Collagen fiber hydrogel composites and methods

Abstract

Among other things, provided herein are collagen-based hydrogel composites, compositions thereof, and their uses for soft tissue repair or cosmetic and reconstructive purposes. TIFF2026502902000007.tif42128

Description

[Technical Field]

[0001] This application claims the benefit of priority to U.S. Provisional Application No. 63 / 435,403, filed December 27, 2022, which is incorporated herein by reference in its entirety.

[0002] Field Provided herein are soft tissue devices, compositions, and uses, including hydrogel composites, for, inter alia, fat grafting and soft tissue regeneration. [Background technology]

[0003] background Soft tissue defects resulting from trauma, tumor resection, or congenital anomalies have been treated using fat grafting. For example, in 2019, there were 600,000 fat grafting procedures performed on the face worldwide, with substantially more performed on other anatomical regions. Tissue regrowth can require a suitable matrix for cells to attach, migrate, proliferate, differentiate, and organize into new tissue. For example, natural extracellular matrices (ECMs) have been used at repair sites. However, no materials currently exist that can withstand the mechanical and structural loads of cells during healing. Summary of the Invention

[0004] overview In one aspect, a polymer (e.g., hyaluronic acid (HA))-based fiber-hydrogel composite is provided in which a porous hydrogel is combined with dispersed fibers to mimic the fibrous structure of the body's extracellular matrix. In a preferred aspect, the composite can instantly fill voids within the body, promote cellular infiltration, macrophage polarization to a pro-healing phenotype, promote angiogenesis, and / or enable permanent soft tissue remodeling. In a preferred aspect of the present disclosure, a hydrogel composite comprising fibers (e.g., nanofibers or microfibers) is provided that can mimic the natural extracellular matrix.

[0005] In one aspect, the present disclosure provides a fiber-hydrogel composite comprising fibers (e.g., nanofibers or microfibers) or microfibers comprising one or more extracellular matrix proteins (ECM), a hydrogel material such as hyaluronic acid (HA), and a crosslinker. The hydrogel material (e.g., HA) is suitably covalently bonded to the fibers by the crosslinker to form a composite network.

[0006] In certain aspects, the present disclosure provides a fiber-hydrogel composite comprising fibers (e.g., nanofibers or microfibers) or microfibers comprising one or more extracellular matrix proteins (ECM), a hydrogel material such as hyaluronic acid (HA), and a crosslinker, wherein the hydrogel material (e.g., HA) is bonded to the fibers by the crosslinker to form a composite network.

[0007] Other suitable materials that can be utilized as hydrogel materials in the composites and compositions of the present invention include materials that have one or preferably multiple functional groups (e.g., hydroxyl groups, amine groups, and thiol groups) that can react with a crosslinker and can be formed into a gel. Hydrogel materials that have hydroxyl groups available for reaction with a crosslinker can be preferred in various aspects.

[0008] Exemplary suitable and preferred hydrogel materials include hyaluronic acid (HA), collagen, chitosan, alginate, polyvinyl acetate (PVA), gelatin, polyethylene glycol (PEG) and other glycol ethers, chondroitin sulfate, or cellulosic materials.

[0009] In certain aspects, the hydrogel component of the compositions of the present invention (fiber-hydrogel composites) comprises one or more of hyaluronic acid (HA), collagen, chitosan, alginic acid, polyvinyl acetate (PVA), gelatin, polyethylene glycol (PEG) and other glycol ethers, and cellulose materials. In certain aspects, the hydrogel component of the compositions of the present invention comprises hyaluronic acid (HA) and one or more of collagen, chitosan, alginic acid, polyvinyl acetate (PVA), gelatin, polyethylene glycol (PEG) and other glycol ethers, chondroitin sulfate, or cellulose materials. In certain aspects, the hydrogel component of the compositions of the present invention comprises one or more of hyaluronic acid (HA), chitosan, alginic acid, polyvinyl acetate (PVA), gelatin, polyethylene glycol (PEG) and other glycol ethers, chondroitin sulfate, or cellulose materials. In certain aspects, the hydrogel component of the compositions of the invention comprises hyaluronic acid (HA) and one or more of chitosan, alginate, polyvinyl acetate (PVA), gelatin, polyethylene glycol (PEG) and other glycol ethers, chondroitin sulfate, or a cellulose material. In certain aspects, the hydrogel component of the compositions of the invention comprises, consists essentially of, or consists of one or more hyaluronic acid (HA) materials.

[0010] In suitable aspects, the hydrogel component and the collagen fiber component are distinct materials, i.e., the hydrogel component and the collagen fiber component differ in composition, molecular weight, among other differences. In certain aspects, the hydrogel material does not include collagen material.

[0011] In certain aspects, the composition or composite includes one or more collagen materials. In one aspect, a fiber-hydrogel composite is provided that includes fibers (e.g., nanofibers or microfibers) or microfibers that include (1) one or more collagen materials, (2) one or more hydrogel materials, such as hyaluronic acid (HA) or other materials, and (3) a crosslinker. In a preferred aspect, the hydrogen material (e.g., HA) is bonded to the fiber by the crosslinker to form a composite network.

[0012] In certain embodiments, the collagen of the composites or compositions (fiber-hydrogel composites) of the invention comprises recombinant collagen, human collagen, and / or recombinant human collagen, respectively. In various aspects, the collagen is selected from the group consisting of collagen types I, II, III, IV, V, VI, VII, VIII, IX, X, XI, XII, XIII, XIV, XV, XVI, XVII, XVIII, XIX, XX, XXI, XXII, XXIII, XXIV, XXV, XXVI, and XXVII. In some aspects, the collagen is collagen of one collagen type, free of any other collagen type; in other aspects, the collagen is a specified or unspecified mixture of more than one collagen type.

[0013] In aspects, the collagen material is present in the composition (fiber hydrogel composite) in an amount that provides at least about 2 mg / mL or 0.2 wt.% to 200 mg / mL or at least about 20 wt.% or 20 wt.% of one or more collagen materials based on the total weight of the composite of the composition at gelation. In aspects, a suitable composite or composition may contain from about 1 mg / mL or 0.1 wt.% to about 150 wt.% of one or more collagen materials based on the total weight of the composite of the composition in the finished gel after fabrication.

[0014] The inventors have also found that if the amount of hyaluronic acid (or other hydrogel material) in the composition exceeds a certain level, the composition can gel more efficiently. Thus, in certain preferred compositions, hyaluronic acid is present in an amount greater than 1 wt.% at gelation based on the total weight of the composition, including amounts of HA of up to or at least 1.1, 1.2, 1.3, 1.4, 1.5, 1.6, 1.7, 1.8, 1.9, 2.0, 2.2, 2.4, 2.5, 2.6, 2, 8, 3, 3, 2, 3, 4, 3.6, 3.8, 4.0, 5.0, 6.0, 7.0, 8.0, 9.0, 10.0, 11.0, or 12.0 wt.% based on the total weight of the composition at gelation, or an equivalent concentration of HA in the finished gel after manufacture with appropriate swelling during manufacture of 0-fold (no conversion) to 10-fold (thus reducing the concentration to 1 / 10 of the gelation concentration).

[0015] The inventors have also found that different crosslinking agents can provide enhanced gelation effects at different amounts. Thus, when using a vinyl sulfone agent (e.g., divinyl sulfone (DVS)), a strong gel can be produced even when the HA concentration at gelation is reduced to about 2.0% based on the total weight of the composition. When using a glycidyl ether (e.g., 1,4-butanediol diglycidyl ether (BDDE)) crosslinker, a larger amount of hydrogel material such as HA may be required at gelation, such as at least about 8, 10, or 12 wt% HA based on the total weight of the composition, to provide efficient gelation.

[0016] In one aspect, the present disclosure provides a soft tissue device comprising the fiber-hydrogel composite described herein.

[0017] In one aspect, the present disclosure provides an implant for promoting vascularization comprising the fiber-hydrogel composite described herein.

[0018] In one aspect, the present disclosure provides an implant for adipose tissue formation comprising the fiber-hydrogel composite described herein.

[0019] In one aspect, the present disclosure provides an implant for vasculature formation comprising the fiber-hydrogel composite described herein.

[0020] In one aspect, the present disclosure provides a kit comprising the fiber-hydrogel composite described herein.

[0021] The present disclosure may also provide a simpler manufacturing process for fiber-hydrogel composites that does not require a spacer in the crosslinker. With enhanced thermal stability in both the hydrogel and fiber components, the process may include autoclaving, allowing for terminal sterilization of the product, improving cost, risk, and regulatory burden. The high thermal stability of the gels of the present invention also provides storage stability at ambient temperatures. In several aspects, the compositions are effective against therapeutic targets.

[0022] In one aspect, the present disclosure provides a method of producing the fiber-hydrogel composite described herein, comprising contacting a crosslinker with fibers comprising one or more extracellular matrix proteins (ECM) and one or more hydrogel materials, such as hyaluronic acid (HA), to obtain a fiber-hydrogel composite, wherein the one or more hydrogel materials, such as HA, are bonded to the fibers and to itself by the crosslinker to form a composite network.

[0023] In one aspect, the present disclosure provides a method of forming adipose tissue formation in a subject, comprising administering to the subject a fiber-hydrogel composite described herein.

[0024] In one aspect, the present disclosure provides a method of delivering cells or tissues in a subject, comprising encapsulating one or more cells or tissues in a fiber-hydrogel composite described herein to form a suspension; and applying the suspension to a target site in the subject.

[0025] In one aspect, the present disclosure provides a method of delivering adipose tissue in a subject, comprising encapsulating one or more adipose tissues and a fiber-hydrogel composite described herein to form a suspension; and applying the suspension to a target site in the subject.

[0026] In one aspect, the present disclosure provides a method of delivering a pharmaceutical agent in a subject, comprising combining a pharmaceutical agent and a fiber-hydrogel composite described herein to form a mixture; and applying the mixture to an intended delivery site.

[0027] In one aspect, the present disclosure provides a dermal filler comprising the fiber-hydrogel composite described herein.

[0028] In one aspect, the present disclosure provides a soft tissue device comprising the fiber-hydrogel composite described herein.

[0029] In one aspect, the present disclosure provides an implant comprising the fiber-hydrogel composite described herein.

[0030] In one aspect, the present disclosure provides a kit including the fiber-hydrogel composite described herein and an applicator.

[0031] As will be understood, the terms fiber hydrogel composite, composition, or composition and composite of the present invention are used interchangeably.

[0032] Other aspects of the invention are disclosed below.

[0033] Where applicable or not specifically denied, it is contemplated that any one of the embodiments described herein can be combined with any other one or more embodiments, even if the embodiments are described under different aspects of the invention. [Brief explanation of the drawings]

[0034] [Figure 1A] Figure 1 depicts a schematic diagram of the preparation of an example of a fiber-hydrogel composite by crosslinking hyaluronic acid (HA) and collagen nanofibers (electrospun and pre-crosslinked) with divinyl sulfone (DVS) or 1,4-butanediol diglycidyl ether (BDDE) to generate a nanofiber-hydrogel composite (NHC). [Figure 1B] Morphological differences between hyaluronic acid (HA) with a storage modulus (G') equivalent to 100 Pascals (Pa) and NHC with a G' equivalent to 250 Pa (G' = 100 Pa) are shown. [Figure 1C] 1 shows a scanning electron microscope (SEM) photograph of electrospun bovine type I collagen nanofibers with an average diameter of approximately 600 nm. [Figure 1D] This indicates that NHC can be injected through a 27-gauge needle. [Figure 1E] 1 shows a scanning electron microscope (SEM) image of a collagen fiber-hyaluronic acid hydrogel composite showing an intimately connected fibrous structure in the hydrogel phase and a porous structure. Scale bar: 50 microns. [Figure 2A] Figure 1 shows the crosslinking kinetics of HA hydrogels with 3.52 w / v% DVS at pH 12.4, 12.7, 13.0, and 13.3 prepared with different NaOH concentrations. [Figure 2B]Figure 1 shows the crosslinking kinetics of HA hydrogels at pH 12.7 using different DVS concentrations (2.34 w / v%, 3.52 w / v%, and 4.68 w / v%). [Figure 2C] HA hydrogels and NHCs were prepared with different DVS crosslinker concentrations (2.34 w / v% and 3.52 w / v%) and fiber amounts (0, 1, and 3 w / v%), with or without interfacial bonding. The reaction pH was maintained at pH 12.7 (n = 3–9). Statistical significance was calculated by two-way ANOVA with Dunnett's post-hoc test. Asterisks indicate comparisons between each material group. ns, P > 0.05, * P < 0.05, ** P < 0.01, *** P < 0.001, **** P < 0.0001. Data are presented as mean ± SEM. [Figure 2D] HA hydrogels and NHCs with different fiber lengths sorted through cell strainers of different sizes (40 μm, 100 μm, and no filtration) are shown. [Figure 2E] Figure 1 shows the G' range of HA hydrogels and NHCs at different crosslinker concentrations. [Figure 2F] G' of different sets of HA hydrogels and NHCs before and after specialization. [Figure 2G] G' of different sets of HA hydrogels and NHCs before and after autoclaving. [Figure 3A]Figures 3A-3D show the enhanced adhesion and migration of human adipose-derived stem cells (hADSCs) in NHCs in both 2D and 3D settings. Cell morphology was visualized by staining F-actin with Alexa Fluor 568 phalloidin. Cell nuclei were stained with 4',6-diamidino-2-phenylindole (DAPI). Statistical significance was calculated by two-way ANOVA with Dunnett's post-hoc test. Asterisks indicate comparisons between each material group. ns, P > 0.05, * P < 0.05, ** P < 0.01, *** P < 0.001, **** P < 0.0001. Data are presented as mean ± SEM. Figure 3A represents the results of live / dead staining showing the percentage of viable hADSCs in 250 Pa NHCs and 100 and 250 Pa HA hydrogel controls at days 1, 4, and 7 (n = 3–4). [Figure 3B] Figures 3A-3D show the enhanced adhesion and migration of human adipose-derived stem cells (hADSCs) to NHCs in both 2D and 3D cultures. Cell morphology was visualized by staining F-actin with Alexa Fluor 568 phalloidin. Cell nuclei were stained with 4',6-diamidino-2-phenylindole (DAPI). Statistical significance was calculated by two-way ANOVA with Dunnett's post-hoc test. Asterisks indicate comparisons between each material group. ns, P > 0.05, * P < 0.05, ** P < 0.01, *** P < 0.001, **** P < 0.0001. Data are shown as mean ± SEM. Figure 3B shows the percentage of adherent cells in HA hydrogels and NHCs with hADSCs, demonstrating cell adhesion after 4, 24, and 48 hours of culture (n = 4). [Figure 3C]Figures 3A-3D show the enhanced adhesion and migration of human adipose-derived stem cells (hADSCs) in NHCs in both 2D and 3D settings. Cell morphology was visualized by staining F-actin with Alexa Fluor 568 phalloidin. Cell nuclei were stained with 4',6-diamidino-2-phenylindole (DAPI). Statistical significance was calculated by two-way ANOVA with Dunnett's post-hoc test. Asterisks indicate comparisons between each material group. ns, P > 0.05, * P < 0.05, ** P < 0.01, *** P < 0.001, **** P < 0.0001. Data are shown as mean ± SEM. Figure 3C shows the enhanced migration and spreading of cells inside 250 Pa NHCs at day 7 using hADSC cell spheroids. Scale bar, 50 µm. [Figure 3D] Figures 3A-3D show the enhanced adhesion and migration of human adipose-derived stem cells (hADSCs) on NHCs in both 2D and 3D. Cell morphology was visualized by staining F-actin with Alexa Fluor 568 phalloidin. Cell nuclei were stained with 4',6-diamidino-2-phenylindole (DAPI). Statistical significance was calculated by two-way ANOVA with Dunnett's post-hoc test. Asterisks indicate comparisons between material groups. ns, P > 0.05, * P < 0.05, ** P < 0.01, *** P < 0.001, **** P < 0.0001. Data are presented as mean ± SEM. Figure 3D shows the cell spreading and migration of hADSCs on fiber mats, which is not observed in the non-fiber hydrogel control. [Figure 4-1] Figure 4A represents macroscopic images of NHC gels during sample collection, demonstrating in vivo NHC retention at postoperative days (POD) 7, 14, 56, and 180 (scale bar = 1 cm). Figure 4B represents quantitative measurements of the retained shape of 100 and 250 Pa HA hydrogels and 250 Pa NHCs using calipers at POD 7, 14, 56, and 180 with an initial injection of 200 μL (n = 3). [Figure 4-2] Figure 4C depicts enhanced host cell infiltration in 250 Pa NHCs in vivo at POD 7. Scale bar, 500 μm. [Figure 4-3] Figure 4D depicts signs of neotissue, vascular, and adipose tissue formation in 250 Pa NHCs at POD 56. Scale bar, upper image = 500 µm; lower image = 100 µm. [Figure 4-4] Figure 4E shows a quantitative analysis of the invasion area, defined as the percentage of the area invaded by host cells within the entire injected matrix (n = 3). Figure 4F shows a quantitative analysis of the host cell density (n = 3). [Figure 5-1] Figure 5A shows a tile-scan immunofluorescence image demonstrating enhanced infiltration and expression of CD68+ pan-macrophages (red), CD38+ M1 macrophages, and CD163+ M2 macrophages at POD 7. Scale bar, tile scan = 200 µm. Figure 5B shows a magnified Z-stack immunofluorescence image demonstrating enhanced infiltration and expression of CD68+ pan-macrophages (red), CD38+ M1 macrophages, and CD163+ M2 macrophages at POD 7. Scale bar, Z-stack image = 50 µm. Figures 5C-5D show relatively low expression of CD38+ M1 macrophages (Figure 5C) and high expression of CD163+ M2 macrophages (Figure 5D) in 250 Pa NHCs at POD 56. Scale bar, 200 µm. [Figure 5-2]Figures 5E–5H show quantitative depictions of infiltrated CD68+ pan-macrophages (Figure 5E), CD38+ M1 macrophages (Figure 5F), CD163+ M2 macrophages (Figure 5G), and the ratio of M2 to M1 macrophages (Figure 5H) by analyzing tile scan images (n = 3) at POD 56. Statistical significance was calculated by two-way ANOVA with Dunnett's post-hoc test. Asterisks indicate comparisons between each material group. ns P > 0.05, * P < 0.05, ** P < 0.01, *** P < 0.001, **** P < 0.0001. Data are shown as mean ± SEM. Figure 5I shows a graph of gene expression levels for the tested genes separated by category: macrophage (CD68, Nos2, CD86, Arg1, CD163, CD38), inflammatory (TNF-α, IL1b), and anti-inflammatory (TGF-β, IL10, IL13). Expression is normalized to d7 250 Pa HA. ​​Samples were pooled together and tested in triplicate (n = 3). Statistical significance was calculated by two-way ANOVA with Dunnett's post-hoc test. Asterisks indicate comparisons between each material group. ns P > 0.05, * P < 0.05, ** P < 0.01, *** P < 0.001, **** P < 0.0001. Data are presented as mean ± SEM. [Figure 6-1] Figure 6A shows the ingrowth of host endothelial cells and smooth muscle in injected 100 and 250 Pa HA hydrogels and 250 Pa NHCs at POD 7, 14, and 56. Endothelial cells were stained with RECA-1, smooth muscle was stained with α-SMA (green), and all infiltrating cell nuclei were stained with DAPI. Scale bars, 200 μm for tile scans (left); 50 μm for Z-stack images (right). [Figure 6-2]Figure 6B shows a quantitative analysis of the endothelial cell infiltration area within the hydrogel and composite (n = 6). Figure 6C shows a quantitative analysis of the endothelial cell infiltration rate within the hydrogel and composite (n = 6). Figure 6D shows a quantitative analysis of the vascular density within the hydrogel and composite (n = 6). Figure 6E shows the gene expression levels of the angiogenic genes tested (CD31 and VEGF-α). Expression is normalized to d7 250 Pa HA. ​​Samples were pooled together and tested in triplicate (n = 3). Statistical significance was calculated by two-way ANOVA with Dunnett's post-hoc test. Hash keys (#) represent comparisons between the 100 Pa HA hydrogel and the 250 Pa NHC. Asterisks (*) indicate comparisons between the 250 Pa HA control and either the 100 Pa HA or the 250 Pa NHC. ns P > 0.05, # or *P < 0.05, ## or **P < 0.01, ## or ***P < 0.001, #### or ****P < 0.0001. Data are presented as mean ± SEM. [Figure 7-1] Figure 7A shows NHC-mediated adipocyte and adipose tissue formation. Figure 7A shows H&E-stained images of 100 and 250 Pa HA and 250 Pa NHC gels at day 180. Scale bar = 100 μm. Figure 7B shows high- and low-magnification H&E-stained images of 250 Pa NHC gels at day 180. Scale bar = 5 mm for the upper low-magnification image and 100 μm for the higher-magnification image. Arrows in the higher-magnification image on the left indicate blood vessels, and arrows in the higher-magnification image on the right indicate the interface between NHCs and adipocytes. Figure 7C shows immunocytochemical images of 250 Pa NHCs at days 14, 56, and 180. Adipocytes were stained with perilipin-1 (green), and all infiltrating cell nuclei were stained with DAPI (blue). Scale bar = 200 μm. [Figure 7-2]Figure 7D shows improved adipocyte and adipose tissue formation in injected 250 Pa NHCs at POD 14 and 56 compared to 100 and 250 Pa gels without fibers. Adipocytes were stained with Acrp-30, adipose tissue was stained with perilipin-1, and all infiltrating cell nuclei were stained with DAPI (blue). Scale bars, 200 μm for tile scans (left); 50 μm for Z-stack images (right). Figure 7E shows the juxtaposition of perilipin-1+ adipose tissue formation (green) with RECA-1+ endothelial cells (red) at POD 14 (left column) and 56 (right column). Scale bars, 200 μm. [Figure 7-3] Figure 7F shows the infiltration of CD107a+ perivascular progenitor cells (gray) within NHCs and their close localization with RECA-1+ endothelial cells (green) and Acrp-30+ adipocytes (red) within NHCs at POD 14. Scale bar, 200 μm. Figure 7G shows quantitative analysis of the density of adipose tissue within NHCs and HA hydrogels by Acrp30+ staining (n = 3). Figure 7H shows quantitative analysis of the area of ​​adipocyte conversion within NHCs and HA hydrogels at day 180 (n = 3). Figure 7I shows quantitative analysis of the density of adipocytes within NHCs and HA hydrogels by perilipin staining at days 14, 56, and 180 (n = 3). Figure 7J shows quantitative analysis of the diameter of adipocytes within NHCs and HA hydrogels by perilipin staining at days 14, 56, and 180 (n = 3). Statistical significance was calculated by one-way ANOVA with Dunnett's post-hoc test for comparison between groups. ***P < 0.001, ****P < 0.0001. Data are presented as mean ± SEM. [Figure 8] HA concentration requirements for forming a strong gel. [Figure 9A] The live / dead cell viability percentages of HUVECs in 250 Pa NHC and 100 and 250 Pa HA hydrogel controls at days 1, 4, and 7 are represented (n = 3–4). [Figure 9B]Figure 1 shows the percentage of adherent cells in HA hydrogels and NHCs with HUVECs, showing cell adhesion after 4, 24, and 48 h of culture (n = 4). [Figure 9C] Figure 1 shows the close localization of CD163+ M2-like pro-regenerative macrophages with host endothelial cells inside 250 Pa NHCs on POD 14. [Figure 9D] Figure 1 shows the infiltration of CD107a+ progenitor cells (gray) and their correlation with RECA-1+ endothelial cells and Acrp-30+ adipocytes within (A) 100 Pa HA, (B) 250 Pa HA, and (C-D) 250 Pa NHCs at POD 14. Scale bar, 200 μm. [Figure 9E] 1 shows immunocytochemistry images of 250 Pa NHCs at days 14, 56, and 180 showing the juxtaposition of adipocytes (stained for perilipin-1) with blood vessels (stained for RECA-1). [Figure 9F] Figure 1 shows immunocytochemistry images of 250 Pa NHCs at day 180 showing juxtaposition of blood vessels (RECA-1 positive) with preadipocytes (Pref-1 positive). [Figure 9G] Immunocytochemistry images of 250 Pa NHCs at days 14, 56, and 180 showing the juxtaposition of endothelial cells (stained for RECA-1) with perivascular cells (stained for PDGFRα+). [Figure 9H] Graphs depict quantitative analysis of PDGFRα perivascular cell (left) and preadipocyte (right) densities in 250 Pa NHCs at days 14, 56, and 180. Statistical significance was calculated by one-way ANOVA with Dunnett's post-hoc test for comparisons between groups. ***P < 0.001, ****P < 0.0001. Data are presented as mean ± SEM. [Figure 10] Figure 1 shows collagen fibers electrospun from porcine type 1 atelocollagen (Nitta) suitable for the preparation of fiber-hydrogel composites. [Figure 11]Figure 1 shows collagen fibers electrospun from porcine gelatin (Sigma) suitable for the preparation of fiber-hydrogel composites. The fibers were formed as ribbons, demonstrating additional morphological options. [Figure 12] Figure 1 shows collagen fibers electrospun from recombinant collagen (Vecollan supplied by Evonik) suitable for the preparation of fiber-hydrogel composites. [Figure 13] Figure 1 shows an optical micrograph of electrospun recombinant collagen (Vecollan by Evonik) stained with picrosirius red colorant after crosslinking. After crosslinking, it is stable in aqueous environments and can be molded into composites with hyaluronic acid, retaining its fiber morphology after gelation and autoclave sterilization. [Figure 14] Figure 1 shows collagen fibers electrospun from recombinant collagen (Demulcent SFA by Jland) suitable for the preparation of fiber-hydrogel composites. [Figure 15] Figure 1 shows an optical micrograph of electrospun recombinant collagen (Demulcent SFA by Jland) stained with picrosirius red colorant after crosslinking. After crosslinking, it is stable in aqueous environments and can be molded into composites with hyaluronic acid, retaining its fibrous morphology after gelation and autoclaving. DETAILED DESCRIPTION OF THE INVENTION

