Polymeric biomaterials for tissue regeneration and methods of use thereof

Abstract

A biomaterial comprising a polymer and an elastic recombinant protein, wherein the polymer comprises at least one cross-linkable functional group, and a method of preparing said biomaterial. In certain embodiments, the hydrogel is gelatin methacryloyl (GelMA) and the elastic recombinant protein is elastin-like polypeptide (ELP).

Description

POLYMERIC BIOMATERIALS FOR TISSUE REGENERATION AND METHODS OF USE THEREOFCROSS-REFERENCE TO RELATED APPLICATIONS

[0001] This application claims priority to and the benefit of U.S. Provisional Application No. 63 / 584,343, filed September 21, 2023, the disclosure of which is incorporated herein by reference herewith in its entirety.STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT

[0002] This invention was made with government support under grant number DK134868 awarded by the National Institutes of Health-National Institute of Diabetes and Digestive and Kidney Diseases (NIH-NIDDK) and grant number HL140618 awarded by the National Institutes of Health (NIH). The government has certain rights in the invention.SUMMARY OF THE INVENTION

[0003] The present invention provides, in part, a biomaterial comprising a polymer and an elastic recombinant protein, wherein the polymer comprises at least one cross-linkable functional group.

[0004] In one embodiment, the polymer is a hydrogel selected from the group consisting of gelatin, collagen, chitosan, polyethyleneimine, hyaluronic acid, bovine serum albumin (BSA), dextran, derivatives thereof, and combinations thereof. In one embodiment, the polymer is a synthetic polymer selected from the group consisting of polycaprolactone (PCL), polylactic acid (PLA), polyglycolide (PGA), poly(lactic-co-glycolic) acid (PLGA), polypropylene oxide (PPO), polyethylene glycol (PEG), poly(amino acid), derivatives thereof, copolymers thereof, and combinations thereof. In one embodiment, the polymer is functionalized with a substituent selected from the group consisting of acrylate, methacrylate, alkene, carbonyl, and combinations thereof.

[0005] In one embodiment, the cross-linkable functional group is selected from the group consisting of acrylate, methacrylate, amino, mercapto, epoxy, hydroxy, isocyanate, amide, carboxyl, pyrrolidone, siloxane, oxetane, nitrile, ethylene ether, derivatives thereof, and combinations thereof.

[0006] In one embodiment, the polymer is gelatin methacryloyl (GelMA). In one embodiment, the elastic recombinant protein is elastin-like polypeptide (ELP). In one embodiment, the ratio of the polymer to the elastic recombinant protein is at least 1 : 1. In one embodiment, the biomaterial further comprises a therapeutic agent.

[0007] In one embodiment, the biomaterial is biodegradable. In one embodiment, the biomaterial is a biomimetic of lower urinary tract tissue. In one embodiment, the biomaterial has an ultimate tensile strength of at least 15 kPa. In one embodiment, the biomaterial has a linear percent extensibility of at least 100%. In one embodiment, the biomaterial has an ultimate tensile strength of at least 15 kPa and a linear percent extensibility of at least 100%.

[0008] In one embodiment, the biomaterial has a suture retention strength of at least 8 kPa.

[0009] The present invention further provides, in part, a method of preparing a biomaterial comprising the steps of: providing a solution comprising a polymer, wherein the polymer comprises at least one cross-linkable functional group; extruding the solution to provide a scaffold; exposing the scaffold to an initiator; and cross-linking the scaffold to provide the biomaterial.

[0010] In one embodiment, the solution further comprises an elastic recombinant protein. In one embodiment, the solution comprises at least 3% (w / v) of the polymer. In one embodiment, the solution comprises at least 1% (w / v) of the elastic recombinant protein. In one embodiment, the solution further comprises a therapeutic agent.

[0011] In one embodiment, the step of extruding the solution is performed using electrospinning. In one embodiment, the step of exposing the scaffold to an initiator further comprises the step of submerging the scaffold in a second solution comprising a therapeutic agent.

[0012] In one embodiment, the initiator is a photoinitiator. In one embodiment, the step of crosslinking the scaffold comprises the step of irradiating the scaffold with light having a wavelength range of 100 to 400 nm. In one embodiment, the step of crosslinking the scaffold initiates the formation of a carbon-carbon bond or a sulfur-sulfur bond.BACKGROUND OF THE INVENTION

[0013] Current lower urinary tract (LUT) reconstructive surgeries such as neobladder creation, augmentation cystoplasty, or proximal hypospadias repair are high-risk, complex procedures. For example, augmentation cystoplasties have reported 10-year median costs of $51 ,798 ($51 ,798-5327,483 95% CI) and 30-day complication rates of 33% (Du, K. et al., 2015, J. Pediatr. Surg., 50, 1535; Kim, S. J. et al., 2022, J. Pediatr. Urol. 18, 314) Similarly, proximal hypospadias repairs are widely variable, with more than 300 described repair options (Ceccarelli, P. L. et al., 2021, BMC Pediatr., 21, 523). However, long-term complications and risk of redo surgery are up to 49-68% in recent series (Long, C. J. et al., 2017, J. Urol., 197, 852; Stanasel, I. et al., 2015, J. Urol., 194, 512; McNamara, E. R. et al., 2015, J. UroL, 194, 1080). One limitation of these repairs is that local tissue options can be either effected by the primary diagnosis or may be inadequate to complete the reconstruction. In this setting, the patient may require a secondary tissue source to complete the repair. Such grafts can be limited in availability, have undesirable tissue properties (e.g., mucous production by bowel), and risk donor site morbidity (Choi, J. J. E. et al., 2020, Clin. Exp. Dent. Res., 6, 602; Rahmati, S. et al., 2022, Cell, 24, 506; Cunnane, E. M. et al., 2021, Biomaterials, 269, 120651). Thus, there is an ongoing need for an improved tissue alternatives with optimized biologic and mechanical properties to support healthy LUT regeneration and enable early return of function (Versteedgen L. R. et al., 2017, Acta Biomat., 52, 1).

[0014] Lower urinary tract (LUT) tissues (e.g., urinary bladder, urethra) have remarkable elasticity, a key finding that supports long-term urine storage, voiding, and penile tumescence. This property fails to be reproduced by current tissue alternatives applied in LUT surgical reconstruction, such as augmentation cystoplasty or urethroplasty procedures. This tensile mismatch of vastly different substrates can minimize early functionality and increase regionalwound tension, thereby increasing the risk of perioperative complications. Additionally, the mechanical niche is an important parameter effecting wound healing by controlling cell fate and extracellular matrix (ECM) deposition.

[0015] An essential aspect of bladder and urethral tissue that enables lifelong cyclic voiding and tumescence is its remarkable elasticity. Specifically, initial work demonstrated that the urethra can extend to up to 300% or more of its original length. However, current tissue and graft sources (e.g., bowel, foreskin, decellularized matrices) failed to recapitulate this key aspect of LUT tissue mechanics. An example of this is one-ply small intestinal submucosa (SIS) which in both prior studies and recent work has been found to have a UTS that is more than 10-fold greater than healthy LUT tissue (Shi, L. et al., 2013, Int. J. Bums Trauma, 3, 173). On a macroscopic surgical scale, this mechanical mismatch resulted in increased risk of failure at and near suture lines, increasing the risk of urine leaks, infection, perforation, and dehiscence. To prevent these complications, prolonged catheterization is used as a diversion mechanism. However, catheter presence decreases early mobilization and return to baseline activities, increases risk of unplanned healthcare visits, and is a source of local and systemic infection (Igawa, Y. et al., 2008, Int. J. UroL, 15, 481). Put together, on a macroscopic scale, mismatched tissue elasticity is a major contributor to a wide range of perioperative complications and decreases early return to voiding for individuals requiring tissue alternative sources in LUT reconstructive procedures.

[0016] On the microscopic scale, tissue viscoelasticity and its interaction with dynamic and static mechanical forces in the substrate likewise plays a significant role in wound healing. The mechanical niche is an important parameter controlling cell fate (Park, J. S. et al., 2011, Biomaterials, 32, 3921). Cell ingrowth, morphology, and phenotype are effected by scaffold topography, as illustrated by a recent urologic study in which a low-density polarized collagen graft with matched tensile properties to native tissue had improved smooth muscle regeneration and patency as compared to a high-density graft following urethroplasty in an animal model (Larsson, H. M. et al., 2018, Scientific Reports, 8, 10057). Furthermore, wound healing and immune response to materials, including macrophage polarization, can be differentially directed by changes in the urinary bladder matrix including changes in structure, mechanics, and biology that occur in disease states (Sadtler, K. et al., 2017, Semin. Immunol., 29, 14). Although moreclosely studied in its role in tumor biology, where increased tissue stiffness was commonly observed and associated with tumor progression (Ahmadian, N. et al., 2023, J. Biomater. Appl., 38, 179) these factors have been less fully characterized and analyzed in terms of their effects on healthy LUT tissue regeneration. Emerging techniques such as 3-dimensional (3D) hydrogel cultures and microfluidic platforms have demonstrated the importance of microenvironment mechanics, surface topography, porosity, and fibrous architecture as interactive with cell biology beyond those characteristics observed in 2-dimensional (2D) cell cultures (Carvalho, E. M. et al., 2023, Acta Biomater., 163, 146).

[0017] Historically, LUT tissue engineering has been successful in healthy small animal models; however, in larger animal and human models synthetic, biologic, and hybrid materials have had far less success (Chan, Y. Y. et al., 2020, Nature Reviews Urology, 17, 162). Key challenges that remain include the need for materials that achieve the unique mechanics and structure of the target tissue. Optimization of wound healing and vascularization in larger animal models with translational urethral length and graft size is also required to ensure long-term favorable clinical outcomes. Additionally, reproducible synthesis methods (e.g., as compared to biologies such as decellularized tissue that vary by animal and site) are needed to ensure a consistent product for surgical use.

[0018] Urinary tract reconstructive procedures have a high failure rate, and patients undergoing these procedures may experience complications such as urinary tract infections, urine leak, fistulae, and urethral strictures. Current tissue repair options fail to recapitulate the biology, mechanics, and structure of healthy native tissue, thereby increasing this risk. Furthermore, these tissues are taken from donor sites with limited graft availability and potential for donor site morbidity. There is a need for a biomaterial that enhances regeneration, restore early function, and is designed for tissue handling in a manner that replicates native tissue.

[0019] There is currently no FDA-approved biologic to replace urethral or bladder tissues. However, current FDA-approved commercial alternatives, such as Cook Biodesign small intestinal submucosa, have a mean tensile strength of 5 MPa, approximately 10-fold higher than the lower urinary tract tissue. As the mechanical niche is an important parameter controlling cell fate and extracellular matrix (ECM) deposition, having tensile matched scaffolds for cell seedingand tissue regeneration are anticipated to decrease localized fibrosis and minimize risk of tension and decrease risk of early return of function when applied in urinary tract reconstructive procedures.

[0020] Thus, there is a need in the art for novel materials for use as biocompatible and tunable alternatives for LUT replacement tissues. The present invention addresses this need.BRIEF DESCRIPTION OF THE DRAWINGSThe foregoing purposes and features, as well as other purposes and features, will become apparent with reference to the description and accompanying figures below, which are included to provide an understanding of the invention and constitute a part of the specification, in which like numerals represent like elements, and in which:

[0021] Figure 1 depicts a graphical abstract illustrating the study format of the present invention, including tensile tissue characterization, scaffold synthesis, reproduction of highly elastic tissue properties with developed biomaterials, and in vitro cell seeding and evaluation of urothelial, smooth muscle, and fibroblast cytocompatibility with synthesized biologic scaffolds.

[0022] Figure 2, comprising Figure 2A through Figure 2E, depicts representative mechanical properties of small intestinal submucosa (SIS) versus lower urinary tract tissue. Figure 2A depicts the tensile modulus of Cook 1 -ply Biodesign (SIS) versus male rabbit bladder and anterior urethral tissue (N=7 per cohort). Figure 2B depicts the ultimate strength of Cook 1- ply Biodesign small intestinal submucosa (SIS) versus male rabbit bladder and anterior urethral tissue (N=7 per cohort). Figure 2C depicts the extensibility of Cook 1 -ply Biodesign (SIS) versus male rabbit bladder and anterior urethral tissue (N=7 per cohort). *p < 0.05, **p < 0.01, ***p < 0.001, ****p < 0.0001. Figure 2D depicts representative histology (hematoxylin and eosin staining (H&E)) adjacent to tensile evaluated segments of bladder tissue. Figure 2D depicts representative histology (H&E) adjacent to tensile evaluated segments of urethral tissue.

[0023] Figure 3, comprising Figure 3A through Figure 3G, depicts representative mechanical properties of photocrosslinkable nanofibrous scaffolds compared to adult male rabbiturethra and bladder tissue. Figure 3 A depicts the tensile modulus of tissue (n=7) versus varied ratios of elastin-like polypeptide (ELP) / gelatin methacrolyl (GelMA) (ELP / GelMA) scaffolds (n=5). Figure 3B depicts the ultimate strength of tissue (n=7) versus varied ratios of ELP / GelMA scaffolds (n=5). Figure 3C depicts the percent extensibility of tissue (n=7) versus varied ratios of ELP / GelMA scaffolds (n=5). Figure 3D depicts a representative 8thcycle tensile curve of urethral tissue (n=8) versus 5E5G candidate scaffold (n=4). Figure 3E depicts a representative energy loss of urethral tissue (n=8) versus 5E5G candidate scaffold (n=4). Figure 3F depicts an illustrative suture line failure during stretching of sutured scaffold to tissue. Figure 3G depicts the ultimate suture retention strength of scaffold to scaffold and scaffold to tissue configurations (n=3-4). *p < 0.05, **p < 0.01.

[0024] Figure 4, comprising Figure 4A through Figure 4H, depicts a representative schematic of gelatin methacryloyl (GelMA) / elastin-like peptide (ELP) (GelMA / ELP) fibrous scaffold formation and chemical structure. Figure 4A depicts gelatin methacrylation to form GelMA. Figure 4B depicts the chemical structure of ELP, indicating the presence of thiol groups. Figure 4C depicts ELP / GelMA hydrogel formation schematic diagrams including crosslinking mechanisms. Figure 4D depicts nuclear magnetic resonance (NMR) characterization, including methacrylization and crosslinking. Figure 4E depicts a representative scanning electron microscopy (SEM) image of an electrospun 0E10G scaffold. Figure 4F depicts a representative SEM image of an electrospun 5E5G scaffold. Figure 4G depicts fiber diameter measurements from SEM images. Figure 4H depicts porosity measurements from SEM images, ns = not significant, **p < 0.01.

[0025] Figure 5, comprising Figure 5A through Figure 5F, depicts representative degradation and swelling behavior of varied material concentrations. Figure 5A depicts a representative degradation profile for a 0E10G scaffold with and without collagenase and elastase. Figure 5B depicts a representative degradation profile for a 5E5G scaffold with and without collagenase and elastase. Figure 5C depicts the representative swelling ratio of a 0E10G scaffold at 4, 20, and 37 °C. Figure 5D depicts the representative swelling ratio of a 5E5G scaffold at 4, 20, and 37 °C. Figure 5E depicts a comparison of the degradation profile of varied compositions of GelMA / ELP scaffolds (0E10G, 3E7G, 5E5G, and 7E3G) at selected time points in phosphate buffered saline (PBS). Figure 5F depicts a comparison of the degradation profile ofvaried compositions of GelMA / ELP scaffolds (0E10G, 3E7G, 5E5G, and 7E3G) at selected time points in combined collagenase and elastase. *p < 0.05, **p < 0.01, ***p < 0.001, ****p < 0.0001.

[0026] Figure 6, comprising Figure 6A through Figure 6M, depicts representative in vitro cytocompatibility of human lower urinary tract smooth muscles on nanofibrous scaffolds, day 1 to 7 endpoints. Figure 6A through Figure 6D depict representative live / dead immunofluorescent stain (Scale bar = 800 pm). Figure 6E through Figure 6H depict representative Actin / Dapi immunofluorescent stains. Figure 61 depicts quantitative cell viability Presto Blue analysis measured in relative fluorescence units (RFU) (N=4 completed in quadruplicate). Figure 6J depicts quantitative cell proliferation (N=4 completed in quadruplicate). Figure 6K depicts quantitative cell distribution within the scaffold (n=9 tiles per condition). Figure 6L depicts representative day 7 Actin / Dapi confocal isometric projection of cell scaffold demonstrating depth penetration and cell proliferation for 0E10G scaffold. (lOx, Scale bar = 200 pm). Figure 6M depicts representative day 7 Actin / Dapi confocal isometric projection of cell scaffold demonstrating depth penetration and cell proliferation for 5E5G scaffold. (lOx, Scale bar = 200 pm). *p < 0.05, **p < 0.01, *** p < 0.001, ****p < 0.0001; mean ± SEM for all bar graphs.

