Composite material for tissue restoration

Abstract

To provide composite materials and methods that restore lost soft tissue volume while promoting soft tissue regeneration.SOLUTION: A composite material can include a gel and at least one nanostructure disposed within the gel. A method for healing a soft tissue defect can include applying a composite material to a soft tissue defect, where the composite material includes a gel and a nanostructure disposed within the gel. A method for manufacturing a composite material for use in healing soft tissue defects can include providing a gel and disposing nanofibers within the gel.SELECTED DRAWING: Figure 1A

Description

[Technical field]

[0001] CROSS-REFERENCE TO RELATED APPLICATIONS This is an international patent application claiming the benefit of priority under 35 U.S.C. § 119(e) to U.S. Provisional Patent Application No. 62 / 038,030, entitled "Composite Material for Tissue Repair," filed August 15, 2014, which is incorporated herein by reference in its entirety.

[0002] Technical Field The present disclosure relates to composite materials and methods for restoring lost soft tissue volume and promoting soft tissue regeneration. [Background technology]

[0003] Soft tissue defects resulting from trauma, oncological resection or congenital malformations are difficult to treat by conventional means. Current therapies, such as tissue reorganization or tissue transplantation, cause donor site defects. Other therapies, such as prosthetic implantation, result in fibrosis and encapsulation. Also, existing strategies for promoting tissue ingrowth are insufficient for treating soft tissue defects. Currently used acellular matrices result in fibrous, flat sheet-like tissues, rather than the soft, three-dimensional tissues required for ideal reconstruction. Finally, fat grafting can restore soft tissue defects, but its widespread use is hindered by the variable survival rate of grafts and limited restoration volume. An ideal approach to soft tissue reconstruction would facilitate the regeneration of soft tissues such as adipose tissue in vivo, followed by implantation of said tissue to promote regeneration. However, adipose tissue regeneration requires an appropriate matrix for cells to attach, migrate, proliferate, differentiate and organize into the new tissue. The repair site is largely devoid of native extracellular matrix (ECM). Therefore, re-creation of a synthetic matrix that not only immediately restores lost tissue volume but also modifies the microenvironment, supports host cell infiltration, and promotes soft tissue regeneration is an essential challenge when repairing soft tissue defects using adipose tissue-based reconstruction.

[0004] Hydrogels offer several advantages as filler materials for soft tissue reconstruction, however, a high crosslinking density is usually required to obtain sufficient mechanical properties. However, under these conditions, host tissue cells (e.g., adipocyte precursors and endothelial precursor cells) are unable to penetrate and grow into the scaffold. In the case of degradable hydrogels, the formation of scar-forming fibrous tissue is typical, as host tissue ingrowth occurs too slowly, or at least at a slower pace than the resorption of the fibrous material.

[0005] Recently, functionalized nanofibers have been developed that serve as ECM mimics to support various cellular activities. FDA-compliant synthetic biodegradable poly-α-esters, such as polycaprolactone (PCL) or poly(lactide-co-glycolide) (PLGA), can be used to create nanofibers by a method known as electrospinning. Biodegradable sutures and implants prepared from such polymers are widely used clinically due to their excellent track record of biocompatibility. A variety of nanofibers of different diameters and topographies have been developed for stem cell engineering applications. However, such nanofibers do not provide macroscopic structures, making them difficult to use as 3D structural scaffolds.

[0006] In view of the various problems associated with such conventional methods and systems, there remains a need in the art for improved solutions to the healing of soft tissue defects. The present disclosure provides a solution to this need that solves the various problems recognized in the art. Summary of the Invention

[0007] The present invention is based, at least in part, on the identification of structural scaffold composites having polymeric fibrous components with improved properties (e.g., improved qualities for soft tissue reconstruction, as described in further detail below).

[0008] In one aspect, the present invention provides a scaffold composite comprising polymeric fibers having an average diameter of about 100 nm to about 8000 nm covalently bonded to a hydrogel material, the ratio of said fibers to said hydrogel material being about 1:10 to about 10:1 on a component weight basis, or about 1 to 50 mg / mL on a concentration basis.

[0009] In one embodiment, the polymeric fibers comprise a biocompatible, biodegradable polyester. Optionally, the polymeric fibers comprise polycaprolactone.

[0010] In another embodiment, the hydrogel material is present in the composite of a functional network.

[0011] In a further embodiment, the ratio of the fibers to the anhydrous hydrogel material is from about 1:10 to about 10:1.

[0012] In another embodiment, the polymeric fibers comprise non-woven polymeric fibers.

[0013] In some particular embodiments, the polymeric fibers comprise polycaprolactone electrospun fibers. Optionally, the polymeric fibers comprise a synthetic polymeric material comprising poly(lactic-co-glycolic acid), polylactic acid and / or polycaprolactone, or a combination thereof.

[0014] In one embodiment, the composite is formulated to be substantially biocompatible. Optionally, the polymeric fibers comprise biological polymeric materials including silk, collagen, chitosan, and / or combinations thereof.

[0015] In one embodiment, the hydrogel material comprises hyaluronic acid. Optionally, the hydrogel material comprises a hydrogel material comprising poly(ethylene glycol), collagen, dextran, elastin, an alginate, fibrin, an alginate, hyaluronic acid, poly(vinyl alcohol), or derivatives thereof, or combinations thereof.

[0016] In certain embodiments, the hydrogel material comprises an engineered tissue extracellular matrix.

[0017] In one embodiment, the processed tissue extracellular matrix is ​​derivable from adipose tissue.

[0018] In another embodiment, the scaffold composite comprises non-woven polycaprolactone fibers.

[0019] In one embodiment, the hydrogel material comprises hyaluronic acid substantially covering at least a portion of the outer surface of the polycaprolactone fibers.

[0020] In some particular embodiments, the hydrogel material is bonded to the outer surface of the polymer fiber.

[0021] In another embodiment, the scaffold composite further comprises a crosslinkable moiety present in an amount effective to introduce bonds between the polymer fibers and the hydrogel material.

[0022] In certain embodiments, the scaffold complex comprises a plurality of pores present on or within a surface, the pores extending within 1 cm of the surface. 2 The pores are present in a concentration of at least about 50 pores per surface, and at least 80% of the pores on the surface have an average pore diameter that is at least about 5 microns.

[0023] In a further embodiment, the composite scaffold further comprises a cross-linkable moiety present in an amount effective to induce cross-linking between the polycaprolactone fibers and hyaluronic acid.

[0024] Optionally, the scaffold complex promotes tissue growth and cellular infiltration when implanted into a target tissue present in a human subject.

[0025] In certain embodiments, the scaffold composite is substantially biodegradable when implanted in human tissue.

[0026] In one embodiment, the scaffold composite is substantially non-biodegradable when implanted in human tissue.

[0027] In another embodiment, the scaffold complex further comprises a therapeutic agent selected from a cell, a small molecule, a nucleic acid, and a polypeptide.

[0028] According to another aspect of the present invention there is provided an implantable biomaterial comprising a scaffold composite of the present invention.

[0029] In certain embodiments, the implantable material is substantially acellular and / or substantially free of polypeptides.

[0030] In one embodiment, the implantable material is formulated for administration by injection.

[0031] In another embodiment, the implantable material is formulated for subcutaneous administration.

[0032] According to a further aspect of the invention, there is provided a kit comprising the implantable material of the invention.

[0033] According to a further aspect of the present invention there is provided a medical device for maintaining tissue shape in a subject undergoing a surgical procedure comprising an effective amount of a scaffold composite and / or implantable material of the present invention to result in maintenance of the shape of the tissue when administered to a subject.

[0034] Another aspect of the present invention provides a method for preparing an implant for tissue or cartilage repair, comprising the steps of: providing an acellular three-dimensional scaffold comprising polymeric fibers oriented to provide a plurality of pores, where at least a portion of the polymeric fibers are crosslinked to other polycaprolactone fibers; disposing a composition comprising a hydrogel material onto the polymeric fibers to form a composite; and reacting or stabilizing the composite to form a stabilized implant, thereby preparing the implant.

[0035] Optionally, the tissue comprises soft tissue.

[0036] According to a further aspect of the present invention, there is provided a method for preparing an implant for tissue or cartilage repair comprising the steps of: providing an acellular three-dimensional scaffold comprising polymeric fibers oriented to provide a plurality of pores; disposing a composition comprising a hydrogel material onto the polymeric fibers to form a composite; and reacting or stabilizing the composite to form a stabilized implant, wherein at least a portion of the polymeric fibers are crosslinked to the hydrogel material.

[0037] In some particular embodiments, the three-dimensional scaffold comprises reactive polycaprolactone fibers.

[0038] According to a further aspect of the present invention, there is provided a method for preparing an implant for tissue or cartilage repair comprising the steps of: providing an acellular three-dimensional scaffold comprising polymeric fibers oriented to provide a plurality of pores; disposing a composition comprising a hydrogel material onto the polymeric fibers to form a composite; and reacting or stabilizing the composite to form a stabilized implant, wherein at least a portion of the polymeric fibers are crosslinked to the hydrogel material.

[0039] According to a further aspect of the present invention there is provided a method for eliminating a tissue defect resulting from trauma or surgical intervention, comprising expanding said tissue, said expanding said tissue comprising implanting an effective amount of a scaffold complex of the present invention into said tissue, thereby expanding it.

[0040] According to another aspect of the present invention, there is provided a method for reducing or reversing tissue loss due to an age-related disease, disorder or condition, comprising expanding a tissue comprising said tissue, wherein expanding said tissue comprises embedding an effective amount of a scaffold complex of the present invention into said tissue, thereby expanding it.

[0041] Optionally, the tissue defect comprises pleural tissue, muscle tissue, skin, or a combination thereof.

[0042] In at least one embodiment, the present invention provides a composite material comprising a gel and at least one nanostructure disposed in the gel. The gel can be a hydrogel or any other suitable gel. The nanostructure can be a nanofiber or any other suitable nanostructure. The nanostructure can be covalently bonded to the gel. The nanostructure can be made of polycaprolactone (PCL) or any other suitable material.

[0043] In at least another aspect, the present invention provides a method for healing a soft tissue defect comprising applying a composite material to the soft tissue defect, the composite material comprising a gel and nanostructures disposed in the gel.

[0044] In yet another aspect, the present invention provides a method of making a composite material for use in healing soft tissue defects, comprising providing a gel and disposing nanofibers in the gel.

[0045] Where applicable, or unless specifically disclaimed, it is contemplated that any one embodiment described herein can be combined with any other one or more embodiments, even if the embodiments are described in different aspects of the invention.

[0046] These and other embodiments are disclosed or are obvious from and are encompassed by the following detailed description. [Brief description of the drawings]

[0047] The following detailed description, provided by way of example and not intended to limit the invention to the specific embodiments described, can be best understood in conjunction with the accompanying drawings.

[0048] [Figure 1A] FIG. 1A illustrates the structure of one embodiment of a composite according to the present disclosure, showing nanostructures disposed within a gel, and in particular, covalent bonding between the nanostructures and functional groups of the gel.

[0049] [Figure 1B] FIG. 1B shows an optical microscope image of the fully swollen composite shown in FIG.

[0050] [Figure 1C] FIG. 1C is an image of the macroscopic appearance of the hydrated composite shown in FIG.

[0051] [Figure 1D] FIG. 1D shows a scanning electron microscopy (SEM) image of the dehydrated composite shown in FIG. 1, showing ultrastructural similarities to ECM.

[0052] [Figure 2A] FIG. 2A shows the stress-strain curve of one embodiment of the composite of FIG. 1 plotted against HA hydrogel alone, demonstrating improved elastic modulus compared to hydrogels of the same crosslink density.

[0053] [Figure 2B] FIG. 2B shows fatigue testing demonstrating that the composite embodiment of FIG. 2A retains similar robust mechanical integrity compared to conventional hydrogels.

[0054] [Figure 3-1] 3A and 3B show fluorescent overlay (FIG. 3A) and phase contrast images (FIG. 3B) of ASCs cultured for 4 days in nanofiber-HA hydrogel composites.

[0055] [Figure 3-2] Figures 3C and 3D show fluorescent overlay (Figure 3C) and phase contrast images (Figure 3D) of ASCs cultured in regular HA hydrogels for 4 days.

[0056] [Figure 4] 4A and 4B show fluorescent images and overlays (FIG. 4A) and phase contrast images (FIG. 4B) of ASCs migrating from spheroids along aligned 650 nm nanofibers.

[0057] [Figure 5A] FIG. 5A is a photograph showing the appearance of the nanofiber-hydrogel composite in situ under the rat inguinal fat pad.

[0058] [Figure 5B] FIG. 5B shows an H&E stained image of a section of tissue surrounding the composite harvested 2 weeks after implantation.

[0059] [Figure 5C]FIG. 5C shows H&E stained images showing cellular infiltration of tissue sections taken from the composite-tissue interface at 4 weeks.

[0060] [Figure 6-1] FIG. 6A shows the synthesis scheme of polycaprolactone (PCL) fiber-HA hydrogel composite.

[0061] [Figure 6-2] Figure 6B shows a schematic diagram of the composite structure with interfacial bonding between the PCL fibers and the HA chain network.

[0062] FIG. 6C shows optical images illustrating the overall appearance of the as-prepared cylindrical fiber-HA hydrogel composite (left) and HA hydrogel (right) with the same dimensions (scale bar = 5 mm).

[0063] FIG. 6D shows an optical image of the same set of samples after lyophilization and rehydration.

[0064] FIG. 6E shows an SEM image of a cross-section of the HA hydrogel (scale bar = 40 μm).

[0065] Figure 6F shows an SEM image of the cross section of a PCL fiber-HA hydrogel composite (scale bar = 100 μm).

[0066] FIG. 6G shows an SEM image of a cross-section of decellularized native adipose tissue (scale bar=10 μm).

[0067] [Figure 7] Figure 7A shows the effect of fiber diameter and interfacial bonding on the reinforcement compressive modulus of HA hydrogels. HA hydrogels and composites were prepared based on 4.5 mg / ml HA. Stress values ​​were measured at 50% strain. *p<0.05 (Student's t-test).

[0068] Figure 7B shows the effect of fiber diameter and interfacial bonding on the reinforcement compressive modulus of PEG hydrogels. PEG hydrogels and composites were prepared based on 30 mg / ml PEGSH and 20 mg / ml PEGDA, and 1.0 μm PCL fibers were used for the synthesis of fiber-PEG hydrogel composites. Stress values ​​were measured at 50% strain. *p<0.05 (Student's t-test).

[0069] [Figure 8] Figure 8A shows the effect of interfacial bond density and fiber diameter on the enhanced shear storage modulus of HA hydrogels. *p<0.05 (Student's t-test).

[0070] FIG. 8B shows the effect of interfacial bond density and fiber diameter on the enhanced shear storage modulus of PEG hydrogels. Shear storage modulus values ​​were measured at a frequency of 1 Hz. *p<0.05 (Student's t-test).

[0071] Figure 8C shows the effect of interfacial bond density and fiber diameter on the enhanced shear storage modulus of HA hydrogels. Shear storage modulus values ​​were measured at a frequency of 1 Hz. *p<0.05 (Student's t-test).

[0072] FIG. 8D shows the effect of interfacial bond density and fiber diameter on the enhanced shear storage modulus of HA hydrogels. Shear storage modulus values ​​were measured at a frequency of 1 Hz. *p<0.05 (Student's t-test).

[0073] [Figure 9] Figure 9A shows the effect of fiber loading on the shear storage modulus of HA hydrogels. HA hydrogels and composites were synthesized with 10 mg / ml HA. The shear storage modulus is measured at a frequency of 1 Hz. The blue arrows indicate the condition for both composites with a molar ratio of SH groups to (DA+MAL) groups of 1:2. *p<0.05 (Student's t-test).

[0074] FIG. 9B shows the effect of fiber loading on the shear storage modulus of HA hydrogels. HA hydrogels and composites were synthesized with 4.5 mg / ml HA. The shear storage modulus is measured at a frequency of 1 Hz. The blue arrows indicate the condition for both composites where the molar ratio of SH groups to (DA+MAL) groups is 1:2. *p<0.05 (Student's t-test).

[0075] [Figure 10] Figure 10A shows the mechanical strength of fiber-HA hydrogel composites at different frequencies. The shear storage modulus of the HA hydrogel and composites is measured against shear loading at different frequencies.

[0076] Figure 10B shows the mechanical strength of fiber-HA hydrogel composites at different rehydration. Comparison of compressive stress of composites before and after rehydration (strain = 40%).

[0077] Figure 10C shows the mechanical strength of fiber-HA hydrogel composites at different load cycles. The compressive stress of the HA hydrogel and the corresponding composites is measured versus load cycle (strain = 25%).

[0078] [Figure 11] Figure 11A shows the migration ability of human adipose-derived stem cells (hASCs) in HA hydrogels at day 27. The two composites and the HA hydrogel control were selected to show similar compressive modulus around 1.9 kPa. F-actin and nuclei of hASCs were stained with Alexa Fluor® 568 phalloidin (red) and DAPI (blue), respectively. Nanofibers were labeled with Alexa Fluor® 647 (white). Scale bar = 100 μm.

[0079] Figure 11B shows the migration ability of human adipose-derived stem cells (hASCs) in nanofiber-HA hydrogel composites at day 27. The two composites and the HA hydrogel control were selected to show similar compressive modulus around 1.9 kPa. F-actin and nuclei of hASCs were stained with Alexa Fluor® 568 phalloidin (red) and DAPI (blue), respectively. Nanofibers were labeled with Alexa Fluor® 647 (white). Scale bar = 100 μm.

[0080] Figure 11C shows the migration ability of human adipose-derived stem cells (hASCs) in RGD-nanofiber-HA hydrogel composites at day 27. The two composites and the HA hydrogel control were selected to show similar compressive modulus around 1.9 kPa. F-actin and nuclei of hASCs were stained with Alexa Fluor® 568 phalloidin (red) and DAPI (blue), respectively. Nanofibers were labeled with Alexa Fluor® 647 (white). Scale bar = 100 μm.

[0081] FIG. 11D shows the migration ability of human adipose-derived stem cells (hASCs) in RGD-nanofiber-HA hydrogel composites at day 27. The two composites and the HA hydrogel control were selected to show similar compressive modulus around 1.9 kPa. The yellow arrows in (d) and (e) indicate cells adhering to the fibers or fiber aggregates. F-actin and nuclei of hASCs were stained with Alexa Fluor® 568 phalloidin (red) and DAPI (blue), respectively. Nanofibers were labeled with Alexa Fluor® 647 (white). Scale bar = 20 μm.

[0082] Figure 11E shows the migration ability of human adipose-derived stem cells (hASCs) in nanofiber-HA hydrogel composites at day 27. The two composites and the HA hydrogel control were selected to show similar compressive modulus around 1.9 kPa. The yellow arrows in (d) and (e) indicate cells adhering to the fibers or fiber aggregates. F-actin and nuclei of hASCs were stained with Alexa Fluor® 568 phalloidin (red) and DAPI (blue), respectively. Nanofibers were labeled with Alexa Fluor® 647 (white). Scale bar = 20 μm.

[0083] Figure 11F shows the migration ability of human adipose-derived stem cells (hASCs). A schematic diagram of hASCs spheroids within a composite structure with interfacial bonding between PCL fibers and HA chain networks is shown.

[0084] [Figure 12-1] Figure 12A shows the implanted fiber-HA hydrogel composite and HA hydrogel-mediated tissue regeneration at day 30. Macroscopic images (scale bar = 2 mm) of the composite before (insets) and after implantation under the inguinal fat pad are shown. The white asterisk indicates the implanted matrix.

[0085] Figure 12B shows tissue regeneration mediated by implanted fiber-HA hydrogel composites and HA hydrogel at day 30. Macroscopic images of HA hydrogel before (inset) and after implantation (scale bar = 2 mm) under the inguinal fat pad are shown. White asterisks indicate the implanted matrix.

[0086] Figure 12C shows tissue regeneration mediated by implanted fiber-HA hydrogel composite and HA hydrogel at 30 days. H&E and Masson's trichrome staining images of (i) native adipose tissue, (ii) healed tissue after sham surgery, (iii, v) fiber-HA hydrogel-embedded tissue, and (iv, vi) HA hydrogel-embedded tissue at 14 and 30 days are shown. In the images, H=HA hydrogel, C=fiber-HA hydrogel composite, B=brown adipose tissue, yellow arrow=blood vessel. Scale bar=200 μm.

[0087] [Figure 12-2] Figure 12D shows tissue regeneration mediated by implanted fiber-HA hydrogel composite and HA hydrogel at 30 days. H&E and Masson's trichrome stained images of (i) native adipose tissue, (ii) healed tissue after sham surgery, (iii,v) fiber-HA hydrogel-embedded tissue, and (iv,vi) HA hydrogel-embedded tissue at 14 and 30 days are shown. Blue staining with Masson's trichrome stain indicates all collagen in the examined tissues. In the images, H=HA hydrogel, C=fiber-HA hydrogel composite, B=brown adipose tissue, yellow arrow=blood vessel. Scale bar=200 μm.

