Double component mandrel for electrospun stentless, multi-leaflet valve fabrication

Abstract

To allow superior control of electrodeposition on indented portions of an electrodeposition target.SOLUTION: A method of making a valve structure includes a step of rotating a mandrel and an electrodeposition target attached to the mandrel about a rotational axis of the mandrel. The target includes an exterior surface having at least one conductive surface portion and at least one non-conductive surface portion. The method also includes a step of electrodepositing a polymer matrix of a biodegradable, biocompatible polymer composition onto the at least one conductive surface portion and the at least one non-conductive surface portion of the exterior surface of the rotating electrodeposition target to form the valve structure.SELECTED DRAWING: Figure 1A

Description

[Technical Field]

[0001] (CROSS-REFERENCE TO RELATED APPLICATIONS) This application claims priority to U.S. Provisional Patent Application No. 62 / 126,040, filed February 27, 2015, which is incorporated herein by reference in its entirety.

[0002] The present disclosure is generally directed to mandrel structures and polymer electrodeposition methods useful, for example, in the fabrication of prosthetic heart valves and other biostructures. [Background technology]

[0003] Congenital valve defects remain a burden for pediatric cardiac patients. 8 / 1000 infants are born with congenital heart defects each year, affecting approximately 1,000,000 Americans. Valve repair / replacement procedures can have limited durability and require permanent antithrombogenic therapy (e.g., Coumadin, Pradaxa, Xarelto, etc.). Most importantly, current materials for heart valve repair or replacement (e.g., bovine pericardium, pyrolytic titanium, etc.) are limited by their inability to grow or remodel.

[0004] Tissue-engineered valves offer the potential to overcome these drawbacks by creating biostructures that undergo bodily growth, have reduced thrombogenicity, and possess appropriate coaptation levels under physiological conditions. However, tissue-engineered valves remain limited by the availability of biocompatible scaffold materials with desirable degradation and biomechanical properties. While the majority of previous in vivo studies available in the tissue engineering literature have involved seeded / non-seeded nonwoven scaffolds (e.g., polyglycolic acid (PGA):poly-L-lactic acid (PLLA) blends), these prostheses suffer from several limitations, including tissue shrinkage over time in vivo, progressive calcification, and valve regurgitation. The amorphous structure and the absence or limited control over material fiber architecture are among the most important contributors to these limitations.

[0005] Electrospun valve leaflet fabrication, as presented in International Publication No. WO 2011 / 150328, offers the opportunity to fabricate valve leaflets with controlled anisotropy. However, conventional fiber deposition onto flat or cylindrical targets does not allow for the attainment of curvilinear fiber distributions that mimic the collagen microarchitecture of native valves, the creation of concave leaflets at rest that mimic native anatomy, or the fabrication of fully assembled multi-leaflet stentless valves. Similarly, fiber deposition onto complex geometries does not allow for control of leaflet bending stiffness (out-of-plane mechanics) or full control over the leaflet's mechanical anisotropy (in-plane mechanics).

[0006] WO 2010 / 041944 describes a method for manufacturing an electrospun valve. However, the described method and device do not allow for the fabrication of a unified valve-plus-conduit structure with good control over the physical parameters of the leaflets, such as anisotropy, fiber orientation, stiffness, and thickness, parameters required for the formation of a useful prosthetic valve.

[0007] Electrospinning of various materials is described, for example, in International Publication No. WO 2010 / 041944 and International Publication No. WO 2011 / 150328, as well as U.S. Patent Application Publication No. US 2008 / 0268019 and U.S. Patent Application Publication No. US 2008 / 0109070. In its most general sense, electrodeposition, such as electrospinning, is the deposition of polymer fibers from an electrically charged nozzle onto an oppositely charged target, where an electric field induces the formation and flow of fibers toward the target. The target may be a rotating object, called a mandrel, or a non-rotating surface. Movement of the nozzle and / or target using standard two-dimensional or three-dimensional stages, robots, motors, etc., including rotational motion, generates relative motion between the nozzle and target. Controlled electrodeposition of polymer compositions onto targets, such as those on a rotating mandrel, presents significant technical constraints. When the target, e.g., a mandrel, includes a concave portion (a depression extending toward the axis of rotation), the rotation and longitudinal movement of the mandrel relative to the nozzle delivering the polymer prevents adequate control over thickness, density, anisotropy, and fiber quality within the plane of the concave portion. [Prior art documents] [Patent documents]

[0008] [Patent Document 1] International Publication No. 2011 / 150328 [Patent Document 2] International Publication No. 2010 / 041944 [Patent Document 3] U.S. Patent Application Publication No. 2008 / 0268019 [Patent Document 4] U.S. Patent Application Publication No. 2008 / 0109070 [Patent Document 5] U.S. Patent Application Publication No. 2002 / 02390 [Patent Document 6] U.S. Patent No. 4,902,508 [Patent Document 7] U.S. Patent No. 4,956,178 [Patent Document 8] U.S. Patent No. 5,281,422 [Patent Document 9] U.S. Patent No. 5,352,463 [Patent Document 10] U.S. Patent No. 5,372,821 [Patent Document 11] U.S. Patent No. 5,554,389 [Patent Document 12] U.S. Patent No. 5,573,784 [Patent Document 13] U.S. Patent No. 5,645,860 [Patent Document 14] U.S. Patent No. 5,771,969 [Patent Document 15] U.S. Patent No. 5,753,267 [Patent Document 16] U.S. Patent No. 5,762,966 [Patent Document 17] U.S. Patent No. 5,866,414 [Patent Document 18] U.S. Patent No. 6,890,562 [Patent Document 19] U.S. Patent No. 6,890,563 [Patent Document 20] U.S. Patent No. 6,890,564 [Patent Document 21] U.S. Patent No. 6,893,666 [Patent Document 22] U.S. Patent Application Publication No. 2008 / 0260831 [Patent Document 23] U.S. Patent No. 5,216,115 [Patent Document 24] U.S. Patent Publication No. 2011 / 0082545 [Non-patent literature]

[0009] [Non-Patent Document 1] "Biaxial Mechanical Evaluation of Planar Biological Materials" by M. Sacks in Journal of elasticity and the physical science of solids, 07-2000, Vol. 61, Issues 1-3, pp 199-246.

Non-patent Document 2

Non-patent Document 3

Non-patent Document 4

Non-patent Document 5

[0010] The apparatus and methods provided herein overcome these limitations by incorporating an electrodeposition target like collection mandrel design that allows for superior control of electrodeposition into the recessed portions of the target. [Means for solving the problem]

[0011] In some aspects, the apparatus includes a mandrel having a cylindrical surface and a concave surface. By patterning conductive and non-conductive or non-conductive insulating materials on the target, polymer deposition can be similarly patterned. Accordingly, provided herein are electrodeposition targets having insulating and non-insulating surfaces, as well as methods for making electrodeposited objects using the targets.

[0012] In the devices described herein, a collecting mandrel is used to fabricate fully multi-leaflet, stentless valve prostheses of variable shape (mitral, aortic, pulmonary, tricuspid, or pathological anomalies) and variable dimensions. In some embodiments, the designs described herein provide the following capabilities: (a) control of the mechanical anisotropy (plane mechanics) of the valve leaflets by varying the tangential velocity of the mandrel; (b) control of the bending stiffness (out-of-plane mechanics) of the valve leaflets by varying the linear velocity of the mandrel; (c) control of the leaflet microfibril orientation (e.g., curved fibers, where the primary direction of alignment is circumferential in the abdominal region and changes axially toward the coaptation region); (d) the possibility of constructing concave leaflets that mimic native anatomy, where the shape, thickness, and size can be varied to replicate healthy or pathological valve anatomy in humans or animals; and (e) the possibility of constructing fully assembled multi-leaflet stentless valves without a valve conduit.

[0013] With the electrospinning mandrel described herein, material deposition is concentrated only in the desired area by utilizing a non-conductive (e.g., insulating) material. In the following examples, the material used is acrylonitrile butadiene styrene (ABS), and the conductive deposit is made of aluminum (material utilized: Aluminum 6061-T651). Previous embodiments, such as those described in PCT Publication No. WO 2010 / 041944 A1, suffer from excessive material accumulation radially adjacent to the polymer injection. In contrast, the devices and methods provided herein enable the fabrication of stentless valves by concentrating the deposition in the ventral region of the valve leaflet. The thickness values of the fabricated valve leaflets range from 40 to 300 microns.

[0014] In some aspects, the devices and methods described herein provide: (a) a dual-component design made from a non-metallic component (shield or insulator) and a metallic target; (b) control of the mechanical anisotropy (plane mechanics) of the valve leaflets by varying the tangential velocity of the mandrel; (c) control of the bending stiffness (out-of-plane mechanics) of the valve leaflets by varying the linear velocity of the mandrel; (d) control of the leaflet microfibril orientation (e.g., curved fibers, where the primary direction of alignment is circumferential in the abdominal region and changes axially toward the coaptation region); (e) the ability to construct concave leaflets that mimic native anatomy, shape, and dimensions, which can be varied to replicate healthy or pathological valve anatomy in humans or animals; and (f) the ability to construct fully assembled multi-leaflet stentless valves without a valve conduit.

[0015] In some aspects, the apparatus provided herein includes an electrodeposition target. The target includes a surface, the surface including a pattern of conductive and non-conductive portions, the target mounted on a mandrel having an axis of rotation, and a spindle electrically connected to the conductive portions of the target. The mandrel is rotated during use. In another aspect, the target includes a support disposed about the axis of rotation of the mandrel, a conductive insert including a plurality of ridges extending longitudinally from the support and a plurality of recessed portions between the ridges, and a non-conductive layer on at least a portion of the support and at least a portion of the ridges.

[0016] In some embodiments, the insert includes two concave portions, which are symmetrical or asymmetrical about the axis of rotation of the mandrel. In other embodiments, the concave portions of the insert have the shape of a normal or pathological valve leaflet, such as the shape and size of the cusps (leaflets) of a normal or pathological human or animal mitral, tricuspid, aortic, or pulmonary valve. In some embodiments, three leaflet portions are used to create a tricuspid valve. In other embodiments, two leaflet portions are used to create a bicuspid valve.

[0017] In some aspects, a prosthetic valve formed from a matrix of polymer fibers comprises a tubular (meaning cylindrical, but can have a circular, elliptical, or any closed shape in cross section perpendicular to the longitudinal axis) support portion defining an opening and having a longitudinal axis, and at least two concave leaflets extending longitudinally from the support portion, each leaflet having a concave central portion, a periphery around the concave central portion, a proximal end connected to the support portion, and a distal end longitudinally distal to the support portion, the peripheral portions of adjacent leaflets being partially coapted at and adjacent the support portion forming a coaptation surface between the adjacent leaflets.

