WO2010048281A1

WO2010048281A1 - Composite biomimetic materials

Info

Publication number: WO2010048281A1
Application number: PCT/US2009/061477
Authority: WO
Inventors: Elliot L. Chaikof; Jeffrey Caves; Mark Allen
Original assignee: Emory University; Georgia Tech Research Corporation
Priority date: 2008-10-21
Filing date: 2009-10-21
Publication date: 2010-04-29

Abstract

In an embodiment, composite materials having a fiber of relatively high Young's modulus is embedded in an elastin mimetic material with a relatively low Young's modulus. In an embodiment, the fiber is crimped so as to provide a composite material having a non-linear stress-strain curve, with an initial Young's modulus at lower strains that is less than the Young's modulus at higher strains. The composite materials are optionally formed into medical devices. Further provided are methods of making composite materials, making a collagen fiber, and crimping fibers such as collagen fibers.

Description

COMPOSITE BIOMIMETIC MATERIALS

CROSS-REFERENCE TO RELATED APPLICATIONS

[001] This application claims priority of U.S. Provisional Patent Application Serial No. 61/107,204 filed October 21 , 2008, which is hereby incorporated by reference in its entirety to the extent not inconsistent with the disclosure herein.

BACKGROUND OF THE INVENTION

[002] The invention generally relates to composite materials, particularly a first relatively elastic material that mimics the mechanical properties of elastin, that supports a second fiber material having a relatively high Young's modulus that mimics the mechanical properties of collagen. Exemplary proteins used in the first material include elastin-mimetic proteins and in the second material include spun collagen from synthetic or animal sources. Related methods of producing and using the same, such as in medical devices and/or medical procedures, are provided. In an aspect, the second fiber material is, in an unstressed state, crimped.

[003] The concept of biomimetics applied to the development of soft tissue substitutes that mimic or expand upon the biomechanics of the extracellular matrix is employed. Because the key non-living constituents of biological structures are fibrous networks of collagen and elastin, it may be possible to fabricate suitable artificial structures from fibrous collagen and elastin analog materials. As a first step, we investigate the fabrication of structures without living cells. The elastin analogs, elastin-mimetic triblock copolymers, are previously reported [2] and have been a topic of continuing research [3- 7]. Because a suitable analog for the fibrous collagen component of soft tissues is not readily available; provided herein are methods for producing artificial collagen fiber. To provide additional context to the materials disclosed and claimed herein, the structure and mechanics of arteries and heart valve leaflet, the application of collagen as a biomaterial, and the properties of elastin-mimetic triblock copolymers are briefly described.

[004] The elasticity and structure of arteries. The arterial side of the circulatory system carries blood from the heart to the tissues of the body. To perform this role smoothly, continuously, and efficiently, arteries act as both conduits and pressure reservoirs. As the left ventricle contracts, arteries expand under pressure, and as the ventricle relaxes, arterial diameters recoil. Thus, arteries are compliant tubes that elastically absorb a portion of the energy from ventricular contraction, and return this energy to aid in blood circulation when the ventricle relaxes. This role is most pronounced in the aortic arch, which has been described as the "auxiliary pump" of the circulatory system [8]. The central arteries are precisely tapered along the length of the aortic arch, thoracic aorta, and abdominal aorta and in terms of decreasing size and increasing stiffness. The taper appears to result in destructive interference between outgoing and reflected pressure waves, further smoothing pressure and reducing the load on the heart [9, 10].

[005] The arterial wall is a laminated structure, consisting of the inner tunica intima followed by the tunica media, and the tunica adventitia. The intima is only 0.2 - 2 μm thick. In many arteries, the intima consists of a single layer of endothelial cells supported by the basil lamina, a collagenous layer. In some arteries, the intima includes more connective tissue and some smooth muscle cells. By mass, the aortic intima is reported to be 65% water, 16% type I collagen, 8% type III collagen, 5% cellular components, and lesser amounts of other collagens and proteoglycans [8].

[006] The intimal layer is followed by an elastic sheet, the internal elastic lamina, and then by the media. Of the three layers, the media contributes the most to arterial mechanics. The structure-function units of the media are concentrically arranged sheets termed medial lamellar units (MLU). The MLU of the mammalian aorta has been studied most extensively. An aortic MLU is a -15 μm thick sheet consisting of one or occasionally two layers of smooth muscle cells surrounded by dense sheets of elastic fibers. The elastic sheet boundaries of the MLU appear severely wrinkled when cross- sections are taken in the absence internal pressure. However, when vessels are fixed under a small amount of internal pressure, below physiologic pressure, the vessel diameter expands considerably, the wall thins, and the elastic sheets become smooth concentric arcs [11]. Close examination reveals that the elastic layers demarcating two MLU are double layers; one layer is associated with each MLU. Bundles of collagen fibers run between the MLU. Collagen bundles tend follow a wavy or helical path even when fixed at physiologic pressures that eliminate the wrinkled path of the elastic layer. Most of the bundles run circumferentially around the aorta, within plus or minus about 5°. All bundles associated with a particular MLU are essentially parallel; variability in orientation is thought to be between rather than within the MLU. In smaller aortas, such as the rat and immature rabbit, a single MLU can wrap entirely around the circumference of the vessel to form a ring. In larger animals such as pigs, a single MLU attains a maximal circumferential length of 4 to 5 mm, and does not encircle the entire aorta [12]. The width of an MLU, measured along the axis of the aorta, varies with vessel and animal size from 100-200 μm in the rabbit to 1 -3 mm in the pig. At bend or branch points in the vessel, the concentric MLU layout is abandoned for more complex structural arrangements [12].

[007] The aortic MLU is considered a functional unit because it is specialized to bear a tension of 1.1 to 3.0 N/m in most mammals [13]. This calculation was motivated by the observations that (i) the aortic MLU thickness is roughly constant, and (ii) the thickness of the aortic media is closely related to aortic diameter and total wall tension. Therefore, when total wall tension is divided by number of MLU through the thickness of the media, a roughly constant tension per MLU results. Consequently, the rat aorta has an average of approximately 5 MLU layers while the sow has 72, corresponding to their different aortic wall tensions [14]. The difference in wall tension is directly related to the difference in vessel diameter.

[008] A recent study using serial block-face scanning election microscopy to examine rat abdominal aorta MLU has permitted enhanced structural resolution and volume calculations [15]. These estimates revealed that the MLU volume was 29% elastin, 24% smooth muscle cell, and 47% collagen and ground substance. Of the elastin, 71 % was part of the lamellar sheets, 27% occurred as fine fibers protruding into the MLU, and 2% as thick radial struts spanning between the lamellar sheets. Smooth muscle cells inside the MLU were encased in cages of thin elastic fibers. The cell nuclei were elliptical, aligned circumferentially, but tilted 20° in the radial direction. The cytoplasm shape was highly irregular. Collagen was present between elastic layers in dense, parallel, coiled bundles of 24 ± 15 fibers, and as thinner bundles and single fibers. Within the MLU, collagen fibers were closely associated with the smooth muscle cell nucleus.

[009] After a final elastin sheet, the external elastic lamina, comes the adventitial layer. The adventitia is largely a collagen fiber network, but also contains fibroblasts, elastin, and the vessels the supply blood to the arterial tissue, the vasa vasorum.

[010] Arterial structure and composition varies substantially throughout the vascular tree. The largest central vessels are elastic arteries, which tend to be more compliant, contain more elastin in the media, and a relatively small adventitia. The next group includes the carotid, coronary, iliac, and brachial arteries and others, known as the muscular arteries. These contain more smooth muscle cells in the media and a proportionally larger adventitia. The smallest arteries are the arterioles, which contain abundant smooth muscle, but less elastin, except for a thick internal elastic lamina. [011] The mechanics and structure of aortic valve leaflets. The three leaflets of the aortic valve are thin, semi-circular tissues. The thickness of the human aortic leaflet varies significantly, from the central belly region, which is approximately 0.250 mm, to the Node of Arantius, approximately 1.330 mm thick [16]. The functional implications of thickness are significant because the flexural rigidity of the leaflet is closely tied to thickness. The leaflet consists of three distinct tissue layers, the ventriculahs, spongiosa, and fibrosa. The central spongiosa is a loose, semifluid layer thought to facilitate shear deformation between the other two layers [17]. Because the spongiosa is largely water and glycosaminoglycans (GAGs), it cannot bear significant mechanical loads [18]. The fibrosa layer, termed the "backbone of the leaflet," contains some elastin [19], but consists predominantly of crimped collagen fibers running in a circumferential direction [17]. In addition, the fibrosa is textured with macroscopic wrinkles, or corrugations, so the aortic face of the leaflet is quite rough. The ventriculahs layer is thinner than the fibrosa, but contains significantly more elastin. The layer is not ridged or wrinkled, making the ventricular side of the leaflet much smoother than the aortic side. The layer is contiguous with the endocardial lining of the left ventricle [18].

[012] The leaflets stretch dramatically during normal function: surface area increases by up to 50% when the valve closes. The strain is anisotropic, with typical strains of 9% in the circumferential direction and 24% in the radial direction [18]. The large radial stretching during valve closure allows significant areas of contact, or coaptation, between neighboring leaflets. These areas can be up to 40% of total leaflet area [20]. When the valve is closed, the leaflets must withstand about 100 mm Hg of pressure, and coaptation areas play a stabilizing role. They also seal the valve against backflow. Importantly, initial radial stretching occurs with relatively little stress, but as stretching continues the leaflet stiffness sharply increases. This increase occurs at a level of strain known as the transition point. Higher stiffness prevents further deformation, stabilizing the leaflet against prolapse. When the valve opens, much of the leaflet strain is recovered. Recovery of strain diminishes leaflet area, opens a wide orifice for flow, and prevents the development of sharp folds or wrinkles in the leaflet. Therefore, it is accepted that leaflet stretch facilitates a highly stable position during valve closure, and recovery of stretch permits smooth, non-obstructive opening [20, 21]. In summary, leaflet anisotropy and the characteristic transition point strains are the salient features revealed in tensile tests of leaflets. In addition to this behavior in tension, bending behaviors and elasticity are important features of valve mechanics. [013] Many elements of leaflet mechanics may be traced to tissue microstructure. First, anisotropy is caused by the circumferential orientation of collagen fibers and the macroscopic wrinkles in the fibrosa. Collagen orientation makes the leaflet stiff in the circumferential direction and compliant in the radial direction, while the macroscopic wrinkles increase radial compliance because they are positioned to unfurl as the leaflet stretches in the radial direction [20]. Second, elasticity of the leaflet is mostly attributed to the elastin-rich ventriculahs [22]. In the unstressed leaflet, the fibrosa is actually prestressed in compression and the ventriculahs is in tension [23]. This observation strongly suggests that the venthcularis functions to elastically retract the fibrosa after the leaflet is extended. Third, transition point behavior is related to microstructural features such as the crimp of the collagen fibers. Crimp allows the fibers to stretch easily during initial deformation, but resist further deformation as they are pulled taut. In addition to the collagen crimp, realignment of the collagen fibers during deformation also contributes to the increased stiffness. Fourth, during flexure, the laminated structure of the leaflet is important. Slippage, or shear deformations, between layers is facilitated by the spongiosa, and prevents the accumulation of large compressive stresses near the concave surface as the leaflet bends [24].

[014] Networks of collagen and elastin proteins act in concert to bear leaflet stress and permit deformation. Elastin is highly compliant and capable of relatively large deformations. Therefore, it is associated with leaflet strain before the transition point, when substantial deformation occurs and the mechanical modulus is very low. Collagen is relatively stiff and inextensible, so it is associated with the stiffening of the leaflet [25]. Microscopic observations suggest that elastin fibers and sheets surround and connect the collagen fiber bundles [19]. The close association between elastin and collagen suggests elastin recruits collagen fiber bundles as the tissue approaches transition point strain. During recovery of deformation, the elastin network functions to recoil the collagen fiber network into its relaxed, slightly compressed, state [19]. Broom referred to this mechanism as a complementary system because the two fiber networks seem evolved to share stresses and transfer loading, generating highly specialized mechanical behavior. Broom also noted that collagen, like elastin, is an elastic material. Although elastin is probably mostly responsible for elastic recoil below the lock-up strain, collagen contributes elasticity at high strains [25]. These observations all confirm that, in addition to the mechanical properties of the constituent materials, the three-dimensional microstructure of the two fiber networks determines leaflet mechanical response. [015] Fibrillar collagen as a biomaterial. The collagens are a family of at least 19 proteins characterized by triple helical macromolecular structure and by the structural role they play in the extracellular matrix (ECM) [26]. The collagen types are most broadly categorized as fibril forming and non-fibril forming, and further differentiated by specific functions and tissue distributions. Collagens consist of three amino acid chains, the α-chains. To assume triple helical structure, the α-chains each coil in a left-handed helix and simultaneously coil about one another in a right-handed super helix. The α- chains may be identical or distinct, but generally consist of the GIy-X-Y repeated amino acid sequence, where X and Y are frequently hydroxylated lysine and proline residues. Many of the prolines and lysines are enzymatically hydroxylated after the α-chains have been transcribed but before formation of triple helical structure. Following hydroxylation and in some cases additional modifications, triple helical fibrillar collagen monomers are secreted into the extracellular space as procollagens. After secretion, propeptides regions are cleaved and the collagen monomers self-assembly into fibrils and are enzymatically crosslinked [26]. Type I collagen is the most abundant fibrillar collagen; often when the collagen type is not specified, type I is implied.

[016] The available methods to extract type I collagen from animal tissue result in different yields and alterations to the material at the fibrillar and macromolecular level. Fibrillar extraction methods mechanically separate tissue into fibril dispersions, removing non-collagenous components with the aid of selective proteolysis and washes [27]. Native assembly and crosslinking is not extensively disrupted as the insoluble fibrils are processed into suspensions. Other methods exist to isolate and purify collagen into solutions of monomeric collagen or small macromolecular aggregates. From some tissues, soluble collagen can be extracted at very low yield with dilute salt solutions. Higher yields are achieved with dilute acid solutions, which disrupt native aldimine crosslinks [27]. Following acid extraction, collagen may be separated from other tissue components by centrifugation and filtration. Sequential filtration stopping at the 0.45 μm level yields aggregates of an average of 2.85 molecules, while filtration to the 0.22 μm level yields monomeric collagen [28]. Solutions completely free of aggregates are difficult to obtain, and aggregates are usually present in the absence of 0.22 μm filtration. Alkali and enzymatic protocols more aggressively disrupt native crosslinks, cleave most of the non-helical ends of the collagen monomer, termed the telopeptides, and result in higher yields [27]. After removal of the telopeptide regions, the material is referred to as atelocollagen. Solutions of both acid-solublized and atelocollagen can be triggered to reassemble into fibrillar structures.

[017] In addition to the isolation of animal collagens, homologous, autologous, and recombinant human collagens are either in use or development. Homologous and autologous collagens are used clinically as for tissue augmentation, although the supply of both is inevitably limited [29]. Recombinant human collagens have been generated, and may be coexpressed with the enzyme prolyl 4-hydroxylase. This enzyme converts proline to 4-hydroxyproline, after which recombinant collagens have been shown to adopt triple helical structure and self-assemble into fibrils [30].

[018] Collagen has been extensively studied as a biomaterial and successfully employed clinically. A small percentage of patients will mount an immune response against bovine or porcine collagen implants, although the risk of adverse reaction can be minimized by screening for collagen allergy in advance [31]. Collagen is prepared for implant either by techniques that retain the fibrous architecture of the ECM or by protocols that purify collagen into a solution or dispersion and then process it into a physical form, such as an injectable, fiber, gel, sheet, membrane, or coating. Processes that retain the native ECM architecture include steps that destroy and remove the remnants of living cells, and crosslink and sterilize the remaining ECM. Many non- collagen components remain after processing. Example devices include bioprosthetic heart valves and small instestinal submucosa, a sheet used to repair and reinforce soft tissue. Preparation of an implant from collagen solutions or dispersions generally consists of precipitation, sterilization, and crosslinking. During precipitation, collagen monomers and aggregates can be triggered to self-assemble to approximate the fibrillar assembly in the ECM; alternatively solutions can be air dried, freeze dried, or otherwise deposited without fibrillar structure [32]. Capacity to recreate the complex and spatially varying density and orientations of native fibril networks is limited. Significantly, fibrillar structure determines the mechanical response of many tissues and in some cases may regulate cell proliferation, migration, and matrix synthesis [33].

[019] Elastin-mimetic protein thblock copolymers. Several elastin-mimetic protein polymers based on variations of the [VPGVG] pentapeptide repeat sequence of native elastin have been designed and evaluated. Results have been positive with regards to biostability in the absence of chemical crosslinks [3], nanofiber formation and tunable mechanical properties [4], drug-release and micelle formation [5, 6], and nonthrombogenic coatings [7].

[020] The properties of an elastin-mimetic protein polymer may be tailored by modifying of the (VPGVG) consensus repeat sequence, combining different repeat sequences, and/or introducing other bioactive peptide sequences. Adjustment of the repeat sequence alters the solubility and mechanical character. For example, the elastin pentapeptide displays temperature dependent solubility in water; it is extended and solvated at low temperatures but collapses and aggregates when warmed above the transition temperature (Tt). The value of 7>depends on the polarity of the sequence, and is usually adjusted by altering the identity of the fourth residue in the pentapeptide [2]. Increasing polarity of the fourth residue strengthens polymer solvent interactions, predictably increasing Tt. The mechanical character of the sequence has been adjusted from elastic to plastic by changing the third residue from glycine to alanine.

[021] Application of the block copolymer design framework has generated additional control over the properties of elastin-mimetics. Hydrophilic and hydrophobic protein polymer blocks were created from variations of the consensus pentapeptide repeat and combined as diblocks and triblocks. The 209 kDa thblock used in various examples disclosed herein, LysB10, has a midblock 28 with repeats of the elastic, hydrophilic (VPGAG)2VPGEG(VPGAG)2 sequence of pentapeptides. Both endblocks contain 170 repeats of the plastic, hydrophobic (IPAVG) sequence. The thblock has the capacity to form reversible virtual crosslinks and chemical crosslinks [3]. Dissolving the protein polymer in cool (4°C) water and warming it introduces virtual crosslinks as the hydrophobic endblocks self-associate as they approach room temperature. The midblock remains elastic and soluble at room and body temperatures. Consequently, self-association transforms concentrated solutions of the material into elastomeric gels. Chemical crosslinks may be introduced by modification of lysine residues that are available between the mid- and endblocks of LysB10 [3]. The addition of chemical crosslinks is expected to enhance biostability and has been shown to enhance strength and resistance to creep.

SUMMARY OF THE INVENTION

[022] The disclosure herein includes, inter alia, a composite material comprising a synthetic elastin-mimetic protein, and various fibers useful for incorporation into the synthetic elastin-mimetic proteins, that are biocompatible and useful for medical applications including as implantable devices. Further, the elastin mimetic proteins and/or fibers can have selectable physical characteristics so that the composite material (and specifically the medical devices/procedures comprising the composite material) may be tailored to better match the physical environment in which the materials are to be implanted. Also disclosed are a variety of related methods for making the composite materials and material components thereof, selectively tuning one or more physical characteristics of the materials, methods of casting the elastin mimetic protein into a film, methods of incorporating fibers such as collagen fibers or crimped fibers into the film, and methods of crimping fibers.

[023] The microstructure and mechanics of collagen and elastin protein fiber networks dictate the mechanical responses of all soft tissues and related organ systems. Disclosed herein, are materials and related methods employed to meet or exceed native tissue biomechanical properties through mimicry of these extracellular matrix components with synthetic collagen fiber and a recombinant elastin-like protein polymer. Significantly, this effort led to the development of a framework for the design and fabrication of protein-based tissue substitutes with enhanced strength, resilience, anisotropy, and more.

[024] We begin with developing a spinning process for scalable production of synthetic collagen fiber. Fiber having controlled cross-sectional geometry, size and physical characteristics are produced from an ultrafiltered collagen solution at relatively high production rates. The starting collagen concentration, flowrate, and needle size is optionally adjusted to control fiber size. Collagen fiber is characterized with mechanical analysis, micro-differential scanning calorimetry, transmission electron microscopy, second harmonic generation analysis, and subcutaneous murine implant.

[025] We subsequently describe the scalable, semi-automated fabrication of elastin- like protein sheets reinforced with synthetic collagen fibers that can be positioned in a precisely defined three-dimensional hierarchical pattern. Multilamellar, fiber-reinforced elastic protein sheets are constructed with controlled user-selected fiber orientation and volume fraction, depending on end application of the composite material. Structures are analyzed with scanning electron microscopy, transmission electron microscopy, and digital volumetric imaging. The effect of fiber orientation and volume fraction on Young's Modulus, yield stress, ultimate tensile stress, strain-to-failure, and resilience is evaluated in uniaxial tension. Increased fiber volume fraction and alignment with applied deformation significantly increased Young's Modulus, resilience, and yield stress.

[026] Highly extensible, elastic tissues display a functionally important mechanical transition from low to high modulus deformation at a strain dictated by the crimped microstructure of native collagen fiber. We report the fabrication of dense arrays of microcrimped synthetic collagen fiber embedded in elastin-like protein lamellae that mimic this aspect of tissue mechanics. Microcrimped fiber arrays are characterized with scanning electron microscopy, confocal laser scanning microscopy, and uniaxial tension analysis. Crimp wavelength is selected as desired, and examples are provided for a wavelength of 143 ± 5 μm. The degree of crimping is varied as desired, and examples are specifically provided from 3.1 % to 9.4%, and corresponded to mechanical modulus transitions at 4.6% and 13.3% strain. The crimping is maintained under repeated cycles of tensile loading. 50,000 cycles of tensile loading do not substantially alter microcrimp morphology.

[027] The composites are particularly useful in medical devices, such as implantable medical devices. One application relates to artificial blood vessles or small-diameter vascular grafts. Exemplified herein is a series of small-diameter vascular grafts comprising elastin-like protein reinforced with controlled volume fractions and orientations of collagen fiber. A pressure-diameter system is developed and implemented to further characterize the effects of fiber distribution on graft mechanics. By varying fiber characteristics, orientation, volume fraction, a desired design is achieved satisfying various target properties such as suture retention strength of 173 ± 4 g-f, burst strength of 1483 ± 143 mm Hg, and compliance of 5.1 ± 0.8 %/100 mm Hg.

[028] In an embodiment, the invention is a composite biomimetic material comprising a first material having an elastin-mimetic protein, wherein the first material is formed into an elastomehc film. Embedded in the film is a second fibrous material, such as a material that is a plurality of collagen fibers. In an aspect, the elastin-mimetic protein is selected from the group consisting of: LysB10 (SEQ ID NO:26), B10 (SEQ ID NO:9), R1 (SEQ ID NO:44), R2 (SEQ ID NO:46) and R4 (SEQ ID NO:34), whose sequences are provided in Table 11 , and various combinations thereof. In an aspect, the fibers are collagen fibers such as synthetic or animal-derived collagen that are continuous spun fibers that extend a length of the first material and are aligned in the first material in at least one preferential direction. The collagen fiber optionally has a conformational structure that mimics in vivo fibrillar collagen, and specifically a triple helical conformational structure. The fibers embedded or supported by the film can be described by one or more physical parameters, such as length, cross-sectional area and/or Young's modulus. In an embodiment, an individual fiber has a Young's modulus that is higher than the Young's modulus of the film. In an embodiment, each a plurality of collagen fibers are connected so that an "individual" fiber observed by eye is actually a plurality of collagen fibers.

[029] A continuous fiber is advantageous for use with the composites provided herein because they can span the entire length of the elastomehc first material film. Alternatively, the continuous fiber may be processed, such as oriented and/or cut into sections that are shorter than the entire length of the film. In an embodiment, at least one or all of the fibers traverse the film, such as the entire length, or if the fiber(s) is oriented, traverses from one edge to the other edge at a fiber angle. In an aspect, the fibers are evenly spaced with respect to each other. In another aspect, the spacing is not uniform, but instead varies such that there is spatially-varying distribution of collagen fibers with respect to position in the film. Such spatial variation is useful where it is desired for the composite material to have a spatially-varying physical parameter, such as Young's modulus. For example, if it is desired the material be stiff around a perimeter, more fiber may be located around the perimeter region or be oriented in a manner to provide increased stiffness. Similarly, if a material is desired to be highly elastic around the edges, but relatively stiff in a central region, the fibers may be provided at a higher density in the middle region. In this manner, any spatial distribution of a physical parameter, such as Young's modulus, may be achieved as desired.

[030] In another aspect, any of the composites described herein related to a fiber material that has, when the composite is unstrained, an undulating or a crimped geometry, such as undulations that are generated by crimping one or more of the fibers. This aspect is advantageous as the undulations provide additional capability of controlling mechanical behavior of the composites. For example, until the composite is sufficiently stretched to straighten the undulations in the fibers, the first material's elasticity dominates the composite's material properties. When, however, the stretch is sufficient to straighten the fiber's undulations, the larger Young's modulus of the fibers will dominate the composite's material properties. The undulations, therefore, better model and mimic behavior of various biological materials that are made from collagen and elastin constituents, such as blood vessels and heart tissue (leaflets, valves). [031] Although the disclosed composite materials can be used with any fiber, in a preferred embodiment the fiber is a collagen fiber such as a wet spun and banded fiber. Any of the methods for producing the wet spun collagen fiber disclosed herein may be used to generate collagen fiber for incorporation with the first material.

[032] In an aspect, an individual collagen fiber has a cross-sectional area that is selected from a range that is greater than or equal to 75 μm² and less than or equal to 8000 μm²; and a Young's modulus that is selected from a range that is greater than or equal to 400 MPa.

[033] In an embodiment, the composite biomimetic material has a fiber that is at least partially crimped. For example, a specific portion of the fiber may be processed to provide crimping, whereas other portions may be unchmped (e.g., central region crimped and end regions not crimped, or vice versa). The chmped/unchmped may be provided in a user-selected pattern along the longitudinal direction of the fiber. Alternatively, the entire length of the fiber may be provided with crimps.