[0088] [Figure 13] FIG. 13A shows a schematic illustration of the preparation of surface-modified fibers with MAL via a PAA grafting method.

[0089] FIG. 13B shows the average density of carboxyl groups on the fibers after PAA grafting with 3 and 10% (v / v) acrylic acid (*p<0.05, n=6).

[0090] [Figure 14] FIG. 14 shows the shear storage modulus of HA hydrogels with various SH to DA molar ratios prepared with 4.5 mg / ml HA-SH.

[0091] [Figure 15] Figure 15A shows the shear storage modulus of fiber-HA hydrogel composites prepared with different amounts of fibers. The average diameter of the fibers is 686 nm. The MAL surface density on the fibers is 100 nmol / mg, and the composites were prepared with 4.5 mg / ml HA-SH and 5 mg / ml PEGDA. The blue arrow indicates that the molar ratio of SH groups to (DA+MAL) groups is 1 to 2. *p<0.05 (n=3).

[0092] Figure 15B shows the shear storage modulus of fiber-PEG hydrogel composites with various amounts of loaded fiber. *p<0.05 (n=3).

[0093] [Figure 16] FIG. 16 shows the average pore size of HA hydrogels and nanofiber-HA hydrogel composites estimated based on cross-sectional SEM images (*p<0.05).

[0094] [Figure 17] Figure 17A shows cell infiltration and tissue ingrowth into fiber-HA hydrogel composites at day 14. Sectioned tissues were stained for total collagen (blue) by H&E. Labels: C=fiber-HA hydrogel composite, yellow arrows=vessels. Scale bar=50 μm.

[0095] Figure 17B shows cell infiltration and tissue ingrowth into fiber-HA hydrogel composites at day 14. Sectioned tissues were stained with Masson's Trichrome for total collagen (blue). Labels: C = fiber-HA hydrogel composite, yellow arrow = blood vessel. Scale bar = 50 μm.

[0096] Figure 17C shows cell infiltration and tissue ingrowth into the fiber-HA hydrogel composite at day 30. Sectioned tissue was stained for total collagen (blue) by H&E. Labels: C = fiber-HA hydrogel composite, yellow arrow = blood vessel. Scale bar = 50 μm.

[0097] Figure 17D shows cell infiltration and tissue ingrowth into fiber-HA hydrogel composite at day 30. Sectioned tissue was stained for total collagen (blue) with Masson's Trichrome. Labels: C = fiber-HA hydrogel composite, yellow arrow = blood vessel. Scale bar = 50 μm.

[0098] [Figure 18] FIG. 18 shows SEM images of cross-sections of decellularized adipose tissue (upper panel) and fiber-HA hydrogel composite (lower panel).

[0099] [Figure 19] Figure 19A shows the migration ability of hASCs at day 4 in HA hydrogels (G'=24.85μ 2.92 Pa). HA hydrogels were fabricated using 2.5mg / ml HA-SH and 5.0mg / ml PEGDA. Scale bar=100μm.

[0100] Figure 19B shows the migration ability of hASCs at day 4 in 1.0 μm fiber-HA hydrogel composites (G'=32.29μ 2.16 Pa). The composites were fabricated with 2.5 mg / ml HA, 5.0 mg / ml PEGDA and 10 mg / ml fibers. Scale bar=100 μm.

[0101] Figure 19C shows the migration ability of hASCs at day 4 in 286 nm fiber-HA hydrogel composites (G' 39.56μ 1.26 Pa). The composites were fabricated with 2.5 mg / ml HA, 5.0 mg / ml PEGDA and 10 mg / ml fibers. Scale bar = 100 μm.

[0102] [Figure 20] Figure 20A shows an injectable formulation. The fiber-hydrogel composite can be formulated for injectable applications.

[0103] FIG. 20B shows that the injectable composite is stable immediately after injection.

[0104] FIG. 20C shows that the injectable composite remains non-dispersible in water and retains its shape and volume.

[0105] Figure 20D shows cell infiltration and tissue ingrowth into the injectable fiber-HA hydrogel composite at day 30, indicating extensive cellular remodeling and adipocyte formation. Sectioned tissue was stained with H&E. Labeling: c = fiber-HA hydrogel composite. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0106] The present invention relates to a composite material comprising a hydrogel and nanostructures for use in a method for soft tissue reconstruction, the present invention also relates to a method for repairing or reconstructing soft tissue injury using a composition comprising a hydrogel and nanostructures disposed within the hydrogel, and in another aspect, the present invention also relates to a method for fabricating a composition for use in soft tissue reconstruction, the composition comprising a hydrogel and nanostructures disposed within the hydrogel.

[0107] The following is a detailed description of the present invention provided to assist those skilled in the art in carrying out the present invention. Modifications and variations may be made by those skilled in the art in the embodiments described herein without departing from the spirit or scope of the present invention. Unless otherwise defined, all scientific and technical terms used herein have the same meaning as commonly understood by those skilled in the art to which the present invention belongs. The technical terms used herein in the description of the present invention are only for the purpose of describing specific embodiments and are not intended to limit the present invention. All publications, patent applications, patents, figures and other references mentioned herein are expressly incorporated by reference in their entirety.

[0108] Although any methods and materials similar or equivalent to those described herein can also be used in the practice or testing of the present invention, the preferred methods and materials are now described. All publications mentioned herein are incorporated by reference to disclose and describe the methods and / or materials in connection with which the publications are cited.

[0109] Unless otherwise defined, all scientific and technical terms used herein have the same meaning as commonly understood by those skilled in the art to which the present invention belongs. The following references (the entire disclosures of which are incorporated herein by reference) provide those skilled in the art with many common definitions of the terms used in the present invention (unless otherwise defined herein): Singleton et al., Dictionary of Microbiology and Molecular Biology (2nd ed. 1994); The Cambridge Dictionary of Science and Technology (Walker, ed., 1988); The Glossary of Genetics, 5th ed., R. Rieger et al. (ed.), Springer Verlag (1991); and Hale & Marham, the Harper Collins Dictionary of Biology (1991). Generally, the molecular biology method steps and the like described herein or inherent herein are common methods used in the art. Such standard procedures can be found, for example, in reference manuals such as Sambrook et al., (2000, Molecular Cloning--A Laboratory Manual, 3rd ed., Cold Spring Harbor Laboratories); and Ausubel et al., (1994, Current Protocols in Molecular Biology, John Wiley & Sons, New York).

[0110] The following terms may have the meanings ascribed to them below unless otherwise specified. However, it is understood that other meanings known or understood by those skilled in the art are also possible and are within the scope of the present invention. All publications, patent applications, patents, and other references mentioned herein are incorporated by reference in their entirety. In case of conflict, the present specification (including definitions) will control. In addition, the materials, methods, and examples are illustrative only and are not intended to be limiting.

[0111] definition As used herein, a "scaffold complex" encompasses any covalent association of two components: polymeric fibers and hydrogel material. The scaffold complex contains the polymeric fibers and hydrogel material in a "functional network", meaning that the interaction between the components provides a chemical, biochemical, biophysical, physical, or physiological benefit. The functional network may also include additional components, such as cells, biological materials (e.g., polypeptides, nucleic acids, lipids, carbohydrates), therapeutic compounds, synthetic molecules, and the like. In certain embodiments, the scaffold complex promotes tissue growth and cell infiltration when implanted into a target tissue present in a human subject.

[0112] As used herein, the term "hydrogel" is a type of "gel" and refers to a water-swellable polymeric matrix that is composed of a three-dimensional network of macromolecules (e.g., hydrophilic polymers, hydrophobic polymers, blends thereof) held together by covalent or non-covalent crosslinks and can absorb significant amounts of water (e.g., 50%, 60%, 70%, 80%, 90%, 95%, 96%, 97%, 98%, 99% or more per non-water molecule) to form an elastic gel. The polymeric matrix can be formed of any suitable synthetic or natural polymeric material. As used herein, the term "gel" refers to a solid three-dimensional network that spans the volume of a liquid medium and is ensnare by surface tension effects. This internal network may be caused by physical bonds (physical gels) or chemical bonds (chemical gels) as well as crystallites or other junctions that remain intact in the extender liquid. Virtually any fluid can be used as an extender, e.g., water (hydrogels), oil, and air (aerogels). By both weight and volume, gels are predominantly fluid in composition and therefore exhibit densities similar to those of their constituent liquids. Hydrogels are a type of gel in which water is used as the liquid medium.

[0113] The definition of "hydrophobic" and "hydrophilic" polymers is based on the amount of water vapor absorbed by the polymer at 100% relative humidity. According to this classification, a hydrophobic polymer is one that absorbs only up to 1% water at 100% relative humidity ("rh"), while a moderately hydrophilic polymer is one that absorbs 1-10% water, a hydrophilic polymer can absorb more than 10% water, and a hygroscopic polymer is one that absorbs more than 20% water. A "water-swellable" polymer is one that absorbs more than at least 50% of its own weight in water when immersed in an aqueous medium.

[0114] The term "crosslinked" as used herein refers to a composition that contains intra- and / or intermolecular crosslinks, whether caused by covalent or non-covalent bonds, which may be direct or involve a crosslinking agent. "Non-covalent" bonds include both hydrogen bonds and electrostatic (ionic) bonds.

[0115] The term "polymer" includes linear and branched polymer structures, and also crosslinked polymers and copolymers (which may or may not be crosslinked), thus including, for example, block copolymers, alternating copolymers, random copolymers, etc. As used herein, an "oligomer" is a compound having a molecular weight of less than about 1000 Da, preferably less than about 800 Da. Polymers and oligomers may be naturally occurring or may be obtained from synthetic sources.

[0116] Soft tissue reconstruction Millions of people suffer from catastrophic soft tissue loss due to tumor eradication, trauma, aging or congenital deformities every year. Loss of tissue, such as skin, fat and muscle, results in significant functional and cosmetic impairments that are difficult to treat by conventional means. As an example, over 300,000 partial mastectomies are performed in the United States each year, resulting in disfiguring breast scars due to loss of breast soft tissue. Existing options for soft tissue restoration have significant drawbacks. Autologous tissue flaps require the transfer of soft tissue from another part of the body in a lengthy surgical procedure, which renders the donor site defective {LoTempio 2010.Plastic and Reconstructive Surgery,126(2),393-401;Patel 2012.Annals of Plastic Surgery,69(2),139-144}. Prosthetic implants are prone to fibrosis and encapsulation in response to the foreign body {Calobrace 2014 Plastic and Reconstructive Surgery,134(1 Suppl),6S-11;Tsoi 2014.Plastic and Reconstructive Surgery,133(2),234-249}. Fat grafting, which involves the placement of fat cells harvested by liposuction, is limited to small volumes and is hampered by poor graft survival rates {Kakagia 2014 Surgical Innovation,21(3),327-336;Largo 2014 British Journal of Plastic Surgery,67(4),437-448}. Finally, injectable hydrogel soft tissue fillers can be used, but these are only suitable for small defects and the volume restoration they provide is temporary {Young 2011.Acta Biomaterialia,7(3),1040-1049;Varma 2014 Acta Biomaterialia,10(12),4996-5004}. A new generation of tissue engineering solutions has been proposed that focuses on the use of hydrogel scaffolds as templates to regenerate soft tissues (such as adipose tissue) at the reconstruction site.

[0117] Current tissue engineering approaches to soft tissue reconstruction Adipose-derived stem cells (ASCs) have been identified in the wound environment surrounding soft tissue defects {Salibian 2013 Archives of plastic surgery 40.6:666-675}. These cells can differentiate into soft tissues such as adipose if supported by an appropriate matrix microenvironment. Therefore, strategies to fill repair sites with functional materials have the potential to regenerate new tissues using endogenous ASCs. Hydrogels have been widely investigated as structural scaffold matrices for the regeneration of tissue defects due to their three-dimensional (3D) nature and elastic properties similar to those of soft tissues. Various methods have been used to fabricate hydrogel scaffolds that maintain their volume and shape against physical stresses imposed by the surrounding tissues, with an elastic modulus similar to that of native adipose tissue (~2 kPa) {Alkhouli 2013 American Journal of Physiology.Endocrinology and Metabolism,305(12),E1427-35;Sommer 2013 Acta biomaterialia 9.11(2013):9036-9048}. This requires high crosslinking density and small average pore size {Ryu 2011 Biomacromolecules 12.7(2011):2653-2659;Khetan 2013 Nature Materials,12(5),458-465;Li 2014 Journal of Neurotrauma,31(16),1431-1438}, but results in low cell infiltration and poor regeneration. The ability of hydrogel scaffolds to promote cell infiltration is key to successful soft tissue repair. Lack of vascular infiltration contributes to the failure of large volume fat grafts and other tissue engineering attempts. There are currently no materials available that can fill the lost volume of soft tissue defects and promote early vascularization and ASC differentiation to regenerate soft tissue.

[0118] Hydrogel Matrix In the past few years, Li and Wen have reported the development of a laminin-derived loop peptide (CCRR IKVAVA hyaluronic acid (HA) hydrogel with optimized pore size and elastic modulus (10-100 Pa) for stem cell transplantation has been developed by conjugating with hyaluronic acid hydrogel (WLC, 10 μM). They have shown that the hydrogel supports robust neural stem cell (NSC) migration and neurite emergence from differentiated cells {Li 2014 Journal of Neurotrauma,31(16),1431-1438}. In a rat controlled cortical injury (CCI) model of traumatic brain injury, the hydrogel promoted significant vasculature network formation when injected 3 days after CCI injury, filling the lesion site (>10 mm) 4 weeks to 6 months after implantation. This improved angiogenesis was due to the hydrogel's ability to retain and present tissue-secreted growth factors, particularly vascular endothelial growth factor (VEGF). Also, literature reports have revealed that small HA degradation fragments of 3-10 disaccharide units are potent regulators of endothelial cell proliferation, migration, tubule formation, and angiogenesis {Slevin 2002 Journal of Biological Chemistry,277(43),41046-41059}. A recent study tested the efficacy of the HA hydrogel to generate human fetal tissue-derived NSC spheroids at the brain lesion site after CCI injury. After implantation, the HA hydrogel led to robust angiogenesis within the scaffold matrix. Regenerated blood vessels grew into the lesion and penetrated into the implanted matrix to support the engraftment and growth of neural progenitor cells. Even though such studies were not for adipose tissue regeneration, such results confirmed the special ability of this optimized HA hydrogel composition in promoting host vascular ingrowth. More importantly, the hydrogel matrix is ​​sufficiently porous to allow robust cell migration inside the hydrogel matrix. However, it is not feasible to use this HA hydrogel directly for soft tissue reconstruction because its mechanical properties are not high enough to maintain the integrity of the implantation site (surrounding adipose tissue has a modulus more than 10 times higher). Increasing the crosslinking density to improve the modulus results in insufficient permeability for cell infiltration and migration. New strategies are needed to enhance the mechanical properties without significantly reducing the average pore size of the bulk hydrogel. Provided are hydrogel materials comprising and / or isolated from processed tissue extracellular matrices, such as extracellular matrices derived from and / or derivable from adipose tissue.

[0119] Structural skeletal complex. Provided herein is a scaffold composite suitable for use as a medical device incorporated into the tissue of a human subject, the composite being administered to the human subject, for example by injection or implantation. The scaffold composite generally contains polymeric fibers having an average diameter of about 10 nm to about 10,000 nm, for example about 100 nm to about 8000 nm, or about 150 nm to about 5,000 nm, or about 100, 150, 200, 250, 300, 350, 400, 450, 500, 600, 700, 800, 900, 1,000, 1,500, 2,000, 2,500, 3,000, 3,500, 4,000, 4,500, 5,000, 5,500, 6,000, 6,500, 7,000, 7,500, or 8,000. As provided herein, the ratio of polymeric fibers to hydrogel material can be determined by any means known in the art. For example, the ratio of polymeric fibers to hydrogel material can be about 1:100 to about 100:1, such as about 1:50 to about 50:1, or 1:10 to about 10:1, such as 1:5 to about 5:1, such as about 1:3 to about 3:1, based on the weight of the components. The ratio of polymeric fibers to hydrogel material can also be based on concentration, such as a given weight of polymeric fibers per unit volume of hydrogel material. For example, the concentration can be about 1 to 50 mg / mL. The hydrogel material is typically disposed on the polymer fibers, such as attached to the outer surface (or one of the outer surfaces depending on the composition and shape) of the polymer fibers. The scaffold complex is generally not a uniform solid material. Instead, the scaffold complex includes a plurality of pores present on or within the surface of the scaffold complex. The presence, size, distribution, frequency and other parameters of the pores can be tailored during the creation of the scaffold complex. The pore size can be from about less than 1 micron to 100 microns, e.g., 1, 2, 3, 4 5, 10, 15, 20, 30, 40, 50, 60 70, 80, 90 or 100 microns, and the size can be narrowly tailored, e.g., such that at least 40%, e.g., 50%, 60%, 70%, 80%, 90%, 95% or more than 95% of the pores are of the desired size or within the desired size range.

[0120] The scaffold complexes of the invention are suitable for incorporation into tissues of a human subject and are therefore generally "biocompatible" (meaning capable of interacting with biological systems (e.g., those found in a human subject) without inducing a pathophysiological response in and / or by the system). In some embodiments, the scaffold complexes are provided for permanent retention in tissue. Alternatively, the scaffold complexes are provided for temporary retention in a human subject and are substantially biodegradable. Preferably, the polymeric fibers comprise a biocompatible, biodegradable polyester. In one preferred embodiment, the polymeric fibers comprise polycaprolactone.

[0121] As provided herein, a preferred form of interaction for the composite comprising a polymer fiber and a hydrogel generally includes crosslinkable moieties present in an amount effective to introduce bonds between the polymer fiber and the hydrogel material, e.g., an amount effective to induce crosslinks between the polycaprolactone fiber and hyaluronic acid.

[0122] Design of structural frameworks for soft tissue repair The composite concept has been widely used as a material reinforcement mechanism. For example, adding hydroxyapatite particles to hydrogels can increase stiffness {Wu 2008 Materials Chemistry and Physics 107.2(2008):364-369}, and elongated particles can increase the tensile modulus of the composite even more {Yusong 2007 Journal of Materials Science,42(13),5129-5134}. Electrospun nanofiber meshes have been widely used as tissue engineering substrates due to their topographical similarity to native ECM. Of particular importance, the decellularized ECM of adipose tissue is highly fibrous and porous in nature (Figure 6G) {Young 2011.Acta Biomaterialia,7(3),1040-1049}. Several recent studies have attempted to recreate this fibrous component by introducing fragmented poly(lactide) (PLA) or chitosan fibers into polyethylene glycol (PEG), polyacrylamide or alginate hydrogels {Coburn 2011 Smart Structures and Systems,7(3),213;#37;Zhou 2011 Colloids and Surfaces B:Biointerfaces,84(1),155-162;Shin 2015 Journal of Materials Chemistry}. When mixed with a hydrogel precursor solution, the fragmented fibers are incorporated into the hydrogel during the gelation process, creating 3D structures. Such fiber-embedded hydrogels have shown improved mechanical properties compared to the corresponding hydrogels. However, no in vivo host cell infiltration studies have been reported. Also, such hydrogels are non-degradable, and adhesive ligands are required for adipocyte adhesion and differentiation.

[0123] Design of nanofiber-hydrogel composites To obtain fiber reinforcement while maintaining high porosity within the hydrogel phase, we provide electrospun fiber-hydrogel composites that provide superior properties compared to other scaffolds. In addition to blending nanofibers with hydrogel matrix as previously reported {Coburn 2011 Smart Structures and Systems,7(3),213}, we introduce interfacial bonding between the fiber surface and the crosslinked network of hydrogel (Figure 6). Such composite design not only allows for stronger mechanical reinforcement from the solid fiber component, but also allows for independent fine tuning of the bulk mechanical properties and average pore size / porosity of the hydrogel phase, allowing for both optimal cell infiltration properties and structural integrity. It is further envisioned that the fibers can be used as a preferred cell adhesion substrate for ASCs and endothelial progenitor cells, thus acting as a guide to support cell migration and ASC differentiation.