[0018] In some aspects, a method of fabricating a valve structure includes electrodepositing a matrix of a biodegradable, biocompatible polymeric composition onto an electrodeposition target.

[0019] In the assembled mandrel structure, the mandrel, annular region, and ridge are covered with insulating ABS. The mandrel consists of three parts: a shield made from the insulator, a conductive target, and a removable axial piece. A polymer, such as poly(ester urethane) urea (PEUU), is electrodeposited around the conductive portion of the target, with some overlap with the insulating cover of the annular portion. The mandrel is placed in a chuck and rotated and moved longitudinally. In some embodiments, the spatial position and relative orientation of the polymer nozzle and mandrel may be controlled manually or by computer using a standard robot or stage, while the mandrel is rotated and the electrodeposition nozzle is not rotated around the mandrel. [Brief explanation of the drawings]

[0020] These and other features and characteristics of the present disclosure, as well as the method of operation and function of the associated elements of structure and the combination of parts and economies of manufacture, will become more apparent from a consideration of the following description and appended claims, taken in conjunction with the accompanying drawings, all of which form a part of this specification, and in which like reference characters indicate corresponding parts in the various views. It is to be expressly understood, however, that the drawings are for the purpose of illustration and description only and are not intended as a definition of the limits of the invention.

[0021] [Figure 1A] FIG. 1A is a schematic diagram of a fully assembled, dual-component, three-leaflet version of the mandrel, according to one embodiment of the present invention. [Figure 1B] FIG. 1B is an exploded view of the dual-component mandrel of FIG. 1A, including, from left to right, a non-conductive plastic shield, a primary collection target, a conductive shaft, and an axial support. [Figure 1C] FIG. 1C is a perspective view of the non-conductive plastic shield of the mandrel of FIG. 1A. [Figure 1D]FIG. 1D is a perspective view of a conductive insert having a three-leaflet heart valve shape of the mandrel of FIG. 1A. [Figure 1E] FIG. 1E is a perspective view of the conductive shaft of the mandrel of FIG. 1A. [Figure 1F] FIG. 1F is a perspective view of an axial support of the mandrel of FIG. 1A. [Figure 2A] FIG. 2A is a side schematic view of a fully assembled, dual-component, two-leaflet version of the mandrel, according to one embodiment of the present invention. [Figure 2B] FIG. 2B is an exploded view of the dual-component mandrel of FIG. 2A. [Figure 2C] FIG. 2C is a perspective schematic view of the non-conductive plastic shield of the mandrel of FIG. 2A. [Figure 2D] FIG. 2D is a perspective schematic view of a conductive insert of the mandrel of FIG. 2A. [Figure 2E] FIG. 2E is a schematic diagram of a variation of the mandrel of FIGS. 1A and 2A. [Figure 2F] FIG. 2F is a schematic diagram of a variation of the mandrel of FIGS. 1A and 2A. [Figure 2G] FIG. 2G is a schematic diagram of a variation of the mandrel of FIGS. 1A and 2A. [Figure 2H] FIG. 2H is a photograph of a prosthetic mitral valve fabricated on a mandrel essentially as shown in FIG. 2F. [Figure 2I] FIG. 21 is a photograph of a prosthetic mitral valve fabricated on a mandrel essentially as shown in FIG. 2F. [Figure 2J] FIG. 2J is a photograph of a prosthetic tricuspid valve fabricated on a mandrel according to FIG. 2E, including a sewing ring. [Figure 3] FIG. 3 is a perspective schematic view of an engineered tricuspid valve made with the mandrel of FIG. 1A. [Figure 4A] FIG. 4A is a perspective schematic view of an engineered bicuspid valve fabricated using the mandrel of FIG. 2A. [Figure 4B] FIG. 4B is a schematic diagram of an engineered bicuspid valve fabricated using the mandrel of FIG. 2A when the valve is in the closed position. [Figure 4C] FIG. 4C is a schematic diagram of an engineered bicuspid valve fabricated using the mandrel of FIG. 2A when the valve is in the closed position. [Figure 4D] FIG. 4D is a schematic diagram of the engineered bicuspid valve made with the mandrel of FIG. 2A when the valve is in the open position. [Figure 4E] FIG. 4E is a schematic diagram of the engineered bicuspid valve made with the mandrel of FIG. 2A when the valve is in the open position. [Figure 5] Figures 5A-5D illustrate the microfiber deposition process. Figure 5A is a photographic image of the dual-component mandrel before polymer fiber deposition. Figure 5B is a photographic image of the dual-component mandrel after 3 hours of polymer fiber deposition, showing selective fiber deposition onto the primary collection target. Figure 5C is a photographic image of the top view of a tri-leaflet valve removed from the mandrel, showing the valve at rest with the leaflets coapted and immersed in PBS (phosphate-buffered saline). Figure 5D is a schematic diagram of the electrospinning manufacturing setup for tri-leaflet cardiac fabrication, showing the mandrel position and its two modes of motion, i.e., mandrel rotational speed ω and mandrel linear speed υ, as well as the voltage generator and injector / pump device. [Figure 6A] FIG. 6A is a photographic image of the side view of a three-leaflet valve removed from the mandrel. [Figure 6B] Figure 6B is a photographic image of the dissected leaflet, which exhibits a concave shape at rest. This process produces a leaflet with a physiological curvature defined by the geometry of the dual-component mandrel described herein. [Figure 7A] FIG. 7A is a graph showing the thickness of engineered heart valves fabricated using the dual-component mandrel described herein in relation to deposition time (independent fabrication for each data point) and the corresponding data points for native porcine tricuspid valves, where the data are presented as mean±standard error (mean±st.e.), demonstrating the scalability of the process with respect to construct thickness by monitoring leaflet thickness versus deposition time. [Figure 7B]Figure 7B shows a thickness map of the thickness distribution of a native porcine tricuspid valve across the leaflet area. Comparison of thickness maps (native vs. prosthetic) shows comparable leaflet thickness values across the entire surface. [Figure 7C] Figure 7C shows a thickness map of the engineered tricuspid valve's thickness distribution across the leaflet area 3 hours after fabrication. Comparison of thickness maps (native vs. prosthetic) shows comparable leaflet thickness values across the entire surface. [Figure 8] Figures 8(A)-(I) provide representative photographic images of engineered valves for nine configurations studied by varying the mandrel tangential velocities V, V2, and V3 (to control anisotropy) and raster velocities R, R1, and R2 (to control bending modulus), demonstrating the feasibility of employing the mandrel designs described herein for different fabrication configurations to control the valve leaflets in both in-plane and out-of-plane mechanisms. The mandrel designs were tested in nine configurations covering the operating region of interest for valve applications (mandrel tangential velocities from 0.3 to 3 m / s, and raster linear velocities from 0 to 2.5 cm / s). [Figure 9A] Figure 9A is a graph showing the in-plane mechanical response of engineered valves tested in biaxial tensile tests in isostress mode for nine configurations obtained by varying the mandrel tangential velocity (ω, controlling for anisotropy) and raster velocity (υ, controlling for bending modulus), n = 3 mean ± standard error (n = 3 mean ± st.e.). [Figure 9B] Figure 9B shows that the anisotropy ratio (AR), defined as the ratio of mechanical strain between the longitudinal and circumferential directions, was used as a metric of anisotropy. The AR versus rotational speed summarizes the results in Figure 9A and shows values for the valve configurations in Figure 9A as well as native porcine tricuspid valves. The appropriate mandrel speeds required to fabricate tricuspid valves with native mechanics (~1.5 m / s) were identified at 0.3, 1.5, and 3 m / s by linear interpolation of the AR. [Figure 9C]Figure 9C shows the biaxial response (valve radial = mandrel longitudinal, valve leaflet circumferential = mandrel circumferential) of an engineered tricuspid valve replicating the AR of a native porcine tricuspid valve, using data generated with a custom-made biological tissue biaxial testing apparatus and protocol previously described in "Biaxial Mechanical Evaluation of Planar Biological Materials" by M. Sacks in Journal of Elasticity and the Physical Science of Solids, July 2000, Vol. 61, Issues 1-3, pp. 199-246. Testing was performed under quasi-static conditions at room temperature, with the specimens continuously immersed in PBS during testing. A preconditioning free-float state was used as the reference configuration, followed by an equiaxial biaxial protocol with a peak of 400 kPa. [Figure 10] Figure 10 is a graph showing the leaflet bending coefficient for different values of mandrel tangential velocity (0.3, 1.5, 3 m / s) and raster velocity (0, 0.25, 2.5 cm / s), n = 3 mean ± standard error. [Figure 11A] Figure 11A provides multiphoton microscopy images of native porcine tricuspid valve microstructure showing the collagen fiber network identified by acquiring collagen second-harmonic generation. A 500 μm × 500 μm × 100 μm volume was analyzed in five different valves and at five different locations within the valve leaflets, including the coaptation plane (upper left TL and upper right TR), the ventral region (center C), and the transition zone between the two (left middle leaflet ML and right middle leaflet MR). The main direction of alignment (n = 5 independent leaflets) is also indicated by white arrows. Note the rotation of the fibers from the coaptation plane to the ventral region. [Figure 11B] FIG. 11B provides a multiphoton microscopy image of the engineered valve leaflet using the same imaging and digital analysis techniques as FIG. 11A. [Figure 12]Figure 12 (A)-(F) provide photographic images showing in vitro testing of the engineered valve leaflet coaptation and suture retention. Electrospun valves were implanted ex vivo into the tricuspid position of a native porcine heart, the right ventricle was gradually filled with saline, and pressure values were monitored with a Millar pressure transducer (mikro-Cath™, Millar, Houston, TX). Simultaneously, images of the coapting leaflets were acquired for healthy native porcine valves (A, B, C) and engineered valves (D, E, F), demonstrating adequate leaflet coaptation and adequate suture retention with Δρ > 30 mmHg. [Figure 13]Figures 13(A) and 13(B) provide graphical representations of valve function during in vitro testing of (A) the dynamics of a state-of-the-art commercial prosthetic valve (n = 5 Carpentier-Edwards® Duraflex™) and (B) the dynamics of an engineered valve (n = 3). Pulsatile flow across the valve was generated by a commercially available ventricular assist device, a thoracic percutaneous VAD system, operating at a constant frequency of 70 beats per minute. Two cameras continuously detected valve motion, and pressure and flow were recorded by sensors. Digital image processing was performed to detect orifice area using a custom Matlab code (Mathworks®, Natick, MA). Figures 13(A) and 13(B) show the detected orifice area (white) for the Carpentier-Edwards® and engineered valves, respectively, during systole. Figure 13(C) is a bar graph representation of the bending deformation index (BDI) for the Carpentier-Edwards bioprosthetic and engineered valves in Figures 13(A) and 13(B), respectively. BDI, a widely adopted metric for bending stiffness (see Non-Patent Document 2, "In vitro hydrodynamics, cusp-bending deformation, and root distensibility for different types of aortic valve-sparing operations: Remodeling, sinus prosthesis, and reimplantation" by A. Erasmi et al. in The Journal of Thoracic and Cardiovascular Surgery, Volume 130, Issue 4, October 2005, pp. 1044-1049), was calculated at mid-diastole. Figure 13(D) is a bar graph representation of the geometric orifice area (GEO) comparison of the Carpentier-Edwards bioprosthetic valve and the engineered valves of Figures 13(A) and 13(B). GEO was calculated from the image processing shown in (A) and (B) at peak systole.Figure 13(E) is a bar graph representation of the peak systolic pressure for the Carpentier-Edwards bioprosthetic valve and the engineered valves of Figures 13(A) and 13(B). Figure 13(F) is a bar graph representation of the mean systolic pressure for the bioprosthetic valve and the engineered valves of Figures 13(A) and 13(B). Figure 13(G) is a bar graph representation of the mean pressure drop across the Carpentier-Edwards bioprosthetic valve and the engineered valves of Figures 13(A) and 13(B) during a complete cycle including systole and diastole. Figure 13(H) is a bar graph representation of the mean flow rate across the Carpentier-Edwards bioprosthetic valve and the engineered valves of Figures 13(A) and 13(B) during a complete cycle including systole and diastole. [Figure 14] Figure 14 shows photographic representations of (A) a flow duplicator for assessing valve function under physiological flow conditions, using a 40% glycerol solution to mimic blood viscosity, (B) an engineered valve retention device with white arrows pointing to the engineered valve, and (C) a schematic diagram of the flow duplicator, where 1) is a desktop computer, 2-4) represent the pressure and flow signal acquisition system, 5) is a preload sensor, 6) is a flow meter, 7) is a capacitor, 8) is a pressure sensor, 9) is a flanged valve holder, 10) is a side camera, 11) is a thoracic percutaneous VAD system (Thoratec Corporation, Pleasanton, CA), 12) is a front camera, and 13) is a VAD controller. DETAILED DESCRIPTION OF THE INVENTION