[034] The crimping may be further described as having a wavelength and/or amplitude. In an embodiment, the crimped portion has a wavelength selected from a range that is greater than or equal to 50 μm and less than or equal to 1 mm, and amplitude selected from a range that is greater than or equal to 20 μm and less than or equal to 1 mm. Varying the amplitude and/or wavelength provides the ability to vary the amount of crimping or undulations, thereby affecting the transition strain of the composite where the modulus transitions from a low-modulus state to a high-modulus state.

[035] The undulating or crimped portion can have any number of geometric shapes, so long as there is defined strain at which the fiber unchmps. For example, the geometric shape can be described as spiral, helical, sinusoidal wave, sawtooth or ridged.

[036] In an embodiment the at least partially crimped portion has a crimp magnitude that is greater than or equal to 2%, such as between about 2% and 15%. In an aspect, the fiber is crimped to provide a composite biomimetic material having one or more of: a transition point strain selected from a range that is greater than or equal to 1 % and less than or equal to 20%; a compliant Young's modulus selected from a range that is less than or equal to 15 MPa; and a rigid Young's modulus selected from a range that is greater than or equal to 20 MPa.

[037] In an aspect, the composite biomimetic material relates to having one or more physical parameters that are anisotropic. For example, the Young's modulus may have a magnitude that depends on the direction of the applied stress based on the alignment direction the fibers. Optionally, the magnitude of a physical parameter may vary in a plurality of directions, such as by a second fibrous material that is aligned with respect to the first fibrous material, thereby providing two preferential directions defined by a fiber angle. In this embodiment, fiber angle refers to a relative angle between the fiber directions and, for a two preferentially aligned fiber system, the angle can be defined as greater than 0° and less than 180°. Alternatively, fiber angle may be defined relative to a direction or axis of the first material film.

[038] In another embodiment, the composite biomimetic material may be further described in terms of a fibrous material average spacing distance between adjacent fibers. Increasing this spacing distance results in an attendant decrease in fiber volume. In an aspect, the average spacing distance is selected from a range that is greater than or equal to 0.05 mm and less than or equal to 1 mm.

[039] Any of the composite biomimetic materials may be further described by one or more physical parameters such as suture retention strength, burst strength, compliance, Young's modulus, transition strain. In an embodiment, the medical device is a vascular graft having a suture retention strength between about 120 g-f and 400 g-f, mechanical modulus transition between about 2% and 15%, burst strength between about 1000 mm Hg and 2000 mm Hg, and compliance between about 3%/100 mmHg and 10%/100 mm Hg. In an embodiment, the composite biomimetic material is further described as having an ultimate tensile strength that is greater than or equal to 2 MPa and a strain to failure that is greater than or equal to 12%.

[040] The composite biomimetic materials described herein are particularly useful in medical devices. Accordingly, in an aspect the invention relates to any of the composite biomimetic materials described herein formed into a medical device. In an aspect, the medical device is a soft tissue patch, a dermal filler, a hernia patch, a valve leaflet or a vascular graft. In another aspect, the composite biomimetic material further comprises one or more of a drug, a growth factor, a polysacharride, a living cell, or a combination thereof supported by or connected to the first material, the second material, or both. [041] In an embodiment, the composite biomimetic material is formed into a specific shape, such as a sheet or a tubular cylinder. One advantage of the materials and processes described herein is that any size material may be made, including relatively large size composite materials that may be subsequently processed or cut into suitable sizes, such as useful implant sizes. In an aspect, the sheet or tubular cylinder has a length that is greater than or equal to 1 cm, and the fibers are continuous and the continuous fibers individually span at least 90% of the length of the sheet. In an aspect, the fibers individually span the entire length of the sheet.

[042] In another embodiment, the invention relates to a multilayer material comprising a plurality of layers, wherein each layer comprises any of the composite biomimetic materials described herein. Multilayer materials provide further control over the size and bulk mechanical property of the composite material. For example, a composite material having a specific fiber alignment may have multiple layers stacked in different directions to provide a bulk mechanical property distribution different than the anisotropy exhibited by a single layer. In an aspect, the number of layers in the multilayer is selected from a range that is greater than or equal to 2 and less than or equal to 100. In an aspect, the multilayer is laminated by a bottom surface layer and a top surface layer, wherein each of said bottom surface layer and top surface layer is the elastin-mimetic protein formed into a film without the second material.

[043] Optionally, the layers are bonded to adjacent layers, such as bonded by an adhesive and/or bonded by modulating the temperature of the material, especially by cooling the material to about 4°C and then warming to about 25°C. The sheet or tube may be subject to subsequent thermal treatment, such as the exposure to warm temperatures of 40-80⁰C for approximately 4 hrs to increase its strength. The sheet or tube may be chemically fixed, including for materials having crimped fibers, for example with a solution of glutaraldehyde or other crosslinker, to increase its strength and stability. For example, a multi-layer tube suitable for blood vessel replacement may be prepared by molding a flat sheet of about 10-200 μm, especially, 50-100 μm, of elastin- mimetic protein polymer, especially the material LysB10 or LysB10 mixed with other recombinant proteins, around a fibrous material such as a layout of collagen fiber. After molding, the sheet may be trimmed and wrapped about a central mandrel or tube. For example, the sheet may be trimmed to 8 x 5 cm, and wrapped about a Teflon tube with an outer diameter of 4 mm to create a wrap with approximately six layers. The layers may be bonded with an adhesive or by cooling the material to about 4°C and then warming to about 25°C. The tube may be thermally annealed at 60⁰C in PBS for 4 hrs and fixed with 0.5% glutaraldehyde in PBS for 24 hrs. This results in a tube with an inner diameter of approximately 4 mm, a wall thickness of approximately 800 μm, and desired bursting strength, suture retention strength, and compliance properties. Optionally, living cells or other biological materials are included with the composite material, such as between the layers.

[044] The tube or other shape material may be prepared with specific fiber layouts to generate the desired properties. For example, if the fibers are arranged so that they are loosely spaced (0.30 mm average distance between the fibers in a given layer of the tube) the bursting pressure of the tube will be lower than if the fibers are closely spaced (0.15 mm average distance between the fibers). More loosely spaced fibers also make the compliance of the tube under internal pressure higher. In addition, if the fibers are arranged at a small angle with the axis of the tube, for example 15°, the compliance will be higher but the burst pressure will be lower than if the fibers are arranged at a larger angle with the axis of the tube, for example 30°. In one examplary layout, the fibers are spaced at an average distance of 0.15 mm and oriented at an angle of 22.5° to the tube axis. This produces a tube with bursting strength of approximately 1500 mm-Hg, compliance of over 5%/100 mm-Hg, and suture retention strength over 170 grams-force.

[045] In another example, a sheet suitable for use as a surgical patch is prepared. In one preferable example, 50 μm layers of LysB10 may be molded about collagen fiber layouts or patterns by adhering the fiber to an ultrasoft sheet of polyurethane, covering the fibers with a solution of LysB10 in cold water, pressing a sheet of acrylic plastic over a solution to spread it into a thin, uniform layer. The layer may be gelled by warming to about 25°C. Several layers, for example 8 to 10, may be stacked, cooled to about 4°C, and then warmed to about 25°C. Lamination of ten layers forms a laminate that is approximately 0.4 mm thick, with individual layers that are well-bonded. Layers of LysB10 molded without fiber may be added in the same way as layers with fiber. Layers without fiber may be added as the outmost layers to prevent unraveling of the fibers from the inner layers of the laminate. The sheet may be prepared with specific fiber layout to generate the desired properties. For example, adding more fiber will increase the yield strength, stiffness, resilience, and suture retention strength of the sheet. Orienting the fiber in predominantly one direction will create a sheet that is stiff in that direction but compliant in the perpendicular direction. Similarly, spatially varying fiber spacing, orientation and/or crimp will provide correspondingly spatial variations in stiffness/elasticity.

[046] Any of the composites provided herein are used as a device for implantation in human or animal. For example, in tubular or cylindrical form, the composite may be used as an artificial blood vessel, tendon, or ligament. As a flat sheet or an assembly of flat sheets, the device may be an artificial heart valve, heart valve leaflet, wound dressing, or surgical patch such as a hernia patch.

[047] In an aspect, the composite biomimetic material has a thickness, and in particular a film thickness selected from a range that is greater than or equal to 30 μm and less than or equal to 1 mm.

[048] In an aspect, the second material has a volume fraction (relative to the volume of the composite) that is greater than or equal to 1 % and less than or equal to 30%. In an aspect, the fibers are uniformly distributed in said first material.

[049] In another embodiment, the invention is a composite biomimetic material comprising a first material comprising an elastin-mimetic protein formed into a film; and a second fibrous material comprising a plurality of continuous spun collagen fibers, wherein at least a portion of the continuous spun collagen fibers are crimped. The collagen fibers in an uncrimped state have a Young's modulus that is higher than the Young's modulus of the film and are aligned in at least one preferential direction. In contrast, the collagen fibers in a crimped state may not make a significant contribution to the composite's Young's modulus, but instead must uncrimp, such as at a strain that is greater than the strain transition, before significantly contributing to the composite material's Young's modulus.

[050] In an aspect, any of the composite biomimetic materials relate to a first material having LysB10 as the elastin-mimetic protein (WO 2008/033847 filed Sept. 11 , 2007).

[051] In another embodiment, the invention is a method of making any of the composite materials provided herein. In an aspect the method relates to providing a fibrous material on a first support surface and introducing a solution of the elastin- mimetic material over the fibrous material. In an aspect, the solution of the elastin- mimetic material is subsequently processed by any means known in the art to provide an elastic material or an elastomer, such as by gelling, crosslinking, polymerizing, or drying the solution to form the film having fibrous material embedded therein. The cross-linking to transition a solution to a gel may be by a thermally controlled sol-gel process. For example, the first material may be transitioned from a solution to a gel by changing the temperature, for example by increasing the temperature from about 4°C to about 25°C.

[052] To embed fibers in a flat sheet by casting, the fiber layout may be fastened to a flat supporting sheet, such as a sheet of plastic or glass, and a solution of matrix material may be poured over the fiber layout, and allowed to dry.

[053] To embed fibers in a flat sheet by molding, the fiber layout may be fastened to a flat supporting sheet, such as a sheet of plastic or glass, and a solution of matrix material may be poured over the fiber layout, and an additional flat sheet such as a sheet of plastic or glass, may be placed over the matrix material solution to press it into a thin layer. The thin layer may be transitioned from a solution to a gel by changing the temperature, for example by increasing the temperature from about 4°C to about 25°C.

[054] To embed fibers in a tube by molding, the fibers may be arranged about a shaft, for example by winding, and the shaft may be enclosed in a mold so that an annular void between the shaft and the mold is created. The void may be filled with a solution of matrix material, which may be transitioned from a solution to a gel by changing the temperature, for example by increasing the temperature from about 4°C to about 25°C.

[055] In another embodiment, any of the methods further comprise pressing the introduced solution of the elastin-mimetic material by a second support surface that faces the first support surface, wherein the first and second support surfaces are separated by the elastin-mimetic material and the fibrous material. In this manner, various composite material shapes may be directly made such as by first and second support surfaces having curved or shaped surfaces. For example, a first support surface that corresponds to a surface of a shaft or cylinder provides the capability of making hollow tubes of composite materials. A fibrous material at least partially wound around the shaft provides composite materials shaped as a hollow tube, such as would be suitable as a vascular graft or artificial blood vessel. Alternatively, arbitrary shapes may be made from a single large surface area composite material. For example, artificial blood vessels may be made by wrapping a sheet of a composite material about a cylindrically-shaped object of a desired diameter. The wrapping may comprise multiple wrappings, thereby providing a multilayer composite material. [056] Methods provided herein are particularly amenable for additional processing to select a desired physical parameter for the composite material. For example, a physical parameter such as stiffness, strength, or Young's modulus may be selectively adjusted by varying water absorbency of the first support surface. The surface may be porous and absorbent to water, such as an ultrasoft polyurethane, or alternatively a non- absorbent material such as polycarbonate or glass. An absorbent mold will increase the stiffness and strength when molding an elastin-mimetic matrix material, potentially by lowering the water content during molding. A non-absorbent mold will decrease strength but increase elasticity and compliance, potentially by retaining the initial water content during molding. The water absorbency is optionally varied by adjusting the porosity of the first support surface.

[057] In another aspect, the invention relates to methods of generating crimps in the fibrous material. In one embodiment, the crimps are generated by providing a first stretchable sheet at a first level of strain, stretching the first stretchable sheet to a second level of strain that is greater than the first level of strain, attaching the fibrous material to the stretchable sheet at the second level of strain, and relaxing the stretchable sheet to which the fibrous material is attached to a third level of strain that is less than the second level of strain, thereby generating crimps in the fibrous material.

[058] In an embodiment, the fibrous material is attached to a second stretchable sheet that faces the first stretchable sheet, wherein the fibrous material is positioned between the first and second stretchable sheets.

[059] In an embodiment, the first stretchable sheet, the second stretchable sheet, or both, have a contact surface that has a pattern of relief features that provides a plurality of contact points with the fibrous material, wherein the fibrous material remains fixed in position relative to the contact points during a change in strain, and portions of the fibrous material between the contact points are undulated. Optionally, portions of the fibers are reversibly bonded to the stretchable sheet to provide a pattern of attachment sites, thereby providing undulations in the fiber between the attachment sites.

[060] In an embodiment, the pattern of relief features generates a surface shape on the stretchable sheet that is a sinusoidal wave, rounded ridges sawtooth and chamfered rectangular. In an aspect, the relief feature is a micro-sized feature having at least one dimension less than 1 mm. In an aspect, at least one dimension is less than or equal to 300 μm. The contact surface optionally has grooves having a receiving volume to receive deformed fibers into the grooves.

[061] In an aspect, the crimped fibers are fixed, such as by chemical fixing, cross- linking or temperature fixing. As used herein, "fixed" refers to the functionality of stabilizing crimps or undulations, without adversely affecting the capability of fibers to uncrimp under an applied stress, and recrimp when the applied stress is removed.

[062] In an aspect, any of the collagen fibers used in any of the composite materials have, in an uncrimped state, a Young's modulus that is at least ten times greater than the Young's modulus of a film made from the first elastin-mimetic protein containing material.

[063] In another embodiment, the invention is a method of making a collagen fiber by providing a collagen material in solution and extruding the collagen-containing solution into a wet spinning buffer at an extrusion flow-rate to form a gel fiber. The gel fiber is passed through a rinse bath and the rinsed gel fiber passed through a dryer to provide a dry collagen fiber. The dried fiber is continuously collected and incubated in an incubation bath. After incubation the fiber is rinsed and dried. The collagen-containing solution is, in an aspect, monomeric collagen having a concentration that is greater than 4 mg/mL. The extrusion flow-rate, in an aspect, is greater than or equal to 0.5 mL/min.

[064] In an embodiment, the rinsed incubated and collected fiber drying step is performed with the fiber under tension.

[065] In an embodiment, a material, such as a material that can affect a biological outcome is introduced to the collagen fiber. In an aspect, the collagen material in solution and/or the incubation bath further comprises one or more of such a material, including a growth factor, a drug, a protein or a polysaccharide.

[066] In an aspect, any of the methods or compositions provided herein are directed to a continuous fiber having a length that is greater than or equal to 1 m.

[067] In another embodiment, the invention is a method of generating crimps in a biomimetic fibrous material, such as collagen fibers, by providing a first stretchable sheet at a first level of strain and attaching a biomimetic fibrous material to the stretchable sheet at the first level of strain. The attachment is by any means known in the art, such as by clamping or bonding. The stretchable sheet to which the fibrous material is attached is stretched to a second level of strain that is greater than the first level of strain and then the stretchable sheet to which the fibrous material is attached is relaxed to a third level of strain that is less than the second level of strain, thereby generating crimps in the biomimetic fibrous material.

[068] In an aspect the fibrous material comprises collagen fibers, such as a plurality of aligned fibers having a length that is greater than or equal to 10 cm.

[069] In another embodiment the biomimetic fibrous material is attached to a second stretchable sheet that faces the first stretchable sheet, wherein the fibrous material is positioned between the first and second stretchable sheets and the second stretchable sheet is correspondingly strained to the first level, the second level and the third level of the first stretchable sheet. For example, the fibers may be clamped between the two stretchable sheets. In an aspect, the first and third levels of strains are approximately 0% (e.g., no applied stress) and the second level of strain is selected from a range that is between about 3% and 50%, or between about 5% and 30%. In an aspect, the third level of strain may have some residual strain, even without any applied force, arising from the interaction of the stiffer fibers with the elastic sheets.

[070] In an aspect, the clamping or attachment may be accomplished by providing a shaped surface on one of the elastic sheets that when clamped to the opposing elastic sheet, provides a fixed connection between the sheet and the fibers. In this manner the relaxation in strain of the elastic sheets generates crimps or undulations in the fibers. For example, the first stretchable sheet, the second stretchable sheet, or both, have a contact surface that has a pattern of relief features that provides a plurality of contact points with the fibrous material, wherein the fibrous material remains fixed in position relative to the contact point during changes in strain. In an aspect, the plurality of contact points provide crimping having a wavelength selected from a range that is greater than or equal to 50 μm and less than or equal to 1 mm, and an amplitude that is selected from a range that is greater than or equal to 50 μm and less than or equal to 1 mm. In an aspect, the relief features are made of a material that is sufficiently hard to prevent unwanted mechanical deformations during changes in the applied stress and corresponding strain, such as from a polyurethane elastomer or other elastomer with durometer of about 7OA or more.

[071] In an embodiment, the pattern of ridges and grooves is microscopic in scale. For example, the height of the ridges may be about 100 μm, and the spacing between the ridges may be about 100-200 μm. In an embodiment, the ridges are made from a polyurethane elastomer with a durometer of 7OA, spaced at 100-200 μm, with a width at the base of the ridge of 50-100 μm, a width at the peak of 10-60 μm, and a height of 40- 150 μm.

[072] The fibers are optionally made from a biologically compatible material, such as a biopolymer or from collagen, a blend containing collagen, or a recombinant collagen. The fibers are optionally softened by treating them with a plasticizer or exposing them to an elevated temperature before relaxing the strain from the sheets of elastic material to deform the clamped fiber or fibers. Collagen fibers are optionally softened by hydrating with water or other aqueous solution. The fibers may be chemically fixed or crosslinked, for example collagen fibers may be fixed with glutaraldehyde or other crosslinking agents. The fibers may be cooled or frozen. The plasticizer may be removed, for example collagen fibers may be dried. Also provided are sheets of manufactured fiber, such as fiber comprised of collagen or other biopolymer, shaped into periodic waves or crimps on the microscopic scale.

[073] In an aspect, the method further comprises embedding the crimped biomimetic fibrous material into an elastic material having a Young's modulus that is at least ten times less than the biomimetic fibrous material Young's modulus in an unchmped configuration. In an aspect, the elastic material is an elastin-mimetic material, such as an elastin-mimetic material formed from any one or more of the elastin-mimetic proteins disclosed herein.

[074] In another embodiment, the invention is a method of making continuous collagen fibers, such as by providing a collagen material in solution, extruding the collagen solution into a wet spinning buffer at an extrusion flow-rate to form a gel fiber and passing the gel fiber through a rinse bath. The rinsed gel fiber is passed through a dryer to provide a dry collagen fiber which is continuously collected. The collected dry fiber is incubated in an incubation bath, and then rinsed and dried to provide a continuous collagen fiber having a triple helical native conformational structure.

[075] For example a solution of monomeric collagen in 10 mM HCI at a concentration of about 5 mg/ml may be extruded into a wet spinning buffer (WSB) of 10 wt% poly (ethylene glycol) M_w = 35000, 4.14 mg/ml monobasic sodium phosphate, 12.1 mg/ml dibasic sodium phosphate, 6.86 mg/ml TES sodium salt, 7.89 mg/ml sodium chloride, with pH adjusted to 8.0. In one embodiment, the collagen solution may be extruded through a needle into a length of fluoropolymer tubing that contains the WSB. The WSB may be flowing, for example at a rate of 1 ml/min. The gel fiber from the WSB may be passed through a bath to rinse it, for example a bath of 70% ethanol or isopropanol in water that is 2 meters long. The rinsed gel fiber can be passed through air to dry it. The rinsed and dried fiber is continuously collected until a significant length is created. For example the fiber may be collected on a rotating frame or cylinder and transferring a length of the collected fiber into an incubation bath for a period, for example by leaving the fiber on the frame or cylinder and immersing it in the incubation bath. The length of fiber may be rinsed, for example in water for 5 min, and passing the incubated, rinsed fiber through air to dry it and collect it.

[076] The incubation step is formulated to drive the assembly of collagen molecules or aggregates into banded collagen fibrils. For example, the incubation may comprise a buffer known as fiber incubation buffer (FIB). FIB may comprise 7.89 mg/ml sodium chloride, 4.26 mg/ml dibasic sodium phosphate, 10 mM Tris, dissolved in water with the pH adjusted to 7.4 and may last for 48 hrs.

[077] The final drying step is optionally performed with the fiber under tension, to enhance the alignment of banded collagen fibrils within the fiber.

[078] In an aspect, prior to extrusion in wet spinning buffer, a chilled (4°C) solution of monomehc collagen or oligomeric collagen multimers is co-eluted with a chilled carrier solution at 4°C, such that the final concentration of the collagen solution is consistent with the composition of FIB buffer, as described above. The final collagen solution in FIB buffer is gradually heated to 37°C during a period of flow at a defined shear rate prior to extrusion in wet spinning buffer.

[079] The starting collagen material in solution may be obtained from any number of sources, such as a solution of collagen fibrils generated from homogenized collagen gels or other processes known in the art. The solution of collagen fibrils is pumped at a defined shear rate prior to extrusion in wet spinning buffer. In any of the methods or compositions provided herein, growth factor(s), drug(s), protein(s), or polysaccharide(s) are added to the solution of collagen monomers, multimers or fibrils prior to extrusion in wet spinning buffer and/or to the collagen fibers prior to or with the incubation step. [080] In an aspect, the elastin-mimetic component of the composite comprises a triblock protein copolymer having hydrophobic end block regions separated by a hydrophilic center block. In various aspects of the invention, chemical cross-linking sites are provided for further tuning of the material's physical parameters, such as by selective incorporation of lysine residues in the protein. In addition, manipulation of the center and end block regions (relative to each other) provides another mechanism for tuning one or more physical parameters. For example, the respective lengths and/or the hydrophobicity/hydrophilicity are increased or decreased to alter a physical parameter. In an embodiment, the invention is a triblock protein copolymer A-B-C, where the end blocks A and C are hydrophobic and the central block B is hydrophilic. In an embodiment, the central block provides elasticity to the protein, and the end block provides plasticity to the protein, thereby providing elastin mimetic characteristics.

[081] In an embodiment, the first material is a synthetic protein copolymer triblock comprising end hydrophobic blocks (SEQ ID NO:23 and/or SEQ ID NO:24) separated by a central hydrophilic block, with a plurality of cross-linkable sites (SEQ ID NO:25), for example the protein having the sequence of LysB10 (SEQ ID NO:26):

[VPAVGKVPAVG(IPAVG )₄][(IPAVG) ₅]s3 -X- [VPAVGKAAKVPGAGVPAVG(IPAVG )₄][(IPAVG) ₅]₃₃ [IPAVGKAAKA]

wherein X is (SEQ ID NO:25) IPAVGKAAKVPGAG][(VPGAG)2VPGEG(VPGAG)2]28

[082] LysB10 is a particularly suitable protein component for the first material in the composite biomimetic material because the lysine units (K) are available for cross- linking of the elastin triblock. Accordingly, in an aspect any of the materials or methods disclosed herein relate to a first material wherein the elastin-mimetic protein is LysB10 (SEQ ID NO:26) that is cross-linked, thereby forming a film in which the fibrous material, such as a plurality of collagen fibers, is suspended.

[083] In another embodiment, the invention is an isolated and purified nucleic acid sequence, that encodes for any one or more of the first endblock (SEQ ID NO:23), the second endblock (SEQ ID NO:24), the central block (SEQ ID NO:26), repeated any number of times as desired, such as from about 10 to 50, or about 28 as exemplified, or the protein LysB10 (SEQ ID NOs:26 or SEQ ID NO:33), and mixtures of any of the endblocks and central blocks as disclosed herein repeated any number of times to form copolymers having more than 3 blocks. [084] In an embodiment, the invention is a synthetic protein triblock copolymer comprising first and second end hydrophobic blocks separated by a central hydrophilic block, wherein:

the central block comprises the sequence:

(IPGAG)(VPGAG)VPGEG(VPGAG)a[(VPGAG)_bVPGEG(VPGAG)c]d

the first and second end blocks each independently comprise the sequence:

[VPAVG(I PAVG)_x] [(I PAVG)_y]z

and wherein: a has a value from about 1 to about 10; b has a value from about 1 to about 10; c has a value from about 1 to about 10; d has a value from about 10 to about 50; x has a value from about 1 to about 10; y has a value from about 1 to about 10; and z has a value from about 20 to about 100.

[085] Optionally, the blocks are separated by one or more residues capable of facilitating cross-linking, such as lysine residues.

[086] The first and second endblocks of any of the proteins provided herein have the same amino acid sequence or have a different amino acid sequence.

[087] In an embodiment, at least one the first and second endblocks of the protein comprises the sequence (SEQ ID NO:6, which itself is made from a plurality of 5-mers from SEQ ID NOs:4-5):

[VPAVG(IPAVG)₄]KIPAVG) ₅]₃3

[088] In an embodiment, the central block of any of the proteins provided herein comprise the sequence (SEQ ID NO:7, which itself is made from a plurality of 5-mers from SEQ ID NOs:1 -3):

(IPGAG)(VPGAG)VPGEG(VPGAG)₂ [(VPGAG )₂VPGEG(VPGAG)₂]₂o

[089] In an embodiment, the protein triblock copolymer comprises the sequence of B10 (SEQ ID NO:9):

[VPAVG(IPAVG M(IPAVG) ₅]₃3 - X - [VPAVG(IPAVG M(IPAVG) ₅]₃3

wherein X = (IPGAG)(VPGAG)VPGEG(VPGAG)₂ [(VPGAG)₂VPGEG(VPGAG)₂]₂o [090] In an aspect, any of the proteins disclosed herein are further characterized in terms of the relative lengths of the endblocks to the central block. For example, the protein is described as having an end block length parameter corresponding to the total number of amino acids in the first and second end blocks, and a central block length parameter corresponding to the number of amino acids in the central block. In this aspect, a ratio of the end block length parameter to the central block length parameter has a selected value, wherein the ratio has a value that is about 1 , greater than 1 , greater than 1.5, from about 1 :1 to about 10:1 , or about 2:1 to about 10:1.