[0124] technological innovation In some specific embodiments, the key innovation is the nanofiber-hydrogel composite design with interfacial bonding between the nanofiber surface and the hydrogel network (Figure 6A). This engineered composite has the potential to dramatically improve the mechanical properties of hydrogels without significantly reducing the average pore size of the hydrogel phase. The introduction of interfacial bonding can provide superior mechanical strength benefits compared to simply physically blending the two components. This study projects the range of mechanical properties (compressive and shear moduli) that can be achieved with electrospun polycaprolactone (PCL) fiber-HA hydrogel composites as opposed to blends. The second innovation is the demonstration that such nanofiber-hydrogel composites repair soft tissue defects. Preliminary characterization showed that the composite shares structural features with adipose tissue (Figure 6) {Christman, 2012 US Patent Publication No. 20120264190; Young 2011.Acta Biomaterialia,7(3),1040-1049}. We hypothesized that this composite would provide structural integrity and mechanical properties important for soft tissue regeneration, and this study demonstrated its versatility and efficiency compared to hydrogels.

[0125] Successful completion of this project will provide an off-the-shelf solution for restoration of missing soft tissue volume, particularly for large defects where establishment of vascular networks, maintenance of tissue repair site integrity, promotion of cell migration and organization, and recruitment of host cells are all crucial for sustained tissue repair. Extensive clinical track record for the material components used in this composite design, i.e., HA hydrogel and biodegradable polyester fibers, coupled with this preliminary data on tissue compatibility, suggests excellent tissue compatibility and a straight path to regulatory approval for clinical application.

[0126] Features: In some embodiments, the present invention provides interfacial bonding between the nanofibers and the polymer network in the hydrogel component, which is important for the formation of a "true" composite. It has been demonstrated that blending such fibers with hydrogels does not provide the same degree of mechanical enhancement. There are also previous reports on the use of nanofiber-hydrogel blends. In other words, it is this interfacial bonding that importantly distinguishes this novel work from the prior art. Furthermore, the interfacial bonding can include covalent bonds as shown in this document, as well as secondary bonds, such as hydrogen bonds and electrostatic charge interactions.

[0127] This is also the first work in the art to show isotropic reinforcement, i.e., the composite becomes stronger in all directions necessary to replace the missing volume of any geometry. Nanofiber mats or designs with small numbers of aligned filaments are anisotropic in nature. The present designs can constitute both isotropic and anisotropic materials.

[0128] In the work presented herein, at least in certain embodiments, the components used to form the composite are defined as a hydrogel network having a pore size and porosity sufficient for cell migration and host tissue ingrowth, and nanofibers containing a sparsely packed array of polymer fibers having diameters in the range of 50 nm to 10 μm.

[0129] Gel / Hydrogel Components The hydrogel composites of the present invention may include any type of suitable hydrogel component. The present invention contemplates nanostructure / gel composites that include any suitable gel component, for example, any suitable hydrogel component known in the art. Gels and / or hydrogels of any suitable synthetic or natural materials may be formed.

[0130] For example, the polymeric component of the gel and / or hydrogel may include cellulose esters such as cellulose acetate, cellulose acetate propionate (CAP), cellulose acetate butyrate (CAB), cellulose propionate (CP), cellulose butyrate (CB), cellulose propionate butyrate (CPB), cellulose diacetate (CDA), cellulose triacetate (CTA), etc. Such cellulose esters are described in U.S. Patent Nos. 1,698,049, 1,683,347, 1,880,808, 1,880,560, 1,984,147, 2,129,052, and 3,617,201, and may be prepared using techniques known in the art, or may be obtained commercially. Commercially available cellulose esters suitable herein include CA 320, CA 398, CAB 381, CAB 551, CAB 553, CAP 482, and CAP 504, all available from Eastman Chemical Company, Kingsport, Tenn. Such cellulose esters typically have a number average molecular weight of about 10,000 to about 75,000.

[0131] Cellulose esters include and comprise a mixture of cellulose and cellulose ester monomer units; for example, commercially available cellulose acetate butyrate includes cellulose acetate as well as cellulose butyrate monomer units and non-esterified cellulose units.

[0132] The gels / hydrogels of the present invention may also be composed of other water-swellable polymers, such as acrylate polymers, which are generally formed from acrylic acid, methacrylic acid, methyl acrylate, ethyl acrylate, methyl methacrylate, ethyl methacrylate and / or other vinyl monomers. Suitable acrylate polymers are the copolymers available under the trade name "Eudragit" from Rohm Pharma (Germany) mentioned above. The Eudragit series E, L, S, RL, RS and NE copolymers are available solubilized in organic solvents, in aqueous dispersions or as dry powders. Preferred acrylate polymers are copolymers of methacrylic acid and methyl methacrylate, such as the Eudragit L and Eudragit S series polymers. Particularly preferred such copolymers are Eudragit L-30D-55 and Eudragit L-100-55 (the latter copolymer is a spray-dried form of Eudragit L-30D-55 that can be reconstituted with water). The molecular weight of the Eudragit L-30D-55 and Eudragit L-100-55 copolymers is approximately 135,000 Da and the ratio of free carboxyl groups to ester groups is approximately 1:1. The copolymers are generally insoluble in aqueous liquids having a pH below 5.5. Another particularly suitable methacrylic acid-methyl methacrylate copolymer is Eudragit S-100, which differs from Eudragit L-30D-55 in that the ratio of free carboxyl groups to ester groups is approximately 1:2. Eudragit S-100 is insoluble at a pH below 5.5, but unlike Eudragit L-30D-55, it is sparingly soluble in aqueous liquids having a pH in the range of 5.5 to 7.0. The copolymers are soluble at pH 7.0 and above. Also, Eudragit L-100 may be used, which has a pH-dependent solubility profile between that of Eudragit L-30D-55 and Eudragit S-100 insofar as it is insoluble at a pH below 6.0.Those skilled in the art will recognize that Eudragit L-30D-55, L-100-55, L-100 and S-100 may be substituted with other acceptable polymers having similar pH-dependent solubility characteristics.

[0133] Any gel / hydrogel composition described herein may be modified to contain an active agent, thereby serving as an active agent delivery system when applied in the context of delivery of the active agent to a body surface (e.g., tissue repair site). The release of the active agent "loaded" into the hydrogel composition of the present invention typically involves both absorption of water and desorption of the agent by a swelling-controlled diffusion mechanism. The active agent-containing hydrogel composition may be used, by way of example, in transdermal drug delivery systems, wound dressings, topical pharmaceutical formulations, implantable drug delivery systems, oral dosage forms, and the like.

[0134] Suitable active agents that can be incorporated into the hydrogel compositions of the invention and delivered systemically (e.g., transdermally, orally, or in other dosage forms suitable for systemic administration of drugs) include, but are not limited to: stimulants; analgesics; anesthetics; anti-arthritic drugs; respiratory drugs, e.g., anti-asthma drugs; anti-cancer drugs, e.g., antineoplastic drugs; anticholinergic drugs; anticonvulsants; antidepressants; antidiabetic drugs; antidiarrheal drugs; anthelmintics; antihistamines; antihyperlipidemic drugs; antihypertensive drugs; anti-infective drugs, e.g., antibiotics and antivirals; anti-inflammatory drugs; antimigraine preparations; antiemetics; antiparkinsonian drugs; antipruritics; antipsychotic drugs; antipyretics; antispasmodics; antituberculous drugs; antiulcer drugs; antivirals; anti-anxiety drugs; appetite suppressants; attention deficit disorder (ADD) and attention deficit hyperactivity disorder (ADHD) drugs; cardiovascular preparations, e.g., These include calcium channel blockers, antianginals, central nervous system (CNS) drugs, beta-blockers and antiarrhythmics; central nervous system stimulants; cough and cold preparations, e.g., decongestants; diuretics; genetic material; herbal remedies; hormone dissolving drugs; hypnotics; hypoglycemic agents; immunosuppressants; leukotriene inhibitors; mitotic inhibitors; muscle relaxants; narcotic antagonists; nicotine; nutritional agents, e.g., vitamins, essential amino acids and fatty acids; ophthalmic drugs, e.g., antiglaucoma agents; parasympatholytics; peptide drugs; psychostimulants; sedatives; steroid drugs, e.g., progestogens, estrogens, corticosteroids, androgens and anabolic agents; smoking cessation drugs; sympathomimetics; tranquilizers; and vasodilators (e.g., for general coronary, peripheral and cerebral). Specific active agents useful with the adhesive compositions of the present invention include, but are not limited to, anabasine, capsaicin, isosorbide dinitrate, aminostigmine, nitroglycerin, verapamil, propranolol, cilaborin, foridone, clonidine, cytisine, phenazepam, nifedipine, fluacizin, and salbutamol.

[0135] For topical drug administration and / or medicated cushions (eg, medicated foot pads), suitable active agents include, by way of example only:

[0136] Bacteriostatic and bactericidal agents: Suitable bacteriostatic and bactericidal agents include, by way of example only: halogen compounds, such as iodine, iodopovidone complex (i.e., a complex of PVP and iodine (also known as "povidine" and available from Purdue Medical, Inc.); mercury compounds such as phenylmercury borate or merbromin; alkylmercury compounds such as thiomercury compounds; phenols such as thymol, o-phenylphenol, 2-benzyl-4-chlorophenol, hexachlorophene, and hexylresorcinol; and organic nitrogen compounds such as 8-hydroxyquinoline, chlorquinaldol, clioquinol, ethacridine, hexetidine, chlorhexedine, and ambazone.

[0137] Antibiotics: suitable antibiotics include, but are not limited to, lincomycin antibiotics (referring to a type of antibiotic originally collected from streptomyces lincolnensis), tetracycline antibiotics (referring to a type of antibiotic originally collected from streptomyces aureofaciens), and sulfur-based antibiotics, i.e., sulfonamides.Exemplary antibiotics of lincomycin include lincomycin, clindamycin, and related compounds, such as those described in U.S. Patent Nos. 3,475,407, 3,509,127, 3,544,551 and 3,513,155, and their pharmacologically acceptable salts and esters. Exemplary antibiotics of the tetracycline family include tetracycline itself, chlortetracycline, oxytetracycline, tetracycline, demeclocycline, rolitetracycline, methacycline, and doxycycline, as well as their pharmaceutically acceptable salts and esters, particularly acid addition salts such as hydrochlorides. Exemplary sulfur-based antibiotics include, but are not limited to, the sulfonamides sulfacetamide, sulfabenzamide, sulfadiazine, sulfadoxine, sulfamerazine, sulfamethazine, sulfamethizole, sulfamethoxazole, as well as their pharmaceutically acceptable salts and esters, such as sodium sulfacetamide.

[0138] Analgesics: Suitable analgesics include local anesthetics, such as, but not limited to, acetamidoeugenol, alphadolone acetate, alphaxalone, amucaine, amoranone, amylocaine, benoxinate, betoxycaine, biphenamine, bupivacaine, burethamine, butacaine, butaben, butanilicaine, buthalital, butoxycaine, carticaine, 2-chloropropane, tetracycline ... Caine, cinchocaine, cocaethylene, cocaine, cyclomethycaine, dibucaine, dimethisoquin, dimethocaine, diperadon, dyclonine, ecgonidine, ecgonine, ethyl aminobenzoate, ethyl chloride, etidocaine, etoxadrol, β-eucaine, euprocin, phenalcomine, fomocaine, hexobarbital, hexylcaine, hydroxy Dione, hydroxyprocaine, hydroxytetracaine, isobutyl p-aminobenzoate, kentamine, leucinocaine mesylate, levoxadrol, lidocaine, mepivacaine, meprylcaine, metabutoxycaine, methohexital, methyl chloride, midazolam, myrtecaine, nepine, octacaine, orthocaine, oxethazaine, parethoxycaine, phenacaine, phencyclidine, pheno The analgesics referred to herein are: tetracaine, lidocaine, and prilocaine.

[0139] Other topical agents that can be delivered using the hydrogel compositions of the present invention as a drug delivery system include: antifungal agents such as undecylenic acid, tolnaftate, miconazole, griseofulvin, ketoconazole, ciclopirox, clotrimazole, and chloroxylenol; keratolytic agents such as salicylic acid, lactic acid, and urea; vessicants such as cantharidin; antiacne agents such as organic peroxides (e.g., benzoyl peroxide), retinol, and the like. amides (e.g., retinoic acid, adapalene, and tazarotene), sulfonamides (e.g., sodium sulfacetamide), resorcinol, corticosteroids (e.g., triamcinolone), alpha-hydroxy acids (e.g., lactic acid and glycolic acid), alpha-keto acids (e.g., glyoxylic acid), and antibacterial agents specifically indicated for the treatment of acne, e.g., azelaic acid, clindamycin, erythromycin, meclocycline, minocycline, nadifloxacin, skin lightening and bleaching agents such as hydroquinone, kojic acid, glycolic acid and other alpha-hydroxy acids, artocarpin, and certain organic peroxides; agents for treating warts such as salicylic acid, imiquimod, dinitrochlorobenzene, dibutyl squarate, podophyllin, podophyllotoxin, cantharidin, trichloroacetic acid, bleomycin, cidofovir, adefovir, and analogs thereof; and anti-inflammatory agents such as corticosteroids and nonsteroidal anti-inflammatory drugs (NSAIDs) (wherein NSAIDs include ketoprofen, flurbiprofen, ibuprofen, naproxen, fenoprofen, benoxaprofen, indoprofen, pirprofen, carprofen, oxaprozin, pranoprofen, suprofen, alminoprofen, butibufen, fenbufen, and tiaprofenic acid).

[0140] For wound dressings, suitable active agents are those useful in treating wounds, including, but not limited to, bacteriostatic and bactericidal compounds, antibiotics, analgesics, vasodilators, tissue healing promoters, amino acids, proteins, proteolytic enzymes, cytokines and polypeptide growth factors.

[0141] For topical and transdermal administration of some active agents, and in wound dressings, it may be necessary or desirable to incorporate a penetration enhancer into the hydrogel composition to increase the rate of penetration of the agent into or through the skin. Suitable enhancers include, for example, sulfoxides, such as dimethyl sulfoxide (DMSO) and decyl methyl sulfoxide; ethers, such as diethylene glycol monoethyl ether (commercially available as Transcutol) and diethylene glycol monomethyl ether; surfactants, such as sodium laurate, sodium lauryl sulfate, cetyltrimethylammonium bromide, benzalkonium chloride, poloxamers (231, 182, 184), Tweens (20, 40, 60, 80) and lecithin (U.S. Pat. No. 4,783,450); 1-substituted azacycloheptan-2-ones, particularly 1-n-dodecylcyclaza-cycloheptan-2-one (Nelson Research & Development Co., Irvine, Calif., available under the trademark Azone; see U.S. Pat. Nos. 3,989,816, 4,316,893, 4,405,616, and 4,557,934); alcohols, such as ethanol, propanol, octanol, decanol, benzyl alcohol, and the like; fatty acids, such as lauric acid, oleic acid, and valeric acid; fatty acid esters, such as isopropyl myristate, isopropyl palmitate, methyl propionate, and ethyl oleate; polyols and their esters, such as propylene glycol, amides and other nitrogen-containing compounds, such as urea, dimethylacetamide (DMA), dimethylformamide (DMF), 2-pyrrolidone, 1-methyl-2-pyrrolidone, ethanolamine, diethanolamine, and triethanolamine; terpenes; alkanones; and organic acids, particularly salicylic acid and salicylates, citric acid, and succinic acid. Mixtures of two or more enhancers may also be used.

[0142] In certain other embodiments, the composite compositions of the present invention, which include a gel (e.g., hydrogel component) and nanostructures, may also include optional additional additive components. Such components are known in the art and may include, for example, fillers, preservatives, pH adjusters, softeners, thickeners, pigments, dyes, refractile particles, stabilizers, reinforcing agents, anti-adherents, pharmaceutical agents (e.g., antibiotics, angiogenic agents, anti-fungals, immunosuppressants, antibodies, etc.), and penetration enhancers. Such additives and their amounts are selected in a manner that does not significantly interfere with the desired chemical and physical properties of the hydrogel composition.

[0143] Absorbent fillers can be advantageously incorporated to control the degree of hydration when the adhesive is on the skin or other body surface. Such fillers can include microcrystalline cellulose, talc, lactose, kaolin, mannitol, colloidal silica, alumina, zinc oxide, titanium oxide, magnesium silicate, magnesium aluminum silicate, hydrophobic starch, calcium sulfate, calcium stearate, calcium phosphate, calcium phosphate dihydrate, mesh and nonwoven paper, and cotton materials. Other suitable fillers are inert, i.e., substantially non-absorbent, and include, for example, polyethylene, polypropylene, polyurethane polyetheramide copolymers, polyesters and polyester copolymers, nylon, and rayon.

[0144] The composition may also include one or more preservatives, including, by way of example only, p-chloro-m-cresol, phenylethyl alcohol, phenoxyethyl alcohol, chlorobutanol, 4-hydroxybenzoic acid methyl ester, 4-hydroxybenzoic acid propyl ester, benzalkonium chloride, cetylpyridinium chloride, chlorhexidine diacetate or gluconate, ethanol, and propylene glycol.

[0145] The composition may also include a pH adjusting compound.The compound useful as a pH adjusting agent includes but is not limited to glycerol buffer, citrate buffer, borate buffer, phosphate buffer, or may also include citrate-phosphate buffer to ensure that the pH of the hydrogel composition is the same as that of the body surface of an individual.

[0146] Also, the composition may contain suitable softening agent.Suitable softening agents include citric acid esters, such as triethyl citrate or triethyl acetyl citrate, tartrate esters, such as dibutyl tartrate, glycerol esters, such as glycerol diacetate and glycerol triacetate; phthalic acid esters, such as dibutyl phthalate and diethyl phthalate; and / or hydrophilic surfactants, preferably hydrophilic nonionic surfactants, such as sugar fatty acid partial esters, polyethylene glycol fatty acid esters, polyethylene glycol fatty alcohol ethers, and polyethylene glycol sorbitan fatty acid esters.

[0147] The composition may also contain a thickening agent.Preferred thickening agent herein is naturally occurring compound or its derivative, for example: collagen; galactomannan; starch; starch derivatives and hydrolysates; cellulose derivatives, such as methylcellulose, hydroxypropylcellulose, hydroxyethylcellulose and hydroxypropylmethylcellulose; colloidal silicic acid; and sugars, such as lactose, saccharose, fructose and glucose.Synthetic thickening agent, such as polyvinyl alcohol, vinylpyrrolidone-vinyl acetate-copolymer, polyethylene glycol and polypropylene glycol, may also be used.

[0148] In some specific embodiments, the hydrogel composite of the present invention, which comprises a hydrogel and a nanostructure, further comprises a component that promotes angiogenesis. A challenge before the present invention to achieve clinically significant soft tissue regeneration is that the regenerated tissue should preferably be revascularized. Therefore, any material that promotes soft tissue regeneration should preferably also promote angiogenesis. One way to achieve this is through the use of heparin-containing hydrogel components that can act as growth factor binding sites to enrich and retain growth factors that promote angiogenesis and tissue formation.

[0149] In various other embodiments, the composite materials of the present invention may be based on hyaluronic acid (HA) as they are hydrogel materials. HA is a non-sulfated linear polysaccharide with disaccharide repeating units that constitute the hydrogel component. HA is also a non-immunogenic natural component of the extracellular matrix of human tissues and is widely used as a dermal filler in cosmetic and reconstructive procedures.

[0150] Degradation of HA is aided by natural hyaluronidases, whose expression is elevated in areas of tissue injury and inflammation. Importantly, studies have shown that degradation fragments of HA of 3-10 disaccharide units are potent regulators of endothelial cell proliferation, migration, tubule formation, and angiogenesis. These biological functions of HA are thought to be mediated by CD44 in a pathway involving Ras and PKC. Blocking CD44 / HA interactions with anti-CD44 antibodies reduced human microvascular endothelial cell proliferation and migration in vitro. HA hydrogels have been investigated in various cell and tissue injury models as promising matrices for cell delivery. Such hydrogels may act as a protective and supportive scaffold for cells and may also reduce scar formation. Thus, HA may have a crucial role in enhancing tissue regeneration by promoting cell infiltration and by promoting angiogenesis.

[0151] First, the material has three-dimensional integrity and consistency similar to that of natural adipose tissue. This makes it suitable for off-the-shelf restoration of missing soft tissue volume. Second, the material can be preferably deposited with multiple flexible nanofibers that can act as a substrate for migration of adipocytes and endothelial progenitor cells. Third, the material has sufficient porosity to allow such progenitor cells to rapidly infiltrate and integrate into the scaffold rather than forming a fibrous capsule around it. Fourth, the HA hydrogel component provides compressibility and volume expansion, as well as important angiogenesis triggers. Fifth, the nanofiber and hydrogel components are biodegradable, allowing them to be replaced by regenerated soft tissue. Sixth, all component materials have a high safety track record in numerous FDA-approved devices, potentially lowering regulatory hurdles for clinical application.