[0035] Detailed Description The following detailed description is given by way of example and is not intended to limit the invention solely to the particular embodiments described, and may be best understood in conjunction with the accompanying drawings.

[0036] In one aspect, a pre-reacted beaded composite material comprising a hydrogel and nanostructures for use in a method for soft tissue reconstruction is provided. The invention also relates to a soft tissue device comprising the beaded composite material for cell and tissue delivery for cosmetic, reconstructive, and cellular treatments.

[0037] In certain preferred aspects, composite materials are provided that are capable of recruiting, capturing, encapsulating, binding, and / or embedding specific tissue constituents, including, but not limited to, adipocytes, other mesenchymal cells, or mesenchymal stem cells. In certain preferred aspects, composite materials are provided that are capable of recruiting, capturing, encapsulating, binding, and / or embedding specific tissue constituents, including, but not limited to, adipose tissue.

[0038] In a further preferred aspect, a method is provided for repairing or reconstructing soft tissue injuries using a composition (such as a soft tissue device) comprising a scaffold composite comprising a biomaterial covalently linked to biodegradable fibers, such as collagen.

[0039] In an additional aspect, a method of making a composition for use in soft tissue reconstruction is provided, the composition comprising a hydrogel and a nanostructure disposed therein. In certain aspects, the present invention also relates to methods of making a composition for use in cell and tissue delivery for cosmetic treatments, reconstructive treatments, and cell therapy.

[0040] The following is a detailed description of the present invention provided to assist those skilled in the art in carrying out the present invention. Those skilled in the art may make modifications and variations to the embodiments described herein without departing from the spirit or scope of the present invention. Unless otherwise defined, all technical and scientific terms used herein have the same meaning as commonly understood by those skilled in the art to which the present invention belongs. The terminology used in the description of the present invention herein is only for describing specific embodiments and is 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.

[0041] 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.

[0042] Unless otherwise defined, all technical and scientific terms used herein have the meanings that are commonly understood by those skilled in the art to which this invention belongs.The following references, the entire disclosures of which are incorporated herein by reference, provide those skilled in the art with many of the general definitions of the terms used in this 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. (eds.), Springer Verlag (1991); and Hale & Marham, the HarperCollins Dictionary of Biology (1991).In general, the molecular biology methods and similar procedures described herein or specific herein are common methods used in the art. Such standard techniques can be found in reference manuals such as, for example, Sambrook et al., (2000, Molecular Cloning—A Laboratory Manual, Third Edition, Cold Spring Harbor Laboratories); and Ausubel et al., (1994, Current Protocols in Molecular Biology, John Wiley & Sons, New York).

[0043] 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 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 the case of conflict, the present specification, including definitions, will control. Furthermore, the materials, methods, and examples are illustrative only and are not intended to be limiting.

[0044] definition The terms "a" and "an" refer to one or to more than one (i.e., to at least one) of the grammatical object of the article. By way of example, "an element" means one element or more than one element.

[0045] As used herein, "about" can mean plus or minus less than 1, or 1, 2, 3, 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15, 16, 17, 18, 19, 20, 25, 30, or more than 30 percent, depending on the context, as would be known or understood by one of ordinary skill in the art.

[0046] As used herein, a "subject" or "subjects" or "individuals" can include, but is not limited to, mammals, such as humans or non-human mammals, e.g., domesticated, agricultural, or wild animals, as well as birds and aquatic animals. In certain embodiments, a subject is a human patient or an animal that is the subject of medical treatment.

[0047] As used herein, the term "hydrogel" is a type of "gel" and refers to a water-swellable polymeric matrix consisting of a three-dimensional network of macromolecules (e.g., hydrophilic polymers, hydrophobic polymers, or blends thereof) held together by covalent or non-covalent crosslinks, which can absorb substantial amounts of water (e.g., 50%, 60%, 70%, 80%, 90%, 95%, 96%, 97%, 98%, 99%, or greater than 99% per non-water molecular unit) to form an elastic gel. Hydrogels may contain "water-swellable" polymers that absorb at least 50% or more of their weight in water upon immersion in an aqueous medium. The polymeric matrix may be formed from any suitable synthetic or naturally occurring polymeric material. As used herein, the term "gel" refers to a solid three-dimensional network that spans the volume of a liquid medium and retains it via surface tension effects. This internal network structure can result from physical bonds (physical gels) or chemical bonds (chemical gels) and crystallites or other junctions that remain intact within the elongated fluid. Virtually any fluid can be used as the elongated material, including water (hydrogels), oil, and air (aerogels). By weight and volume, gels exhibit densities similar to those of their constituent liquids because they are predominantly fluid in composition. Hydrogels are a type of gel that uses water as the liquid medium.

[0048] In certain embodiments, the hydrogel is a composite or composite material. The term "composite," as used herein, includes any association, bond, or attachment of two or more components. In some embodiments, a "hydrogel composite," as used herein, includes at least polymer fibers and a hydrogel material. A hydrogel composite contains polymer fibers (e.g., collagen, gelatin, etc.) and a hydrogel material (e.g., hyaluronic acid (HA)).

[0049] The term "functional network," as used herein, means that the interaction between components provides a chemical, biochemical, biophysical, physical, or physiological benefit. Additionally, functional networks can include additional components, including cells, biological materials (e.g., polypeptides, nucleic acids, lipids, carbohydrates), therapeutic compounds, synthetic molecules, and the like. In certain embodiments, the scaffold composite promotes tissue growth and cell infiltration when implanted into a target tissue present in a human subject.

[0050] The terms "nanofiber" or "microfiber" can be used interchangeably to refer to fibers that are several thousand nanometers in diameter, such as 1 micron to 10 microns.

[0051] The term "nanofiber," as used herein, refers to a fibrous material having at least one dimension (e.g., length or width) of less than about 999 nm, less than about 900 nm, less than about 800 nm, less than about 700 nm, less than about 600 nm, less than about 500 nm, less than about 400 nm, less than about 300 nm, less than about 200 nm, less than about 100 nm, less than about 50 nm, or less than about 10 nm. In some embodiments, nanofibers can have a length of less than about 999 nm, less than about 900 nm, less than about 800 nm, less than about 700 nm, less than about 600 nm, less than about 500 nm, less than about 400 nm, less than about 300 nm, less than about 200 nm, less than about 100 nm, less than about 50 nm, or less than about 10 nm. In some embodiments, the nanofibers can have a width of less than about 999 nm, less than about 900 nm, less than about 800 nm, less than about 700 nm, less than about 600 nm, less than about 500 nm, less than about 400 nm, less than about 300 nm, less than about 200 nm, or less than about 100 nm.

[0052] The term "microfiber," as used herein, refers to a fibrous material having at least one dimension (e.g., length or width) of less than about 100 μm, less than about 90 μm, less than about 80 μm, less than about 70 μm, less than about 60 μm, less than about 50 μm, less than about 40 μm, less than about 30 μm, less than about 20 μm, or less than about 10 μm. In some embodiments, microfibers may have lengths of less than about 10 μm, less than about 8 μm, less than about 7 μm, less than about 6 μm, less than about 5 μm, less than about 4 μm, less than about 3 μm, less than about 2 μm, or less than about 1 μm. In certain embodiments, microfibers may have lengths of about 1 μm to 10 μm, about 5 μm to 10 μm, or about 8 μm to 10 μm. In certain embodiments, microfibers may have lengths of about 8 μm or 10 μm.

[0053] The term "fiber-hydrogel composite," as used herein, refers to a composite comprising at least fibers, such as nanofibers, microfibers, or combinations thereof (e.g., polymeric fibers, or nanofibers made from extracellular matrix proteins), and a hydrogel component (e.g., HA), which forms a functional network. Furthermore, the terms "nanofiber-hydrogel composite," "fiber-hydrogel composite," "hydrogel composite," "composite," or "composite" are used interchangeably herein to refer to such a composite comprising at least fibers (e.g., collagen or gelatin nanofibers) and a hydrogel component (e.g., HA).

[0054] The term "crosslinked" herein refers to a composition containing intramolecular and / or intermolecular crosslinks, whether occurring through covalent or noncovalent bonds, which may be direct or may involve crosslinkers. "Noncovalent" bonds include both hydrogen bonds and electrostatic (ionic) bonds.

[0055] The term "polymer" includes linear and branched polymer structures, and also encompasses crosslinked polymers and copolymers (which may or may not be crosslinked), and thus includes block copolymers, alternating copolymers, random copolymers, and the like. Compounds referred to herein as "oligomers" are polymers having a molecular weight of less than about 1000 Da, preferably less than about 800 Da. Polymers and oligomers can be naturally occurring or obtained from synthetic sources.

[0056] The term "extracellular matrix protein" ("ECM"), as used herein, refers to proteins or their macromolecular matrix containing proteins, which mimic a three-dimensional network of proteins and reside outside cells (e.g., animal or mammalian cells or plant cells) to provide structural and biochemical support or adhesion to surrounding cells. Exemplary ECMs can include collagen, gelatin, elastin, decellularized matrix, or derivatives thereof.

[0057] As used herein, the term "biodegradable" refers to a material that can be broken down by biological means in a subject.

[0058] As used herein, the terms "implantable" or "injectable" mean capable of being formulated for implantation into or on a subject, such as via a syringe or device into the subject. For example, implantable means can include syringe delivery, versus mesh, paste, or the like.

[0059] As used herein, the term "soft tissue" refers to tissue that connects, supports, or surrounds other structures and organs of the body. Soft tissue includes muscles, tendons, ligaments, fascia, nerves, fibrous tissue, fat, blood vessels, and synovial membranes.

[0060] As used herein, the term "stable" refers to a material property that does not significantly degrade under given conditions (such as room temperature) over a given time frame. In certain aspects, "stable" refers to a material property that does not significantly degrade at a particular temperature or after treatment at a particular temperature. For example, when a material is said to be stable during autoclaving, it is intended that the material does not significantly degrade during the autoclaving process (e.g., steaming) or over the range of temperatures during the autoclaving process (e.g., steaming).

[0061] As used herein, the term "autologous" refers to any material derived from the same individual into which it is subsequently reintroduced.

[0062] As used herein, the term "allogeneic" or, alternatively, "allogeneic" refers to any material derived from a different animal of the same species as the individual into whom the material is being introduced or from a different patient than the individual into whom the material is being introduced.

[0063] As used herein, the term "functionalized" refers to a material that has been homogeneously or heterogeneously modified (e.g., chemically modified) to have a functional chemical moiety associated therewith. In some cases, the functional chemical moiety has the ability to react to allow for the formation of a covalent or non-covalent bond. In some cases, the functional chemical moiety can impart improved properties to the material.

[0064] Fiber-hydrogel composite Provided are "fiber-hydrogel composites," "hydrogel composites," or "composites" formed by combining hydrogel materials or other biomaterials (e.g., cells, macrophages, cell digests, or cell debris) with fibers (e.g., nanofibers or microfibers), including those having diameters in the range of 1 nm to 10 microns.

[0065] In particular, the fibers (eg, nanofibers or microfibers) suitably comprise one or more extracellular matrix proteins (ECM).

[0066] Specifically, the fibers (e.g., nanofibers or microfibers) comprise one or more collagen materials. In preferred aspects, the collagen material may comprise recombinant collagen material, including human recombinant collagen material.

[0067] In one aspect, the fiber-hydrogel composite includes (i) fibers (e.g., nanofibers or microfibers) containing one or more extracellular matrix proteins (ECM), specifically one or more collagen materials; (ii) hyaluronic acid (HA); and (iii) a crosslinker, where the HA is bonded (e.g., covalently or noncovalently) to the fibers by the crosslinker to form a composite network.

[0068] In certain embodiments, the fibers (e.g., nanofibers or microfibers) have an average diameter of less than about 10 μm, less than about 5 μm, less than about 4 μm, less than about 3 μm, less than about 2 μm, less than about 1 μm, less than about 900 nm, less than about 800 nm, less than about 700 nm, less than about 600 nm, less than about 500 nm, less than about 400 nm, less than about 300 nm, less than about 200 nm, less than about 100 nm, less than about 50 nm, less than about 10 nm, less than about 5 nm, or less than about 1 nm. In certain embodiments, the fibers have an average diameter in the range of about 10 nm to 5 μm, about 100 nm to 5 μm. In certain embodiments, the fibers have an average diameter in the range of about 1 nm to 1,000 nm, about 1 nm to 500 nm, or about 1 nm to 100 nm.

[0069] In certain embodiments, the fibers (e.g., microfibers) have a length of greater than about 1 μm, greater than about 5 μm, greater than about 10 μm, greater than about 20 μm, greater than about 30 μm, greater than about 40 μm, greater than about 50 μm, greater than about 60 μm, greater than about 70 μm, greater than about 80 μm, greater than about 90 μm, greater than about 100 μm, greater than about 200 μm, greater than about 300 μm, greater than about 400 μm, greater than about 500 μm, greater than about 600 μm, greater than about 700 μm, greater than about 800 μm, greater than about 900 μm, or about 1 mm. In certain embodiments, the fibers (e.g., microfibers) have a length within the range of about 1-1,000 μm, about 10-500 μm, or about 100-500 μm. The diameter and length of the fibers (eg, nanofibers or microfibers) can be determined using optical (including fluorescent) or electron microscopy.

[0070] In certain embodiments, the fibers may be cylindrical (having a cross section with similar width and height) or ribbon-like (having a cross section with a width greater than a height) in shape.

[0071] Preferably, the fibers (e.g., nanofibers or microfibers) can have an aspect ratio in the range of at least about 10 to about at least 10,000. Because the diameter of the fibers is so small, it will be appreciated that the fibers have a high surface area per unit of mass. This high surface area:mass ratio allows the fiber-forming solution or liquid to be transformed from a liquid or solvated fiber-forming material into a solid fiber in a fraction of a second.

[0072] In certain embodiments, the fibers (e.g., nanofibers or microfibers) are functionalized. In certain embodiments, the fibers (e.g., nanofibers or microfibers) are functionalized with groups including hydroxyl, amino, carboxyl, thio, acrylate, sulfonate, phosphate, maleimide, amide, and modified forms thereof (e.g., activated or protected forms). In certain embodiments, any functional group that can react with the crosslinker, particularly the epoxy and vinylsulfonyl groups of the crosslinker, can be used. Examples of functional groups that are reactive with the epoxy or vinylsulfonyl groups of the crosslinker include hydroxyl, carboxyl, thiol, or amino groups.