[091] In another aspect, any of the proteins are described in terms of the amount of isoleucine, such as a mole fraction of isoleucine of greater than about 18%, between about 18% to about 25%, or about 20%.

[092] In an embodiment, any of the proteins are hydrated. Such hydration provides the capacity of at least one of the end hydrophobic blocks to form physical crosslinks that provide improved mechanical stability under sustained or repeated mechanical loading such as, for example, the sustained repeated load experienced by the blood vessel wall, a tissue, or an organ in a living system.

[093] In an embodiment, any of the proteins are described in terms of any one or more of a physical parameter. In an aspect of this embodiment, any of the proteins have an inverse transition temperature, such as a transition temperature that is between about 15⁰C and about 27⁰C, or selected from a range that is between about 19⁰C and about 23⁰C.

[094] In another embodiment, the invention is a hydrated film or fiber network comprising any of the proteins disclosed herein. Optionally, the film or fiber network is cast from a solution comprising TFE or water, such as by electrospinning, and the film or fiber network has a cast temperature. The cast temperature may be of any value so long as suitable elastin-mimetic materials having suitable mechanical properties are obtained, such as a cast temperature selected from a range that is between about 2°C and about 35°C. In an aspect, any of these films or fiber networks is formed into a tissue engineering scaffold capable of supporting cell growth. A useful property of the proteins disclosed herein is their capacity of having a user-selected physical parameter by selection of appropriate amino acids, amino acid sequences and amino acid configurations. For example, the film or fiber network of any of the proteins optionally have a tunable physical parameter, such as a physical parameter that is a: Young's modulus; an ultimate tensile stress; strain at failure; resilience; and creep resistance. Of course, any of the materials described herein may be subject to any one or more postprocessing techniques known in the art to further effect a change in one or more physical parameters (e.g., post-processing that changes porosity), or may be incorporated with fibers such as collagen fibers to further affect a change in the physical parameter.

[095] The ability to tune one or more physical property parameters of the film or fiber network that is made from any of the disclosed proteins provides the capability of tailoring the material to a particular application. For example, any of the films or fiber networks is formed into a medical device that may be implanted into the body, such as a vascular graft. Depending on the location of the vascular graft, however, the desired mechanical properties can be very different. Some applications may require resistance to high loads, other low lows, and others a repeated cycling of loads. An embodiment of the present invention provides the ability to tune any one or more of these parameters by varying one or more of end block to central block length, end block hydrophobicity, center block hydrophilicity, degree of cross-linking, and fiber orientation and geometry.

[096] In an embodiment, the invention is a medical device comprising any of the proteins provided herein, such as LysB10, B9, B10, R1 , R2 or R4, and films thereof operably connected to a fiber or fiber network, such as collagen fibers, including crimped fibers, embedded in the film. Examples of medical devices of particular utility include, but are not limited to, an artificial blood vessel; a stent; a graft; a wound dressing; an embolic agent; and a drug delivery device. Any of the medical devices may have a protein, film, or fiber network comprising a protein of the present invention that at least partially coats one or more surfaces of the medical device. In an aspect the protein, film, or fiber network of the medical device retains physical integrity under sustained mechanical load.

[097] In another embodiment, the film or fiber network has a cast temperature is greater than the inverse transition temperature. In an embodiment, any of the proteins comprise one or more chemical cross-linking sites flanking each block. "Chemical cross-linking" refers to covalent interactions, van der Waals interactions, dipole-diople interactions and/or hydrogen bonding interactions within the proteins that provide the capability of effecting a measurable change in one or more physical parameters, and is different from the "physical cross-linking" arising from the physical interaction of hydrophobic and hydrophilic regions which causes conformational changes. In an embodiment, the chemical cross-linking site comprises an amino acid that is lysine. Lysine can be suitably processed to mediate chemical cross-linking, such as by gluteraldehyde or a photocross-linkable acrylate functionalized lysine.

[098] In another embodiment, the invention is nucleic acid sequence that encodes the any one or more of the first endblock, the second endblock (SEQ ID NO:14), the central block (SEQ ID NO:15) and/or any of the proteins disclosed herein.

[099] In an embodiment, the nucleic acid sequence encodes the protein having the amino acid sequence of B10 (SEQ ID NOs:9-10), or any blocks thereof (DNA cross- referenced as SEQ ID NOs:11 -17,19 or repeating combinations thereof).

[0100] In an embodiment, the invention is a synthetic protein copolymer thblock having a plurality of chemically cross-linkable sites, such as the protein of SEQ ID NO:33 or:

K[(IPAVG)₅]₂₆-KK[(VPGAG)₄(VPGEG)]₂₆KK-[(I PAVG)₅J₂₆ KK

[0101] In an aspect, the invention is a synthetic protein copolymer thblock comprising end hydrophobic blocks separated by a central hydrophilic block, said protein comprising the sequence of R4 or SEQ ID NO:34:

VPAVGKVPAVG[(IPAVG)₅]i6 (IPAVGIPAVG)KAAK(VPGAGVPGIG) [(VPGIG)₅]i5 (VPGIGVPAVG)KAAK(VPGAGVPAVG) [(IPAVG)₅]i6 IPAVGVPAVGKAAKA

[0102] In another embodiment, the invention is an isolated and purified nucleic acid sequence, the sequence encoding for any one or more of the first endblock, the second endblock, the central block and/or the entire R4 protein, such as the nucleic acid sequence of SEQ ID NO:42.

[0103] In another embodiment, the invention is a peptide capable of establishing elastic-like behavior when incorporated into an elastin-mimetic protein, such as a peptide comprising the sequence R1 :

K_a[(VPGIG)_b]_cK_d

Wherein a has a value from about 1 to about 5; b has a value from about 1 to about 10; c has a value from about 5 to about 50; d has a value from about 1 to about 5.

[0104] In an aspect R1 has the amino acid sequence of SEQ ID NO:44: K[(VPGIG)₅]i₅KK

[0105] In an embodiment, the invention is a peptide capable of establishing plastic-like behavior when incorporated into an elastin-mimetic protein, such as a peptide comprising the sequence of R2:

K_a[(IPAVG)_b]_cK_d

[0106] In an aspect, R2 has the amino acid sequence of SEQ ID NO:46:

K[(IPAVG)₅]i₆KK

[0107] In another embodiment, the invention comprises a multi-block elastin mimetic protein having the formula:

R2-R1 -R2 or (R2-R1 )_n ;

R1 and R2 are as defined above and wherein n is greater than or equal to 2, or is selected from a range that is between 2 and 10

[0108] In an aspect, R1 comprises the sequence of SEQ ID NO:44 and R2 comprises the sequence of SEQ ID NO:46:

([(IPAVG)₅]i6)-KK[(VPGIG)₅]i5KK-([(IPAVG)₅]i6)KK

[0109] In an embodiment, the invention is a medical device, cell, tissue, or organ comprising any one or more of the proteins disclosed herein, such as any one or more of B9 (SEQ ID NO:50), B10 (SEQ ID NOs:9,26, 33), R1 (SEQ ID NO:44), R2 (SEQ ID NO:46), or R4 (SEQ ID NO:34), any combinations thereof, or spun fiber or fiber networks thereof. In an embodiment, the protein is one or more of B10, R1, R2, or R4. One example of a medical device is a vascular graft, such as a shunt. The graft or shunt optionally comprises a base scaffold material that is coated and/or impregnated with any one or more of the proteins or films and/or fiber networks thereof. One example is a shunt that is made of ePTFE. In an aspect, the coating is a multi-layer coating. In an embodiment, the medical device comprises a woven collagen graft. [0110] In another embodiment, the invention is an embolic agent, wherein the embolic agent comprises one or more of the proteins of the present invention, such as any one or more of the amino acid sequences in Table 11 alone or in combination with each other, or SEQ ID NOs:9, 10, 26, 33, 34, 44, 46, 47, 48, 50, B9, B10, R1 , R2, R4, or a blend thereof. In an aspect, the embolic agent has an inverse transition temperature, said temperature selected from a range that is between about 19⁰C and about 23⁰C. Such an inverse temperature may be used to readily administer the embolic agent in a liquid form, and upon administration, the embolic agent gels or solidifies.

[0111] In an embodiment, the invention is a method of applying an embolic agent to a patient in need of an embolic agent by providing an embolic agent, wherein the embolic agent is any of the proteins disclosed herein, such as B9, B10, R1 , R2, R4 or mixtures thereof. The embolic agent is applied to the patient. In an aspect, the embolic agent is applied in a solid or a gel form. Alternatively, the embolic agent is injectable and has an inverse phase transition temperature that is less than the environment in which the agent is applied, so that upon or after application said embolic agent undergoes a phase transition from liquid to a gel or solid form. In an aspect, the patient in need suffers from a cardiovascular defect. One example of such a defect is a neurovascular aneurysm.

[0112] In another embodiment, the invention is a method of producing a fiber network having improved mechanical properties from a triblock copolymer of any of the proteins provided herein, or any mixture thereof, along with fibers such as collagen fibers that are optionally crimped. Examples of specific triblock copolymers are amino acid sequences selected from the group consisting of LysB10, B10, B9, R1 , R2, R1 -R2, R4. In an aspect, the method improves a mechanical property that is an elastic modulus, and the elastic modulus increases by at least 30% compared to a nonannealed fabric. In an aspect, the annealing temperature is greater than 50⁰C. In another aspect, the method of annealing generates a decrease in water swelling ratio, selected from a range that is between 30% and 70%, or about 50%. Optionally, the method further comprises preconditioning the fiber network by repeated stress-relaxation cycling. In an aspect, the number of repeats is less than 10, such as between the range of about 4 and about 8.

[0113] Without wishing to be bound by any particular theory, there can be discussion herein of beliefs or understandings of underlying principles or mechanisms relating to embodiments of the invention. It is recognized that regardless of the ultimate correctness of any explanation or hypothesis, an embodiment of the invention can nonetheless be operative and useful.

BRIEF DESCRIPTION OF THE DRAWINGS

[0114] Figure 1. Wet spinning system for making collagen fibers. A syringe pump extruded WSB (i) though a bubble trap (iv) and into a coagulation column (v). The pump also drove the flow of the collagen solution (ii) through a needle (iii) and into the column. As the collagen stream emerged from the needle, it aggregated into a gel-like fiber due to the surrounding WSB. Flowing WSB carried the collagen fiber down the column, into the 70% ethanol rinse (vi). Short fiber segments were collected from the rinse with a hand-operated frame (vii). An automated roller system (viii) was installed to collect 30 to 60 m of continuous fiber.

[0115] Figure 2. System for continuous fiber collection and drying. After FIB incubation, pipe segments were rinsed in ddH20 for 15 minutes and transferred to the collection and drying system. The pipe segment was partly immersed in 70% ethanol and the AC fiber was transferred through air to a second pipe segment. The fiber dried under tension as it traveled between the two pipe segments.

[0116] Figure 3. Micro-differential scanning calohmetry of MRTC (long dash), AC-FIB (solid), AC+FIB (dotted), RTT (dash-dot), and AC+FIB+GLUT (short dash). Representative results from three experiments are shown.

[0117] Figure 4. Selected comparisons of UTS and major diameter of crosslinked MC with varied wet spinning parameters. All differences in mean diameter resulting from altered spinning parameters are statistically significant (p<0.05), while all differences in UTS are not. Error bars represent standard deviations.

[0118] Figure 5. Representative stress-strain data for automatically collected fiber.

[0119] Figure 6. Incubation in FIB results in the assembly of collagen fibrils. Banded collagen fibrils are not visible in axial sections of the untreated fiber (A, B), but are broadly evident following incubation (C, D).

[0120] Figure 7. Mechanical annealing during incubation enhances fibril alignment. Lower magnification axial sections of MC-0%, MC-15%, and MC-30% (left column) showed a trend toward enhanced fibril alignment in the stretched samples (middle and bottom rows). However, alignment is not uniform: in high magnification images of the samples, regions of low fibril alignment (center column) and high alignment (right column) can be identified.

[0121] Figure 8. . Fibrillar structure of AC imaged by TEM. Axial sections reveal an aligned pattern of fibrils, often displaying banding (A). Fiber cross sections comprise tightly packed fibril cross sections (B, C).

[0122] Figure 9. The second harmonic generated by FIB treated manually collected fiber and rat-tail tendon. A cluster of three wet spun fibers displayed a clear SHG signal but only short discontinuous fibrillar substructure was noted (A). The signal from rat-tail tendon revealed fibrillar structure (B). Scale bars are 20 μm.

[0123] Figure 10. Morphology of explanted fiber bundles after 6 weeks. Upper three panels show crosslinked bundles while the lower panels are uncrosslinked bundles. Sections are stained with Gomori Trichrome (A, B, D, E) or HE (E, F). Original magnifications were 10X (A, D) or 4OX (B, C, E, F).

[0124] Figure 11. Macrophage distribution in crosslinked and uncrosslinked fiber bundles. Macrophages are present inside the crosslinked bundles of fiber (A, 10x, and B, 4Ox) but collected primarily around the perimeter of the uncrosslinked bundles (C, 10x, and D, 4Ox).

[0125] Figure 12. A process for embedding a fiber layout in an elastin-like protein matrix. Fiber is wound about rectangular frames to obtain the desired fiber orientation angle, θ, and average spacing, ω (A). The fiber layouts are transferred to sheets of ultrasoft polyurethane (B), and the LysB10 solution is distributed over the layout (C). An acrylic sheet is placed over the layout to spread the LysB10 solution into a film reinforced with the fiber layout (D).

[0126] Figure 13. Fiber is reacted in a co-axial pipe system (A) and dried by transferring it through air to a second roller (B).

[0127] Figure 14. Fiber composite sheet geometry. Fibers made an angle of ± θ with the x direction of the sheet. The y direction is across the width of the sheet, and the z direction is through the sheet thickness. Fiber spacing, ω, is measured in the y direction. [0128] Figure 15. Representative plot from fiber orientation analysis of a layout with a fiber orientation of 15° and volume fraction of 7%. Intensity peaks at -12.0 and 15.6° correspond to the nominal fiber orientations of ±15°.

[0129] Figure 16. Fiber layout reconstructed by digital volumetric imaging. The x and y directions are as indicated. Scale bar is 500 μm.

[0130] Figure 17. Transmission electron microscopy of a multilamellar sheet. Stain localized in irregularly shaped 50 - 200 nm areas that speckled the LysB10 matrix. Z sections of the composite displayed wet spun collagen fibers with a banded fibrillar structure, generally aligned with the overall fiber axis (A, B). In the x section views of the sheet, the collagen fiber comprises densely packed fibrils in cross section (C, D).

[0131] Figure 18. Scanning electron microscopy of a multilamellar sheet. As illustrated, the fiber cross-section will appear circular or elliptical when the sheet is sectioned along the x- or y-plane, respectively (A). Synthetic collagen fibers can be visualized within a cross-section through the x-plane of the composite sheet (B, C). The exterior of the fibers display a fibhllated texture (arrows, D). Fibers protruding within a cross-section through the y-plane appear elliptical due to the oblique angle they made with the section plane. Sectioning artifacts appeared in (C) as vertical microgrooves and in (E) as feathered horizontal ridges in the protein polymer. The number and spacing of the ridges indicated that they did not correspond directly to the lamellar interfaces. Some fibers appear beneath the z-plane of the sample, and in rare instances, protrude through the surface (arrows, F).

[0132] Figure 19. Uniaxial mechanical response of a composite sheet with (A) or without collagen fiber (B) over a period of cyclic loading to 8% strain. Collagen fiber was oriented at 15° and a volume fraction of 17% in (A). Between the first loading cycle (•) and sixteenth loading cycle (o), the material became less stiff, and 1 -2% residual strain was introduced. The difference between intermediate cycles (— ) diminished as the cycle number increased.

[0133] Figure 20. Stress-strain response of fiber composite sheets of varying fiber orientation, but with fixed fiber volume fraction. Composites with fiber angles of 0° (•), 15° (---), 90° (— ), and without fiber (o) were tested at low and high strains (A, B).

Photographs of collagen fiber layouts for 0, 15, and 90° layouts are shown in C, D, and E respectively. Scale bars are 2 mm. [0134] Figure 21. Stress-strain response at varying fiber volume fraction. Composites with average fiber fractions of 17% (•), 7% (— ), and 3% (o), and without fiber ( — ) are tested at low and high strains (A, B). Photographs of collagen fiber layouts are taken after staining (C, D, and E). Scale bars are 2 mm.

[0135] Figure 22. Dependence of resilience on fiber volume and orientation.

Resilience increased with increasing fiber fraction (A). Layouts with fibers closer to the loading direction, with orientations of 15° and 0°, are more resilient than layouts with a fiber orientation of 90°, perpendicular to the loading direction (B). Bars without letter labels are significantly different from all other bars (p<0.05). Bars with the same letter are not statistically different from each other.

[0136] Figure 23. Dependence of mechanical properties on fiber fraction and orientation. Increased fiber fraction and alignment to the loading direction increased modulus (A, B). Increased fiber fraction elevated the yield stress, while adjusting fiber orientation from 15° to 0° did not significantly change yield stress (C, D). Increased fiber fraction did not significantly enhance UTS (F). Alignment of fibers in the loading direction results in greater ultimate stress compared to alignment perpendicular to the load (E). Bars without labels are significantly different from all other bars (p<0.05). Bars with the same letter are not statistically different from each other.

[0137] Figure 24. Four modes of failure are observed for composite sheets that reflected patterns of fiber orientation and volume fraction. At fiber orientations close to the loading direction and fiber fractions of 7 or 17%, samples exhibit failure soon after a single yielding (Mode 1 in panels A, B). At fiber fractions of 3%, the fiber network yielded at several different locations and levels of strain before tensile failure (Mode 2). Fiber oriented perpendicular to loading generated smooth deformation followed by failure at moderate strain (Mode 3). Samples without fiber deformed smoothly and tended to fail at higher levels of strain (Mode 4).

[0138] Figure 25. Examples of microhdge profile patterns, including triangular (A), rectangular (B), and chamfered rectangular (C).

[0139] Figure 26. Fabrication of the triangular microridge template. A silicon wafer (i) was spin-coated with negative photoresist (ii), and exposed to inclined ultraviolet light through a photomask (iii, A). The repeating strip pattern of the photoresist and the incident angle of the ultraviolet light determines the 3D micropattern of the UV crosslinked photoresist (B). Polyurethane is micromolded over the photoresist to generate the flexible template (C, D).

[0140] Figure 27. Fabrication of defined patterns of parallel chamfered rectangular microridge arrays. A layer of positive photoresist is patterned into strips on a silicon wafer with traditional photolithography techniques (A). The strips serve as a mask for inductively coupled plasma etching, yielding rectangular micro-trenches in the silicon (B). To fabricate chamfered rectangular microhdges, anisotropic wet etching in an aqueous KOH bath at is performed (C). After parylene coating, PDMS is molded over the template, followed by molding of polyurethane onto the PDMS to yield the parallel chamfered rectangular microridge membrane template (D, E, F). Rectangular microridged templates without the chamfered geometry are generated by a similar process, except after step (B) the silicon is coated with parylene and PU is cast over the silicon template.

[0141] Figure 28. Exemplary microcrimping system and method. A system, diagrammed in an exploded view (A), comprises the clamping assembly (i), microridged template membrane (ii), the base membrane (iii), and the lead screw assembly (iv). Pre-extension is applied to the base membrane (iii) with the lead screw assembly (iv). Then, the collagen fiber array is applied to the pre-extended base membrane and hydrated. The microridged template membrane is subsequently applied, with the same degree of pre-extension as the base membrane, and clamped with (i). The pre- extension is relaxed in both the base and the template membranes with the lead screw assembly to generate the microchmped geometry. Panel (B) illustrates a single collagen fiber clamped between the two pre-extended membranes. After the pre- extension is simultaneously relaxed in both membranes, the fiber becomes microcrimped (C).

[0142] Figure 29. Microridge design considerations. The shape of the fiber contacting region of the triangular microridge was too sharp, leading to relatively sharp grooves across the fiber (A). This raised concerns of partial fiber disruption. Tall, thin microridges are unstable, and collapsed. This results in non-uniform and unsymmethcal microcrimp (B). If the rectangular microridges are too short, there is insufficient overhead space for the fiber to fully deform (C). If the microridges are too wide, the fiber contained relatively long, flat segments, limiting the degree of crimp that could be imparted to the fiber (D). Chamfered rectangular microfeatures are selected because the ample overhead space, the relative stability against microridge collapse, and the blunt fiber contacting region (E).

[0143] Figure 30. Development of a microcrimping process. The triangular micro-ridge (A, optical micrograph, scale = 100μm) inflicts damaging grooves on the collagen fiber (B, SEM, scale = 200μm). When crimped without clamping, fibers often slip sideways instead of crimping (C, SEM, scale = 500μm). The tops of the chamfered rectangular microridge either tilted over or completely collapsed when the template membrane was fabricated from 3OA durometer polyurethane (D and E, SEM, scales = 200 and 100 μm).

[0144] Figure 31. Scanning electron microscopy of a microchmped collagen fiber array (scale bar is 200 μm). Synthetic collagen fiber provided herein is geometrically similar to native collagen.

[0145] Figure 32. Dependence of crimp morphology on pre-extension. Fiber arrays microcrimped with 30, 40, 45, and 50% pre-extension of the chamfered rectangular template demonstrated varying morphologies (A-D, respectively. Scale = 50 μm). For 45 and 50% pre-extension, the crimp shape appeared flattened rather than bending upwards in a smooth arc. Evidence that the fibers buckled during crimping (white arrows) was more severe in the 50% pre-extension sample.

[0146] Figure 33. Scanning electron micrographs and 3D reconstructions of hydrated, embedded fibers crimped with 15 (A, D), 30 (B, E), and 40% (D, F) pre-stretch. Scale bars 200 μm.

[0147] Figure 34. Determination of the degree of fiber crimp by rotating 3D images obtained from confocal laser scanning microscopy. Fibers are crimped by 15 (A) or 30% (B) pre-extension. The degree of crimp is defined as the difference in lengths between the straight fiber length (white line) and the path of the crimped fiber (red line) divided by the straight fiber length. Scales are 200 μm.

[0148] Figure 35. (A) Uniaxial stress-strain behavior for composite lamellae containing microcrimped fibers aligned parallel to the direction of the imposed load. The degree of crimp influenced the mechanical response. Non-crimped fiber (solid), 15% pre-stretch (dotted), and 30% pre-stretch (dashed). (B) Stress-strain response of a composite membrane reinforced with fibers in which crimp was induced by a pre-extension of 30%. [0149] Figure 36. Effect of cyclic tensile loading on crimp. Three-dimensional reconstructions of crimped fiber before loading (A, D), after 15 cycles (B, E), and 1000 cycles (C, F) of loading to 10% strain demonstrates that the crimp shape is generally preserved. Scales bars 200 μm.

[0150] Figure 37. Fabrication of a fiber reinforced small diameter vascular graft from oriented synthetic collagen fiber arrays embedded in an elastin protein polymer matrix. (A) Parallel arrays of fiber are created by winding about a frame. (B) Two such arrays are oriented at the desired angle and transferred to a glass sheet. (C) Fiber arrays are surrounded with precision shims and a solution of elastin protein polymer is applied before a polycarbonate sheet is pressed over the fibers to spread the solution into a thin film (D) The gelled film is then rolled about a Teflon tube to create a six-layered tube. (E) Schematic illustrates average fiber spacing (d) and angle (θ).

[0151] Figure 38. System to assess vascular graft pressure-diameter response and burst pressure. The graft (i) is suspended in PBS at 37°C with the lower end plugged with a 5 g weight (ii). As a syringe pump (iii) inflates the graft, a transducer (iv) reports the pressure, and changes to the graft diameter are monitored by video (v).

[0152] Figure 39. Result of Inverse Fast Fourier Transform analysis of fiber orientation. Peaks corresponding to 25.6° and -21.2° represent the primary fiber orientations in this layout, from a graft design with a fiber fraction of 7.3% and nominal angle of 22.5°.

[0153] Figure 40. Vascular grafts fabricated with 15 (A), 22.5 (B), and 30° (C) collagen fiber layouts. Collagen fibers are stained with von Gieson. Cross-sectional view of a prototype (D). Scale bar 2 mm.

[0154] Figure 41. Scanning electron microscopy of a prototype of design 6. Synthetic collagen fibers appear close to the surface of the graft exterior (A, B). Delineations or seams between the six layers of the wrapped film do not appear in the cross-sections of the graft wall (C, D). Compared to the graft exterior, collagen fibers do not appear as close to the surface of the lumen (E, F) with rare exception (grooves visible in F). Scale bar 200 μm.

[0155] Figure 42. Representative pressure-diameter responses for composite vascular grafts. (A) Increasing fiber density at a fixed 30° fiber angle yielded prototypes with enhanced burst pressure. (B), Increasing fiber angle at a fixed fiber fraction of 6 to 7% yielded prototypes with decreased compliance. [0156] Figure 43. Dependence of suture retention, compliance, and bursting strength on fiber spacing (A, B, C) and angle (D, E, F). In plots A, B, and C, NA indicates data for design 1 , without fiber reinforcement.

[0157] Figure 44 Mechanical response of the protein polymer composite compared to human fascia. The 25° fiber orientation produces a mechanical response similar to that of human linea alba (I and S refer to tissue strips from the infraumbilical or supraumbilical regions of the linea alba, while O and T refer to samples tested in the oblique or transverse orientations. Linea alba data adapted from Grassel, D., et al., J Surg Res, 2005. 124(1 ): 118-25.

[0158] Figure 45 Abdominal defect repair. Appearance of the multilamellar elastin- composite following implant (A) and at 8 weeks (B). None of the repaired defects demonstrated hernia formation for the duration of the study (C), as compared to unrepaired defects (D).

