scope of the invention as set forth in the claims. For example, the order in which the various described method steps are performed can often be changed in alternative embodiments, and in other alternative embodiments, one or more method steps may be entirely omitted. Optional features of the various device and system embodiments may be included in some embodiments and not in others. Accordingly, the foregoing description is provided primarily for illustrative purposes and should not be construed as limiting the scope of the invention as set forth in the claims.

[0144]

[0181] The examples and diagrams included herein are illustrative only and not limiting, showing specific embodiments in which the subject matter may be practiced. As described, other embodiments may be utilized and derived from them, such that structural and logical replacements and changes are made without departing from the scope of the present disclosure. Such embodiments of the subject matter of the present invention may be referred to herein, merely for convenience, individually or collectively by the term "invention," and in fact, where two or more inventions are disclosed, it is not intended to spontaneously limit the scope of the present application to any single invention or inventive concept. Accordingly, although specific embodiments have been illustrated and described herein, any alternative arrangements calculated to achieve the same purpose may be substituted for the specific embodiments shown. The present disclosure is intended to embrace any and all adaptations or variations of the various embodiments. Combinations of the above-described embodiments, as well as other embodiments not specifically described herein, will be apparent to those skilled in the art upon reference to the above description.

Claims

1. An implant comprising: an implant body having an upper support surface; an anchor base; and a cellulose-reinforced hydrogel, wherein the cellulose-reinforced hydrogel comprises: a crosslinked cellulose nanofiber network structure fixed on the upper support surface of the implant body; and a gap hydrogel portion within the gap region of the crosslinked cellulose nanofiber network structure, the gap hydrogel portion having a crystallinity of 20% or more. The implant.

2. The implant according to claim 1, wherein the gap hydrogel portion comprises polyvinyl alcohol (PVA).

3. The implant according to claim 1, wherein the cellulose-reinforced hydrogel contains at least 20% by weight of water.

4. The implant according to claim 1, wherein the cellulose-reinforced hydrogel has a tensile strength of more than 40 MPa.

5. The implant according to claim 1, wherein the crosslinked cellulose nanofiber network structure is chemically crosslinked.

6. The implant according to claim 1, wherein the crosslinked cellulose nanofiber network structure comprises bacterial cellulose.

7. The implant according to claim 1, wherein the cellulose-reinforced hydrogel has a compressive strength of more than 59 MPa.

8. The implant according to claim 1, wherein the crosslinked cellulose nanofiber network structure is fixed on the upper support surface by a clamp.

9. The implant according to claim 8, wherein the crosslinked cellulose nanofiber network structure comprises one or more sheets of bacterial cellulose (BC) held on the upper support surface by a clamp fixed to the lip or rim of the upper support surface.

10. A method of forming an implant having a cellulose-reinforced hydrogel, the method comprising: attaching a crosslinked cellulose nanofiber network structure to the upper support surface of the implant; penetrating a hydrogel component into the gap region of the crosslinked cellulose nanofiber network structure to form the cellulose-reinforced hydrogel; and annealing the cellulose-reinforced hydrogel such that the crystalline substance of the hydrogel component has a crystallinity of 20% or more.

11. The method according to claim 10, wherein the hydrogel component comprises polyvinyl alcohol (PVA).

12. ​ ​ The method according to claim 10, wherein the step of annealing the cellulose-reinforced hydrogel includes a step of heating the cellulose-reinforced hydrogel.

13. The method according to claim 10, wherein the step of annealing the cellulose-reinforced hydrogel includes a step of heating the cellulose-reinforced hydrogel to reduce the water content of the cellulose-reinforced hydrogel.

14. The method according to claim 12, wherein the cellulose-reinforced hydrogel is heated to a temperature in the range of 90 to 140 °C.

15. The method according to claim 10, wherein the step of annealing the cellulose-reinforced hydrogel includes a step of rehydrating the cellulose-reinforced hydrogel.

16. The method according to claim 15, wherein the step of rehydrating the cellulose-reinforced hydrogel includes a step of increasing the water content of the cellulose-reinforced hydrogel to at least 20 wt%.

17. The method according to claim 10, further comprising a step of removing the excess hydrogel component from the surface of the crosslinked cellulose nanofiber network structure.

18. The method according to claim 17, wherein the excess hydrogel component is removed manually or by shaping the cellulose-reinforced hydrogel.

19. The method according to claim 10, wherein the crosslinked cellulose nanofiber network structure comprises bacterial cellulose (BC).

20. The method according to claim 10, wherein the step of attaching the crosslinked cellulose nanofiber network structure to the upper support surface includes a step of clamping the crosslinked cellulose nanofibers at the peripheral portion of the upper support surface.

21. An implant for knee joint surface reconstruction, comprising an upper support surface including a cellulose-reinforced hydrogel, wherein the cellulose-reinforced hydrogel comprises a cellulose nanofiber network structure and a hydrogel component impregnated in the cellulose nanofiber network structure, the hydrogel component having a crystallinity of 20% or more. The implant.

22. The implant according to claim 21, wherein the hydrogel component comprises polyvinyl alcohol (PVA).

23. The implant according to claim 21, wherein the cellulose-reinforced hydrogel contains at least 20% by weight of water.

24. ​ The implant according to claim 21, wherein the cellulose-reinforced hydrogel has a tensile strength of more than 40 MPa.

25. The implant according to claim 21, wherein the cellulose-reinforced hydrogel is attached to the metal base of the upper support surface with a shear strength of more than 0.2 MPa.

26. The implant according to claim 21, wherein the upper support surface has a coefficient of friction that is not statistically greater than the coefficient of friction (COF) of cartilage.

27. A cellulose-reinforced hydrogel, comprising a cellulose nanofiber network structure and a gap hydrogel portion within the gap region of the cellulose nanofiber network structure, wherein the gap hydrogel portion contains polyvinyl alcohol (PVA) and the hydrogel component has a crystallinity of 20% or more. The cellulose-reinforced hydrogel.

28. The cellulose-reinforced hydrogel according to claim 27, wherein the cellulose nanofiber network structure contains at least 20% by weight of water.

29. The cellulose-reinforced hydrogel according to claim 27 or 28, having a tensile strength of 40 MPa or more.

30. The cellulose-reinforced hydrogel according to any one of claims 27 to 29, having a compressive strength of 59 MPa or more.

Images