[0152] The gel / hydrogel / nanostructure composite of the present invention may also include tissue repair agents, such as several growth factors, such as epidermal growth factor (EDF), PDGF, and nerve growth factor (NGF). For example, the composition may include EGF. Epidermal growth factor (EGF) was discovered after the observation that skin wounds in laboratory mice seemed to heal faster when the mice were allowed to lick the wounds. This was not simply due to some bactericidal factors in saliva (such as lysozyme). It was shown that a specific growth factor (now known as EGF) was responsible. EGF is identical to urogastrone and has angiogenic properties. Transforming growth factor-α (TGF-α) is very similar, binds to the same receptor, and is even more effective in stimulating the regeneration of epithelial cells (epithelialization).

[0153] Thus, the hydrogels of the present invention, which contain EGF / TGF, may be advantageously used in accelerating wound healing and reducing scar formation in burns, keloids (especially for burns), skin living dressings, and in the treatment of chronic leg ulcers.

[0154] Tissue repair agents useful in the present invention include several growth factors, such as epidermal growth factor (EDF), PDGF, and nerve growth factor (NGF). Generally, growth-promoting hormones affect one to four types of tissue. Many of the preparations developed from such proteins are aimed at one or another wound repair, but also have other indications. Some of the most important tissue growth factors are further described below.

[0155] The gel / nanostructure compositions of the present invention may also include one or more growth factors that may be useful in the tissue repair methods and other applications of the present invention.

[0156] For example, the present invention contemplates the inclusion of PDGF in the compositions of the present invention. Platelet-derived growth factor (PDGF) is a mitogen for almost all mesenchymal-derived cells, i.e., blood, muscle, bone, cartilage and connective tissue cells. It is a dimeric glycoprotein that exists as an AA or BB homodimer, or as an AB heterodimer. Like many growth factors, PDGF is currently believed to be a member of a large family of factors. In addition to PDGF, this family includes the homodimeric factors vascular endothelial growth factor (VEGF) and placental growth factor (PIGF), VEGF / PIGF heterodimers, and connective tissue growth factor (CTGF), which are PDGF-like factors secreted by human vascular endothelial cells and fibroblasts. Along with NGF, TGF-β, and glycoprotein hormones such as human chorionic gonadotropin hormone (hCG), PDGF is currently classified as a member of the cysteine ​​knot growth factor superfamily. All such factors may be used with the hydrogels of the present invention.

[0157] PDGF is produced by platelets and released during the process of blood clotting. It is just one of the growth factors derived from these cells. PDGF attracts fibroblasts and leukocytes to the site of injury and stimulates the growth of replacement connective tissue (mainly fibroblasts and smooth muscle cells). It stimulates cell division of various cells (e.g., those that produce collagen), thus promoting angiogenesis. It also stimulates mitogenesis, vasoconstriction, chemotaxis, enzyme activity, and calcium mobilization.

[0158] Platelet-derived growth factors can be used to restore bone and soft tissue regeneration during certain treatments with the compositions of the present invention, as well as to accelerate the healing process of chronic and acute wounds.Hence, the hydrogel / nanostructure compositions of the present invention can advantageously include a platelet-derived growth factor cocktail.

[0159] The hydrogel / nanostructure compositions of the present invention can be used in gene therapy, for example, for localized delivery of the PDGF gene. Plasmid DNA encoding PDGF is incorporated into the hydrogel matrix, causing granulation tissue fibroblasts, which arise in the viable tissue surrounding the wound and serve as targets for introduction and expression of the plasmid gene, to proliferate and migrate into the matrix.

[0160] The hydrogel / nanostructure composition of the present invention may also include VEGF to promote angiogenesis. Vascular endothelial growth factor (VEGF - also known as vascular permeability factor) is another vascular growth factor and a multifunctional angiogenic cytokine. It contributes to angiogenesis (blood vessel growth) both indirectly and directly by stimulating the proliferation of endothelial cells at the microvascular level, causing the cells to migrate and to cause general expression modifications. VEGF also increases the permeability of theses endothelial cells, causing them to release plasma proteins outside the vascular lumen, thereby causing changes in the area that contribute to angiogenesis.

[0161] The composition of the present invention may also include FGF. Fibroblast growth factor (FGF) is actually a family of peptides of at least 19, 14, 18 kD that belong to the heparin-binding growth factor family and is mitogenic for cultured fibroblasts and vascular endothelial cells. It is also angiogenic in vivo, and this angiogenicity is enhanced by TNF. FGF can be used in a similar manner to EGF. bFGF (also known as FGF-2) is involved in the regulation of human megakaryocytopoiesis, and FGF has been shown to be effective in stimulating endothelial cell formation and supporting connective tissue repair.

[0162] The hydrogel / nanostructure compositions may also include keratinocyte growth factor (KGF) (also known as FGF-7) for use in wound healing and other disorders involving the destruction of epithelial cells.

[0163] Transforming growth factors (TGFs) have the ability to transform a variety of cell lines, conferring, for example, the ability to grow in culture for more than a limited number of generations, growth in multilayers rather than monolayers, and the acquisition of abnormal karyotypes. There are at least five members of the TGF family, the two most widely studied being TGF-α and TGF-β. The former is mitogenic and angiogenic for fibroblasts and endothelial cells, and promotes bone resorption. The composition may also include TGF. TGF-β is a general mediator of cell regulation, a potent inhibitor of cell growth, and inhibits the proliferation of many cell types. TGF-β can antagonize the mitogenic effects of other peptide growth factors and also inhibit the proliferation of many tumor cell lines. TGF-β also has angiogenic effects and promotes collagen formation in fibroblasts. Indications for the hydrogel of the present invention include chronic skin ulcers, such as neurotrophic foot ulcers in diabetic patients. Other areas include wound healing, bone repair, and immunosuppressive diseases.

[0164] The hydrogel / nanostructure composition of the present invention can be used, for example, to carry suitable cells. To maximize efficacy, cells can be incorporated into the gel immediately prior to application to the wound or other suitable area. Suitable cells include autologous fibroblasts and keratinocytes, which are primarily responsible for the formation of the dermis and epidermis. Separate gels, each containing one cell type, can be applied sequentially or together, or one gel can contain both cell types, although this is generally less preferred.

[0165] The hydrogel / nanostructure compositions of the present invention may beneficially include collagen. In this form, collagen is unlikely to serve a useful structural function, but rather, for example, serves primarily as a sacrificial protein when proteolytic activity is undesirably high, thereby helping to inhibit the softening of healthy tissue.

[0166] Certain enzymes may also be included in the hydrogel / nanostructure composition. Enzymes are used for debridement of both acute and chronic wounds. Debridement is the removal of non-viable tissue and foreign material from the wound and is a naturally occurring event in the wound repair process. During the inflammatory phase, neutrophils and macrophages digest and remove "spent" platelets, cellular debris, and avascular damaged tissue from the wound area. However, when a significant amount of damaged tissue accumulates, this natural process becomes catastrophic and insufficient. As a result, the accumulation of necrotic tissue creates a significant phagocyte demand on the wound, slowing wound healing. Therefore, debridement of necrotic tissue is a specific goal of transsurface therapy and a key component of optimal wound management.

[0167] For example, enzymes can be incorporated into the hydrogels of the present invention for topical application to provide selective debridement. Suitable enzymes can be derived from a variety of sources, such as krill, crustaceans, papaya, milk, and bacteria. Suitable commercially available enzymes include collagenase, papain / urea, and a combination of fibrinolysin and deoxyribonuclease.

[0168] Enzymes for use in the present invention generally act in one of two ways: directly digesting scab components (e.g., fibrin, bacteria, leukocytes, cellular debris, serous exudate, DNA); or dissolving the collagen "anchors" that secure the avascular tissue to the underlying wound environment.

[0169] Optionally, Dakin's solution may be included in the hydrogel of the present invention, typically to provide antibacterial effects and odor control. As a debridement agent, Dakin's solution is non-selective due to its cytotoxic properties. Dakin's solution denatures proteins, making them easier to remove from the wound. It also facilitates the detachment of scabs, which aids in debridement by other methods. If the purpose is debridement, the hydrogel containing Dakin's solution should be changed twice a day. Protection of the skin around the wound is typically provided, for example, with an ointment, a liquid skin protectant film dressing, or a solid skin protectant oblate.

[0170] The gels of the invention may be delivered by any suitable method, for example by delivery via a syringe or bellows pack (single dose delivery systems) or a multi-dose system, such as a pressurized delivery system or a "bag in the can" type system (such as those published in WO 98 / 32675). An example of a bellows pack is shown in published UK design registration number 2082665.

[0171] Therefore, the present invention also extends to a single dose delivery system for the treatment of wounds comprising a gel according to the present invention. The present invention also extends to a pressurized delivery system comprising a gel according to the present invention, and a hydrogel according to the present invention pressurized in an aerosol container that can form a spray upon release of pressure. The use of such a delivery means allows the gel to be delivered to areas of the patient that are otherwise difficult to reach by direct application, for example the patient's back when the patient is lying down.

[0172] In some particular embodiments, it may be advantageous to make the hydrogel compositions of the present invention electrically conductive for use in biomedical electrodes and other electrotherapy situations, i.e., for attaching electrodes or other electrically conductive members to a body surface. For example, the hydrogel compositions may be used to attach transcutaneous nerve stimulation electrodes, electrosurgical return electrodes, or EKG electrodes to the skin or mucosal tissue of a patient. Such applications involve modification of the hydrogel compositions to contain electrically conductive species. Suitable electrically conductive species include ionically conductive electrolytes, particularly those commonly used in the manufacture of electrically conductive adhesives used for application to the skin or other body surfaces, including ionizable inorganic salts, organic compounds, or a combination of both. Examples of ionically conductive electrolytes include, but are not limited to, ammonium sulfate, ammonium acetate, monoethanolamine acetate, diethanolamine acetate, sodium lactate, sodium citrate, magnesium acetate, magnesium sulfate, sodium acetate, calcium chloride, magnesium chloride, calcium sulfate, lithium chloride, lithium perchlorate, sodium citrate, and potassium chloride, as well as mixtures of redox couples, such as ferric and ferrous salts (e.g., sulfate and gluconate). Preferred salts are potassium chloride, sodium chloride, magnesium sulfate, and magnesium acetate, with potassium chloride being most preferred for EKG applications. Virtually any amount of electrolyte may be present in the adhesive composition of the present invention, but electrolytes, if present, are preferably present at a concentration ranging from about 0.1 to about 15 wt.% of the hydrogel composition. The procedures described in U.S. Patent No. 5,846,558 to Nielsen et al. for fabricating biomedical electrodes may be adapted for use with the hydrogel compositions of the present invention, the disclosure of which regarding manufacturing details is incorporated by reference. Other suitable fabrication procedures may be used as well, as will be appreciated by those skilled in the art.

[0173] Crosslinking In some specific applications, particularly where high adhesive and cohesive strength is desired, the polymers of the gels / hydrogels of the invention may be covalently crosslinked. While the present disclosure contemplates that crosslinking may be desired between the polymers of the gel / hydrogel components, it is also contemplated that crosslinking may be desired between the polymers and the nanostructure components of the gels / hydrogels of the composite materials of the invention. The present invention contemplates any suitable means for crosslinking polymers to each other and to the nanostructure components of the gels / hydrogels of the present invention. The polymers of the gels / hydrogels may be covalently crosslinked to other polymers or nanostructures, either by intramolecular or intermolecular bonds or by covalent bonds. In the former case, there are no covalent bonds connecting the polymers to each other or to the nanostructures, while in the latter case, there are covalent crosslinks connecting the polymers to each other or to the nanostructures. Crosslinks may be formed by any suitable means, for example, by using heat, radiation, or chemical curing (crosslinking) agents. The degree of crosslinking should be sufficient to eliminate or at least minimize cold flow under compression. Crosslinking also encompasses the use of a third molecule, a "crosslinker," used in the crosslinking process.

[0174] In thermal crosslinking, a free radical polymerization initiator is used, which can be any known free radical generating initiator commonly used in vinyl polymerization. Preferred initiators are organic peroxides and azo compounds, which are generally used in an amount of about 0.01 wt.% to 15 wt.%, preferably 0.05 wt.% to 10 wt.%, more preferably about 0.1 wt.% to about 5%, and most preferably about 0.5 wt.% to about 4 wt.% of the polymerizable material. Suitable organic peroxides include dialkyl peroxides such as t-butyl peroxide and 2,2 bis(t-butylperoxy)propane, diacyl peroxides such as benzoyl peroxide and acetyl peroxide, peresters such as t-butyl perbenzoate and t-butyl per-2-ethylhexanoate, perdicarbonates such as dicetyl peroxydicarbonate and dicyclohexyl peroxydicarbonate, ketone peroxides such as cyclohexanone peroxide and methyl ethyl ketone peroxide, and hydroperoxides such as cumene hydroperoxide and tert-butyl hydroperoxide. Suitable azo compounds include azobis(isobutyronitrile) and azobis(2,4-dimethylvaleronitrile). The temperature for thermal crosslinking depends on the actual components and can be easily estimated by those skilled in the art, but typically ranges from about 80C to about 200C.

[0175] Crosslinking can also be carried out using radiation, typically in the presence of a photoinitiator. The radiation can be ultraviolet light, alpha radiation, beta radiation, gamma radiation, electron beam and x-rays, although ultraviolet light is preferred. Useful photosensitizers are triplet sensitizers of the "hydrogen abstraction" type, such as benzophenone and substituted benzophenones and acetophenones, e.g., benzil dimethyl ketal, 4-acryloxybenzophenone (ABP), 1-hydroxy-cyclohexyl phenyl ketone, 2,2-diethoxyacetophenone and 2,2-dimethoxy-2-phenylaceto-phenone, substituted α-ketols, e.g., 2-methyl-2-hydroxypropiophenone, benzoin ethers, e.g., benzoin methyl ether and benzoin isopropyl ether, substituted benzoin ethers, e.g., anisoin methyl ether, aromatic sulfonyl chlorides, e.g., 2-naphthalenesulfonyl chloride, photoactive oximes, e.g., 1-phenyl-1,2-propanedione-2-(O-ethoxy-carbonyl)-oxime, thioxanthones, e.g., alkyl- and halogen-substituted thioxanthones, e.g., 2-isopropylthioxanthones, 2-chlorothioxanthones, 2,4 Examples include dimethylthioxanone, 2,4 dichlorothioxanone, and 2,4-diethylthioxanone, as well as acylphosphine oxides. Radiation having a wavelength of 200-800 nm, preferably 200-500 nm, is preferred for use herein, and in most cases low intensity ultraviolet light is sufficient to induce crosslinking. However, with hydrogen abstraction type photosensitizers, high intensity UV exposure may be required to effect sufficient crosslinking. Such exposure may be provided by mercury lamp processors, such as those available from PPG, Fusion, Xenon, and the like. Crosslinking may also be induced by irradiation with gamma radiation or electron beams. Appropriate irradiation parameters, i.e., type and dose of radiation used to effect crosslinking, will be apparent to those skilled in the art.

[0176] Suitable chemical curing agents (also referred to as chemical crosslinking "accelerators") include, but are not limited to, polymercaptans such as 2,2-dimercaptodiethyl ether, dipentaerythritol hexa(3-mercaptopropionate), ethylene bis(3-mercaptoacetate), pentaerythritol tetra(3-mercaptopropionate), pentaerythritol tetrathioglycolate, polyethylene glycol dimercaptoacetate, polyethylene glycol di(3-mercaptopropionate), trimethylolethane tri(3-mercaptopropionate), trimethylolethane trithioglycolate, trimethylolpropane tri(3-mercaptopropionate), trimethylolpropane trithioglycolate, dithioethane, di- or trithiopropane, and 1,6-hexanedithiol. Crosslinking promoters are added to uncrosslinked hydrophilic polymers to promote their covalent crosslinking, or to blends of uncrosslinked hydrophilic polymers and complementary oligomers to effect crosslinking between the two components.

[0177] The polymer and / or nanostructure may also be crosslinked before mixing with the complementary oligomer. In such cases, it may be preferable to synthesize the polymer in crosslinked form by mixing and copolymerizing the monomeric precursor of the polymer with a multifunctional comonomer. Examples of monomeric precursors and corresponding polymeric products are: N-vinylamide precursor for poly(N-vinylamide) products; N-alkylacrylamide for poly(N-alkylacrylamide) products; acrylic acid for polyacrylic acid products; methacrylic acid for polymethacrylic acid products; acrylonitrile for poly(acrylonitrile) products; and N-vinylpyrrolidone (NVP) for poly(vinylpyrrolidone) (PVP) products. The polymerization may be carried out in bulk, suspension, solution or emulsion. Solution polymerization is preferred, with polar organic solvents such as ethyl acetate and lower alkanols (e.g., ethanol, isopropyl alcohol, etc.) being particularly preferred. For the preparation of hydrophilic vinyl polymers, the synthesis is typically carried out by free radical polymerization methods in the presence of the above-mentioned free radical initiators. Polyfunctional comonomers include, for example, bisacrylamide, acrylic or methacrylic acid esters of diols such as butanediol and hexanediol (1,6-hexanediol diacrylate is preferred), other acrylates such as pentaerythritol tetraacrylate, and 1,2-ethylene glycol diacrylate, and 1,12-dodecanediol diacrylate. Other useful multifunctional crosslinking monomers include oligomeric and polymeric multifunctional (meth)acrylates, such as poly(ethylene oxide) diacrylate or poly(ethylene oxide) dimethacrylate; polyvinyl crosslinkers, such as substituted and unsubstituted divinylbenzene; and difunctional urethane acrylates, such as EBECRYL 270 and EBECRYL 230 (acrylate-containing urethanes of weight average molecular weight 1500 and weight average molecular weight 5000, respectively, both available from UCB, Smyrna, Ga.), and combinations thereof. When chemical crosslinkers are used, the amount used is preferably such that the weight ratio of crosslinker to hydrophilic polymer ranges from about 1:100 to 1:5. Chemical crosslinking is optionally combined with radiation curing to obtain high crosslink density.

[0178] Nanostructure The nanostructure components of the present invention may be in any suitable form, such as fibers, filaments, mesh segments, branched filaments or networks, sheets, or shaped particles. Nanostructures may also contain any suitable chemical functional groups that facilitate covalent or non-covalent crosslinking between the nanostructures of the present invention and the polymers of the hydrogel. Methods, techniques, and materials are well known in the art for the fabrication and functionalization of nanostructures.

[0179] In some specific embodiments, microfabrication methods are used to fabricate the nanostructures of the present invention. In various embodiments, the devices of the present disclosure can be assembled and / or manufactured using any suitable microfabrication technique. Such methods and techniques are widely known in the art.

[0180] Microfabrication methods that can be used to fabricate the nanostructures disclosed herein include lithography; etching techniques such as laser, plasma etching, photolithography, or chemical etching, such as wet chemical, dry, and photoresist removal; or by solid freeform techniques such as three-dimensional printing (3DP), stereolithography (SLA), selective laser sintering (SLS), ballistic particle manufacturing (BPM) and fused deposition modeling (FDM); by micromachining; silicon thermal oxidation; electroplating and electroless plating; diffusion techniques such as boron, phosphorus, arsenic, and antimony diffusion; ion implantation; film deposition such as evaporation (filament, e-beam, flash, and shadow and step coverage), sputtering, chemical vapor deposition (CVD), epitaxy (gas, liquid, and molecular beam), electroplating, screen printing, lamination, or combinations thereof. Jaeger. See Microlithography, Micromachining & Microfabrication (SPIE Optical Engineering Press, Bellingham, Wash. 1997). The choice of material used as the mold determines how the surface is configured to form the branched structure.

[0181] For example, state of the art fabrication techniques, Micro-Electro-Mechanical Systems (MEMS), using photolithographic techniques and methods derived from the semiconductor industry, may be used. More recently developed methods include "soft lithography" (Whitesides et al, Angew chem. Int eds., 37;550-575, (1998)) and microfluidic tectonics (U.S. Patent No. 6,488,872, Beebe et al., Nature;404:588-59 (2000)). Reviews and other discussions of polymer microdevice fabrication include Madou, MJ Fundamentals of Microfabrication: The Science of Miniaturization; 2nd ed.; CRC Press: Boca Raton, 1997; Becker, H., and Locascio, LE "Polymer microfluidic devices." Talanta, 56(2):267-287, 2002; Quake, SR, and Scherer, A. "From micro- to nanofabrication with soft materials." Science, 290(5496):1536-1540, 2000; and Whitesides, GM, and Stroock, AD "Flexible methods for microfluidics." Physics Today, 54(6):42-48, 2001, each of which is incorporated herein by reference.

[0182] The nanostructures of the present invention can also be fabricated by electrostatic spinning (also called electrospinning). The electrospinning technique of liquids and / or solutions from which fibers can be formed is well known and has been reported in several patents, such as U.S. Pat. Nos. 4,043,331 and 5,522,879. Electrospinning generally involves introducing a liquid into an electric field, which causes fibers to be produced from the liquid. Such fibers are generally drawn toward a conductor of attractive potential and collected. During the conversion of the liquid into fibers, the fibers harden and / or dry. This hardening and / or drying can be caused by cooling the liquid (i.e., if the liquid is normally solid at room temperature); evaporation of a solvent, such as dehydration (physically induced hardening); or hardening mechanisms (chemically induced hardening).