[0159] Figure 46 Histology of abdominal repair materials. The appearance of non- implanted multilamellar protein composite sheets and the porcine dermis product is shown in (A) and (D), respectively (100x). After 8 weeks, the elastin-like protein component appeared largely absent, except in rare areas. (B, 4Ox, elastin-like protein and synthetic collagen fiber indicated with solid and dashed arrows, respectively). In regions where cells and fibrous tissue replaced the elastin-like protein, the spacing between synthetic collagen fibers increased (C, 100x). The dense collagen of the porcine dermis product appeared to have separated, with cell and tissue ingrowth between implant fragments (right side of E, 4Ox, and in G, 100x). The arrows in (G) indicate implant fragments. In (E) the host-implant interface is apparent, with the abdominal wall at left in and the porcine dermis at right. Identifiable fragments of the porcine dermal product were largely absent from many regions of the harvested patches (F, 20Ox). Scale bars 200 μm.

DETAILED DESCRIPTION OF THE INVENTION

[0160] The following acronyms are used herein: θxωy - Fiber layout in a composite; x refers to the fiber angle and y is the spacing in mm; Tm - Apparent temperature of melting; ΔH - Apparent enthalpy; μDSC - Micro-differential scanning calohmetry; CF - Continuous fiber; CF-FIB Continuous fiber treated with fiber incubation buffer; CF+FIB Continuous fiber not treated with fiber incubation buffer; DF - Discontinuous fiber; DF- 0%, DF-15%, DF-30%: Discontinuous fiber with an applied strain during mechanical annealing; DMSO Dimethyl sulfoxide; DMTA Dynamic mechanical thermal analyzer; DVI Digital volumetric imaging; FIB Fiber incubation buffer; F/B Forward-to-backward ratio in second harmonic generation experiments; HE - Hematoxylin eosin stain; ICP Inductively coupled plasma; LSM Laser scanning microscopy; MRTC Monomehc rat tail tendon collagen; PBS Phosphate buffered saline; PDMS Polydimethylsiloxane; PEG

Poly(ethylene glycol); PFCS Protein fiber composite sheet; PU Polyurethane; PVC Polyvinyl chloride); RTT Rat tail tendon; SHG Second harmonic generation; SEM Scanning electron microscopy; TRITC Tetramethylrhodamine isothiocyanate; WSB Wet spinning buffer; UTS Ultimate tensile strength.

[0161] As used herein "film" refers to a layer of an elastic material, such as an elastomeric material. The film can have a thickness ranging from relatively small coating to relatively thick layers. In an aspect, the thickness is selected from a range that is greater than or equal to 10 μm and less than or equal to 1 mm, and any subranges therein.

[0162] "Embedded" refers to a fiber material that is at least partially covered by the first material. In an aspect, embedded refers to a fiber material that is completely covered by the first material.

[0163] "Continuous spun fibers" refers to collagen fibers made from a solubilized collagen starting material and that are collected on a spinning receiving system. In this manner, long-length collagen fibers are collected and large surface-area composites can be made with continuous collagen fibers spanning the footprint. Accordingly, the collagen fibers are said to "extend a length" of the first material such as the entire length of the first material. In the case where the fibers are oriented, such as not parallel with an edge of a rectangular sheet, the collagen fibers embedded in the first material may actually be longer than the length of the first material. A "preferential direction" refers to the alignment of the fibers and, in particular, where the fibers are on aligned within at least a 10°, 5°, or 3° range with respect to each other. Fibers are said to be "aligned" when they have a direction that is within at least 2° of the other aligned fibers. In an aspect, there can be multiple preferential directions or multiple alignments.

[0164] "Undulating" refers to a fiber that has a z-coordinate position that varies along the longitudinal axis. In other words, an undulating fiber has vertical bends. Crimp is used in a similar manner, but also refers to a process used to generate the undulations. The crimps or undulations accommodate a certain level of strain before the inherent properties of the fiber are exerted. In this manner, at lower strain levels the first material's physical properties (e.g., low Young's modulus) predominate as the fibers uncrimp, while at higher strain levels when the fibers are uncrimped the fibers' inherent physical properties (e.g., high Young's modulus) dominate.

[0165] "Crimp magnitude" refers to the strain at which the fiber unchmps. It can be quantitatively defined as the ratio of the length of the crimped geometry at rest to the corresponding end-to-end straight line length at rest. Accordingly, a fiber with a larger crimp magnitude requires a correspondingly larger strain to uncrimp the fiber.

[0166] A "spatially-varying" pattern refers to a user selected placement of fibers. For example, the crimp geometry may vary with position along the first material. Similarly, the density of fibers in the first material may vary such as by defining regions where the fibers are widely separated and other regions where the fibers are tightly packed. In addition, collagen fibers having different physical parameter(s) (e.g., diameter, Young's modulus) may be selectably positioned as desired.

[0167] "Synthetic" refers to an isolated artificial protein that is not normally made by an organism. A synthetic protein may be made by an organism or manufactured outside an organism. For example, the protein may be a recombinant protein in that a organism has been genetically engineered to express the protein or a precursor thereof.

[0168] "Thblock" refers to a protein having at least three distinct regions, such as a hydrophobic central block that separates end blocks that tend to be more hydrophilic. Optionally, a thblock amino acid sequence has additional material inserted between one or more of the blocks or at the block ends. For example, a cross-linkable amino acid or modified amino acid that is capable of cross-linking may be inserted between the blocks to facilitate cross-linkage manipulation. Spacers such as amino acid spacers may be included with the cross-linkable amino acids to further influence cross-linking. Such chemical cross-linking may be in addition to the physical cross-linking that tends to occur naturally with the amphilic thblocks and provides ability to tailor a mechanical property to the end-application to which the protein may be used.

[0169] "Creep" refers to a mechanical property of a material that is time-dependent. In particular, creep relates to the tendency of a material to permanently deform in response to an applied force or stress applied over time, or a time-dependent deformation of the material under stress. [0170] "Inverse transition temperature" refers to the property where a material is a liquid at a lower temperature, but changes state to a gel or solid at a higher temperature. The temperature at which such a change of state begins is referred to as the "inverse transition temperature" and is useful for assisting in placement of an embolic agent into a cardiovascular defect as a liquid initially that later changes to a gel or solid, thereby providing therapeutic benefit.

[0171] "Young's modulus" is a mechanical property of a material, device or layer which refers to the ratio of stress to strain for a given substance. Young's modulus may be provided by the expression;

wherein E is Young's modulus, L₀ is the equilibrium length, ΔL is the length change under the applied stress, F is the force applied and A is the area over which the force is applied.

[0172] "Elastomer" refers to a biopolymer material which can be stretched or deformed and return to its original shape without substantial permanent deformation. Elastomers commonly undergo substantially elastic deformations. An "elastomeric film" refers to an elastomer material formed into a layer having a defined thickness.

[0173] "Physical parameter" refers to a property of the protein or material made from the protein and includes mechanical parameters provided herein (e.g., Young's modulus, bending modulus, stiffness, compressibility, ultimate tensile stress, strength, fracture or failure strain, resilience, compliance, permeability, swelling ratio, dimension, and other parameters and particularly those parameters used in the art to describe biological systems and materials). A "tunable physical parameter" refers to a parameter that can be controllably adjusted by any of the methods disclosed herein or that depends on the structure or sequence of the proteins that make up a film or fiber network. For example, adjusting the properties of the end and/or central blocks (e.g., length, hydrophobicity) permits tuning of a physical parameter that describes the environment or surrounding tissue in which the film or fiber network is to be used or implanted into (e.g., a blood vessel or a portion of the cardiovascular system). Optionally, further tuning is accomplished by any processing or post-processing known in the art or by orienting or spatially distributing fibers thereby providing further control of the mechanical properties of the medical device.

[0174] "Embolic agent" refers to a material that is capable of physically impacting blood flow or altering hemodynamics in and around a blood vessel. The embolic agent may be applied to a blood vessel or blood vessel wall, such as a wall rupture or aneurysm, in a liquid form that subsequently gels or solidifies, thereby displacing or preventing further blood flow in a region. Alternatively, the embolic agent may be applied as a gel, semisolid or solid in a blood vessel or blood vessel wall, such as a wall rupture or aneurysm to provide a therapeutic benefit.

[0175] Example 1 : Large Scale Production of continuously spun Synthetic collagen fiber.

[0176] Collagen fiber serves well-known structural roles in many tissues, and has been a candidate for tissue repair and replacement for decades. The earliest version of manufactured collagen fiber, catgut suture, consisted of sheep or bovine intestine, chemically and mechanically processed into strands that were ground and polished to create suture [34]. The processing introduced non-uniformities and potential mechanical failure points. Catgut also generated variable tissue reactivity and often a highly inflammatory response [35]. Following catgut, reconstituted collagen fiber processes were developed to improve thread uniformity and reduce non-collagen tissue remnants to decrease the inflammatory response. In these protocols, tendon was processed into gels or dispersions consisting of collagen fibrils and extruded it into acetone-based solutions to generate fiber [36, 37]. However, animal studies indicated inferior performance of chromic reconstituted collagen compared to chromic catgut, seemingly due to suture fragmentation and chemical irritation [38].

[0177] Subsequent to reconstituted fibers, others processed collagen into more highly pure fiber explicitly for medical fabrics and tissue engineering. Collagen was formulated into dispersions or solutions and extruded into aqueous buffers instead of acetone. After an incubation period, fibers were dehydrated in alcohol and air-dried [39]. Before fiber fabrication, these protocols purified the collagen into ultrafiltered solutions of aggregates of a few molecules or monomehc collagen, instead of the collagen fibril dispersions of earlier processes [40, 41]. These and similar methods also employed longer buffer incubations, totaling 30 min [42] to 48 hr [43]. Longer incubation provided ample time for collagen molecules or aggregates to reassemble into fibrils. These incubations usually limited the length the self-assembled collagen fiber to short 10-20 cm segments instead of continuous fiber. For example, if fiber was continuously extruded at 1 m/min, and 30 minutes of incubation were required, then 30 meters of the nascent protein filament had to be threaded through a series of incubation baths without breaks or tangles. To avoid this difficulty, long incubations were truncated in continuous fiber spinning systems [44-46]. However, in one such case, inferior properties were reported when comparing continuous fiber to short fiber segments made with longer incubations [44]. In another instance, electron micrographs showed that banded collagen fibrils had only self-assembled on the outer shell of continuously spun fiber [45].

[0178] In this example, we report a process in which the requirements for continuous fiber spinning are decoupled from the long incubations required for fibril self-assembly. The continuous spinning system includes an extrusion tube and 2 m rinsing bath. After the fiber is dried, collected, and stored on rollers it is subjected to a separate off-line incubation step. The decoupled system results in scalable production of continuous fiber, with evidence of fibril formation throughout the fiber cross-section.

[0179] MATERIALS and METHODS: Isolation and purification of monomehc collagen. Acid-soluble, monomehc rat-tail tendon collagen (MRTC) is obtained from Sprague- Dawley rat tails following Silver and Trelstad [47]. Frozen rat tails (Pel-Freez Biologicals, Rogers, AK) are thawed at room temperature and tendon is extracted with a wire stripper, immersed in 10 mM HCI (pH 2.0; 150 ml_ per tail) and stirred for 4 hr at room temperature. Soluble collagen is separated by centhfugation at 30,000 g and 4°C for 30 minutes followed by sequential filtration through P8, 0.45 μm, and 0.2 μm membranes. Addition of concentrated NaCI in 10 mM HCI to a net salt concentration of 0.7 M, followed by 1 hr stirring and 1 hr centhfugation at 30,000 g and 4°C, precipitates the collagen. After overnight re-dissolution in 10 mM HCI the material is dialyzed against 20 mM phosphate buffer for at least 8 hr at room temperature. Subsequent dialysis is performed against 20 mM phosphate buffer at 4°C for at least 8 hr and against 10 mM HCI at 4°C overnight. The resulting MRTC solution is stored at 4°C for the short-term or frozen and lyophilized.

[0180] Production of a synthetic collagen microfiber by continuous co-extrusion. A modified wet spinning device facilitates collagen fiber production (Figure 1). A collagen solution (2, 5, or 7.5 mg/ml in 10 mM HCI) and wet spinning buffer (WSB: 10 wt% poly (ethylene glycol) Mw = 35000, 4.14 mg/mL monobasic sodium phosphate, 12.1 mg/mL dibasic sodium phosphate, 6.86 mg/mL TES (N-ths (hydroxymethyl) methyl-2- aminoethane sulfonic acid sodium salt), 7.89 mg/mL sodium chloride, pH = 8.0) are extruded with a dual syringe pump (Harvard Apparatus, Holliston, MA). The collagen solution emerges through a 0.1 or 0.4 mm inner diameter blunt-tipped needle into the center of a vertical tube (1.6 mm inner-diameter x 1 m long fluoropolymer tubing) at rates 0.03, 0.06, or 0.1 mL/min. Wet spinning buffer simultaneously advances through a bubble trap and down the fluoropolymer tube at a rate of 1.0 mL/min. As it exited the extrusion needle, the collagen coagulates into a gel-like fiber and is carried downward by the WSB stream.

[0181] Upon emergence from the fluoropolymer tube, the fiber enters a 2 meter-long rinsing bath of 70% ethanol in water. Initially, 5 to 10 meter samples of manually collected fiber (MC) are collected by hand on rectangular frames. After optimization, automatically collected fiber (AC) is produced and collected by winding it out of the rinsing bath onto segments of polyvinyl chloride (PVC) pipe that rotated and translated automatically.

[0182] Fiber incubation and drying. After spinning, the fiber is placed in fiber incubation buffer (FIB: 7.89 mg/mL sodium chloride, 4.26 mg/mL dibasic sodium phosphate, 10 mM Tris, pH = 7.4) [48] at 37°C for 48 hr. Manually collected fiber is incubated on rectangular frames, while automatically collected fiber is incubated directly on PVC pipe segments. Rectangular frames containing MC are subsequently rinsed for 15 min in ddH₂0 and 2 min in 70% ethanol before drying in air. Pipe segments containing AC fiber are rinsed in ddH₂0 for 15 min before drying and collecting the fiber under tension with an automated system (Figure 2).

[0183] Fiber crosslinking. Some fibers are left in a desiccator on a ceramic plate above a pool of 25% (w/v) glutaraldehyde in water for 18 - 24 hr. To remove excess glutaraldehyde, fibers are subjected to three 4 hr rinses, gently rocking in PBS. Fibers are stored in PBS for immediate testing or rinsed in water and air dried.

[0184] Optimization of wet spinning parameters. Spinning parameters, including the concentration of the collagen solution, collagen solution flow rate, and extrusion needle size are varied and the diameter and strength of the resulting MC fiber is measured. Nine combinations of wet spinning parameters are investigated before the conditions for AC fiber production are selected (Table 1). For each set of parameters that consistently produced fiber, samples are crosslinked with glutaraldehyde vapor and rinsed before diameter measurement and mechanical testing. Initial experiments shows that decreasing collagen flow rate and concentration lowers fiber diameter, resulting in more frequent fiber failure during processing. Conditions are selected for automatically collected fiber production to minimize fiber diameter while eliminating breaks during fiber spinning.

[0185] Mechanical annealing of collagen fibers. Some samples of fiber are subject to a mechanical annealing protocol similar to that previously reported by Pins et al. [49]. Specifically, fiber is collected manually from the spinning system on expandable rectangular frames and transferred to 37°C FIB solution for 24 hr. Adjustment to the expandable frames imparts 0, 15, or 30% strain to the fibers followed by 24 additional hours in buffer. These samples, referred to as MC-0%, MC-15%, and MC-30%, are rinsed and dried identically to non-annealed MC fiber and analyzed with transmission electron microscopy (TEM).

[0186] Analysis of fiber diameter and cross-sectional area. Optical microscopy permits measurement of the crosslinked, hydrated fiber diameter. Preliminary examination reveals that the fiber cross-section is somewhat flattened and resembles an ellipse. Each fiber segment is twisted to display the widest and narrowest cross-section dimensions beneath a 2Ox objective and three measurements of each are recorded per sample. After analyzing six to ten samples from each type of fiber, the cross-sectional area is approximated using:

"cross-section ~ TT (LJma_jor * Uminorj / 4

where Dmapr and D_mιnor represent the diameters of the major and minor axes of the fiber cross-section.

[0187] Mechanical responses of single collagen fibers. A Dynamic Mechanical

Thermal Analyzer V (DMTA V, Rheometric Scientific, Piscataway, NJ) enables tensile strength analysis of crosslinked, hydrated fiber samples. Fiber is prepared by gluing either end of a segment of dry fiber between thin plastic shims with cyanoacrylate glue. Samples are hydrated in PBS overnight and loaded in the DMTA by clamping the plastic shims in the instrument grips. The DMTA is inverted with the sample submerged in a thermally jacketed beaker of 37°C PBS. Samples are loaded with a gauge length of 6 to 9 mm, pulled to 5% strain and relaxed in four preconditioning cycles, and pulled to failure at a rate of 5 mm/min. The force and strain at failure are recorded and engineering ultimate tensile stress (UTS) is calculated using the cross-sectional area estimate for each fiber type.

[0188] Microdifferential scanning calohmetry (μDSC). Automatically collected fiber before or after the FIB treatment (AC-FIB and AC+FIB), after FIB treatment and glutaraldehyde crosslinking (AC+FIB+GLUT), and rat tail tendon (RTT) are pressed into 5 -8 mg pellets, dried under vacuum, massed, and hydrated for 10 hrs in 0.5 ml_ of PBS at 5°C. Similarly, MRTC is lyophilized, massed, and hydrated. Analysis in a μDSC III (Setaram Instrumentation, Caluire, France) permits the investigation of triple-helix unfolding as samples are heated from 5 to 90⁰C at 0.5°C/min and returned to 5°C. Heat flow per unit mass is plotted against furnace temperature. The denaturation temperature, T_m, is determined from the peak of the denaturation transition, while the area of the transition is used to calculate the enthalpy of denaturation, ΔH. Lack of a similar thermal transition in a second identical heating cycle confirms complete denaturation in the initial heating cycle.

[0189] Analysis of Second Harmonic Generation (SHG). The SHG system comprises an Olympus 1X71 inverted microscope coupled to a Ti-Sapphire laser (Spectra Physics, Mountain View, CA). A 60 x 1.2 NA (0.28 mm) water immersion objective focuses and collects the backward signal, while a 1.4 NA oil immersion condenser (Olympus, Center Valley, PA) collects the forward signal. High sensitivity photomultiplier tube devices (Hamamatsu Photonics, Hamamatsu City, Japan) detects both the forward and backward signals. In preparation, AC fiber is hydrated in PBS for 2 hrs and mounted beneath a cover slip with PVA-Dabco mounting media. Ten-micron cryosections of RTT are also mounted with PVA-Dabco. Spectrometer analysis of the emitted signal confirmed the presence of a steep intensity peak at 480 nm, half the excitation wavelength of 960 nm, indicative of SHG. The forward-to-backward (F/B) signal ratios are calculated for both rat tail tendon and continuous fiber.

[0190] Transmission electron microscopy. As previously described [48], bundles of approximately ten parallel, uncrosslinked fibers are tied at both ends with 4-0 prolene suture, hydrated for 1 hr in PBS, rinsed three times in 0.1 M cacodylate buffer (pH = 7.4), fixed (2.5% glutaraldehyde in 0.1 M cacodylate buffer, pH 7.4) for 90 minutes, and washed in 0.1 M cacodylate water followed by ddH₂0. Samples are partly dehydrated in 30% ethanol followed by staining en bloc with filtered 2% uranyl acetate in 50% ethanol. Preparation continued with ethanol series dehydration, 100% resin infiltration, and polymerization for 3 days at 60⁰C. Using an RMC MT-7000 ultramicrotome (Boeckeler, Tucson, AZ) and a diamond knife, ultrathin sections (60-80 nm) are cut to reveal cross and axial perspectives of the fiber interior. Following a post-stain with 3% uranyl acetate and Reynold's lead citrate, sections are examined and photographed with a JOEL JEM-1210 TEM (JOEL, Tokyo, Japan) at 90 kV. Samples investigated included MC-0%, MC-15%, MC-30%, and AC. The diameter of collagen fibrils in continuous fiber is estimated from 40 measurements taken from four 50,00Ox magnification images of the fiber cross-section, and the result is expressed as the mean and standard deviation.

[0191] Murine subcutaneous implant studies. The host response and biostability of glutaraldehyde crosslinked and uncrosslinked continuous fiber is assessed by subcutaneous implant in 10 wk old, 25-30 g, inbred male C57BL/6 mice (Jackson Laboratory, Bar Harbor, ME). To prepare the implants, bundles of 160 fibers, 2 cm in length, are aligned and tied at both ends with 7-0 prolene suture. Crosslinked bundles are exposed to glutaraldehyde vapor for 18 hr, followed by a 24 hr rinse in sterile water and a 1 hr incubation in PBS immediately before the implant. Uncrosslinked bundles are incubated in PBS for 1 hr immediately prior to implant. Three bundles of crosslinked and uncrosslinked fiber are implanted for a 6 wk period.

[0192] Following sedation with ketamine (95 mg/kg, IM) and xylazine (5 mg/kg, IM), a subcutaneous pouch is created through a dorsal midline incision and the fiber bundle implanted. After 3 or 6 wks, animals are sacrificed, fiber bundles excised with overlying skin, and samples photographed to qualitatively assess gross local tissue responses. All samples are fixed overnight in 10% neutral buffered formalin and processed for parafin embedding. Five-micron sections are stained with hematoxylin and eosin (HE) to visualize tissue morphology and Gomoh's Trichrome to distinguish collagen, lmmunohistochemical staining is performed with rat monoclonal antibody [CI:A3-1] which recognizes the F4/80 antigen expressed by murine macrophages (Abeam, Inc., Cambridge, MA).

[0193] Statistics. Tests for statistically significant differences between the means of two groups is conducted with the Student's f-test (two-tailed, homoscedastic). Tests between three or more groups are conducted with the one-way ANOVA followed by the Tukey HSD test. [0194] RESULTS: Development of a scalable system for continuous collagen fiber formation. For AC fiber production, PVC pipe segments (outer diameter = 48 mm) are typically rotated at 6 rpm and translated at 6 mm/minute, leading to the deposition of consecutive loops of 15.2 cm of fiber along the 20 cm length of the pipe segment. Each pipe segment is therefore loaded with about 30 m of fiber in 33 min. Although production rate is not maximized, this supply proved adequate for characterization and production of experimental collagen fiber assemblies.

[0195] Collagen fibers can be produced without loss of triple helical structure. The μDSC study (Table 2, Figure 3) demonstrates that wet spinning without FIB treatment does not change ΔH, and increases T_m by 8.4°C on average. Treatment with FIB increased ΔH to the level of RTT and slightly increased T_m. Samples of RTT display a substantially higher T_m.

[0196] Collagen fiber size is influenced by selection of wet spinning parameters. The average major fiber cross-section dimension increases significantly (p < 0.05) with increasing needle size, collagen flow rate, and collagen concentration (Figure 2.4).

Average UTS of manually collected samples is between 54 and 90 MPa. Differences in UTS are not significant at the p < 0.05 level for the samples compared in Figure 4. Automatically collected fibers display the highest mean UTS (Figure 5, Table 3).

[0197] Collagen fibers can be produced as a close packed assembly of axially oriented D-periodic fibrils. The axial sections of FIB-treated MC samples reveals banded fibrils while the sections from untreated samples do not (Figure 6). With no axial stretching (MC-0%), the fibrils are disorganized, although isolated areas of alignment with the fiber axis could be identified. Fibers mechanically annealed to 15 or 30% display qualitatively more alignment (Figure 7). The automatically collected fiber sections display an aligned, densely packed fibril structure even without mechanical annealing (Figure 8). Fibril diameter average and standard deviation are 54 ± 13 nm. Fiber samples also display an SHG signal (Figure 9). The F/B ratio is 0.039 for continuous fiber and 3.75 for rat-tail tendon.

[0198] Murine subcutaneous implant studies. After six weeks, all samples could be readily located and explanted. Uncrosslinked fiber bundles are firmly adhered to the skin, while crosslinked bundles are not adhered. Photographs before and after the implanting indicate that the length of the crosslinked bundles is unchanged, while the uncrosslinked bundles shrink to approximately 30% of the original length during the implant period. Fiber diameter and bundle size are similar for both crosslinked and uncrosslinked samples (Figure 10). Gomori's Thchrome stained crosslinked and uncrosslinked fibers red and blue, respectively. More fibrous collagen is deposited around the perimeter of the uncrosslinked bundles, while the crosslinked bundles contain more fibrous material in the bundle interior, surrounding individual fibers. Fibers near the perimeter of the uncrosslinked bundles are smaller than those in the bundle center, lmmunohistochemistry demonstrate that macrophages are also located predominantly at the perimeter of uncrosslinked bundles, and within the crosslinked bundles (Figure 11).

[0199] DISCUSSION: This example describes large scale production of collagen fiber with fibrillar substructure that closely mimics that of native collagen fibers. In the first stage of spinning, a purified, acid-soluble, MRTC solution and a buffered PEG solution are continuously co-extruded through fluoropolymer tubing to form a fiber that is passed into a rinsing trough, through air, and onto a collection roller. Although the first spinning stage failed to reformulate the banded, fibrillar structure of native collagen, TEM demonstrates that an additional 48 hr incubation drives fibrillar self-assembly throughout the fiber cross-section. Without specialized equipment, the system spun 60 m/hr of fiber, which compares favorably to fiber production rates of 100 m/hr, previously reported for experimental collagen textile production systems that did not display fibrillogensis throughout the fiber cross-section [45].