[0022] The use of numerical values in the various ranges specified in this application, unless otherwise stated, are stated as approximations, as if the word "about" were appended to both the minimum and maximum values within the stated range. In this manner, slight variations above and below the stated ranges can be used to achieve substantially the same results as values within the range. Also, unless otherwise indicated, the disclosure of these ranges is intended as a continuous range, including every value between the minimum and maximum values.

[0023] As used herein, the terms "comprising," "comprise," or "comprised," and variations thereof, are meant to be open-ended. The terms "a" and "an" are intended to refer to one or more.

[0024] As used herein, the term "target" means a point on the surface of a charged object at or adjacent to which polymeric material is to be electrodeposited, such that in the absence of an electrical insulator, electrodeposition would be affected by the presence of an electrical charge at that point during electrodeposition. As such, target does not include a portion of the surface of a charged object where, in the absence of an electrical insulator, an electrical charge at that portion of the surface of the charged object would not affect electrodeposition.

[0025] As used herein, "treatment" or "treating" a wound or defect means the administration of any appropriate regimen, procedure and / or method of administration of a composition, device or structure to a patient for the purpose of achieving a desired clinical / medical endpoint, including attracting progenitor cells, healing the wound, correcting the defect, etc.

[0026] As used herein, the term "patient" or "subject" refers to any member of the animal kingdom, including, but not limited to, humans, and "animal" refers to all mammals, including, but not limited to, humans.

[0027] A biodegradable polymer composition is one in which the polymer and its degradation products are substantially non-toxic to cells or organisms within acceptable tolerances, substantially non-carcinogenic, substantially non-immunogenic, and without substantial toxic effects in biological systems, e.g., living organisms (patients). Non-limiting examples of degradation mechanisms within biological systems include chemical reactions, hydrolysis reactions, and enzymatic cleavage.

[0028] As used herein, the term "polymer composition" refers to a composition comprising one or more polymers. As a class, "polymers" include, but are not limited to, homopolymers, heteropolymer-co-polymers, block polymers, and block copolymers, and can be both natural and synthetic. Homopolymers contain one type of building block or monomer, while copolymers contain two or more types of monomer. The term "(co)polymer" and similar terms refer to either homopolymers or copolymers.

[0029] A polymer "comprises" or "is derived from" a recited monomer if that monomer is incorporated into the polymer. Thus, the incorporated monomer that the polymer comprises is not the same as the monomer prior to incorporation into the polymer, at least in that certain groups are lost and / or modified upon incorporation into the polymer scaffold. A polymer is said to comprise a particular type of bond if that bond is present in the polymer.

[0030] As used herein, a "fiber" is an elongated, slender thread-like and / or filament-like structure. A "matrix" is any two- or three-dimensional arrangement of elements, either ordered (e.g., in a woven or non-woven mesh) or randomly arranged (typically a mat of fibers produced by electrospinning), and may be isotropic or anisotropic.

[0031] As used herein, the term "polymer" refers to both synthetic and biological polymer components. A "biological polymer" is a polymer that can be obtained from a biological source, such as mammalian or vertebrate tissue, as in the case of certain extracellular matrix-derived (ECM-derived) compositions described herein. Biological polymers may be modified by additional processing steps. Polymers generally include, for example, but are not limited to, monopolymers, copolymers, polymer blends, block polymers, block copolymers, crosslinked polymers, non-crosslinked polymers, linear, branched, comb-like, star-shaped, and / or dendritic polymer(s), where the polymers can be formed into any useful form, for example, but not limited to, hydrogels, porous meshes, fibers, woven meshes, or nonwoven meshes formed by electrospinning.

[0032] By "biodegradable" or "bioerodible," it is meant that the polymer, once implanted and placed in contact with bodily fluids and tissues, typically, and often preferably, partially or completely degrades over a period of hours, days, weeks, or months via chemical reactions with bodily fluids and / or tissues. Non-limiting examples of such chemical reactions include acid / base reactions, hydrolysis reactions, and enzymatic cleavage. The biodegradation rate of the polymer matrix may be engineered, optimized, or otherwise tailored to ensure that the matrix degrades over a useful period of time. Typically, one or more polymers are selected to degrade in situ over a period of time to optimize mechanical adjustment of the tissue. For example, for abdominal wall repair, it is desirable for the matrix to dissolve over at least one week, and preferably longer. More importantly, the matrix must retain its supportive capacity until tissue remodeling occurs, such as for at least two to eight weeks or longer.

[0033] The valve structures described herein can be fabricated from any biocompatible material. In some examples below, the valve structures are fabricated from urethanes, specifically poly(ester-urethane)urea (PEUU), synthesized using putrescine as a chain extender and the described two-step solvent synthesis method. The valve structures were fabricated using PEUU by electrospinning. PEUU's characteristics include high elasticity and mechanical strength combined with controllable biodegradability and cell adhesion. The polymer composition finds use in many in vivo scenarios, such as cardiac components in abdominal wall repair and vascular grafts. Alternative chemistries include non-thrombogenic chemical components added to polyurethanes, allowing the non-degradable polyurethane to be used as a permanent structure without in-situ remodeling. Additional biodegradable polymer compositions are known in the art and exhibit suitable strength and elasticity for use with or in place of the described PEUU.

[0034] The valve structure optionally comprises biopolymeric components such as biodegradable elastomeric polymer components and / or extracellular matrix (ECM) gel.

[0035] In one embodiment, the valve structure is fabricated from a synthetic polymer composition. In another embodiment, the polymer composition combines a synthetic polymer with an ECM gel, as described in U.S. Patent Application Publication No. 2002 / 02390. While the ECM gel component is useful for promoting cell growth (including, but not limited to, one or more of colonization, proliferation, invasion, cell viability, differentiation, and tissue repair), it lacks sufficient strength for use as a structural tissue repair scaffold in a patient. When the synthetic polymer and ECM gel are mixed, any ratio of biodegradable elastomeric polymer to ECM gel that exhibits adequate tensile strength and elasticity while providing excellent cell infiltration may be used; for example, a useful ratio of polymer to gel is 70%-85% to 15%-30%, including increments therebetween. This can be achieved by co-depositing the biodegradable elastomeric polymer and ECM gel via electrospinning. For example, the synthetic biodegradable elastomeric polymer is electrospun and the ECM gel is sprayed, e.g., electrosprayed.

[0036] In the broadest sense, to produce an ECM gel, by one non-limiting example, ECM-derived scaffold material, e.g., decellularized or devitalized tissue, is coagulated and solubilized to form a hydrogel. In one example, the solubilized hydrogel is not dialyzed. Solubilization can be achieved by digestion with appropriate degradative enzymes, such as the endoproteases trypsin, chymotrypsin, pepsin, papain, and elastase. In a specific non-limiting example, a method for making such a gel includes (i) disrupting the extracellular matrix; (ii) solubilizing the intact, undialyzed, and / or uncrosslinked extracellular matrix in an acidic solution (e.g., 0.01 N HCl) at a pH of, e.g., about 2.0, to produce a digestion solution; (iii) increasing the pH of the digestion solution to 7.2-7.8 to produce a neutralized digestion solution; and (iv) gelling the solution at a temperature greater than about 25°C.

[0037] "ECM materials" are materials made from extracellular matrix-containing tissues, including decellularized or devitalized tissues. ECM materials may be used to generate gels according to the methods, compositions, and devices described herein (see generally U.S. Patent No. 4,902,508; U.S. Patent No. 4,956,178; U.S. Patent No. 5,281,422; U.S. Patent No. 5,352,463; U.S. Patent No. 5,372,821; U.S. Patent No. 5,554,389; U.S. Patent No. 5,573,784; U.S. Patent No. 5,573,784). See U.S. Patent No. 5,645,860, U.S. Patent No. 5,771,969, U.S. Patent No. 5,753,267, U.S. Patent No. 5,762,966, U.S. Patent No. 5,866,414, U.S. Patent No. 6,890,562, U.S. Patent No. 6,890,563, U.S. Patent No. 6,890,564, and U.S. Patent No. 6,893,666.