[0183] Electrospinning has typically involved the use of fibers to create mats or other nonwoven materials, as disclosed, for example, in U.S. Patent No. 4,043,331. Nanofibers ranging from 50 nm to 5 micrometers in diameter can be electrospun into nonwoven or aligned nanofiber meshes. Due to their small fiber diameter, electrospun fabrics inherently have very high surface areas and small pore sizes. These properties make electrospun fabrics promising candidates for several applications, such as: membranes, scaffolds for tissues, and other biomedical applications.

[0184] Electrospun fibers can be produced with very small diameters. Parameters that affect the diameter, consistency and uniformity of electrospun fibers include the concentration (loading) of polymeric material and crosslinker in the fiber-forming formulation, the applied voltage, and the needle-to-collector spacing. According to one embodiment of the invention, the nanofibers have diameters ranging from about 1 nm to about 100 μm. In other embodiments, the nanofibers have diameters ranging from about 1 nm to about 1000 nm. Additionally, the nanofibers may have aspect ratios ranging from at least about 10 to about at least 100. It will be appreciated that due to the very small diameter of the fibers, the fibers have a high surface area per unit mass. This high surface area to mass ratio allows the fiber-forming solution or liquid to be transformed from liquid or solvated fiber-forming material into solid nanofibers in a fraction of a second.

[0185] The polymeric material used to form the nanofiber / nanostructure of the present invention can be selected from any fiber-forming material that is compatible with the crosslinking agent. Depending on the intended application, the fiber-forming polymeric material can be hydrophilic, hydrophobic or amphiphilic. In addition, the fiber-forming polymeric material can be a thermoresponsive polymeric material.

[0186] Synthetic or natural biodegradable or non-biodegradable polymers may comprise the nanofibers / nanostructures of the invention. A "synthetic polymer" refers to a polymer that is synthetically prepared and contains monomer units that do not occur in nature. For example, a synthetic polymer may contain non-natural monomer units such as acrylate or acrylamide units. Synthetic polymers are typically formed by conventional polymerization reactions, such as addition polymerization, condensation polymerization, or free radical polymerization. Synthetic polymers may also include those that have natural monomer units, such as natural peptide, nucleotide, and sugar monomer units, in combination with non-natural monomer units (e.g., synthetic peptide, nucleotide, and sugar derivatives). These types of synthetic polymers may be made by standard synthesis techniques, such as solid phase synthesis, or by recombinant techniques, where permitted.

[0187] A "natural polymer" refers to a polymer prepared either naturally, recombinantly, or synthetically, in which the polymeric backbone is composed of naturally occurring monomeric units. In some cases, a natural polymer may be modified, engineered, derivatized, or otherwise treated to alter the chemical and / or physical properties of the natural polymer. In such cases, the term "natural polymer" is modified to reflect the alteration to the natural polymer (e.g., a "derivatized natural polymer" or a "deglycosylated natural polymer").

[0188] Nanofiber materials can include, for example, both addition and condensation polymeric materials, such as polyolefins, polyacetals, polyamides, polyesters, ethers and esters of cellulose, polyalkylene sulfides, polyarylene oxides, polysulfones, modified polysulfone polymers, and mixtures thereof. Exemplary materials within this general category include polyethylene, poly(ε-caprolactone), poly(lactate), poly(glycolate), polypropylene, poly(vinyl chloride), polymethyl methacrylate (and other acrylics), polystyrene and its copolymers (e.g., ABA-type block copolymers), poly(vinylidene fluoride), poly(vinylidene chloride), polyvinyl alcohol (crosslinked and uncrosslinked forms) with various degrees of hydrolysis (87%-99.5%). Exemplary addition polymers tend to be glassy (Tg above room temperature). This is the case for polyvinyl chloride and polymethyl methacrylate, polystyrene polymer compositions, or alloy or low crystallinity polyvinylidene fluoride and polyvinyl alcohol materials.

[0189] In some embodiments of the invention, the nanofiber / nanostructure material is a polyamide condensation polymer. In a more specific embodiment, the polyamide condensation polymer is a nylon polymer. The term "nylon" is a generic term for all long chain synthetic polyamides. Another nylon may be made by polycondensation of ε-caprolactam in the presence of small amounts of water. This reaction forms nylon-6, a linear polyamide (made from a cyclic lactam also known as ε-aminocaproic acid). Additionally, nylon copolymers are also envisioned. Copolymers may be made by combining various diamine compounds, various diacid compounds and various cyclic lactam structures in a reaction mixture, and then forming nylons with the monomeric materials randomly located in the polyamide structure. For example, nylon 6,6-6,10 material is a nylon made from a blend of hexamethylene diamine and C6 and C10 diacids. Nylon 6-6,6-6,10 is a nylon produced by copolymerization of epsilon aminocaproic acid, hexamethylenediamine, and a blend of C6 and C10 diacid materials.

[0190] Block copolymers can also be used as nanofiber materials. In preparing the composition for the preparation of nanofibers, the solvent system can be selected so that both blocks are soluble in the solvent. An example is ABA (styrene-EP-styrene) polymer or AB (styrene-EP) polymer in methylene chloride solvent. Examples of such block copolymers are Kraton-type AB and ABA block polymers, such as styrene / butadiene and styrene / hydrogenated butadiene (ethylene propylene), Pebax-type ε-caprolactam / ethylene oxide, and Sympatex-type polyester / ethylene oxide and polyurethane of ethylene oxide and isocyanate.

[0191] Addition polymers such as polyvinylidene fluoride, syndiotactic polystyrene, copolymers of vinylidene fluoride and hexafluoropropylene, polyvinyl alcohol, polyvinyl acetate, amorphous addition polymers such as poly(acrylonitrile) and its copolymers with acrylic acid and methacrylate, polystyrene, poly(vinyl chloride) and its various copolymers, poly(methyl methacrylate) and its various copolymers are soluble at low pressure and temperature and can therefore be solution spun relatively easily. Highly crystalline polymers such as polyethylene and polypropylene generally require high temperature and high pressure solvents when wet spun.

[0192] Nanofibers may also be formed from polymeric compositions that include two or more polymeric materials in the form of polymer mixtures, alloys, or crosslinked chemically bonded structures. Blending two related polymeric materials may result in nanofibers with beneficial properties. For example, a high molecular weight polyvinyl chloride may be blended with a low molecular weight polyvinyl chloride. Similarly, a high molecular weight nylon material may be blended with a low molecular weight nylon material. Additionally, different common polymeric species may be blended. For example, a high molecular weight styrene material may be blended with a low molecular weight high impact polystyrene. A nylon-6 material may be blended with a nylon copolymer, such as a nylon-6;6,6;6,10 copolymer. Additionally, a polyvinyl alcohol having a low degree of hydrolysis, such as 87% hydrolyzed polyvinyl alcohol, may be blended with a fully or super hydrolyzed polyvinyl alcohol having a degree of hydrolysis of 98-99.9% and above. All such materials in the form of mixtures may be crosslinked using an appropriate crosslinking mechanism. Nylon may be crosslinked using a crosslinking agent that is reactive with the nitrogen atoms of the amide bonds. Polyvinyl alcohol materials can be crosslinked with hydroxy-reactive materials such as monoaldehydes, e.g., formaldehyde, urea, melamine-formaldehyde resins and their analogs, boric acid and other inorganic compounds, dialdehydes, diacids, urethanes, epoxides, and other known crosslinking agents that react to form covalent bonds between polymer chains, substantially improving molecular weight, chemical resistance, overall strength, and resistance to mechanical degradation.

[0193] Biodegradable polymers can also be used in the preparation of the nanostructures of the present invention. Examples of synthetic polymers of the types that have been investigated as biodegradable materials include polyesters, polyamides, polyurethanes, polyorthoesters, polycaprolactone (PCL), polyiminocarbonates, aliphatic carbonates, polyphosphazenes, polyanhydrides, and copolymers thereof. For example, specific examples of biodegradable materials that can be used in connection with implantable medical devices include polylactides, polyglycolides, polydioxanones, poly(lactide-co-glycolide), poly(glycolide-co-polydioxanones), polyanhydrides, poly(glycolide-co-trimethylene carbonate), and poly(glycolide-co-caprolactone). Blends of these polymers with other biodegradable polymers can also be used.

[0194] In some embodiments, the nanofibers are non-biodegradable polymers. Non-biodegradable means that the polymer generally cannot be degraded non-enzymatically, hydrolytically or enzymatically. For example, a non-biodegradable polymer is one that is resistant to degradation that can be caused by proteases. Non-biodegradable polymers can include either natural or synthetic polymers.

[0195] The inclusion of a crosslinking agent in the composition forming the nanofibers allows the nanofibers to be compatible with a wide range of support surfaces. Crosslinking agents may be used alone or in combination with other materials that provide desired surface properties.

[0196] Suitable crosslinkers include either monomeric (small molecule materials) or polymeric materials with at least two latent reactive activating groups that can form covalent bonds with other materials when subjected to an energy source, such as radiation, electrical energy, or thermal energy. In general, latent reactive activating groups are chemical entities that respond to the application of a specific external energy or stimulus to generate active species, resulting in covalent bonds with adjacent chemical structures. Latent reactive groups are groups that retain their covalent bonds under storage conditions, but form covalent bonds with other molecules when activated by an external energy source. In some embodiments, the latent reactive groups are those that form active species, such as free radicals. Such free radicals may include nitrenes, carbines, or ketones that absorb externally applied electrical, electrochemical, or thermal energy to an excited state.Various examples of known or commercially available latent reactive groups are reported in U.S. Patents 4,973,493; 5,258,041; 5,563,056; 5,637,460; or 6,278,018.

[0197] For example, commercially available multifunctional photo-crosslinkers based on trichloromethyltriazine (available from either Aldrich Chemicals, Produits Chimiques Auxiliaires et de Syntheses (Longjumeau, France), Shin-Nakamura Chemical Industry, Midori Chemicals Co., Ltd. or Panchim SA (France)) may be used. The eight compounds include 2,4,6-tris(trichloromethyl)-1,3,5-triazine, 2-(methyl)-4,6-bis(trichloromethyl)-1,3,5-triazine, 2-(4-methoxynaphthyl)-4,6-bis(trichloromethyl)-1,3,5-triazine, 2-(4-ethoxynaphthyl)-4,6-bis(trichloromethyl)-1,3,5-triazine, 4-(4-carboxylphenyl)-2,6-bis(trichloromethyl)-1,3,5-triazine, 2-(4-methoxyphenyl)-4,6-bis(trichloromethyl)-1,3,5-triazine, 2-(1-ethene-2-2'-furyl)-4,6-bis(trichloromethyl)-1,3,5-triazine and 2-(4-methoxystyryl)-4,6-bis(trichloromethyl)-1,3,5-triazine.

[0198] Methods of Use and Exemplary Embodiments The gel / hydrogel / nanostructure compositions of the present invention may be advantageously used in numerous tissue repair situations, as well as other applications, such as coatings on catheters and other surgical devices and implants, and may also be used to deliver active agents, such as antibiotics, growth factors, and immunosuppressants, as described herein.

[0199] In some specific embodiments, the present invention provides a method for healing a soft tissue defect comprising applying a composite material to the soft tissue defect, the composite material comprising a gel and nanostructures disposed in the gel.

[0200] It will be appreciated that the advantageous properties of the hydrogel / nanostructure compositions described herein include: 1) providing easy characterization and quality control; 2) integrating with existing tissue matrix; 3) directly incorporating into newly formed matrix; 4) directly encapsulating cells and bioactive factors; 5) maintaining biocompatibility; 6) controlled bioabsorption; 7) being easily molded into complex anatomical shapes due to the great structural rigidity of the nanostructures; and 8) the ability to exhibit the mechanical properties of native tissue (e.g., articular cartilage).

[0201] In one example of application, the hydrogel / nanostructure composite composition of the present invention may be used to repair cartilage tissue. Current biologically based surgical procedures for cartilage repair include autologous chondrocyte implantation, drilling, abrasion chondroplasty, microfracture, and mosaicplasty. All of these procedures only treat the lesions of the articular cartilage at the focal site, and not the cartilage exposed at the articular surface, as seen, for example, in severe osteoarthritis and rheumatoid arthritis. These procedures also use either cartilage tissue plugs harvested from the patient or expanded cultured chondrocytes to fill the cartilage defect. Such tissue or chondrocytes are expected to integrate with the existing cartilage matrix and fill the defect by synthesizing a completely de novo material, e.g., newly synthesized hyaline cartilage, with the biomechanical properties of normal cartilage. However, all such procedures promote the formation of reparative tissue (fibrocartilage) rather than true hyaline cartilage, and further involve mechanical damage to the fibrocartilage that is believed to predispose the joint to osteoarthritis. Furthermore, the availability of endogenous cartilage as a repair material is quite limited, and its harvesting presents risks and morbidity to the patient itself. As is evident from the above discussion, the resulting hydrogel / nanostructure compositions disclosed herein represent viable materials for a promising new treatment in patients suffering from cartilage degenerative diseases.

[0202] As described herein, the hydrogel / nanostructure compositions of the present invention can be prepared with a wide variety of properties suitable for any number of synthetic tissue implants or augmentations as well as other clinical applications. As previously described, the materials of the present invention can be used to repair cartilage defects resulting from either injury or disease. Injury defects that can be so repaired can be sports-related or accident-related, and can involve only the cartilage surface layer or can include the underlying subchondral bone. Disease defects that can be repaired using the compositions described herein include those resulting from osteoarthritis and rheumatoid arthritis. Whether injury- or disease-related, such defects can be in either mature cartilage or growth plate cartilage. Hydrogel formulations for synthetic growth plate cartilage may require the inclusion of a non-substituted structural scaffold material to allow for controlled bioabsorption of the growing biomaterial.

[0203] Another area in which the hydrogel / nanostructure compositions described herein may be useful is the repair, reconstruction or augmentation of cartilaginous and soft tissues of the head and neck. The availability of biomaterials for soft tissue augmentation and head and neck reconstruction remains a fundamental challenge in the field of plastic and reconstructive surgery. Considerable research and investment has been made into the development of materials with suitable biocompatibility and longevity. The results of this research have not been promising. When placed in immune-competent animals, the structural integrity of currently proposed materials has been shown to be compromised as the scaffold is absorbed. Furthermore, conventional synthetic materials offer excellent longevity but present certain unavoidable pitfalls. For example, silicone includes safety and long-term immune-related effects issues. The synthetic polymers PTFE (Gore-Tex) and Silastic offer low tissue reactivity but do not offer tissue integration and may represent a risk of long-term foreign body infection and extrusion. The materials described in this application are useful for preparing synthetic soft tissue scaffold materials for augmentation or repair of soft tissue defects in the head and neck. In particular, hydrogel / nanostructure compositions that are non-inflammatory, non-immunogenic, and can be prepared to have an appropriate degree of viscoelasticity (see description herein) could be used as effective implantable scaffold materials.

[0204] The hydrogel / nanostructure compositions of the present invention may also be used as novel biocompatible and biocompliant materials for use in head and neck reconstructive procedures to repair cartilage or bony defects, often secondary to trauma or congenital defects, for example, for preparing cartilage implants. Specific applications for the ear include otoplasty and ear reconstruction, which are often performed to repair cartilage defects due to trauma, neoplasms (i.e., squamous cell carcinoma, basal cell carcinoma, and melanoma), and congenital defects, such as microtia. Specific applications for the nose include cosmetic and reconstructive procedures of the nose and nasal septum. Dorsal hump augmentation, nasal tip, shields, and spreader grafts are often used in cosmetic rhinoplasty. Cartilage is required for repair in nasal reconstruction after trauma, neoplasms, autoimmune diseases (e.g., Wegener's granulomatosis with polyangiitis), or congenital defects. Septal perforation for management is difficult and often the procedure is unsuccessful. Cartilage grafts would be ideal for such applications, since autologous or donor cartilage is often unavailable. Specific applications for the throat include laryngotracheal reconstruction, which in children typically requires harvesting of costal cartilage, which is not done without morbidity. Auricular and septal cartilage are often inadequate for this application. Synthetic cartilaginous materials prepared from the hydrogels disclosed herein can be synthesized to suit each of the aforementioned applications based on fine-tuning of the hydrogel synthesis parameters, such as reagent concentration, substitution, and crosslinking rate. Laryngotracheal reconstruction is usually performed for airway narrowing due to subglottic or tracheal stenosis. The etiology can be traumatic (i.e., trauma from intubation, or tracheotomy) or idiopathic. Other possibilities include use in chin and cheek augmentation and lower eyelid eversion repair, in addition to numerous craniofacial applications. It should be noted that such applications may not require cartilage with the exact mechanical properties of articular cartilage, and it may be desirable to include cell populations or bioactive agents.

[0205] The hydrogel / nanostructure compositions described herein may also be used for repair and narrowing of the nasal passages to prevent chronic accumulation of fluid in the nasal passages, which usually leads to infection and crust formation after overly aggressive surgical resections. Another promising application is in laryngotracheal reconstruction in both children and adults as a result of laryngotracheal injury, for example, due to intubation during surgical procedures (e.g., cardiovascular surgery). The hydrogel / nanostructure compositions described herein may also be used to provide annular ring replacements to protect the carotid artery after cancer resections in the neck - the compositions of the present invention may be placed between the carotid artery and the skin as a protective barrier for the carotid artery from loss of the skin barrier. As a protective coating during neuronal regrowth of resected nerves - in many cases, fibrous tissue forms faster than the neuronal regrowth, preventing its eventual formation. Placing the nerve endings in a preformed tube of the hydrogel / nanostructure composition of the present invention may eliminate the formation of fibrous tissue at the regrowth site.

[0206] The hydrogel / nanostructure compositions of the present invention may also be used for the repair of soft tissue defects in any internal or external organ. For example, the materials of the present invention may be used for use in chin and cheek augmentation and lower eyelid eversion repair, in addition to numerous cranial and facial applications. For cosmetic and reconstructive purposes outside the head and neck, for example, as breast implants for breast augmentation, for example, as wound sealants to fill voids after lymph node removal (i.e., for cancer) in the breast or neck, and to seal lymphatic vessels and relieve uncontrolled drainage into the resection site, which can lead to infection and other complications.

[0207] In addition to the uses described above, the hydrogel / nanostructure compositions described herein can be used in other tissue engineering applications to create synthetic orthopedic tissues, including but not limited to bone, tendons, ligaments, menisci, and intervertebral discs, using strategies and methodologies similar to those described above for the synthesis of artificial forms of cartilage. The hydrogel / nanostructure compositions can also be used to create synthetic non-orthopedic tissues, including but not limited to vocal cords, vitreous bodies, heart valves, liver, pancreas, and kidneys, using strategies and methodologies similar to those described above for the synthesis of artificial forms of cartilage.

[0208] Another area in which the hydrogel / nanostructure compositions disclosed herein may be used is in gastrointestinal applications where it is necessary to treat or prevent the formation of scar tissue or strictures in the abdominal or gastrointestinal organs. There are already several formulations in various clinical and FDA approved stages, generally referred to as "hydrogels", that are designed or intended to be useful in the treatment and prevention of scar formation and / or stricture formation. The material of the present invention is distinguished over other known hydrogels in that the disclosed herein may include nanostructures that may provide support, shape and strength to the hydrogel material. The hydrogel / nanostructure compositions disclosed herein may be used in similar applications to those in which already known hydrogels are used or intended to be used, such as the following: treatment of strictures or scar formation in the gastrointestinal tract. The treatment involves injection of the hydrogel material to prevent scar formation at the site of a predicted stricture, or to prevent recurrence of stricture at the site of an existing stricture after treatment to dilate the strictured GI tract.

[0209] The materials of the present invention may also be used to treat esophageal strictures. Esophageal strictures are a common complication of gastroesophageal reflux disease (GERD). GERD is caused by acid, bile and other harmful stomach contents refluxing into the esophagus and damaging the lining cells of the esophagus. Approximately 7-23% of GERD patients develop esophageal strictures or fibrous scarring of the esophagus. Esophageal scarring may also be caused by ablation therapy used to treat Barrett's esophagus. The main complication of such ablation therapy is that the ablation injury extends too deeply into the esophageal wall, resulting in scarring or stricture of the esophagus. Esophageal strictures prevent normal swallowing and are a major cause of patient morbidity. The materials described herein may be used to treat or prevent esophageal strictures resulting from GERD, Barrett's esophagus and esophageal ablation therapy.