[0073] Nanofibers may include, but are not limited to, nanofibers, nanotubes, nanofilaments, mesh sections, branched filaments, or networks. Nanofibers may also include any suitable chemical functional groups to facilitate covalent or non-covalent crosslinking between the nanofibers and the polymers of the hydrogels of the present invention. Methods, techniques, and materials for making and functionalizing nanofibers are well known in the art. Nanofibers may be made using any microfabrication method. In various embodiments, devices of the present disclosure may be assembled and / or manufactured using any suitable microfabrication technique. Such methods and techniques are widely known in the art.

[0074] Specifically, the fibers (e.g., nanofibers or microfibers) in the fiber-hydrogel composite comprise one or more extracellular matrix proteins (ECM). In certain embodiments, the fibers (e.g., nanofibers or microfibers) suitably comprise one or more selected from collagen, gelatin, elastin, elastin-like polypeptides, tropoelastin, decellularized matrix, and hyaluronic acid. In certain embodiments, the fibers (e.g., nanofibers or microfibers) comprise one or more of bovine type I collagen, gelatin, or derivatives. In certain embodiments, the fibers (e.g., nanofibers or microfibers) comprise one or more collagens, including bovine type I collagen and recombinant collagen, which may be human recombinant collagen material.

[0075] In certain embodiments, the fiber-hydrogel composite may include a natural extracellular matrix for the fibers (e.g., nanofibers or microfibers). In certain embodiments, the fiber-hydrogel composite may include a synthetic extracellular matrix for the fibers (e.g., nanofibers or microfibers). In certain embodiments, the one or more ECMs include collagen nanofibers, which may be naturally obtained or synthesized. In certain embodiments, the collagen nanofibers include type I bovine collagen nanofibers or fragments thereof. In certain embodiments, the collagen nanofibers may be obtained from natural sources or may be made or prepared from a composition (resin composition) comprising collagen. For example, the collagen nanofibers may be formed by electrospinning, centrifugal spinning, blow spinning, or a combination thereof. Specifically, the collagen nanofibers are preferably prepared by electrospinning.

[0076] "Human recombinant collagen" or "human collagen material" or similar terms may suitably refer to collagen types I, II, III, IV, V, VI, VII, VIII, IX, X, XI, XII, XIII, XIV, XV, XVI, XVII, XVIII, XIX, XX, XXI, XXII, XXIII, XXIV, XXV, XXVI, and XXVII, or their component chains. The collagen may be one type of collagen free of any other type, or may be a mixture of collagen types. Suitably, the collagen comprises or consists essentially of a collagen selected from the group consisting of type I collagen, type III collagen, and mixtures thereof. Human recombinant collagen may suitably be produced, for example, by culturing a non-human organism to express at least one human gene encoding a collagen. Suitable collagens may or may not contain hydroxyproline residues or telopeptide sequences.

[0077] Human recombinant collagen can be prepared by any suitable method known in the art, including those disclosed in U.S. Patent Nos. 5,962,648 and 5,593,859 and WO2004 / 078120.Suitably, collagen is recombinantly produced by culturing cells transfected with at least one gene encoding a collagen-constituting polypeptide and a gene encoding the oc and subunits of the post-translational enzyme prolyl 4-hydroxylase, and purifying the resulting collagen monomers therefrom.The recombinant collagen solution can then be subjected to polymerization or cross-linking conditions.

[0078] Bovine collagen may suitably be a mixture of type I collagen (85%) and type III collagen (15%). The advantage of recombinant collagen is that type I collagen and type III collagen are produced independently of each other, thus allowing for the production of any combination of type I and type III collagen. The compositions and composites of the present invention may suitably contain human type I collagen and human type III collagen in any ratio. For example, the compositions and composites may contain human type I collagen and human type III collagen in a weight ratio of 100:0, 80:20, 60:40, 50:50, 40:60, 20:80, or 0:100, or any ratio therebetween. Preferably, the weight ratio of human type I collagen to human type III collagen is greater than about 50:50, preferably greater than about 70:30, e.g., about 80:20. Suitably, type I human recombinant collagen constitutes at least about 75% by weight of the total human recombinant collagen in the material.

[0079] As discussed above, in certain aspects, the collagen of the compositions and composites of the present invention is recombinant collagen, human collagen, or recombinant human collagen, respectively. In various embodiments, the collagen is selected from the group consisting of collagen types I, II, III, IV, V, VI, VII, VIII, IX, X, XI, XII, XIII, XIV, XV, XVI, XVII, XVIII, XIX, XX, XXI, XXII, XXIII, XXIV, XXV, XXVI, and XXVII. In some embodiments, the collagen is collagen of one collagen type without any other collagen type; in other embodiments, the collagen is a specified or unspecified mixture of more than one collagen type.

[0080] In one aspect, the composites and compositions of the present invention include collagen, and the composite is about 2.3 mm thick. 2 In other aspects, the composite has a surface area of ​​about 2.5, 2.75, 3.0, 3.25, 3.5, 3.75, 3.8, 4.0, and 4.4 m / g collagen.2 In certain aspects, a composite comprising collagen is provided, the composite having a surface area greater than about 4.0 m / g collagen. 2 / g collagen has a larger surface area.

[0081] In further aspects, the collagen is human collagen, recombinant collagen, recombinant human collagen, and / or collagen type I. In a preferred aspect, the collagen is recombinant human collagen type I. In certain aspects, the collagen is selected from the group consisting of collagen types I, II, III, IV, V, VI, VII, VIII, IX, X, XI, XII, XIII, XIV, XV, XVI, XVII, XVIII, XIX, XX, XXI, XXII, XXIII, XXIV, XXV, XXVI, and XXVII.

[0082] Fibers (e.g., nanofibers or microfibers) may preferably be produced by electrostatic spinning (also known as electrospinning). The electrospinning process generally involves introducing a liquid into an electric field, which causes the liquid to produce fibers. These fibers are generally attracted to a conductor with an attractive potential for collection. During the liquid-to-fiber transformation, the fibers solidify and / or dry. This solidification and / or drying can be caused by cooling the liquid (i.e., in this case, the liquid is usually solid at room temperature); by evaporation of the solvent, for example, by dehydration (physically induced solidification); or by a hardening mechanism (chemically induced solidification). Electrostatically spun fibers can be produced with very thin diameters. Parameters that affect the diameter, consistency, and uniformity of electrospun fibers include the polymeric material and crosslinker concentration (loading) in the fiber-forming combination, the applied voltage, and the needle-to-collector distance.

[0083] Electrospun fibers (e.g., collagen nanofibers) can offer superior properties, such as high porosity in the hydrogel phase and mechanical reinforcement from the solid fiber component, which can be beneficial for optimal cell infiltration properties and structural integrity.

[0084] In certain embodiments, collagen fibers are prepared by electrospinning fibers in a solution or suspension containing 1,1,1,3,3,3-hexafluoro-2-propanol (HFIP) solvent. In other embodiments, collagen fibers are prepared by electrospinning in a solution or suspension containing a different solvent, such as trifluoroethanol (TFE), trifluoroacetic acid (TFA), acetic acid, ethanol, or a phosphate mixture. In certain embodiments, the suitable solvent may include at least one of 1,1,1,3,3,3 hexafluoro-2-propanol (HFIP), 2,2,2-trifluoroethanol (TFE), and a mixture of water and acetic acid. Other solvents that may be used or combined with other solvents in the electrospinning of natural matrix materials such as collagen fibers include acetamide, N-methylformamide, N,N-dimethylformamide (DMF), dimethyl sulfoxide (DMSO), dimethylacetamide, N-methylpyrrolidone (NMP), ethyl acetate, acetonitrile, trifluoroacetic anhydride, 1,1,1-trifluoroacetone, maleic acid, and hexafluoroacetone.

[0085] The choice of solvent depends on the characteristics of the synthetic polymer being electrodeposited, such as secondary forces that stabilize polymer-polymer interactions, and the solvent's ability to replace them with strong polymer-solvent interactions. In the case of polypeptides such as collagen, and in the absence of covalent crosslinks, the primary interchain secondary forces are (1) Coulombic forces resulting from the attraction of fixed charges on the backbone and dictated by the primary structure (e.g., lysine and arginine residues are positively charged at physiological pH, while aspartic acid or glutamic acid residues are negatively charged); (2) dipole-dipole interactions resulting from the interaction of permanent dipoles (hydrogen bonds commonly found in polypeptides are the strongest such interactions); and (3) hydrophobic interactions resulting from the association of nonpolar regions of the polypeptide due to the low tendency of nonpolar species to interact favorably with polar water molecules.

[0086] Therefore, solvents or solvent combinations that can compete favorably for these interactions can dissolve or disperse polypeptides. For example, HFP and TFE possess highly polar hydroxyl groups adjacent to very hydrophobic fluorinated regions. Without wishing to be bound by theory, it is believed that the alcohol moieties can hydrogen bond with peptides and also solvate the backbone charges, thereby reducing intermolecular Coulombic interactions. Furthermore, the hydrophobic moieties of these solvents may interact with hydrophobic domains in polypeptides, helping the latter resist the tendency to aggregate through hydrophobic interactions. Due to their lower overall polarity compared to water, solvents such as HFP and TFE may compete less for intramolecular hydrogen bonds that stabilize secondary structures such as alpha helices. Consequently, alpha helices in these solvents are believed to be stabilized by stronger intramolecular hydrogen bonds. Stabilization of polypeptide secondary structures in these solvents is believed to be desirable, particularly in the case of collagen and elastin, to maintain proper collagen fibril formation during electrospinning. In some embodiments, solvents are selected based on their tendency to induce helical structure in electrospun protein fibers, thereby allowing collagen or other protein monomers to undergo polymerization and form helical polymers that mimic natural collagen fibrils. Examples of such solvents include halogenated alcohols, preferably fluorinated alcohols (HFP and TFE), hexafluoroacetone, chloroalcohols in combination with aqueous solutions of mineral acids, and dimethylacetamide, preferably containing lithium chloride. HFP and TFE are more preferred. In some embodiments, water is added to the solvent.

[0087] In certain embodiments, collagen nanofibers have been prepared using vapor-phase glutaraldehyde, an alternative fiber stabilizer. Treatment with glutaraldehyde results in crosslinking of collagen fibers, as the aldehyde groups of glutaraldehyde react with free lysine or hydroxylysine groups on the collagen fibers to form Schiff base structures. For example, a 6-hour vapor-phase glutaraldehyde treatment results in increased tensile strength, elasticity, stretchability, and stability of collagen fibers. In certain embodiments, collagen nanofibers have been prepared using alternative collagen stabilizers or crosslinkers, such as D-ribose. As disclosed in U.S. Pat. No. 4,971,954, the entire contents of which are incorporated herein, D-ribose can crosslink collagen fibers to produce a nontoxic and nonimmunogenic matrix.

[0088] In certain embodiments, collagen nanofibers have been prepared using DVS (divinyl sulfone) as a fiber stabilizer, either alone or in series with another stabilizer, such as vapor-phase glutaraldehyde. This can result in crosslinking of the collagen fibers, stabilizing the physical structure of the fibers prior to reaction during gelation. By adjusting the selection of fiber crosslinking conditions (including solvent selection), it is possible to crosslink the fibers while leaving abundant functional groups for subsequent reaction to form interfacial bonds during composite gelation.

[0089] In certain embodiments, the EDC and the fiber (e.g., nanofiber or microfiber) can be crosslinked (e.g., via a crosslinking moiety or directly linked). For example, the mode of interaction between the EDC and the fiber (e.g., nanofiber or microfiber) can be effective in introducing a bond (e.g., a covalent bond) between them.

[0090] In certain embodiments, a hydrogel material such as HA can be covalently bonded to fibers (e.g., nanofibers or microfibers). For example, a hydrogel material such as HA can be covalently bonded to recombinant or type I bovine collagen nanofibers or fragments thereof. In certain embodiments, a crosslinker generates an interfacial bond between the collagen nanofiber and HA. Due to bonds and interactions (e.g., covalent, non-covalent, or ionic bonds), the collagen nanofibers can be held within or within the interior space of the fiber-hydrogel composite (e.g., within the composite network). In certain embodiments, the crosslinker can react with the hydroxyl groups of the hydrogel material (e.g., HA) and the amino groups of the collagen nanofiber to form a composite network. For example, the interfacial bond between the collagen nanofiber and the hydrogel material (e.g., HA) can increase the stiffness of the composite, even at relatively low fiber loading densities.

[0091] Due to interfacial bonding and the formation of a composite network, the fiber-hydrogel composite may have increased cell permeability and / or maintain storage modulus.

[0092] In certain embodiments, the storage modulus of the fiber-hydrogel composite is at least about 10 Pa, at least about 20 Pa, at least about 30 Pa, at least about 40 Pa, at least about 50 Pa, at least about 60 Pa, at least about 70 Pa, at least about 80 Pa, at least about 90 Pa, at least about 100 Pa, at least about 150 Pa, at least about 200 Pa, at least about 250 Pa, at least about 300 Pa, at least about 400 Pa, or at least about 500 Pa. In certain embodiments, the storage modulus of the fiber-hydrogel composite is in the range of about 1 to about 1,000 Pa, about 20 to about 800 Pa, about 100 to about 500 Pa, or about 150 to about 500 Pa. In certain embodiments, the storage modulus of the fiber-hydrogel composite is in the range of about 0.5 to about 30 kPa. Storage modulus values, as referred to herein, may be determined by the procedure set forth in Example 3 below, which involves measuring the linear viscoelastic region of a composition sample by strain sweeping at increasing shear strain amplitudes at a set frequency (1 Hz) using an analytical tool such as an AR2 (TA Instruments) using an 8 mm parallel plate geometry with a 0.5 mm gap at 25°C.

[0093] In certain embodiments, as discussed, alternative hydrogel phases such as collagen, chitosan, alginate, PVA, gelatin, PEG or other glycol ethers, cellulose, or cellulosic materials may be used in place of or in combination with hyaluronic acid (HA).

[0094] In certain embodiments, the fiber-hydrogel composite can be stable at temperatures and pressures suitable for terminal sterilization (e.g., autoclaving). For example, the fiber-hydrogel composite can be stable at temperatures between about 100°C and 131°C for at least 30 minutes (less time required at 131°C). In certain embodiments, the composite can have increased thermal stability and / or be shelf-stable at ambient temperatures. Thus, for example, autoclave sterilization can be used to sterilize the gel in the final manufacturing step without significantly altering the mechanical behavior of the composite, reducing manufacturing costs, risks, or regulatory burdens.

[0095] A preferred form of interaction between the fibers (e.g., nanofibers or microfibers) and the hydrogel components includes cross-linking moieties, typically present in an amount effective to introduce bonds between the fibers (e.g., nanofibers or microfibers) and the hydrogel material, e.g., to induce cross-linking between collagen nanofibers and hyaluronic acid.

[0096] For example, if high cohesion strength is desired, the fibers (e.g., nanofibers or microfibers) and hydrogel components can be covalently crosslinked. For example, HA hydrogel polymer components can be covalently crosslinked to the fibers either intramolecularly, intermolecularly, or via covalent bonds. In the former case, there are no covalent bonds connecting the polymers to each other or to the nanostructure, while in the latter case, there are covalent crosslinks connecting the polymers to each other or to the nanostructure. Crosslinks can be formed using any suitable means, including the use of 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 can also involve the use of a third molecule, a "crosslinker," utilized in the crosslinking process.

[0097] The "crosslinker" or "crosslinking agent" may suitably comprise one or more selected from difunctional epoxide-based crosslinkers, 1,4-butanediol diglycidyl ether (BDDE), divinyl sulfone (DVS), PEG molecules with terminal epoxide functional groups, and HA reactive agents. Preferably, the crosslinking agent comprises DVS or BDDE. In certain embodiments, the use of a PEG crosslinking agent, which may contain terminal functional groups such as epoxide or vinyl sulfone groups, introduces crosslinks between fibers (e.g., nanofibers or microfibers) as well as between fibers (e.g., nanofibers or microfibers) and hydrogels, thereby extending the durability of the composite network and adjusting the crosslinking density. In certain embodiments, the crosslinking agent does not include any spacers within its structure.

[0098] Crosslinking can also be achieved using radiation, typically in the presence of a photoinitiator. The radiation can be ultraviolet, alpha, beta, gamma, electron beam, and x-ray radiation, with ultraviolet radiation being preferred. Useful photosensitizers are "hydrogen abstraction" type triplet sensitizers, including benzophenones and substituted benzophenones and acetophenones such as benzyl dimethyl ketal, 4-acryloxybenzophenone (ABP), 1-hydroxy-cyclohexylphenyl ketone, 2,2-diethoxyacetophenone, and 2,2-dimethoxy-2-phenylacetophenone; substituted alpha-ketols such as 2-methyl-2-hydroxypropiophenone; benzoin ethers such as benzoin methyl ether and benzoin isopropyl ether; and anisoin methyl ether. These include substituted benzoin ethers of the formula (I), aromatic sulfonyl chlorides such as 2-naphthalenesulfonyl chloride, photoactive oximes such as 1-phenyl-1,2-propanedione-2-(O-ethoxycarbonyl)-oxime, thioxanthones including alkyl- and halogen-substituted thioxanthones such as 2-isopropylthioxanthone, 2-chlorothioxanthone, 2,4-dimethylthioxanone, 2,4-dichlorothioxanone, and 2,4-diethylthioxanone, and acylphosphine oxides. Radiation having wavelengths of 200-800 nm, preferably 200-500 nm, is preferred for use herein; in most cases, low-intensity UV light is sufficient to induce crosslinking. However, with hydrogen abstraction-type photosensitizers, higher-intensity UV exposure may be required to achieve sufficient crosslinking. Such exposure can be provided by mercury lamp processors such as those available from PPG, Fusion, Xenon, and others. Crosslinking can also be induced by gamma or electron beam irradiation. Appropriate irradiation parameters, i.e., the type and dose of radiation used to induce crosslinking, will be apparent to those skilled in the art.

[0099] Suitable chemical curing agents, also referred to as chemical crosslinking "accelerators," include, but are not limited to, 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, dithiopropane, or trithiopropane, and polymercaptans such as 1,6-hexanedithiol. Adding a cross-linking promoter to a non-cross-linked hydrophilic polymer promotes its covalent cross-linking, or adding a cross-linking promoter to a blend of a non-cross-linked hydrophilic polymer and a complementary oligomer provides cross-links between the two components.

[0100] In certain embodiments, the concentration of the crosslinker (e.g., DVS or BDDE) ranges from about 0.01 v / v% to about 10 v / v%, 0.05 v / v% to about 10 v / v%, 0.05 v / v% to about 5 v / v%, about 0.2 v / v% to about 10 v / v%, about 0.5 v / v% to about 5.0 v / v%, or about 0.5 v / v% to about 2.5 v / v%, based on the total volume of the fiber-hydrogel composite. In certain embodiments, the concentration of DVS ranges from about 0.01 v / v% to about 10 v / v%, 0.05 v / v% to about 10 v / v%, 0.05 v / v% to about 5 v / v%, about 0.2 v / v% to about 10 v / v%, about 0.5 v / v% to about 5.0 v / v%, or about 0.5 v / v% to about 2.5 v / v%, based on the total volume of the fiber-hydrogel composite.

[0101] In certain embodiments, the concentration of HA ranges from about 0.1 w / v% to about 10 w / v%, about 0.2 w / v% to about 10 w / v%, about 0.2 w / v% to about 5.0 w / v%, about 0.5 w / v% to about 10.0 w / v%, about 0.5 w / v% to about 5.0 w / v%, about 0.5 w / v% to about 2.0 w / v%, or about 0.8 w / v% to about 2.0 w / v%, based on the total volume of the fiber-hydrogel composite.

[0102] In certain embodiments, the concentration of one or more ECMs ranges from about 0.1 to about 50 w / v%, about 0.1 to about 40 w / v%, about 0.1 to about 30 w / v%, about 0.1 to about 20 w / v%, about 0.1 to about 10 w / v%, about 1 to about 10 w / v%, about 1 to about 5 w / v%, about 1 to about 3 w / v%, or about 1.5 to about 3 w / v% based on the total volume of the fiber-hydrogel composite. In certain embodiments, the fiber loading density of the collagen nanofibers ranges from about 0.1 w / v% to about 10 w / v%, about 1 w / v% to about 10 w / v%, about 1 w / v% to about 5 w / v%, or about 1 w / v% to about 3 w / v% based on the total volume of the fiber-hydrogel composite.