[0200] Techniques for producing collagen fiber may be categorized either as processes that yield continuous fiber or short discontinuous fiber segments suitable for laboratory analysis. Excluding catgut production and older methods of spinning fibril dispersions, several scalable, continuous fiber-spinning methods have recently been proposed [44- 46]. Silver and Kato compared discontinuous and continuous methods and observed that fiber produced by a continuous process displayed somewhat inferior mechanical properties and faster biodegradation [44]. Both their continuous and discontinuous processes consisted of extrusion into a fiber formation buffer at 37⁰C, a rinse in isopropyl alcohol, a rinse in water, and air drying. However, during the continuous process the temperature of the fiber formation buffer was noted to fall below 37⁰C, and all of incubations and rinses were shorter than in the discontinuous process. Also, less tension was applied to the fiber during the air drying step of the continuous process. These factors, in particular, shorter incubation time in the fiber formation buffer, may not permit optimal collagen self-assembly and fibril formation, contributing to sub-optimal mechanical properties and biostability [44]. Pins and coworkers have also reported a discontinuous fiber process using longer incubation periods, and were able to generate fibrils throughout the fiber cross-section [41 , 43]. In another discontinuous fiber system, Zeugolis and colleagues recently demonstrated fibril formation after three relatively short incubations of 10-15 min [42, 50-52]. Adaptation of this protocol to continuous spinning may be feasible but would be more complex in nature with a lower overall production rate, given the requirement for a total incubation time of up to 45 min. That group studied fibril formation in fibers extruded from both pepsin and acid extracted collagen. Pepsin treated collagen can be harvested in higher yield and contains fewer antigenic determinants due to cleavage of most of the telopeptide regions and, thus, may offer certain advantages as compared to acid extracted collagen. Unfortunately, fibril banding and density is significantly reduced in discontinuous fibers produced from pepsin extracted collagen [42].

[0201] Others have developed a microfluidic system to characterize fibril formation by real-time x-ray microdiffraction and birefringence during collagen fiber spinning [53]. Data generated from this technology may eventually enable the selection of temperatures, flowrates, and buffer systems for scalable production of fiber with aligned, banded fibrils without any additional incubation required. However, at present no process for spinning purified collagen solutions into continuous fibers with adequate fibrillar structure is available. By appending an off-line incubation step to a relatively simple continuous spinning process, we have created a process that is both scalable and results in fibrillogenesis throughout the fiber.

[0202] Thermal and optical analysis demonstrates conservation of collagen triple helical structure. Triple helical structure defines the collagen family of proteins, is a prerequisite for fibril self-assembly, and, critically, shields antigenic and proteolytic cleavage sites [31]. Use of unsuitable solvents and high temperatures, for example in early collagen wet spinning processes and electrospinning [54], destabilizes the triple helix and may be expected to compromise biocompatibility, biostability, and strength of the resulting material. In this study, μDSC demonstrated that spinning monomehc collagen into fiber, before fibrillogenesis, did not disrupt triple helical structure, and raised the apparent helix T_m by 8.4°C. This increase may largely be attributed to the enhanced stability of the collagen aggregated in the fiber compared to the monomehc collagen in solution. However, the different hydrating solutions used during the calohmetry measurement also caused a portion of the T_m difference. The fiber was hydrated with PBS (pH 7.4), while the monomeric collagen was dissolved in 10 mM HCI (pH 2.0). The stability of the triple helix is related to pH, and increasing pH from 2 to 7.4 which would account for more than a 2°C increase in T_m [55, 56]. Comparing results for fibers before and after fibrillogenesis demonstrates that fibril formation moderately raised T_m and elevated ΔH by 15%. The greater thermal stability observed after fibrillogenesis is likely due to hydrophobic interactions in the fibril and greater surface energy associated with the melting of a larger fibrillar structure [56]. The T_m of AC+FIB was 12°C less than that noted for native tendon, since the tendon is further stabilized by native covalent crosslinks.

[0203] SHG emission is consistent with retention of triple helical structure and fibril assembly. Second harmonic generation occurs when laser light passes through a molecularly noncentrosymmetric, highly polahzable material. The wavelength is halved and the frequency is doubled in a coherent optical process [57]. SHG in extruded fiber and other reformulated collagens confirms the retention of triple helix structure with loss of triple helical structure eliminating the SHG signal [54]. Although the collagen triple helix is necessary to produce an SHG signal, this feature alone, in the absence of fibrillar assembly is insufficient for signal generation. For example, triple-helical, non- fibrillar collagen (type IV) does not produce SHG [58, 59]. Furthermore, reports have proposed that only a portion of the collagen in a fibril, possibly in the fibril's outer shell, can generate SHG scattering [60, 61]. High SHG forward to backward signal ratios have been correlated with increased collagen fibril thickness. This has been attributed to the accumulation of the forward signal as light travels through the fibril with limited accumulation of the backward signal due to destructive interference [60]. Therefore, the low F/B ratio observed in continuous fiber may be due to lower fibril diameters compared to RTT. The reconstituted fibril diameters in this study are 54 ± 13 nm, while the fibrils from RTT are reported to have a bimodal diameter distribution with average diameters of 70 and 250 nm [62].

[0204] Wet spinning parameters can be adjusted to optimize continuous fiber diameter and strength. Optical microscopy reveals that continuous fibers are produced with an elliptical cross-section with a ratio of major to minor axis diameters of approximately 2:1. Reduction of collagen extrusion rate or the collagen concentration reduced the diameter of the fiber, presumably because in both cases less collagen is supplied to produce a given length of fiber. Reducing the diameter of needle used for extrusion of the collagen solution by a factor of four led to only slight reduction of fiber diameter, probably because the decrease in needle size did not change the flow rate of the collagen solution. The smallest fibers produced are 12 by 25 μm in cross-section, however, parameters ultimately selected for automatically collected fiber yield larger fibers to eliminate breaks during wet spinning. Dimensions of fiber produced by other strategies are summarized in Table 4.

[0205] The ultimate tensile strength of the automatically collected fiber is compared to collagen fiber in other reports in Table 4. Automatically collected fibers are on average stronger than manually collected fibers of a similar diameter in this study. Qualitatively, TEM shows enhanced fibril alignment and density in AC, possibly generated by greater tension applied during the spinning, drying and collection of AC. After drying and collection, the length of the dried fiber is approximately 5% more than the length of the incubated, hydrated fiber. However, the total strain induced in the drying step includes both the 5% observed strain and the reduction in fiber length that would occur if the fiber is dried without tension, at least 5 to 10%. Therefore, the drying and collecting step imparts 10 to 15% strain to the fiber. Indeed, mechanical drawing and related techniques have been shown to increase strength in other collagen fiber systems [41 , 63]. In addition to mechanical force, others have employed magnetic and electric fields [64, 65], magnetic beads [66], thermal convection currents [67], blow drying [68], microfluidics [53, 69], and cell-based collagen synthesis and assembly [70-74] to align collagen. The goals of these studies are related either to enhanced mechanics or contact guidance in cell culture and tissue engineering systems.

[0206] Local tissue responses to collagen fibers. Crosslinked fiber implants display short-term biostability and a local inflammatory response consistent with previous reports [75, 77]. While others have quantified biodegradation as a reduction in the number of intact fibers [78], we also observed an overall shortening of the uncrosslinked fiber bundles by approximately 70%, while the length of crosslinked bundles did not appreciably change. However, the uncrosslinked bundles could still be identified and explanted. In contrast, others have found that even after dehydrothermal-cyanamide crosslinking, many fibers were completely degraded after 10 days and all fibers were degraded after 6 wk in a rat intramuscular implant [75]. Similarly, Kato and Silver reported that crosslinked fiber bundles were degraded 2 weeks after implantation in the subcutaneous tissue of a rat animal model [44]. The greater stability we observe may be partly due to our use of a mouse model rather than any property of AC fiber. Khouw et al. compared the response of mice and rats to subcutaneous collagen implants and found more giant cells and substantially increased evidence of implant phagocytosis in rats when compared to mice [79].

[0207] Although collagen fiber has been manufactured for decades, techniques that are easily scalable for textile production, begin with a purified solutions rather than fibril dispersions, and yield fibrillar assembly throughout the fiber have been slow to materialize. Here we provide a process for continuous spinning followed by off-line incubation. This system does not disrupt the triple helical macromolecular structure, and triggered fibhllogenesis. Bundles of crosslinked fiber generated a typical healing response in a mouse implant. Uncrosslinked fiber exhibited relative biostablility.

[0208] Example 2: ANISOTROPIC PROTEIN POLYMER LAMELLAR ELASTIC STRUCTURES

[0209] In this example, we describe a new, scalable, semi-automated, fabrication process that yields an elastic or elastomehc material with stiffer fibers controllably positioned in the elastic material. In one example, an elastin-like protein sheet is reinforced with synthetic collagen fibers that can be positioned in a precisely defined three-dimensional hierarchical pattern. An artificial collagen fiber reinforced composite within an elastin-like matrix protein that displays many of the biomechanical properties of native tissues is presented in this example. Significantly, the flexible nature of the fabrication process lends itself to varying fiber orientation and volume fraction within and between individual lamellae of a planar sheet made of a plurality of lamellae. Moreover, it is evident that such structures can be used as acellular tissue analogues or incorporated within schemes that integrate cells within the analogue prior to or after implantation in vivo.

[0210] The composition and hierarchical structure of collagen and elastin protein fiber networks dictates the mechanical responses of all soft tissues and related organ systems. In turn, the biomechanical properties of a tissue dictate a variety of performance characteristics that affect function and durability, including local cellular behavior. As a prototypical soft tissue, the composite nature of the vascular wall was first highlighted in 1957 when Roach and Burton demonstrated that as pressure is increased within an iliac artery, the vessel initially behaves as a highly compliant tube, which displays a rapid increase in material stiffness as the physiologic range of normal blood pressure is exceeded. The highly compliant responses at relatively low pressures could be attributed to elastin, while the collagen fiber network was identified as the primary feature that dictated increasing tissue stiffness at high pressure [80]. This unique arrangement of collage and elastin networks afforded a tissue of substantial strength that is both compliant and resilient. Significantly, in the case of arterial blood vessels, these properties contribute to the dampening of peak pulsatile blood pressure, reduction of the mechanical work of the heart, and enhanced resistance of blood vessels to fatigue and catastrophic failure in response to repetitive cyclic loading forces [81].

[0211] The vessel wall provides a useful starting point for the consideration of the integrated structure and significance of collagen and elastin networks. However, nearly all other soft tissues are dependent upon the presence of such networks, whose uniquely site-specific composition and structure profoundly influences organ specific function. For example, as heart valve leaflets close during ventricular diastole and are subject to increasing strain, the mechanical response transitions from compliant to stiff due to the mechanics and geometry of a coordinated network of collagen and elastin fibers. Broom described the process of collagen fiber straightening and aligning with applied stress and the role of elastin in returning the collagen to its relaxed formation, as a complementary deformation processes [82]. This relationship is thought to facilitate efficient stretching of leaflets as the valve closes, providing large coaptation regions that limits retrograde bloodflow, as well as rapid leaflet opening in response to forward flow [83].

[0212] The biological and biomechanical significance of the extracellular matrix (ECM) has motivated the development of analogues as candidate scaffolds to replace diseased or damaged tissues, including blood vessels, heart valves, skin, as well as fascia and tendon. Two design decisions, sourcing of materials and the fabrication method, are inherent to each approach. In practice, investigators have most often addressed the challenge of generating a matrix substitute through the use of decellulahzed xenogeneic tissues, which have been used clinically in various forms for nearly 50 years. Although clinically effective devices have been processed from xeneogeic tissues, including bovine mesenteric vein and bioprosthetic porcine heart valves, the flexibility in designing a tissue to meet a desired set of conditions is limited by a fixed set of starting tissue properties, as well as matrix composition and architecture. Natural tissue variability and potential contaminants also pose challenges. [0213] Alternative approaches, that remain largely preclinical in nature, include the use of living cells to assemble extracellular matrices, as well as other fabrication methods, such as molding, lyophilizing, and electrospinning to process purified native proteins. Cell-based extracellular matrix assembly strategies include molding of cells in degradable polymers [84] and biopolymers [85] and blood vessel substitutes fabricated from rolled sheets of cells and their endogenous matrix [86]. Living tissue substitutes offer numerous advantages, but acknowledged limitations include long production times, cell sourcing, and the inability to create devices that display prolonged shelf life. Moreover, although collagen fiber networks produced by cells can be generated with some degree of alignment, the capacity to assemble fiber composites containing a substantial elastin network has not achieved nor has it been possible to create 3-D structures with precisely defined architecture that provides the flexibility to tailor related biomechanical responses [86-88]. The development of a convenient process for the large-scale production of continuous synthetic collagen fibers composed of D-pehodic fibrils facilitate the investigation of such an approach.

[0214] Several elastin-mimetic protein polymers based on variations of the [VPGVG] peptide repeat sequence of native elastin have been synthesized. In particular, recombinant triblock copolymers containing an elastomehc, hydrophilic midblock flanked by rigid hydrophobic, plastic-like, endblocks have emerged as a promising biomatehal platform with the capacity to undergo either reversible, physical crosslinking or chemical crosslinking. These materials exhibit excellent biostability in vivo, even in the absence of chemical crosslinks [3], can be processed as nanofibers with tunable mechanical properties [4], micelles [5, 6], or used for drug-release applications. Primate shunt studies have confirmed that these elastin protein polymers can form non-thrombogenic blood contacting coatings [7]. In this example, we employ an elastin-mimetic protein polymer with an integrated, oriented system of synthetic collagen fibers to fabricate APPLES (Anisotropic Protein Polymer Laminated Elastic Structures).

[0215] MATERIALS and METHODS: Synthesis of a recombinant elastin-mimetic triblock protein polymer. Genetic engineering, expression, purification, and characterization of the elastin-mimetic protein polymer, designated LysB10, has been described elsewhere [3] (see also WO 2008/033847 published March 20, 2008, hereby specifically incorporated by reference). Briefly, the flanking 75 kDa endblocks of the protein polymer contained 33 repeats of the hydrophobic pentapeptide sequence [IPAVG]₅, and the central 58 kDa midblock consisted of 28 repeats of the elastic, hydrophilic sequence [(VPGAG)₂VPGEG(VPGAG)₂]. The sequences between blocks and at the C terminus include the residues, [KAAK], provides amine groups for chemical crosslinking.

[0216] The protein polymer sequence is contained a single contiguous reading frame within the plasmid pET24-a, which is used to transform the E. coli expression strain BL21 (DE3). Fermentation is performed at 37°C in Circle Grow (QBIOgene) medium supplemented with kanamycin (50 μg/mL) in a 100 L fermentor at the Bioexpression and Fermentation Facility, University of Georgia. Cultures were incubated under antibiotic selection for 24 hr at 37°C. Isolation of the LysB10 comprises breaking the cells with freeze / thaw cycles and sonication, a high speed centrifugation (20,000 RCF, 40 min, 4°C) with 0.5% poly(ethyleneimine) to precipitate nucleic acids, and a series of alternating warm / cold centhfugations. Each cold centrifugation (20,000 RCF, 40 min, 4°C) is followed by the addition of NaCI to 2M to precipitate the protein polymer as it incubated for 25 min at 25°C. This is followed by warm centrifugation (9500 RCF, 15 min, 25) and resuspension of the pellet in cold, sterile PBS on ice for 10 - 20 min. After 6 to 10 cycles, when minimal contamination is recovered in the final cold centrifugation, the material is subject to a warm centrifugation, resuspended in cold sterile PBS, dialyzed, and lyophilized.

[0217] Isolation and purification of monomehc collagen. Acid-soluble, monomehc rat- tail tendon collagen (MRTC) is obtained from Sprague-Dawley rat tails following Silver and Trelstad [47]. Frozen rat tails (Pel-Freez Biologicals, Rogers, AK) are thawed at room temperature and tendon is extracted with a wire stripper, immersed in 10 mM HCI (pH 2.0; 150 ml_ per tail) and stirred for 4 hr at room temperature. Soluble collagen is separated by centrifugation at 30,000 g and 4°C for 30 minutes followed by sequential filtration through P8, 0.45 μm, and 0.2 μm membranes. Addition of concentrated NaCI in 10 mM HCI to a net salt concentration of 0.7 M, followed by 1 hr stirring and 1 hr centrifugation at 30,000 g and 4°C, precipitated the collagen. After overnight re- dissolution in 10 mM HCI the material is dialyzed against 20 mM phosphate buffer for at least 8 hr at room temperature. Subsequent dialysis is performed against 20 mM phosphate buffer at 4°C for at least 8 hr and against 10 mM HCI at 4°C overnight. The resulting MRTC solution is stored at 4°C for the short-term or frozen and lyophilized.

[0218] Production of a synthetic collagen microfiber by continuous co-extrusion. A modified wet spinning device facilitates collagen fiber production. A collagen solution (5 mg/mL in 10 mM HCI) and wet spinning buffer (WSB: 10 wt% poly (ethylene glycol) M_w = 35000, 4.14 mg/mL monobasic sodium phosphate, 12.1 mg/mL dibasic sodium phosphate, 6.86 mg/mL TES (N-ths (hydroxymethyl) methyl-2-aminoethane sulfonic acid sodium salt), 7.89 mg/mL sodium chloride, pH 8.0) are extruded with a dual syringe pump (Harvard Apparatus, Holliston, MA). The collagen solution emerges through a 0.4 mm inner diameter blunt-tipped needle into the center of a vertical tube (1.6 mm inner-diameter x 1 m long fluoropolymer tubing) at 0.08 mL/min. Wet spinning buffer simultaneously advances through a bubble trap and down the fluoropolymer tube at a rate of 1.0 mL/min. As it exits the extrusion needle, the collagen coagulates into a gel- like fiber and is carried downward by the WSB stream. Upon emergence from the fluoropolymer tube, the fiber enters a 2 meter-long rinsing bath of 70% ethanol in water. Continuous fiber is produced and collected by winding it out of the rinsing bath onto segments of polyvinyl chloride (PVC) pipe that rotates and translates automatically. After spinning, the fiber is placed in fiber incubation buffer (FIB: 7.89 mg/mL sodium chloride, 4.26 mg/mL dibasic sodium phosphate, 10 mM Tris, pH = 7.4) [48] at 37°C for 48 hr. Fiber is incubated directly on the PVC pipe segments used for collection. Subsequently, pipe segments containing continuous fiber are rinsed in ddH₂0 for 15 min before drying and collecting the fiber under tension with an automated system.

[0219] Fabrication of anisotropic protein polymer lamellar elastic structures from synthetic collagen fibers within an elastin-like protein polymer matrix. Several APPLES are designed and fabricated by winding defined collagen fiber layouts onto rectangular frames and implementing the transition temperature fiber embedding and lamination protocol illustrated in Figure 12. To arrange the fiber with the desired spacing and angle, the frame translation speed, translation distance, and rotation speed are computed with a MATLAB script. An automated linear actuator (Velmex, Inc, Bloomfied, NY) and a DC gear motor translate and rotate the frames. After winding, each fiber layout is transferred onto a sheet of ultrasoft polyurethane, secured with tape, and photographed. Images from at least three regions of each layout enabled the measurement of average fiber spacing and angle.

[0220] Solutions of LysB10 are prepared at 10 wt% concentration in ice-cold ddH₂O. Argon is bubbled through the solutions, followed by centhfugation at 4°C and 50Og for 5 min to remove bubbles. To embed the fiber layouts, precision 50 μm thick plastic shims (Precision Brand, Inc., Downers Grove IL) are placed around the layouts, and all embedding materials are cooled to 4°C. The LysB10 solution is distributed over the fibers and a sheet of acrylic is pressed on top of the solution. The fibers and the LysBIO solution are located within a 50 μm space, sandwiched between the acrylic sheet and polyurethane base that are separated by precision shims. The embedding assembly is left at 4°C for one hour to allow the LysBIO solution to hydrate the fiber layout, followed by transfer of the assembly to 37°C incubator for 30 min. When the polyurethane and acrylic sheets are peeled apart, the fiber layout remained embedded in a solid film of LysBIO, adherent to the polyurethane base. After a 5-minute incubation in 37°C ddH₂O, the fiber-reinforced film can be separated from the polyurethane base.

[0221] Each APPLES design comprises a stack of ten 40 μm thick layers. The eight central layers contained embedded fiber while the top and bottom layers contained only LysBIO. Ten-layer stacks are covered with plastic wrap to prevent drying, cooled to 4°C for 12 hr, and transferred to 37°C for 30 min to facilitate interlamellar bonding with formation of a cohesive sheet. The sheet is removed, rinsed in 37°C PBS for 30 min, and crosslinked in 0.5% glutaraldehyde in PBS for 24 hr at 37°C. Vigorous shaking in PBS for 6 hr at 37°C with three buffer changes serves to remove excess glutaraldehyde.

[0222] Analysis of collagen fiber orientation and volume fraction. Six different fiber layouts are analyzed. To study fiber spacing, we prepare samples with fiber orientations of 15° and spacing equal to 0.15 mm, 0.45 mm, or 1.3 mm, as well as a sample without fiber. Fiber volume fraction is calculated from measurements of average fiber spacing taken from digital photographs of the fiber layouts, the fiber diameter of 40 μm, and the thickness of the multilamellar sheet (10 layers x40 μm/layer = 400 μm). The fiber spacings above corresponded to 17%, 7%, 3%, and 0% fiber volume fractions. The effect of fiber orientation is demonstrated by setting fiber volume to 17% and adjusting fiber orientation to 0°, 15°, or 90°. The primary orientations of the fiber layouts are measured from digital photographs of the fiber layouts using the Inverse Fast-Fourier Transform tool from ImageJ software [89].

[0223] Digital volumetric imaging of fiber orientation and packing density. For digital volumetric imaging (DVI), fiber is first conjugated to tetramethyl rhodamine isothiocyanate (TRITC) [90]. Fiber is wound about a PVC pipe segment placed inside a larger pipe. This arrangement creates a 100 mL annular volume in which 20 - 4O m of fiber can be reacted without tangles or breaks. A 1 mg/mL solution of TRITC in DMSO is added to a 0.1 M sodium carbonate solution to a concentration of 0.05 mg/mL. This solution is added between the pipe segments and stirred for 12 hr at 4°C, after which the fiber is rinsed four times with ddH₂O for 2 hr and for 5 min with 70% ethanol, and then dried as it is transferred to a second pipe segment (Figure 13). Composite sheets re-enforced with TRITC-conjugated fiber are prepared for DVI by serial dehydration in ethanol and xylene. Samples are embedded in Spurr's epoxy modified with an optical opacifier, Sudan Black B, and imaged with a DVI Microimager (Microscience Group, Inc. Redwood City, CD)

[0224] Transmission electron microscopy. Transmission electron microscopy is used to investigate the ultrastructure of the composite. Samples of the APPLES are rinsed twice in 0.1 M cacodylate buffer (pH 7.4), fixed (2.5% glutaraldehyde in 0.1 M cacodylate buffer, pH 7.4) for 90 minutes, washed in 0.1 M cacodylate water and then dH₂0, postfixed with 1 % osmium tetroxide for one hour, and stained en bloc with filtered 2% uranyl acetate in 50% ethanol. Samples are then dehydrated with an ethanol series, pre-infiltrated with propylene oxide, and embedded in Spurr's epoxy. Using an RMC MT-7000 ultramicrotome and a diamond knife, ultrathin sections (60-80 nm) were cut to display planes perpendicular to the x, y, and z directions (Figure 14). For the TEM analysis, these planes are referred to as the x, y, and z sections. Sections were post- stained with 3% uranyl acetate and 2% lead citrate, and examined and photographed using a JOEL JEM-1210 TEM at 90 kV.

[0225] Figure 14 is a schematic illustrating a first material 100 comprising an elastin- mimetic protein formed into a film and a second fibrous material 200 that is supported by the film 100, and in this embodiment the second material 200 is collagen fibers that are embedded in the first material film 100. The fibers are aligned in two preferential directions to define a fiber angle θ, relative to an axial direction, x.

[0226] Scanning electron microscopy. Samples for scanning electron microscopy are cut with a razor blade to expose x and y sections of the sheet. The z face, or the top of the sheet, is also imaged. Samples are critical point dried (E3000, Energy Beam Sciences, Inc., East Granby, CT), sputter coated with gold (Emscope SC-500, Emitech, Kent, England), and examined and photographed with a DS-150 F scanning electron microscope (Topcon Co., Tokyo, Japan) operated at 15 kV.

[0227] Mechanical responses of anisotropic protein polymer lamellar structures. The effects of fiber orientation and volume fraction on APPLES mechanical properties are evaluated under uniaxial tension. Conditioning and tensile analysis are conducted on a dynamic mechanical thermal analyzer DMTA V (Rheometric Scientific, Inc., Newcastle, DE) with a 15 N load cell in the inverted orientation, so that samples could be immersed in a jacketed beaker of 37 ⁰C PBS. For each fiber layout, 5 to 7 replicates of the APPLES are tested. Samples that are 4 mm in width and 0.4 mm in thickness are mounted on the mechanical testing equipment with gauge lengths of 12 to 13 mm. Force is applied in the x direction indicated in Figure 3.2. Samples are extended to 8% strain for 16 cycles and then to 30% strain. Samples that did not fail when stretched to 30% are remounted on a miniature materials tester, the Minimat 2000 (Rheometric Scientific Inc., Newcastle, DE), and tested to failure. All tests are performed at a rate of 5 mm/min. For each APPLES design, resilience is calculated from the 8% strain data by dividing the area beneath the loading curve by the area beneath the unloading curve and multiplying by 100%, and reported as the mean and standard deviation from all samples. To characterize fiber failure modes, samples are treated after testing with Van Gieson's stain to distinguish the collagen fiber and photographed.

[0228] Suture retention strength of protein fiber composite sheets. Sutures (Prolene 4- 0) are passed through 4 mm square APPLES segments at a distance of 2 mm from the sheet edge. Samples had a fiber orientation of 15° and volume faction of 17%. The APPLES is clamped in the DMTA, and the suture is fastened to the actuating arm of the instrument and pulled at a rate of 1 mm/sec. The maximum force measured before the suture tore out of the sheet is recorded as the suture retention strength, reported in grams-force (g-f). For seven samples, the suture is pulled in the y direction, and for four samples the suture was pulled in the x direction.

[0229] Statistics. Tests for statistically significant differences between the means of two groups are conducted with the Student's f-test (two-tailed, homoscedastic). Tests between three or more groups are conducted with the one-way ANOVA followed by the Tukey HSD test.

[0230] RESULTS and DISCUSSION: Rigid fiber-reinforced composites are widely known to offer a combination of high stiffness, strength, and toughness at low weight. Flexible composites display an alternative property set. This class of materials has long been applied as steel or Kevlar-reinforced rubber composites common in pneumatic tires, and is under investigation for applications such as morphing aircraft wings, flexible body armor, and stretchable electronics [91 -93]. Advanced passive mechanical properties associated with flexible composites includes an enormous usable deformation range, the propensity to store and return strain energy, to limit crack propagation and fatigue, to tailor mechanical anisotropy and Young's Modulus [94], and to engineer nonlinear mechanical responses [95]. The APPLES presented here provide a protein-based biomaterial platform that incorporates the mechanical characteristics of flexible composites.