[0038] In certain embodiments, the ECM material is decellularized tissue prepared from a mammal, such as, but not limited to, vertebrate tissue, including, but not limited to, human, monkey, pig, cow, and sheep. The ECM material can be prepared from any organ or tissue, including, but not limited to, the bladder, intestine, liver, esophagus, and dermis. In one embodiment, the ECM material is decellularized tissue isolated from bladder tissue. The ECM material may or may not include a basement membrane portion of the tissue. In certain embodiments, the ECM material includes at least a portion of a basement membrane. In certain embodiments, the ECM material is prepared from pericardium or valve leaflets obtained from a pig, cow, horse, monkey, or human, such as bovine pericardium or porcine valve leaflets.

[0039] As an example, decellularized tissue is isolated from a removed porcine bladder to create a urinary bladder matrix (UBM). Excess connective tissue and residual urine are removed from the bladder. The serosa, tunica muscularis, submucosa, and most of the muscularis mucosa can be removed by mechanical abrasion or a combination of enzymatic treatment, hydration, and abrasion. Mechanical removal of these tissues can be achieved by abrasion using a longitudinal wiping motion to remove the outer layer (particularly the abluminal smooth muscle layer) and even the luminal portion of the outer mucosa (epithelial layer). Mechanical removal of these tissues can be achieved, for example, by removing the mesenteric tissue with Adson-Brown forceps and Metzenbaum scissors, and then wiping off the tunica muscularis and submucosa using a longitudinal wiping motion with a scalpel handle or other rigid object wrapped in moist gauze. Epithelial cells of the muscularis mucosa can also be dissociated by immersing the tissue in a de-epithelialization solution, such as, but not limited to, hypertonic saline. The resulting UBM contains the basement membrane of the muscularis mucosa and the adjacent lamina propria.

[0040] In another example, epithelial cells are first stripped by immersing the tissue in a de-epithelializing solution, such as, but not limited to, hypertonic saline, for a period ranging from 10 minutes to 4 hours. Exposure to hypertonic saline effectively removes the epithelial cells from the underlying basement membrane. The tissue remaining after the initial stripping procedure includes the epithelial basement membrane and the abluminal tissue layer of the epithelial basement membrane. This tissue is then further processed to remove most of the intraluminal tissue, but not the epithelial basement membrane. The outer serosa, adventitia, smooth muscle tissue, submucosal capsule, and most of the muscularis mucosa are removed from the remaining de-epithelialized tissue by mechanical abrasion or a combination of enzymatic treatment, hydration, and abrasion.

[0041] In one example, decellularized tissue is created by abrading porcine bladder tissue to remove the outer layers, including both the serosa and muscularis, using a longitudinal wiping motion with a scalpel handle and moist gauze. Following eversion of the tissue segment, the luminal portion of the inner mucosa is peeled away from the underlying tissue using the same wiping motion. After these tissues are removed, care is taken to avoid perforating the submucosa, and the resulting ECM material consists primarily of submucosa.

[0042] The ECM material can be decellularized, sterilized, and / or dried by any useful method. The ECM material can be sterilized by any of several standard methods without losing its ability to induce endogenous tissue growth. For example, the material can be sterilized by propylene oxide or ethylene oxide treatment, gamma irradiation (0.05 to 4 mRad), gas plasma sterilization, peracetic acid sterilization, or electron beam treatment. The material can also be sterilized by treatment with glutaraldehyde, which causes cross-linking of the protein material; however, this treatment essentially alters the material so that it is gradually or not resorbed at all, resulting in a different type of host remodeling that more closely resembles the formation of scar tissue or encapsulation rather than contrast remodeling. Cross-linking of the protein material can also be induced using carbodiimide, dehydrothermal, or photooxidation methods. More typically, the ECM is disinfected by immersion in 0.1% (v / v) peracetic acid (σ), 4% (v / v) ethanol, and 96% (v / v) sterile water for 2 hours. The decellularized tissue is then washed twice for 15 minutes with PBS (pH = 7.4) and twice for 15 minutes with deionized water.

[0043] Commercially available ECM materials derived from small intestinal submucosa (SIS) include, but are not limited to, Surgisis®, Surgisis-ES®, Stratasis®, and Stratasis-ES® (Cook Urological; Indianapolis, Indiana) and GraftPatch® (Organogenesis; Canton, Massachusetts). In another example, the ECM material is derived from dermis. Commercially available preparations include, but are not limited to, Pelvicol® (cross-linked porcine dermal collagen sold in Europe as Permacol®, Bard Medical Division, Covington, GA), Repliform® (Microvasive; Boston, Massachusetts), and Alloderm® (LifeCell, Branchburg, New Jersey). In another example, the ECM is derived from the urinary bladder. Commercially available products include, but are not limited to, UBM (Acell, Jessup, Maryland).

[0044] In one non-limiting example, decellularized tissue is freeze-dried, pulverized, and then solubilized with an acidic degradative enzyme. In certain aspects, the decellularized tissue is not dialyzed and / or cross-linked (subjected to a cross-linking method) prior to digestion with the acidic degradative enzyme. The acidic degradative enzyme may be, but is not limited to, pepsin or trypsin, and in one example, is pepsin. Decellularized tissue is typically solubilized at an acidic pH suitable for or optimal for the degradative enzyme, for example, in a 0.01 M HCl solution (pH ∼2), between pH 1.5 and 3. This solution is typically solubilized for 12 to 48 hours by mixing (e.g., stirring, agitating, blending, rotating, tilting, etc.), depending on the type of tissue. Once the decellularized tissue is solubilized, the pH rises to 7.2-7.8, and in one example, the pH rises to 7.4. A base, such as a hydroxyl ion-containing base, including NaOH, may be used to raise the pH of the solution. Similarly, buffers such as isotonic buffers, including but not limited to phosphate buffered saline (PBS), may be used to bring the solution to a target pH, such as physiological pH and ionic conditions, or to help maintain the pH and ionic strength of the gel at a target level. Although gelation proceeds more rapidly at temperatures above 30°C and as the temperature approaches physiological temperature (37°C), the neutralized digestion solution is gelled at any temperature approaching 37°C, typically above 25°C. This method typically does not include a dialysis step prior to gelation and results in a more intact ECM-like matrix that typically gels more slowly at 37°C than comparable collagen or dialyzed ECM constructs.

[0045] ECM gels can be sprayed, for example, as a liquid or hydrogel and may be combined with other polymers as described herein. ECM gels form hydrogels when their temperature is elevated and may have an LCST (lower critical solution temperature) above or below the temperature at which the solution is sprayed. Therefore, they are reverse-gelling, meaning that the gel transition occurs at temperatures above, equal to, or below the temperature at which the ECM gel is sprayed. For example, if a hydrogel is sprayed at room temperature (i.e., approximately 20-25°C) or below, and the LCST of the ECM material is higher than the spray temperature but less than, for example, 37°C, the material can be sprayed and will later gel when warmed. See, for example, U.S. Patent Application Publication No. 2008 / 0260831, the technical disclosure of which is incorporated herein by reference. See also Non-Patent Document 3 (Stankus et al., Hybrid nanofibrous scaffolds from electrospinning of a synthetic biodegradable elastomer and urinary bladder matrix, J Biomater. Sci. Polym. Ed. (2008) 19(5):635-652.) In Non-Patent Document 3 (Stankus' paper), PEUU was mixed with solubilized UBM ECM and electrospun.

[0046] In general, polymeric components suitable for the anatomical artificial structures described herein are any polymers that are biocompatible and biodegradable. In certain non-limiting examples, the biodegradable polymer may comprise a homopolymer, copolymer, and / or polymer blend, including, but not limited to, one or more of the following monomers: glycolide, lactide, caprolactone, dioxanone, and trimethylene carbonate. In other non-limiting examples, the polymer may comprise labile chemical moieties, such as, but not limited to, esters, anhydrides, polyanhydrides, or amides, that are useful for controlling the degradation rate of the scaffold and / or the release rate (if applicable) of a therapeutic agent from the scaffold. Alternatively, the polymer may comprise a polypeptide or biopolymer as a building block that is susceptible to chemical reaction once deployed in situ. In one non-limiting example, the polymer composition comprises a polypeptide comprising the amino acid sequence alanine-alanine-lysine, which confers enzymatic lability to the polymer. In another non-limiting example, the polymer composition may comprise a biopolymer component derived from the extracellular matrix (ECM). For example, as described in more detail below, the polymer composition may comprise the biopolymer collagen so that collagenase present in situ can degrade the collagen. As used herein, a polymer may be an elastomer, meaning that it changes shape upon application of a deforming force and substantially returns to its original shape when the deforming force is removed.

[0047] In another non-limiting example, the synthetic polymer component comprises hydrolytically, chemically, biochemically, and / or proteolytically labile groups, including, by way of non-limiting example, ester moieties, amide moieties, anhydride moieties, specific peptide sequences, and general peptide sequences.

[0048] A number of biocompatible and biodegradable elastomeric (co)polymers are known and established as useful for creating cell growth matrices, including biodegradable poly(ester urethane) urea (PEUU), poly(ether ester urethane) urea (PEEUU), poly(ester carbonate) urethane urea (PECUU), and poly(carbonate) urethane urea (PCUU). Generally, useful (co)polymers comprise monomers derived from alpha hydroxy acids, including polylactide, poly(lactide-co-glycolide), poly(L-lactide-co-caprolactone), polyglycolic acid, poly(di-lactide-co-glycolide), poly(l-lactide-co-dl-lactide); monomers derived from esters, including polyhydroxybutyrate, polyhydroxyvalerate, polydioxanone, and polygalactin; monomers derived from lactones, including polycaprolactone; monomers derived from carbonates, including polycarbonate, polyglyconate, poly(glycolide-co-trimethylenecarbonate), poly(glycolide-co-trimethylenecarbonate-co-dioxanone); and monomers linked through urethane linkages, including polyurethanes, poly(ester urethane) urea elastomers.

[0049] In certain embodiments, the polymers used to create the structures described herein also release therapeutic agents when they degrade within a patient's body. For example, individual building blocks of the polymer can be selected to provide a therapeutic benefit when the building blocks themselves are released in situ through the degradation process. In one example, one of the polymer building blocks is putrescine, which is involved as a substance that induces cell proliferation and differentiation.