[0103] For example, in one embodiment, the concentration of one or more ECMs ranges from about 0.1 to about 20 w / v%, the concentration of HA ranges from about 0.5 to about 10 w / v%, and the concentration of a crosslinker (e.g., DVS or BDDE) ranges from about 0.05 to about 5.0 v / v%, based on the total volume of the fiber-hydrogel composite. Further, in one embodiment, the concentration of one or more ECMs ranges from about 1.5 to about 3.0 w / v%, the concentration of HA ranges from about 0.8 to about 2 w / v%, and the concentration of a crosslinker (e.g., DVS) ranges from about 0.5 to about 2.5 v / v%, based on the total volume of the fiber-hydrogel composite.

[0104] The fiber-hydrogel composites can be formed to have a final device pH in an isotonic solution ranging from about 5.0 to about 9.0, from about 6.0 to about 8.0, or from about 7.0 to about 7.4.

[0105] The fiber-hydrogel composite can be formulated into a sheet or a flowable or injectable fluid. For example, the fiber-hydrogel can be constructed into a solution that can be applied and shaped (or optionally dried) to form a sheet-type gel. Additionally, the fiber-hydrogel composite can be formed into a fluid (e.g., an aqueous suspension or dispersion) that can pass through a 27-gauge or finer needle.

[0106] In certain embodiments, the fiber-hydrogel composite may exhibit monocyte recruitment, monocyte polarization, or both. In certain embodiments, the fiber-hydrogel may accommodate or contain macrophages.

[0107] In certain embodiments, fiber-hydrogel composites are characterized by exhibiting biostimulatory effects such as tissue remodeling, host cell infiltration, cell adhesion, cell migration, angiogenic responses, adipogenic responses, and macrophage polarization toward a pro-healing phenotype. Specifically, fiber-hydrogel composites can have biostimulatory effects due to the low crosslink density of the composite material. Fiber-hydrogel composites can have long-term retention of biostimulatory fibers (e.g., collagen fibers) in a hyaluronic acid (HA) network.

[0108] In certain embodiments, the fiber-hydrogel composite can allow cell infiltration. In certain embodiments, the fiber-hydrogel can accommodate or contain infiltrating macrophages. In certain embodiments, the fiber-hydrogel can accommodate or contain M2 phenotype infiltrating macrophages, due to the collagen nanofibers conditioning the population of M2 phenotype infiltrating macrophages. The fiber-hydrogel composite can regulate host cell infiltration and / or composite shape retention.

[0109] In certain embodiments, the fiber-hydrogel composite can have a tissue remodeling effect induced by the fiber-hydrogel composite without the incorporation of cells and growth factors.

[0110] In certain embodiments, the fiber-hydrogel composites can promote or induce cell migration, ie, attract 1.5-fold more, 2.0-fold more, 2.5-fold more, 3.0-fold more, or 5-fold more host cells into an injection site in vivo compared to a hydrogel control without ECM nanofibers.

[0111] In certain embodiments, the fiber-hydrogel composites can promote or induce an angiogenic response. For example, the fiber-hydrogel composites can promote neovasculature formation. In certain embodiments, the fiber-hydrogel composites can also promote angiogenesis and soft tissue repair in the absence of any exogenous cytokines and cells, thereby providing long-term cellular repair benefits.

[0112] In certain embodiments, the collagen nanofiber composites promote progenitor cell adipogenesis in vitro and / or accelerate adipogenesis in vivo.

[0113] In certain aspects, collagen nanofiber composites promote adipogenesis in vivo in surrounding tissues such as the corneum dartos in rodent models.

[0114] In certain embodiments, the fiber-hydrogel composite is administered to a subject by injection and maintains greater than about 50%, about 60%, about 70%, about 80%, about 90%, or about 95% cell viability 7 days after injection.

[0115] Preferably, fiber-hydrogel composites can be formulated to vary density, gel-to-fiber ratio, and other properties while maintaining sufficient porosity and strength. The ratio of fiber (e.g., nanofiber or microfiber) to hydrogel material can be determined by any means known in the art. For example, the ratio of ECM fiber to hydrogel material (e.g., hyaluronic acid) is 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, or about 1:3 to about 3:1, by weight of components. The ratio of ECM fiber to hydrogel material (e.g., hyaluronic acid) can also be given on a concentration basis, e.g., a given weight of polymer fiber per volume of hydrogel material. For example, the concentration can be about 1 to 50 mg / mL. The hydrogel material is suitably tethered, attached, or disposed within the ECM fibers to generally form a composite network.

[0116] Fiber-hydrogel composites can contain a plurality of pores present on or within the surface of the composite. The presence, size, distribution, frequency, and other parameters of the pores can be adjusted during creation of the composite, hydrogel, or fiber (e.g., nanofiber or microfiber). Pore size can be less than about 1 nm up to 100 μm, including 1, 2, 3, 4, 5, 10, 15, 20, 30, 40, 50, 60, 70, 80, 90, or 100 μm, and the size can be narrowly tailored so that at least 40% of the pores are within the desired size or fall within a desired size range, e.g., 50%, 60%, 70%, 80%, 90%, 95%, or more than 95% of the pores.

[0117] In one aspect, a preferred composition or composite comprises a plurality of pores present on the surface of a layer of the composite or composition, the pores being at least about 50 pores / cm of surface. 2 and at least 80% of the pores have an average surface pore diameter that is at least about 5 microns.

[0118] The fiber-hydrogel composites may be suitable for incorporation into the tissues of a human subject and are therefore generally "biocompatible," meaning they have the ability to interact with biological systems (such as those found in a human subject) without inducing a pathophysiological response therein and / or thereby. In certain embodiments, the fiber-hydrogel composites are provided to be permanently retained in tissue, e.g., organs, nervous tissue, etc. Alternatively, the composites are provided to be transiently retained in a human subject and are substantially biodegradable. Preferably, the ECM fibers may further comprise a biocompatible, biodegradable polymer, e.g., a biocompatible, biodegradable polyester.

[0119] In certain embodiments, the fiber-hydrogel composite is formed into microbeads (spherical or non-spherical) having an average diameter within the range of about 1 μm to about 1000 μm, about 10 μm to about 1000 μm, about 20 μm to about 1000 μm, about 30 μm to about 1000 μm, about 40 μm to about 1000 μm, about 50 μm to about 1000 μm, about 50 μm to about 500 μm, about 50 μm to about 400 μm, or about 100 μm to about 500 μm. The fiber-hydrogel composite can be formed into a particulate formulation (e.g., microbeads or microgels) to allow for the use of higher concentrations of each component and enhanced stability. In certain embodiments, a micronization system may be used in which a preformed fiber-hydrogel composite is physically conditioned, such as by forcing it through one, two, three, or more mesh screens to create a population of non-spherical beads that are relatively similar in shape and size. This multi-screen system allows for strict control over the size of the beads, allowing the user to adjust the size as needed. Such microbeads (e.g., non-spherical) are disclosed in US 2020 / 0069846. For example, fiber-hydrogel composites have been conditioned by applying mechanical shear through meshes with defined sizes in the range of 50 to 400 μm.

[0120] The fiber-hydrogel composite can further include an active agent, thereby acting as an active agent delivery system when applied to a body surface (e.g., a tissue repair site) in a manner that delivers the active agent thereto. Release of the active agent "loaded" into the fiber-hydrogel composite (e.g., HA hydrogel or fiber) typically involves both water absorption via a swelling-controlled diffusion mechanism and desorption of the agent. For example, active agent-containing hydrogel compositions can be used in, for example, transdermal drug delivery systems, wound dressings, topical pharmaceutical formulations, implantable drug delivery systems, oral dosage forms, and the like.

[0121] In certain embodiments, the fiber-hydrogel composite may contain one or more exogenous growth factors and / or cytokines. Exemplary growth factors may include, but are not limited to, epidermal growth factor (EGF), platelet-derived growth factor (PDGF), vascular endothelial growth factor (VEGF), fibroblast growth factor (FGF), transforming growth factor-alpha (TGFα), and / or nerve growth factor (NGF). In certain embodiments, fiber-hydrogel composites containing EGF / TGF may be advantageously used in accelerating wound healing and reducing keloid scar formation (especially for burns), as skin dressings, and treating chronic leg ulcers. In certain embodiments, fiber-hydrogel composites containing VEGF may promote or contribute to angiogenesis (blood vessel growth) both indirectly and directly by stimulating the proliferation and migration of endothelial cells at the microvascular level and altering their overall expression. In certain embodiments, fiber-hydrogel composites containing FGF can promote or induce angiogenesis in vivo, and the angiogenic properties can be enhanced by the co-administration of TNF. For example, the inclusion of FGF-2 in the fiber-hydrogel composite can suitably regulate human megakaryopoiesis or stimulate endothelial cell formation and connective tissue repair. Keratinocyte growth factor (KGF), also known as FGF-7, can also be suitably included in fiber-hydrogel composites for wound healing and other disorders involving epithelial cell destruction. In certain embodiments, fiber-hydrogel composites containing transforming growth factor (TGF) can transform various cell lines, which may have, for example, the ability to grow in culture for more than a limited number of generations, multilayer rather than monolayer growth, and the acquisition of abnormal karyotypes. For example, suitably, the inclusion of TGF-beta in the fiber-hydrogel composite can promote angiogenic effects and collagen formation in fibroblasts, antagonize the mitogenic effects of other peptide growth factors, and inhibit the growth of many tumor cell lines.

[0122] Additional suitable active agents that can be incorporated into the fiber-hydrogel composite and delivered systemically (e.g., using a transdermal, oral, or other dosage form suitable for systemic administration of the drug) include analgesics; anesthesia agents; anti-arthritic agents; respiratory medications, including anti-asthmatic agents; anti-cancer medications, including antineoplastic agents; anticholinergics; anticonvulsants; antidepressants; antidiabetic agents; antidiarrheals; anthelmintics; antihistamines; anti-dyslipidemic agents; antihypertensive agents; anti-infective agents, such as antibiotics and antivirals; anti-inflammatory agents; anti-migraine preparations; antiemetics; anti-Parkinson's agents; antipruritics; antipsychotics; antipyretics; antispasmodics; anti-tuberculosis agents; antiulcer agents; antivirals; anti-anxiety agents; appetite suppressants; attention deficit disorder (ADD) and attention deficit hyperactivity disorder (ADHD) medications; calcium channel blockers, anti-anginal agents, central nervous system These include, but are not limited to, cardiovascular preparations including CNS agents, beta-blockers, and antiarrhythmics; central nervous system stimulants; cold preparations including decongestants; diuretics; genetic materials; herbal medicines; hormonelytics; hypnotics; hypoglycemic agents; immunosuppressants; leukotriene inhibitors; antimitotic agents; muscle relaxants; anesthetic antagonists; nicotine; nutritional supplements such as vitamins, essential amino acids, and fatty acids; eye drops such as antiglaucoma agents; parasympatholytics; peptide drugs; psychostimulants; sedatives; steroids including progestogens, estrogens, corticosteroids, androgens, and anabolic agents; smoking cessation agents; sympathomimetics; tranquilizers; and vasodilators, including general coronary, peripheral, and cerebral. Specific active agents useful in combination with the adhesive compositions of the present invention include, but are not limited to, anabasine, capsaicin, isosorbide dinitrate, aminostigmine, nitroglycerin, verapamil, propranolol, cilabolin, foridone, clonidine, cytisine, phenazepam, nifedipine, fluacizin, and salbutamol.

[0123] In certain embodiments, the fiber-hydrogel composite may also contain additional optional additive components. Such components are known in the art and may include, for example, fillers, preservatives, pH adjusters, softening agents, viscosity enhancers, pigments, coloring agents, refractive particles, stabilizers, reinforcing agents, adhesion-reducing agents, pharmaceutical agents (e.g., antibiotics, angiogenesis promoters, antifungals, immunosuppressants, antibodies, and the like), and permeation enhancers. Such additives and their amounts are selected so that they do not significantly interfere with the desired chemical and physical properties of the hydrogel composition.

[0124] In certain embodiments, the fiber-hydrogel composite may include a pH-adjusting compound. Compounds useful as pH-adjusting agents include, but are not limited to, glycerol buffer, citrate buffer, borate buffer, phosphate buffer, or citrate-phosphate buffer, which may also be included to ensure that the pH of the hydrogel composition is compatible with that of an individual's body surface.

[0125] In certain embodiments, the fiber-hydrogel composite may comprise and deliver an antibody. The term "antibody" is used herein in its broadest sense to include certain types of immunoglobulin molecules that contain one or more antigen-binding domains that specifically bind to an antigen or epitope. For example, the fiber-hydrogel composite may comprise an intact antibody (e.g., an intact immunoglobulin), an antibody fragment, a multispecific antibody, a monoclonal antibody, a chimeric antibody, or a humanized antibody.

[0126] In certain embodiments, the fiber-hydrogel composite can include cells for delivery. For example, the cells can be derived from the subject to whom they are administered, from a cell line derived from a source other than the subject to whom they are administered, from a human source, or from a humanized animal source. In certain embodiments, the cells can include, but are not limited to, stem cells, adipose cells or tissue, or neural cells.

[0127] In certain embodiments, the fiber-hydrogel composite may also include small molecules for delivery, which may cause pharmacological activity or another direct effect in the diagnosis, cure, mitigation, treatment, or prevention of disease, or may affect the structure or function of the body.

[0128] Specifically, the fiber-hydrogel composite may further include components that promote angiogenesis, such as growth factors or cells. For example, heparin-containing hydrogel components may be used, which can serve as growth factor binding sites to concentrate and retain growth factors that promote angiogenesis and tissue formation.

[0129] In certain embodiments, the fiber-hydrogel composite may further include one or more enzymes that may be used in debridement of both acute and chronic wounds. Enzymes may be incorporated, for example, by directly digesting components of carrion (e.g., fibrin, bacteria, leukocytes, cellular debris, serous exudates, DNA) or by dissolving collagen "anchors" that secure avascular tissue to the underlying wound bed.

[0130] In certain embodiments, the fiber-hydrogel composite can be delivered by any suitable method, such as via a syringe or bellows pack (single-dose delivery system), or via a multi-dose system, such as via a pressurized delivery system or a "bag-in-the-can" type system. Administration can be extended to single-dose delivery systems containing the fiber-hydrogel composite for wound treatment, pressurized delivery systems (e.g., aerosol sprays) containing the fiber-hydrogel composite.

[0131] In certain embodiments, it may be advantageous to render the fiber-hydrogel composite electrically conductive for use in biomedical electrodes and other electrotherapeutic settings, i.e., for attaching electrodes or other electrically conductive members to a body surface. For example, the fiber-hydrogel composite may be used to attach a transcutaneous nerve stimulation electrode, an electrosurgical return electrode, or an EKG electrode to a patient's skin or mucosal tissue. Such applications involve modifying the fiber-hydrogel composite to contain conductive species. Suitable conductive species are ionically conductive electrolytes, particularly those commonly used in the manufacture of conductive adhesives for application to 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, redox couples such as 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 ferric and ferrous salts (such as sulfate and gluconate). Preferred salts are potassium chloride, sodium chloride, magnesium sulfate, and magnesium acetate, with potassium chloride being most preferred for EKG applications. While virtually any amount of electrolyte can be present in the adhesive compositions of the present invention, any electrolyte is preferably present at a concentration within the range of about 0.1 to about 15 wt. % of the hydrogel composition. The techniques 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 is incorporated by reference for manufacturing details. Other suitable fabrication techniques may also be used, as will be understood by those skilled in the art.

[0132] In some embodiments, the hyaluronic acid is replaced with carboxymethylcellulose (CMC) and the collagen fibers are replaced with cellulose-based fibers. The hydroxyls in both components allow for easy modification, and both can be autoclaved.

[0133] Soft Tissue Devices Provided herein are soft tissue devices or implants comprising the fiber-hydrogel composites described herein.

[0134] In certain embodiments, the implants may be used to promote angiogenesis.

[0135] In certain embodiments, the implant can be used for adipose tissue formation. For example, the implant can further comprise a biologically active material suitable for fat transplantation, which can differentiate into soft tissue such as fat when supported, for example, with an appropriate matrix microenvironment. In certain embodiments, the biologically active material comprises a population of adipose cells, autologous adipose cells, allogeneic cells, genetically modified allogeneic cells, stem cells, mesenchymal stem cells, genetically modified stem cells, genetically modified allogeneic induced pluripotent stem (iPS) cells, genetically modified low-immunogenic pluripotent stem cells, adipose stromal vascular cells, adipose tissue, autologous adipose tissue, lipoaspirate tissue, derivatives thereof, or combinations thereof. In some embodiments, the biologically active material comprises adipose tissue.

[0136] In certain aspects, the implants may be used for vasculogenesis.

[0137] In certain embodiments, a soft tissue device or implant may suitably include microbeads (e.g., spherical or non-spherical) made from a fiber-hydrogel composite in a volume that is approximately 25% to 75% of the total volume of the soft tissue device. In certain embodiments, a soft tissue device or implant may suitably include microbeads (e.g., spherical or non-spherical) in a volume that is approximately 30% to 70% of the total volume of the soft tissue device. In certain embodiments, a soft tissue device or implant may suitably include microbeads (e.g., spherical or non-spherical) in a volume that is approximately 35% to 65% of the total volume of the soft tissue device. In certain embodiments, a soft tissue device or implant may suitably include microbeads (e.g., spherical or non-spherical) in a volume that is approximately 40% to 60% of the total volume of the soft tissue device. In certain embodiments, the soft tissue device or implant may suitably contain microbeads (eg, spherical or non-spherical) in a volume that is about 50% of the total volume of the soft tissue device.

[0138] In certain embodiments, a soft tissue device or implant may suitably include microbeads (e.g., spherical or non-spherical) in an amount of about 10% to 90% of the total weight of the soft tissue device. In certain embodiments, a soft tissue device or implant may suitably include microbeads (e.g., spherical or non-spherical) in an amount of about 20% to 80% of the total weight of the soft tissue device. In certain embodiments, a soft tissue device or implant may suitably include microbeads (e.g., spherical or non-spherical) in an amount of about 30% to 70% of the total weight of the soft tissue device. In certain embodiments, a soft tissue device or implant may suitably include microbeads (e.g., spherical or non-spherical) in an amount of about 40% to 60% of the total weight of the soft tissue device.

[0139] In certain embodiments, the microbeads are gelled or cured before the biologically active material is added to the microbeads. In certain embodiments, the microbeads are gelled or cured after the biologically active material is added to the microbeads. The microbeads can be gelled by chemical reaction (e.g., cross-linking) or UV irradiation so that the polymer cures or polymerizes to form nanofibers.

[0140] In certain embodiments, the soft tissue device or implant may further comprise a compound selected from the group consisting of growth factors, angiogenesis-stimulating compounds, immunomodulators, inflammation inhibitors, and combinations thereof, which may be mediated in or separately present in the fiber-hydrogel composite. In certain embodiments, the soft tissue device may further comprise one or more compounds having therapeutic, angiogenic, anti-angiogenic, anti-inflammatory, antibacterial, antihistamine effects, and combinations thereof, which may be mediated in or separately present in the fiber-hydrogel composite.

[0141] In certain embodiments, the soft tissue device has a tan delta value of less than about 0.27. Tan delta is the rheological loss modulus divided by the storage modulus, meaning that a lower tan delta number corresponds to a more "solid-like" as opposed to "liquid-like" material. Tan delta can also represent a rheological property, which can vary based on the oil or fat content of a material or substance.

[0142] In certain embodiments, the microbeads are substantially stable at room temperature for at least about 1 week, at least about 2 weeks, at least about 1 month, at least about 2 months, at least about 3 months, at least about 4 months, at least about 5 months, or at least about 6 months.

[0143] In certain embodiments, the soft tissue device is stable at a temperature of 37°C.

[0144] The soft tissue device or implant can be used for implantation or injection administration near a target tissue, in which the fiber-hydrogel composite or microbeads (e.g., spherical or non-spherical) made from the fiber-hydrogel composite can be administered topically or subcutaneously.