[0027] Figure 7, comprising Figure 7A through Figure 7M, depicts representative in vitro cytocompatibility of human lower urinary tract urothelial cells on nanofibrous scaffolds, day 1 to 7 endpoints. Figure 7A through Figure 7D depict representative live / dead immunofluorescent stain (Scale bar = 800 pm). Figure 7E through Figure 7H depict representative Actin / Dapi immunofluorescent stains. Figure 71 depicts quantitative cell viability Presto Blue analysis measured in relative fluorescence units (RFU) (N=4 completed in quadruplicate). Figure 7J depicts quantitative cell proliferation (N=4 completed in quadruplicate). Figure 7K depicts quantitative cell distribution within the scaffold (n=9 tiles per condition). Figure 7L depicts representative day 7 Actin / Dapi confocal isometric projection of cell scaffold demonstrating depth penetration and cell proliferation for 0E10G scaffold. (lOx, Scale bar = 200 pm). Figure 7M depicts representative day 7 Actin / Dapi confocal isometric projection of cell scaffold demonstrating depth penetration and cell proliferation for 5E5G scaffold. (lOx, Scale bar = 200 pm). *p < 0.05, **p < 0.01, *** p < 0.001, ****p < 0.0001; mean ± SEM for all bar graphs.

[0028] Figure 8, comprising Figure 8A through Figure 8M, depicts representative in vitro cytocompatibility of fibroblast cells on nanofibrous scaffolds, day 1 to 7 endpoints. Figure 8A through Figure 8D depict representative live / dead immunofluorescent stain (Scale bar = 800 pm). Figure 8E through Figure 8H depict representative Actin / Dapi immunofluorescent stains. Figure 81 depicts quantitative cell viability Presto Blue analysis measured in relative fluorescence units (RFU) (N=4 completed in quadruplicate). Figure 8J depicts quantitative cell proliferation (N=4 completed in quadruplicate). Figure 8K depicts quantitative cell distribution within the scaffold (n=9 tiles per condition). Figure 8L depicts representative day 7 Actin / Dapi confocal isometric projection of cell scaffold demonstrating depth penetration and cell proliferation for 0E10G scaffold. (lOx, Scale bar = 200 pm). Figure 8M depicts representative day 7 Actin / Dapi confocal isometric projection of cell scaffold demonstrating depth penetration and cell proliferation for 5E5G scaffold. (lOx, Scale bar = 200 pm). *p < 0.05, **p < 0.01, *** p < 0.001, ****p < 0.0001; mean ± SEM for all bar graphs.

[0029] Figure 9 depicts in vivo subcutaneous implantation of 0E10G, 5E5G, and SIS cohorts at day 1 to 2-month surgical endpoints.

[0030] Figure 10, comprising Figure 10A through Figure 10E, depicts n vivo subcutaneous implantation, immune marker analysis. Figure 10A depicts H&E staining of tissues with implant. Figure 10B depicts immunofluorescence of combined CD68, MPO, and CD3 markers of tissues adjacent to implant. Figure 10C through Figure 10E depicts quantified immunofluorescent cells in comparison to sham images at each surgical endpoint for CD68, MPO, and CD3, respectively.DETAILED DESCRIPTION

[0031] The present invention relates in part to the unexpected tensile behavior and biocompatibility of the highly elastic, biomimetic, biodegradable, suturable materials described herein. In certain embodiments, the materials are synthesized through the photocrosslinking of two polymers, gelatin methacrolyl (GelMA) and recombinant elastin-like polypeptide (ELP). In certain embodiments, the resulting biopolymer possesses tunable mechanical and structural properties. The resultant materials can be designed to possess the range of tensile properties of native healthy tissue. In certain embodiments, the material is non-cytotoxic to lower urinary tractand / or genital-derived human cell lines. In certain embodiments, the material can serve as an alternative to current tissue repair options, with use in clinical translation for reconstructive surgery in the urinary tract.

[0032] In certain embodiments, the material comprises biomaterials which can be naturally derived, with modifications that enable tunability through photocrosslinking, or modification of the relative ratios of the two components, such as ELP and GelMA. Varied nanotopography can be manufactured using electrospinning. The present invention can be used in combination with biomaterials in urinary tract applications, with the potential to apply to other internal elastic, luminal organs that require cyclic functionality (e.g., cardiovascular, pulmonary, gastrointestinal) in addition to the lower genitourinary tract (e.g., urethra, bladder, pelvic floor, vagina). These materials may be applied in a cell-seeded (bioengineered containing cells prior to implantation) versus unseeded fashion and have been demonstrated to have cytocompatibility. Additionally, ELP can be modified such as through methacrylation to improve crosslinking within the hydrogel.Definitions

[0033] It is to be understood that the figures and descriptions of the present invention have been simplified to illustrate elements that are relevant for a clear understanding of the present invention, while eliminating, for the purpose of clarity, many other elements found in related systems and methods. Those of ordinary skill in the art may recognize that other elements and / or steps are desirable and / or required in implementing the present invention. However, because such elements and steps are well known in the art, and because they do not facilitate a better understanding of the present invention, a discussion of such elements and steps is not provided herein. The disclosure herein is directed to all such variations and modifications to such elements and methods known to those skilled in the art.

[0034] Unless defined otherwise, all technical and scientific terms used herein have the same meaning as commonly understood by one of ordinary skill in the art to which this invention belongs. Although any methods and materials similar or equivalent to those described herein can be used in the practice or testing of the present invention, exemplary methods and materials are described.

[0035] As used herein, each of the following terms has the meaning associated with it in this section.

[0036] The articles “a” and “an” are used herein to 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.

[0037] “About” as used herein when referring to a measurable value such as an amount, a temporal duration, and the like, is meant to encompass variations of ±20%, ±10%, ±5%, ±1%, and ±0.1% from the specified value, as such variations are appropriate.

[0038] Throughout this disclosure, various aspects of the invention can be presented in a range format. It should be understood that the description in range format is merely for convenience and brevity and should not be construed as an inflexible limitation on the scope of the invention. Accordingly, the description of a range should be considered to have specifically disclosed all the possible subranges as well as individual numerical values within that range. For example, description of a range such as from 1 to 6 should be considered to have specifically disclosed subranges such as from 1 to 3, from 1 to 4, from 1 to 5, from 2 to 4, from 2 to 6, from 3 to 6 etc., as well as individual numbers within that range, for example, 1, 2, 2.7, 3, 4, 5, 5.3, 6 and any whole and partial increments therebetween. This applies regardless of the breadth of the range.

[0039] As used herein, the term “abnormal,” when used in the context of organisms, tissues, cells or components thereof, refers to those organisms, tissues, cells or components thereof that differ in at least one observable or detectable characteristic (e.g., age, treatment, or time of day) from those organisms, tissues, cells or components thereof that display the “normal” (expected) respective characteristic. Characteristics that are normal or expected for one cell or tissue type might be abnormal for a different cell or tissue type.

[0040] As used herein, a “disease” is a state of health of an animal wherein the animal cannot maintain homeostasis, and wherein if the disease is not ameliorated then the animal’s health continues to deteriorate.

[0041] As used herein, a “disorder” in an animal is a state of health in which the animal is able to maintain homeostasis, but in which the animal’s state of health is less favorable than it would be in the absence of the disorder. Left untreated, a disorder does not necessarily cause a further decrease in the animal’s state of health.

[0042] As used herein, a disease or disorder is “alleviated” if the severity of a symptom of the disease or disorder, the frequency with which such a symptom is experienced by a patient, or both, is reduced.

[0043] The terms “treatment”, “treating” and the like are used herein to generally mean obtaining a desired pharmacological and / or physiological effect. The effect may be prophylactic in terms of completely or partially preventing a disease or symptom thereof and / or may be therapeutic in terms of partially or completely curing a disease and / or adverse effect attributed to the disease. The term “treatment” as used herein covers any treatment of a disease in a subject and includes: (a) preventing a disease related to an undesired immune response from occurring in a subject which may be predisposed to the disease; (b) inhibiting the disease, i.e., arresting its development: or (c) relieving the disease, i.e., causing regression of the disease.

[0044] The term “derivative” refers to a small molecule that differs in structure from the reference molecule, but may retain or enhance the essential properties of the reference molecule and may have additional properties. A derivative may change its interaction with certain other molecules relative to the reference molecule. A derivative molecule may also include a salt, an adduct, tautomer, isomer, or other variant of the reference molecule.

[0045] As used herein, the terms “patient,” “subject,” “individual” and the like are used interchangeably, and refer to any animal, or cells thereof whether in vitro or in situ, amenable to the methods described herein. In a non-limiting embodiment, the patient, subject or individual is a human.

[0046] As used herein, the term “therapeutic” treatment refers to a treatment administered to a subject who exhibits signs of pathology, for the purpose of diminishing or eliminating those signs.

[0047] As used herein, the term “treating a disease or disorder” means reducing the frequency with which a symptom of the disease or disorder is experienced by a patient. Disease and disorder are used interchangeably herein.

[0048] As used herein, the term “therapeutically effective amount” refers to an amount that is sufficient or effective to prevent or treat (delay or prevent the onset of, prevent the progression of, inhibit, decrease or reverse) a disease or condition described or contemplated herein, including alleviating symptoms of such disease or condition.

[0049] The terms “effective amount” and “pharmaceutically effective amount” refer to a sufficient amount of an agent to provide the desired biological result. That result can be reduction and / or alleviation of a sign, symptom, or cause of a disease or disorder, or any other desired alteration of a biological system. An appropriate effective amount in any individual case may be determined by one of ordinary skill in the art using routine experimentation.

[0050] A “therapeutically effective amount” refers to that amount which provides a therapeutic effect for a given condition and administration regimen. In particular, “therapeutically effective amount” means an amount that is effective to prevent, alleviate or ameliorate symptoms of the disease or prolong the survival of the subject being treated, which may be a human or non-human animal. Determination of a therapeutically effective amount is within the skill of the person skilled in the art.

[0051] The terms “patient,” “subject,” “individual,” and the like are used interchangeably herein, and refer to any animal, or cells thereof whether in vitro or in situ, amenable to the methods described herein. In certain non-limiting embodiments, the patient, subject or individual is a human. In certain embodiments, a subject is a mammal such as a non-primate (e.g., cows, pigs, horses, cats, dogs, rats, etc.) and a primate (e.g., monkey and human), or a human.

[0052] As used herein, the term “pharmaceutical composition” refers to a mixture of at least one compound of the invention with other chemical components and entities, such as carriers, stabilizers, diluents, dispersing agents, suspending agents, thickening agents, and / or excipients. The pharmaceutical composition facilitates administration of the compound to an organism. Multiple techniques of administering a compound exist in the art including, but notlimited to, intravenous, oral, aerosol, parenteral, ophthalmic, pulmonary and topical administration.

[0053] As used herein, the term “hydrogel” or “aquagel” refers to a network of oligomers or polymer chains that are water-insoluble, sometimes found as a colloidal gel in which water is the dispersion medium.

[0054] The term “biodegradable” includes polymers, compositions and formulations, such as those described herein, that are intended to degrade during use. Biodegradable polymers typically differ from non-biodegradable polymers in that the former may be degraded during use. In one embodiment, such use involves in vivo use, such as in vivo therapy. In another embodiment, such use involves in vitro use. In general, biodegradation involves the degradation of a biodegradable polymer into its component subunits, or digestion, e g., by a biochemical process, of the polymer into smaller, non-polymeric subunits. Two types of biodegradation may generally be identified. For example, biodegradation may involve cleavage of bonds (whether covalent or otherwise) in the polymer backbone. In such biodegradation, monomers and oligomers typically result, and even more typically, such biodegradation occurs by cleavage of a bond connecting one or more of subunits of a polymer. Further, biodegradation may involve cleavage of a bond (whether covalent or otherwise) internal to side chain or that connects a side chain to the polymer backbone. For example, a therapeutic agent or other chemical moiety attached as a side chain to the polymer backbone may be released by biodegradation. In one embodiment, at least one type of biodegradation may occur during use of a polymer. As used herein, the term “biodegradation” encompasses all known types of biodegradation.

[0055] The term "biomimetic", as used herein, refers to substances, or processes that mimic any naturally occurring biological function or process of any naturally occurring biological compound.

[0056] As used herein, the terms “incorporated” and “encapsulated” refer to a therapeutic agent and a polymeric composition, such as a composition disclosed herein. In one embodiment, these terms include incorporating, formulating or otherwise including such agent into a composition which allows for fast or sustained release of such agent in the desired application. The terms may contemplate any manner by which a therapeutic agent or other material isincorporated into a polymer matrix, including for example: attached to a monomer of such polymer (by covalent or other binding interaction) and having such monomer be polymerized to give a polymeric formulation; distributed throughout the polymeric matrix; appended to the surface of the polymeric matrix (by covalent or other binding interactions); encapsulated inside the polymeric matrix, and the like. The term “co-incorporation” or “co-encapsulation” refers to the incorporation of a therapeutic agent or equivalent and at least one other therapeutic agent or equivalent in a subject composition.

[0057] As used herein, the terms “biocompatible polymer” and “biocompatibility” when used in relation to polymers are recognized in the art. For example, biocompatible polymers include polymers that are generally neither toxic to the host, nor degrade (if the polymer degrades) at a rate that produces monomeric or oligomeric subunits or other byproducts at toxic concentrations in the host. In one embodiment, biodegradation generally involves degradation of the polymer in a host, e.g., into its monomeric subunits, which may be known to be effectively non-toxic. Intermediate oligomeric products resulting from such degradation may have different toxicological properties, however, or biodegradation may involve oxidation or other biochemical reactions that generate molecules other than monomeric subunits of the polymer. Consequently, in one embodiment, toxicology of a biodegradable polymer intended for in vivo use, such as implantation or injection into a patient, may be determined after one or more toxicity analyses. It is not necessary that any subject composition have a purity of 100% to be deemed biocompatible; indeed, it is only necessary that the subject compositions be biocompatible as set forth above. Hence, a subject composition may comprise polymers comprising 99%, 98%, 97%, 96%, 95%, 90%, 85%, 80%, 75% or even less of biocompatible polymers, e.g., including polymers and other materials and excipients described herein, and still be biocompatible.

[0058] As used herein, the terms “peptide,” “polypeptide,” and “protein” are used interchangeably, and refer to a compound comprised of amino acid residues covalently linked by peptide bonds. A protein or peptide must contain at least two amino acids, and no limitation is placed on the maximum number of amino acids that can comprise a protein’s or peptide’s sequence. Polypeptides include any peptide or protein comprising two or more amino acids joined to each other by peptide bonds. As used herein, the term refers to both short chains, which also commonly are referred to in the art as peptides, oligopeptides and oligomers, for example,and to longer chains, which generally are referred to in the art as proteins, of which there are many types. “Polypeptides” include, for example, biologically active fragments, substantially homologous polypeptides, oligopeptides, homodimers, heterodimers, variants of polypeptides, modified polypeptides, derivatives, analogs, fusion proteins, among others. The polypeptides include natural peptides, recombinant peptides, synthetic peptides, or a combination thereof.

[0059] As used herein, the term “polymerization” or “cross-linking” refers to at least one reaction that consumes at least one functional group in a monomeric molecule (or monomer), oligomeric molecule (or oligomer) or polymeric molecule (or polymer), to create at least one chemical linkage between at least two distinct molecules (e.g., intermolecular bond), at least one chemical linkage within the same molecule (e.g., intramolecular bond), or any combination thereof. A polymerization or cross-linking reaction may consume between about 0% and about 100% of the at least one functional group available in the system. In one embodiment, polymerization or cross-linking of at least one functional group results in about 100% consumption of the at least one functional group. In another embodiment, polymerization or cross-linking of at least one functional group results in less than about 100% consumption of the at least one functional group.

[0060] As used herein, the term “ultimate tensile strength” refers to the value of stress applied to a material which just causes its rupture.

[0061] As used herein, the term “linear percent extensibility” indicates the percentage of which a material may be stretched or elongated in a single direction. In certain embodiments, the linear percent extensibility is calculated by dividing the total distance that the material is stretched by the original length of the material and multiplying it by 100. Throughout the disclosure, the term “extensibility” may be used interchangeably with “elongation”.

[0062] As used herein, the term “suturable” means such a state that a material can be fixed at a part of the material, for instance, two or more points thereof to an implanted part in a living body, for example, defective sites in the lower urinary tract tissue, such as the bladder or urethra, or the pericardium, cerebral dura mater, pleura, intestinal chorion and the like, or a resected part in the lungs, liver or the like, using a general suture thread for surgery. Further, italso means that the membrane can be maintained in a state that the membrane does not easily deviate, drop, or move from the suture-fixed position or deform for a certain period of time.

[0063] As used herein, the term “suture retention strength” refers to the value of the maximum strength or force required to pull a suture out of a material or prothesis or cause the material or prothesis to fail. The term can be used interchangeably with “anastomotic strength”.