[0050] The biodegradable polymer may be, but is not limited to, a homopolymer, copolymer, and / or polymer blend. According to certain embodiments, the polymer comprises one or more monomers, including, but not limited to, glycolide, lactide, caprolactone, dioxanone, and trimethylene carbonate. According to certain embodiments, the polymer comprises a polymer derived from an alpha-hydroxy acid, polylactide, poly(lactide-co-glycolide), poly(L-lactide-co-caprolactone), polyglycolic acid, poly(dl-lactide-co-glycolide), poly(l-lactide-co-dl-lactide), lactone monomer, polycaprolactone, carbonate linkage, polycarbonate, polyglyconate, poly(trimethylene carbonate), poly(glycolide-co-trimethylene carbonate), poly(glycolide-co-trimethylene carbonate-co-dioxanone). Examples of polymers include polymers containing urethane linkages, polyurethanes, poly(esterurethane)ureas, poly(etheresterurethane)urea elastomers, poly(carbonateurethane)ureas, polycarbonateurethanes, polyesterurethanes, polymers containing ester linkages, polyalkanoates, polyhydroxybutyrates, polyhydroxyvalerates, polydioxanone, polygalactin, natural polymers, chitosan, collagen, elastin, alginates, cellulose, hyaluronic acid, and gelatin. In one embodiment, the polymer composition comprises poly(esterurethane)urea with collagen in an amount of about 25% to about 75% by weight. The polymer composition may also comprise elastin-collagen or a mixture thereof, for example, but not limited to, a mixture of collagen and elastin in an amount of about 25% to about 75% by weight, approximately the same as in one embodiment. In one non-limiting embodiment, the polymer is polycaprolactone. In another embodiment, the polymer comprises polycaprolactone diol. In yet another embodiment, the polymer is a triblock copolymer comprising polycaprolactone, poly(ethylene glycol) and polycaprolactone blocks.

[0051] In another non-limiting example, the polymer composition includes a biopolymer component derived from the ECM. For example, the polymer composition may include the biopolymer collagen, such that collagenase present in situ can degrade the collagen. As an example, the polymer composition may include one or both of collagen and elastin. Collagen is a common ECM component and typically degrades in vivo faster than many synthetic biodegradable polymers. Therefore, manipulating the collagen content in the polymer composition can be used as a method to modify the in vivo bioerosion rate. Collagen may be present in the polymer composition in any useful range, including, but not limited to, about 2% to about 95% by weight, e.g., about 25% to about 75% by weight, including all ranges and points therebetween, and including about 40% to about 75% by weight, including about 75% and about 42.3% by weight. Elastin may also be incorporated into the polymer composition to enhance elasticity. Elastin may be present in the polymer composition in any useful range, including, but not limited to, about 2% to about 50% by weight, including all ranges and points therebetween, including from about 40% to about 42.3% by weight, including all integers and points therebetween and their equivalents. In one non-limiting example, collagen and elastin are present in the polymer composition in approximately equal amounts. In another example, the total collagen and elastin content in the polymer composition ranges from about 2% to about 95% by weight, such as from about 25% to about 75% by weight, including all ranges and points therebetween, including about 40% to about 75% by weight, including about 75% and about 42.3% by weight.

[0052] In one non-limiting example, the polymer composition comprises a biodegradable poly(ester urethane) urea elastomer (PEUU). PEUUs can be made by reacting a diol with a diisocyanate to form a prepolymer, and then reacting the prepolymer with a diamine. A non-limiting example of such a PEUU is an elastomeric polymer made from polycaprolactone diol (Mw 2000) and 1,4-diisocyanatobutane using a diamine chain extender such as putrescine. One non-limiting example, or method, for making a PEUU polymer is a two-stage polymerization process whereby polycaprolactone diol (Mw 2000), 1,4-diisocyanatobutane, and a diamine are combined in a 2:1:1 molar ratio. In the first step to form a prepolymer, a 15% by weight solution of 1,4-diisocyanatobutane in DMSO (dimethyl sulfoxide) is continuously stirred with a 25% by weight solution of polycaprolactone diol in DMSO. Stannous octoate is then added, and the mixture is allowed to react at 75°C for 3 hours. In the second step, the prepolymer is reacted with a diamine to extend the chain and form a polymer. In one example, the diamine is putrescine, which is added dropwise with stirring and allowed to react at room temperature for 18 hours. In one example, the diamine is lysine ethyl ester, which is dissolved in DMSO with triethylamine and added to the prepolymer solution and allowed to react at 75°C for 18 hours. After the two-step polymerization process, the polymer solution is precipitated in distilled water. Next, the wet polymer is immersed in isopropanol for 3 days to remove unreacted monomers. Finally, the polymer is dried under vacuum at 50°C for 24 hours.

[0053] In another non-limiting example, the polymer composition comprises a poly(ether ester urethane) urea elastomer (PEEUU). For example, but not by way of limitation, PEEUU can be produced by reacting a polycaprolactone-b-polyethylene glycol-b-polycaprolactone triblock copolymer with 1,4-diisocyanatobutane and putrescine. In one non-limiting example, PEEUU is obtained by a two-step reaction using a 2:1:1 reactant stoichiometry of 1,4-diisocyanatobutane:triblock copolymer:putrescine. According to one non-limiting example, a triblock polymer can be prepared by reacting poly(ethylene glycol) and ε-caprolactone with stannous octoate at 120°C for 24 hours under a nitrogen atmosphere. The triblock copolymer is then washed with ethyl ether and hexane and then dried in a vacuum oven at 50°C. In the first step to form the prepolymer, a 15 wt% solution of 1,4-diisocyanatobutane in DMSO is continuously stirred with a 25 wt% solution of triblock copolymer in DMSO. Stannous octoate is then added, and the mixture is allowed to react at 75°C for 3 hours. In the second step, putrescine is added dropwise to the prepolymer solution with stirring, and the reaction is allowed to continue at room temperature for 18 hours. The PEEUU polymer solution is then precipitated with distilled water. The wet polymer is immersed in isopropanol for 3 days to remove unreacted monomers. It is vacuum dried at 50°C for 24 hours.

[0054] In another non-limiting example, the polymer composition comprises poly(ester carbonate)urethane urea (PECUU) or poly(carbonate)urethane urea (PCUU), as described, for example, in Hong et al., "Tailoring the degradation kinetics of poly(ester carbonate urethane) urea thermoplastic elastomers for tissue engineering scaffolds: Biomaterials, 31 (2010) 4249-4258." Poly(ester carbonate urethane) urea (PECUU) is synthesized, for example, using a blend of soft segments of polycaprolactone (PCL) and poly(1,6-hexamethylene carbonate) (PHC) and a hard segment of 1,4-diisocyanatobutane (BDI) with chain extension by putrescine. Different molar ratios of PCL and PHC may be used to achieve different physical properties. Putrescine is used as a chain extender by a two-stage solvent synthesis method. In one example, the molar ratio of (PCL + PHC):BDI:putrescine was determined to be 1:2:1. Varying molar ratios of PCL and PHC (e.g., 100 / 0 (yielding PEUU), 75 / 25, 50 / 50, 25 / 75, and 0 / 100 (yielding PCUU)) were completely dissolved in DMSO in an argon-protected three-neck flask. BDI was then added to the solution, followed by four drops of Sn(Oct)2. The flask was placed in an oil bath at 70°C. After 3 hours, the prepolymer solution was cooled to room temperature, and then the putrescine / DMSO solution was added dropwise to the stirred solution. The final polymer solution concentration was controlled to be approximately 4% (w / v). The flask was then placed in an oil bath and kept at 70°C overnight. The polymer was precipitated in an excess amount of cold deionized water and then vacuum dried at 60°C for 3 days. The polyurethaneureas synthesized from the different PCL / PHC molar ratios defined above are designated as PEUU, PECUU 75 / 25, PECUU 50 / 50, PECUU 25 / 75, and PCUU, respectively. In practice, the yield of the final product using this method is about 95%.

[0055] Diamines and diols are useful building blocks for preparing the (co)polymer compositions described herein. Diamines, as described above, have the structure "H2N-R-NH2," where "R" is an aliphatic or aromatic hydrocarbon or a hydrocarbon containing aromatic and aliphatic regions. The hydrocarbon may be linear or branched. Examples of useful diamines are putrescine (R = butylene) and cadaverine (R = pentylene). Useful diols include polycaprolactone (e.g., Mw 1000-5000) and multiblock copolymers such as polycaprolactone-PEG copolymers, including polycaprolactone-b-polyethylene glycol-b-polycaprolactone triblock copolymers of various sizes. Other building blocks for useful diols include, but are not limited to, glycolides (e.g., polyglycolic acid (PGA)), lactide, dioxanone, and trimethylene carbonate. Diisocyanates have the general structure OCN-R-NCO, where "R" is an aliphatic or aromatic hydrocarbon or a hydrocarbon containing aromatic and aliphatic regions. The hydrocarbon may be straight or branched chain.

[0056] In a further example, the polymer composition can include polyethylene terephthalate (PET, e.g., DACRON®). Notably, PET is less biodegradable and more rigid than the copolymers described above. PET scaffold structures are essentially made using the methods described herein for PEUU and other polymer compositions described herein. Polymer concentrations and infusion rates can be varied to accommodate different properties of PET compositions, such as, without limitation, 20% w / v for PET in HFIP at a 12 mL / h infusion rate, as used in the following examples.

[0057] In another embodiment, the polymer composition comprises polyarylate tyrosine (TPA). Like PET, TPA is less biodegradable and more rigid than the polyurethane copolymers described above. TPA scaffold structures are essentially made using the methods described herein for PEUU and other polymer compositions. The polymer concentration and infusion rate may be varied to accommodate different properties of the TPA composition, for example, without limitation, for 12% w / v TPA in HFIP at an infusion rate of 20 mL / h. Tyrosine polyarylates are generally made from fatty acids and tyrosine-derived diphenols. Non-limiting examples of useful fatty acids include succinic acid, adipic acid, sebacic acid, and dicarboxylic acid chlorides or anhydrides. Non-limiting examples of tyrosine-derived diphenols include desaminotyrosyl-tyrosine alkyl esters (DTE), where the alkyl is one of, for example, ethyl, hexyl, and octyl. As an example, poly(DTE-co-27.5DT succinate ester) is used. TPA and methods for producing TPA are described, for example, in U.S. Patent No. 5,216,115 and U.S. Patent Publication No. 2011 / 0082545, which disclose useful TPAs and whose technical disclosures are incorporated herein by reference.Further references disclosing TPA compositions and methods of making and using these compositions include Fiordeliso, J. et al., "Design, synthesis, and preliminary characterization of tyrosine-containing polyarylates: new biomaterials for medical applications, J. Biomater Sci. Polym. Ed. 1994;5(6):497-510," Huang, X. et al., "A library of L-tyrosine-derived biodegradable polyarylates for potential biomaterial applications, part I: synthesis, characterization and accelerated hydrolytic degradation, J. Biomater Sci. Polym. Ed. 2009;20(7-8):935-55," and Bourke, S. L. et al., "Polymers derived from the amino acid L-tyrosine: polycarbonates, polyarylates, and copolymers with poly(ethylene glycol) Adv. Drug Deliv. Rev. 2003 Apr. 25;55(4):447-66."