[0145] kit Provided herein are kits that include the fiber-hydrogel composites described herein and a suitable applicator. In certain embodiments, the applicator may include a syringe or injection needle.

[0146] In certain embodiments, the kit may include a vial containing one or more components selected from salts, buffers, and therapeutic agents. For example, the components in the vial may include, but are not limited to, lidocaine HCl at a concentration of about 3.0 mg / mL, potassium chloride at a concentration of 0.185 mg / mL, potassium dihydrogen phosphate at a concentration of 0.185 mg / mL, sodium chloride at a concentration of 7.40 mg / mL, and sodium phosphate dibasic at a concentration of 1.06 mg / mL.

[0147] In certain embodiments, the kit includes a fiber-hydrogel composite in the form of microbeads or microgels. Preferably, the microbeads have an average size in the size range of 50 to 400 μm.

[0148] How to use In one aspect, methods for producing the fiber-hydrogel composites described herein are also provided. The methods can include contacting a crosslinker with fibers (e.g., nanofibers or microfibers) containing one or more extracellular matrix proteins (ECM) and a hydrogel material, such as hyaluronic acid (HA), to obtain the fiber-hydrogel composite. The hydrogel material (e.g., HA) is suitably bonded to the fibers (e.g., nanofibers or microfibers) by the crosslinker to form a composite network.

[0149] In certain embodiments, the method may include adjusting the crosslinking conditions. In certain embodiments, the crosslinking conditions are adjusted to basic conditions, and the pH of the crosslinking conditions is in the range of about 9 to about 14, about 10 to about 14, about 10 to about 13, or about 10 to about 12. In certain embodiments, the pH of the crosslinking conditions is in the range of about 10 to about 14, or about 12 to about 13.3. For example, the pH of the crosslinking conditions is about 12.4, about 12.7, about 13.0, or about 13.3.

[0150] In certain embodiments, the contacting step is carried out for about 30 minutes to about 4 hours. In certain embodiments, the contacting step is carried out for more than about 30 minutes, more than about 1 hour, more than about 1.5 hours, more than about 2 hours, more than about 2.5 hours, more than about 3 hours, more than about 3.5 hours, or up to about 4 hours. In certain embodiments, the contacting step is carried out for about 30 minutes, about 1 hour, about 1.5 hours, about 2 hours, about 2.5 hours, about 3 hours, about 3.5 hours, or up to about 4 hours. In certain embodiments, the contacting step is carried out at a temperature of about 37 degrees Celsius.

[0151] In certain embodiments, the contacting step is carried out for less than about 2 hours, and in such cases, the contacting step may be carried out at an elevated temperature, for example, at a temperature of about 37-100 degrees Celsius, about 40-100 degrees Celsius, about 50-100 degrees Celsius, or about 60-100 degrees Celsius.

[0152] In certain embodiments, the contacting step is carried out for more than one day, in which case the contacting step may be carried out at a lower temperature, such as about 0-37 degrees Celsius, about 0-30 degrees Celsius, about 0-25 degrees Celsius, about 0-20 degrees Celsius, or about 0-15 degrees Celsius.

[0153] The method for producing a fiber-hydrogel composite can further include a step of swelling the fiber-hydrogel composite. In certain embodiments, the swelling step is carried out by incubating the fiber-hydrogel composite in an aqueous buffer (e.g., PBS). In certain embodiments, the buffer can be used at more than 10, 50, 100, 200, 300, 400, 500, or 1000 times the volume of the fiber-hydrogel composite. In certain embodiments, the swelling step is carried out for more than 1 hour, more than 2 hours, more than 5 hours, more than 10 hours, more than 12 hours, more than 15 hours, more than 20 hours, more than 24 hours, more than 30 hours, more than 40 hours, more than 50 hours, more than 60 hours, or more than 72 hours (continuous or discontinuous). Preferably, the swelling step is carried out at room temperature for 24 to 72 hours (continuous or discontinuous).

[0154] The method for producing a fiber-hydrogel composite can further include processing the fiber-hydrogel composite to form microbeads. In certain embodiments, the microbeads have an average diameter within the range of about 1 μm to about 1000 μm, about 10 μm to about 1000 μm, about 20 μm to about 1000 μm, about 30 μm to about 1000 μm, about 40 μm to about 1000 μm, about 50 μm to about 1000 μm, about 50 μm to about 500 μm, about 50 μm to about 400 μm, or about 100 μm to about 500 μm. As described above, the fiber-hydrogel composite can be formed into a microparticulate formulation (e.g., microbeads or microgels), allowing for the use of higher concentrations of each component and enhanced stability. In certain embodiments, a micronization system may be used, in which a preformed fiber-hydrogel composite is physically conditioned, such as by forcing it through one, two, three, or more mesh screens to create a population of non-spherical beads that are relatively similar in shape and size. In certain embodiments, the processing step may include mechanical grinding or mechanical screening, preferably by applying shear using a mesh. This two-screen system allows for tight control over the size of the beads, allowing the user to adjust the size as needed. Such microbeads (e.g., non-spherical ones) are disclosed in US 2020 / 0069846. For example, a fiber-hydrogel composite is conditioned by applying mechanical shear through a mesh with a defined size in the range of 50 to 400 μm.

[0155] In certain embodiments, the fiber-hydrogel composites produced herein can be in the form of two-phase or single-phase gels. For example, the fiber-hydrogel composites can be two-phase gels with similar bead sizes created by forcing the gel through a screen. Alternatively, the fiber-hydrogel composites can be single-phase gels with a continuous distribution of gel bead sizes created by homogenizing the gel during the crosslinking reaction or by mechanically disrupting the gel.

[0156] The method for producing a fiber-hydrogel composite can further include sterilizing the fiber-hydrogel composite. Any sterilization method or process known in the art can be used, including, but not limited to, sterilization. For example, the bagged composite can be sterilized by autoclaving at 118°C for approximately 5 to 30 minutes.

[0157] The method of producing a fiber-hydrogel composite can further include fabricating the fiber-hydrogel composite in a sheet or in an injectable fluid.

[0158] In certain embodiments, the method can further comprise incubating infiltrating macrophages in the fiber-hydrogel composite. In certain embodiments, the fiber-hydrogel composite containing collagen nanofibers can condition a population of M2 phenotype infiltrating macrophages.

[0159] Methods of treating a subject using the fiber-hydrogel composite are also provided.

[0160] In one aspect, the present disclosure provides a method for forming adipose tissue in a subject. The method includes administering to the subject a fiber-hydrogel composite described herein. The fiber-hydrogel composite induces prolonged in vivo retention or enhanced host cell infiltration in the subject. In certain embodiments, the fiber-hydrogel composite includes cells or growth factors. In certain embodiments, the fiber-hydrogel composite does not include cells or growth factors.

[0161] In certain embodiments, the subject exhibits neovasculature formation driven by M2 macrophage polarization. In certain embodiments, the subject exhibits an approximately 2.5-fold increase in α-SMA+ cells at the administration site. In certain embodiments, the subject exhibits localization of endothelial cells and CD163+M2 macrophages at the administration site. In certain embodiments, the subject exhibits an approximately 1.5-fold increase in CD68+ pan-macrophages at the administration site by post-operative day 7. In certain embodiments, the subject exhibits an approximately 2-fold increase in CD68+ pan-macrophages at the administration site by post-operative day 14. + It presents with an increase in pan-macrophages.

[0162] In one aspect, the present disclosure also provides a method for delivering cells or tissues in a subject, the method comprising encapsulating one or more cells or tissues in a fiber-hydrogel composite described herein to form a suspension; and applying the suspension to a target site in the subject.

[0163] In another aspect, the present disclosure also provides a method for delivering adipose tissue to a subject. The method includes encapsulating one or more adipose tissues in the fiber-hydrogel composite described herein to form a suspension; and applying the suspension to a target site in the subject. In certain embodiments, the one or more adipose tissues suitably comprise adipose-derived stem cells, adipocytes, or a combination thereof.

[0164] In one aspect, the present disclosure also provides a method for delivering a pharmaceutical agent in a subject, the method comprising combining the pharmaceutical agent and a fiber-hydrogel composite described herein to form a mixture; and applying the mixture to an intended delivery site.

[0165] As discussed above, the fiber-hydrogel composites described herein can be advantageously used in numerous tissue repair situations, as well as in other applications, such as providing coatings on catheters and other surgical devices and implants. The fiber-hydrogel composites can also be used to deliver active agents described herein, such as antibiotics, growth factors, and immunosuppressants. For example, the present disclosure provides methods for healing soft tissue defects, including applying a fiber-hydrogel composite to the soft tissue defect.

[0166] It will be appreciated that the advantageous properties of the fiber-hydrogel composites described herein include: (1) the ability to provide easy characterization and quality control; (2) the ability to integrate with existing tissue matrices; (3) the ability to directly incorporate into newly formed matrices; (4) the ability to directly contain cells and bioactive factors; (5) the ability to maintain biocompatibility; (6) the ability to control bioresorption; (7) the ability to easily mold into complex anatomical shapes due to the increased structural rigidity provided by the nanostructure; (8) the ability to exhibit the mechanical properties of natural tissues such as articular cartilage; (9) the ability to crosslink the fiber and hydrogel phases in a single step; and (10) the ability to retain mechanical properties after terminal sterilization by autoclave.

[0167] In another aspect, fiber-hydrogel composites can be used to repair cartilage tissue. Current biologically based surgical procedures for cartilage repair include autologous chondrocyte implantation, drilling, abrasion chondroplasty, microfracture, and mosaic arthroplasty. All of these procedures treat only localized articular cartilage damage and do not address exposed cartilage surfaces, such as those seen in severe osteoarthritis and rheumatoid arthritis. Furthermore, they use either cartilage tissue plugs harvested from the patient or expanded chondrocytes to fill the cartilage defect. These tissues 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) that possesses the biomechanical properties of normal cartilage. However, all such procedures promote the formation of repair tissue (fibrocartilage) rather than true hyaline cartilage, which may incur additional mechanical damage and predispose the joint to osteoarthritis. Furthermore, the availability of endogenous cartilage as a repair material is extremely limited, and its acquisition presents its own risks and morbidity to the patient. As is evident from the foregoing discussion, the resulting hydrogel / nanofiber compositions disclosed herein are viable materials for offering new therapeutic prospects in patients suffering from cartilage degenerative diseases.

[0168] As described herein, fiber-hydrogel composites can be prepared with a wide variety of properties suitable for any number of synthetic tissue implants or augmentations and other clinical applications. As previously described, fiber-hydrogel composites can be used to repair cartilage defects resulting from either injury or disease. Defects resulting from injuries that can be repaired in this way can be sports- or accident-related, and can involve only the superficial cartilage layer or can include the underlying subchondral bone. Disease-related defects that can be repaired using the compositions described herein include those resulting from osteoarthritis and rheumatoid arthritis. Whether resulting from injury or disease, such defects can be in either mature or growth cartilage plates. Formulations for hydrogels for synthetic growth cartilage plates may require the inclusion of a non-substituting scaffolding material to allow for controlled bioabsorption of the biomaterial during growth.

[0169] Another area in which the fiber-hydrogel composites described herein may be useful is the repair, reconstruction, or augmentation of cartilage 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 fields of plastic and reconstructive surgery. Considerable research and investigation has been conducted into the development of materials with suitable biocompatibility and longevity. The results of this research have not been encouraging. When placed in immunocompetent animals, the structural integrity of fiber-hydrogel composites has been shown to be compromised as the framework is absorbed. Furthermore, while traditional synthetic materials offer excellent longevity, they have presented certain unavoidable pitfalls. For example, silicone raises safety and long-term immune-related concerns. The synthetic polymers PTFE (Gore-Tex) and silastic exhibit low tissue reactivity but do not result in tissue integration, potentially posing a long-term risk of foreign body infection and expulsion. It is believed that fiber-hydrogel composites may be useful in preparing synthetic soft tissue scaffolds for augmentation or repair of soft tissue defects in the head and neck. Specifically, hydrogel / nanofiber compositions are non-inflammatory and non-immunogenic, can be formulated to have an appropriate degree of viscoelasticity (see discussion herein), and can be used as effective implantable scaffolding materials.

[0170] Furthermore, fiber-hydrogel composites can be used as novel biocompatible and bioadaptable materials for preparing cartilage implants, which are frequently used in head and neck reconstructive procedures to repair cartilage or bone defects secondary to trauma or congenital anomalies. Applications specific to the ear include otoplasty and auricular reconstruction, which are often performed to repair cartilage defects resulting from trauma, neoplasms (i.e., squamous cell carcinoma, basal cell carcinoma, and melanoma), and congenital defects such as microtia. Applications specific to the nose include cosmetic and reconstructive procedures of the nose and nasal septum. Rhinoplasty, tip grafts, shield grafts, and spreader grafts are frequently used in cosmetic rhinoplasty. Nasal reconstruction after trauma, neoplasms, autoimmune diseases such as Wegener's granulomatosis, or congenital defects requires reparative cartilage. Nasal septum perforations are difficult to manage and frequently result in failed procedures. Because autologous or donor cartilage is often unavailable, cartilage grafts would be ideal for these applications. Pharynx-specific applications include laryngotracheal reconstruction, which typically requires the harvesting of rib cartilage in children, which is associated with morbidity. Auricular cartilage and nasal septum cartilage are often unsuitable for this application. Synthetic cartilage materials prepared from the hydrogels disclosed herein can be synthesized to meet each of the aforementioned applications by adjusting hydrogel synthesis parameters such as reagent concentration, substitution, and crosslinking rate. Laryngotracheal reconstruction is typically performed for airway narrowing due to subglottic stenosis or tracheal stenosis. The etiology can be traumatic (i.e., intubation trauma or tracheotomy) or idiopathic. Other possibilities include chin and cheek augmentation and lower eyelid eversion repair, in addition to numerous craniofacial applications. Note that for these applications, the strict mechanical properties of articular cartilage may not be necessary. The inclusion of cell populations or bioactive agents may also be desirable.

[0171] The fiber-hydrogel composites described herein can be used to repair and narrow nasal passages to prevent chronic fluid accumulation in the nasal passages, which usually leads to infection and crust formation after excessively large surgical resections. Another promising application is in laryngotracheal reconstruction in both children and adults, for example, as a result of laryngotracheal injury caused by intubation during surgical procedures such as cardiovascular surgery. The fiber-hydrogel composites described herein can also be used to provide cricoid cartilage replacements to protect the carotid artery after cancer-directed neck resections—the fiber-hydrogel composite can be placed between the carotid artery and the skin as a protective barrier against loss of the skin barrier. As a protective coating during neuronal regeneration of resected nerves—fibrous tissue often forms faster than regenerating neurons, preventing their eventual formation. Placing nerve endings within a fiber-hydrogel composite preformed tube can eliminate fibrous tissue formation from the regeneration site.

[0172] Fiber-hydrogel composites can also be used to repair any internal or external organ soft tissue defect. For example, fiber-hydrogel composites can be used in chin and cheek augmentation and lower eyelid eversion repair, in addition to numerous craniofacial applications. For cosmetic and reconstructive purposes in areas other than the head and neck, for example, as breast implants for breast augmentation surgery, and as wound sealants to fill voids left after lymph node removal (i.e., due to cancer) in the breast or neck, sealing lymphatic vessels and reducing uncontrolled fluid drainage into the excision site, which can lead to infection and other complications.

[0173] In addition to the uses described above, the fiber-hydrogel composites described herein can be used in other tissue engineering applications to generate 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 cartilage in artificial forms. The fiber-hydrogel composites can also be used to create synthetic non-orthopedic tissues, including but not limited to, vocal cords, vitreous humor, heart valves, liver, pancreas, and kidneys, using strategies and methodologies similar to those described above for the synthesis of cartilage in artificial forms.

[0174] Another area in which the fiber-hydrogel composites disclosed herein can be used is in gastrointestinal applications where it is necessary to treat or prevent the formation of scar tissue or strictures in abdominal or gastrointestinal organs. Several products, commonly referred to as "hydrogels," designed or intended to be useful in treating and preventing scarring and / or stricture formation already exist and are in various stages of clinical and FDA approval. The fiber-hydrogel composites disclosed herein have an advantage over other known hydrogels in that they may contain nanostructures that can provide support, shape, and strength to the hydrogel material. The fiber-hydrogel composites disclosed herein can be used in applications similar to those in which known hydrogels are used or intended to be used, including for the treatment of strictures or scarring in the gastrointestinal tract. Treatment involves injection of the fiber-hydrogel composite into a site of anticipated stricture to prevent scarring, or into an existing stricture site after treatment to dilate the narrowed GI tract and prevent the stricture from recurring.

[0175] The fiber-hydrogel composites described herein can also be used to treat esophageal strictures. Esophageal strictures are a common complication of gastroesophageal reflux disease (GERD). GERD is caused by the reflux of acid, bile, and other harmful stomach contents into the esophagus, damaging the esophageal lining cells. Approximately 7-23% of GERD patients develop esophageal strictures or fibrous scarring of the esophagus. Esophageal scarring is also caused by resection therapies used to treat Barrett's esophagus. A major complication of such resection therapies is that the resection lesion extends too deeply into the esophageal wall, resulting in esophageal scarring or strictures. Esophageal strictures prevent normal swallowing and are a major cause of patient morbidity. The materials described herein can be used to treat or prevent esophageal strictures resulting from GERD, Barrett's esophagus, and esophagectomy.

[0176] Fiber-hydrogel composites can also be used to treat Crohn's disease, which causes strictures or scarring that close or narrow the intestinal lumen, preventing normal bowel function. For example, fiber-hydrogel composites can be useful in treating or preventing such strictures.

[0177] The fiber-hydrogel composite can also be used in methods for treating primary sclerosing cholangitis (PSC). PSC is a rare disease of the liver's bile ducts. Bile ducts form a branching network within the liver and exit the liver through two major branches that merge into the common bile duct, which drains bile from the liver and gallbladder into the duodenum. Although the bile duct is very narrow, typically only 2 mm in diameter at its most distal end, it still must drain several liters of bile from the liver into the duodenum daily. Any blockage of these ducts can lead to a serious condition known as jaundice, which allows many toxins, especially hemoglobin breakdown products, to accumulate in the body. PSC is a scarring or structuring disease of the intrahepatic bile ducts and the extrahepatic bile ducts that connect the liver to the small intestine. Bile duct strictures in PSC can be treated or prevented using the hydrogel / nanofiber compositions of the present invention.

[0178] Fiber-hydrogel composites can also be used to treat chronic pancreatitis. Chronic pancreatitis is a chronic inflammatory disease of the pancreas that can involve scarring or narrowing of the pancreatic duct. Such narrowing blocks the drainage of pancreatic juice, which normally must exit the pancreas through a system of ducts or drainage channels and enter 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 severe complications, such as the pancreas self-digesting and the formation of life-threatening abdominal infections and / or abscesses. Pancreatic narrowing due to chronic pancreatitis can be treated or prevented using the hydrogel of the present invention.

[0179] The fiber-hydrogel composites can also be used to treat gallstone-induced bile duct and pancreatic duct strictures. Gallstones are a very common disorder, and their major complication is the formation of bile duct and pancreatic duct strictures, which can be treated or prevented using hydrogels. Regarding the treatment of ischemic bowel disease, the intestine is prone to scarring or stricture formation when its blood supply is impaired. Impaired blood flow, known as ischemia, can be caused by many conditions, including cardiovascular disease, atherosclerosis, hypotension, hypovolemia, renal or liver disease-induced hypoalbuminemia, vasculitis, drug-induced disease, and many others. The end result of all of these causes can be intestinal strictures, which close the intestine and prevent its normal function. The hydrogel / nanofiber composites of the present invention can be used to treat or prevent ischemic bowel strictures.

[0180] The fiber-hydrogel composites can also be used to treat radiation-induced intestinal strictures. Radiation therapy for cancer is associated with numerous pathologies, not the least of which is the formation of intestinal strictures. The fiber-hydrogel composites can be used to treat or prevent radiation-induced intestinal strictures.