[0231] Structural analysis of a protein polymer composite. Multilamellar sheets are initially translucent and colorless, but acquire a slightly tan color after glutaraldehyde crosslinking. Fiber angle and spacing are measured from photographs of the fiber layouts and fiber volume fraction calculated by assuming an average fiber diameter of 40 μm, a sheet thickness of 400 μm, and the presence of eight fiber-reinforced layers in the composite sheet (Table 5). Observed fiber orientation and spacing based on image analysis of 2-D photographic images are close to expected values and consistent with the three-dimensional geometry of the fiber layout reconstructed from digital volumetric imaging (Figures 15 and 16). Transmission electron microscopy reveal that fibers are comprised of axially aligned, D-pehodic fibrils resembling native collagen (Figure 17). High resolution SEM images demonstrate that composite sheets are uniformly bonded without evidence of voids or delamination between individual 40 μm membrane layers within each sheet (Figure 18). Figure 18A is a schematic of a multilayer material 300 comprising a plurality of layers 310 , wherein each layer comprises collagen fibers 200 suspended in an elastin-mimetic material 100 formed into a film.

[0232] Uniaxial mechanical responses of composite sheet. Samples are initially preconditioned to 8% strain to obtain stable mechanical responses. Preconditioning enhances resilience, and introduces a small amount of residual strain (Figure 19). The effect of fiber fraction and orientation is apparent in average tensile responses (Figures 20, 21). When fiber is oriented parallel or at 15° to the loading direction, stress-strain curves are linear between 2 and 12 or 14% strain. The initial non-linearity up to 2% strain reflects artifact due to sample misalignment and slack due to residual strain induced by preconditioning. Young's Modulus is measured from a linear fit between 4 and 10% strain. The slope of the stress-strain response above a strain of 12 to 14 % decreased for samples in which fibers were orientated at 0° or 15°. This transition reflected yielding of the fiber network at average stresses 1.31 to 2.87 MPa, depending on the fiber layout. Substantial hysteresis between the loading and unloading curves is observed upon yielding of the fiber network due to irrecoverable loss of mechanical energy in this loading cycle (Figures 2OB and 21 B). Similar yielding behavior is not observed in designs without fiber, or when fibers are perpendicular to the loading direction (Figure 2OB (open circles)). Mechanical parameters including yield stress, Young's Modulus, ultimate tensile stress, and strain-to-failure for all fiber layouts are summarized in Table 6. Designs with greater fractions of fiber, or with fiber parallel rather than perpendicular to the loading direction, demonstrate enhanced resilience (Figure 22). Increasing fiber fraction (e.g., second material volume fraction) also elevated the Young's Modulus and yield stress but did not significantly alter ultimate tensile stress (Figure 23). Layouts with fiber orientation more closely aligned to the loading directions (e.g., 0° and 15° fiber orientation) also display increased Young's Modulus (Figure 23B) and ultimate tensile strength (Figure 23E). Collectively, mechanical analysis indicates the capacity to increase Young's Modulus by up to a factor of five, and enhance resilience from 53.1 ± 1.4% to 76.1 ± 2.9%. Resilience, or elastic efficiency, measures the energy stored and returned during a loading-unloading cycle. The high resilience of many tissues contributes to their ability to transmit energy and resist fatigue during cyclic loading, and native collagen and elastin are both highly resilient [81].

[0233] Sample analysis after rupture reveals four modes of failure in tension related to fiber volume fraction and orientation (Figure 24). A primary failure of the elastin-like protein matrix is observed when the construct is fabricated without fibers or fibers that are oriented such that they are not the primary load bearing element of the sample (e.g., 90° orientation). Typically, a 4.08 ± 0.80 MPa stress and 314 ± 26% strain results in the failure of the elastin matrix analogue. When fibers oriented perpendicular to the loading direction are added to the elastin-like matrix, the ultimate tensile stress and the strain-to- failure decreased to 1.99 ± 0.28 MPa and 184 ± 38%, respectively, although the stress- strain curve are similar prior to tensile failure (compare Modes 3 and 4 in Figure 24A). Layouts with 7 or 17% fiber fraction oriented at 0 or 15° to the loading direction demonstrate the fiber network yielding behavior discussed above, followed by strain-to- failure values of 20 to 50% (see Mode 1 ). The relatively lower strain-to-failure in these samples occur because the fiber network yielded in one or two discrete locations. After yielding, it is likely that local strains in these regions are higher than the overall sample strain, resulting in tensile failure in these regions. Voids left in the elastin-like protein matrix after fibers yield and debond from the matrix would also contribute to tensile failure [96]. A distinct failure mode is observed when fiber fraction is lowered to 3% (Mode 2). In this case, the fiber network demonstrates several yielding events with increasing strain before failure at strains of 182 ± 135%. This pattern occurs when the stress transfers from the ruptured fibers to the matrix is low enough that the matrix itself does not fail, and is known to occur at low fiber volume fractions [96]. Suture pulled out of the protein polymer flex composite in the y direction at 124 ± 8 gf, and in the x direction at 170 ± 36 gf.

[0234] Aqueous solutions of the thblock elastin-mimetic protein polymer are capable of a sol-gel transition, which facilitates incorporation of a fiber layout into a single 40 μm thick membrane and subsequent bonding of a multilayer membrane stack. The fiber layouts in this example comprise closely spaced fibers with two predominate orientation angles, resembling several native tissues including the linea alba of the anterior abdominal wall [97], small intestinal submucosa [98, 99], and the annulus fibrosis of intervertebral discs [100]. The laminated geometry resembles cell sheet tissue engineering methods investigated by others [86, 101] and provides the capacity to incorporate living cells at controlled spatial intervals through the sheet thickness. The use of a recombinant protein matrix and synthetic protein microfibers affords control over compliance, resilience, strength, and anisotropy not available with cell sheet tissue engineering strategies and obviates the long culture time required for the development of cell-secreted collagen layers.

[0235] The development of substitute extracellular matrix proteins, and technologies to produce the desired geometries in three dimensions, provides the capability of improved non-living and tissue-engineered implants. The flexible, elastic protein polymer fiber composites developed in this example represents an important step for that application. Through the development of a fiber embedding and lamination process, composite sheets with controlled mechanical anisotropy, and increased modulus, yield strength, resilience, and suture retention strength are created.

[0236] Example 3: Microcrimping Fibers

[0237] Collagen fibers exhibit a waviness, or crimp, in a diversity of tissues including tendon [102], ligament, intestine, blood vessel [103], heart valve leaflet, intervertebral discs, the intra-articular disc of the temporomandibular joint [104], and others [100]. The wavelength of crimp varies between 10 to 200 μm, and the shape has been characterized as a planar zig-zag [100], a planar sinusoid [103], and a 3D helix [105, 106]. In several instances, researchers have observed that crimp disappears as soft tissues are stretched, and simultaneously the tissue transitions from low to high stiffness [82, 102, 107]. Crimp is thought to represent redundancy in the collagen network, which only partially contributes to the overall resistance to deformation at low stretch levels. At higher stretch, collagen fibers un-crimp and/or rotate into alignment with the direction of tension, stiffening the tissue by bearing an increased share of the load. Consequently, crimping is one of the features of collagen fiber architectures that allow tissues to be both compliant and strong. This combination contributes to significant biomechanical phenomena such as the efficient opening and sealing of heart valves [83], the propensity of tendon to smoothly absorb load, and the compliance of arteries.

[0238] In Example 2, protocols to fabricate laminated composites of collagen fiber embedded in elastin-mimetic protein are provided. Although that fabrication scheme increases strength and resilience and presents the capacity to tailor mechanical anisotropy, the composites do not display the mechanical transition point behavior observed in many native tissues. Although the laminated sheets differ from native protein fiber networks in several respects, the lack of transition point is largely due to absence of crimp in the collagen fiber component.

[0239] The biomechanical significance of collagen fiber architectures motivates the development of a soft tissue substitute replicating crimp structure and function. In this example we present a method to fabricate microchmped sheets of collagen fiber embedded in elastin-mimetic protein.

[0240] MATERIALS and METHODS: Synthesis of a recombinant elastin-mimetic triblock protein polymer. Genetic engineering, expression, purification, and characterization of the elastin-mimetic protein polymer, designated LysB10, has been described elsewhere [3]. Briefly, the flanking 75 kDa endblocks of the protein polymer contained 33 repeats of the hydrophobic pentapeptide sequence [IPAVG]₅, and the central 58 kDa midblock contains 28 repeats of the elastic, hydrophilic sequence

[(VPGAG)₂VPGEG(VPGAG)₂]. The sequences between blocks and at the C terminus include the residues, [KAAK], provides amine groups for chemical crosslinking.

[0241] The protein polymer sequence is contained a single contiguous reading frame within the plasmid pET24-a, which is used to transform the E. coli expression strain BL21 (DE3). Fermentation is performed at 37°C in Circle Grow (QBIOgene) medium supplemented with kanamycin (50 μg/mL) in a 100 L fermentor at the Bioexpression and Fermentation Facility, University of Georgia. Cultures are incubated under antibiotic selection for 24 hr at 37°C. Isolation of the LysB10 comprises breaking the cells with freeze / thaw cycles and sonication, a high speed centrifugation (20,000 RCF, 40 min, 4°C) with 0.5% poly(ethyleneimine) to precipitate nucleic acids, and a series of alternating warm / cold centhfugations. Each cold centrifugation (20,000 RCF, 40 min, 4°C) is followed by the addition of NaCI to 2M to precipitate the protein polymer as it incubated for 25 min at 25°C. This is followed by warm centrifugation (9500 RCF, 15 min, 25) and resuspension of the pellet in cold, sterile PBS on ice for 10 - 20 min. After 6 to 10 cycles, when minimal contamination is recovered in the final cold centrifugation, the material is subject to a warm centrifugation, resuspended in cold sterile PBS, dialyzed, and lyophilized.

[0242] Isolation and purification of monomehc collagen. Acid-soluble, monomehc rat- tail tendon collagen (MRTC) is obtained from Sprague-Dawley rat tails following Silver and Trelstad [47]. Frozen rat tails (Pel-Freez Biologicals, Rogers, AK) are thawed at room temperature and tendon is extracted with a wire stripper, immersed in 10 mM HCI (pH 2.0; 150 ml_ per tail) and stirred for 4 hr at room temperature. Soluble collagen is separated by centrifugation at 30,000 g and 4°C for 30 minutes followed by sequential filtration through P8, 0.45 μm, and 0.2 μm membranes. Addition of concentrated NaCI in 10 mM HCI to a net salt concentration of 0.7 M, followed by 1 hr stirring and 1 hr centrifugation at 30,000 g and 4°C, precipitates the collagen. After overnight re- dissolution in 10 mM HCI the material is dialyzed against 20 mM phosphate buffer for at least 8 hr at room temperature. Subsequent dialysis is performed against 20 mM phosphate buffer at 4°C for at least 8 hr and against 10 mM HCI at 4°C overnight. The resulting MRTC solution is stored at 4°C for the short-term or frozen and lyophilized.

[0243] Production of a synthetic collagen microfiber by continuous co-extrusion. A modified wet spinning device facilitates collagen fiber production. A collagen solution (5 mg/mL in 10 mM HCI) and wet spinning buffer (WSB: 10 wt% poly (ethylene glycol) Mw = 35000, 4.14 mg/mL monobasic sodium phosphate, 12.1 mg/mL dibasic sodium phosphate, 6.86 mg/mL TES (N-tris (hydroxymethyl) methyl-2-aminoethane sulfonic acid sodium salt), 7.89 mg/mL sodium chloride, pH 8.0) are extruded with a dual syringe pump (Harvard Apparatus, Holliston, MA). The collagen solution emerges through a 0.4 mm inner diameter blunt-tipped needle into the center of a vertical tube (1.6 mm inner-diameter x 1 m long fluoropolymer tubing) at 0.08 mL/min. Wet spinning buffer simultaneously advanced through a bubble trap and down the fluoropolymer tube at a rate of 1.0 mL/min. As it exited the extrusion needle, the collagen coagulates into a gel- like fiber and is carried downward by the WSB stream. Upon emergence from the fluoropolymer tube, the fiber enters a 2 meter-long rinsing bath of 70% ethanol in water. Continuous fiber is produced and collected by winding it out of the rinsing bath onto segments of polyvinyl chloride (PVC) pipe that rotates and translates automatically. After spinning, the fiber is placed in fiber incubation buffer (FIB: 7.89 mg/mL sodium chloride, 4.26 mg/mL dibasic sodium phosphate, 10 mM Tris, pH = 7.4) [48] at 37°C for 48 hr. Fiber is incubated directly on the PVC pipe segments used for collection. Subsequently, pipe segments containing continuous fiber are rinsed in ddH₂0 for 15 min before drying and collecting the fiber under tension with an automated system.

[0244] Before microchmping, the fiber is arranged into dense parallel sheets by winding about rectangular frames. The frames are rotated at 40 rpm by a DC gearmotor and translated at 7 mm/min by an automated linear actuator (Velmex, Inc, Bloomfied, NY). Digital photographs of the array of fibers on the frame indicate an average fiber spacing of 190 ± 10 μm. Two additional fiber layers are wound onto the frame, over the first layer, to reduce the average fiber spacing to 63 μm. For single-fiber mechanical testing, only ten fibers are crimped at a time. In this case the fibers were separated by 1 to 2 mm, so that after crimping individual fibers can be readily obtained.

[0245] Fabrication of defined patterns of parallel microridge arrays in flexible template membranes. Three microridge profiles patterns are designed and tested for the microcrimping system (Figure 25). The triangular microridge template (Figure 25A) is fabricated by spin coating a uniform layer of SU-8 2050 negative photoresist

(MicroChem Corp, Newton, MA) onto a silicon wafer. A photomask is applied over the photoresist, and the SU-8 is exposed to 45° inclined ultraviolet light for 70 seconds. Exposure to inclined UV through the photomask crosslinks SU-8 in a 3D pattern consisting of an array of parallel micro-trenches with triangular cross-sections. After removal of the unexposed SU-8 with developer, a 1 μm parylene coating is vapor deposited. Polyurethane solution (PMC 121-30® and PMC® 780, Smooth-On, Inc., Easton, PA) is cast over the parylene-coated SU-8 micro-trench mold and allowed to crosslink for 24 hr, yielding the flexible triangular microridge template (Figure 26).

[0246] Rectangular (Figure 25B) and chamfered rectangular (Figure 25B) profiles are generated by an alternative process (Figure 27). For the generation of rectangular microridges, a layer of positive photoresist (AZ® 4620, Clahant Corp.) is patterned into repeating strips on a silicon wafer, which served as a mask for inductively coupled plasma etching. Inductively coupled plasma etching generates rectangular micro- trenches in the regions of the silicon wafer not shielded by photoresist, at an etch rate of 0.6um/min following the Bosch process. After removal of the photoresist with an acetone rinse and cleaning of the wafer in piranha solution, a layer of parylene is vapor deposited over the micro-trenched silicon wafer. Following parylene deposition, polyurethane is cast over the wafer and allowed to crosslink for 24 hr, generating the flexible rectangular microhdge template membrane.

[0247] The chamfered rectangular microridge template membrane is produced by a similar process, with the additional step of anisotropic wet etching in a KOH aqueous bath (40 wt%, 70⁰C) and an additional molding step. The anisotropic etch is performed subsequent to the inductively coupled plasma etch, converting the rectangular micro- trench geometry into the desired chamfered geometry. After etching parylene is coated over the silicon and polydimethylsiloxane (PDMS, Dow Corning Sylgard 184) is cast over the patterned silicon wafer created a negative of the desired profile. Subsequently, parylene is coated over the PDMS and PU was cast over the coated PDMS to yield the flexible chamfered rectangular microridge template. These templates are fabricated from PU of 3OA and 7OA durometer.

[0248] Method for scalable microcrimping of synthetic collagen fibers. The microcrimping system comprises a lead screw assembly, an ultrasoft smooth viscoelastic base membrane (60 OO durometer Sorbothane, Sorbothane, Inc., Kent, OH), the microridged template membrane, and a clamping assembly (Figure 28). The base membrane is fastened to the lead screw assembly and a pre-extension of 15 to 55% tensile strain is applied with the lead screw. A parallel array of collagen fiber is transferred from the rectangular winding frame onto the extended base membrane and fastened with tape. The fiber is hydrated with ddH₂0 for 15 min, excess water removed, and the flexible microridged template membrane is applied over the hydrated fiber array. When the microridged template is applied, it is manually extended to the same tensile strain as the pre-extended base membrane, and fastened to the lead screw assembly. Application of the clamping system then secures the collagen fiber between the pre- extended microridged template and the pre-extended base membrane with a normal templating force. Adjustment to the lead screw assembly relaxed the pre-extension simultaneously in both membranes and introduced the microcrimped geometry into the collagen fiber array. The system is frozen at -80⁰C for 2 hr, warmed to -20°C for 4 hr, and then the clamping assembly and the microridged template membrane are removed. The microcrimped collagen fiber array remains on the base membrane, and is transferred to a room-temperature desiccator saturated with vapor from a 25% glutaraldehyde solution to crosslink the fiber. The base membrane and microchmped fiber array remain frozen when placed in the dessicator, so that the hydrated, crimped shape of the fiber is largely held in place as the collagen began to crosslink. After one day, the fiber and the base membrane are removed from the dessicator and allowed to dry in air, yielding a dense, parallel array of microcrimped synthetic collagen fiber. For single-fiber mechanical tests, 8 mm lengths of dried fiber are removed from the base membrane with tweezers and mounted on plastic frames. Initial observations demonstrate that the amount of pre-extension tensile strain, from 15 to 55%, largely dictates the amount of deformation imparted to the fiber and thus the morphology of the microcrimp.

[0249] Formation of an elastin-like protein polymer lamella with an integrated array of microcrimped collagen fibers. A crimped fiber array, arranged on the base membrane and surrounded by 250 μm thick precision shims (250 μm, Precision Brand, Inc., Downers Grove IL), is frozen at -80⁰C. A 10 wt% solution of protein polymer is applied to the frozen crimped fiber array after it is initially purged with argon and centhfuged for 5 min at 4°C and 500 g to remove bubbles. An acrylic sheet is used to spread the protein polymer solution as a thin film, which when incubated at room temperature for 25 min gelled around and embedded the crimped fibers. This fiber-reinforced lamella is separated from base membrane after a 5 minute incubation in PBS at 37°C.

Microcrimped fiber films are crosslinked in 0.5% glutaraldehyde for 24 hr at 37°C and then rinsed in PBS for 2 hr at 37°C, which is repeated three times. This yields 80 μm- thick films of crosslinked protein polymer with an embedded array of parallel, microcrimped fiber.

[0250] Evolution of the microchmping system and microridge design. Features of the microridges, including the width, height, spacing, and the stability of the microridge are important to the microcrimping process (Figures 29 and 30). The application of a normal templating force with the clamp assembly proved to be important to maintain parallel alignment of the fibers during microcrimping. We initially fabricate the microridge template membrane from a soft PU (3OA durometer) in order to limit potential fiber damage from a stiff, inflexible ridge. However, significant microridge deformation is observed after removal of PU mold following fiber crimping (Figure 30). PU of higher durometer (70A) offered sufficient material flexibility to allow for defined membrane pre- extension along with adequate rigidity to prevent microhdge collapse upon application of a normal templating force.

[0251] Microscopic analysis of fiber crimp. Microcrimped fiber arrays are prepared for scanning electron microscopy by sputter coating with gold (Emscope SC-500, Emitech, Kent, England), and examined and imaged with a DS-150 F scanning electron microscope (Topcon Co., Tokyo, Japan) operated at 15 kV. Fiber is prepared for confocal laser scanning microscopy (CLSM) by conjugation to tetramethyl rhodamine isothiocyanate (TRITC) [90]. For TRITC conjugation, fiber is wound about a PVC pipe segment placed inside a larger pipe. This arrangement creates a 100 ml_ annular volume in which up to 40 m of fiber can be reacted without tangles or breaks. A 1 mg/mL solution of TRITC in DMSO is added to a 0.1 M sodium carbonate solution to a concentration of 0.05 mg/mL. This solution is added between the pipe segments and stirred for 12 hr at 4°C, after which the fiber is rinsed four times with ddH2O for 2 hr and for 5 min with 70% ethanol and dried in air. Fiber conjugated to TRITC is microcrimped and embedded in elastin-like protein polymer as described above. Samples are examined with an LSM 510 Confocal (Carl Ziess Microimaging, Oberkochen, Germany) using a 543 nm Helium-Neon laser and a 10x objective. Three-dimensional projections are created from stacks of 25 to 35 optical slices taken at 3 to 6 μm intervals using LSM Image Broswer software (Carl Ziess Microimaging, Oberkochen, Germany). Projections of the crimped fiber are rotated to depict the profile of the microchmp. The degree of crimp, C, of the embedding and hydrating the fiber assembly is defined as:

C = (/_C - /_S) / /_S X 100%

where /_c is the length of a line that traced the center of the crimped fiber and /_s is the straight-line distance along the path of the fiber. Alternatively, crimp is quantitatively defined as the amount of strain that may be applied to the fiber before the intrinsic properties of the fiber begin to hinder strain. In other words, crimp may correspond to how much strain it takes to uncrimp, unfold or unwind the fiber.

[0252] Mechanical analysis. Fiber lamella samples, 5 mm in width, 80 μm in thickness and with a gauge length of 12 to 13 mm, are mounted on a dynamic mechanical thermal analyzer (DMTA V, Rheometric Scientific, Inc., Newcastle, DE) with a 15 N load cell in an inverted orientation and immersed in a jacketed beaker of PBS at 37°C. Samples are oriented such that the direction of tensile stress is parallel to the embedded fibers. Three to four samples of embedded, uncrimped fiber, fiber microcrimped with 15% pre- extension, and fiber microcrimped with 30% pre-extension are tested. Samples are allowed to equilibrate in PBS for 5 min and strained to failure. Engineering stress and strain are reported. The transition point strain, the level of strain at which the material progressed from compliant deformation to high modulus deformation, is quantified for all stress-strain curves. The transition point strain is defined as the x-intercept of a straight line fit to the last 4% strain prior to sample yielding.

[0253] To assess the effect of cyclic loading on microchmp morphology, samples are examined by confocal laser scanning microscopy before and after 15 and 1000 cycles of loading to 10% strain. The failure strength of individual crimped fibers is also measured in tension. Non-embedded fibers are allowed to dry after microcrimping and crosslinking, removed from the base membrane and mounted on a plastic frame with cyanoacrylate glue. Fibers microcrimped with 15 and 30% pre-extension are evaluated (n=8, n=4 respectively). Fibers are hydrated in PBS for 2 hr, and mounted in the DMTA. The plastic frame supporting the fiber is cut away and fibers are pulled to failure. All tests are performed at a rate of 5 mm/min in PBS at 37⁰C.

[0254] RESULTS and DISCUSSION: Controlled deformation of a flexible template dictates periodic microcrimp morphology. Scanning electron microscopy indicates that after the crimping, vapor crosslinking, and drying on the base membrane, a regular crimp pattern is introduced into the fiber arrays (Figures 31 and 32). Upon pre- extension of the template to 30 or 40% beyond the resting length, the crimp comprises relatively smooth arcs, while deforming to 45 and 50% led to some sagging of the crimp peaks. At low magnification, the crimp pattern appears consistent over a scale of several millimeters (Figure 33). However, occasional irregularities at the peak of the crimp suggest that in some areas the fiber may have buckled. In addition, angular imprints are observed at contact points along the fiber, although less severe than those produced by triangular microhdge features.

[0255] Three-dimensional reconstructions from CLSM of embedded fiber lamellae demonstrates that crimp geometry is preserved after fiber embedding in the elastin- mimetic protein polymer and imaging in the hydrated state (Figure 33). A second fibrous material 400 comprising a plurality of collagen fibers is embedded in a first material 300 comprising an elastin-mimetic protein formed into a film. The induced crimp curvature depends on the level of applied membrane pre-extension. For example, the degree of crimp as measured from CLSM reconstructions is 3.1 ± 0.4 % and 9.4 ± 2.9 % for fibers crimped by membrane deformation of 15 and 30%, respectively (Figure 34). Ideally, the applied pre-extension and the degree of crimp imparted to the fiber should be equivalent; however reduced crimp might be anticipated if the fiber length is not constant during microcrimping. Specifically, fiber length may decrease during the crosslinking and drying steps. These steps are designed to crosslink the fiber while the hydrated microchmp morphology is frozen in place. However, during the 24 hr crosslinking period, the fiber arrays thawed and partly dried. The real crosslinked geometry is therefore in between the swelled, hydrated geometry and the contracted, dry geometry. If partial drying reduced fiber length, it would also reduce the degree of crimp. The hydrated geometry may not be fully restored during the embedding step if the elastin-like polymer gels before fully hydrating the fiber. The embedding step may also reduce the degree of crimp if the protein polymer lamella swells after release from the base membrane. Swelling would effectively pull a portion of the crimp out of the fiber. Notably, the microchmped fiber wavelength before embedding is 127 ± 5 μm, as observed by SEM, and the wavelength hydrated, embedded microcrimped fiber is 143 ± 5 μm, observed by CLSM. This relative increase of 13% may indicate swelling of the protein polymer lamella that could reduce the degree of crimp.

[0256] Microcrimping alters the mechanical response of fiber-reinforced elastin-like matrix composites (Table 7). Composites display a transition point strain between low (e.g., compliant) and high modulus regimes at an extension that is dictated by the degree of fiber crimp (Figure 35). The observed level of strain at the calculated transition point is 1.1 ± 0.2 %, 4.6 ± 0.9 %, and 13.3 ± 0.7 % for fibers that were non- crimped or had been subject to a pre-extension of 15 % or 30%, respectively. Non- crimped fibers display a transition at very low strain due to imperfect sample loading and alignment. Notably, the strain at which each transition occurred is very close to the measured degree of crimp (0% vs 1.1 %; 3.1 % vs 4.6%; 9.4% vs 13.3%). Mechanical testing also demonstrates that the crimp structure is not lost during cyclic loading. Samples subjected to 15 and 1000 loading cycles do not demonstrate a notable change in the degree of crimp (Figure 36; 10.2 ± 2.0 %, 9.4 ± 2.9 %, and 8.8 ± 1.4% for zero, 15, and 1000 cycles).