[0210] The composite materials of the present invention may also be used to treat Crohn's disease, which causes strictures or scarring that block or narrow the intestinal lumen, preventing normal intestinal function. The materials of the present invention may be useful for treating or preventing such strictures.

[0211] The composite material may also be used in a method for treating primary sclerosing cholangitis (PSC). PSC is a rare disease of the bile ducts of the liver. The bile ducts form a branching network in the liver and leave the liver through two major branches that join in a common bile duct that drains bile from the liver and gallbladder into the duodenum. Although the bile ducts are very narrow in diameter, usually only up to 2 mm at their largest distal portion, they must normally drain many liters of bile from the liver into the duodenum every day. If such ducts become blocked, a serious condition known as jaundice can result, which causes many toxins and especially hemoglobin degradation products to accumulate in the body. PSC is a scarring or structuring disease in the bile ducts in the liver and the extrahepatic bile ducts that connect the liver and the small intestine. Bile duct strictures in PSC can be treated or prevented by the hydrogel / nanostructure composition of the present invention.

[0212] The composite material of the present invention may also be used to treat chronic pancreatitis. Chronic pancreatitis is a chronic inflammatory disease of the pancreas that can be complicated by scarring or narrowing of the pancreatic duct. Such narrowing blocks the drainage of pancreatic juice, which normally must leave the pancreas through a duct system or drainage channel into the small intestine. Pancreatic juice contains many digestive enzymes and other elements important for normal digestion and nutrient absorption. Blockage or narrowing of the pancreatic duct due to chronic pancreatitis can lead to the pancreas becoming autodigested, leading to severe complications such as life-threatening abdominal infections and or abscesses. Pancreatic narrowing in chronic pancreatitis may be treated or prevented by the hydrogel of the present invention.

[0213] The composition according to the present invention can also be used to treat gallstone-induced strictures of the bile duct and pancreatic duct. Gallstones are a very common disorder, and their main complication is the formation of strictures of the bile duct and pancreatic duct, which can be treated or prevented by the hydrogel. Due to ischemic bowel disease. The intestine is prone to scarring or stricture formation when blood supply is impaired. Impaired blood flow is called ischemia and can be caused by many pathologies, such as cardiovascular disease, atherosclerosis, hypertension, hypovolemia, renal or liver disease-induced hypoalbuminemia, vasculitis, drug-induced diseases, and many others. The end result of all such pathologies can result in intestinal strictures that block the intestine and prevent its normal function. The hydrogel / nanostructure composite of the present invention can be used to treat or prevent ischemic bowel strictures.

[0214] The compositions of the present invention can also be used to treat radiation-induced intestinal strictures. Radiotherapy of cancer is associated with numerous morbidities, the most important of which is the formation of intestinal strictures. The hydrogel composites of the present invention can be used to treat or prevent radiation-induced intestinal strictures.

[0215] In addition to creating synthetic tissue or repairing natural tissue, the hydrogel / nanostructure composites disclosed herein may also be used to provide coatings on non-biological structures or devices (e.g., surgical instruments) or ceramic or metallic prostheses used for surgery or other procedures for in vivo implantation. Such coatings could provide a barrier between the non-biological device material and living tissue. The role of hydrogels as barriers for non-biological devices includes, but is not limited to: 1) inhibition of absorption of macromolecules and / or cells onto the surface of the non-biological device that may lead to protein fouling or thrombosis on the device surface; 2) imparting a non-toxic, non-inflammatory, non-immunogenic, biologically compatible surface to devices made of otherwise non-biologically compatible materials; 3) compatibility with device function, e.g., diffusion of glucose to a glucose sensor, transmission of mechanical force for a pressure sensor, or endothelialization of a vascular graft or stent; 4) enhancement of device function, e.g., providing a charge barrier to the size barriers present in MEMS-based artificial nephrons; 5) incorporation into the non-biological device of viable cell populations entrapped within an aqueous, physiologically compatible environment; and 6) inclusion into the device of drugs or bioactive factors designed to promote angiogenesis, epithelialization, or endothelialization, e.g., growth factors, antivirals, antibiotics, or adhesion molecules.

[0216] Based on the foregoing, the hydrogel / nanostructure composites of the present invention can be used to provide non-allergenic coatings on a variety of implantable devices, such as implantable glucose sensors for diabetes management, and can be used to provide: charge barriers for the development of MEMS-based artificial nephrons; aqueous, physiologically compatible environments in which resident kidney cells, such as podocytes, can be incorporated within MEMS-based artificial nephron designs; and coatings for implantable MEMS devices designed for a variety of purposes, including but not limited to, drug delivery, mechanical sensing, and as biodetection systems.

[0217] The hydrogel / nanostructure composites of the present disclosure, particularly hyaluronan-based hydrogels, may also be covalently attached to silicon-based devices, for example, by first covalently attaching the primary amines of tyramine to the silicon surface to obtain a hydroxyphenyl-coated surface chemistry. This may use the same chemistry used to attach DNA modified with free amines to silicon surfaces. The HA-based hydrogel is then covalently attached to the hydroxyphenyl-coated surface by the same peroxidase-driven chemistry used in the preferred crosslinking mode above.

[0218] The hydrogel / nanostructure composites may also be used to coat non-biological cardiovascular devices such as catheters, stents, and vascular grafts. This could include devices made from materials not commonly used due to their biological incompatibility, but which have superior design features than currently used devices. Bioactive factors may be incorporated into the hydrogel to promote endothelialization or epithelialization of the hydrogel and, therefore, of the implanted device.

[0219] Although specific examples and uses of the hydrogel / nanostructure composites of the invention are described herein, such specific uses are not intended to be limiting. The hydrogel / nanostructure composites of the invention may be used in any application generally used for known hydrogels, and are particularly useful for the repair and / or regeneration of soft tissue anywhere in the body.

[0220] Reference is now made to the drawings, in which like reference numbers identify like structural features or aspects of the subject disclosure. For purposes of explanation and illustration, without limitation, an illustration of one embodiment of a biodegradable composite according to the present disclosure is shown in FIG. 1A and generally designated by the reference character 100. The systems and methods described herein can be used to enhance healing of soft tissue defects.

[0221] See generally Figures 1A-1D. The biodegradable composite 100 can include a gel 103 reinforced with nanofibers 101, which combines the advantages of both the gel 103 and the nanofibers 101. The gel 103 can be any suitable material, including, but not limited to, a hydrogel. The nanofibers 101 can be made of any suitable nanomaterial, such as polycaprolactone (PCL) or any other suitable material, and can be of any suitable shape and / or size. The composite 100 includes high porosity (e.g., to mediate cell adhesion and migration) while maintaining sufficient mechanical properties (e.g., to maintain integrity and tissue support).

[0222] In at least some embodiments, the nanofibers 101 are covalently bonded to the hydrogel 103 forming one or more polymer chains. The covalent bond between the hydrogel 103 and the nanofibers 101 can result in a material that has a combination of ideal properties superior to the component materials used alone or as a simple blend.

[0223] Figure 2A shows the stress-strain curve of one embodiment of the composite of Figure 1 plotted against HA hydrogel alone, showing improved elastic modulus compared to hydrogels of the same crosslink density. As shown, the elastic modulus of the test composite 100 (4.5 mg / ml HA, 10 mg / ml PEG-DA, 6.75 mg / ml PCL fibers) was 750 Pa compared to 320 Pa for the hydrogel alone of the same density. Figure 2B shows fatigue testing demonstrating that the composite shown in Figure 1 retains similar robust mechanical integrity compared to regular hydrogels.

[0224] See Figures 3A-3B. The composite 100 was shown to support migration of adipose tissue-derived stem cells (ASCs). GFP-labeled ASCs from liposuction aspirates were cultured into spheroids and then seeded into the composite or hydrogel.

[0225] Figures 3A and 3B show fluorescence and overlay (Figure 3A) with phase contrast image (Figure 3B) of ASCs cultured in nanofiber-HA hydrogel composite for 4 days. The cells migrate outward, extending long processes and tracks. In contrast, ASCs cultured in HA hydrogel alone, shown in Figures 3C and 3D, did not show significant cell migration.

[0226] 4A and 4B show a fluorescent image and overlay (FIG. 4A) (with a phase contrast image (FIG. 4B)) of ASCs migrating from a spheroid along aligned 650 nm nanofibers 101, demonstrating a strong migratory response to the presence of nanofibers 101.

[0227] Example 1: Preparation of nanofiber-hydrogel composites. Nanofibers were fabricated by electrospinning PCL (polycaprolactone, 80k from Sigma Aldrich). The nanofibers were spun into a random mesh. The spinning parameters were a 10% wt solution of PCL in 90% / 1% w / w DCM-DMF, passed through a 27-gauge blunt needle 15 cm from the target metal plate at a flow rate of 0.6 ml / hr. The needle voltage was +10 kV and the target plate was negatively biased at a voltage of -3 kV. 1 mL of solution was spun per round.

[0228] The fibers are then functionalized through a multi-step process. Briefly, the fibers are plasma treated to have reactive groups on the fiber surface to which acrylic acid is attached by UV initiation. The acrylate groups are then reacted with EDC and a diazimine to form a primary amine. The amine can then be reacted with SMCC to attach a maleimide group that can easily react with the thiol groups of the hydrogel.

[0229] The functionalized fiber mesh was cut into sections of 60 mg or less. A 60 mg sample was immersed in ethanol and then added to a ceramic mortar that was partially filled with liquid nitrogen. The fiber sample became very rigid. While keeping the sample cool enough to maintain rigidity, the fiber sheet was cut into approximately 5 mm x 5 mm sections with scissors. Once the entire sheet was cut, the fibers were ground in a mortar and pestle for approximately 20 minutes while keeping the mortar partially filled with liquid nitrogen. The fiber slurry was then poured into ethanol. Approximately 1 mg of a surfactant was added to the slurry to help inhibit fiber entanglement. The suspension was centrifuged at 300 G for minutes and the supernatant was discarded. The fibers were allowed to dry overnight. The fibers were then weighed into a second centrifuge tube so that the exact concentration of fibers could be suspended. The fibers were then sterilized by immersing in ethanol, centrifuged, the supernatant was discarded, and the fibers were allowed to dry overnight in a biohazard cabinet. The fibers are then resuspended in deionized water to the desired concentration (usually 15 mg / mL).

[0230] To form the hydrogel composite, one vial of Glycosil hyaluronic acid is rehydrated with 1 mL of fiber-suspension to obtain a solution of 15 mg / mL fibers and 10 mg / mL hyaluronic acid. To 900 μL of this solution, 100 μL of 10% PEG-DA stock solution is added to give a final concentration of 13.6 mg / mL fibers, 9 mg / mL hyaluronic acid, and 10 mg / mL PEG-DA. This is the formulation for the first in vivo example, but other formulations have been made by varying the concentrations of the components.

[0231] The resulting composite was milky in color, whereas the hydrogel without fibers was transparent (Figure 1C). The composite gel maintained its shape and had good handling properties, whereas the hydrogel alone group was prone to tearing. The fibers in the hydrogel were dispersed and ranged in length from tens to hundreds of microns (Figure 1B). SEM images of cross sections of broken freeze-dried sample composites show intimate bonding between the fibers and hydrogel components as well as high dispersed fiber density (Figure 1D).

[0232] material and method material Thiol-containing hyaluronic acid (HA) was purchased from ESI BIO (Alameda, CA). Poly(ethylene glycol) diacrylate was purchased from Laysan Bio, Inc. (Arab, AL). The following were obtained from Sigma: poly(e-caprolactone), ethylamino-maleimide, acrylic acid, toluidine blue O, N-hydroxysuccinimide (NHS), cysteine, bovine serum albumin (BSA), acetic acid, and Triton™ X-100. Dulbecco's modified Eagle's medium (DMEM), fetal bovine serum (FBS), penicillin / streptomycin, Alexa Fluor® 568 Phalloidin, and 4',6-diamidino-2-phenylindole (DAPI) were purchased from Invitrogen Life Technologies. Ethyl(dimethylaminopropyl)carbodiimide (EDC) was obtained from AnaSpec, Inc. (Fremont, CA). All other chemicals and reagents were of analytical grade.

[0233] method Electrospinning of PCL nanofibers for rheological experiments: To fabricate PCL fibers with two different diameters, 11.0 and 8.5% (w / v) PCL solutions were prepared in a mixture of dichloromethane and dimethylformamide (9:1, v / v) and in a mixture of chloroform and methanol (3:1, v / v), respectively. Each homogeneous PCL solution was loaded into a syringe with a 27G metal needle. Then, electrospinning was performed with the following parameters: a feed rate of 1.0 ml / h, an applied positive voltage of 15 kV to the metal needle, and a distance of 12 cm between the needle end and ground. The morphology of the fibers was observed using a field emission scanning electron microscope (FESEM, JEOL 6700F), and the fiber diameter was measured on the FESEM images using ImageJ software (US National Institutes of Health, Bethesda, MD).

[0234] In vivo composite electrospinning: Spinning conditions: 16% w / v PCL (95% 45.000 Mn PCL, 5% 80,000 Mn PCL, both from Sigma) in a solvent mixture of dichloromethane and dimethylformamide (9:1, w / w). Fibers were spun at a rate of 5.25 ml / hr from a 27-gauge blunt needle 10 cm away from the face of the grounded wheel (spinning at 1000 rpm). The applied voltage was 15 kV, and the electrospinning pump was passed back and forth at 2 mm / sec in a constant orbit for 140 passes over a travel distance of 85 mm (approximately 4 hours). The fiber sheet was then cut into individual sheets of 14 cm diameter for functionalization.

[0235] Preparation of surface-functionalized fibers with MAL: To surface functionalize the fibers with MAL, carboxyl groups were introduced onto the fiber surface by grafting poly(acrylic acid) (PAA) according to the literature [Interface Focus 2011,1,725-733] (with minor modifications). Briefly, the fibers were plasma treated at room temperature for 10 min in oxygen atmosphere under 280 mmHg to induce free radicals on the fiber surface. Then, 10 ml of 3 or 10% (v / v) acrylic acid solution (in 0.5 mM NaIO3) containing 70 mg of fibers was immersed in UV (36 mW / cm) for photopolymerization of PAA onto the fiber surface (PAA-fiber). 2 The PAA-fibers were exposed to a 20-milliliter (DYMAX Light Curing Systems 5000 Flood, Torrington, CT) for 90 seconds. After incubation of the PAA-fibers at room temperature for 20 minutes, the PAA-fibers were washed three times with 20 ml of deionized water to remove unreacted acrylic acid. After the PAA-fibers were completely air-dried, the density of carboxyl groups on the PAA-fibers was measured by the toluidine blue O (TBO) assay, assuming that TBO interacts with the carboxyl groups on the fibers in a 1:1 molar ratio [J Biomed Mater Res 2003, 67, 1093-1104]. Briefly, the PAA-fibers (1 × 1 cm 2) was immersed in 20 μl of 50% (v / v) ethanol, then completely immersed in 1 ml of 0.5 mM TBO solution (in 0.1 mM NaOH (pH 10)) and reacted at room temperature for 5 h with gentle shaking. After washing with 0.1 mM NaOH (pH 10), the TBO adsorbed on the surface of the PAA-fiber was desorbed with 1 ml of 50% (v / v) acetic acid for 1 h at room temperature with vigorous shaking. The optical density of the supernatant was then measured at 633 nm using a microplate reader (BioTeck Synergy2, Winooski, VT). TBO in 50% (v / v) acetic acid was used as a standard.

[0236] Fiber fragments were prepared by grinding the PAA-fibers using a cryo-mill (Freezer / Mill 6770, SPEX SamplePrep, Metuchen, NJ) with the following parameters: 1 min milling and 3 min cooling in liquid nitrogen for 10 cycles. After collecting the PAA-fiber fragments in a 50 ml conical tube, the PAA-fiber fragments were thoroughly dispersed in a 10 ml mixture of isopropyl alcohol (acohol) and distilled water (1:1, v / v) and modified with aminoethyl-MAL on the fiber surface. Briefly, the PAA-fibers were added to NHS and EDC to activate the carboxyl groups of PAA on the fibers. The molar ratios of carboxyl groups to NHS and EDC were 1:4 and 4, respectively. The activation was carried out at room temperature with gentle shaking. After 1 h, aminoethyl-MAL was added to the carboxyl-activated fibers at a molar ratio of carboxyl groups to aminoethyl-MAL of 1:2. The reaction was then carried out at room temperature for 12 h with gentle shaking. After washing three times with distilled water, the surface-functionalized fibers with MAL were freeze-dried, where the density of MAL on the fibers was assumed to be completely substituted by MAL for all carboxyl groups on the fiber surface.

[0237] Preparation of fiber-HA hydrogel composites: To prepare fiber-HA hydrogel composites, thiol-containing HA and PEGDA were completely dissolved in PBS (pH 7.4) to the desired concentrations of 12.5 mg / mL and 100 mg / mL, respectively. MAL-fibers with the desired concentration of 25 mg / mL were completely dispersed in PBS (pH 7.4). Then, suspensions of nanofibers, HA, PEG-DA and PBS were added successively to the desired final concentrations of the formulation. After the composite precursor solution was mixed uniformly, 100 μL of the composite precursor solution was poured into a mold (φ=8 mm) for rheological testing and incubated at 37° C. for 2 h to gel. For compression testing, 200 μL of the precursor solution was added into a cylindrical Teflon mold (φ=6.35 mm, h=6.35 mm) and incubated as above. To observe the cross-sectional morphology of fiber-HA hydrogel composites and HA hydrogels using FESEM, the composites and HA hydrogels were dehydrated by successive ethanol washes (50%, 70%, 80%, 90%, 100% and 100% ethanol for 10 min each) followed by either critical point drying (Samdri-795, Tousimis, Rockville, MD) or chemical drying (HDMS). Samples were freeze-fractured in liquid nitrogen to expose the internal pore structure. The structures were sputter-coated with a 10 nm platinum layer (Hummer 6.2 Sputter System, Anatech UDA, Hayward, CA) and then imaged with a field emission SEM (JEOL 6700F, Tokyo, Japan).

[0238] For preparation of composites for in vivo animal testing, thiol-containing HA was reconstituted in PBS to 12.5 mg / mL. PEG-DA was dissolved in PBS to 100 mg / mL. MAL-fibers were resuspended in sterile PBS to 25 mg / mL. The fibers were first combined with HA solution and reacted for 10 minutes, then combined with PEG-DA to obtain the desired final concentration. The suspension was then immediately pipetted into a cylindrical Teflon mold (McMaster-Carr, Robbinsville, NJ) and 300 μL was pipetted into a cylindrical mold with a diameter of 11.125 mm and a height of 3 mm for the in vivo sample. The gel was then placed in a 37°C incubator and allowed to gel overnight.

[0239] To confirm the effect of interfacial bonding between thiol groups of HA and MAL on fibers, MAL on fibers was quenched with cysteine ​​to prepare quenched fiber-HA hydrogel composites. Briefly, 1 mg of fibers was dispersed in 1 ml of cysteine ​​solution (in PBS (pH 8.0)), and then the molar ratio of MAL to cysteine ​​was 1:2. After quenching MAL for 12 h at room temperature with gentle shaking, the MAL-quenched fibers were washed five times with 1 ml of distilled water to remove unreacted cysteine ​​and lyophilized.

[0240] Mechanical properties of fiber-HA hydrogel composites: Compression Test. The hydrogel precursor suspension was pipetted into a cylindrical Teflon mold (McMaster-Carr, Robbinsville, NJ) and 200 μL was pipetted into a cylindrical mold with a diameter of 6.35 mm and a height of 6.35 mm for compression testing. The gel was then placed in a 37° C. incubator to gel overnight. The gel was removed from the mold and immediately tested using an Endura TEC mechanical testing instrument (ELF 3200 Series, BOSE ElectroForce, Eden Prairie, MN) by unconfined uniaxial compression between two parallel plates. Samples were compressed to 50% strain and the elastic modulus was determined from the slope of the linear portion of the stress-strain curve from 10% strain to 20% strain. Samples were tested in triplicate and three samples per group were tested to measure the average compressive modulus. To measure the compressive modulus of the rehydrated fiber-HA hydrogel composites, the composites were freeze-dried and rehydrated in 1 ml of PBS (pH 7.4) for 24 h at 37° C. For fatigue testing, the compressed samples were cycled from 0 to 25% strain at 0.1 Hz.

[0241] Rheological testing. The shear storage modulus (G') of various fiber-HA composites was measured using an oscillatory rheometer (ARES-G2 Rheometer, TA Instruments, New Castle, DE) with parallel plates (φ = 8 mm). A frequency sweep was used to monitor the variation of G' from 1 Hz to 10 Hz at a constant strain of 10%.