[0064] The material described herein addresses key limitations in current repairs and historic tissue engineered strategies. Specifically, a novel platform was created for further refined mechanical and structural property optimization, with a goal of improved wound healing and vascularization with reproducible and refinable materials. By applying systematic evaluation and selection criteria, the scaffold developed allows for modifications by layer specific to each cell line that address structural and functional requirements, creating micro-niche structures that support healthy LUT tissue regeneration (Figure 1).Biomaterials

[0065] The present invention relates to, in part, a biomaterial comprising a polymer and an elastic recombinant protein, wherein the polymer comprises at least one cross-linkable functional group.

[0066] In certain embodiments, the polymer is selected from any FDA-approved hydrogel, including, but not limited to, hydrogels selected from the group consisting of gelatin, collagen, chitosan, polyethyleneimine, hyaluronic acid, bovine serum albumin (BSA), dextran, derivatives thereof, and combinations thereof. In one embodiment, the polymer is derived from collagen. In one embodiment, the polymer is gelatin. In one embodiment, the polymer is derived from fish, fish skin, porcine tendons, and / or bovine tendons.

[0067] The term "gelatin" as used herein refers to a proteinaceous material derived from collagen. In the context of the present invention, "gelatin" also refers to equivalent substances, such as synthetic analogues of gelatin. In general, gelatin can be classified as alkaline gelatin, acidic gelatin, or enzymatic gelatin. Alkaline gelatin can be obtained by treating collagen with abase, such as sodium hydroxide or calcium hydroxide. Acidic gelatin can be obtained by treating collagen with an acid (e.g., hydrochloric acid). Enzymatic gelatin can be obtained by treating collagen with an enzyme (e.g., a hydrolase). Since gelatin may be in the form of a hydrogel, factors affecting the degradation properties of the hydrogels mentioned herein may also apply to gelatin.

[0068] In certain embodiments, the polymer is selected from any FDA-approved synthetic polymer, including, but not limited to, synthetic polymers selected from the group consisting of polycaprolactone (PCL), polylactic acid (PLA), polyglycolide (PGA), poly(lactic- co-glycolic) acid (PLGA), polypropylene oxide (PPG), polyethylene glycol (PEG), poly(amino acid), polyanhydrides, polyurethanes, poly(trimethylene carbonate) (PTMC), polyesters, polyphosphazene, polyvinyl alcohol (PVA), poly(3-hydroxybutyrate-co-3 -hydroxy valerate) (PHBV), polyolefins, such as polypropylene (PP) and polyethylene (PE), derivatives thereof, copolymers thereof, and combinations thereof.

[0069] In one embodiment, the hydrogel is functionalized with a substituent selected from the group consisting of acrylate, methacrylate, alkene, carbonyl, and combinations thereof. In one embodiment, the hydrogel is gelatin functionalized with an acrylate. In one embodiment, the hydrogel is gelatin methacryloyl (GelMA).

[0070] In one embodiment, the cross-linkable functional group is selected from the group consisting of acrylate, methacrylate, amino, mercapto, epoxy, hydroxy, isocyanate, amide, carboxyl, pyrrolidone, siloxane, oxetane, nitrile, ethylene ether, derivatives thereof, and combinations thereof. In one embodiment, the biomaterial comprises an acrylate. In one embodiment, the biomaterial comprises a mercapto group.

[0071] In one embodiment, the elastic recombinant protein is an elastin-like polypeptide (ELP) or a resilin-like polypeptide (RLP). In one embodiment, the elastic recombinant protein has the structure (VPGXG)n, wherein n can be any integer. In one embodiment, the elastic recombinant protein has the structure ([VPGVG]x[IPGVG]y)z, wherein x is an integer from 1 to 50, y is an integer from 1 to 50, and z is an integer from 1 to 50. In one embodiment, the elastic recombinant protein has the structure ([VPGVG]4[IPGVG])i4. In one embodiment, the elastic recombinant protein is derived from Escherichia coli (E. coli).

[0072] In one embodiment, the ratio of the polymer to the elastic recombinant protein is at least 1 : 1. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 10: 1. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 10:2. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 10:3. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 10:4. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 10:5. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 10:6. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 10:7. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 10:8. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 10:9. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 9: 1. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 9:2. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 9:3. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 9:4. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 9:5. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 9:6. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 9:7. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 9:8. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 8:1. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 8:2. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 8:3. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 8:4. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 8:5. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 8:6. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 8:7. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 7:1. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 7:2. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 7:3. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 7:4. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 7:5. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 7:6. In oneembodiment, the ratio of the polymer to the elastic recombinant protein is about 6:1 . Tn one embodiment, the ratio of the polymer to the elastic recombinant protein is about 6:2. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 6:3. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 6:4. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 6:5. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 5:1. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 5:2. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 5:3. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 5:4. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 4:1. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 4:2. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 4:3. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 3:1. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 3:2. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 2:1. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 1: 1.

[0073] In one embodiment, the ratio of the polymer to the elastic recombinant protein is less than 1 : 1. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 1 : 10. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 2: 10. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 3: 10. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 4: 10. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 5: 10. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 6: 10. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 7: 10. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 8: 10. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 9: 10. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 1 :9. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 2:9. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 3:9. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 4:9. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 5:9. In oneembodiment, the ratio of the polymer to the elastic recombinant protein is about 6:9. Tn one embodiment, the ratio of the polymer to the elastic recombinant protein is about 7:9. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 8:9. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 1 :8. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 2:8. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 3:8. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 4:8. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 5:8. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 6:8. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 7:8. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 1 :7. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 2:7. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 3:7. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 4:7. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 5:7. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 6:7. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 1 :6. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 2:6. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 3:6. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 4:6. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 5:6. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 1 :6. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 2:6. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 3:6. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 4:6. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 5:6. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 1 :5. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 2:5. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 3:5. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 4:5. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 1 :4. In oneembodiment, the ratio of the polymer to the elastic recombinant protein is about 2:4. Tn one embodiment, the ratio of the polymer to the elastic recombinant protein is about 3:4. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 1 :3. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 2:3. In one embodiment, the ratio of the polymer to the elastic recombinant protein is about 1:2.

[0074] In one embodiment, the biomaterial further comprises a therapeutic agent. Exemplary therapeutic agents include one or more growth factors (e.g., vascular endothelial growth factor (VEGF), platelet-derived growth factor (PDGF), transforming growth factor beta (TGF-P), fibroblast growth factor (FGF), epidermal growth factor (EGF), and keratinocyte growth factor); one or more stem cells (e.g., adipose tissue-derived stem cells and / or bone marrow-derived mesenchymal stem cells), steroids (e.g., steroids to prevent edema), agents to prevent post-inflammatory skin hyperpigmentation (e.g., hydroquinone, azelaic acid, kojic acid, mandelic acid, or niacinamide); one or more painkillers (e.g., paracetamol / acetaminophen, aspirin, dextropropoxyphene napsylate, co-codamol, opioids (e.g., morphine, codeine, oxycodone, hydrocodone, dihydromorphine, pethidine, buprenorphine, tramadol, etc.). , or methadone), fentanyl, procaine, lidocaine, tetracaine, dibucaine, benzocaine, p- butylaminobenzoic acid 2-(diethylamino)ethyl ester HC1, mepivacaine, piperocaine, dyclonine, or venlafaxine; one or more antibiotics (e.g., cephalosporins, bacitracin, polymyxin B sulfate, neomycin, bismuth tribromocarbonate, or polysporin); one or more antifungal agents (e.g., nystatin), one or more anti-inflammatory agents (e.g., nonsteroidal anti-inflammatory drugs (NSAIDs), e.g., ibuprofen, ketoprofen, flurbiprofen, piroxicam, indomethacin, diclofenac, sulindac, naproxen, aspirin, ketorolac, or tacrolimus), cyclooxygenase and / or cyclosporine (COX-2)-specific inhibitors (COX-2 inhibitors, e.g., rofecoxib (Vioxx®), etoricoxib, and celecoxib (Celebrex®)), glucocorticoid drugs, specific cytokines targeting T-lymphocyte function, steroids (e.g., glucocorticoids (e.g., corticosteroids such as aldosterone, beclomethasone, betamethasone, cortisone, deoxycorticosterone acetate, dexamethasone, fludrocortisone acetate, hydrocortisone, methylprednisolone, prednisone, prednisolone, or triamcinolone) or mineralocorticoid drugs (e.g., aldosterone, corticosterone, or deoxycorticosterone)), or immunoselective anti-inflammatory derivatives (e.g., phenylalanine, one or more antimicrobial agents (e.g., chlorhexidine gluconate, iodine (e.g., tincture of iodine, povidone iodine, or Lugol's iodine), or silver, such as silver nitrate (e.g., as a 0.5% solution),silver sulfadiazine (e.g., as a cream), or Ag in one or more useful carriers (e.g., alginates such as Acticoat®, which comprises a nanocrystalline silver coating on high density polyethylene, available from Smith & Nephew, London, UK, or Silvercel®, which comprises a mixture of alginate, carboxymethylcellulose, and silver-coated nylon fibers, available from Systagenix, Gatwick, UK).+; foams (e.g., Contree® Foam, which contains soft hydrophilic polyurethane foam and silver, available from Coloplast A / S, Humlebaek, Denmark); hydrocolloids (e.g., Aquacel® Ag, which contains ionic silver and hydrocolloids, available from Conva Tec, Skillman, New Jersey); or hydrogels (e.g., Medline® Foam, which contains ionic silver and hydrocolloids, available from Conva Tec, Skillman, New Jersey). Silvasorb®, which contains silver in an ionic state, commercially available from Automotive Industries, Inc. (Mansfield, Mass.); one or more germicides (e.g., alcohols such as ethanol (e.g., 60-90%), 1-propanol (e.g., 60-70%), and mixtures of 2-propanol / isopropanol); boric acid; calcium hypochlorite; hydrogen peroxide; manuka honey and / or methylglyoxal; phenolic (carbolic acid) compounds (e.g., sodium 3,5-dibromo-4-hydroxybenzenesulfonate, tri chlorophenyl methyl iodosalicyl, or triclosan); polyhexanide compounds (e.g., polyhexamethylene biguanide (PHMB)); quaternary ammonium compounds (e.g., benzalkonium chloride (BAC), benzethonium chloride (BZT), bromide cetyltrimethylammonium (CTMB), cetylpyridinium chloride (CPC), chlorhexidine (e.g., chlorhexidine gluconate) or octenidine (e.g., octenidine dihydrochloride); sodium bicarbonate; sodium chloride; sodium hypochlorite (e.g., optionally in combination with boric acid in Dakin's solution); or triarylmethane dyes (e.g., brilliant green); one or more antiproliferative agents (e.g., sirolimus, tacrolimus, zotarolimus, biolimus, or paclitaxel); one or more emollients; one or more hemostatic agents (e.g., collagen, such as microfibrillar collagen, chitosan, calcium-containing zeolite, cellulose, anhydrous aluminum sulfate, silver nitrate, potassium alum, titanium dioxide, fibrinogen, epinephrine, calcium alginate, poly-N-acetate, tilglucosamine, thrombin, clotting factor(s) (e.g., II, V, VII, VIII, IX, X, XI, XIII, or von Willebrand factor, and its activated forms), procoagulants (e.g., propyl gallate), antifibrinolytics (epsilon aminocaproic acid or tranexamic acid), etc.); one or more procoagulants (e g., any of the hemostatic agents described herein, desmopressin, clotting factor(s) (e.g., II, V, VII, VIII, IX, X, XI, XIII, or von Willebrand factor, and its activated forms), procoagulants (e.g., propyl gallate), antifibrinolytics (e.g., epsilon aminocaproic acid), etc ); one or more anticoagulants (e.g., heparin or a derivative thereof, such as low molecular weight heparin, fondaparinux, or idraparinux; aspirinantiplatelet agents such as dipyridamole, ticlopidine, clopidogrel, or prasugrel; factor Xa inhibitors such as direct acting factor Xa inhibitors (e.g., apixaban or rivaroxaban); thrombin inhibitors such as direct acting thrombin inhibitors (e.g., argatroban, bivalirudin, dabigatran, hirudin, lepirudin, or ximelagatran; or coumarin derivatives or vitamin K antagonists (e.g., warfarin (Coumadin), acenocoumarol, atromentin, phenindione, or phenprocoumon); one or more immunomodulators including corticosteroids and nonsteroidal immunomodulators (e.g., NSAIDS as described herein); one or more proteins; one or more vitamins (e.g., vitamin A, vitamin C, and / or vitamin E).

[0075] In one embodiment, the biomaterial is non-cytotoxic. In one embodiment, the biomaterial is biodegradable. In one embodiment, the biomaterial is suturable. In one embodiment, the biomaterial is a biomimetic of lower urinary tract tissue.

[0076] In one embodiment, the biomaterial has an ultimate tensile strength of at least about 10 kPa. In one embodiment, the biomaterial has an ultimate tensile strength of at least about 15 kPa. In one embodiment, the biomaterial has an ultimate tensile strength of at least about 20 kPa. In one embodiment, the biomaterial has an ultimate tensile strength of at least about 25 kPa. In one embodiment, the biomaterial has an ultimate tensile strength of at least about 30 kPa. In one embodiment, the biomaterial has an ultimate tensile strength of at least about 40 kPa. In one embodiment, the biomaterial has an ultimate tensile strength of at least about 50 kPa. In one embodiment, the biomaterial has an ultimate tensile strength of at least about 60 kPa. In one embodiment, the biomaterial has an ultimate tensile strength of at least about 70 kPa. In one embodiment, the biomaterial has an ultimate tensile strength of at least about 80 kPa. In one embodiment, the biomaterial has an ultimate tensile strength of at least about 90 kPa. In one embodiment, the biomaterial has an ultimate tensile strength of at least about 100 kPa. In one embodiment, the biomaterial has an ultimate tensile strength of at least about 110 kPa. In one embodiment, the biomaterial has an ultimate tensile strength of at least about 120 kPa. In one embodiment, the biomaterial has an ultimate tensile strength of at least about 130 kPa. In one embodiment, the biomaterial has an ultimate tensile strength of at least about 140 kPa. In one embodiment, the biomaterial has an ultimate tensile strength of at least about 150 kPa. In one embodiment, the biomaterial has an ultimate tensile strength of at least about 160 kPa. In one embodiment, the biomaterial has an ultimate tensile strength of at leastabout 170 kPa. In one embodiment, the biomatenal has an ultimate tensile strength of at least about 180 kPa. In one embodiment, the biomaterial has an ultimate tensile strength of at least about 190 kPa. In one embodiment, the biomaterial has an ultimate tensile strength of at least about 200 kPa.

[0077] In one embodiment, the biomaterial has a linear percent extensibility of at least about 10%. embodiment, the biomaterial has a linear percent extensibility of at least about 20% In one embodiment, the biomaterial has a linear percent extensibility of at least about 30%. In one embodiment, the biomaterial has a linear percent extensibilit y of at least about 40%. In one embodiment, the biomaterial has a linear percent extensibility of at least about 50%. In one embodiment, the biomaterial has a linear percent extensibility of at least about 60%. In one embodiment, the biomaterial has a linear percent extensibility of at least about 70%. In one embodiment, the biomaterial has a linear percent extensibility of at least about 80%. In one embodiment, the biomaterial has a linear percent extensibility of at least about 90%. In one embodiment, the biomaterial has a linear percent extensibility of at least about 100%. In one embodiment, the biomaterial has a linear percent extensibility of at least about 110%. In one embodiment, the biomaterial has a linear percent extensibility of at least about 120%. In one embodiment, the biomaterial has a linear percent extensibility of at least about 130%. In one embodiment, the biomaterial has a linear percent extensibility of at least about 140%. In one embodiment, the biomaterial has a linear percent extensibility of at least about 150%. In one embodiment, the biomaterial has a linear percent extensibility of at least about 160%. In one embodiment, the biomaterial has a linear percent extensibility of at least about 170%. In one embodiment, the biomaterial has a linear percent extensibility of at least about 180%. In one embodiment, the biomaterial has a linear percent extensibility of at least about 190%. In one embodiment, the biomaterial has a linear percent extensibility of at least about 200%. In one embodiment, the biomaterial has a linear percent extensibility of at least about 210%. embodiment, the biomaterial has a linear percent extensibility of at least about 220%. In one embodiment, the biomaterial has a linear percent extensibility of at least about 230%. In one embodiment, the biomaterial has a linear percent extensibility of at least about 240%. In one embodiment, the biomaterial has a linear percent extensibility of at least about 250%. In one embodiment, the biomaterial has a linear percent extensibility of at least about 260%. In one embodiment, the biomaterial has a linear percent extensibility of at least about 270%. In oneembodiment, the biomaterial has a linear percent extensibility of at least about 280%. In one embodiment, the biomaterial has a linear percent extensibility of at least about 290%. In one embodiment, the biomaterial has a linear percent extensibility of at least about 300%.