[0181] In addition to creating synthetic tissue or repairing natural tissue, the fiber-hydrogel composites disclosed herein can also be used to provide coatings for non-biological structures or devices to be used in surgery or otherwise for in vivo implantation, such as surgical instruments or ceramic or metallic prostheses. Such coatings would provide a barrier between the non-biological device material and living tissue. The role of fiber-hydrogel composites as barriers for non-biological devices includes, but is not limited to, (1) preventing the adsorption of macromolecules and / or cells onto the surface of non-biological devices, which can lead to protein contamination or thrombosis on the device surface; (2) providing a non-toxic, non-inflammatory, non-immunogenic, and biologically compatible surface to devices made from materials that are not otherwise biologically compatible; (3) compatibility with device function, such as glucose diffusion to a glucose sensor, mechanical force transmission to a pressure sensor, or endothelialization of a vascular graft or stent; (4) enhancing device function, such as providing a charge barrier to existing size barriers in MEMS-based artificial nephrons; (5) incorporating live cell populations into non-biological devices that are trapped within an aqueous, physiologically compatible environment; and (6) inclusion of drugs or bioactive factors, such as growth factors, antivirals, antibiotics, or adhesion molecules, designed to promote angiogenesis, epithelialization, or endothelialization of the device.

[0182] Based on the foregoing, the fiber-hydrogel composites can be used to provide non-allergenic coatings for a variety of implantable devices, including implantable glucose sensors for diabetes management. Furthermore, the fiber-hydrogel composites can be used to provide a charge barrier for the development of MEMS-based artificial nephrons, an aqueous, physiologically compatible environment in which embedded kidney cells, such as podocytes, can be incorporated into 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 biodetection systems.

[0183] The fiber-hydrogel composites of the present disclosure (specifically, containing HA) can be covalently attached to silicon-based devices by first covalently attaching the silicon surface to a hydroxyphenyl-coated surface chemistry, for example, via the primary amine of tyramine. This can be the same chemistry used to attach DNA modified with free amines to silicon surfaces. The HA-based fiber-hydrogel composite is then covalently bonded to the hydroxyphenyl-coated surface using the same peroxidase-driven chemistry used in the preferred crosslinking mode described above.

[0184] Fiber-hydrogel composites can also be used to coat non-biological cardiovascular devices such as catheters, stents, and vascular grafts. These would include devices made from materials not previously used due to their biological incompatibility, and which have superior design features than currently used devices. Incorporation of bioactive factors into the hydrogel can promote endothelialization or epithelialization of the hydrogel and, in turn, of the implanted device.

[0185] Although specific examples and uses for the hydrogel / nanostructured composites of the present invention are described herein, such specific uses are not intended to be limiting. The hydrogel / nanostructured composites of the present invention can be used in any application commonly used for known hydrogels, and are particularly useful for the repair and / or regeneration of soft tissue anywhere in the body. [Example]

[0186] Introduction Crosslinked hydrogels incorporating bioactive components show promise as adipose tissue substitutes in soft tissue repair. Collagen, a bioactive and biostimulable component, has been widely used as a filler and wound dressing, but its volume retention and resulting inflammatory response have been drawbacks. In this study, we report nanofiber-hydrogel composites (NHCs) incorporating biostimulable collagen nanofibers and hyaluronic acid (HA) networks with mechanical reinforcement. By adjusting the crosslinking conditions and collagen nanofiber loading density, we developed NHC constructs capable of attracting 2.5-fold more host cells into the injection site in vivo compared with hydrogel controls. Furthermore, the long-term stimulatory effect of NHC retention conditioned infiltrating macrophages toward a pro-regenerative M2 phenotype and promoted neovasculature formation. Consequently, improved adipose tissue formation was observed in NHCs near the vasculature, a sign of accelerated tissue remodeling. This highlights the benefits of long-term retention of biostimulated collagen fibers in a hyaluronic acid (HA) network and offers a new approach to soft tissue engineering. Surgeries such as tumor resection and wound repair, as well as burns, often result in the loss of adipose tissue. The absence of such tissue can cause reduced functionality and patient inconvenience [1, 2]. Traditional means of restoring adipose tissue have certain drawbacks, including cost and potential morbidity if surgery is involved. Autologous tissue repair, which requires surgery to transfer tissue from another part of the body, is prone to donor-site deformity [3]. Allogeneic transplantation can be associated with several potentially fatal complications, such as graft failure and graft-versus-host disease [4]. Although synthetic implants have gained more acceptance in adipose tissue repair, they still have the disadvantage of a foreign body response followed by host tissue necrosis [5].

[0187] Naturally derived materials, such as collagen, silk, and extracellular matrix, have been engineered as scaffolds to enhance adipose tissue reconstruction [6, 7]. These materials have attracted attention due to their biocompatibility and controllable degradation rate. While various studies have shown that scaffolds developed from these materials have robust biostimulatory effects that attract ingrowth cells, they come with numerous tradeoffs, including high lot-to-lot variability, limited control over design parameters, strong regulatory oversight, high cost, and the potential for deformation or loss of structural integrity [6].

[0188] Hydrogels are the most popular tissue restoration scaffold matrix in the field of soft tissue repair due to their biomimetic characteristics that closely resemble the tissue microenvironment and their ability to adapt to different adipose tissue needs, such as volume and porosity [8, 9]. Various studies have shown that hydrogel scaffolds with a storage modulus (G') similar to that of natural soft tissue (100-500 Pa) can promote volume and shape retention [10, 11]. However, the elastic and viscoelastic properties of hydrogels come at the expense of a high crosslinking density, which, if the hydrogel is not sufficiently biodegradable, can reduce the rate of cell infiltration and host tissue ingrowth, impairing regeneration and tissue remodeling

[12] . Among various hydrogel products, hyaluronic acid (HA) is widely used as a filler and foam for soft tissue augmentation due to its biocompatibility as an endogenous polysaccharide, its moderate biodegradability, and physical properties similar to those of soft tissues such as adipose, fibrous, and nerve tissues [ 13 , 14 ].

[0189] Studies have shown that the skewed polarization of macrophages activated by the mechanical properties of scaffolds predicts the extent of architectural remodeling, from fibrotic responses to matrix remodeling (M1 macrophage-dominated pro-inflammatory responses vs. M2 macrophage-dominated pro-regenerative responses) [15, 16]. Therefore, the ability of hydrogel scaffolds to promote macrophage infiltration and programming in acute and chronic inflammation is essential during tissue remodeling. Electrospun nanofiber meshes have been widely used as regenerative substrates to mimic the extracellular matrix of host tissue

[17] . Studies have shown that macrophages interacting with electrospun fiber meshes result in a bias toward M2 macrophage polarization

[18] . However, such nanofiber mat devices have numerous disadvantages. As solid, two-dimensional mats, the scaffolds are non-injectable, have inherently anisotropic mechanical properties, and are not suitable for volumetric applications. These interrelated variables also limit device design options. For example, for mats, scaffold porosity is dictated by fiber diameter and orientation, which also determines the preferred mechanical properties of the scaffold; consequently, mechanical properties and porosity cannot be easily modified independently of each other.

[0190] Therefore, to maximize therapeutic efficacy, our group's earlier work demonstrated a robust therapeutic system fabricated by chemically attaching hyaluronic acid (HA) to poly(ε-caprolactone) (PCL) nanofibers, enhancing the material's porosity and cell permeability while maintaining a storage modulus compatible with adipose tissue [19, 20]. In this study, we report the synthesis and characterization of a hydrogel-nanofiber composite system using hyaluronic acid (HA) covalently attached to electrospun bovine type I collagen nanofiber fragments to meet critical requirements for macrophage attraction and programming, host endothelial cell infiltration and distribution, and eventual adipose tissue formation. By using divinyl sulfone (DVS) as a cross-linker reactive with hydroxyl and amino groups, we introduced interfacial bonding between the HA hydrogel and collagen nanofibers to form a composite network

[21] . This system leverages improved mechanical properties over nanofiber-hydrogel blends, superior injectability due to a loose crosslinking density, and the feasibility of terminal sterilization by autoclaving. The material's biocompatibility, cell adhesion, and cell migration potential were evaluated in vitro using human mesenchymal stem cells (MSCs) and human umbilical vein endothelial cells (HUVECs). In vivo studies were performed using subcutaneous injections into the backs of SD rats, using H / E staining to analyze host cell infiltration, and immunofluorescence with various markers to examine macrophage attraction and conditioning, angiogenesis, and adipose tissue regeneration.

[0191] Example 1: Collagen fiber production Bovine type I collagen solution was purchased from Advanced Biomatrix. After first lyophilizing the bovine collagen solution overnight to obtain collagen powder, type I collagen solution (8% w / v) was prepared in 1,1,1,3,3,3-hexafluoro-2-propanol (HFIP) at room temperature for 6 hours to create a viscous, cloudy electrospinning solution. Electrospinning was performed using the following parameters: a flow rate of 5 mL / h; an applied voltage of 20–25 kV to a 22-G metal needle; a collection distance of 12.5 cm; and a metal collector rotation speed of 900 rpm. These parameter settings resulted in an average fiber diameter of approximately 600 nm (Figure 1C). Using carbodiimide chemistry, the fibers were immersed in an ethanol solution (95% v / v%) containing 50 mM 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC) and 20 mM N-hydroxysuccinimide (NHS) for 24 hours. After crosslinking, the fibers were washed three times for 5 minutes in 0.75% glycine solution to remove excess reagent and quench the activated fiber surface. The collagen fibers were then disassembled into fragments using cryomilling (Freezer / Mill 6770, SPEX SamplePrep). The fragments were filtered through different cell strainers (40 and 100 μm) to achieve a relatively uniform fiber length.

[0192] Example 2: Preparation of collagen fiber-HA hydrogel composite The proposed NHC construct consists of three components: a hyaluronic acid (HA) network, bovine type I collagen nanofibers, and a divinyl sulfone (DVS) cross-linker (Figure 1A). Before incorporating collagen nanofibers (obtained in Example 1) into the HA network during cross-linking, we first optimized the cross-linking conditions for the HA gel phase alone.

[0193] Sodium hyaluronate (MW 1.5 MDa) was purchased from LifeCore. HUVEC and vascular endothelial cell culture media were purchased from Lonza. All other chemical reagents were purchased from Sigma-Aldrich. All other cell culture reagents and additives were obtained from Invitrogen. HA was dissolved in distilled water at a stock concentration of 25 mg / mL. DVS concentrations were calculated as a ratio to hydroxyl groups in HA (e.g., 1.17 w / v%, 2.34 w / v%, and 4.68 w / v%). The stock HA solution was diluted to 2 w / v% using distilled water and sodium hydroxide to obtain four different pH values ​​(12.4, 12.7, 13.0, and 13.3), with other parameters kept the same (2 w / v% HA, 37°C, 3 h reaction time). Multiple samples were prepared and the reaction time or gelation kinetics was performed by measuring the mechanical properties at various time points (30 min, 1 h, 2 h, 3 h, 4 h, 8 h, and 16 h) to determine the time point at which stiffness plateaued. The crosslinking of NHCs followed the same conditions as for HA hydrogels, but different fiber densities (0, 1, and 3 w / v%) were added to the mixed precursors to test the gelation kinetics of NHCs. After gelation of the hydrogels and NHCs, dialysis was performed against pH 7.4 phosphate buffer using a dialysis membrane (6000-8000 MWCO, Spectrum) for 48 h to remove unreacted DVS, balance the pH, and allow the samples to swell for further study. Mechanical properties were then measured again after swelling. Microgels were then generated using stainless steel wire cloth discs, as previously reported, to reach a gel particle size of approximately 100 μm.

[0194] We found that cross-linking HA networks at a pH of 12.7 using DVS chemistry was effective in forming robust cross-linked hydrogels while minimizing degradation of HA molecules and collagen fibers during gelation (Figure 2A). The reaction pH had a significant effect on the resulting gels in the pH range of 12 to 13.3. While other parameters (37°C, 2% HA w / v concentration, 2.93% DVS w / v concentration) were kept the same, the reaction pH was set at 12.4, 12.7, 13.0, and 13.3 using different NaOH concentrations (0.001 M, 0.01 M, 0.1 M, 1 M). Reactions at pH 13.3 and 13.0 showed dramatic degradation after 2 hours of reaction, resulting in two unstable hydrogels with poor reproducibility.

[0195] The crosslinking time was optimized to approximately 2 hours to reach the maximum storage modulus and limit degradation (Figure 2B). After adjusting the DVS chemistry, we introduced nanofibers into the HA network while crosslinking to create interfacial bonds between the HA network and the nanofibers. We observed a reinforcement effect by comparing HA and NHC at similar crosslinking densities (Figure 1B). The close association between the fiber and gel phase resulted in a robust, porous gel (Figure 1E). Furthermore, the composite could easily pass through a 27-gauge needle after crosslinking (Figure 1D). To further quantitatively examine this reinforcement effect, we measured the storage modulus G0' of the HA hydrogel phase (normalized to the storage modulus control), G' of the total NHC, and G' of the hydrogel-nanofiber blend without interfacial bonds. In rheological tests with fiber loading densities ranging from 1 to 3 w / v%, the G' of the composites ranged from 1.5 to 4 times higher than that without interfacial bonding, and the G' difference increased with increasing fiber loading and crosslinker concentration (Figure 2C). Furthermore, we also investigated the effect of fiber length on stiffness enhancement by sorting out large fiber fragments using different cell strainers (40 μm, 100 μm, no sorting). Counterintuitively, gels with fiber fragments between 40 μm and 100 μm in length produced the greatest stiffness enhancement (Figure 2D), but this relative enhancement was minimized at the highest crosslinker concentration.

[0196] As a result, this preparation process allowed us to generate NHCs with post-crosslinking G' values ​​ranging from 450 Pa to 1500 Pa (Figure 2E) and post-swelling G' values ​​ranging from 150 Pa to 1000 Pa. To mimic the soft tissue microenvironment, HA control (G' = 100 Pa and G' = 250 Pa) and composite (G' = 250 Pa and G' = 100 Pa) were generated and characterized into microgels with diameters of approximately 100 μm (Figure 1A). The storage moduli of all three groups were measured, and these storage moduli were not significantly different from the initial crosslinked bulk gel (Figure 2F). Finally, before utilizing the prepared materials for subsequent in vitro and in vivo studies, we autoclaved the hydrogels and composites and observed that the G' measurements did not decrease significantly after terminal sterilization, demonstrating the application potential of this material. Autoclaving was performed to sterilize the hydrogels and NHCs after gelation. Briefly, the gels were subjected to an autoclave cycle at 118 °C with a 5-minute sterilization step. The entire sterilization cycle took 30 minutes to complete

[22] . After sterilization, the mechanical properties of each gel were again measured using rheological testing, demonstrating excellent thermal stability, allowing for terminal sterilization and a robust shelf life (Figure 2G).

[0197] Example 3: NHC mechanical characterization Rheological characterization of hydrogels and NHCs was performed as previously described. Using a rheometer (AR2, TA Instruments), various shear mechanical properties of hydrogels and NHCs were measured at 25 °C using an 8 mm parallel plate geometry with a 0.5 mm gap. The linear viscoelastic region of the gels was measured by strain sweeping at a set frequency (1 Hz) with increasing shear strain amplitude. Three key properties were storage modulus (G'), loss modulus (G''), and tan delta (G'' / G') (Figure 2A-G). The injection force of HA and NHC bulk gels and microgels was measured by extruding 1 mL of sample through a 27-gauge NIPRO needle in a 1 mL BD syringe using an Instron load frame (34SC-05, Norwood, MA). Statistical analysis data are presented as mean ± SEM. Data were analyzed by Student's t test (unpaired and two-tailed), one-way or two-way ANOVA (analysis of variance), followed by Dunnett's or Bonferroni post hoc tests, as appropriate, using GraphPad Prism software (GraphPad Software Inc.). Values ​​were considered significantly different at P < 0.05.

[0198] Example 4: Characterization of NHCs in vitro cell culture Human umbilical vein endothelial cells (HUVECs) were purchased from Lonza and cultured in EGM (Lonza) and incubated at 37°C under 5% CO2. Human adipose-derived stem cells (hADSCs) were maintained in Dulbecco's modified Eagle's medium containing 10% fetal bovine serum and incubated at 37°C under 5% CO2 before passage 3. Uniform 200 μm-sized cell spheroids were generated using agarose hydrogel microwells as previously described by the inventors. For three-dimensional (3D) culture of hADSC spheroids, plates were coated with crosslinked hydrogel or NHC and centrifuged at 300 g for 5 minutes

[19] . The hADSC spheroids mixed with hydrogel or NHC were then seeded on the preformed gel and cultured for 7 days.

[0199] In vitro evaluation Cell viability was analyzed using a Live / Dead cell viability kit (Sigma-Aldrich). Cells were seeded (50 μL, 5 x 10 cells) per cell type into HA hydrogel- or NHC-coated 96-well plates containing 150 μL of medium. 6 Cells were cultured at a concentration of 1000 cells / mL for 1, 4, and 7 days. Cell adhesion assays were similarly performed by seeding cells into pre-coated wells. After 4, 24, and 48 hours, the percentage of non-adherent cells was determined by measuring the number of cells in the medium suspension. Cell migration and spreading of hADSCs in 2D and 3D cultures were investigated by immunostaining. Cells were fixed with 4% w / v paraformaldehyde for 10 minutes and then stained with phalloidin 568. Cell nuclei were counterstained with 4',6-diamidino-2-phenylindole dihydrochloride (DAPI, Molecular Probes). All images were acquired using an LSM 780 confocal microscope.

[0200] To further characterize the versatile capabilities of this composite construct, we first conducted in vitro studies by culturing hADSCs and HUVECs on the three materials we prepared: 100 Pa HA, 250 Pa HA, and 250 Pa NHC (G = 100 Pa). Using a live / dead assay, we assessed the cell viability of cells encapsulated in the three materials. We observed that the cell viability of both hADSCs and HUVECs was greater than 90% for all three materials and showed no significant decrease on days 4 and 7 compared to day 1 (ns, P > 0.05; Figures 3A and 9A). We then investigated cell attachment in a 2D culture procedure and observed that 250 Pa NHC showed improved cell attachment at 4 hours after cell seeding, with the difference minimizing after 48 hours (Figures 3B and 9B).

[0201] To evaluate the improvement of cell migration due to the nanofiber component, we first cultured hADSCs on 100 Pa HA hydrogels and collagen fiber mats. As expected, cells cultured on fiber mats exhibited significant cell spreading behavior, while cells cultured on HA hydrogels clustered and failed to migrate (Figure 3D). We then investigated cell migration behavior in a 3D culture system in which hADSC spheroids were encapsulated in 100 Pa HA control and 250 Pa HA (G0' = 100 Pa) for 7 days. As a result, no significant cell migration was observed in the 100 Pa HA control, consistent with our previous study showing limited cell migration behavior in HA hydrogels

[23] . In contrast, cell spheroids in 250 Pa NHCs exhibited sprouting behavior consistent with cell behavior in 2D culture (Figure 3C).

[0202] Example 5: Subcutaneous injection in rats To investigate NHC-mediated macrophage polarization, angiogenesis, and adipogenesis in vivo, NHCs and hydrogels were injected into the subcutaneous space of Sprague-Dawley (SD) rats (6–8 weeks old). Three replicates were performed for each formulation at a volume of 200 μL per injection. Rats were sacrificed 7, 14, and 56 days after injection. Explant size was measured using a caliper to obtain a rough outline of the retention volume. The injected explants were then immediately reserved for immunohistochemistry and stored in 4% w / v paraformaldehyde (PFA) for 3 days after the animals were sacrificed. The explants were then serially dehydrated, embedded in paraffin, and sectioned to obtain slices within 10 μm. For histological staining, tissue slices embedded in paraffin were deparaffinized and then stained with hematoxylin and eosin (H / E stain) and Masson's trichrome stain for histological analysis. For quantitative analysis, three slices were stained per injection. The area of ​​cellular infiltration in the histological images was determined in ImageJ by selecting the embedded area, converting it to grayscale, and measuring the dark area (since the initial embedded material is lighter than the surrounding tissue, simple thresholding is possible).