[0257] Fibers are weaker after crimping with an observed failure strength of 2.2 ± 0.5 and 2.1 ± 1.2 g-f for fibers crimped at 15 or 30% pre-stretch, as compared to 8.8 ± 1.7 g- f for non-crimped fibers. Although crimped fiber composites are weaker than their corresponding counterparts composed of non-crimped fibers, it is significant that composite membranes containing fibers crimped at 30% pre-stretch displayed an ultimate tensile strength of 2.08 ± 0.73 MPa, exceeding the strength of many native tissues, such as human urinary bladder (270 ± 140 kPa) [108], pulmonary artery (385 ± 45 kPa) [109], and aorta (1.72 ± 0.89 MPa) [110].

[0258] Oriented arrays of synthetic collagen fibers are created with a microcrimped structure similar in scale to naturally occurring collagen crimp. After embedding microcrimped collagen fiber arrays in a matrix consisting of recombinant, elastin-mimetic protein polymer film, crimp geometry is largely retained. The designed composites demonstrate transition points between low and high modulus regions at a strain that can be predicted by the degree of crimp. The observed mechanical responses for this acellular tissue analogue is similar to that observed for a number of native tissues.

[0259] Example 4: Medical Device - Artificial Blood Vessel

[0260] This example provides an artificial blood vessel from elastin-mimetic protein polymer reinforced with collagen fiber.

[0261] The lack of a clinically successful replacement for the diseased small-diameter artery represents a challenge and opportunity for the fields of biomaterials and tissue engineering. Small to medium (<4 to <7 mm) prosthetic vascular grafts occlude due to perianastomotic intimal hyperplasia and surface thrombogenicity. lntimal hyperplasia, the formation of pannus tissue that narrows the lumen of the graft, is driven in part by a compliance mismatch between stiff prosthetics and compliant native artery. Disrupted flow and shear stresses are also driving factors [1 11]. In this example, we develop compliant vascular grafts from recombinant elastin-mimetic protein polymers, reinforced with collagen fiber, as a platform from which to address these challenges.

[0262] Improved small-diameter vascular graft technology has been broadly sought and reviewed [112-114]. Research may be categorized as that related to alternative synthetic materials, decellularized allo- and xenogenic tissues, and cell-assembled or tissue engineered matrices. Synthetic materials, especially improved polyurethane, remain an area of interest despite set backs related to in vivo surface chemical modification, hydrolysis, and oxidative biodegradation encountered by early polyurethane formulations [114]. New polyurethanes exhibit improved biostability, and drug release strategies or endothelial cell seeding may lower the thrombogenic potential of those grafts. However, concerns remain regarding the carcinogenic potential of degradation products [114], delay of endothelialization caused by drug elution, and the propensity for infection that accompanies any long-term synthetic implant.

[0263] Decellulahzed tissues modified for vascular conduits have included vascular tissues such as human umbilical vein [115-117] and bovine carotid artery [118, 119], and the adaptation of non-vascular tissues, in particular porcine small intestinal submucosa (SIS). Enzymatic and detergent extraction of cells followed by glutaraldehyde crosslinking has been applied to vascular conduits to prevent antigenicity and biodegradation. However, poor patency rates and handling characteristics have limited the use of both human umbilical vein and bovine carotid artery. Tissue heterogeneity, incomplete cell extraction, biodegradation, and the potential risk of viral transmission from animal tissue may impede the application of decellulahzed tissues.

[0264] Technologies for cell-assisted matrix assembly have progressed, although many challenges remain. Those protocols begin with cells seeded on tubular biodegradable polymer scaffolds, cells suspended in molded biopolymer gels, or cell sheets rolled into tubes. The primary mechanical goal has been the strength to support suturing and arterial pressure levels. To this end, scaffolds were subject at least to 7 - 16 weeks of culture and maturation, yielding burst pressures ranging form 800 - 3500 mm-Hg [86, 88]. Although strong, these versions of the grafts were much stiffer than native artery. In this regard, arterial mechanics depend largely upon the amount and microstructure of collagen, elastin, and smooth muscle cells [120]. Elastin is credited with contributing elastic recoil and compliance to cardiovascular tissues, although more specifically the collagen and elastin fiber networks may both need to be present with the correct complementary microstructure [82]. Vascular tissue engineers have thus investigated the cell-assisted assembly of elastin with mixed success [86-88], and evidence of organized, concentric sheets of elastin and the resulting resilience and compliance has not been reported. In addition to mechanical issues, cell-assisted matrix assembly technologies must overcome challenges related to the immunologic challenges of allogenic cells, as well as scale-up and quality issues associated with long incubation times.

[0265] Several alternative fabrication technologies have been explored in small- diameter vascular graft research to address strength and compliance issues, include filament winding, electrospinning, molding, and sheet wrapping. Filament winding, in which a filament is wrapped about a rotating, translating inner mandrel, allows the compliance and strength of the graft to be adjusted by modulation of the filament angle and density [121 , 122]. Like conventional textile grafts, a filament wound structure is permeable and must be filled or coated with a second material. One potential disadvantage of the wound filament technique is that as the graft diameter expands under pulsatile flow, it may contract axially, repeatedly stressing the anastomosis sites [32]. Balancing high compliance and kink resistance with wound filament structures also presents a challenge because circumferentially oriented fibers prevent kinking but restrict compliance while fibers aligned more closely to the axis of the graft do the opposite [121].

[0266] In the electrospinning process, a high strength electric field pulls a jet of a charged polymer solution from an extrusion needle, through an air gap, and onto a grounded collecting target. The solution evaporates as it travels through the air gap, transforming the jet into a solid micro- or nanofiber that collects on the target as a nonwoven fibrous mat. Tubular structures for vascular grafts have been created by electrospinning onto a rotating, translating mandrel. Fiber orientation may be achieved through an expanding library of strategies related to motion of the electrospinning target or shaping of the electric field [123]. In principle, adjustment of fiber orientation should allow control over graft compliance, similar to filament winding. Like other textile techniques, the tubular mesh is highly permeable and must be coated or sealed with a second material. Electrospun tubes have also been reinforced with wound filament, increasing the bursting strength [124]. Depending upon fabrication conditions, electrospinning can be a slow and inefficient process, even with multi-jet spinning heads [125]. Groups have electrospun elastin or elastin-mimetic recombinant proteins to serve as one element in an arterial substitute [4, 124, 126-129]. However, another of the seemingly most appropriate materials for a protein-based vascular graft scaffold, pure collagen, has not been electrospun without the use of solvents that denature the protein [54]. Collagen may be blended with polymers or biopolymers and electrospun from non- denaturing solvents, although these fibers may not have the strength of pure collagen [128, 130].

[0267] Polymers and some biopolymers such as collagen, fibrin, and elastin can be molded to create tubular structures. Changing the wall thickness of molded tubes results in limited compliance adjustment. Compliance and strength of molding structures has also been modulated by controlling the orientation of fibril networks that develop during the molding process [131] and excimer laser ablation to generate arrays of pores [132]. The desired combination of compliance and strength can also be tuned by surrounding a soft, elastic tube with one or more stiffer tubes [133].

[0268] Sheet wrapping approaches are clearly required for grafts fabricated from flat materials such as cell layers [134] and SIS. That technique also provides the opportunity to apply a variety of 2D fabrication techniques, and mimic the laminar structure of the native vessel wall. Automated wrapping devices may increase control and repeatability in those constructs [101]. Tubular elastin scaffolds, extracted from porcine carotid arteries, have been wrapped with SIS [135]. The SIS wrap enhanced burst pressure and suture retention, and although the compliance was not reported numerically the structure was visually similar to native artery under pulsatile conditions.

[0269] Provided herein is a sheet wrapping process that has some characteristics of filament winding. Sheets of the recombinant elastin-mimetic protein, LysB10, are reinforced with oriented collagen fiber described in Example 1 , and wrapped to create a tubular structure. This method created a multi-layer tube reinforced by helical fiber arrays with controlled angle and spacing.

[0270] MATERIALS and METHODS: Fabrication of a small diameter vascular graft. Synthetic collagen fibers are arranged into parallel arrays, embedded within a thin membrane of a recombinant elastin analogue, and rolled into multilayered tubes (Figure 37). Synthetic collagen fiber is wet spun from rat-tail tendon collagen, as described in Example 2, about a rectangular frame rotated by a DC gearmotor and translated by an automated linear actuator (Velmex, Inc, Bloomfied, NY). Rotation and translation speeds are adjusted to control fiber spacing. Fibers undergo vapor phase glutaraldehyde crosslinking by placement in a desiccator containing a 25% (w/v) glutaraldehyde solution for 24 hrs. Two fiber arrays are then transferred to a glass plate and secured with tape. A protractor beneath the plate is used to align the two arrays to a desired fiber angle or orientation.

[0271] The elastin-mimetic protein polymer, LysB10, is prepared as described in Example 3. Solutions of LysB10 are prepared at 10 wt% concentration in ice-cold ddH₂O. Argon is bubbled through the solutions, followed by centrifugation at 4°C and 500 g for 5 min. To embed the fiber layouts, precision 130 μm thick plastic shims (Precision Brand, Inc., Downers Grove IL) are placed around the layouts, and all embedding materials are cooled to 4°C. The LysB10 solution is distributed over the fibers and a sheet of polycarbonate is pressed on top of the solution. The fibers and the LysBIO solution are located within the 130 μm space, sandwiched between the polycarbonate sheet and a glass plate that are separated by precision shims. The embedding assembly is incubated for one hour at 4°C, followed by a 20 min incubation at room temperature. The glass and polycarbonate are pulled apart and the film is separated and trimmed to 5 by 8 cm. The polycarbonate and glass plates are separated affording a 100 μm thick fiber-reinforced protein polymer film.

[0272] A 5 x 8 cm film is rolled about a 4 mm diameter Teflon tube to form a 5 cm long, six-layer tube, which is then wrapped in a thermoplastic film. The assembly is incubated at 4 ⁰C overnight to promote interlayer bonding, and then centhfuged at 200 g and 4°C for 5 min to remove trapped air bubbles. The assembly is incubated at 37 ⁰C for 180 min, detached from the Teflon mandrel, and hydrated in 37°C PBS for 30 min. Constructs are then thermally annealed at 60⁰C in PBS for 4 hrs. All constructs are cross-linked in 37°C PBS containing 0.5 % (w/v) glutaraldehyde for 24 hrs and rinsed for 12 hrs in PBS (Table 8).

[0273] Measurements of vascular graft fiber orientation, spacing and wall thickness. The spacing and orientation of the synthetic collagen fibers are measured from photographs of planar fiber arrays prior to embedding in protein polymer (Figure 37B). Orientation is measured from digital photographs of the fiber layouts using the Inverse Fast-Fourier Transform tool from ImageJ software [89]. After completing vascular graft fabrication, samples are stained with Van Gieson for 5 min, rinsed, and photographed. After exposure to Van Gieson strain, the organization of collagen fibers can be observed on the exterior of the graft. Additionally, three rings are sectioned from each graft, photographed, and wall thickness measured at six points around the circumference of each ring.

[0274] Graft pressure-diameter responses and burst pressure. Vascular graft compliance and burst pressure are evaluated using the system diagrammed in Figure 38. Grafts are positioned vertically in an acrylic box and submerged in 37°C PBS. As grafts are inflated with PBS supplied by a syringe pump (Harvard Apparatus, Holliston, MA) at 4 mL/min, a 3CCD camera (Dage-MTI, Michigan City, IN) with a 10x macro video zoom lens (Edmund Optics, Barrington, NJ) records video at 30 frames per second and pressure is recorded with a pressure transducer (WIKA, Lawrenceville, GA). A PC equipped with data and image acquisition cards (PCI-1405 and PCI-6220, National Instruments, Austin, TX) acquires the video and pressure data. A Labview program synchronized the video and pressure data, collecting video frames and sampling the corresponding pressure measurements at 30 Hz. A MATLAB routine is used to quantify the initial graft outer diameter (D₀) and the inflated diameter (D) from every video frame and calculated the percent change in diameter [DID₀) corresponding to each pressure measurement. Each graft is preconditioned with 20 inflations to 250 mm Hg and video taken of the 21 st inflation. Grafts are then inflated to failure while video and pressure data are recorded. Compliance (C), the percent change in outer diameter (DID₀) per 100 mm Hg of applied pressure, is calculated as:

C = 1 / b - 100 (1 )

where b is the slope of a line fit to the pressure vs. DID₀ curve between 80 and 120 mm Hg.

[0275] Suture retention strength. Two 4 mm-long segments are cut from each graft and then sectioned into thirds around the circumference to generate three 4 x 4 mm squares of material. Samples are mounted in a dynamic mechanical thermal analyzer (DMTA V, Rheometric Scientific, Inc., Newcastle, DE) with a 15 N load cell in the inverted orientation, so that samples can be immersed in a jacketed beaker filled with PBS at 37 ⁰C. Prolene suture (4-0) is passed through the sample and fastened to the actuating arm of the instrument. The sample is oriented with the suture 2 mm away from the edge, and pulled parallel to the central axis of the graft. The suture is pulled at a rate of 1 mm/sec and the maximum force measured before the suture tore out is recorded in grams-force (g-f). Five to six samples from each design are tested and the data expressed as mean ± standard deviation.

[0276] Scanning electron microscopy. Samples for scanning electron microscopy are cut with a razor blade to expose the lumenal surface and the cross-section of the graft wall. Samples are critical point dried (E3000, Energy Beam Sciences, Inc., East

Granby, CT), sputter coated with gold (Emscope SC-500, Emitech, Kent, England), and examined and photographed with a DS-150 F scanning electron microscope (Topcon Co., Tokyo, Japan) operated at 15 kV.

[0277] RESULTS and DISCUSSION: Development of a fabrication scheme to create vascular graft composites with controlled fiber orientation and spacing. The fabrication scheme facilitates the generation of vascular grafts of varying fiber content and architecture. The dimensions of exemplified grafts, including fiber spacing and orientation are summarized in Table 9 (Figures 39-41). Although the angle and spacing measured from the fiber layout are close to the nominal values, photographs of the stained prototypes suggests more variability in the final fiber arrangement. Irregularities in the fiber layout may have been introduced due to swelling of the fiber- reinforced films before rolling them into the graft, stretching of the films during graft rolling, and shrinkage of the structure during thermal annealing. Despite some irregularity, the mechanical properties presented below demonstrated a repeatable dependence on fiber layout.

[0278] Fiber architecture dictates mechanical behavior of composite vascular grafts. Mechanical responses, including burst pressure, compliance, and suture retention strength are summarized in Table 10. Representative burst data is illustrated in Figure 42, and the relationship between mechanical behavior of the vascular graft and fiber angle and spacing is presented in Figure 43. Thermal annealing enhances burst pressure and suture retention while reducing compliance, consistent with our prior observations that thermal annealing can increase the strength and Young's modulus of this protein polymer [136]. At a fixed fiber orientation (30°), decreasing average fiber spacing lead to increased fiber density with enhanced burst pressure and suture retention, but lower overall compliance. Likewise, with fiber spacing fixed and fiber angle increased, burst pressure and suture retention increases, but compliance decreases. Given the trade-offs of thermal annealing, fiber orientation, and fiber spacing, Design 6, with a fiber orientation of 22.5 and volume fraction of 7.3%, is selected as the best match for target mechanical properties.

[0279] Compliance matching in vascular bypass technology may reduce intimal hyperplasia and lead to increased patency [137-139]. Notably, arterial compliance varies broadly with age, sex, diet, smoking, and position in the vascular tree, ranging between 3 and 25 %/100 mm-Hg [140, 141]. This range highlights the advantage of platforms with the capacity to tailor compliance. Design 6 approaches the compliance of several native arteries (Table 10).

[0280] Burst pressures of native vein and artery (2000-3000+ mm-Hg) are often cited as benchmarks for bypass grafts although, even in a hypertensive emergency, blood pressure rarely exceeds 240 mm-Hg. Instead of representing anticipated conditions in vivo, the high strength of native vessel probably reflects a proxy measurement for the capacity of native vessels to resist fatigue in the face of hypertension, arteriosclerosis, or aneurysm. In the case of bypass grafts, high bursting strength suggests a greater resistance to damage from suture-line stress, biaxial stress, and fatigue. Grafts designed to biodegrade and remodel require even greater bursting strength to compensate for the anticipated structural alterations. Here we select 1000 mm-Hg as a target well above physiologic conditions, in consideration of data demonstrating the biostability of a physically crosslinked elastin-mimetic thblock even in the absence of chemical crosslinks [3].

[0281] This example provides a series of protein-based, small-diameter vascular grafts and experimental characterization of mechanical performance. The angle and density of fiber modulates the suture retention strength, bursting strength, and compliance of the grafts. Iterative adjustment of the fiber layout demonstrates the capacity to meet our mechanical targets, providing a platform from which the additional challenges of small- diameter vascular bypass can be addressed.

[0282] Example 5: Implanted medical device formed of a composite material.

[0283] Matching the stiffness of human linea alba. A composite material with +/- 25° fiber orientation and 0.15 mm average fiber spacing is developed to closely match the mechanical behavior of human abdominal fascia. The average uniaxial mechanical response is compared to literature data for human linea alba in Figure 44.

[0284] Ventral hernia repair model. Full-thickness abdominal wall defects in 8 wk old male Wistar rats are repaired with the APPLES surgical patch design described above (n=5) or a commercially available biologic implant (Permacol™, crosslinked porcine dermal collagen, n=3). Anesthesia is induced and maintained with isoflurane (2.5% and 1.5%, respectively) inhalation. A 3-cm vertical midline incision centered between the xiphoid and the pubis is created and skin is decollated from the muscular and fascial layers. A 1.5 x 2.0 cm rectangular full-thickness ventral abdominal wall defect consisting of muscles, fascia, and peritoneum is created and repaired with the implant materials using an onlay technique. Animals are monitored continuously for 1 to 2 hours postoperatively and then daily for the course of the study for signs of hernia recurrence or infection.

[0285] Abdominal wall repair with protein fiber composites. All animals survived the implant period without evidence of hernia recurrence, overlying skin ulceration, or bowel obstruction. Grossly, both patch materials appear to have supported native tissue ingrowth and firmly integrated with the abdominal wall (Figure 45). A region in the center of one of the five protein polymer composite patches appeared thin and translucent. Although no outward bulging is observed, perimeter measurements revealed area increases of 78 ± 53% and 72 ± 53% (n=5 and 3, p=0.89) protein polymer composite and porcine dermal collagen patches, respectively. Omental adhesions requiring sharp separation (grade III) are noted for all implants, although no other tissues or organs were involved. Strength of integration tensions were 0.62 ± 0.23 N/mm and 0.88 ± 0.17 N/mm (n=8 and 4, p=0.07) for elastin and tissue-derived patches. During testing, samples tore in either the implant region or the abdominal muscle rather than at the interface. Histological analysis of the porcine dermal collagen product reveas that the implanted collagen is fragmented after 8 weeks, with cells and host repair tissue in between and surrounding the fragments. In many fields, fragments of the crosslinked dermal collagen could not be discerned, and the material appears to be completely replaced with repair tissue. Histology of the protein polymer composite implants reveals that in many fields the elastin-like protein component of the implant is no longer visible, although in isolated regions the material persisted with an absence of cellular infiltration. The synthetic collagen fiber component is still largely present. In regions where the elastin-like protein is absent, the distance between collagen fibers increases, with cells and repair tissue between the fibers (Figure 46).

[0286] Protein polymer fiber composites in abdominal wall repair. Recombinant proteins derived from elastin sequences have been investigated as membranes to prevent adhesion and fibrosis, and for topical delivery of therapeutics [1 -3]. With the addition of reinforcing collagen fibers, the mechanical properties of protein polymer composites became potentially suitable for abdominal wall repair. The composites could be readily sutured and provided adequate mechanical support for the duration of the 8 wk implant, even though they were considerably thinner than the porcine dermal collagen sheets (0.4 and 1.0 mm, respectively). Porcine dermal collagen and the elastin-like portion of the composite are both substantially degraded over the 8 wk period, and similar results for the strength of integration and increased size of the repaired region are noted for both implant types. The observed porcine dermal collagen degradation agrees qualitatively with the manufacturer's claim that the material "balances the rate of degradation with the rate of tissue ingrowth [4]." Others report a variety of tissue responses to crosslinked porcine dermal collagen, ranging from cases of implants "melting" within a few weeks [5], to partial degradation [6], to no degradation [7, 8]. Highly variable tissue responses are observed within a single study of abdominal wall repair in primates, suggesting that the stability of crosslinked dermal collagen may be exquisitely sensitive to the local wound environment [9].

[0287] A range of degradation behaviors is also reported for elastin-like polymers. Protein polymers based on (GVGVP) and (GEGVP) sequences have persisted in vivo for between 2 weeks to about 6 months, depending upon chemical modification to the amino acid side chains [2, 10]. Sheets of poly(GVGVP) crosslinked with cobalt 60 radiation and implanted in the abdominal cavity of rats were reported not to absorb within 6 months [1]. Subcutaneous and intraperitoneal implants of LysB10, the protein polymer in this study, demonstrates a high degree of biostability with limited local inflammatory activity after three weeks in a mouse model [11]. The greater extent of degradation observed in this experiment may be due to the inflammatory environment resulting from excision of the abdominal wall tissues.

[0288] References for Example 5:

1. Hoban, L. D., et al., Use of polypentapeptides of elastin to prevent postoperative adhesions: efficacy in a contaminated peritoneal model. J Surg Res, 1994. 56(2): p. 179-83.

2. Alkalay, R. N., et al., Prevention of postlaminectomy epidural fibrosis using bioelastic materials. Spine (Phila Pa 1976), 2003. 28(15): p. 1659-65.

3. Wang, N.Z., et al., Skin concentrations of thromboxane synthetase inhibitor after topical application with bioelastic membrane. J Vet Pharmacol Ther, 2004. 27(1 ): p. 37- 43.

4. Permacol(TM) biologic implant: sound decisions in soft tissue repair, in Covidien. 2008.

5. Alwitry, A., S.J. Burns, and L. C. Abercrombie, Orbital implant exposure treatment with porcine dermal collagen patching. Orbit, 2006. 25(3): p. 253-6.

6. Macleod, T.M., et al., Histological evaluation of Permacol as a subcutaneous implant over a 20-week period in the rat model. Br J Plast Surg, 2005. 58(4): p. 518-32. 7. Kaleya, R. N., Evaluation of implant/host tissue interactions following intraperitoneal implantation of porcine dermal collagen prosthesis in the rat. Hernia, 2005. 9(3): p. 269-76.

8. Ayubi, F. S., et al., Abdominal wall hernia repair: a comparison of Permacol and Surgisis grafts in a rat hernia model. Hernia, 2008. 12(4): p. 373-8.

9. Sandor, M., et al., Host response to implanted porcine-derived biologic materials in a primate model of abdominal wall repair. Tissue Eng Part A, 2008. 14(12): p. 2021 - 31.

10. Urry, D.W., Elastic molecular machines in metabolism and soft-tissue restoration. Trends Biotechnol, 1999. 17(6): p. 249-57.

11. Sallach, R. E., et al., Elastin-mimetic protein polymers capable of physical and chemical crosslinking. Biomatehals, 2009. 30(3): p. 409-22.

[0289] SUMMARY: We present a process for the scalable production of synthetic collagen fiber, by the spinning of collagen fiber and subsequent assembly of fibrillar ultrastructure. Fiber with an elliptical cross-section of 53 ± 14 by 21 ± 3 μm and an ultimate tensile strength of 90 ± 19 MPa is produced at 60 meters per hour from a sterile-filtered monomeric collagen solution. The spinning concentration, flowrate, and needle size can be adjusted to control the size of the spun fiber. Micro-differential scanning calorimetry demonstrates that the triple helical macromolecular structure is preserved after spinning. Second harmonic generation analysis suggests fibrillar structure and transmission electron microscopy confirms the presence of banded, self- assembled fibrils of 53 ± 14 nm diameter, largely aligned with the fiber axis. Six week subcutaneous murine implants of glutaraldehyde crosslinked bundles demonstrates little degradation, but infiltration of macrophages. Uncrosslinked bundles present after six weeks but displayed more degradation, with macrophages localized largely around the bundle perimeter. Provides is a process for the scalable production of collagen fiber with a self-assembled fibrillar structure and sufficient strength for use in flexible composite tissue substitutes.

[0290] Collagen fiber spinning can be scaled-up by optimizing any number of process parameters. For example, referring to Figures 1 and 2, extrusion rate may be increased. Both the collagen and buffer can be extruded faster to increase production rate. In addition, different pump types may be used. In particular, the syringe-type pumps used in the examples have a confined volume, so the process is stopped when the buffer syringe (i) runs out. Another type of pump, including specifically non-syringe pumps as known in the art can avoid this delay. Furthermore, other types of non- syringe pumps can also permit the buffer to be re-circulated to limit waste. The pipe segment in (viii) can be longer and/or of a larger diameter to collect more fiber. Parallel processes may be used to extrude more fibers, such as by extruding through multiple spinnerets / needles (iii). This could be useful to produce a multi-filament yarn or if the geometry of the system is altered to avoid tangles multiple monofilaments may be spun in parallel. Additional fiber drawing (stretching) step. Adding a step for controlled stretching of the fiber can improve the fiber structure and strength. There is some stretching (estimated 15%) in the stage depicted in Figure 2, and additional stretching may be provided. Further improvement may be obtained by altering the collagen extrusion solution. Currently, the solution in (ii) comprises collagen monomers. These are extruded into a fiber, but do not self-assemble into native collagen fibrils until the off- line incubation step. An alternative strategy to simplify fiber production is to initiate the self-assembly of collagen monomers into fibrils before spinning. Then a dispersion of fibrils can be spun into a fiber without the off-line incubation step, thereby substantially simplifying fiber production.