[0078] In one embodiment, the biomaterial has a linear percent extensibility ranging from 10% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from20% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from30% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from40% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from50% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from60% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from70% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from80% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from90% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from100% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 110% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 120% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 130% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 140% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 150% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 160% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 170% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 180% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 190% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 200% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 210% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 220% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 230% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 240% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 250% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 260% to 300%. In one embodiment, the biomaterial has a linear percentextensibility ranging from 270% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 280% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 290% to 300%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 100% to 200%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 110% to 200%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 120% to 200%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 130% to 200%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 140% to 200%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 150% to 200%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 160% to 200%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 170% to 200%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 180% to 200%. In one embodiment, the biomaterial has a linear percent extensibility ranging from 190% to 200%.

[0079] In one embodiment, the biomaterial has a suture retention strength of at least about 1 kPa. In one embodiment, the biomaterial has a suture retention strength of at least about 2 kPa. In one embodiment, the biomaterial has a suture retention strength of at least about 3 kPa. In one embodiment, the biomaterial has a suture retention strength of at least about 4 kPa. In one embodiment, the biomaterial has a suture retention strength of at least about 5 kPa. In one embodiment, the biomaterial has a suture retention strength of at least about 6 kPa. In one embodiment, the biomaterial has a suture retention strength of at least about 7 kPa. In one embodiment, the biomaterial has a suture retention strength of at least about 8 kPa. In one embodiment, the biomaterial has a suture retention strength of at least about 9 kPa. In one embodiment, the biomaterial has a suture retention strength of at least about 10 kPa. In one embodiment, the biomaterial has a suture retention strength of at least about 15 kPa. In one embodiment, the biomaterial has a suture retention strength of at least about 20 kPa. In one embodiment, the biomaterial has a suture retention strength of at least about 25 kPa. In one embodiment, the biomaterial has a suture retention strength of at least about 30 kPa. In one embodiment, the biomaterial has a suture retention strength of at least about 35 kPa. In one embodiment, the biomaterial has a suture retention strength of at least about 40 kPa. In one embodiment, the biomaterial has a suture retention strength of at least about 45 kPa. In oneembodiment, the biomaterial has a suture retention strength of at least about 50 kPa. In one embodiment, the biomaterial maintains a suture retention strength of at least about 5 kPa over at least a day.Method of Making Crosslinked Biomaterials

[0080] The present invention further relates to, in part, a method of preparing a biomaterial comprising the steps of: providing a solution comprising a polymer, wherein the polymer comprises at least one cross-linkable functional group; extruding the solution to provide a scaffold; exposing the scaffold to an initiator; and cross-linking the scaffold to provide the biomaterial.

[0081] In one embodiment, the solution comprises about at least 1% (w / v) of the polymer. In one embodiment, the solution comprises about at least 2% (w / v) of the polymer. In one embodiment, the solution comprises about at least 3% (w / v) of the polymer. In one embodiment, the solution comprises about at least 4% (w / v) of the polymer. In one embodiment, the solution comprises about at least 5% (w / v) of the polymer. In one embodiment, the solution comprises about at least 6% (w / v) of the polymer. In one embodiment, the solution comprises about at least 7% (w / v) of the polymer. In one embodiment, the solution comprises about at least 8% (w / v) of the polymer. In one embodiment, the solution comprises about at least 9% (w / v) of the polymer. In one embodiment, the solution comprises about at least 10% (w / v) of the polymer. In one embodiment, the solution comprises about at least 15% (w / v) of the polymer. In one embodiment, the solution comprises about at least 20% (w / v) of the polymer. In one embodiment, the solution comprises about at least 25% (w / v) of the polymer. In one embodiment, the solution comprises about at least 30% (w / v) of the polymer. In one embodiment, the solution comprises about at least 35% (w / v) of the polymer. In one embodiment, the solution comprises about at least 40% (w / v) of the polymer. In one embodiment, the solution comprises about at least 45% (w / v) of the polymer. In one embodiment, the solution comprises about at least 50% (w / v) of the polymer. In one embodiment, the solution comprises about at least 55% (w / v) of the polymer. In one embodiment, the solution comprises about at least 60% (w / v) of the polymer. In one embodiment, the solution comprises about at least 65% (w / v) of the polymer. In oneembodiment, the solution comprises about at least 70% (w / v) of the polymer. In one embodiment, the solution comprises about at least 75% (w / v) of the polymer. In one embodiment, the solution comprises about at least 80% (w / v) of the polymer. In one embodiment, the solution comprises about at least 85% (w / v) of the polymer. In one embodiment, the solution comprises about at least 90% (w / v) of the polymer. In one embodiment, the solution comprises about at least 95% (w / v) of the polymer.

[0082] In certain embodiments, the solution further comprises an elastic recombinant protein. In one embodiment, the solution comprises about at least 1% (w / v) of the elastic recombinant protein. In one embodiment, the solution comprises about at least 2% (w / v) of the elastic recombinant protein. In one embodiment, the solution comprises about at least 3% (w / v) of the elastic recombinant protein. In one embodiment, the solution comprises about at least 4% (w / v) of the elastic recombinant protein. In one embodiment, the solution comprises about at least 5% (w / v) of the elastic recombinant protein. In one embodiment, the solution comprises about at least 6% (w / v) of the elastic recombinant protein. In one embodiment, the solution comprises about at least 7% (w / v) of the elastic recombinant protein. In one embodiment, the solution comprises about at least 8% (w / v) of the elastic recombinant protein. In one embodiment, the solution comprises about at least 9% (w / v) of the elastic recombinant protein. In one embodiment, the solution comprises about at least 10% (w / v) of the elastic recombinant protein. In one embodiment, the solution comprises about at least 15% (w / v) of the elastic recombinant protein. In one embodiment, the solution comprises about at least 20% (w / v) of the elastic recombinant protein. In one embodiment, the solution comprises about at least 25% (w / v) of the elastic recombinant protein. In one embodiment, the solution comprises about at least 30% (w / v) of the elastic recombinant protein. In one embodiment, the solution comprises about at least 35% (w / v) of the elastic recombinant protein. In one embodiment, the solution comprises about at least 40% (w / v) of the elastic recombinant protein. In one embodiment, the solution comprises about at least 45% (w / v) of the elastic recombinant protein. In one embodiment, the solution comprises about at least 50% (w / v) of the elastic recombinant protein. In one embodiment, the solution comprises about at least 55% (w / v) of the elastic recombinant protein. In one embodiment, the solution comprises about at least 60% (w / v) of the elastic recombinant protein. In one embodiment, the solution comprises about at least 65% (w / v) of the elastic recombinant protein. In one embodiment, the solution comprises about at least 70% (w / v) of theelastic recombinant protein. In one embodiment, the solution comprises about at least 75% (w / v) of the elastic recombinant protein. In one embodiment, the solution comprises about at least 80% (w / v) of the elastic recombinant protein. In one embodiment, the solution comprises about at least 85% (w / v) of the elastic recombinant protein. In one embodiment, the solution comprises about at least 90% (w / v) of the elastic recombinant protein. In one embodiment, the solution comprises about at least 95% (w / v) of the elastic recombinant protein.

[0083] In certain embodiments, the solution comprises an organic solvent. In one embodiment, the solution is a polar solvent. In one embodiment, the solution is a polar aprotic solvent. In one embodiment, the solution is one used in the art to solubilize peptides. Throughout the present disclosure, organic solvents of the invention may include, but are not limited to, hexafluoroisopropanol (HFIP), acetonitrile (ACN), dimethylformamide (DMF), tetrahydrofuran (THF), 1,4-di oxane, dimethylsulfoxide (DMSO), N-methyl-2-pyrrolidone (NMP), acetic acid, formic acid, tri fluoroacetic acid, alcohols, such as methanol, propanol, isopropanol, ethane diol, diethylene glycol, and propanediol, alkanes, such as pentane, hexane, heptane, octane, cyclohexane, benzene, toluene, ethylbenzene, acetone, ketones, such as methyl ethyl ketone, methyl isobutyl ketone, and cyclohexanone, acetophenone, ethers, such as diethyl ether, methyl acetate, ethyl acetate, isopropyl acetate, glycol ethers, such as propylene glycol ether, ethylene glycol ether, and ethylene glycol ethyl ether, chloroform, di chloroethane (DCE), nitrobenzene, morpholine, pyridine, acetonitrile, phenol, and combinations thereof.

[0084] In certain embodiments, the step of extruding the solution is performed using any method of polymer extrusion presently known in the art. In certain embodiments, the method of extrusion is one used to produce fibers. Exemplary methods include, but are not limited to, electrospinning, including but not limited to alternating current electrospinning, nozzle-free electrospinning, high-throughput electrospinning, wire electrospinning, bubble electrospinning, ball electrospinning, high speed electrospinning, plate edge electrospinning, bowl electrospinning, hollow tube electrospinning, rotary cone electrospinning, spiral coil electrospinning, electroblowing, co-axial electrospinning, emulsion electrospinning, and melt electrospinning, wet solution spinning, melt spinning, and emulsion spinning. In one embodiment, the step of extruding the solution is performed using electrospinning.

[0085] In certain embodiments, the solution further comprises a therapeutic agent. Thus, the therapeutic agent may become embedded in the scaffold upon extrusion. In certain embodiments, the step of exposing the scaffold to an initiator further comprises the step of submerging the scaffold in a second solution comprising a therapeutic agent. Thus, the scaffold may effectively absorb the therapeutic agent. Exemplary therapeutic agents which may be utilized are discussed elsewhere herein. In certain embodiments, the second solution further comprises an organic solvent. Exemplary organic solvents are described elsewhere herein.

[0086] In certain embodiments, the initiator is used to initiate a cross-linking reaction between the polymer and the elastic recombinant protein. Exemplary initiators that can be used in the present invention include, but are not limited to, photoinitiators and radical initiators including but not limited to halogens, azo compounds, organic peroxides, and inorganic peroxides. In one embodiment, the initiator is Irgacure 2959. In one embodiment, the initiator can be activated by irradiating the initiator with light. In one embodiment, the initiator can be activated by irradiating the initiator with light having a wavelength ranging from 100 nm to 400 nm. In one embodiment, the initiator can be activated by irradiating the initiator with ultraviolet (UV) light. In certain embodiments, the step of crosslinking the scaffold initiates the formation of a carbon-carbon bond, a carbon-sulfur bond, or a sulfur-sulfur bond.

[0087] In one embodiment, the step of exposing the scaffold to an initiator is performed by immersing the scaffold in an initiator solution. In one embodiment, the initiator solution comprises a solvent. In one embodiment, the initiator solution comprises an organic solvent. Exemplary organic solvents are described elsewhere herein. In one embodiment, the initiator solution comprises about at least 0.1% (w / v) initiator. In one embodiment, the initiator solution comprises about at least 0.2% (w / v) initiator. In one embodiment, the initiator solution comprises about at least 0.3% (w / v) initiator. In one embodiment, the initiator solution comprises about at least 0.4% (w / v) initiator. In one embodiment, the initiator solution comprises about at least 0.5% (w / v) initiator. In one embodiment, the initiator solution comprises about at least 0.6% (w / v) initiator. In one embodiment, the initiator solution comprises about at least 0.7% (w / v) initiator. In one embodiment, the initiator solution comprises about at least 0.8% (w / v) initiator. In one embodiment, the initiator solution comprises about at least 0.9% (w / v) initiator. In one embodiment, the initiator solution comprises about at least 1.0% (w / v) initiator.EXPERIMENTAL EXAMPLESThe invention is further described in detail by reference to the following experimental examples. These examples are provided for purposes of illustration only, and are not intended to be limiting unless otherwise specified. Thus, the invention should in no way be construed as being limited to the following examples, but rather, should be construed to encompass any and all variations which become evident as a result of the teaching provided herein.Without further description, it is believed that one of ordinary skill in the art can, using the preceding description and the following illustrative examples, make and utilize the present invention and practice the claimed methods. The following working examples therefore are not to be construed as limiting in any way the remainder of the disclosure.Example 1 : Highly Elastic, Suturable Biomaterial Development for the Lower Urinary Tract

[0088] The application of engineered biomimetic materials offers a promising alternative to autologous grafts for repair or replacement of the bladder or urethra affected by malignant, neurologic, traumatic, infectious, or congenital conditions. Lower urinary tract (LUT) reconstructive surgeries, such as neobladder creation, augmentation cystoplasty, or proximal hypospadias repair, are high-risk and complex procedures. Recent studies report complication rates of 33-68%, in part due to abnormal or inadequate local bladder or urethral tissue (Du, K. et al., 2015, J. Pediatr. Surg., 50, 1535; Kim, S. J. et al., 2022, Urology, 18, 314; Stanasel, I. et al., 2015, J. Urol., 194, 512; Long, C. J. et al., 2017, J. Urol., 197, 852; Mcnamara, E. R. et al., 2015, J. Pediatr. Urol., 11, 209). In such cases, secondary tissue sources are often required. However, current gold-standard autologous grafts such as foreskin, buccal mucosa, or bowel have limited availability, possess undesirable tissue properties (e.g., mucous production by bowel), and increase the risk of donor site morbidity (Choi, J. J. E. et al., 2020, Clin. Exp. Dent. Res., 6, 602; Rahmati, S. et al., 2022, Cell J., 24, 506; Cunnane, E. M. et al., 2021, Biomaterials, 269, 120651). Thus, there is an ongoing need for tissue alternatives with optimized biological and mechanical properties to support healthy LUT regeneration, thereby enabling early return offunction and reducing postoperative morbidity (Versteegden, L. R. et al., 2017, Acta Biomat., 52, 1).

[0089] Historically, LUT tissue engineering has been successful to replace or repair urethral or bladder tissue in healthy small animal models. However, in larger animal and human models synthetic, biologic, and hybrid materials have had limited translatability due to lack of vascularization and tissue regeneration, inconsistent functional outcomes, and need for repeat surgical reconstruction (Chan, Y. Y. et al., 2020, Nature Reviews Urology, 17, 162). The most common historic approach in LUT tissue engineering has been to apply cell-seeded or unseeded decellularized matrices to enhance tissue regeneration. However, these biologic scaffolds have tensile and structural properties that substantially differ from native LUT tissue. An example is one-ply small intestinal submucosa (SIS; Cook Biodesign™) which has a mean tensile strength of 5MPa, approximately 10-fold higher than the urethra. Furthermore, these acellular matrix products have limited reproducibility due to organism-, tissue- and processing-level effects on each individual product. To date, a mechanically matched FDA-approved clinical product is not available for lower urinary tract tissue repair or replacement.

[0090] The pronounced differences in the tensile properties of both current autografts and engineered substrates as compared to native LUT tissue have both microscopic and macroscopic effects following surgical reconstruction. On a microscopic level, the mechanical niche is an important parameter controlling cell fate and extracellular matrix (ECM) deposition. Tissue mechanics and topography impact wound healing, influencing cell ingrowth, morphology, and phenotype (Park, J. S. et al., 2011, Biomaterials, 32, 3921). In the urinary tract, a low-density polarized collagen graft with tensile properties matching native tissue improved smooth muscle regeneration and patency compared to a high-density graft following urethroplasty in an animal model (Larsson, H. M. et al., 2018, Scientific Reports, 8, 10057). Recent development of micropatterned scaffolds likewise improved urothelial regeneration and vascularization in an animal model. Furthermore, wound healing and immune response to materials, including macrophage phenotype, can be directed by bladder extracellular matrix structure, mechanics, and biologic modifications that occur in disease states (Shi, L. et al., 2013, International Journal of Burns and Trauma, 3, 173). On a macroscopic scale, the elasticity in particular of bladder and urethral tissue is crucial for lifelong cyclic voiding and tumescence. This mismatch can lead tocomplications like urine leaks, infection, and dehiscence, necessitating prolonged catheterization, which further increases the risk of complications (Igawa, Y. et al., 2008, International Journal of Urology, 15, 481). Future substrates for use in urethral or bladder reconstructive procedures must account for the macroscopic and microscopic effects of biomaterial mechanics and structure.

[0091] From a biologic standpoint, biomimetic reconstruction strategies for various tissues are a growing field in tissue engineering. Scaffolds with biomimetic characteristics have shown great potential for promoting the reconstruction of damaged tissues and restoring their functionality. However, significant challenges remain in regenerating intricate tissue-based architectures. For example, an ideal tissue-engineered graft for the urethra should exhibit excellent biocompatibility to enhance cell adhesion and feature a highly permeable, three- dimensional (3D) fiber structure that mimics the extracellular matrix (ECM). This design facilitates cell migration and cell-cell adhesion, providing a protective barrier against urine. Additionally, the graft should promote vascularization to support cell metabolism and ensure an adequate blood supply. It must also withstand mechanical forces during surgical procedures. By meeting these criteria, a tissue-engineered graft can closely replicate the native urethra in both structure and function.