[0203] For immunofluorescence staining, tissue slices embedded in paraffin were deparaffinized and rehydrated to activate the surface for immunostaining. Briefly, explants were immersed in xylene, 100% ethanol, 95% ethanol, and 70% ethanol for 10 min each, and then placed in IHC buffer in a boiling pot for 20 min to activate the tissue surface. The sections were then permeabilized with 0.5% Triton X-100 solution and blocked with 4% donkey serum in PBS for 2 h. The samples were then incubated with primary antibodies (Table 1) overnight at 4°C. Cy3 and Cy5 affinity secondary antibodies (Jackson ImmunoResearch Laboratories), including Cy3-conjugated donkey anti-mouse (Cat. No. 715-165-151), Cy3-conjugated donkey anti-goat (Cat. No. 705-165-147), and Cy5-conjugated donkey anti-rabbit (Cat. No. 711-175-152), were applied to the sections for 2 hours at room temperature. The sections were then washed three times with PBS and counterstained with DAPI for 15 minutes. Six random fields of each specimen were imaged using a ZEISS LSM 780 confocal microscope for quantitative analysis.

[0204] Table 1. Primary antibodies used in this study TIFF2026502902000002.tif90128

[0205] To investigate the biostimulatory effects mediated by collagen fiber components in vivo, we subcutaneously injected 200 μL of NHC (G' = 250 Pa and G' = 100 Pa) and two hydrogel controls (G' = 100 Pa and 250 Pa) into the dorsum of SD rats. After sacrificing the rats on POD 7, 14, and 56, we measured the shape retention of the materials by macroscopic imaging. Measurements showed that the three material groups underwent initial swelling, retaining a volume of approximately 300–400 μL by POD 56, followed by a small volume loss for the 250 Pa NHC and 250 Pa HA control (although the volume was still larger than the initial injection volume), but a large volume loss for the 100 Pa HA control group (Figures 4A and 4B). Furthermore, histological images revealed increased cell infiltration in NHCs at POD 7 (Figure 4C), with the cell infiltration area being approximately 17% compared with approximately 6% in the two hydrogel controls (P < 0.01). Cell infiltration continued to increase in NHCs at POD 14, with approximately 60% of the area covered by infiltrated cells, significantly higher than that of the hydrogel controls (P < 0.0001; Figures 4E and 4F). Cell density was also measured based on H&E staining images, which showed a consistent trend in infiltrated cells in NHCs compared with those in the hydrogel controls at POD 7 (P < 0.05) and POD 14 (P < 0.01) (Figures 4E and 4F). The continuous increase in cell infiltration within the retained NHCs at POD 7 and POD 14 suggests that such infiltrating cells may be recruited by the presence of biostimulatory collagen fibers, and this effect is maintained due to the good shape retention of the material.

[0206] At POD 56, H / E images and cell density revealed enhanced angiogenesis and adipogenesis, paralleled by some neotissue formation (Figure 4D). Cell density analysis revealed that the cell count within the 250 Pa NHC was approximately 1500 cells / mm. 2The results showed that the 250 Pa NHC-injected area reached a plateau at 100 Pa and was not significantly different from the 100 Pa HA control (ns, P > 0.05; Fig. 4F). Without being bound by theory, this may indicate that the area injected with 250 Pa NHC had already undergone tissue remodeling, thus allowing fewer additional cells to infiltrate into the NHC.

[0207] To understand how the biostimulatory effect of NHCs can enhance the tissue remodeling process, we investigated whether this response correlates with macrophage infiltration and polarization. Activated macrophages exhibited two phenotypes: pro-inflammatory M1 and pro-regenerative M2. Three markers, CD68, CD38, and CD163, were used in our study to identify pan-, M1-, and M2-macrophages, respectively [24-26]. At POD 7, 250 NHCs had approximately twice the CD68 expression compared to 100 Pa HA (P < 0.05) and 250 Pa HA (P < 0.01). + It was able to recruit pan-macrophages (Figures 5A and 5E). + M1 macrophages (P < 0.001) and CD163 + Both M2 macrophages (P < 0.05) began to be activated at higher levels in NHCs, which was confirmed by statistical analysis (Figures 5B, 5F, and 5G). At POD 14, CD68 + and CD38 + A similar trend was observed for cells, but CD163 in 250 Pa NHCs + M2 macrophage expression was found to be increased compared to the 100 Pa HA control (P < 0.01) and 250 Pa HA control (P < 0.001) (Figures 5B, 5D, and 5G). At POD 56, infiltrating CD68 in the three groups was consistent with the histological findings by H / E. +The number of pan-macrophages showed no significant difference (Figure 5E), suggesting that the enhanced macrophage recruitment in the NHC group seen at earlier time points had ceased by this time point. Furthermore, the CD38 expression in NHCs + The density of M1 macrophages was 72.9 ± 19.8 cells / mm 2 A significant decrease was observed, with 216.0 ± 52.0 and 192.0 ± 17.8 cells / mm , respectively, while the two hydrogel controls showed 216.0 ± 52.0 and 192.0 ± 17.8 cells / mm , respectively. 2 (P < 0.05; Figures 5E and 5G). Furthermore, this decrease in the M1 phenotype in 250 Pa NHCs was due to the CD163 + M2 macrophages were 331.0 ± 36.6 cells / mm 2 This was accompanied by an increased expression of 213.6 ± 19.4 cells / mm 2 This was in contrast to the 100 Pa HA control, which showed a significant reduction in macrophage infiltration (P < 0.05; Figures 5F and 5G). The improved macrophage infiltration at POD 7 in NHCs and the skewed polarization with a higher M2 / M1 ratio at POD 56 (Figure 5H) suggest that the porous hydrogel network reinforced by bioadhesive collagen nanofibers (as seen in Figure 1E) may accelerate the polarization of recruited macrophages toward a pro-regenerative phenotype, thereby hastening the tissue remodeling observed in the histological images (Figure 4D).

[0208] The pattern of early regenerative inflammation induced by NHCs was confirmed by quantitative gene expression analysis. The pan-macrophage-associated gene CD68 was upregulated in NHCs at days 7, 14, and 56 compared with HA controls, indicating a continuous biostimulatory effect for macrophage recruitment (Figure 5I). Furthermore, the M1-specific genes NOS1 and CD38 showed an early 5-fold increase in NHC matrices compared with HA controls at day 7 and subsequently decreased (Figure 5I). In contrast, the M2-specific genes Arg1 and CD163 were expressed early in NHC matrices and persisted through intermediate time points (Figure 5I).

[0209] Example 6: Biostimulatory NHC-mediated endothelial invasion and angiogenesis To examine the angiogenesis observed in histological images, angiogenesis was investigated in vivo as an aspect of tissue remodeling. Immunofluorescence was used to detect the angiogenesis of host vascular endothelial cells (RECA-1 + The infiltration area and rate of endothelial cells in the NHC group were analyzed

[27] . On POD 7, the NHC group showed the highest endothelial cell infiltration area: 250 Pa NHC: 9.2 ± 6.1%, 100 Pa HA: 1.7 ± 0.6% (P < 0.05), and 250 Pa HA: 0.2 ± 0.1% (P < 0.05); Figures 6A and 6B. The endothelial cell infiltration rate in the NHC group also accelerated to 67.5 ± 7.2 μm / day, which was significantly faster than those in the 100 Pa HA group (28.8 ± 5.9, P < 0.0001) and 250 Pa HA groups (8.7 ± 2.7, P < 0.0001) (Figures 6A and 6C). At POD 14, similar trends were observed regarding the endothelial cell invasion rate and invasion area, and the invasion rate in NHCs remained high at 66.5 ± 10.4 μm / day, indicating that endothelial cells began to diffuse widely within the NHCs (Figures 6A and 6C).

[0210] As of POD 56, RECA-1 + Cells plus α-SMA + By examining the vasculature, well-formed blood vessels were studied

[28] . As expected, α-SMA inside NHCs was + The vascular density was significantly higher than in the 100 Pa (P < 0.001) and 250 Pa (P < 0.0001) hydrogels (Figure 6D). Although the endothelial cell infiltration rate decreased to 33.6 ± 3.7 μm / day in the NHC, which may be due to the pre-existing distribution of host endothelial cells throughout the entire NHC construct, the infiltrated area increased to 64.9 ± 9.6% compared to the relatively small area in the hydrogel control (P < 0.0001) (Figures 6A, 6B, and 6C).

[0211] To understand whether programmed macrophage polarization altered local inflammatory behavior at the gene expression level, we used qPCR to measure a small number of pro-inflammatory genes (TNF-α, IL-1β) and pro-regenerative genes (IL-10 and TGF-β1). The qPCR results revealed that these two pro-inflammatory genes were upregulated (3.7-fold and 7.9-fold, respectively) in the NHC matrix at the early 7-day time point, followed by a low downregulation (1.7-fold and 0.4-fold, respectively) in the NHC matrix (Figure 6E). At the same time, we selected two anti-inflammatory genes, IL-10 and TGF-β1, to measure regenerative conditions. Similar to the observations in macrophage polarization, IL-10 was upregulated in the NHC matrix, indicating the development of regenerative behavior in the NHC matrix, whereas TGF-β1 measurements showed no significant differences among the three groups (Figure 6E).

[0212] The ingrowth of host endothelial cells in biostimulated NHCs corresponds to the enhanced cell infiltration and macrophage recruitment and conditioning observed in histological images. To understand the correlation between the angiogenic response and the skewed macrophage polarization within NHCs, we investigated the effects of RECA-1 on POD 14. + and CD163 + Further immunofluorescence of the cells revealed that a greater than 20% of the CD163 + Results showed that M2 macrophages localized near endothelial cell clusters, suggesting that early vasculature formation may be enhanced by pro-regenerative macrophages (FIG. 9C).

[0213] Example 7: Enhanced adipose tissue formation and its correlation with mobilized progenitor cells Our previous work suggested that adipose tissue begins to form within NHCs 28 days after in vivo injection

[19] . At 180 days, this trend continued to the point where the implanted NHC gel was substantially converted into vascularized adipose tissue, whereas the HA gel control was not (Figure 7A). The resulting mixture of adipose tissue and remaining NHC gel maintained volume at the implant site without fibrosis, possessed numerous functional blood vessels, and was tightly integrated with the remaining implanted gel (Figure 7B). The progressive conversion to adipose tissue was verified using immunofluorescence, and perilipin staining demonstrated an increase in lipid-laden mature adipocytes in the gel over time (Figure 7C).

[0214] The improved adipose tissue formation within NHCs indicates an acceleration of the tissue remodeling process promoted by biostimulatory NHCs. To investigate the mechanism of how adipose tissue formed within NHCs, we used immunofluorescence to characterize the connections between cell types over time. First, we co-stained the adipose tissue marker adiponectin (Acrp-30) with perilipin-1 within NHCs (characterizing mature adipogenesis). Results showed that Acrp-30 expression was also elevated within NHCs compared to HA controls at day 14, when approximately 160 cells / mm were present within the composite. 2 At a density of Acrp-30 +Adipocytes were already infiltrating at a rate approximately five times faster than that seen in the fiberless HA controls (P < 0.001), a trend that continued at day 56 (Figure 7D). Adipogenesis stained with perilipin-1 showed a similar trend (P < 0.0001), with numerous fat particles forming at the periphery of the injected 250 Pa NHCs. At POD 56, fat particles began to appear at the periphery of the 100 Pa HA controls, while more than a dozen such fat particles were distributed within the NHC interior (P < 0.0001 (Figures 7D and 7I)). Meanwhile, in the NHCs, adipocyte density continued to increase over time, reaching 134.5 ± 14.0 cells / mm2 by day 180, fivefold higher than either hydrogel control (P < 0.0001). 0.0001; Figure 7H). At day 180, not only were there more adipocytes in NHCs, but they were also significantly larger and contained mature lipid droplets (Figure 7J). These findings suggest that NHCs not only support adipocyte formation but also create a microenvironment favorable for adipocyte growth and distribution. Interestingly, Acrp-30 could be expressed on endothelial cells and adipocytes, and was observed very close to adipocytes in remodeling NHCs, depending on the vasculature (Figure 7B). Based on these observations, we hypothesized that adipocyte formation may have some correlation with vascular invasion and formation.

[0215] To further investigate this correlation, we performed immunofluorescence of Reca-1+ endothelial cells and perilipin-1+ adipocytes. At POD 14, close localization of endothelial cells and adipocytes was observed at the periphery of 100 Pa HA and 250 Pa NHCs, and at POD 56, it was widely distributed throughout 250 Pa NHCs (Figure 7E). Furthermore, Acrp-30 + Depending on the adipocyte signal, RECA-1 +They even intermingled with host endothelial cells, indicating a strong correlation between adipogenesis and angiogenesis. Previous studies have shown that vascular progenitor cells have the potential to differentiate into adipocytes and promote tissue remodeling [31, 32]. Therefore, we added the CD107a marker to identify infiltrating perivascular progenitor cells. CD107a in 250 Pa NHCs + In addition to increased pericyte infiltration, results on POD 14 revealed the closely localized signals of RECA-1, Acrp-30, and CD107a, indicating that progenitor cells recruited by biostimulated NHCs could promote angiogenic and adipogenic responses ( Figures 7F and 9D ).

[0216] As described above, we observed the close localization of endothelial cells and adipocytes within 250 Pa NHCs at days 14, 56, and 180 (Figure 9E). Therefore, to understand the correlation between vasculature formation and adipocyte differentiation or origin, we performed co-staining of Pref-1+ preadipocytes and Reca-1+ endothelial cells. Confocal images showed that Pref-1+ preadipocytes also localized near endothelial cells (Figure 9F). Furthermore, quantitative analysis showed that almost no Pref-1+ cells were observed within NHCs at day 14. On the other hand, approximately 5% of cells began to express preadipocyte markers at day 56, and this percentage remained consistent until day 180 (Figure 9H). These observations suggest that the cell population that begins to differentiate into adipocytes is small, and the origin of these adipocytes remains unknown.

[0217] Several recent studies have identified perivascular cell populations, specifically PDGFRα+ cells, as adipocyte stem cells capable of differentiating into preadipocytes and ultimately mature adipocytes. Remarkably, immunofluorescence imaging showed that the number of PDGFRα+ cells in NHCs continued to increase from days 14 to 180 and that they localized in close proximity to endothelial cells (Figures 9G and 9H). These results, combined with the kinetics of tissue remodeling, suggest a self-sufficient mechanism by which NHCs generate adipose tissue in vivo. Specifically, NHCs recruit and polarize macrophages in a pro-regenerative direction, triggering angiogenesis and the recruitment of adipocyte precursors into the perivascular space.

[0218] Example 8: HA concentration is crucial for successful gelation The gelation system may depend on the HA concentration and DVS concentration. When the HA concentration was 0.5 w / v% (5 mg / mL), the solution did not gel, regardless of the crosslinker concentration. At 1.0 w / v%, the solution did not gel at the low crosslinker concentration, and only gelled very weakly at the high crosslinker concentration, while at an HA concentration of 2 w / v% (20 mg / mL), the solution gelled strongly. (W / v% is weight percent based on the component weight in grams per 100 mL of solution; 2 wt% can also be expressed as 20 mg / mL. Weight per volume is equivalent to weight per weight (w / w%) if the composition has a density of 1 g / mL, which is approximately the density of these examples.) (Figure 8).

[0219] Example 9: The pre-swelling step is essential for successful sterilization of composites by autoclave To achieve the required thermal stability for autoclaving, dialysis and pH neutralization steps are necessary. Fiber-hydrogel composites were prepared using a formulation that produced a strong gel: 2 wt% HA, 1x DVS, 37°C, pH 13, and a fiber loading of 3 wt% into the composite. While the gelled sample was subsequently autoclaved, it was not dialyzed. After autoclaving, the resulting NHC was liquefied.

[0220] Example 10: Alternative Crosslinker: BDDE Crosslinked Composite Instead of DVS Crosslinked Composite To evaluate whether fiber hydrogel composites could be produced using alternative crosslinkers used in many commercial products, we attempted to produce gels using BDDE instead of DVS. Although using the same reaction conditions as the successful DVS-crosslinked gels described previously, we were unable to produce stable crosslinked gels, even with HA concentrations similar to those used in the DVS group. In these DVS-crosslinked groups, a 2.0 w / v% HA concentration was required to form strong gels. In contrast, groups crosslinked with 2, 4, 6, and 8 w / v% HA (all mixed with 1 v / v% BDDE at pH 13 (0.1 M NaOH solution) and reacted at 40°C) did not form gels. The group using 10 w / v% HA formed strong gels after 16 h of reaction, forming gels with a 300 Pa resistance. This is a 5-fold higher HA concentration than that required for DVS crosslinking.

[0221] Example 11: Fiber-hydrogel composites using recombinant collagen samples Electrospinning and crosslinking of Demulcent SFA Similar to the process used to produce electrospun porcine type 1 atelocollagen (Nitta) (Figure 10) and porcine gelatin (Sigma) (Figure 11) used in the preparation of fiber-hydrogel composites, we also produced electrospun collagen fibers using recombinant collagen Vecollan (Evonik) (Figure 12) or recombinant collagen Demulcent SFA (Jland Biotech) (Figure 14) and efficiently crosslinked these fibers, as shown in optical micrographs of the electrospun fibers after staining with picrosirius red colorant (Figures 13 and 15). These collagen fibers are stable in aqueous environments and can be formed into composites with HA and retain their fiber morphology after gelation and autoclaving.

[0222] As an example, a 13.2 wt% Demulcent SFA (human recombinant collagen; Jland Biotech) collagen solution dissolved in 2,2,2-trifluoroethanol was extruded from a 5 mL syringe through a 27-G blunt-tip needle at a rate of 2.0 mL / h. The potential was set at 15 kV at the needle tip and -4.5 kV at the collector, with the drum rotating at approximately 140 rpm. The fibers were then treated with 1% DVS in ethanol for 24 hours and 0.7% glutaraldehyde in ethanol for 20 hours.

[0223] Both DVS and glutaraldehyde can crosslink proteins through amine groups on the polypeptide chain. They target the amine groups at the N-terminus of peptides and the surface lysines of proteins. On the other hand, the use of DVS or glutaraldehyde alone does not efficiently crosslink either Vecollan or Demulcent SFA. Electrospun recombinant protein fibers are crosslinked via a two-step process. The resulting fibers have limited swelling in water and in gel composites.

[0224] Fiber Swelling Test When the fibers were suspended in water and stained with picrosirius red, minimal swelling was observed, indicating effective cross-linking of the collagen fibers (Figure 15).

[0225] Hydrogel composite preparation 2.08 g of HA stock solution was mixed with 0.42 mL of water and 15 mg of cross-linked collagen fibers, followed by 50 mL of 5 M NaOH. μ Prepare the hydrogel composite by adding 66 μL of DVS. Incubate the mixture at 30°C for 2 hours. Neutralize the reaction with 20 mM NaH2PO4 solution. Dialyze the gel composite against PBS, then centrifuge at 2500 x g and autoclave using the liquid cycle at 121°C for 30 minutes.

[0226] Rheometry tested according to Example 3 TIFF2026502902000003.tif46161

[0227] Consideration The inventors herein disclose, inter alia, mechanically reinforced nanofiber-hydrogel composites (NHCs) that not only function as soft tissue substitutes but also incorporate biostimulatory collagen nanofibers to promote host cell infiltration, macrophage conditioning, and eventual adipose tissue repair. The interfacial bonding between the HA network and collagen nanofibers provides an effective network to retain the collagen fibers within the composite, enhancing composite stiffness at relatively low fiber loading densities. Our results demonstrate that the materials and compositions of the present invention (including NHC gels) are able to attract host cell infiltration, condition infiltrating macrophages, enhance angiogenic responses, and, most importantly, accelerate adipogenesis compared to HA controls. Furthermore, the feasibility of terminal sterilization, compatibility with recombinant collagen, and excellent injectability of NHCs demonstrate the material's potential applications.