[0242] Migration of hASCs in fiber-HA hydrogel composites: Human adipose-derived stem cells (hASCs) were cultured in high-glucose DMEM containing 10% FBS, 1% penicillin / streptomycin, and 1 ng / ml bFGF. The culture medium was changed three times a week for optimal growth. To prepare hASC spheroids, 50 μl of hASC solution (5.6 × 10 5hASC spheroids were prepared by pouring 1000 μg / ml of hASC cells into a casting micromolded agarose gel (MicroTissues® 3D Petri Dish® Micromolded Spheroids, 96-well) and incubating at 37° C. for 24 hours with gentle shaking.

[0243] HA and PEGDA were completely dissolved in PBS (pH 7.4) at final concentrations of 4.5 and 2.5 mg / ml for HA and 5.0 mg / ml for PEGDA. Fibers pre-wetted with 20 μl of 50% (v / v) ethanol were completely dispersed in PEGDA at a final concentration of 10.0 mg / ml, and then HA was added into the mixture of fibers and PEGDA. 30 μl of the composite precursor solution was poured into each well of a 96-well tissue culture plate and incubated at 37°C for 1 h for crosslinking to avoid the hASC spheroids reaching onto the surface of the tissue culture plate. Then, 50 μl of the composite precursor solution was poured into each well with 3-5 hASC spheroids. After 1 h of crosslinking at 37°C, 200 μl of fresh medium was added into each well, and the medium was changed every 2-3 days. To observe the migrating cells from the hASC spheroids inside the composites, F-actin and nuclei of hASCs were stained with Alexa Flour® 568 Phalloidin and DAPI, respectively. Briefly, after 4 days of culture, the composites with hASC spheroids were fixed with 100 μl of 4% (v / v) paraformamide at room temperature overnight. Then, after washing three times with PBS (pH 7.4), the composites were incubated with 100 μl of 1% (w / v) BSA (in PBS) at 4°C overnight to suppress non-specific staining, and washed three times with PBS. Subsequently, the composites were incubated with 100 μl of 0.1% (v / v) Triton-X 100 (in PBS) at room temperature for 1 h. After washing three times with PBS, 100 μl of 160 nM Alexa Fluor® 568 Phalloidin was added to each composite and incubated at room temperature for 4 h. Then, after removing the supernatant, the composite was incubated with 100 μl of 0.5 μg / ml DAPI at room temperature for 1 h. After washing three times with PBS, the migrated hASCs were observed using a confocal laser microscope (CLSM, Carl Zeiss LSM780, Germany) at excitation 561 nm and emission 570-600 nm for Alexa Fluor® 568 Phalloidin, and excitation 405 nm and emission 385-420 nm for DAPI.

[0244] In vivo performance of fiber-hydrogel composites: Thiol-containing HA was reconstituted in PBS to 12.5 mg / mL. PEG-DA was dissolved in PBS to 100 mg / mL. MAL-fibers were resuspended in sterile PBS to 25 mg / mL. The fibers were first combined with HA solution and allowed to react for 10 minutes, then combined with PEG-DA to obtain the desired final concentration. The suspension was then immediately pipetted into a cylindrical Teflon mold (McMaster-Carr, Robbinsville, NJ) and 300 μL was pipetted into a cylindrical mold with a diameter of 11.125 mm and a height of 3 mm. The gel was then placed in a 37° C. incubator and allowed to gel overnight. Two formulations were selected to match the 2 kPa stiffness of adipose tissue. The HA-alone formulation was 10 mg / mL PEG-DA and 9 mg / mL HA-SH, and the HA-fiber composite formulation was 5 mg / mL PEG-DA, 5 mg / mL HA-SH, and 12.5 mg / mL dispersed nanofibers.

[0245] To test the biocompatibility of the composite nanomaterial scaffolds, they were implanted under the inguinal fat pads of Sprague-Dawley rats and observed for various lengths of time. Under volatile anesthetic administration, a 1 cm incision was made on both sides just proximal to the inguinal fold. After blunt dissection of the subcutaneous tissue, the inguinal fat pad was exposed. It was elevated with meticulous hemostasis using electrocautery, carefully preserving the nutrient vessels. The scaffold was implanted under the fat pad on the right side of the animal. The left side was left without an implant and served as a sham-operated control. Both sides were closed in a standard layered fashion. The animals were observed for 7, 14, 30 and 90 days. At the time of collection, the animals were sacrificed and the inguinal fat pads with and without the scaffold were exposed and fixed in 4% PFA. Specimens were embedded and sectioned for standard hematoxylin and eosin staining.

[0246] statistical analysis All results are expressed as mean and standard deviation. Statistical significance between pairs of groups was determined by performing one-way ANOVA using SigmaPlot 12.0 software (SPSS); a value of p<0.05 was considered statistically significant.

[0247] Any other suitable methods of making the embodiments of the composite 100 disclosed herein are contemplated herein.

[0248] Example 2: Compression testing of nanofiber-hydrogel composites. For compression testing, fiber-hydrogel samples were formed as cylinders with a diameter of 8.5 mm and a height of approximately 4 mm and cured overnight at 37° C. in a mold. Elastic modulus was determined by compression testing using a Bose EnduraTEC ELF 3200 (Eden Prairie, MN). Samples were uniaxially compressed between two parallel plates and compressed to 50% strain. Elastic modulus was determined by measuring the slope of the first linear region. Two groups of samples with the same hydrogel formulation were tested, with and without fibers. The hydrogel-only sample was formed with 4.5 mg / mL thiol-containing hyaluronic acid (Gylcosan Glycosil) and 10 mg / mL PEG-DA (polyethyl-glycol diacrylate, molecular weight 3350). The fiber-hydrogel composite group had the same hydrogel concentration but additionally contained 6.75 mg / mL PCL nanofibers (with surfaces functionalized with maleimide groups that can readily react with thiol-containing hyaluronic acid).

[0249] A representative stress-strain diagram can be seen in FIG. 2A. The hydrogel only group had a modulus of 320 Pa, while the fiber-hydrogel composite had a higher modulus of 750 Pa. The high stiffness of the fiber-hydrogel composite can be seen by the high stress values ​​at any strain value. The presence of functionalized nanofibers greatly increased the strength and stiffness of the material. Thus, the overall composite structure can be of a stiffness matched to the target tissue, but the hydrogel component can have a lower crosslink density than would be necessary to obtain the same stiffness without the benefit of nanofibers. This should result in a better cellular response at a given implant stiffness.

[0250] The sample groups were then tested by repeated compression (20 cycles) to 25% strain at 0.1 Hz. A representative diagram can be seen in FIG. 2B. This shows that the hydrogels and composites can withstand repeated compression and that the composites are persistently stiffer than the fiberless group.

[0251] Example 3: Cell-material interactions. To test the cellular response to the composite hydrogels, the migration ability of adipose-derived stem cells (ASCs) was tested in different formulations of hydrogels with and without fibers.

[0252] ASCs were transfected to express GFP, and the cells were then formed into spheroid aggregates by seeding overnight in alginate molds made by Microtissues molding. Cells were seeded as spheroids so that cell movement could be better assessed. This is because spheroids are an independent point source where migrating cells can be easily measured. The spheroids were mixed into the hydrogel, then pipetted into a 96-well plate and allowed to harden. The cells were then imaged over the next few days to observe their migration. As the concentration of hyaluronic acid and PEG-DA decreased, the cells were able to migrate progressively further due to their increasing pore size, respectively. At the same hydrogel density (4.5 mg / mL hyaluronic acid and 2.5 mg / mL PEG-DA), the cells were able to migrate better in samples with dispersed nanofibers (12 mg / mL, Figures 3A and 3B) than in samples without (shown in Figures 3C and 3D). This indicates that the presence of functionalized nanofibers not only improved the mechanical properties of the nanofibers but may also help improve cell migration.

[0253] To clearly demonstrate that ASCs were strongly influenced by the presence of nanofibers, ASC spheroids were cultured on sheets of aligned nanofibers without hydrogel, and after 96 h, the cells (green in Figures 3C and 3D) visibly migrated out of the spheroids along the same axis of the aligned nanofibers (shown in Figure 3D).

[0254] Example 4: Tissue compatibility of nanofiber-hydrogel composites. To test the biocompatibility of the composite nanomaterial scaffolds, they were implanted under the inguinal fat pad of Sprague-Dawley rats and observed for various lengths of time. Under volatile anesthetic administration, 1 cm incisions were made bilaterally just proximal to the inguinal folds.

[0255] Figure 5A is a photograph showing the appearance of the nanofiber-hydrogel composite in situ under a rat inguinal fat pad. Figure 5B shows an H&E stained image of a tissue section surrounding the composite collected 2 weeks after implantation. A lobe of mesenchymal cells stained dark pink with eosinophils is shown migrating into the nanomaterial (stained light pink).

[0256] Figure 5C shows an H&E stained image of a tissue section taken from the composite-tissue interface at 4 weeks showing cellular infiltration. The mesenchymal tissue surrounding the implant site is stained deep pink with eosin. The nanomaterial appears light pink. Infiltrating pink mesenchymal cells can be seen at the interface as well as putative adipocytes in evident round lacunae.

[0257] After blunt dissection of the subcutaneous tissue, the inguinal fat pad was exposed. It was elevated, carefully preserving the nutrient vessels, with meticulous hemostasis achieved using electrocautery. The scaffold was implanted under the fat pad on the right side of the animal. The left side received no implant and served as a sham-operated control. Both sides were closed in a standard layered fashion. Animals were observed for 2, 4 and 6 weeks. At harvest time, animals were sacrificed and the inguinal fat pad with and without the scaffold was exposed and fixed in 4% PFA. Specimens were embedded and sectioned for standard hematoxylin and eosin staining. At early time points (2 weeks), mesenchymal cells from the wound environment were found to have infiltrated the material, suggesting that the material had sufficient porosity to allow native cell ingrowth (dark pink staining in Figure 5B).

[0258] Importantly, cellular ingrowth occurred even in the absence of exogenous growth factors. The presence of cells infiltrating the material, rather than simply surrounding it, distinguishes this composite nanomaterial from other alloplastic materials currently in use. The latter materials are surrounded by a fibrous capsule and are therefore less desirable for soft tissue reconstruction. At later time points (4 weeks), cellular ingrowth is even more evident, with follicular regions that may represent nascent adipocyte differentiation (dark pink staining and obvious circles in FIG. 5C).

[0259] Example 5: Design of fiber-HA hydrogel composites The fibers form fibrous structures, which are often found in natural extracellular matrices, and can support cell migration and reinforce the initially poor mechanical properties of the hydrogel. By introducing interfacial bonds between the hydrogel and the fibers (Figure 6A, 6B), the composites are strengthened (Figure 6) without decreasing the average pore size and porosity, which can significantly impair cell migration. It was also predicted that the mechanical properties could be fine-tuned by controlling the density of interfacial bonds between the hydrogel and the fiber surface. Here, surface-functionalized fibers were prepared using maleimide (MAL) to introduce interfacial bonds with thiol-containing hyaluronic acid (HA-SH) (Figure 6). The surface of electrospun poly(ε-caprolactone) (PCL) fibers was treated with O 2 After treatment with plasma to induce free radicals on the surface, poly(acrylic acid) (PAA) was grafted. The carboxyl groups were activated by coupling reagents NHS and EDC, and then N-(2-aminoethyl)maleimide was reacted with the activated carboxyl groups (Figure 13). Subsequently, the MAL-functionalized fibers were introduced into a hydrogel precursor solution composed of HA-SH and PEGDA to fabricate fiber-hydrogel composites. The thiol groups of HA were used to form a gel by reacting with both the MAL groups on the fibers and the DA groups of the PEG linker. Interestingly, the cross-section of the fiber-hydrogel composites showed a fibrous 3D structure with higher porosity compared to the cross-section of the HA hydrogel with similar cross-linking density (Figure 6). The resulting composites showed a uniform distribution of nanofibers throughout both the width and height of the composite, allowing for isotropic reinforcement. Additionally, the rehydrated fiber-HA hydrogel composite showed 99.34% volume recovery after freeze-drying, whereas the HA hydrogel showed 70.17% volume recovery (Figure 6D).

[0260] Example 6: Compressive modulus of fiber-HA hydrogel composites First, it was confirmed that the composites have the greatest stiffness (under shear) when the reactive groups are equal on a molar basis. The thiol groups on the HA can react with either the MAL groups on the nanofibers or the acrylate groups on the PEG-DA. Therefore, the gels showed optimal shear storage modulus when the molar ratio of SH to (DA+MAL) was approximately 1:1. Therefore, this ratio was maintained in all subsequent tests. Uniaxial compression tests were performed on the gels to evaluate the elastic modulus of the HA hydrogel and fiber-HA hydrogel composites (Figure 7). The reinforcing effect of the functionalized nanofibers can be confirmed at compressive stresses at strains up to 50% (Figure 7A). The compressive stress was 3.1 times greater for the 1.0 μm fiber group than the hydrogel-only group, indicating a mechanical reinforcing effect. The 286 nm fiber group showed an even more pronounced reinforcing effect, with the compressive stress being 4.2 times greater at 50% strain. Interestingly, the stiffening effect of the 286 nm fibers was greatly reduced by only 1.3 times compared to the hydrogel when the maleimide groups were quenched prior to gelation, confirming that the interfacial bonding between the fibers and hydrogel is crucial to the reinforcement effect of functionalized fibers. Furthermore, when the 286 nm fibers were not functionalized prior to forming the composite, the reinforcement effect disappeared, resulting in a composite that was barely stiffer than the hydrogel alone. The same reinforcement effect could be seen by incorporating higher concentrations of HA and PEG-DA into the composite to incorporate a stiffer gel (Figure 7). This interfacial bonding also shows a dose response in the stiffening of the composite gel, as the addition of progressively more maleimide groups to the nanofiber surface resulted in a progressively stiffer material, making the importance of interfacial bonding more evident. The composites were also tested for changes in mechanical properties before and after dehydration and rehydration. Mechanical testing was performed under compression on gels with and without functionalized nanofibers of two different maleimide densities. The gels were then freeze-dried and then fully rehydrated and tested again for compression. All samples maintained their stiffness after rehydration, indicating that the composites may be suitable for clinical use as freeze-dried products. The HA-alone gels seemingly maintained their stiffness, but unlike the fiber-containing group, the gels themselves were significantly compacted during the dehydration-rehydration process. The composite gels were also subjected to loading cycles to test for fatigue effects, and representative figures are shown in Figure 10. Under cyclic loading up to 25% strain, the composite gels maintained their stiffness over time and were consistently stiffer than the hydrogel alone.

[0261] Example 7: Shear storage modulus of fiber-HA hydrogel composites In addition to the high compressive modulus, the fiber-HA hydrogel composites exhibited significantly higher shear storage modulus than HA hydrogel alone (Figure 8A). The shear storage modulus of the composites with 286 nm fibers was higher than that of the composites with 686 nm fibers (Figure 8C). It was also confirmed that the shear storage modulus of the composites, as well as the modulus under compression tests, was increased by increasing the surface density of maleimide on the 286 nm fibers (Figure 8D). By introducing fibers with 62 nmol / mg MAL on the surface, the composites showed a 1.3-fold increase in shear storage modulus compared to that of HA hydrogel alone. Furthermore, the shear storage modulus of the composites with 147 nmol / mg MAL on the fibers was 1.8-fold higher than that of the 62 nmol / mg MAL group, showing a clear dose response to the corresponding 2.4-fold increase in MAL surface density on the fibers. When the MAL groups on the fibers were quenched prior to gelation, the shear storage modulus was correspondingly decreased compared to that of the unquenched fibers, similar to that seen in the compression tests. Furthermore, when the frequency was increased to 10 Hz, the shear storage modulus of the composites was maintained, but both the HA hydrogel alone and the composites with quenched fibers showed a decrease in shear storage modulus at 10 Hz compared to 1 Hz. The shear storage modulus of the composites increased with increasing MAL surface density on the fibers regardless of the surface area (diameter) of the fibers, indicating that the previously observed effect of fiber diameter on stiffness may be a function of maleimide density (Figure 8D). A linear regression was obtained (R2 = 0.93) from the correlation between MAL surface density and shear storage modulus. Furthermore, the composites showed a dose response to fiber loading, as the shear storage modulus of the composites increased with increasing weight ratio of functionalized fiber to hydrogel component (Figure 9).

[0262] Example 8: Cell migration in fiber-HA hydrogel composites in vitro We hypothesized that fiber-HA hydrogel composites would enhance cell migration compared to HA hydrogels because (i) the high porosity of the composites with large pore size provides space for cell migration when they have the same mechanical properties, and (ii) the ECM-mimicking fibrous structure within the composites can inherently guide cell migration. Therefore, to demonstrate this hypothesis, we seeded spheroids of human adipose-derived stem cells (hASCs) as model cells and cultured the HA hydrogels and mimic tissue masses inside the composites, then cultured the hASC spheroids for 27 days (Figure 11). ASCs were chosen because of their presence in adipose tissue and their importance in both angiogenesis and adipogenesis. Although the composites have a similar Young's modulus of 1.9 kPa as HA hydrogels, the pore size of the composites is 2.08 times larger than that of HA hydrogels (Figure 16). Thus, it was clearly observed that hASCs migrated three-dimensionally inside the composites (Fig. 11B-11E), because the larger pores could accommodate cell migration and hASCs maintained their spheroid shape in HA hydrogels without any cell migration (Fig. 11A). Notably, cell migration was further enhanced when the composite fibers were modified with cell adhesion peptide RGD (Fig. 11C). However, in an in vivo situation, this difference should be smaller, as additional adhesion triggers would be provided by the diffusion of factors from the local environment into the composites. In some cases, it was observed that some fibers formed slight aggregates during gelation due to hydrophobic interactions between PCL fibers, and cell populations preferentially adhered to the fiber aggregates inside the composites (Fig. 11D and 11E). Furthermore, at the same HA and PEG-DA concentrations (Figure 19), the composites showed enhanced cell migration compared to the fiber-free group, indicating that the nanofibers themselves can inherently help guide cell migration regardless of porosity.

[0263] Example 9: Tissue Response and Host Tissue Infiltration To investigate the therapeutic potential of such composite implants, they were tested in vivo in a rat fat pad model. The implant formulations were formulated to achieve the same initial stiffness of 2 kPa as the composite gel and the target adipose tissue. Thus, the HA-gel only implant formulation had higher concentrations of both thiol-containing HA and PEG-DA to match the stiffness of the fiber-composite group. Despite this higher concentration, the HA-only implants were unable to maintain their shape and volume during the study. Under macroscopic observation after 4 weeks, the HA-only implants were elongated and significantly smaller in volume. Given their macroscopic appearance and lack of histological infiltration, the HA-only system cannot be optimized to promote cell infiltration and maintain a given shape. However, the fiber-gel composite implants maintained their original shape well under macroscopic observation after 90 days in vivo. However, it was noted that histological observations showed that the composite was completely infiltrated to such an extent that it was difficult to determine the interface between the implant and the native tissue.

[0264] A soft tissue defect model has been developed in Lewis rats in which the inguinal fat pad was exposed using microsurgical techniques, elevated, and a preformed composite placed underneath. This well-defined model is ideal for addressing all elements of the Aim 3 hypothesis and at a scale appropriate for R21 testing. While it does not directly demonstrate the ability of such a composite to repair large defects, it establishes proof of concept, confirms all essential functional parts of the composite design, and lays the foundation for a large animal model to test large defect repair in a more clinically relevant model.

[0265] In a pilot study, PCL nanofiber-HA hydrogel composites and HA hydrogels of similar modulus were implanted under the inguinal fat pad of 8-12 week old male Lewis rats (n=3 / time point). Both HA hydrogel and composite groups showed good tissue compatibility at 14 and 30 days after implantation (Figure 12, similar observations at POD 14 and POD 30. POD=days after surgery). Histology at POD 30 did not show a high level of inflammatory response compared to the sham-operated group. H&E and Masson's trichrome staining showed native fat septum formation and cellular infiltration into the composite, capillary formation at the periphery, and regeneration of the glandular and adipocyte portions of native fat (Figure 12). On the other hand, the HA hydrogel control showed no cell infiltration, formation of a thin fibrous tissue sheet, and a foreign body response. The HA hydrogel was prepared to have a pressure of 2 kPa to ensure sufficient mechanical properties. This result emphasizes the importance of scaffold porosity for cell infiltration.

[0266] At early time points (2 weeks), mesenchymal cells from the wound environment were found to have infiltrated the material, suggesting that the material had sufficient porosity to allow natural cellular ingrowth (dark pink staining in FIG. 12). Importantly, cellular ingrowth occurred even in the absence of exogenous growth factors. The presence of cells infiltrating the material, rather than simply surrounding it, distinguishes this composite nanomaterial from other alloplastic materials currently in use. The latter materials are surrounded by a fibrous capsule and are therefore less desirable for soft tissue reconstruction. At later time points (4 weeks), cellular ingrowth is even more evident, with follicular regions that may represent nascent adipocyte differentiation.