[0092] Hydrogels, with their high water content and permeability, show promise in mimicking native tissue microenvironments (Igawa, Y. et al., 2008, International Journal of Urology, 15, 481). Emerging techniques such as 3D hydrogel cultures and microfluidic platforms have demonstrated the importance of surface topography, porosity, and fibrous architecture in directing and enhancing tissue neovascularization and regeneration (Carvalho, E. M., 2023, Acta Biomater., 163, 146; Barkefors, I. et al., 2009, Lab Chip, 9, 529). However, their clinical application is limited by inadequate mechanical properties and unpredictable degradation rates. Collagen has been explored for its potential to reconstruct various urologic tissues, given its prominence in the native ECM, in which collagen plays a crucial role in maintaining structural integrity and directing cell behavior. For instance, collagen-based scaffolds have been used for the repair of urethral defects in a rabbit model, demonstrating significant regeneration of urethral tissue with well-organized smooth muscle layers (Sadder, K. et al., 2017, Semin. Immunol., 29, 14). Additionally, collagen has been incorporated into hybrid scaffolds with synthetic polymers to improve mechanical properties and biocompatibility for bladder augmentation (AhmadianKia,N. et al., 2023, J. Biomater. Appl., 38, 179). Silk fibroin (SF)-based scaffolds have also been applied to engineer patient-specific implants for bladder augmentation (Carvalho, E. M., 2023, Acta Biomater. 163, 146). While these examples underline the promising role of biomimetic components for LUT tissue reconstruction, challenges remain regarding their mechanical strength, saturability, in vivo degradation rates, and implantation.

[0093] To address the need for biomimetic, non-immunogenic, tensile-matched scaffolds that promote the proliferation of LUT cell lines and tissue regeneration, developed herein is a photocrosslinked suturable elastic fibrous hydrogel composed of gelatin methacryloyl (GelMA) and ELP using an electrospinning technology. GelMA, with inherent arginine-glycine-aspartic acid (RGD) sequences, supports cell adhesion (Klotz, B. J. et al., 2016, Trends Biotechnol., 34, 394), while its modification with methacrylamide allows tunable mechanics and degradation (Yue, K. et al., 2016, Trends Biotechnol., 34, 394). ELP, a recombinant protein analogous to naturally-occurring elastin, is biocompatible and biodegradable, ELP possesses intrinsic structural disorder that contributes to elasticity (Chan, Y. Y. et al., 2020, Nature Reviews Urology, 17, 162). In this study, GelMA-ELP composite hydrogels were designed to match the mechanical properties of LUT tissue . Electrospinning was used as an engineering modality to enable control of scaffold topography and contribute to scaffold elasticity. Additionally, the ability of the engineered scaffolds to physically support tissue repair was confirmed through suture retention and ex vivo bladder repair tests. In vitro evaluation of cytocompatibility with urothelial, smooth muscle, and fibroblast cell lines have confirmed the biocompatibility. Finally, in vivo rat implantation studies evaluated the biocompatibility of these novel biomaterials in animal models as illustrated in Figure 1.Mechanical characterization of the ELP GelMA fibrous scaffolds compared to LUT tissue and commercial decelhilarized matrix

[0094] Mechanical properties including the stiffness and elasticity of the ECM play an important role in governing numerous cellular processes and facilitating the development of new urologic tissues (Rossetto, V. J. et al., 2013, Acta Vet Scand., 55, 39; Vo, T. N. et al., 2012, Advanced Drug Delivery Reviews, 64, 1292). With new insights into the influence of mechanical properties on cell behavior, the experiments conducted herein characterize and evaluate themechanical properties of the engineered hydrogels synthesized with different concentrations of GelMA (i.e., 3, 5, 7 and 10% (w / v)) and ELP (i.e., 0, 3, 5, 7% (w / v)) compared to LUT tissues and commercial ECM, by performing tensile and cyclic tensile tests.

[0095] One-ply Cook Biodesign® small intestinal submucosa (SIS) was selected as a commercial product for comparative analysis due to its FDA-approved indication for reinforcement of soft tissues and prior clinical studies supporting its use in corporal body (penile) grafting and bladder neck support during human LUT reconstructive surgeries (Misseri, R. et al., 2005, J. Urol., 174, 1680). It has also demonstrated cytocompatibility with LUT cell lines in prior studies of its use for LUT tissue engineering (Davis, N. F. et al., 2011, Urology, 77, 1007). Static tensile evaluation of 1 -ply Cook Biodesign® SIS (Figure 2A, Figure 2B, Figure 2C) demonstrated that SIS significantly differed from LUT tissue with a tensile modulus of 16041.64 ± 1890.92 kPa, p<0.0001. The extensibility % of SIS (67.83 ± 5.61) was significantly lower than LUT tissue (p<0.01), signifying that SIS is a stiffer, less extensible substrate than LUT tissue.

[0096] LUT tissue samples were obtained from an established animal model of anterior urethroplasty, the New Zealand White male rabbit. One sample was taken immediately post- euthanization from two lower urinary tract sites (anterior urethra, supratrigonal bladder) for use in tensile characterization. Hematoxylin and eosin staining (H&E) images demonstrated intact layered tissue in histologic characterization, verified with samples adjacent to those utilized in tensile testing (Figure 2D, Figure 2E). In static tensile testing, the tensile modulus (urethra mean 58.89 ± standard error of the mean (SEM) 16.54 kPa) did not differ between the bladder (46.81 ± 16.01 kPa) and urethral tissues (p>0.05) (Figure 3A, Figure 3B). Ultimate tensile strength likewise did not differ between the two sites (urethra 99.82 ± 24.14; bladder 138.85 ± 40.56 kPa). While both urethral and bladder samples exhibited remarkable stretchability on static tensile testing, the urethra had significantly higher percent extensibility (%E, 322.59 ± 71.51) as compared to bladder tissue (194.54 ± 21.35) (p<0.01) (Figure 3C). These findings established the highly elastic nature of the LUT tissues and provided target tensile properties for scaffold development.

[0097] Mechanical properties of the engineered hydrogels synthesized with different concentrations of ELP and GelMA were elucidated by performing static tensile (n=5) and cyclic tensile (n=4) tests. Across concentrations in static testing, scaffolds exhibited a mean tensile modulus that ranged from 12.44 ± 2.08 kPa to 25.63 ± 4.77 kPa, without significant difference from bladder or urethral tissue, regardless of polymer concentration. However, 5E5G scaffolds with a mean tensile modulus (i.e., 25.63 ± 4.77 kPa) were most readily handled by surgeons in a qualitative evaluation of tissue-like qualities for suture placement (Figure 3A). The ultimate tensile strength of engineered hydrogels ranged from 19.43 ± 3.08 kPa to 27.48 ± 5.32 kPa, without a significant difference based on respective polymer concentrations between scaffolds and in comparison, to bladder tissue. However, among the polymer concentrations, 5E5G was the only material with UTS (40.96 ± 6.68 kPa) that did not significantly differ from urethral tissue (Figure 3B). While 0E10G demonstrated a significant difference from urethral %E, across tested polymer concentrations with the addition of ELP this difference was no longer observed (Figure 3C). Put together, the addition of ELP improved tensile and qualitative handling properties of the resultant biomaterials in their approximation of native LUT tissue tensile parameters evaluated. Based on tensile testing results, the primary selected candidate material was 5E5G, with GelMA alone used as a second well-characterized biomaterial for subsequent in vitro and in vivo testing.

[0098] Cyclic testing evaluated 10 cycles of extension and relaxation of anterior urethral tissue (n=8 samples evaluated per cohort) versus candidate scaffold material. The 8thcycle is demonstrated as a representative curve with substantial overlap observed from the candidate scaffold materials and was representative of cycles 2-10. Calculated energy loss in cycle 8 is demonstrated in Figure 3E (29.09 ± 6.95%). This evaluation illustrated preserved tensile properties of LUT tissue across multiple deformation cycles with low energy loss. Cyclic tensile testing (10 cycles per sample) of the hybrid hydrogels demonstrated energy loss of 44.86 ± 15.63% for the ELP / GelMA hydrogels (Figure 3E), without a significant difference from urethral tissue. Specifically, low energy dissipation was observed during the loading and unloading cycles, coupled with the high resilience of ELP / GelMA hydrogels, underscores their potential for applications in dynamic organ systems such as the urinary bladder or urethra.

[0099] T he ability of GelMA / ELP fibrous scaffolds to withstand forces resulting from suturing was measured by suture retention strength, which was assessed by applying a controlled force to the suture until it failed as shown in Figure 3F. When ELPs were incorporated into the hydrogel networks, the suture retention strength of the GelMA / ELP electrospun scaffolds significantly increased. The suture retention strength improved from 9.62 2.75 kPa to 14.62 ± 5.47 kPa for scaffold-to- scaffold suturing and from 7.42 ± 1.95 kPa to 13.29 ± 1.49 kPa for scaffold-to-tissue suturing In addition, there is no statistically significant difference in suture retention strength observed between scaftold-to-scaffold suturing and scaffold-to-tissue suturing for both 0E10G and 5E5G scaffolds (Figure 3G). Overall, the results demonstrate that a diverse array of tunable mechanical properties can be achieved through varied polymer concentrations of GelMA and ELP.

[0100] In summary, mechanical properties of the engineered hydrogels synthesized with different concentrations of GelMA (i.e., 3, 5, 7 and 10% (w / v)) and ELP (i.e., 0, 3, 5, 7% (w / v)) were elucidated by performing static tensile (n=5) and cyclic tensile (N=4) tests (Figure 3). By varying the relative concentration of GelMA and ELP, the tensile tests conducted on the resultant scaffolds probed one aspect of their mechanical tunability, and guided polymer concentration selection based on target criteria, ensuring that tensile properties were closely matched with LUT tissue. Across concentrations, scaffolds exhibited a mean TM that ranged from 12.44 ± 2.08 kPa to 25.63 ± 4.77 kPa, without significant difference from bladder or urethral tissue, regardless of polymer concentration (p>0.05). However, although not a significant difference, 5E5G scaffolds exhibited the highest mean TM (i.e., 25.63 ± 4.77 kPa) and was most readily handled by surgeons in a qualitative evaluation (Figure 3A). Secondly, the ultimate strength of engineered hydrogels ranged from 19.43 ± 3.08 kPa to 27.48 ± 5.32 kPa, without a significant difference between them based on respective polymer concentration. However, among the hydrogels synthesized, 5E5G was the only material with ultimate strength (40.96 ± 6.68 kPa) that did not significantly differ from urethral tissue (Figure 3B). Lastly, across ELP concentrations, scaffolds were elastic, without a significant difference in %E from evaluated tissue. However, there was a trend toward improvement in %E with addition of ELP. Furthermore, 0E10G (GelMA alone) had significantly less %E as compared to urethral tissue (p<0.01). (Figure 3C). Put together, the addition of ELP improved tensile properties, aligning with native LUT tissue tensile studies across parameters evaluated. Based on tensile testing results, the primary selected candidate material was 5E5G. Toensure that comparison was made to well -de scribed hydrogel alone with excellent cytocompatibility, 0E10G was used as a second candidate material for subsequent in vitro testing.

[0101] Cyclic tensile testing of the hybrid hydrogels demonstrated favorable recoverability of scaffold materials throughout a 10 stress-strain cycle evaluation. Specifically, energy loss was determined to be 44.86 ± 15.63% for the ELP / GelMA hydrogels (Figure 3E), without a significant difference from urethral tissue. The relatively low energy dissipation observed during the loading and unloading cycles, coupled with the high resilience of ELP / GelMA hydrogels, underscores their potential for applications in urinary tract reconstruction. Overall, the results demonstrate that a diverse array of tunable mechanical properties for static and cyclic strain environments can be achieved by combining various concentrations of GelMA and ELP.

[0102] The data presented herein support that LUT tissue exhibits remarkable elasticity, significantly greater than SIS. Nanofibrous scaffolds comprised of 5% ELP / 5% GelMA weight per volume (w / v, 5E5G) achieved a tensile modulus (TM), ultimate tensile strength (UTS) and extensibility (%E) in static and cyclic testing that did not significantly differ from bladder or urethral tissue. Scaffolds were suturable, with improved performance in suture tensile testing with the addition of ELP, as compared to GelMA alone. Excellent cell viability was observed from day 1 to 7 across cell lines and scaffolds. Proliferation and spreading of SMCs and Fib was robust, while UC demonstrated less scaffold adherence and proliferation as compared to the other two cell lines.

[0103] Tissue samples (n=7) for each anatomic site (anterior urethra, supratrigonal bladder) were surgically excised and used in tensile characterization. H&E images demonstrated intact layered tissue with anticipated histologic characterization (Figure 2D, Figure 2E). The TM and UTS did not differ between the two tissue types. Specifically, rabbit urethral and bladder tissue exhibited a TM of 58.89 ± 16.54 kPa and 153.6 ± 68.67 kPa, respectively (p>0.05) (Figure 3 A, Figure 3B). While both urethral and bladder samples exhibited remarkable elasticity on static tensile testing, the urethra had significantly higher percent extensibility (%E, 322.59 ± 71.51) as compared to bladder tissue (194.54 ± 21.35) (p<0.01) (Figure 3C). These findingsestablished the significant elasticity of the LUT tissues, and provided target tensile properties (TM, %E) for scaffold development.

[0104] In contrast to native tissues, tensile evaluation of 1 -ply Cook Biodesign SIS (Figure 2A, Figure 2B, Figure 2C) significantly differed from LUT tissue with a TM of 16041.64 ± 1890.92 kPa, pO.OOOl. The %E of SIS (67.83 ± 5.61) was significantly lower than LUT tissue (p<0.01), signifying a stiffer, less elastic substrate than healthy tissue.

[0105] Cyclic testing evaluated 10 cycles of extension and relaxation of anterior urethral tissue (n=8 samples evaluated per cohort) versus candidate scaffold material. The 8thcycle is demonstrated as a representative curve with substantial overlap observed from the candidate scaffold materials and was representative of cycles 2-10. Calculated energy loss in cycle 8 is demonstrated in Figure 3E (29.09 ± 6.95%). The findings of this evaluation illustrated the nature of LUT tissue as it maintains its mechanical properties across multiple deformation cycles.Suture retention test: material to material and material to tissue

[0106] When ensuring potential translatability, effective use of a proposed material in a sutured anastomosis is critical to ensure that the graft can be approximated to native tissue without suture pull-through (Meng, X. et al., 2019, RSC Adv., 9, 21258). One-layer running apposition occurred between (N=3 for scaffold, 4 for tissue cohort) candidate material samples or material to urethral tissue (Figure 3F, Figure 3G). 5E5G scaffolds trended toward improved suture retention strength across both conditions as compared to 0E10G (Figure 3G). Specifically, the suture retention strength improved from 9.62 ± 2.75 kPa to 14.62 ± 5.47 kPa for scaffold-to- scaffold suturing and from 7.42 ± 1 .95 kPa to 13.29 ± 1 .49 kPa for scaffbld-to-tissue suturing with the addition of ELP (0E10G vs 5E5G, respectively, p<0.05 for scaffold to urethra tissue versus scaffold to scaffold categories) (Figure 3G).Synthesis and chemical characterization of ELP / GelMA hydrogel and electrospun scaffolds

[0107] Candidate biologic materials were selected to ensure a viscosity that enabled extrusion for electrospinning, followed by cross-linking to impart strength and preferred degradation parameters. Furthermore, materials were selected with key biologic properties, including anticipated cell binding sites and biocompatibility. To this end, a recombinant proteinwas added to a cross-linkable hydrogel with the goal of developing novel nanofibrous scaffolds with target mechanical properties mimicking native LUT tissue.

[0108] To engineer hybrid composite scaffolds based on GelMA and ELP, GelMA was first formed by adding methacrylic anhydride to gelatin derived from cold water fish due to its better solubility at 4°C compared to gelatin from porcine (Figure 4A) (Misseri, R. et al., 2005, J. Urol., 174, 1680). Following that, a highly elastic photocrosslinkable ELP was successfully synthesized and purified based on previous work (Chan, Y. Y. et al., 2020, Nature Reviews Urology, 17, 162). ELP used in this study possesses a carefully engineered sequence that has been experimentally validated to exhibit enhanced mechanical properties, characterized by remarkable stretchability, along with notable structural stability. Moreover, the strategic inclusion of cysteine residues at both ends of the protein sequence enables efficient photocrosslinking (Figure 4B). To prepare ELP / GelMA prepolymer, different concentrations of GelMA (ranging from 3% to 10% (w / v)) and ELP (ranging from 0% to 7% (w / v)) were combined and dissolved in hexafluoro-2-propanol (HFIP) at room temperature. Then the electrospun scaffold was collected in an aluminum collector following the formation of the fibrous scaffold by crosslinking with a 1% (w / v) solution of Irgacure 2959 as a photoinitiator upon exposure to UV light for 600 s. Upon being subjected to UV radiation, the methacrylate moieties within GelMA interacted with the thiol (-SH) groups present in the cysteine residues of the ELPs, C-C bond formation in Gelatin structure, as well as Michael addition between GelMA and ELP (Figure 4C). This resulted in prompt photocrosslinking and the creation of a 3D hydrogel framework that provides a suitable 3D environment for endothelialization (Hayn, M. H. et al., 2009, Urology, 73, 277; Davis, N. F. et al., 2011, Urology, 77, 1007). The effectiveness of the crosslinking process was demonstrated by observing the insolubility of the crosslinked scaffold when placed in an aqueous solution.1HNMR analysis of uncrosslinked systems indicated the presence of peaks corresponding to methacrylate and methacrylamide groups at approximately 5.3 and 5.7 ppm, respectively (referred to as b and c, highlighted in yellow in Figure 4D). The chemical environments of the H protons in these groups are distinct due to the presence of double bonds, resulting in expected peak splitting. However, in crosslinked samples, these peaks were significantly reduced in intensity, providing evidence of successful linkages. These findings align with previous studies on GelMA and ELP (Unal, G. et al., 2021, Bioengineering & Translational Medicine, 6, el0240). SEM images of ELP / GelMA 5:5 % (w / v),(5E5G), and 10% (w / v) GelMA, 0E10G, demonstrated that electrospun scaffolds are highly porous matts (N=2 scaffolds imaged per cohort) (Figure 4E, Figure 4F). Individual fiber diameter (Figure 4G) within 5E5G (0.79 ± 0.04 pm) scaffolds were significantly narrower than those within 0E10G (1.26 ± 0.05 pm) (p<0.01). However, there was not a significant difference in porosity (Figure 4H) between the two scaffold compositions 0E10G (21% ± 0.17) versus 5E5G (18.91 % ± 1.35).