[0228] Previous studies have examined the effect of hyaluronic acid (HA) hydrogel storage modulus (G') on host cell infiltration and shape retention, but these studies used gel stiffness levels 5-10 times lower than the NHCs used in this example [33, 34]. Our previous work has shown that increasing the storage modulus can significantly impede host cell infiltration and neovascularization

[19] . Therefore, we generated an NHC sample with G' = 250 Pa (G' = 100 Pa) along with two HA hydrogel controls with G' = 100 Pa and G' = 250 Pa. When comparing the two HA controls, we observed that the 250 Pa hydrogel was nearly incapable of permitting host cell infiltration at POD 7 and POD 14, and that angiogenesis and adipogenesis were minimal at POD 56. Although the 100 Pa HA hydrogel showed a relatively high cell density compared to the 250 Pa HA hydrogel at POD 56, its storage modulus was not within the range of most soft tissues (150–500 Pa) and failed to mimic the host tissue architecture

[35] . The 250 Pa NHC had a similar crosslink density to the 100 Pa HA hydrogel, but with a collagen nanofiber loading of 3 w / v%, its storage modulus was enhanced to 250 Pa, which is within the range of 150–500 Pa, indicating its mechanical potential as a soft tissue substitute. Furthermore, the results showed that the biostimulatory collagen fibers recruited more than enough host cells (approximately 2.5 times denser than the 100 Pa HA hydrogel) to the NHC, and the infiltrating cells were uniformly distributed at POD 14, indicating that the gels remained cell-permissive and isotropic with uniform fiber distribution (anisotropic or non-uniform fiber loading would be expected to bias the cellular response toward more clustering in stiffer, fiber-rich regions).

[0229] Concurrently with the recruitment of large numbers of host cells by NHCs, we observed that the NHC material also attracted large numbers of macrophages, with approximately 1.5-fold higher CD68 expression at POD 7 compared to the 100 Pa HA control.+ Pan-macrophages were attracted to NHCs, and this ratio increased approximately twofold at POD 14, suggesting a prolonged biostimulatory effect of collagen nanofibers within NHCs. Our previous work has shown that PCL nanofibers can shift macrophage polarization toward M2

[19] . In this example, we observed that collagen nanofibers can also modulate macrophage status toward the M2 phenotype, as the M2 / M1 ratio in the NHC group was found to be approximately twice that of the HA control at POD 56. Furthermore, the decrease in the M1 phenotype in the NHC group at POD 56 compared to that at POD 14 suggests that such M1 phenotypes can be repolarized to the M2 phenotype within NHC materials. Alternatively, this tendency toward M2 polarization could be explained if M1 cells present at the implant site preferentially migrated away from the implant site, while incoming naive macrophages preferentially polarized toward the M2 phenotype.

[0230] Angiogenesis is thought to be limited in hyaluronic acid (HA) hydrogels [19, 36]. The use of NHCs significantly promoted the rate of endothelial cell infiltration, which was approximately 2.5-fold in NHCs compared with HA controls. α-SMA was also significantly increased in NHCs. + The number of cells also increased by approximately 2.5 times, suggesting an accelerated maturation of newly formed blood vessels. Furthermore, the present inventors found that endothelial cells and CD163 + We also found close localization with M2 macrophages, indicating that neovasculature formation can be promoted by skewed M2 macrophage polarization. Therefore, it appears that biostimulatory collagen nanofibers recruit macrophages and condition them toward an M2 phenotype, thereby promoting neovasculature formation in NHCs.

[0231] Generating soft tissue using hydrogels is challenging because spontaneous host cell infiltration and remodeling are slow processes [37, 38]. In this example, in vivo adipogenesis was accelerated by incorporating collagen nanofibers, which have been shown to promote progenitor cell adipogenesis in vitro

[39] . At POD 14, a few adipocytes began to form around the periphery of the NHC material. At POD 56, adipocytes became more abundant and distributed throughout the NHC. By POD 180, vascularized adipose tissue occupied nearly half of the implant area. Furthermore, these adipocytes were found to be localized in close proximity to endothelial cells. Because early adipose tissue formation requires a blood supply, vasculature structures likely form in close proximity to these adipocytes. We also found that CD107a + We have found that progenitor cells, such as progenitor cells, can promote adipogenesis by direct differentiation into adipocytes. Herein, using immunofluorescence imaging, we found that more CD107a + Cells are recruited to NHCs, which then recruit RECA-1 + Endothelial cells and Acrp-30 + We observed a similar trend of sharing a close location with adipocytes.

[0232] This improvement is possible because synthetic, yet biomimetic, composites or gels can induce macrophage recruitment and polarization, resulting in a controllable inflammatory response. Inflammation is an essential tissue response to injury, infection, and tissue loss to promote homeostasis. Depending on the local host tissue microenvironment and the degree of inflammation, outcomes typically involve incomplete fibrotic tissue formation or scarring. Considering the nature of inflammation in achieving a regenerative outcome, recent studies have shown that a programmed inflammatory response can optimize the regenerative response in different tissues. This is achieved by programming macrophages, which serve as a bridge between inflammation and remodeling. While macrophages also play a central role in the development of chronic inflammation, altering macrophage polarization can be programmed to promote a pro-regenerative outcome.

[0233] In our work, by subcutaneously introducing a biostimulatory collagen NHC matrix, we were able to create a sustained and mild inflammatory microenvironment accompanied by continuous host macrophage recruitment. Remarkably, instead of deteriorating toward chronic inflammation and fibrotic tissue formation, a pro-regenerative phenomenon was observed in the NHC matrix. Pro-regenerative CD163+ macrophages were more polarized in the NHC matrix compared with the HA control. Clearly, the biostimulatory properties of the NHC matrix altered the trajectory of the local inflammatory response. Following macrophage polarization, repair genes such as IL-10 and VEGF-α were upregulated in the NHC matrix, leading to increased angiogenic behavior and soft tissue formation over a 6-month study.

[0234] This research has several implications. First, the NHC induces excellent vascularization and soft tissue repair without the introduction of any exogenous cytokines or cells. Compared to most existing studies that focus on biologics to support repair, this NHC material can itself produce long-term repair effects. Furthermore, in preferred aspects, the composite of the present invention can be provided with a physical mixture or combination of exogenous growth factors or cytokines, which may further improve repair outcomes.

[0235] Another significance of these examples is the investigation of the correlation between polarization of recruited macrophages, induction of angiogenesis, increased adipose tissue formation, and infiltration of progenitor cells.

[0236] A final significant aspect is the applicability of this material, as terminal sterilization using an autoclave does not significantly alter its mechanical behavior. The safety records of HA, collagen, and DVS in clinical use also favor the future clinical application of NHC materials. Compatibility with alternative collagens, such as those of recombinant origin, allows for a fully synthetic product, further improving the applicability of this invention and overcoming the religious, lot-to-lot variability, and allergy issues inherent in animal-derived products.

[0237] In conclusion, we developed granular collagen nanofiber-reinforced hydrogel composites that function as soft tissue substitutes by mimicking the mechanical properties and microstructure of native tissue. The porous structure and biostimulatory collagen fibers of NHCs can promote host cell infiltration. NHCs can further program local inflammation by polarizing infiltrating macrophages toward a pro-regenerative M2-like phenotype, subsequently promoting angiogenesis and neovasculature ingrowth, leading to the formation of angiogenic adipose-like tissue in vivo. This adipogenic outcome was highly correlated with neovasculature in the remodeling matrix, suggesting that PDGFRα+ perivascular cells are a potential source of adipocyte precursors. NHC-programmed regenerative inflammation represents a novel mechanism enabled by our injectable biostimulatory matrix design and may have broad implications for soft tissue reconstruction and many other regenerative therapies. In preferred aspects, the compositions and composites of the present invention can provide soft tissue augmentation and restoration using a shelf-stable, ready-to-use gel (composite) that does not require the incorporation of cells and growth factors.

[0238] In preferred aspects, the beads used in the materials and compositions are in the form of irregular, non-spherical gel beads containing uniformly dispersed fibers (particles with aspect ratios well above 1.0) (with random orientation as opposed to aligned or planar orientation). The fibers can be engineered to decorate the gel bead surface and produce an isotropic gel. The gels can be fabricated into a variety of forms, including injectables for aesthetic or therapeutic use. The composites preferably have high thermal stability, allowing for the use of steam autoclaves for terminal sterilization and providing a long shelf life at ambient temperatures.

[0239] In certain aspects, the mechanical and biostimulatory properties of the composites of the present invention may depend on the high aspect ratio fibrous component, which differs from other particulate morphologies with low aspect ratios, such as spheres.

[0240] As referred to herein, aspect ratio is the length of a particle divided by its width. Preferred fibers of the compositions and composites of the present invention have a diameter or cross-sectional dimension of approximately up to or about 0.6, 0.7, or 0.8 μm, or from about 0.4 or 0.5 μm to 0.8 or 0.9 μm, or from 0.5 or 0.6 μm to 0.7 or 0.8 μm, with even the shortest fiber sections (3 μm long) having an aspect ratio of at least 5, and longer fibers (100 μm or longer) having an aspect ratio of 166 or greater.

[0241] Composite materials of the invention having an average aspect ratio of at least 20 (corresponding to fibers 20 μm long and 1 μm in diameter) are expected to provide improved mechanical and biological responses compared to gels having aspect ratios (shorter and / or wider fiber fragments) below 20. In one evaluation, composite materials with the highest mechanical reinforcement, containing fibers 40-100 μm in length, had fiber aspect ratios of about 60 to about 170.

[0242] References TIFF2026502902000004.tif90159TIFF2026502902000005.tif213160TIFF2026502902000006.tif155160

[0243] equivalent It will be understood that the detailed examples and embodiments described herein are given by way of example only for illustrative purposes and are not to be construed as limiting the invention in any manner. Various modifications or variations therein will occur to those skilled in the art and are therefore considered to be within the spirit and scope of this application and within the scope of the appended claims. For example, the relative amounts of components may be varied to optimize a desired effect, additional components may be added, and / or one or more of the described components may be substituted with similar components. Additional advantageous features and functionality associated with the systems, methods, and processes will be apparent from the appended claims. Moreover, those skilled in the art will recognize, or be able to ascertain using no more than routine experimentation, many equivalents to the specific embodiments of the invention described herein. Such equivalents are intended to be encompassed by the following claims.

Claims

1. 1. A fiber-hydrogel composition for administration to a subject, the composite comprising: Fibers comprising one or more collagen materials; and one or more hydrogel materials Including, the fiber and one or more hydrogel materials are covalently bonded; The fiber-hydrogel composition.

2. 10. The composition of claim 1, wherein the fibers comprise one or more selected from collagen, gelatin, cellulose, modified cellulose, cellulose acetate, HPMC, ethyl cellulose, silk, chitosan, keratin, elastin, elastin-like polypeptides, tropoelastin, and hyaluronic acid.

3. 3. The composition of claim 1 or 2, wherein the fibers are nanofibers or microfibers.

4. 4. The composition of claim 1, wherein the fibers comprise one or more of bovine type I collagen or gelatin and derivatives thereof.

5. The composition of any one of claims 1 to 4, wherein the fibers comprise one or more recombinant collagen materials.

6. The composition of any one of claims 1 to 5, wherein the fibers comprise one or more recombinant human collagen materials.

7. The composition of any one of claims 1 to 6, wherein the collagen fibers comprise type I bovine collagen nanofibers or fragments thereof.

8. The composition of any one of claims 1 to 7, wherein the collagen fibers are electrospun or centrifugally spun.

9. 9. The composition of any one of claims 1-8, wherein the one or more hydrogel materials comprise one or more of hyaluronic acid (HA), collagen, chitosan, alginate, polyvinyl acetate (PVA), gelatin, polyethylene glycol (PEG) and other glycol ethers, chondroitin sulfate, or a cellulosic material.

10. The composition of any one of claims 1 to 8, wherein the one or more hydrogel materials comprise one or more hyaluronic acid (HA) materials.

11. 11. The composition of any one of claims 1-10, wherein the one or more hydrogel materials are covalently bonded to a type I bovine collagen nanofiber sheet or fragment thereof.

12. The composition of any one of claims 1 to 11, wherein the collagen nanofibers are retained within the fiber-hydrogel composite.

13. 13. The composition of any one of claims 1 to 12, further comprising a crosslinker or a reacted form of a crosslinker.

14. 14. The composition of claim 13, wherein the crosslinking agent creates an interfacial bond between the collagen nanofibers and the HA chains.

15. 15. The composition of claim 13 or 14, wherein the crosslinking agent comprises one or more selected from a difunctional epoxide-based crosslinker, 1,4-butanediol diglycidyl ether (BDDE), divinyl sulfone (DVS), a PEG molecule with a terminal epoxide functional group, or an HA reactive agent.

16. 15. The composition of claim 13 or 14, wherein the cross-linking agent comprises DVS or BDDE.

17. 17. The composition of claim 16, wherein the concentration of the DVS ranges from about 0.5 v / v% to about 2.5 v / v% based on the total volume of the fiber-hydrogel composite.

18. 18. The composition of any one of claims 1 to 17, wherein the concentration of the HA ranges from about 0.5 w / v% to about 2 w / v%, based on the total volume of the fiber-hydrogel composite.

19. 19. The composition of any one of claims 1 to 18, wherein the fiber loading density of the collagen nanofibers ranges from about 1 w / v% to about 3 w / v%, based on the total volume of the fiber-hydrogel composite.

20. 20. The composition of any one of claims 1 to 19, wherein the concentration of the fiber ranges from about 0.1 w / v% to about 20 w / v%, the concentration of the hydrogel material ranges from about 0.5 w / v% to about 10 w / v%, and the concentration of the crosslinker ranges from about 0.05 v / v% to about 5 v / v%, based on the total volume of the composition.

21. 21. The composition of claim 20, wherein the concentration of the fiber ranges from about 1.5 w / v% to about 3.0 w / v%, the concentration of the hydrogel material ranges from about 0.8 w / v% to about 2 w / v%, and the concentration of the crosslinker ranges from about 0.5 v / v% to about 2.5 v / v%, based on the total volume of the composition.

22. 22. The composition of any one of claims 1 to 21, wherein the pH of the composition ranges from about 7.0 to about 7.4 in an isotonic solution.

23. 23. The composition of any one of claims 1 to 22, formulated in the form of a sheet or an injectable fluid.

24. 24. The composition of any one of claims 1 to 23, comprising one or more porous structures.

25. 25. The composition of any one of claims 1 to 24, wherein the storage modulus of the composition is in the range of about 100 to about 2 kPa.

26. 26. The composition of claim 25, wherein the storage modulus of the composition is in the range of about 100 to about 500 Pa.

27. 27. The composition of any one of claims 1 to 26, formed into microbeads having an average diameter in the range of about 50 μm to about 1000 μm.

28. 28. A soft tissue device comprising the composition of any one of claims 1 to 27.

29. An implant for promoting angiogenesis, comprising the composition of any one of claims 1 to 27.

30. An implant for adipose tissue formation, comprising the composition of any one of claims 1 to 27.

31. An implant for vasculogenesis, comprising the composition of any one of claims 1 to 27.

32. A composition according to any one of claims 1 to 27; Applicator and Kit including:

33. 33. The kit of claim 32, wherein the applicator is a syringe equipped with a needle or cannula.

34. a vial containing one or more of a salt, a buffer, and a therapeutic agent; 34. The kit of claim 32 or 33, further comprising:

35. 1. A method for producing a fiber-hydrogel composite, comprising: contacting a crosslinker with fibers comprising one or more collagen materials and hyaluronic acid (HA) to obtain a fiber-hydrogel composite; Including, the fibers are nanofibers or microfibers, the HA is bonded to the fibers by the crosslinker to form a composite network; The method.

36. 36. The method of claim 35, further comprising adjusting cross-linking conditions.

37. 36. The method of claim 35, wherein the cross-linking conditions are basic conditions.

38. 38. The method of claim 37, wherein the pH of the cross-linking conditions ranges from about 10 to about 14.

39. 39. The method of any one of claims 35 to 38, wherein said contacting step is carried out for about 30 minutes to about 4 hours.

40. 40. The method of any one of claims 35-39, wherein said contacting step is carried out at a temperature of about 20 to 90 degrees Celsius.

41. 41. The method of claim 40, wherein the contacting step is carried out at a temperature of about 37 degrees Celsius.

42. 42. The method of any one of claims 35 to 41, further comprising processing the fiber-hydrogel composite to specialize it into microbeads.

43. 43. The method of claim 42, wherein the microbeads have an average diameter ranging from about 50 μm to about 1,000 μm.

44. 44. The method of any one of claims 35 to 43, wherein said processing step comprises mechanical grinding or mechanical screening by applying shear using a mesh.

45. The method of any one of claims 35 to 44, further comprising sterilizing the fiber-hydrogel composite.

46. 46. ​​The method of any one of claims 35 to 45, wherein the fibers comprise one or more of collagen, chitosan, alginate, gelatin, and derivatives.

47. 47. The method of claim 46, wherein the fibers comprise collagen nanofibers.

48. 48. The method of claim 47, wherein the collagen nanofibers comprise type I bovine collagen nanofibers or fragments thereof or human collagen material.

49. The method of any one of claims 35 to 48, wherein the collagen nanofibers are electrospun and / or cryomilled.

50. 42. The method of claim 41, wherein the HA is covalently bound to the collagen nanofibers.

51. The method of any one of claims 35 to 50, wherein the crosslinking agent creates an interfacial bond between the collagen nanofibers and the HA.

52. 52. The method of any one of claims 35-51, wherein the crosslinking agent comprises one or more selected from a difunctional epoxide-based crosslinker, 1,4-butanediol diglycidyl ether (BDDE), divinyl sulfone (DVS), a PEG molecule with a terminal epoxide functional group, and an HA reactive agent.

53. 53. The method of any one of claims 35 to 52, wherein the cross-linking agent comprises DVS.

54. 54. The method of claim 53, wherein the concentration of the DVS ranges from about 0.5 v / v% to about 2.5 v / v% based on the total volume of the fiber-hydrogel composite.

55. The method of any one of claims 35 to 54, wherein the concentration of the HA ranges from about 0.5 w / v% to about 2 w / v%, based on the total volume of the fiber-hydrogel composite.

56. The method of any one of claims 35 to 55, wherein the fiber loading density of the collagen nanofibers ranges from about 1 w / v% to about 3 w / v%, based on the total volume of the fiber-hydrogel composite.

57. 57. The method of any one of claims 35 to 56, wherein the concentration of the collagen fiber ranges from about 0.1 w / v% to about 20 w / v%, the concentration of the HA ranges from about 0.5 w / v% to about 10 w / v%, and the concentration of the crosslinker ranges from about 0.05 v / v% to about 5 v / v%, based on the total volume of the fiber-hydrogel composite.

58. 57. The method of any one of claims 35 to 56, wherein the concentration of the collagen fibers ranges from about 1.5 w / v% to about 3.0 w / v%, the concentration of the HA ranges from about 0.8 w / v% to about 2 w / v%, and the concentration of the crosslinker ranges from about 0.5 v / v% to about 2.5 v / v%, based on the total volume of the fiber-hydrogel composite.

59. 59. The method of any one of claims 35-58, wherein the pH of the fiber-hydrogel composite ranges from about 7.0 to about 7.4 in an isotonic solution.

60. 60. The method of any one of claims 35-59, further comprising fabricating the fiber-hydrogel composite into a sheet form or an injectable bead form.

61. The method of any one of claims 35 to 60, wherein the fiber-hydrogel composite comprises one or more porous structures.

62. 62. The method of any one of claims 35-61, wherein the fiber-hydrogel composite has a storage modulus in the range of about 100 to about 500 Pa.

63. 28. A method for forming adipose tissue in a subject, the method comprising administering to the subject a composition according to any one of claims 1 to 27.

64. 64. The method of claim 63, wherein the composition induces prolonged in vivo retention or enhanced host cell infiltration, optionally in the absence of incorporation of cells and growth factors.

65. 65. The method of claim 64, wherein the subject exhibits neovasculature formation driven by M2 macrophage polarization.

66. A method of delivering cells or tissue in a subject, encapsulating one or more cells or tissues in a composition according to any one of claims 1 to 27 to form a mixture; and applying the mixture to a target site in the subject. The method comprising:

67. 67. The method of claim 66, wherein the tissue is processed adipose tissue.

68. 67. The method of claim 66, wherein the cells are one or more types of cells selected from the group consisting of adipose-derived stem cells, adipocytes, or a combination thereof.

69. 1. A method of delivering a pharmaceutical agent in a subject, comprising: combining a pharmaceutical agent and a composition according to any one of claims 1 to 27 to form a mixture; and applying the mixture to the intended delivery site. The method comprising:

70. An implant device of the fiber-hydrogel composite of any one of claims 1 to 27.

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