[0058] In another example, at least one therapeutic agent is added to a scaffold or composition described herein before it is implanted into a patient or otherwise administered to a patient. Generally, therapeutic agents include any substance coated, embedded, imbibed, absorbed, or otherwise attached to a structure, or incorporated into a structure or drug that will provide a therapeutic benefit to a patient. Non-limiting examples of such therapeutic agents include antimicrobial agents, growth factors, emollients, retinoids, and topical steroids. Each therapeutic agent can be used alone or in combination with other therapeutic agents. For example, but not limited to, structures comprising neurotrophic factors or cells expressing neurotrophic factors can be applied to wounds near critical regions of the central nervous system, such as the spine. Alternatively, the therapeutic agent can be blended with the polymer while the polymer is being processed. For example, the therapeutic agent can be dissolved in a solvent (e.g., DMSO) and added to the polymer blend during processing. In another example, the therapeutic agent is mixed with a carrier polymer (e.g., polylactic-co-glycolic acid microparticles) and then processed with an elastomeric polymer. By mixing the therapeutic agent with the carrier polymer or the elastomeric polymer itself, the release rate of the therapeutic agent can be controlled by the rate of polymer degradation.

[0059] In specific, non-limiting examples, the therapeutic agent is a growth factor, such as a neurotrophic factor or an angiogenic factor, optionally produced using recombinant technology. Non-limiting examples of growth factors include basic fibroblast growth factor (bFGF), acidic fibroblast growth factor (aFGF), vascular endothelial growth factor (VEGF), human vascular endothelial growth factor-165 (hVEGF), and the like. 165), vascular endothelial growth factor A (VEGF-A), vascular endothelial growth factor B (VEGF-B), hepatocyte growth factor (HGF), insulin-like growth factors 1 and 2 (IGF-1 and IGF-2), platelet-derived growth factor (PDGF), stromal-derived factor 1 alpha (SDF-1α), nerve growth factor (NGF), ciliary neurotrophic factor (CNTF), neurotrophin-3, neurotrophin-4, neurotrophin-5, pleotrophin protein (neurite outgrowth-promoting factor 1), midkine protein (neurite outgrowth-promoting factor 2), brain-derived neurotrophic factor (BDNF), tumor angiogenic factor (TAF), corticotropin-releasing factor (CRF), transforming growth factors alpha and beta (TGF-α and TGF-β), interleukin-8 (IL-8), granulocyte-macrophage colony-stimulating factor (GM-CSF), interleukins, and interferons. Commercially available preparations of various growth factors, including neurotrophic and angiogenic factors, are available from R&D Systems, Minneapolis, Minnesota; Biovision, Mountain View, California; ProSpec-Tany TechnoGene, Rehovot, Israel; and Cell Sciences®, Canton, Massachusetts.

[0060] In certain non-limiting examples, the therapeutic agent is an antibacterial agent such as, but not limited to, isoniazid, ethambutol, pyrazinamide, streptomycin, clofazimine, rifabutin, fluoroquinolones, ofloxacin, sparfloxacin, rifampin, azithromycin, clarithromycin, dapsone, tetracycline, erythromycin anthraquinone, paromomycin, diclozalil, acyclovir, trifluorouridine, foscarnet, penicillin, gentamicin, ganciclovir, iatronazole, miconazole, Zn-pyrithione, and silver salts such as chloride, bromide, iodide, and periodate.

[0061] In specific, non-limiting examples, the therapeutic agent is a nonsteroidal anti-inflammatory agent such as, but not limited to, salicylic acid, indomethacin, indomethacin trihydrate sodium, salicylamide, naproxen, colchicine, fenoprofen, sulindac, diflunisal, diclofenac, indoprofen, sodium salicylamide; an anti-inflammatory cytokine; an anti-inflammatory protein; a steroidal anti-inflammatory agent; or an anticoagulant such as heparin; or a nitro-fatty acid such as nitro-oleic acid or nitro-conjugated linoleic acid. Other agents that may promote wound healing and / or tissue regeneration may also be included.

[0062] The structures described herein are preferably fabricated by electrospinning a biodegradable elastomeric polymer and co-depositing an ECM gel and / or, where appropriate, a blood product or other liquid by spraying, e.g., electrospraying. Other compounds or components may be incorporated into structures as described herein by any method, such as absorption, adsorption, mixing, etc.

[0063] The deposited biodegradable elastomeric polymer is typically porous. As used herein, the term "porosity" refers to the ratio between the volume of all pores within a polymer composition and the volume of the entire polymer composition. For example, a polymer composition with 85% porosity would have 85% of its volume containing pores and 15% of its volume containing polymer. In certain non-limiting examples, the porosity of the structure is at least 60%, 65%, 70%, 75%, 80%, 85%, or 90%, or any increment therebetween. In other non-limiting examples, the average pore size of the structure is 0.1 to 300 microns, 0.1 to 100 microns, or 1 to 25 microns, including increments therebetween. By way of example and not limitation, structures that act as barriers to bacteria and other pathogens may have an average pore size of less than 0.5 microns or less than 0.2 microns. In one example, the structures described herein are fabricated by electrospinning. Therefore, it is often advantageous to adjust the pore size or porosity by varying the polymer concentration in the electrospinning solution or by varying the spinning distance from the nozzle to the target. For example, but not by way of limitation, the average pore size can be increased by increasing the amount of polymer component in the suspension used for electrospinning, resulting in larger fiber diameters and larger pore sizes. In another non-limiting example, the average pore size can be increased by increasing the spinning distance from the nozzle to the target, resulting in decreased adhesion between the fibers and the looser matrix. When ECM gel is co-deposited during electrospinning, many of the pores (i.e., a large proportion of the pores or interstices) in the deposited polymer are filled with ECM gel.

[0064] In certain embodiments, electrospinning is used to deposit biodegradable elastomeric polymers and, optionally, ECM gels and / or other liquids, such as mammalian blood products, culture buffer solutions, media, and drug products. In its simplest sense, electrospinning involves the deposition of a liquid composition, such as polymer fibers, onto a target surface, triggered by an electric potential. The electrospinning process is well known in the field of tissue engineering and essentially operates as follows: Electrospinning allows for the creation of structures that resemble the scale and fibrous nature of natural extracellular matrix (ECM). ECM is composed of fibers, pores, and other surface features at the submicron and nanometer dimension scale. Such features directly affect cellular interactions with synthetic materials, such as migration and orientation. Electrospinning also allows for the production of oriented fibers, resulting in structures with inherent anisotropy or varying anisotropy in different parts of the structure. These aligned structures can affect cell growth, morphology, and ECM production. For example, Xu et al. found that poly(L-lactide-co-ε-caprolactone) fibers align smooth muscle cells (SMCs). See Non-Patent Document 8 (Xu CY, et al., Aligned biodegradable nanofibrous structure: a potential for blood vessel engineering, Biomaterials 2004 (25) 877-86). Lee et al. applied aligned non-biodegradable polyurethane to mechanical stimulation and showed that cells cultured on aligned scaffolds produced more ECM than those cultured on randomly organized scaffolds. See Non-Patent Document 4Z2 (Lee CH, et al., Nanofiber alignment and direction of mechanical strain affect the ECM production of human ACL fibroblast, Biomaterials 2005 (26) 1261-1270).

[0065] The electrospinning process involves placing a polymer-containing fluid (e.g., a polymer solution, polymer suspension, or polymer melt) into a reservoir equipped with a small orifice, such as a needle or pipette tip, and a metering pump. One electrode of a high-voltage source is placed in electrical communication with the polymer-containing fluid or the orifice, while the other electrode is also placed in electrical communication with a target (typically a collector screen or rotating mandrel). During electrospinning, the polymer-containing fluid is charged by application of a high voltage (e.g., about 3 to 15 kV) to the solution or orifice and then forced through the small orifice by a metering pump that provides a steady flow. The polymer-containing fluid at the orifice typically has a hemispherical shape due to surface tension, but application of a high voltage causes the hemispherical shape of the polymer-containing fluid at the orifice to expand to form a conical shape known as a Taylor cone. When a sufficiently high voltage is applied to the polymer-containing fluid and / or the orifice, the repulsive electrostatic forces of the charged polymer-containing fluid overcome surface tension, causing a jet of charged fluid to emerge from the tip of the Taylor cone and accelerate toward a target that is typically biased between -2 and -10 kV. Optionally, a focusing ring with an applied bias (e.g., 1 to 10 kV) may be used to guide the trajectory of the charged jet of polymer-containing fluid. As the jet of charged fluid moves toward the biased target, it experiences complex whipping and bending motions. If the fluid is a polymer solution or suspension, the solvent typically evaporates during flight, leaving polymer fibers on the biased target. If the fluid is a polymer melt, the molten polymer cools and solidifies during flight, collecting as polymer fibers on the biased target. As the polymer fibers accumulate on the biased target, a nonwoven porous mesh forms on the biased target. Under certain conditions, for example, solutions lacking sufficient viscosity and / or electrospun to a certain tolerance, fibers will not form, but a spray will form and discrete droplets will be deposited on the target instead of fibers: this is electrospray.

[0066] The properties of electrospun structures, such as elastomeric scaffolds, can be tailored by varying the electrospinning conditions. For example, if the biased target is relatively close to the orifice, the resulting electrospun mesh tends to contain nonuniformly thick fibers, such that some regions of the fibers have a "bead-like" appearance. However, as the biased target is moved further away from the orifice, the fibers of the nonwoven mesh tend to be more uniformly thick. Additionally, the biased target can be moved relative to the orifice. In a specific, non-limiting example, the biased target is moved back and forth in a regular, cyclical manner so that the fibers of the nonwoven mesh are substantially parallel to each other. In this case, the resulting nonwoven mesh can have a higher resistance to strain in the direction parallel to the fibers compared to the direction perpendicular to the fibers. In another non-limiting example, the biased target is moved randomly relative to the orifice, resulting in isotropic resistance to strain in the plane of the nonwoven mesh. The target can also be electrospun on a rotating mandrel. In this case, the properties of the nonwoven mesh can be varied by changing the rotation speed. The properties of the electrospun structure can also be varied by changing the magnitude of the voltage applied to the electrospinning system. In one non-limiting example, the electrospinning apparatus includes an orifice biased at 12 kV, a target biased at -7 kV, and a focusing ring biased at 3 kV. Furthermore, a useful orifice diameter is 0.047 inches (0.12 cm) (ID), and a useful target distance is approximately 23 cm. Other electrospinning conditions that can be varied include, for example, but are not limited to, the polymer solution feed rate, solution concentration, polymer molecular weight, injector-mandrel gap distance, and the relative trajectory of the injector and mandrel via the CNN control system.