[0109] In summary, a biomimetic hydrogel was developed, with a goal of creating highly elastic, biodegradable, and suturable nanofibrous scaffolds for future application in LUT reconstruction. Specifically, GelMA was synthesized by adding methacrylic anhydride to gelatin derived from cold water fish due to its better solubility at 4 °C compared to gelatin from porcine sources (Figure 4A) (Noshadi, I. et al., 2017, Biomaterials Science, 5, 2093). Following that, a highly elastic recombinant protein ELP was successfully synthesized and purified. As noted in the illustration, ELP used in this study possesses a carefully engineered peptide sequence that has been experimentally validated to exhibit enhanced elasticity, along with notable structural stability. Moreover, the strategic inclusion of cysteine residues at both ends of the protein sequence enables efficient photocrosslinking (Figure 4B).

[0110] To prepare prepolymer solutions, varied select concentrations of ELP (0% to 7% (w / v)) and GelMA (3% to 10% (w / v)) and were dissolved in HFIP in ambient conditions to enhance crosslinking. Prepolymers were then elctrospun, followed by crosslinking of the nanofibrous sheets in a 1% (w / v) solution of Irgacure 2959 as a photoinitiator. UV light exposure followed to complete crosslinking for a total of 600s (Figure 4C). Due to UV radiation, the methacrylate moieties within GelMA interacted with the thiol (-SH) groups present in the cysteine residues of the ELP adding to chemical self-reaction, resulting in an effectively crosslinked material (Figure 4C) (Burdick, J. A. et al., 2011, Adv. Mater., 23, H41, Li, Y. et al., 2021, Composites Part B: Engineering, 223, 109101).

[0111] Analysis of the 1H NMR spectra for uncrosslinked systems indicated the presence of peaks corresponding to methacrylate and methacrylamide groups at approximately 5.3 and 5.7 parts per million (ppm) respectively (Figure 4D). The chemical environments of the H protons in these groups are distinct due to the presence of double bonds, resulting in expected peak splitting.However, in crosslinked samples, these peaks were significantly reduced in intensity, providing evidence of successful linkages and a decrease in the H proton presence of the uncrosslinked material (Figure 4D). These findings align with previous studies on GelMA and ELP (Unal, G. et al., 2021, Bioengineering & Transla. Med., 6, el0240). The effectiveness of the crosslinking process was additionally demonstrated by observing the insolubility of the crosslinked scaffold when placed in an aqueous solution.Stability analysis of crosslinked, electrospun scaffolds

[0112] Excessive water uptake of hydrogels can potentially weaken the interaction between polymer chains, leading to a reduction in the mechanical properties of engineered scaffolds (Zhao, X. et al., 2016, Adv. Healthc. Mater., 5, 108) Moreover, rapid, and excessive swelling rates of implanted scaffolds can lead to significant morphological changes and the risk of compressing nerves and surrounding tissues (Aubin, H. et al., 2010, Biomaterials, 31, 6941).. For clinical applications, it is imperative that designed polymeric networks degrade into nontoxic byproducts while allowing sufficient time to facilitate autologous cell proliferation and support tissue regeneration. Therefore, the enzymatic degradation and water uptake capacity in GelMA / ELP fibrous hydrogels were evaluated to ensure the stability of the scaffolds.Nanotopography: imaging and quantification of fiber dimensions and scaffold porosity

[0113] SEM images demonstrated that post electrospinning, uncrosslinked scaffolds are highly porous and nanofibrous matts (N=2 scaffolds imaged per cohort). Given their structure, these hydrophilic materials are permeable to media and possess expansive binding sites for cell adherence (Figure 4E, Figure 4F). Individual fiber diameter (Figure 4G) within 5E5G (0.207±0.013 pm) scaffolds were significantly smaller than those in 0E10G (0.300±0.026 pm) (p<0.01). However, there was not a significant difference in porosity (Figure 4H) between the two scaffold compositions 0E10G (26.54 % ± 5.40) versus 5E5G (26.82 % ± 1.87) in their uncrosslinked state.Enzymatic degradation of ELP / GelMA composite

[0114] To replicate the urinary tract microenvironment more precisely, degradations of different composites of ELP / GelMA scaffolds (0E10G, 3E7G, 5E5G, and 7E3G) were characterized in the presence of elastase and collagenase at 37 °C until the scaffolds were completely degraded (Figure 5A, Figure 5B). These enzymes are inherent to the urinary tract and play a critical role in the disassembly and remodeling of ECM proteins, including those integral to bioactive scaffolds (Leach, J. B. et al., 2005, Acta Biomaterialia, 155).. As anticipated due to the relative activity of each enzyme, scaffold degradation rates in collagenase solutions were highest in scaffolds formed by GelMA alone, while in elastase solutions, degradation rates increased with increasing ELP concentrations (Figure 5A, Figure 5B). In addition to evaluation in the presence of enzymes, the weight losses of varying composition scaffolds were evaluated at specified intervals in PBS solution (Figure 5C), revealing a trend toward decrease in scaffold weight loss with increased ELP concentration that was not significantly different between cohorts (Figure 5D). This trend may be attributed to the optimal compositions of GelMA and ELP facilitating additional physical crosslinking and complex polymeric networks. Furthermore, given that elastase has been reported to degrade elastin faster than collagenase degrades collagens under identical conditions (Cunnane, E. M. et al., 2021, Biomaterials, 269), a high ELP concentration can lead to accelerated scaffold degradation driven by elastase degradation of ELP. The presence of enzymes also significantly increased scaffold degradation rates. Scaffolds with ELP compositions of 0, 3, 5, and 7% (w / v) were entirely degraded in PBS after 38, 39, 41, and 45 days, respectively. In contrast, all scaffolds were fully degraded in combined elastase and collagenase solutions within 7 days. Overall, these findings suggest that GelMAZELP fibrous scaffolds exhibit modifiable degradation and have the potential to remain present in the physiological environment during in vitro seeding or implantation for up to one month, which would facilitate cell seeding and early ECM deposition for remodeling and tissue regrowth.

[0115] For clinical applications, it is imperative that designed polymeric networks degrade into non-toxic byproducts while allowing sufficient time to facilitate autologous cell proliferation and support tissue regeneration. Thus, in summart, following electrospinning and crosslinking, hydrophilic proposed scaffold materials with varied concentrations (0E10G, 3E7G, 5E5G, and 7E3G, N=3 per composition and endpoint) were evaluated using degradation testing. Degradation was observed in DPBS versus in select enzymes, collagenase and elastase, selected due to their physiologic occurrence and known direct effect on gelatin- and elastin-basedproducts, respectively (Figure 5A, Figure 5B) (Vo, T. N. et al., 2012, Adv. Drug Deliv. Rev., 64, 1292). In DPBS alone, both 0E10G and 5E5G required approximately 40 days to achieve complete degradation, and approximately 20 days for 50% degradation. As anticipated, scaffold degradation rates in collagenase solutions were particularly increased in the presence of GelMA. In elastase solutions, degradation rates increased most significantly in the presence of ELP (Figure 5A, Figure 5B). When the enzymes were combined, the degradation rates increased further, with nearly 100% degradation at 7 days across all scaffold compositions (Figure 5E, Figure 5F). The mass loss of varying composition scaffolds was also evaluated at specified intervals in DPBS (Figure 5E), revealing an early significant decrease in scaffold weight loss with increased ELP compositions. Scaffolds with ELP compositions of 0, 3, 5, and 7% (w / v) were entirely degraded in PBS after 38, 39, 41, and 45 days, respectively. Overall, these findings suggest that ELP / GelMA nanofibrous scaffolds have the potential to remain present in the physiological environment during in vitro seeding for up to one month which would be anticipated to facilitate cell seeding and early ECM deposition for remodeling and tissue regrowth. However, although these concentrations are anticipated to be supraphysiologic, the potential effect of local enzymatic reactions will require further in vivo evaluation to confirm this degradation timeline.Swelling quantification

[0116] The results indicated that both 0E10G and 5E5G scaffolds exhibited rapid swelling within 1 hour of incubation at 4, 25, and 37 °C. However, there were no significant increases in water uptake observed after 4 hours of incubation for all tested samples (Figure 5E, Figure 5F). Additionally, hydrogels with higher ELP concentrations demonstrated lower swelling compared to those with lower ELP concentrations. This can be attributed to the highly hydrophobic repeating motifs ([[VPGVG]4IPGVG]14) present in the chemical structure of ELP (Lopez Barreiro, D. et al., 2023, ACS Biomater. Sci. Eng., 9, 3796). The hydrophobic interactions, combined with an increase in ELP concentration, likely led to higher crosslinking density, resulting in decreased swelling ratios. Furthermore, no significant difference in the swelling ratio of 0E10G was observed with changes in incubation temperature, whereas 5E5G showed significantly higher swelling at 4°C compared to higher temperatures.

[0117] In summary, as anticipated when using hydrophilic scaffold materials, both 0E10G and 5E5G scaffolds (N=3 per composition and endpoint) exhibited rapid swelling within 1 hour of incubation at 4, 25, and 37 °C. However, the scaffolds remained with stable gross appearance and no significant increases in swelling was observed after 4 hours of incubation across material compositions. Additionally, hydrogels with higher ELP concentrations demonstrated lower swelling compared to those with lower ELP concentrations. Furthermore, no significant difference in the swelling ratio of 0E10G was observed with changes in incubation temperature, whereas 5E5G showed significantly higher swelling at 4 °C compared to physiologic temperatures.In vitro cytocompatibility analysis of ELP / GelMA fibrous scaffolds

[0118] Bioactive scaffolds offer structural reinforcement in tissue reconstruction processes. Also, they play a crucial role in modulating cellular behaviors in multiple ways (Zamani, M. et al., 2020, Front. Bioeng. Biotechnol., 8, 579925). The scaffold composition promotes initial cell attachment, a prerequisite for cell proliferation. Additionally, their porous architecture supports cell migration, tissue formation, and essential nutrient and oxygen exchange (Bridge, J. C. et al., 2015, JoVE, 101, e52986; Amesty, M. V. et al., 2021, Frontiers in Pediatrics, 9). Moreover, scaffold mechanical properties can affect cell phenotype and functions via mechanical signal transduction (Mahdavi, S. S. et al., 2021, Materials Today Communications, 27, 102459). To more accurately mimic the LUT cellular environment three cell lines, including human bladder-derived smooth muscle cells (SMC) (Figure 6), bladder- derived urothelial cells (UCs) (Figure 7), and neonatal preputial (foreskin-derived) fibroblast cells (Fib) (Figure 8) were seeded and maintained up to 7 days in monoculture (Fan, F. et al., 2021, Front. Bioeng. Biotechnol., 9, 718377). Candidate scaffolds were found to be cytocompatible, with live-dead assays demonstrating excellent cell survival without demonstrating apoptotic cells at day 1 to 7 endpoints across for all three cell lines (Figure 6A, Figure 6B, Figure 6E, Figure 6F, Figure 7A through Figure 7D, Figure 8A through Figure 8D). Persistent cell viability for all three cell lines up to 7 days post-seeding was also demonstrated with a Prestoblue assay demonstrating an increase across endpoints for SMCs and Fib, with maintenance of relative fluorescence units (RFU) by UCs over time (Figure 61, Figure 71, Figure 81). Specifically, nuclei counts with DAPI staining increased rapidly (>3-fold) for both scaffoldcompositions in the first 7 days following SMC and Fib scaffold seeding (Figure 6E through Figure 6H, Figure 6J, Figure 7E through Figure 7H, Figure 7J, Figure 8E through Figure 8H, Figure 8J). However, despite the preserved metabolic activity and a lack of demonstrated cytotoxicity, UC nuclei counts on the scaffolds significantly increased from day 1 to 7 postseeding while urothelium proliferated on the petri dish surface. Furthermore, a difference was observed in cell distribution on confocal imaging, regardless of scaffold composition. SMC and UC cohorts significantly increased their depth from day 1 to 7 while Fib instead remained primarily located on the scaffold surface with minimal depth achieved at 7 days. In summary, the results indicate that the proposed fibrous scaffolds exhibit excellent overall cytocompatibility across cell lines, with robust proliferation and cell adhesion in SMC and Fib cohorts. While UC cohorts adhered to the scaffold (Figure 7A, Figure 7B), proliferation was limited (Figure 7C, Figure 7D) in comparison to the SMC and Fib cell lines.

[0119] In summary, three cell lines — SMC (Figure 6), UC (Figure 7), and Fib cells (Figure 8) — were seeded and maintained up to 7 days in monoculture conditions on 0E10G and 5E5G scaffolds. Candidate scaffolds were cytocompatible, with live-dead assays exhibiting profuse green fluorescence with an absence of red IF, signifying cell survival at day 1 to 7 endpoints across cell lines (Figure 6A through Figure 6D, Figure 7A through Figure 7D, Figure 8A through Figure 8D). Persistent cell viability for all three cell lines up to 7 days postseeding was also demonstrated due to a significant increase in RFU with Prestoblue assay for SMCs and Fib, with maintenance of RFU by UCs (Figure 61, Figure 71, Figure 81). Specifically, nuclei counts increased rapidly (>3-fold) for both scaffold compositions the first 5 days following SMC and Fib scaffold seeding (Figure 61, Figure 8J), while nuclei remained present but plateaued in the urothelial cohorts (Figure 7J). Despite preserved metabolic activity and a lack of demonstrated cytotoxicity, UC nuclei counts significantly decreased day 1 to 7 post-seeding, while counts significantly increased in SMC and Fib cohorts. Furthermore, a difference was observed by cell line in depth of penetration on confocal imaging which was not effected by scaffold composition. SMC and UC cohorts significantly increased their depth (Figure 6K through Figure 6M, Figure 7K through Figure 7M) from day 1 to 7 while Fib instead remained primarily located on the scaffold surface with minimal depth achieved at 7 days (Figure 8K through Figure 8M). In summary, the results indicate that the proposed nanofibrous scaffolds exhibit excellent overallcytocompatibility across cell lines, with robust proliferation and cell adhesion in SMC and Fib cohorts.In vivo subcutaneous rat implantation of ELP / GelMA and commercial SIS

[0120] Two male adult healthy Sprague Dawley rats (10-12 weeks of age) were designated for each combination of conditions (SIS, 0E10G, 5E5G, and Sham) and timepoint (days 1, 7, 14, 28, and 56), resulting in a total of 40 rats.

[0121] All 40 animals remained in excellent clinical condition postoperatively, without post-operative infections, wound complications, or mortality. After 2 months post-implantation, 0E10G and 5E5G cohorts demonstrated gross degradation without evident remaining scaffold material and appeared grossly consistent with sham cohort findings. SIS remained palpable with slight tissue deformation present at 2 months post-operatively (Figure 9). None of the cohorts had a visible fibrotic capsule present on gross dissection at any endpoint.