[0291] Example 2 provides a strategy for fabricating anisotropic protein polymer lamellar elastic structures. Sheets comprising eight fiber-reinforced lamellae and exterior capping lamellae without fiber are constructed with controlled fiber orientation and volume fraction. Scanning electron microscopy, transmission electron microscopy, and digital volumetric imaging confirms the structure of the flexible biocomposites. The effect of fiber orientation and volume fraction on Young's Modulus, yield stress, ultimate tensile stress, strain-to-failure, and resilience is evaluated in uniaxial tension. The addition of collagen fiber and the alignment of fiber with the direction of applied force tend to increase Young's Modulus, resilience, and yield stress. This analysis demonstrates a semi-automated fabrication strategy for flexible biomaterial composites with defined resilience, modulus, and yield stress.

[0292] The microcrimping method provided in Example 3 relates to elastin-like protein polymer lamellae reinforced with undulating collagen fibers. Three exemplified profiles for the microridge crimping template are investigated: triangular, rectangular, and chamfered rectangular. The chamfered rectangle design is preferred due to the stability of the microridge against collapse, the relatively large overhead space for crimping, and the minimal damage caused by the fiber contacting region. The wavelength of fiber crimp in a hydrated lamella is 143 ± 5 μm. Alteration of the pre-extension parameter demonstrates the capacity to adjust the degree of crimping from 3.1 % to 9.4%, corresponding to mechanical modulus transitions at 4.6% and 13.3% strain. Cyclic mechanical loading of up to 1000 cycles do not substantially alter the crimp morphology of embedded fibers. This example represents the first process for microchmping of synthetic collagen fibers, and demonstrates the capacity of elastin-like protein embedded fiber arrays to display a defined mechanical transition point response.

[0293] Example 4 provides one application of the materials and fabrication strategies described herein. The example specifically relates to the generation of small diameter vascular grafts. Six vascular graft designs are fabricated with an inner diameter of 4 mm, wall thicknesses of 0.9 mm, fiber volume fractions ranging from 3 to 7%, and fiber orientations of 15, 22.5, and 30° relative to the axial direction. The structure of the graft wall is examined with scanning electron microscopy. A system to perform pressure- diameter analysis at defined levels of axial force is developed and implemented to study the effects of fiber volume fraction and orientation on graft mechanics. Different design variations displayed burst pressures as high as 2760 ± 360 mm Hg, compliance as high as 8.4 ± 1.4 % / 100 mm Hg, and suture retention strengths up to 192 ± 20 g-f. One optimal design, with a fiber orientation of 22.5° and fiber volume fraction of 7.3%, simultaneously satisfied target mechanical properties with suture retention strength of 173 ± 4 g-f, bursting strength of 1483 ± 143 mm Hg, and compliance of 5.1 ± 0.8 %/100 mm Hg. Example 5 is another application, where a composite material corresponding to a medical device that is a surgical patch is implanted in an animal.

[0294] The application of advanced fiber reinforcing strategies can further enhance graft mechanics. For example, the use of microcrimped fiber, non-continuous fiber, or 2D collagen fiber layouts generated by alternative means may lead to further optimization of compliance and strength. The use of a more a compliant elastin-like matrix in combination with higher fiber volume fractions can be explored to identify a material with an improved compliance and strength.

[0295] In summary, design and fabrication of flexible biocomposites for soft tissue substitution is provided. A new technique for the scalable production of synthetic collagen fiber is provided. Recombinant, elastin-like protein polymer lamellae are reinforced with defined densities and orientations of collagen fiber and the capacity to tailor the Young's modulus, resilience, compliance, bursting strength and other mechanical properties of multilamellar sheet and tube structures was demonstrated. A novel approach to generate microcrimped collagen fiber arrays is provided and the mechanics of these structures were analyzed in uniaxial tension. Collectively, this work represents significant progress toward the design and fabrication of synthetic extracellular matrices with defined mechanical properties

[0296] A summary of sequence listings is provided in TABLE 11. In an embodiment, the invention is directed to any one or more of these sequences as the first material corresponding to the elastin-like material that supports the correspondingly stiffer fibers of the second material.

STATEMENTS REGARDING INCORPORATION BY REFERENCE

AND VARIATIONS

[0297] All references throughout this application, for example patent documents including issued or granted patents or equivalents; patent application publications; and non-patent literature documents or other source material; are hereby incorporated by reference herein in their entireties, as though individually incorporated by reference, to the extent each reference is at least partially not inconsistent with the disclosure in this application (for example, a reference that is partially inconsistent is incorporated by reference except for the partially inconsistent portion of the reference).

[0298] The following U.S. patents and published patent applications (followed by attorney docket number) are specifically incorporated by reference to the extent not inconsistent with the present disclosure: 7,244,830 (1 -02); 2004-0110439 (29-01 ); 2004- 0063200 (78-01 ) and 2004-0171545 (133-02); WO 2008/033847 (133-02A).

[0299] The terms and expressions which have been employed herein are used as terms of description and not of limitation, and there is no intention in the use of such terms and expressions of excluding any equivalents of the features shown and described or portions thereof, but it is recognized that various modifications are possible within the scope of the invention claimed. Thus, it should be understood that although the present invention has been specifically disclosed by preferred embodiments, exemplary embodiments and optional features, modification and variation of the concepts herein disclosed may be resorted to by those skilled in the art, and that such modifications and variations are considered to be within the scope of this invention as defined by the appended claims. The specific embodiments provided herein are examples of useful embodiments of the present invention and it will be apparent to one skilled in the art that the present invention may be carried out using a large number of variations of the devices, device components, methods steps set forth in the present description. As will be obvious to one of skill in the art, methods and devices useful for the present methods can include a large number of optional composition and processing elements and steps.

[0300] When a group of substituents is disclosed herein, it is understood that all individual members of that group and all subgroups, including any isomers, enantiomers, and diastereomers of the group members, are disclosed separately. When a Markush group or other grouping is used herein, all individual members of the group and all combinations and subcombinations possible of the group are intended to be individually included in the disclosure.

[0301] Every formulation or combination of components described or exemplified herein can be used to practice the invention, unless otherwise stated. Although nucleotide sequences are specifically exemplified as DNA sequences, those sequences as known in the art are also optionally RNA sequences (e.g., with the T base replaced by U, for example).

[0302] Whenever a range is given in the specification, for example, a physical parameter range (modulus, dimension), strain, stress, a temperature range, a time range, or a composition or concentration range, all intermediate ranges and subranges, as well as all individual values included in the ranges given (e.g., within a range and at the ends of a range) are intended to be included in the disclosure. It will be understood that any subranges or individual values in a range or subrange that are included in the description herein can be excluded from the claims herein.

[0303] All patents and publications mentioned in the specification are indicative of the levels of skill of those skilled in the art to which the invention pertains. References cited herein are incorporated by reference herein in their entirety to indicate the state of the art as of their publication or filing date and it is intended that this information can be employed herein, if needed, to exclude specific embodiments that are in the prior art. For example, when composition of matter are claimed, it should be understood that compounds known and available in the art prior to Applicant's invention, including compounds for which an enabling disclosure is provided in the references cited herein, are not intended to be included in the composition of matter claims herein. [0304] As used herein, "comprising" is synonymous with "including," "containing," or "characterized by," and is inclusive or open-ended and does not exclude additional, unrecited elements or method steps. As used herein, "consisting of excludes any element, step, or ingredient not specified in the claim element. As used herein, "consisting essentially of does not exclude materials or steps that do not materially affect the basic and novel characteristics of the claim. In each instance herein any of the terms "comprising", "consisting essentially of and "consisting of may be replaced with either of the other two terms. The invention illustratively described herein suitably may be practiced in the absence of any element or elements, limitation or limitations which is not specifically disclosed herein.

[0305] One of ordinary skill in the art will appreciate that starting materials, biological materials, reagents, synthetic methods, purification methods, analytical methods, assay methods, and biological methods other than those specifically exemplified can be employed in the practice of the invention without resort to undue experimentation. All art-known functional equivalents, of any such materials and methods are intended to be included in this invention. The terms and expressions which have been employed are used as terms of description and not of limitation, and there is no intention that in the use of such terms and expressions of excluding any equivalents of the features shown and described or portions thereof, but it is recognized that various modifications are possible within the scope of the invention claimed. Thus, it should be understood that although the present invention has been specifically disclosed by preferred embodiments and optional features, modification and variation of the concepts herein disclosed may be resorted to by those skilled in the art, and that such modifications and variations are considered to be within the scope of this invention as defined by the appended claims.

TABLE 1 : Wet Spinning Parameters

1 0.4 0.1 5

2 0.4 0.06 5

3 0.4 0.03 5

4 0.4 0.015 5

5 0.4 0.1 2

6 0.4 0.06 2

7 0.4 0.06 7.5

8 0.1 0.06 7.5

9 0.1 0.03 7.5

AC 0.4 0.08 5

MC sample numbers 4 - 6 were not analyzed further because the low collagen extrusion rate or concentration resulted in frequent breaks during fiber spinning.

TABLE 2: Apparent Temperature and Enthalpy of Collagen Denaturation

Purified collagen (MRTC), continuous fiber without FIB treatment (AC-FIB), continuous fiber with FIB treatment (AC+FIB), continuous fiber with FIB treatment and glutaraldehyde crosslinking (AC+FIB+GLUT), and rat-tail tendon (RTT). Both the process of spinning collagen into fiber and the FIB treatment increased the thermal stability of the material, as shown by the elevated T_m and ΔH values. Values are averages and standard deviations from three scans. Table 3. Size and Mechanical Properties of Collagen Fiber

MC

Sample Needle Extrusion Major Minor

Number ID Rate Concentration Diameter Diameter UTS Modulus Strain-to-

/ AC t^T (mm) (niL/min) (mg/mL) (μm) (μm) (MPa) (MPa) failure (%)

1 0.4 0.1 5 59 ± 10 22 ± 3 54 ±12 556 ±120 11.2 ± 2.9

2 0.4 0.06 5 37 ± 9 17 ± 2 60 ±16 529 ±157 12.4 ± 3.1

3 0.4 0.03 5 25 ± 4 12 ± 2 77 ±16 673 ± 100 12.6 ± 1.8

7 0.4 0.06 7.5 55 ± 6 27 ± 2 55 ±14 575 ± 66 11.2 ± 2.5

8 0.1 0.06 7.5 44 ± 6 25 ± 2 67 ±19 608 ± 54 12.2 ± 3.2

9 0.1 0.03 7.5 29 ± 3 20 ± 2 90 ±37 837 ±133 11.7 ± 3.5

AC 0.4 0.08 5 53 ± 14 21 ± 3 94 ±19 775± 173 14.3 ± 1.9

'Samples 4 through 6 did not consistently produce fiber. Table 4: Summary of Collagen Fiber Diameter and Strength In Prior Reports and in the Current

Example

The strength of hydrated, glutaraldehyde crosslinked fibers is provided when available. XL refers to crosslinking method and Glut, DHT, and EDC refer to glutaraldehyde, dehydrothermal, and 1-Ethyl-3-[3- dimethylaminopropyl]carbodiimide hydrochloride crosslinking. Fiber schemes are indicated as discontinuous fiber (DF) and continuous fiber (CF) processes.

Table 5. Assessment of Fiber Layout

Fiber Orientation Fiber Spacing (mm)

Design Number (°) Volume fraction (%)

Nominal Measured Nominal Measured

1 0 0.8 ±3.7 0.15 0.19 ±0.01 16 .3 ±0.6

2 90 90.8 ±1.4 0.15 0.19 ±0.01 16 .7 ±0.5

3 15 13.4 ±0.9 0.15 0.18 ±0.01 17 .7 ±0.8

4 15 13.2 ±0.8 0.45 0.47 ± 0.02 6. 8 ±0.3

5 15 13.0 ±0.9 1.30 1.03 ±0.03 3. l±O.l

Table 6. Mechanical Properties of APPLES Designs

Volume Fiber Young's Fraction Orientation Modulus Yield Stress Yield Resilience Strain to

Design (%) O (MPa) (MPa) Strain (%) UTS (MPa) (%) failure (%) n

1 16.3 ±0.6 0 33.1 ±3.9 2.87 ±0.56 12 ± 1 3.62 ±1.07 76.1 ±2.9 31±10 5

²t 16.7 ±0.5 90 5.3 ±0.6 1.99 ±0.28 50.8 ±0.6 184 ±38 5

15

3 17.7 ±0.8 26.0 ±4.1 2.58 ±0.37 14 ± 1 2.66 ±0.72 75.8 ±2.0 23 ±3 6

15

4 6.8 ±0.3 20.9 ±2.0 1.87 ±0.34 13 ±2 2.41 ±0.17 71.9±2.1 47 ±7 6

15

5 3. l±O.l 13.9±1.36 1.31 ±0.19 12 ± 1 1.85 ±0.66 66.8 ±1.4 182 ±135 7

6t 6.1 ±0.7 4.08 ± 0.80 53.1 ±1.4 314±26 5 tDesigns without fiber (6) or with fibers oriented perpendicular to loading (2) did not display abrupt yielding points. Table 7. Mechanical Properties of Microcrimped Lamellaef

Pre- Young s Young's extension Modulus 1 Modulus 2 Transition UTS Strain-to-

(%) (MPa) (MPa) (%) (MPa) failure (%) n

0 - 90.8 ± 14.3 1.1 ± 0.2 8.56 ± 1.64 11.5 ± 1.9 4

15 9.1 ± 1.8 62.2 ± 10.4 4.6 ± 0.9 4.93 ± 0.92 13.1 ± 1.6 3

30 2.2 ± 0.5 25.3 ± 4.1 13.3 ± 0.7 2.08 ± 0.73 21.9 ± 3.3 4 tYoung's Modulus 1 and 2 are the slopes of linear fits to the stress-strain response before the transition strain and for the final 4% strain before yielding, respectively.

Table 8. Synthetic Collagen Fiber Architecture in Graft Designs

t No fiber

Table 9. Synthetic Collagen Fiber Layout and Graft Dimensions!

Fiber Spacing (mm) Fiber Angle O Fiber Inner Diameter

L)Q SIgIl Wall (mm)

Nominal Measured Nominal Measured Volume (%) (mm)

1 - - - - 0 3.36 1.08

2 0.23 0.22± 0.02 30 30.8 ± 0.3 5.0 ± 0.5 4.57 ± 0.45 0.96 ± 0.10

3 0.30 0.33 ± 0.02 30 30.2 ± 0.7 3.3 ±0.2 5.02 ± 0.19 0.93 ± 0.10

4 0.15 0.17 ± 0.02 30 30.0 ± 1.0 6.4 ± 0.8 4.77 ± 0.39 0.87 ± 0.07

5 0.15 0.15 ± 0.01 15 14.9 ± 1.0 6.5 ± 0.4 4.82 ± 0.26 0.89 ± 0.10

6 0.15 0.14 ± 0.02 22.5 23.4 ± 1.6 7.3 ± 1.0 4.56 ± 0.15 0.84 ± 0.08 tValues represent the mean and standard deviation from three prototypes of each graft design. Design 1 did not contain collagen fiber. Graft diameter is the pressurized inner diameter at 120 mm Hg. The fiber volume fraction was calculated from the mean graft dimensions, fiber dimensions, and fiber spacing. Error in the fiber volume is propagated from the standard deviation of the fiber spacing measurement and other sources of error were ignored. Inner diameter values were calculated by subtracting twice the wall thickness from the outer diameter of the graft.

Table 10. Mechanical Responses of Composite Vascular Grafts and Arteries i_i E_tn_gn_erere;i .

>- 2 1409 ± 141 3.6 ±1.8 121 ±31

O 2a 649 ± 74 8.4 ±1.4 70 ±18 in 3 755 ± 227 8.4 ±3.4 95 ±23

4 2760 ± 360 2.8 ±0.5 192 ±20

5 893 ± 126 7.1 ±1.2 124 ± 24

6 1483 ± 143 5.1 ±0.8 173 ±4

Target >1000 5-9 180

<u Dahl, et al [88] 803 ± 105 3.5 ±0.2 -

L'Heureux, et al [86] 3468 ± 500 1.5 ±0.3 162 ±15

Porcine common carotid

3320 ±413 18.7±4.1 artery [88]

Human common carotid - 18-26 - artery [141]

Human saphenous vein

1680-2273 0.7-1.5 196 ±2 [86]

Human saphenous vein

5.0 ±6.7

[142]

Human artery (range) [86] 2031-4225 4.5-6.2 200 ± 119

Human common femoral

- 8.3 ±1.8 - artery [143]

<q <u Human proximal superficial femoral artery - 7.2 ±1.5 -

% [143]

Human distal superficial f O,.J ^ +E 0 yj .8 o femoral artery [143]

Human midgenicular

6.1 ±1.1 popliteal artery [143]

Human radial artery [140] - 3-4 -

Human external iliac artery

8.0 ±5.9

[142]

Dog femoral artery [144] - 6.8 - f Design 2a refers to design 2, fabricated without the thermal annealing step. TABLE 11: Summar of Se uences

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Claims

CLAIMS We claim:

1 . A composite biomimetic material comprising: a first material comprising an elastin-mimetic protein formed into an elastomehc film; and a second fibrous material comprising a plurality of collagen fibers embedded in the first material; wherein: the elastin-mimetic protein is selected from the group consisting of: LysB10, B10, R1 , R2 and R4; and the collagen fibers are continuous spun fibers that extend a length of the first material and are aligned in the first material in at least one preferential direction, and an individual fiber has a Young's modulus that is higher than the Young's modulus of the film.

2. The composite biomimetic material of claim 1 , wherein the fiber is configured with an undulating geometry when said fiber is unstrained.

3. The composite biomimetic material of claim 1 , wherein the plurality of collagen fibers are arranged in a spatially-varying pattern in said film.

4. The composite biomimetic material of claim 1 , wherein an individual collagen fiber has a cross-sectional area that is selected from a range that is greater than or equal to 75 μm² and less than or equal to 8000 μm²; and a Young's modulus that is selected from a range that is greater than or equal to 400 MPa.

5. The composite biomimetic material of claim 1 , wherein the fiber is at least partially crimped when said composite biomimetic material is not strained.

6. The composite biomimetic material of claim 5, wherein the at least partially crimped portion has: a wavelength selected from a range that is greater than or equal to 50 μm and less than or equal to 1 mm, and an amplitude selected from a range that is greater than or equal to 20 μm and less than or equal to 1 mm.

7. The composite biomimetic material of claim 5, wherein the crimped portion has a geometric shape that is selected from the group consisting of spiral, helical, sinusoidal wave, sawtooth and ridged.

8. The composite biomimetic material of claim 5, wherein the at least partially crimped portion has a crimp magnitude that is greater than or equal to 2%.

9. The composite biomimetic material of claim 8 having: a transition point strain selected from a range that is greater than or equal to 1 % and less than or equal to 20%, a compliant Young's modulus selected from a range that is less than or equal to 15 MPa; and a rigid Young's modulus selected from a range that is greater than or equal to 20 MPa.

10. The composite biomimetic material of claim 1 having one or more physical parameters that are anisotropic.

1 1 . The composite biomimetic material of claim 10, wherein the second fibrous material is aligned in two preferential directions defined by a fiber angle.

12. The composite biomimetic material of claim 1 , wherein the fibrous material has an average spacing distance between adjacent fibers, said average spacing distance selected from a range that is greater than or equal to 0.05 mm and less than or equal to 1 mm.

13. The composite biomimetic material of claim 1 having an ultimate tensile strength that is greater than or equal to 2 MPa and a strain to failure that is greater than or equal to 12%.

14. The composite biomimetic material of claim 1 formed into a medical device.

15. The composite biomimetic material of claim 14, wherein the medical device is a soft tissue patch, a dermal filler, a hernia patch, a valve leaflet or a vascular graft.

16. The composite biomimetic material of claim 14, further comprising one or more of a drug, a growth factor, a polysaccharide, a living cell, or a combination thereof supported by or connected to the first material, the second material, or both.

17. The composite biomimetic material of claim 1 , wherein the material is formed into a sheet or a tubular cylinder having a length that is greater than or equal to 1 cm, and wherein the fibers are continuous and the continuous fibers individually span at least 90% of the length of the sheet.

18. A multilayer material comprising a plurality of layers, wherein each layer comprises the composite biomimetic material of claim 1 .

19. The multilayer material of claim 18, wherein the number of layers is selected from a range that is greater than or equal to 2 and less than or equal to 100.

20. The multilayer material of claim 19, wherein the multilayer is laminated by a bottom surface layer and a top surface layer, wherein each of said bottom surface layer and top surface layer is the elastin-mimetic protein formed into a film without the second material.

21. The composite biomimetic material of claim 1 , wherein the film has a thickness selected from a range that is greater than or equal to 30 μm and less than or equal to 1 mm.

22. The composite biomimetic material of claim 1 , wherein the second material has a volume fraction that is greater than or equal to 1 % and less than or equal to 30%, and wherein the fibers are uniformly distributed in said first material.

23. A composite biomimetic material comprising: a first material comprising an elastin-mimetic protein formed into a film; and a second fibrous material embedded in said film, said second fibrous material comprising a plurality of continuous spun collagen fibers, wherein at least a portion of the continuous spun collagen fibers are crimped; wherein the collagen fibers in an unchmped state have a Young's modulus that is higher than the Young's modulus of the film and said collagen fibers are aligned in at least one preferential direction.

24. The composite biomimetic material of any of claims 1 - 23, wherein the elastin- mimetic protein is LysBI O.

25. A method of making the material of claim 1 comprising the steps of: providing a fibrous material on a first support surface; and introducing a solution of the elastin-mimetic material over the fibrous material.

26. The method of claim 25 further comprising gelling, crosslinking, polymerizing or drying the solution of the elastin-mimetic material to form the film having fibrous material embedded therein.

27. The method of claim 25, further comprising pressing the introduced solution of the elastin-mimetic material by a second support surface that faces the first support surface, wherein the first and second support surfaces are separated by the elastin-mimetic material and the fibrous material.

28. The method of claim 25, wherein the first support surface corresponds to a surface of a shaft, and the fibrous material is at least partially wound around the shaft.

29. The method of claim 25, further comprising selectively adjusting one or more physical parameters of the composite material by varying water absorbency of said first support surface.

30. The method of claim 29, wherein the water absorbency is varied by adjusting the porosity of the first support surface.

31. The method of claim 25, further comprising generating crimps in the fibrous material, wherein the crimps are generated by: providing a first stretchable sheet at a first level of strain; stretching the first stretchable sheet to a second level of strain that is greater than the first level of strain; attaching the fibrous material to the stretchable sheet at the second level of strain; and relaxing the stretchable sheet to which the fibrous material is attached to a third level of strain that is less than the second level of strain, thereby generating crimps in the fibrous material.

32. The method of claim 31 further comprising fixing the crimped fiber.

33. The method of claim 31 , wherein the fibrous material is attached to a second stretchable sheet that faces the first stretchable sheet, wherein the fibrous material is positioned between the first and second stretchable sheets.

34. The method of claim 33, wherein the first stretchable sheet, the second stretchable sheet, or both, have a contact surface that has a pattern of relief features that provides a plurality of contact points with the fibrous material, wherein the fibrous material remains fixed in position relative to the contact points during a change in strain, and portions of the fibrous material between the contact points are crimped.

35. The method of claim 34, wherein the pattern of relief features generate a surface shape on the stretchable sheet that is selected from the group consisting of: sinusoidal wave; rounded ridges; sawtooth; and chamfered rectangular.

36. The method of claim 34 wherein the relief features have at least one dimension that is less than or equal to 300 μm.

37. The method of claim 34, wherein the contact surface further comprises grooves having a receiving volume to receive deformed fibers into the grooves.

38. The method of claim 25 wherein the collagen fiber has a Young's modulus that is at least ten times greater than the Young's modulus of the film.

39. A method of making a collagen fiber, the method comprising: providing a collagen material in solution; extruding the collagen-containing solution into a wet spinning buffer at an extrusion flow-rate to form a gel fiber; passing the gel fiber through a rinse bath; passing the rinsed gel fiber through a dryer to provide a dry collagen fiber; continuously collecting the dried fiber; incubating the collected fiber in an incubation bath; rinsing the incubated collected fiber; and drying the rinsed incubated and collected fiber.

40. The method of claim 39, wherein the collagen-containing solution comprises monomehc collagen having a concentration that is greater than 4 mg/mL and the extrusion flow-rate is greater than or equal to 0.5 mL/min.

41 . The method of claim 39, wherein the rinsed incubated and collected fiber drying step is performed with the fiber under tension.

42. The method of claim 39, wherein the collagen material in solution and/or the incubation bath further comprises one or more of a growth factor, a drug, a protein or a polysaccharide.

43. The method of claim 39, wherein the continuous fiber has a length that is greater than or equal to 1 m.

44. A method of generating crimps in a biomimetic fibrous material, said method comprising the steps of : providing a first stretchable sheet at a first level of strain; stretching the stretchable sheet to a second level of strain that is greater than the first level of strain; attaching a biomimetic fibrous material to the stretchable sheet at the second level of strain; and relaxing the stretchable sheet to which the fibrous material is attached to a third level of strain that is less than the second level of strain, thereby generating crimps in the biomimetic fibrous material.

45. The method of claim 44, wherein the fibrous material comprises collagen fibers.

46. The method of claim 44, wherein the fibrous material comprises a plurality of aligned fibers having a length that is greater than or equal to 10 cm.

47. The method of claim 44, wherein the biomimetic fibrous material is attached to a second stretchable sheet that faces the first stretchable sheet, wherein the fibrous material is positioned between the first and second stretchable sheets and the second stretchable sheet is correspondingly strained to the first level, the second level and the third level of the first stretchable sheet.

48. The method of claim 47, wherein the first stretchable sheet, the second stretchable sheet, or both, have a contact surface that has a pattern of relief features that provides a plurality of contact points with the fibrous material, wherein the fibrous material remains fixed in position relative to the contact point during changes in strain.

49. The method of claim 47, wherein the plurality of contact points provide crimping having a wavelength selected from a range that is greater than or equal to 50 μm and less than or equal to 1 mm, and an amplitude that is selected from a range that is greater than or equal to 50 μm and less than or equal to 1 mm.

50. The method of claim 44, further comprising embedding the crimped biomimetic fibrous material into an elastic material having a Young's modulus that is at least ten times less than the biomimetic fibrous material Young's modulus in an unchmped configuration.

51. The method of claim 50, wherein the elastic material is an elastin-mimetic material.