[0067] More specifically, with respect to a rotating mandrel, anisotropic matrices, which are matrices or articles that are at least partially anisotropic, can be created by electrospinning the fibers onto the mandrel by biasing the deposition of fibers away from a random, isotropic orientation, resulting in a non-random bias of fiber orientation in a particular orientation, for example, with a circumferential bias (at least a portion of the deposited fibers are non-randomly oriented in the circumferential direction, resulting in anisotropy) or with a longitudinal bias (at least a portion of the deposited fibers are non-randomly oriented in the longitudinal direction, resulting in anisotropy). Fiber bias can be introduced into the electrodeposited article by relative movement of the target and polymer source (e.g., reservoir orifice, needle, pipette tip, etc.). For example, the mandrel target can be rotated at different speeds to generate different degrees of circumferential bias. The mandrel target and / or polymer source may be moved longitudinally, e.g., reciprocated, at different speeds (cycles) and amplitudes while electrospinning to generate varying degrees of longitudinal bias. For example, as shown in Figure 10B, in the illustrated system, a rotational speed of 1.5 m / s generates an anisotropy ratio (AR, a common metric of mechanical anisotropy defined as the ratio of the most flexible axis divided by the mechanical strain in the stiffer axis) that matches the natural anisotropy. The rotational speed of the mandrel and the longitudinal movement of the mandrel and / or polymer source can be easily controlled by a computer by one skilled in the art.

[0068] One measure of fiber orientation is called the fiber orientation index. The orientation index is defined in D'Amore et al., "Characterization of the complete fiber network topology of planar fibrous tissues and scaffolds," Biomaterials 31 (20), 5345-5354 (2010)." The orientation index is calculated by cos 2The orientation index (θ) can be obtained from the average over all fiber segments of (COS OI), where θ represents the angle between the fiber segment and the assumed direction of alignment. The anisotropic portion of the matrix described herein has an orientation index in the range of 0.5 to 0.8.

[0069] In certain embodiments, electrospinning is performed using two or more nozzles, each nozzle supplying a different polymer solution. The nozzles may be biased with different or the same bias to tailor the physical and chemical properties of the resulting nonwoven polymer mesh. Additionally, many different targets can be used. In addition to flat, plate-like targets, it is contemplated to use mandrels or rotating disks as targets.

[0070] When electrospinning is performed using a polymer suspension, the concentration of the polymer component in the suspension may also be varied to modify the physical properties of the elastomeric scaffold. For example, when the polymer component is present at a relatively low concentration, the resulting fibers of the electrospun nonwoven mesh will have a smaller diameter than when the polymer component is present at a relatively high concentration. Without wishing to be bound by theory, it is believed that a lower concentration solution will have a lower viscosity, resulting in faster flow through the orifice and producing thinner fibers. One skilled in the art can adjust the polymer concentration to obtain fibers with the desired properties. Useful ranges for the concentration of the polymer component include 1 wt% to 25 wt%, 4 wt% to 20 wt%, and 10 wt% to 15 wt%, including all increments therebetween.

[0071] In one non-limiting example, the construct is fabricated by electrospraying an ECM gel and / or other liquid while co-electrospinning a polymer suspension comprising synthetic and biological polymer components. In another non-limiting example, the polymer component of the construct is fabricated by electrospinning a polymer suspension comprising synthetic polymer components from one nozzle and a polymer suspension comprising biological polymer components from another nozzle. Non-limiting examples of useful ranges of high voltage to be applied to the polymer suspension are 0.5-30 kV, 5-25 kV, and 10-15 kV.

[0072] If present, the ECM gel component of the structure is sprayed (e.g., pressure sprayed) or electrosprayed simultaneously with the electrospinning of the polymer. Similarly, the liquid component of the wet electrospun layer is sprayed or electrosprayed simultaneously with the polymer component.

[0073] Prosthetic heart valves generally include two sections. The first support section is annular (forming a ring, but not necessarily defining a particular geometric shape such as a circle or cylinder) and serves as an attachment point for the heart valve, providing, for example, a structure for suturing and fixation, as well as an opening for blood flow through the prosthetic valve. The second section includes two or more flexible leaflets that are movable relative to the support section between an open configuration, in which the leaflets allow blood flow in a first direction through the opening, and a closed configuration, in which the leaflets restrict blood flow in a second direction opposite the first direction through the opening. The leaflets are coapted with adjacent leaflets at their edges immediately adjacent the support section to form a coaptation surface, and are uncoapted at a distal portion of the support section to allow blood to flow through the valve when the valve is open. When the valve is closed, the leaflets are concave, meaning that the concave portions extend toward the central axis of the opening in the support section and contact or coapt with adjacent leaflets to form a seal. Unless otherwise indicated, with respect to the mandrels and heart valve structures described herein, concave means curved or extending toward a rotational, longitudinal, or central axis, and convex means curved or extending outward, away from a rotational, longitudinal, or central axis.

[0074] 1A through 1F show different views of an example of a mandrel useful for fabricating a tricuspid valve prosthesis as described herein. Referring to FIG. 1A, a mandrel 10 useful for fabricating a tricuspid valve prosthesis by electrospinning is provided. In FIG. 1A, the mandrel 10 has non-conductive and conductive surfaces and, as shown in the exploded view of FIG. 1B, includes a non-conductive sheath 20, a conductive insert 30, a conductive rod 40 electrically connected to the insert 30, and a removable axial member 50 electrically connected to the rod 40. The axis of rotation is indicated by a dotted line in FIG. 1B. The elements of the mandrel 10 are arranged around the axis of rotation of the mandrel. FIG. 1B is an exploded view of the mandrel 10, showing the individual elements of the mandrel 10. The mandrel 10 has an axis of rotation, or longitudinal axis, and the radial direction, or radius, is perpendicular to any point on the axis of rotation. The radius of the mandrel 10 is measured perpendicular to the axis of rotation. The longitudinal direction is the direction parallel to the longitudinal axis. The circumference of the mandrel is the boundary of a circle centered on the longitudinal axis and perpendicular to the longitudinal axis, and the circumferential direction is the direction along the circumference.

[0075] 1B through 1F, the non-conductive sheath 20 includes a shaft portion 21, a cylindrical portion 22 having a radius, and longitudinal protrusions 23 extending longitudinally from the cylindrical portion 22. The longitudinal protrusions 23 taper in circumferential width from their attachment to the cylindrical portion 22 to their tips 24. The longitudinal protrusions 23 are inwardly biased such that their radii decrease from their attachment to the cylindrical portion 22 to their tips. The reduction in radius due to the inward bias does not exceed 10% of the radius of the cylindrical portion (i.e., the radius of the tips 24 is at least 90%, e.g., 91%, 92%, 93%, 94%, 95%, 96%, 97%, 98%, or 99% of the radius of the cylindrical portion). In an alternative embodiment, the longitudinal protrusions 23 are not inwardly biased. The radial profile of the longitudinal protrusion 23 as it extends from the cylindrical portion 22 to the tip 24 is shown as curved or arcuate, but may also be linear. A ridge 25 is shown that mates with a hole (not shown) in the cylindrical portion of the insert 30 to orient the insert 30 within the sheath 20. While the ridge 25 is shown, it can have any useful shape or configuration so long as it allows for orientation of the insert 30 within the sheath 20 and does not interfere with the function of the mandrel 10 as described herein. The distribution of mass about the axis of rotation of the mandrel 10 is preferably symmetrical or substantially symmetrical, i.e., balanced. The insert 30 is fabricated from a conductive material, such as a metal. The insert 30 fits within the sheath 20 as shown in FIG. 1A. The insert 30 includes a cylindrical portion 31, a first portion 32, and a second portion 33 that extends longitudinally from the first portion 32 opposite the cylindrical portion 31. The first portion 32 and the second portion 33 extend longitudinally from the cylindrical portion 31 so that they contact the inner surface of the longitudinal protrusion 23 of the sheath 20 and have a ridge 34 having a radius slightly smaller than the inner radius of the longitudinal protrusion 23, so that when the insert 30 is inserted into the sheath 20, the longitudinal protrusion 23 of the sheath 20 at least partially covers and insulates the ridge 34 of the insert 30 at the first portion 32 of the insert 30.The ridge 34 of the first portion 32 has a concave arcuate profile with a radius that decreases by no more than 10%, e.g., 1, 2, 3, 4, 5, 6, 7, 8, 9, or 10% or less of the radius of the cylindrical portion 31 of the insert 30. That is, the radii of the cylindrical portion 31 of the insert 30 and the ridge 34 include increments of between 90% and 100%, 95% and 100%, or 99% and 100% of the inner radius, formed on the surface of the cylindrical portion 22 of the sheath 20. In one example, the ridge 34 has a radius that is smaller than the radius of the cylindrical portion 31. The ridge 34 has a peak 35, and in the second portion 33 of the insert 30, the ridge 34 has a radius that decreases from the first portion 32 to the tip 36 of the insert 30, with a concave arcuate profile 37, such as a circular or parabolic profile. The "contour" of the longitudinal protrusions 23 and ridges 34 refers to the longitudinal change in radius of those features, for example, first from the cylindrical portion 31 to the second portion 33 of the insert 30, and then from the first portion 32 to the tip 35 of the ridges 34. Instead, the ridges 34 of the second portion 33 have a linear contour. The first and second portions 32 and 33 of the insert 30 also include curved recessed regions 38 between the ridges 34. The surfaces of adjacent recessed regions 38 on either side of the same ridge 34 of the insert 30 are substantially parallel in at least a portion of the second portion 33 of the insert 30. "Substantially parallel" means that the surfaces are not necessarily perfectly parallel, but will produce coapting leaflets that contact each other in the closed position when used as a target for electrodeposition of a polymer composition to fabricate a prosthetic heart valve. The concave regions and the depicted shapes of the illustrated leaflets are also referred to as "leaflet shapes," meaning that the geometry of the concave regions mimics the shape of the leaflets of a bicuspid or tricuspid heart valve.

[0076] Figures 2A through 2D show different views of one example of a mandrel useful in creating a tricuspid valve prosthesis as described herein. Referring to Figures 2A through 2C, a mandrel 110 is provided. Mandrel 110 includes a structure similar to that of the mandrel of Figure 1A, except that it essentially has an axis of rotation as shown for mandrel 10 of Figure 1B and is used to form a prosthetic bicuspid valve, including a non-conductive insulating sheath 120, a conductive insert 130, a conductive rod 140 electrically connected to insert 130, and a conductive removable axial piece 150 electrically connected to rod 140.