[0122] Excised biomaterials were analyzed using CD68 (macrophage), MPO (neutrophil), and CD3 (T-cell) antigens at each of the 5 timepoints compared to a sham sample without implants (Figure 10A, Figure 10B). Rats implanted with SIS had acute day 1 or 7 significantly (p<0.05) higher marker counts compared to sham tissue analyses (Figure 10C, Figure 10D, Figure 10E). Chronic inflammatory response was likewise seen in the increase in T- cell markers above that of sham specimens at day 56. In contrast, rats implanted using 0E10G or 5E5G had counts for early and late immune markers that did not significantly increase from sham values at any endpoint. In total, 0E10G demonstrated a significant decrease in comparison to sham at 4 endpoints (CD68: d28, MPO: d56, CD3: dl and d28), whereas 5E5G demonstrated a decrease at 7 endpoints (CD68: dl and d28, MPO: dl and d56, CD3: dl, d7, and d28) versus sham values. Overall, the 0E10G and 5E5G cohorts demonstrated excellent biocompatibility throughout the scaffold degradation and regenerative period.Discussion

[0123] Hybrid hydrogels have been developed to optimize the mechanical properties and biological activities of resultant biomimetic scaffolds, rendering them well-suited for intended medical applications (Klotz, B. J. et al., 2016, Trends Biotechnol., 34, 394). Moreover, incontrast to hydrogels with a single polymer network, hybrid hydrogels exhibit greater capacities to emulate versatile properties of native physiological microenvironments (Yue, K. et al., 2015, Biomaterials, 73, 254). GelMA and ELP can be microfabricated to incorporate the composition, morphology and other essential properties of native ECM, thereby facilitating cell proliferation and spreading within GelMA-based scaffolds. In addition, tunable physical characteristics has enabled GelMA to become widely used in tissue engineering applications (Zhu, M. et al., 2019, Sci. Rep., 9, 6963). The genetically encoded synthesis provides complete control over the amino acid sequence, molecular weight, and other physicochemical properties ofELPs, allowing for their wide-ranging utilization in biomedical applications where stretchability is of main importance (Al Halawani, A. et al., 2021, Biomolecules, 11; Tian, D. M. et al., 2022, Mater. Today Bio., 15, 100300; Unal, G. et al., 2021, Bioengineering & Translational Medicine, 6, 10240). Despite the excellent inherent bioactivates of GelMA-based hydrogels, their poor elasticity and suturability have limited their clinical applications in urologic reconstructive surgeries (Biotech, C. et al., 2024, Biodesign Biologic Graft Portfolio). Therefore, it is plausible to postulate that the incorporation of highly elastic ELP into GelMA-based hydrogels could effectively increase the elasticity and resilience of the hydrogel system upon photocrosslinking.

[0124] The findings highlight the remarkable elasticity of LUT tissue, while simultaneously demonstrating the feasibility of matching LUT tissue strength and elasticity in bioinspired, biomimetic scaffold development. From the initial tissue testing set, a resultant range of tensile parameters was defined to ensure ‘matched’ selection of material composition as proposed materials. By further adjusting material composition and fabrication techniques, all three tensile parameters were achieved by the candidate material, ensuring that there was no significant difference observed in modulus, strength, or elasticity as compared to healthy LUT tissue. Furthermore, the 5E5G scaffold demonstrated excellent cyclic behavior during tensile testing, retaining its mechanical characteristics over multiple cycles evaluated. The engineered designed scaffolds in this study are also suturable, non-cytotoxic, and reproducible using synthetic biologic materials. Put together, the novel candidate material far outperforms current commercial alternatives such as SIS in its handling and tensile properties, as well as many historic tissue engineering substrates previously developed for the LUT.

[0125] During the preliminary testing and materials development initial phases, the specific proposed materials and modality of scaffold formation was thoughtfully selected to ensure resultant scaffolds with tunable biologic, structural, and mechanical characteristics. These modifiable materials enable refinement of each matt’s microenvironment, and enable analyses of each of macro and micro effects on tissue regeneration. First, gelatin-based biologies were selected due to their inherent arginine-glycine-aspartic acid (RGD) sequences that provide cell binding sites promoting cell adhesion and proliferation. However, gelatin lacks stability at physiologic temperatures and has weak mechanical properties. Thus, GelMA was used which is gelatin modified with synthetic methacrylamide. Methacrylation allows the material to undergo polymerization via ultraviolet (UV) light, resulting in covalent crosslinking to refine mechanical and degradation tunability while retaining its biocompatibility (Zhao, L. Q. et al., 2016, Adv. Healthc. Mater., 5, 108; Aubin, N. J. et al., 2010, Biomaterials, 31, 6941). Second, the recombinant protein ELP was added due to its homology to elastin and tropoelastin which are intrinsically disordered proteins that naturally impart elasticity to the ECM (Leach, W. J. et al., 2005, Acta Biomaterialia, 155; Cunnane, D. N. et al., 2021, Biomaterials, 269). Although consensus exists between native tropoelastin and ELP in its pentapeptide repeat unit, an exception ‘X’ in the sequence further facilitates modular substitution. In this manner, ELP facilitates development of highly elastic cross-linked scaffolds that can be precisely engineered for cell attachment, proteolytic degradation or addition of biologically active peptides or nanoparticles for therapeutic applications (Lopez Barreiro, D., 2023, ACS Biomater. Sci. Eng., 9, 3796). Finally, electrospinning was used as an engineering modality to enable control of scaffold nanotopography and contribute to scaffold elasticity. The resultant diffusible, porous scaffold structure lends itself effectively to modifications post-production whether secondary to cellseeding or via incorporation of small molecules, drugs or nanoparticles to achieve targeted immunomodulation or angiogenesis (Zamani, S. N. et al., 2020, Front Bioeng. Biotechnol., 8, 579925; Bridge, J. C. et al., 2015, J. Visualized Exp., 101, e52986).

[0126] Urethral and bladder tissue possess critical mechanical properties for physiologic function, as these tissues require prolonged distension and recoil during thousands of voiding cycles throughout the life cycle. This study is novel in its approach to LUT tissue engineering due to its selection of tensile targets derived from urethral and bladder tissue to inform biomaterial development. Through this focus on tunability and enhanced elasticity, the candidate5E5G scaffolds demonstrated no significant difference in modulus, strength, or elasticity as compared to healthy LUT tissue. Furthermore, 5E5G exhibits low energy loss across multiple loading and unloading cycles, critical for the resumption of the physiologic function of distensible organ systems. Put together, the selected ELP / GelMA scaffold outperforms current commercial alternatives such as SIS in its potential to match the tensile properties of healthy LUT tissue. During the preliminary testing and materials development phases, the specific proposed materials and modality of scaffold formation were thoughtfully selected to ensure resultant scaffolds with tunable biological, structural, and mechanical characteristics. Specifically, the addition of ELP to GelMA to electrospun scaffolds enhanced tensile strength and extensibility, suturability, and degradation of the resultant biomaterial. It also ensured cytocompatbility, while minimizing acute and chronic local inflammatory responses postimplantation.

[0127] However, the proposed materials are not only promising due to the tunability and matching of their tensile properties. As the in vitro data indicate, additional parameters beyond tensile properties also contribute to cell growth and spreading throughout the 3D scaffold environment. A key aspect of the select materials and electrospinning as a fabrication method are their versatility, enabling creation of multilayered microenvironments with characteristics carefully tuned to each primary cell line (Amesty, M. V. et al., 2021, Frontiers in Pediatrics, 9). This can occur though modification of crosslinking parameters, material ratios, and spinning techniques (e.g., aligned fibers through use of a mandrel) which enables tunability across multiple dimensions including scaffold presence of cell binding sites, nanotopography, swelling, and degradation (Mahdavi, S. S. et al., 2021, Materials Today Comm., 27, 102459). Furthermore, selecting materials with functional groups such as the modifiable peptide repeat sequence ELP enables a range of biologically active modifications that can improve vascularization, wound healing, and immunomodulation (Fan, F. et al., 2021, Front Bioeng. Biotechnol., 9, 718377). Starting with materials that enable structural, mechanical, and biologically active modifications based on each desired organ and translational application ensures the feasibility of tissue matching across a range of applications, patient demographics, and surgical conditions.

[0128] Although human LUT surgical reconstruction is the future translational target of this scaffold development, the New Zealand White rabbit animal model was selected due to itsaccessible urethra, enabling a tubular or patch defect to be created of 1 to 3 cm length (De Filippo, K. B. et al., 2012, J. Tissue Eng. Regen. Med., 9, 257; Jiang, X. Z. et al., 2017, Biomed. Res. Inti., 1). Due to the initial planned application to anterior urethral defects, only healthy young adult males were selected (4-5 mo is defined as timing of post-pubertal status and the youngest feasible age for transfer to facility) (Sengupta, P. et al., 2019, Inti. J. Preventative Med., 10, 2) as the initial model for evaluation. Additionally, the tensile characteristics of this model closely correspond to a mean TM 34kPa+-0.01 reported in human adult male anterior urethral tissue. Finally, prior studies had established a chronic urethral defect model of urethroplasty (Amesty, M. V. et al., 2021, Frontiers in Pediatrics, 9) to recapitulate certain features of the hypospadias defect which is intended for future in vivo studies of these materials. Put together, the rabbit model is selected as the smallest animal model feasible for a moderately extensive repair allowing evaluation of central vascularization and innervation, with tensile properties that are matched to healthy human urethral tissue.

[0129] Despite being non-cytotoxic, cell growth and proliferation on scaffold materials differed between cell types. Specifically, UC demonstrated less cell proliferation on the scaffold materials but had strong adherence and growth on the underlying well plate. Of note, tissue was evaluated to define tensile target properties of a multilayered intact urethra; however, differing microenvironments are present within urethral and bladder tissue as demonstrate by prior studies using atomic force microscopy to evaluate the layers of the bladder wall (Martinez- Vi dal, L. et al., 2023, Comm. Biol., 6, 217). Thus, it is an anticipated finding that all cell lines may not respond in an equivalent manner to tissue matching in this manner. Furthermore, a balance must be considered between cell binding sites and increased elasticity, which may not be the same ratio for all cell lines, particularly epithelial cell lines that must create a watertight barrier, in contact with one another, in a layered fashion for normal physiologic function. Current plans to mitigate this include evaluation UC encapsulation versus seeding on a hydrogel layer alone, increased fiber diameter, addition of grooves to the fibers to enhance binding sites, or modification to increase the tensile strength of the UC layer. For SMCs and Fib, the effects of alignment on cell behavior will be further evaluated, ensuring that future longer-term in vitro studies have physiological development of contractile tissue constructs using these techniques. Put together, this novel biomaterial testing and development for the LUT provides the groundwork to create microenvironments that recapitulate varied regional requirements. Futuremultilayered scaffolds with or without gradient growth factors are envisaged to enhance proliferation of cells and deposition of functional ECM optimized by cell line.

[0130] Limitations of the study may include those inherent in the development of a novel biomaterial with specifications for a unique organ system. Biologic data has wide variations between individuals. Therefore, ensuring that there is no significant difference in select static and cyclic tensile parameters may or may not provide an adequate approximation to ensure translatability across individuals and disease states. Future studies are planned to evaluate effects of tensile properties on ECM deposition, including evaluation of proteomic and gene expression changes by cell line in varied microenvironments. As previously discussed, it is known that structural and mechanical parameters can likewise affect local immune responses; next steps will include evaluation of this response in vitro and in vivo to ensure that healthy wound healing is facilitated by the presence of developed biomaterials. Certainly, further optimization and evaluation in an in vitro and in vivo environment is required, which is planned by starting with a small animal model to ensure biocompatibility followed by larger models of urethroplasty and augmentation cystoplasty that will enable assessment of functionalization, innervation, and vascularization. In the future, this work illustrates a method of early tensile and structural modifications to ensure that modem biomaterial development techniques are applied in a manner that is informed from the outset by target translational needs.Conclusion

[0131] Highly elastic biomimetic scaffolds can be successfully engineered from biologic materials that mimic the tensile properties of LUT tissue using ELP and GelMA. Future scaffold modifications are ongoing to target urothelial proliferation, and to further define effects of the proposed biomaterials on mechanotransduction and ECM deposition. In the future, these highly elastic nanofibrous scaffolds may provide an alternative tissue source for complex bladder and urethral repairs that recapitulates native tissue structure and mechanics.

[0132] Lower urinary tract tissue is unique in its structure and function, with extraordinary tissue elasticity to support lifelong voiding and penile tumescence. Highly elastic biomimetic scaffolds can be successfully engineered from biologic materials that mimic the tensile properties of LUT tissue using ELP and GelMA. This study is novel in its materials andapproach, ensuring resultant constructs had tensile properties optimized from the outset to match these parameters derived from healthy LUT tissue. Furthermore, the results showed the engineered scaffolds were highly biocompatible in vitro, and could promote robust adhesion and proliferation of SMCs and Fib, with maintenance of urothelial cells. Future scaffold modifications are ongoing to target urothelial proliferation, and to further define effects of structural and mechanical refinement of the proposed biomaterials on mechanotransduction and ECM deposition. In the future, it is anticipated that these highly elastic nanofibrous scaffolds may provide an alternative tissue source for complex bladder and urethral repairs that is readily suturable and recapitulates native healthy LUT structure and mechanics.

[0133] Highly elastic engineered GelMA / ELP and ELP-derivative materials are a novel biomimetic, tensile matched, tunable alternative for application in lower urinary track tissue regeneration. In the future, these scaffolds may provide an alternative strategy for suturable, elastic lower urinary tract replacement tissue that can restore early functionality. These materials can be modified to ensure tuned microenvironment by cell line and primary function of each target location and applied in cell-seeded or unseeded configurations.Experimental Section

[0134] Tissue tensile testing evaluated 1 -ply Cook Biodesign small intestinal submucosa (SIS) and male 18-20 week New Zealand White rabbit anterior urethral and bladder tissue.Highly elastic nanofibrous scaffolds were synthesized by electrospinning select concentrations of two naturally derived polymers: gelatin methacryloyl (GelMA) and elastin-like peptide (ELP), followed by crosslinking. Scaffolds were then characterized using nuclear magnetic resonance (NMR) spectroscopy, degradation, swelling, tensile and suture testing, and scanning electron microscopy (SEM). Human bladder-derived urothelium (UC), smooth muscle cells (SMC), and neonatal foreskin fibroblasts (Fib) were seeded in mono cell culture conditions using cellspecific media onto 1cm2scaffolds in 24-well plates at a IxlO4cells / mL concentration. A systematic evaluation of cell viability, adhesion / spreading, cytotoxicity and proliferation using light / fluorescent and confocal microscopy occurred between day 1 and 7 endpoints. T-tests were used for pairwise comparisons; one-way ANOVA to evaluate between-group differences.Animal model and tissue sampling

[0135] Male New Zealand White rabbit bladder and penile tissue were collected immediately post-euthanasia from 18- to 20-week-old animals as a part of an institutional tissue sharing program (Institutional animal care and use committee (IACUC) Approval Number: 18- 043). Animals had previously undergone a tracheal procedure with 3 days of dexamethasone received 6 weeks prior to LUT tissue extraction at time of euthanasia as per IACUC protocol. There was no additional ongoing medication administration, or effects anticipated to lower urinary tract tissue based on the upper airway surgical procedure performed. For each prosection, the bladder was separated from the urethra at the bladder neck, opened with a vertical anterior incision and the ureteral orifices were identified. Supratrigonal bladder tissue samples were obtained. From the phallus prosection, the external foreskin of the anterior urethra was excised. Next, anterior urethral samples were excised intact from the underlying cavemosal bodies. These samples were then divided dorsally to create longitudinal sheets for testing. All dissections were completed by a pediatric urologist using 2.5x loupe magnification. Each sample was approximately 10mm x 5mm; 2mm sections were obtained proximal to, distal to, and between the sections for histologic structural verification.Histologic tissue evaluation

[0136] Two mm cross sections adjacent to each tissue sample were fixed in 4% paraformaldehyde (PF A), followed by dehydration and xylene preparation. The tissue was then paraffin embedded, and transferred to slides following 5pm sectioning. Hematoxylin and eosin (H&E) staining (ABCAM, ab245880) occurred by manufacturer protocol (ABCAM, H&E Staining Kit 2023). Slides were examined using an Evos M5000 (Invitrogen, Thermo Fisher) microscope, with capture of representative images for verification of intact layered tissue structure.Tissue tensile testing

[0137] Tissue samples were prepared as described in rectangular sections, and tensile testing occurred within 24 to 48 hours of resection. Prior to use, the tissue remained moistened in Dulbecco’s phosphate buffered saline (DPBS, Gibco) and stored at 4°C. At time of testing, the tissue was warmed to room temperature (ambient conditions), while ensuring that it remained moistened with DPBS throughout the evaluation process. Preliminary data included evaluation inradial and longitudinal orientation for both urethral and bladder tissue; however, differences in elasticity were not observed (Chun, S. R. et al., 2019, Engineering a Biomimetic Scaffold for Urethral Tissue Engineering Application). Additionally, urethral tissue was diminutive in the radial direction leading to tensile testing failures within the grips. Due to this, testing proceeded in a longitudinal orientation for the representative tensile data informing materials development. Tissue sample dimensions and thickness were measured using a digital caliper. An Instron 5943 mechanical tester was used to perform all mechanical tests. For the static tensile tests, all samples were fixed between two pieces of double-sided tape and secured within tension grips, followed by tensile stretch at a rate of 1 mm / min until failure. The tensile strain (mm) and load (N) were recorded using the Instron Bluehill software. Tensile modulus (Young’s, TM) was calculated as the tangent slope of the stress-strain curve. Ultimate tensile strength (UTS) and percent elongation (%E) were calculated at the point of maximal strain associated with sample rupture. Cyclic tensile tests (10 cycles / sample; 8 per cohort) were performed at maximum 150% strain and a rate of 1.5% / s by performing 10 cycles of loading and unloading. Cycle 8 was presented as the representative curve; energy loss was calculated by obtaining the area between the loading and unloading curve for cycle 8.Gelatin Methacryloyl (GelMA) Synthesis