[0267] Example 10: Heparin-containing formulations A composite formulation with heparin conjugated to hyaluronic acid was also prepared. This formulation was tested in vivo identically to the preform scaffolds described above. Tissue was harvested at 7, 14, 30 and 90 days (n=3). Many important growth factors, such as bFGF, PDGF and VEGF, have heparin-binding domains. Conjugated heparin may serve two purposes; first, it may bind many of the endogenous growth factors present at the injection site and act as a local reservoir and attraction cue for regenerating tissue. Second, the heparin-containing composite may be used to preload the scaffold with growth factors to better enhance regeneration. Heparin-containing scaffolds showed enhanced angiogenesis compared to non-heparin-containing composite scaffolds at 7 and 14 days, but similar results at 30 and 90 days.

[0268] Example 11: Injectable Formulation The hydrogel-nanofiber composite was also formulated into an injectable variant. 200 μL of the same composition (5 mg / mL thiol-containing HA, 5 mg / mL PEG-DA, 12.5 mg / mL fibers) used in the in vivo preform composite was mixed and partially cured in a syringe for 8-10 minutes. At this point, the composite is a viscous, flowable liquid that can be injected through a surgical needle (Figure 20). Once injected, the composite maintains its shape when inverted and is non-dispersive, shape-retentive and non / low-swelling when immersed in water. To test the biocompatibility of the injectable composite, the suspension is then injected through a 21-gauge needle into the inguinal fat pad of rats. Tissues were then collected at 7, 14, 30, and 90 days (n=3) and analyzed identically to the previous examples. The composite showed extensive cellular remodeling and maintained the dose at day 30 without causing fibrous encapsulation. Early stage adipocytes developing within the composite can be clearly seen.

[0269] Example 12: Consideration of Examples 5 to 11 Hydrogels have been widely studied as filler materials for the regeneration of tissue defects due to their 3D hydration environment and high porosity, which are conducive to cell migration. However, hydrogels have been shown to be inadequate in replacing large volume defects. This is because the relatively weak mechanical properties of hydrogels are insufficient to maintain their volume during the entire period of tissue regeneration, as they can be easily degraded and disintegrated by body fluids and internal and external stresses. To improve the mechanical properties of hydrogels, the main strategies in the art have been (i) increasing the concentration of hydrogel precursors, (ii) increasing the density of the crosslinked network inside the hydrogel, and (iii) introducing reinforcing materials, for example, by embedding hydroxyapatite particles or laminating fiber sheets [Mater Chem Physics, 2008, 107, 364-369, Biomaterials 2006, 27, 505-518, Acta Biomaterialia 2010, 6, 1992-2002]. Unfortunately, such very strong strategies inherently reduce the average pore size and porosity of the resulting hydrogels, preventing cells from migrating into such hydrogels. Therefore, a new mechanism was sought to increase the strength of hydrogels that would still retain high porosity but allow for rapid cell infiltration. Composites were designed by introducing functionalized nanofibers that could contain porosity and increase the overall strength of the hydrogel composite while leaving the hydrogel phase largely intact. The resulting fiber-hydrogel composites are an improvement over previous soft tissue composites due to two key components. First, to increase strength isotropically, the nanofibers needed to be uniformly dispersed in the hydrogel at high loading levels. In the tissue engineering field, electrospun nanofibers have generally been used as flat sheets or mats of fibers. As such, they are typically fabricated into composites by impregnating this mat with a hydrogel precursor solution.

[0270] This severely constrains the dispersion of nanofibers throughout the hydrogel and limits the composite geometry to a 2D sheet or tube. Such geometries are useful for certain applications, e.g., nerve repair or wound dressings, but are poor options for repairing large volume defects. Cryomilling the fiber sheets allowed the average fiber length to be reduced to a length short enough that it could be kept in suspension in aqueous solution. Thus, the sample was then easily pipetted into the hydrogel precursor solution, creating a uniform dispersion of nanofiber fragments throughout the pregelation hydrogel volume. The solution can then be used directly as an injectable formulation or added to a mold to create a structural scaffold gel of any arbitrary geometry, unlike the restricted planar geometry of most electrospun nanofiber meshes. The composite structure of dispersed fibers in the hydrogel also recapitulates the fibrous structure of the extracellular matrix (Figure 6G), providing adhesion sites that can aid cell migration in the composite.

[0271] Second, simply dispersing nanofibers in a hydrogel is insufficient to form a strong composite. The data show that the inclusion of nanofibers by themselves provides very little improvement in the composite modulus, and improvement occurs only when interfacial bonding is introduced. Interfacial bonding is necessary because without the formation of strong bonds between the hydrogel and fiber components, the water and hydrogel components would not be able to cross in front of the fiber components without transferring the load to the stiffer material. Furthermore, such interfaces between different materials can result in delamination and fracture in the composite. Furthermore, the initial hydrophobicity of PCL makes it difficult to disperse in aqueous solutions, as the fibers preferentially clump together and form agglomerates that separate from the suspension. Plasma treatment and subsequent functionalization with carboxylic acid and amine groups greatly increase the hydrophilicity of the fibers, allowing for dispersion. This dramatic increase in mechanical properties only occurred when interfacial bonding was present between the maleimide groups on the fiber surface and the thiol groups on the hyaluronic acid molecules. This covalent strength bond transfers load to the fiber more efficiently during compression or tension, resulting in a stiffer and stronger material. Furthermore, the composites show a strong trend for increased modulus with increasing maleimide density, highlighting its importance in the strength-enhancing mechanism as well as the fine-tunability of reinforcement.

[0272] In this study, it was confirmed that the mechanical properties of fiber-hydrogel composites can be fine-tuned by various factors, such as the total surface area of ​​the fibers, the density of functional maleimide groups on the fiber surface, and the amount of fibers loaded in the hydrogel. First, composites with small diameter fibers exhibited higher compression and shear storage moduli than those with large diameter fibers (Figures 7A and 8C). Similarly, in the literature, a single ultra-high molecular weight polyethylene (UHMWPE) fiber (about 25 μm) plasma-activated with glutaraldehyde showed an approximately 2.36-fold increase in interfacial shear strength in poly(vinyl alcohol) hydrogels compared to 60 bundles of UHMWPE fibers [Acta Biomaterialia 2014,10,3581-3589]. Therefore, it can be considered that reducing the fiber diameter and therefore increasing the specific surface area of ​​the fibers may be effective in improving the mechanical properties of the composites. However, each fiber group had only a small difference in MAL surface density on the fiber (approximately 10-15 nmol / mg), so the effect of the surface area of ​​the fiber alone cannot be conclusively determined. Therefore, secondly, composites were fabricated with fibers of the same diameter but with varying MAL surface density on the fiber (Figure 8). The compressive and shear storage moduli of the composites increased with increasing MAL surface density on the fiber. Composites without interfacial bonding were confirmed to show only a slight enhancement in compressive modulus (Figure 7) by using fibers modified by the PAA step (carboxyl groups on the fiber) but without a further MAL conjugation step. The importance of interfacial bonding was further confirmed by quenching the MAL groups on the fiber with cysteine ​​prior to gelation. Cysteine ​​conjugates with the maleimide groups and prevents interfacial bonding between the fiber and the hydrogel. This leads us to address the effect of interfacial bonding precisely because we processed the fibers identically to the interfacially bonded group but in a different way. Interestingly, the mechanical properties of the composites with MAL-quenched fibers were dramatically reduced (Figures 7A and 8B), and the MAL-quenched fibers exhibited a lower compressive modulus than the HA hydrogel alone at a HA concentration of 10 mg / ml (Figure 7).It may be possible that MAL-quenched fibers weakened the overall composite by easily causing delamination at the fiber-hydrogel interface, as was the case in the previous study [Acta Biomaterialia 2014,10,3581-3589]. Also, fibers without functional groups may act as foreign bodies that inhibit gelation compared to pure hydrogels that are composed of one component or that are completely free of foreign bodies during gelation [JMC B 2015,DOI:10.1039 / C3TB21830A,Journal of Biomedical Materials Research Part A 2010,95(2),564-573]. Moreover, composites with different MAL surface densities confirmed a significant correlation between the shear storage modulus and the density of interfacial bonds (Figure 8C). This study provides strong evidence that the mechanical properties of hydrogels are enhanced and fine-tuned by interfacial bonds. Third, the shear storage modulus of the composites increased with increasing fiber to hydrogel weight ratio (Figure 9). It confirmed that the weight ratio is another variable that can be used to fine-tune the mechanical properties of fiber-hydrogel composites. However, it was confirmed here that with increasing fiber loading, the increase in shear storage modulus started to plateau and even slightly decreased for weight ratios above 0.6. One possibility for this saturation effect could be that the density of interfacial bonds in the composites was reduced to what extent the excess fibers with MAL reacted with the thiol groups of the bulk fraction of HA and prevented the reaction of PEGDA for gelation. Considering that the highest shear storage modulus of HA hydrogels was obtained with equimolar amounts of HA-SH and PEGDA functional groups, and that the shear storage modulus decreased with excess amounts of either HA-SH or DA (Figure 14A), the excess MAL on the fibers might destroy the SH-DA bonds inside the composites with increasing amounts of fibers.

[0273] In general, the implanted biomaterial must withstand numerous internal and external stresses during the regeneration of tissue defects. Although the stresses are not severe and continuous, to mimic such stresses, stress resistance tests were performed under cyclic conditions and high frequency (10 Hz) (Figures 10 and 8). Both HA hydrogel and fiber-HA hydrogel composites withstood repeated compressive strain without any damage or loss of mechanical strength. Notably, the composite with interfacial bonding retained its shear storage modulus at a frequency of 10 Hz, while the shear storage modulus of HA hydrogel and composite without interfacial bonding decreased at 10 Hz. This trend indicates that interfacial bonding with dispersed fibers is crucial for enhancing the mechanical properties of the composites. Also, the fiber-HA composites maintained their dimensions and Young's modulus after being subjected to freeze-drying and subsequent rehydration, whereas the HA-only gels substantially shrank under the same process (Figures 6C and 10). This maintenance of shape, volume and stiffness after dehydration and rehydration is an important feature for the clinical application of this technology, as having a lyophilized form of the composite may allow for easier sterilization and storage of the commercial product.

[0274] An ideal implantable scaffold for soft tissue reconstruction would be one that could immediately fill the defect void, but also serve as a substrate for the body's own cells to grow into the scaffold, proliferate and differentiate into the appropriate tissue phenotype, and ultimately replace the artificial scaffold with normal healthy tissue. Therefore, it is critical that the appropriate cells would be able to migrate within the hydrogel or composite scaffold. To investigate the potential of the appropriate cell type to migrate within the scaffold, hASC spheroids were seeded within HA hydrogel and fiber-HA hydrogel composites, and the cell migration was evaluated. Within HA hydrogel alone, hASC could not migrate because the HA hydrogel was too soft to provide traction for cell migration (Figure 11A) [Biomaterials 2015,42,134-143]. Interestingly, although the shear storage modulus of the composite was similar to that of HA hydrogels, hASCs were able to significantly migrate away from the spheroids inside the composite (Figure 11). One hypothesis is that the fibers inside the composite may provide adhesion sites to guide cell migration, similar to the fibrillar components of the native ECM of adipose tissue. Previously, it was shown that aligned and random fibers can be crucial factors for cell adhesion, proliferation, differentiation and migration in various cell types [Biomaterials 2005,26,2537-2547 / 2006,27,6043-6051 / 2009,30,556-564 / 2010,31,9031-9039,Acta Biomaterialia 2013,9,7727-7736]. In particular, cells were observed to recognize the fibers as a guiding matrix, with their cytoskeleton aligned and following along the underlying fibers [Biomaterials 2006,30,6043-6051 / 2009,30,556-564]. However, the diameter of the fibers within the composite did not affect migrating cells, as they migrated robustly within composites with either 1000 nm or 286 nm nanofibers (Figure 19).

[0275] The effects of porosity and cell migration seen in bench-top testing and in vitro cultured cells translated into striking differences during in vivo testing of the composites. Fiber-free hydrogels formulated to mimic fat, with a stiffness of 2 kPa, had too low porosity for cell infiltration. The cellular response was that the hydrogel was surrounded by a thick collagen layer, without infiltration or remodeling typical of a foreign body response. However, the nanofiber-hydrogel composites had sufficient porosity to encourage cell ingrowth, vascularization, and cellular remodeling, without a foreign body response. This offers the prospect of permanently filling large volume defects in the body with what will eventually become the body's own tissue. The results were even more pronounced with the injectable formulation, which could form a tighter interface with the host tissue, showing signs of robust adipogenesis.

[0276] Conclusion: Dispersion of functionalized nanofibers in hydrogels results in the formation of composite structures that combine the strength of the two components. Interfacial bonding between the nanofibers and hydrogel components is crucial to creating a strong composite while maintaining high porosity and pore size to facilitate tissue and cell ingrowth. The properties of the resulting composites can be easily fine-tuned by varying fiber diameter, fiber loading level, maleimide density level, and loading level of the hydrogel components. This allows for low cross-linking and high porosity at a targeted overall stiffness to enhance cell infiltration and subsequent tissue remodeling. The fibers themselves may also directly improve cell migration by providing adhesion sites similar to those found in the native ECM. The resulting composite implants can be fine-tuned to match the stiffness of native adipose tissue while retaining permeability for cell infiltration and remodeling. This novel composite is strong enough to be readily filled into voluminous defects of any arbitrary shape. The composite implant thus serves as a permissive scaffold for the body's own cells to infiltrate into the composite, form blood vessels, and differentiate into cells such as adipocytes. The scaffold slowly degrades during tissue remodeling until the initial defect void is completely replaced by normal healthy tissue. This composite structure holds great promise for the future of reconstructive and cosmetic surgery.

[0277] Equivalent It will be understood that the detailed examples and embodiments described herein are presented for illustrative purposes only and are not to be considered as limiting the present invention in any way. In view of this, various modifications or changes will be suggested to those skilled in the art and are considered to be within the spirit and scope of this application and are included in the appended claims. For example, the relative amounts of the components may be changed to optimize the desired effect, additional components may be added, and / or one or more of the components described may be replaced with similar components. Further advantageous features and functionalities associated with the systems, methods, and process steps of the present invention will be apparent from the appended claims. Moreover, those skilled in the art will recognize or be able to ascertain, using nominal experimentation, many equivalents to the specific embodiments of the invention described herein. Such equivalents are intended to be encompassed within the scope of the following claims.

Claims

1. 1. A composite scaffold comprising polymeric fibers having an average diameter of about 100 nm to about 8000 nm covalently bonded to a hydrogel material, the ratio of said fibers to said hydrogel material being about 1:10 to about 10:1 on a component weight basis, or about 1 to 50 mg / mL on a concentration basis.

2. 10. The scaffold composite of claim 1, wherein the polymeric fibers comprise a biocompatible, biodegradable polyester.

3. 10. The scaffold composite of claim 1, wherein the polymeric fibers comprise polycaprolactone.

4. The scaffold composite of claim 1 , wherein a hydrogel material is present in said composite of functional networks.

5. 10. The scaffold composite of claim 1, wherein the ratio of fiber to anhydrous hydrogel material is from about 1:10 to about 10:

1.

6. 10. The scaffold composite of claim 1, wherein the polymeric fibers comprise non-woven polymeric fibers.

7. 10. The scaffold composite of claim 1, wherein the polymeric fibers comprise polycaprolactone electrospun fibers.

8. 10. The scaffold composite of claim 1, wherein the polymeric fibers comprise a synthetic polymeric material comprising poly(lactic-co-glycolic acid), polylactic acid and / or polycaprolactone, or a combination thereof.

9. 10. The scaffold composite of claim 1 formulated to be substantially biocompatible.

10. 10. The scaffold composite of claim 1, wherein the polymeric fibers comprise a biological polymeric material selected from the group consisting of silk, collagen and chitosan, or a combination thereof.

11. The scaffold composite of claim 1 , wherein the hydrogel material comprises hyaluronic acid.

12. 10. The scaffold composite of claim 1, wherein the hydrogel material comprises poly(ethylene glycol), collagen, dextran, elastin, alginate, fibrin, alginate, hyaluronic acid, poly(vinyl alcohol), or derivatives thereof, or combinations thereof.

13. The scaffold composite of claim 1 , wherein the hydrogel material comprises an engineered tissue extracellular matrix.

14. 14. The scaffold complex of claim 13, wherein the processed tissue extracellular matrix is ​​derivable from adipose tissue.

15. 10. The scaffold composite of claim 1 comprising non-woven polycaprolactone fibers.

16. 10. The scaffold composite of claim 1, wherein the hydrogel material comprises hyaluronic acid substantially covering at least a portion of the outer surface of the polycaprolactone fibers.

17. The scaffold composite of claim 1 , wherein a hydrogel material is bonded to an outer surface of said polymer fibers.

18. 20. The scaffold composite of claim 17, further comprising a crosslinkable moiety present in an amount effective to introduce bonds between the polymer fibers and the hydrogel material.

19. a plurality of pores present on or within the surface, the pores extending over 1 cm of the surface; ２ 2. The scaffold complex of claim 1, wherein the pores are present at a concentration of at least about 50 pores per surface, and at least 80% of the pores on the surface have an average pore diameter that is at least about 5 microns.

20. 10. The scaffold composite of claim 1, further comprising a cross-linkable moiety present in an amount effective to induce cross-linking between the polycaprolactone fibers and the hyaluronic acid.

21. 10. The scaffold complex of claim 1, which promotes tissue growth and cell infiltration when implanted into a target tissue present in a human subject.

22. 10. The scaffold complex of claim 1, which is substantially biodegradable when implanted in human tissue.

23. 10. The scaffold complex of claim 1, which is substantially non-biodegradable when implanted in human tissue.

24. 10. The scaffold complex of claim 1, further comprising a therapeutic agent selected from a cell, a small molecule, a nucleic acid, and a polypeptide.

25. An implantable biomaterial comprising the scaffold complex of claim 1.

26. 26. The implantable material of claim 25, which is substantially acellular and / or substantially polypeptide-free.

27. 26. The implantable material of claim 25, formulated for administration by injection.

28. 26. The implantable material of claim 25, formulated for subcutaneous administration.

29. A kit comprising an implantable material according to any one of claims 25 to 28.

30. A medical device for maintaining the shape of tissue in a subject undergoing a surgical procedure, comprising an effective amount of a scaffold complex described in claim 1 or an implantable material described in claim 25 to result in maintenance of the shape of the tissue when administered to a subject.

31. 1. A method for preparing an implant for tissue or cartilage repair, comprising the steps of: providing an acellular three-dimensional scaffold comprising polymeric fibers oriented to provide a plurality of pores, wherein at least a portion of the polymeric fibers are crosslinked to other polycaprolactone fibers; disposing a composition comprising a hydrogel material onto the polymeric fibers to form a composite; and reacting or stabilizing the composite to form a stabilized implant, thereby preparing the implant.

32. 32. The method of claim 31 , wherein the tissue comprises soft tissue.

33. 1. A method for preparing an implant for tissue or cartilage repair, comprising the steps of: providing an acellular three-dimensional scaffold comprising polymeric fibers oriented to provide a plurality of pores; disposing a composition comprising a hydrogel material on the polymeric fibers to form a composite; and reacting or stabilizing the composite to form a stabilized implant, wherein at least a portion of the polymeric fibers are crosslinked to the hydrogel material.

34. 34. The method of claim 33, wherein the polymeric fiber comprises polycaprolactone.

35. 34. The method of claim 33, wherein the hydrogel material comprises hyaluronic acid.

36. 34. The method of claim 33, wherein the three-dimensional scaffold comprises reactive polycaprolactone fibers.

37. 1. A method for preparing an implant for tissue or cartilage repair, comprising the steps of: providing an acellular three-dimensional scaffold comprising polymeric fibers oriented to provide a plurality of pores; disposing a composition comprising a hydrogel material on the polymeric fibers to form a composite; and reacting or stabilizing the composite to form a stabilized implant, wherein at least a portion of the polymeric fibers are crosslinked to the hydrogel material.

38. 13. A method for eliminating a tissue defect resulting from trauma or surgical intervention, comprising expanding a tissue comprising the tissue, wherein expanding the tissue comprises embedding an effective amount of a structural scaffold complex according to claim 1 into the tissue, thereby expanding it.

39. A method for reducing or reversing tissue loss due to an age-related disease, disorder or condition, comprising expanding a tissue comprising the tissue, wherein expanding the tissue comprises implanting an effective amount of a structural scaffold complex described in claim 1 into the tissue, thereby expanding it.

40. 40. The method of claim 38 or 39, wherein the tissue defect comprises pleural tissue, muscle tissue, skin, or a combination thereof.

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