[0138] GelMA was synthesized following a previously described method (Nichol, J. W. et al., 2010, Biomaterials, 31, 5535). In summary, 10% weight per volume (w / v) cold water fish- derived gelatin (Sigma-Aldrich) in DPBS was heated to 100 °C until complete dissolution. Subsequently, methacrylic anhydride (Sigma-Aldrich) was gradually added to the gelatin solution under vigorous stirring at 100 °C to modify the lysine groups on the gelatin chains. After 3 hours, the methacrylation reaction was terminated by addition of lOOoC pre-heated DPBS. The solution was then dialyzed in dialysis tubes (Spectrum Laboratories) and immersed in deionized water at 100 °C for 5 days to remove any unreacted methacrylic anhydride. After dialysis, the contents were stored at -80 °C for at least 2 hours. Resultant materials were lyophilized for 5 days to obtain the final GelMA product, which was stored in ambient sterile conditions until electrospinning.Elastin-like Polypeptide (ELP) Synthesis

[0139] The photocrosslinkable ELP sequence for recombinant protein synthesis was edited as previously described (Zhang, Y. N. et al., 2015, Adv. Funct. Mater., 25, 4814). This kanamycin-resistant protein comprises 70 repetitions of the pentapeptide VPGVG, with an isoleucine substitution for the first valine every five pentapeptides, resulting in a pattern of ([VPGVG]4[IPGVG])14. To enable photocrosslinking through the formation of S-S thiol bonds, Lys-Cys-Thr-Ser (KCTS) residues were incorporated on both sides of the ELP sequence. Recombinant ELP was expressed by Escherichia coli (E. coli) as the host organism. This was followed by lysis and purification via inverse transition cycling (Lin, M. et al., 2006, Biochem. J. 398, 577). ELP solutions were then subjected to alternate equilibration and centrifugation above and below their thermal transition temperature (Tt = 29 °C in a 1% w / v solution). This temperature cycling takes advantage of the ELP solubility at low temperatures and precipitation at high temperatures, allowing for the removal of impurities and obtaining a pure ELP product. The purified ELP solution was subsequently dialyzed in a water bath at 4 °C against deionized (DI) water for 7 days following lyophilization until no weight change was observed. Resultant ELP was stored in ambient sterile conditions until electrospinning.Prepolymer solution synthesis and electrospinning formation of nanofibrous mats

[0140] A prepolymer solution was prepared by dissolving select weight percentages of GelMA (3, 5, 7, 10% (w / v)) with or without ELP (0, 3, 5, 7% (w / v)) in 1,1, 1 ,3,3, 3-Hexafluoro-2- Propanal solvent (HFIP) (Sigma-Aldrich). 1.5 mL solution was loaded into a syringe, which was placed into a syringe pump affixed to a high voltage (Glassman High Voltage, Inc., Series EH) custom electrospinning system. A voltage of 25 kV was applied to the needle and aluminum collecting plate (25 cm x 25 cm) positioned 13 cm from the syringe nozzle. The prepolymer solution was extruded at a rate of 1 mL / h. Each nanofibrous scaffold was then removed from the collection plate and placed in the desiccator overnight in ambient conditions to remove remaining solvent. Uncrosslinked scaffolds were stored as needed in dry, ambient conditions. At time of planned analysis or application, the scaffolds were immersed in 1% (w / v) solution of Irgacure 2959 (Sigma-Aldrich) photoinitiator in ethanol for 2 hours. Crosslinking was achieved by applying UV light (6.9 mW / cm2, EXFO OmniCure S2000) for 600s per sample, followed by DPBS wash.Chemical characterization of candidate materials

[0141] Crosslinked and uncrosslinked candidate nanofibrous scaffolds (1 cm x 1 cm) were dissolved in 1 mL dimethyl sulfoxide (DMSO-d6) in 37 °C. 1H NMR (nuclear magnetic resonance) spectroscopy was performed using a 400 mHz Bruker AV400 spectrometer on resultant samples. The degree of methacrylation of polymers was evaluated by NMR analysis as per previously published methods (Nichol, J. W. et al., 2010, Biomaterials, 31, 5536). Providing a comparison for the area under these respective peaks within a crosslinked sample versus an uncrosslinked prepolymer solution quantified the degree of crosslinking achieved.Mechanical characterization of electrospun nanofibrous scaffolds

[0142] GelMA and ELP / GelMA nanofibrous scaffolds were prepared as described, including prepolymer synthesis, electrospinning, and crosslinking. The central portion of each scaffold was sectioned to obtain rectangular samples of approximately 1cm x 2cm. Methods described herein for tissue were applied to complete both static and cyclic testing of these samples.Nanolopography: imaging and quantification of fiber dimensions and scaffold porosity

[0143] Each candidate scaffold sample was mounted on SEM stubs and gold sputter coated (SC7620, Quorum Technologies, UK). The nanotopography of electrospun scaffolds were then characterized by SEM (ZEISS Supra 40VP scanning electron microscope) at an accelerating voltage of 10KV (magnification: 10,000). Imaged was used to quantify porosity by categorizing non-material based porous spaces after removing all labels on each image to maximize analysis area. Fiber diameter was evaluated manually in Imaged by obtaining 10 representative fiber measurements per image.Degradation quantification with versus without enzymatic assays

[0144] Circular samples of consistent size were obtained from the central portion of candidate scaffolds, using a 5mm punch biopsy. The samples were lyophilized, weighed, and recorded as initial dry weight W0 at Oh. Then the samples were incubated until completely degraded at 37°C in DPBS. Samples were immersed (Ma, P. X. et al., 2008, Adv. Drug Deliv.Rev., 60, 184) in 0.1 U / mL porcine elastase solution, 1 .5 U / mL type I collagenase solution, or combined in equal volumes 0.1 U / mL of porcine elastase solution and 5U / mL of type I collagenase solution in DPBS, respectively. At day 1, 2, 4, 7, 10, 14, 21, 28 (extended to day 30, 38, 40 as needed) to the final endpoint of complete degradation, a subset of samples from each candidate material were retrieved, lyophilized, and weighed. The incubation solution was replaced every 3 days throughout the degradation period.

[0145] Percent degradation (D%) was calculated using the equation: % = l-Mt / M0 x 100 where M0 is the initial dry mass of the sample and Mt is the dry mass at time t.Swelling quantification following in vitro hydration

[0146] Circular samples were obtained using a 5mm punch biopsy. Samples were washed with DPBS, lyophilized, and weighed. Each was immersed in DPBS at 4°C, 37°C, or room temperature for 1, 2, 4 and 8 h, respectively (until no increase in weight was observed). Three samples per endpoint were removed and immediately weighed to obtain the hydrated weight.

[0147] Percent swelling ratio (SR%) was calculated using the equation:SR% = (Ms - M0) / M0) x 100%x 100 where M0 is the initial dry weight of the sample and Ms is the swelling weight of the sample at time t (Wang, M. et al., 2011, Tissue Engineering Part B: Reviews, 17, 177; Song, X. et al., 2017, Polymers, 9, 638).Suture retention test: material to material and material to tissue

[0148] Candidate nanofibrous scaffolds were prepared as described through the crosslinking and DPBS wash steps. Rabbit urethral tissue was prepared for tensile testing (distal and proximal longitudinal anterior urethral samples as described. A running single layer approximation of tissue to scaffold or scaffold to scaffold (1cm x 2cm samples) was completed using 5-0 monofilament suture (PDS II, J&J MedTech Ethicon) on an RB-2 needle by a pediatricurologist. Full thickness 2mm width of each sample was included at the suture line. An Instron 5943 mechanical tester was then used to perform the tensile portion of the suture retention test. Sutured samples were fixed between two pieces of double-sided tape within the tension grips. Force was applied at a rate of 1 mm / min. Suture retention strength (endpoint) was defined as the maximum mechanical load (kPa) applied prior to suture pullout from the respective materials or tissue samples.Cell line selection, verification, and expansion

[0149] Three commercial cell lines consisting of human bladder smooth muscle cells (SMC, ATCC PCS-420-012), human bladder urothelium (UC, ATCC PCS-420-010), and human neonatal foreskin fibroblasts (Fib, ATCC BJ CRL-2522) were expanded in manufacturer recommended cell-line specific media (ATCC) and sterile cell culture conditions (37°C, 5% CO2 humidified atmosphere). For quality control, each cell line was verified to exhibit cell-line specific markers including alpha-smooth muscle (oc-SMA, Invitrogen by Thermo Fisher), pan- Cytokeratin (CK, Invitrogen by Thermo Fisher), or Anti-fibroblast antibody, (TE-7, Millipore- Sigma), respectively.Monoculture ofLUT cell lines on candidate biomaterials

[0150] Once expanded (passage 2-5), 104 cells / scaffold of each cell line, respectively, was applied dropwise to each individual candidate scaffold surface within 24-well plates with 500 pL of growth medium (ATCC PCS-100-030, PCS-100-042 for SMC, ATCC PCS-420-032, PCS-420-042 for UC, and ATCC 30-2003, 30-2020 for Fib). After 1 hour, an additional 500 pL of growth medium was added to each well. UC and SMC media included 0.5% penicillin / streptomycin (P / S) (ATCC 30-2300) and 1% (P / S) was included in Fib media.Cultures were maintained at 37°C in a 5% CO2 humidified atmosphere with media replacement each 48h.Cell viability assay

[0151] The viability of SMC, UC, and Fib cell lines, respectively, following seeding of nanofibrous scaffolds were evaluated using a commercial live-dead viability kit (Invitrogen) in accordance with manufacturer instructions (T. Fisher, Live Head Cytotoxicity Assay, 2023).Briefly, 1 mL of Calcein, AM cell permeant dye was combined with 1 pL of BOBO-3 Iodide. At select day 1, 5, and 7 endpoints, 50 pL / mL of staining solution was added to each culture well, with subsequent incubation in dark ambient conditions for 15 minutes preceding imaging.Fluorescent images were acquired using an Evos M5000 microscope at 4x magnification (Invitrogen by Thermo Fisher). Due to the parameters of the assay, viable cells were visualized as green and apoptotic / dead cells as red.Metabolic activity assay

[0152] A commercial PrestoBlue cell metabolic activity assay (Life Technologies) was used in accordance with manufacturer instructions (T. Fisher, Presto Blue assay, 2023) to evaluate the metabolic activity of each cell line, respectively, at days 1, 3, 5 and 7 post-seeding. Briefly, each cell line was incubated in cell-specific growth medium. At each endpoint, 10% PrestoBlue reagent was added to the media for 1 h at 37 °C. After Ihr in ambient conditions, four 100 pL aliquots of PrestoBlue containing solution was removed from each well and transferred into separate wells each of a 96-well plate. The resulting fluorescence was measured (excitation 540 nm; emission 600 nm) using a Synergy HT fluorescence plate reader (BioTek). Solution from wells containing only media and PrestoBlue established the background fluorescence for all experiments. The Relative fluorescence units (RFU) were calculated for each sample by subtracting mean blank values from experimental values.Cell adhesion, proliferation, and spreading

[0153] Surface presence of cells (nuclei) and cytoskeleton on candidate scaffolds were visualized through immunofluorescent (IF) staining of F-actin 54 filaments and cell nuclei. Briefly, at day 1, 5, and 7 post-seeding cells were fixed with 4% (w / v) neutral buffered formalin (Epredia) for 15 min., rinsed with DPBS, and permeabilized in 0.1% (w / v) Triton X-100 (Sigma) for 15 min. After permeabilization and a subsequent DPBS rinse protocol, 2 drops / mL ActinRed™ 555 ReadyProbes™ reagent (Rhodamine phalloidin) (Thermo Fisher) were added to each well. The samples were incubated in actin stain for 30min in ambient conditions. Subsequently, 2 drops / mL NucBlue™ Fixed Cell ReadyProbes™ Reagent (DAPI) (Thermo Fisher) were added to each sample followed by a subsequent 30 min incubation. Fluorescent imaging was carried out as previously described. Confocal images were also obtained day 1- and 7-days post-cell seeding of eachcandidate scaffold to evaluate depth of cell penetration and distribution throughout the nanofibrous 3D structure. Samples were stained for both F-actin and DAPI then embedded in 1% agar. Samples were then imaged with a TCS SPE High-resolution Spectral Confocal Microscope (Leica). Images were processed with Lightning deconvolution software (Leica). Analysis of cellular distribution was completed through measurement of the Z-position of each nucleus as compared to the most superficial nuclei position on each scaffold.Statistical analysis

[0154] The reported data consist of the mean values derived from a minimum of three replicates, along with their respective standard deviations. The statistical analysis was conducted via a two-way ANOVA, followed by a post-hoc Tukey's multiple comparison test, utilizing the GraphPad Prism software (version 8.2.1). The level of statistical significance was indicated as follows: * for p < 0.05, ** for p < 0.01, *** for p < 0.001, and **** for p < 0.0001.

Claims

CLAIMSWhat is claimed is:

1. A biomaterial comprising a polymer and an elastic recombinant protein, wherein the polymer comprises at least one cross-linkable functional group.

2. The biomaterial of claim 1, wherein the polymer is a hydrogel selected from the group consisting of gelatin, collagen, chitosan, polyethyleneimine, hyaluronic acid, bovine serum albumin (BSA), dextran, derivatives thereof, and combinations thereof.

3. The biomaterial of claim 1, wherein the polymer is a synthetic polymer selected from the group consisting of polycaprolactone (PCL), polylactic acid (PLA), polyglycolide (PGA), poly(lactic-co-glycolic) acid (PLGA), polypropylene oxide (PPO), polyethylene glycol (PEG), poly(amino acid), derivatives thereof, copolymers thereof, and combinations thereof.

4. The biomaterial of claim 1, wherein the polymer is functionalized with a substituent selected from the group consisting of acrylate, methacrylate, alkene, carbonyl, and combinations thereof.

5. The biomaterial of claim 1, wherein the cross-linkable functional group is selected from the group consisting of acrylate, methacrylate, amino, mercapto, epoxy, hydroxy, isocyanate, amide, carboxyl, pyrrolidone, siloxane, oxetane, nitrile, ethylene ether, derivatives thereof, and combinations thereof.

6. The biomaterial of claim 1, wherein the polymer is gelatin methacryloyl (GelMA).

7. The biomaterial of claim 1, wherein the elastic recombinant protein is elastin-like polypeptide (ELP).

8. The biomaterial of claim 1, wherein the ratio of the polymer to the elastic recombinant protein is at least 1: 1.

9. The biomaterial of claim 1, wherein the biomaterial further comprises a therapeutic agent.

10. The biomaterial of claim 1, wherein the biomaterial is biodegradable.

11. The biomaterial of claim 1, wherein the biomaterial is a biomimetic of lower urinary tract tissue.

12. The biomaterial of claim 1, wherein the biomaterial has an ultimate tensile strength of at least 15 kPa.

13. The biomaterial of claim 1, wherein the biomaterial has a linear percent extensibility of at least 100%.

14. The biomaterial of claim 1, wherein the biomaterial has an ultimate tensile strength of at least 15 kPa and a linear percent extensibility of at least 100%.

15. The biomaterial of claim 1, wherein the biomaterial has a suture retention strength of at least 8 kPa.

16. A method of preparing a biomaterial comprising the steps of: providing a solution comprising a polymer, wherein the polymer comprises at least one cross-linkable functional group; extruding the solution to provide a scaffold; exposing the scaffold to an initiator; and cross-linking the scaffold to provide the biomaterial.

17. The method of claim 16, wherein the solution further comprises an elastic recombinant protein.

18. The method of claim 16, wherein the solution comprises at least 3% (w / v) of the polymer.

19. The method of claim 17, wherein the solution comprises at least 1% (w / v) of the elastic recombinant protein.

20. The method of claim 16, wherein the solution further comprises a therapeutic agent.

21. The method of claim 16, wherein the step of extruding the solution is performed using electrospinning.

22. The method of claim 16, wherein the step of exposing the scaffold to an initiator further comprises the step of submerging the scaffold in a second solution comprising a therapeutic agent.

23. The method of claim 16, wherein the initiator is a photoinitiator.

24. The method of claim 16, wherein the step of crosslinking the scaffold comprises the step of irradiating the scaffold with light having a wavelength range of 100 to 400 nm.

25. The method of claim 16, wherein the step of crosslinking the scaffold initiates the formation of a carbon-carbon bond or a sulfur-sulfur bond.

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