[0077] 2B-2D, non-conductive sheath 120 includes shaft portion 121, cylindrical portion 122 having a radius, and longitudinal protrusions 123 extending longitudinally from cylindrical portion 122. They taper to a narrower circumferential width from their attachment to cylindrical portion 122 to a tip 124, and unlike mandrel 10 of FIGS. 1A-1F, longitudinal protrusions 123 are not inwardly biased. However, in another embodiment (not shown, but essentially as shown in mandrel 10 of FIGS. 1A-1F), longitudinal protrusions 123 are inwardly biased such that their radius decreases from their attachment to cylindrical portion 122 to their tip. The reduction in radius due to the inward bias is 10% or less of the radius of the cylindrical portion (i.e., the radius of the tip 124 is at least 90%, e.g., 91%, 92%, 93%, 94%, 95%, 96%, 97%, 98%, or 99% of the radius of the cylindrical portion). When biased inward, the radial profile of the longitudinal protrusion 123 as it extends from the cylindrical portion 122 to the tip 124 can be curved, arcuate, or straight. Not shown are protrusions that fit into holes in the cylindrical portion of the insert 130 to align the sheath 120 with the insert 130, as described with respect to FIG. 1C . The distribution of mass about the axis of rotation of the mandrel 110 is preferably symmetrical or substantially symmetrical, i.e., balanced. The insert 130 is fabricated from a conductive material, such as a metal. 2A, the insert 130 fits within the sheath 120. The insert 130 includes a cylindrical portion 131 and a first portion 132 extending longitudinally from the cylindrical portion 131. The first portion 132 includes ridges 134 extending longitudinally from the cylindrical portion 131 and having a radius slightly smaller than the inner radius of the longitudinal protrusions 123 such that they contact the inner surface of the longitudinal protrusions 123 of the sheath 120. As a result, when the insert 130 is inserted into the sheath 120, the longitudinal protrusions 123 of the sheath 120 at least partially cover and insulate the ridges 134 of the insert 130 at the first portion 132 of the insert 130.1A , the ridges 134 of the first portion 132 are depicted as straight lines having the same radius as the cylindrical portion 131, but, similar to the mandrel 10 of FIG. 1A , they may instead be biased inward and have a concave arcuate profile with a decreasing radius that does not exceed 10%, e.g., 1, 2, 3, 4, 5, 6, 7, 8, 9, or 10% of the radius of the cylindrical portion 131 of the insert 130. That is, the radii of the cylindrical portion 131 of the insert 130 and the ridges 134 include increments of between 90% and 100%, 95% and 100%, or 99% and 100% of the inner radius, formed on the surface of the cylindrical portion 122 of the sheath 120. In one example, the ridges 134 have a radius that is smaller than the radius of the cylindrical portion 131. As discussed above, the "contour" of the longitudinal protrusions 123 and the ridges 134 refers to the longitudinal change in radius of these features, for example, as the ridges 134 extend longitudinally from the cylindrical portion 131. The first portion 132 of the insert 130 also includes curved recessed regions 138 between the ridges 134. The surfaces of adjacent recessed regions 138 on either side of the same ridge 134 of the insert 130 are generally parallel distally to the cylindrical portion such that, when used as a target for electrodeposition of a polymer composition to fabricate a prosthetic heart valve, the target produces coapted leaflets that contact each other in the closed position. As discussed above, the depicted shapes of the recessed regions and the depicted leaflets are also referred to as "leaflet shapes," and the geometry of the recessed regions mimics the shape of the leaflets of a bicuspid or tricuspid heart valve.

[0078] In one embodiment, the first portion of the electrodeposition target has substantially the same radius as the cylindrical portion. In one embodiment, the electrodeposition target includes ridges symmetrically spaced about the axis of rotation.

[0079] Figures 2E-2G show alternative embodiments of the mandrels described herein. Figure 2E shows a mandrel 111, essentially as shown in Figure 1A, with a sheath 120 and a sewing ring 125 extending radially around the cylindrical portion (specific reference numbers omitted for clarity). A radially extending sewing ring can be included in any of the mandrel designs described herein to provide additional material for the resulting polymer matrix valve structure created on the mandrel. While the radially extending sewing ring is shown extending perpendicular to the longitudinal axis, it need not extend perpendicular to the longitudinal axis. Figure 2F shows a variation of the mandrel of Figure 2A for creating a bicuspid valve. The mandrel 112 includes a non-conductive sheath 120 and an insert 130 essentially as shown in FIG. 2A (specific reference numbers omitted for clarity), except that a first portion of the insert 130 is radially curved, and the sheath 120 includes a longitudinal protrusion 123 extending around the first portion, including over the distal edge 123′ of the insert. FIG. 2G shows substantially the same structure as FIG. 1A (specific reference numbers omitted for clarity), but with smaller dimensions, including the non-conductive sheath 120 and the longitudinal protrusion 123, including the second portion's raised portion 123′ and covering the entire raised portion of the insert. FIGS. 2H and 21 show the bottom and top views of a bicuspid valve fabricated on the mandrel according to FIG. 2F. FIG. 2J is a photograph of a tricuspid valve fabricated on the mandrel of FIG. 2E, including a sewing ring.

[0080] The mandrel structures shown in Figures 1A-1F and 2A-2G are merely exemplary. The mandrel can have any useful shape and configuration, for example, the mandrel may be fabricated from a single conductive material with a non-conductive coating deposited on a portion thereof to produce the same or a similar pattern to the pattern produced by the assembly shown in Figures 1A-1F and 2A-2G.

[0081] Prosthetic heart valves are fabricated by electrodeposition of polymer fibers onto a mandrel, as described herein. The resulting structure is removed from the mandrel and trimmed as necessary to ensure that openings are formed between the leaflets and that the commissures are the desired length to produce properly coapted leaflets. The matrices produced by electrospinning preferentially have thicknesses between 100 μm and 400 μm. The diameter of the cylindrical portion is the diameter of the native heart valve and can be varied by changing the radius of the cylindrical portion of the mandrel, as will be understood by those skilled in the art. Similarly, the AR can be varied to cover the full range of measured or estimated AR for porcine and human valves, AR = 1-3, corresponding to stress and strain values of 0-500 kPa and 0-40% strain, respectively. Finally, the range of bending modulus of natural tissue (1000-20000) kPa can also be covered by varying the speed for the raster (linear motion of the mandrel).

[0082] In another aspect, a prosthetic tricuspid valve 210 is shown in Figure 3. The valve 210 includes a support portion 220 defining a longitudinal axis 222, an opening 225 extending therethrough, and three concave leaflets 230 extending longitudinally from a distal end 234 of the support portion 220, each leaflet 230 including a concave abdomen or central region 235 and a coaptation surface 236 joining adjacent leaflets 230. The fiber matrix in the central region 235 and the coaptation surface 236 is anisotropic, with different fiber orientations in the central region 235 and the coaptation surface 236, the fiber orientation being more circumferential in the central region 235 than in the coaptation surface 236.

[0083] In another aspect, a bicuspid valve 340 is provided, as shown in FIGS. 4A through 4E. The valve 340 is formed from a fibrous matrix and includes a support portion 341 defining a longitudinal axis 342 and an opening, and two concave leaflets 343 extending longitudinally from the support portion 341, each having a central region 345 and a coaptation surface 346 where the leaflets coapt. FIGS. 4B and 4D provide top views of the bicuspid valve 340 along the longitudinal axis in the closed and open positions, respectively. FIGS. 4C and 4E show the valve 340 along the X in the corresponding FIGS. 4B and 4C. FIGS. 4B and 4C show the valve 340 in a closed configuration, while FIGS. 4D and 4E show the valve in an open configuration with blood flow in the direction of the arrows. It should be noted that the leaflets of a bicuspid valve have a radially curved profile, with one leaflet being larger than the other.

[0084] In any aspect of the valve structure described herein, the identification of a cylindrical portion is merely illustrative and exemplary of one possible geometry of a potential support structure (e.g., support portion) for the valve leaflets. Indeed, in alternative embodiments, the support structure may take any shape useful for supporting the function of the valve leaflets and securing the valve in place, e.g., functioning at least partially as a sewing ring or providing an attached sewing ring for suturing the structure in place during implantation, and / or for attachment to additional support or alignment structures, but a cylindrical shape may be considered the simplest and most appropriate shape for purposes herein. The leaflet portions of the valves shown in Figures 3 and 4A-4E are flexible and concave, and when in a closed configuration or position, they contact (coapt) the distally adjacent leaflets of the cylindrical or support portion of the prosthetic valve to prevent backflow of blood. When in the open configuration, the leaflets dilate openings in the (e.g., cylindrical) support structure, allowing blood flow through the device. It is worth noting that for any valve structure and corresponding mandrel target shape, the valve leaflets do not need to be symmetrical in size, as with the native mitral and tricuspid valves. In any case, the mandrel target shape can be referred to as a leaflet shape, such as a mitral, tricuspid, aortic, or pulmonary valve leaflet shape, in reference to the concave leaflet shape of the insert. Alternatively, it can be a pathological mitral, tricuspid, aortic, or pulmonary valve leaflet shape, referring to the natural or damaged / pathological shape of the leaflets or cusps of a valve in an organism, such as a human or mammal. Pathological shapes are useful in research to study defects in valve structure. Table 1 provides exemplary diameters for heart valves and, therefore, mandrel target diameters for electrospinning. The values in Table 1 are appropriate for humans, pigs, and other animals weighing over 40 kg. For animals weighing less than 40 kg, such as minipigs, sheep, and goats, the diameter should be, for example, 30% smaller. Even in smaller animals such as rabbits and rats, the value should be, for example, 90% smaller.

[0085] [Table 1]

[0086] Electrodeposition of polymer fibers onto the mandrel structures described herein can impart anisotropy to all or portions of a prosthetic heart valve structure, but the primary fiber alignment direction can be varied within the same engineered valve. In contrast, conventional deposition targets, such as rotating mandrels or flat surfaces, will produce materials with no alignment or a consistent alignment direction within the same structure. This aspect is particularly important because the leaflets of native valves are characterized by a consistent level of fiber alignment (AR) within the leaflet, but also by varying primary alignment directions within the same leaflet. For example, the abdominal region of a native valve leaflet is circumferentially oriented, while the coaptation region is approximately longitudinally oriented. Precise fiber deposition on a concave surface allows for the rotation of the primary direction of scaffold fiber alignment from the abdominal region to the coaptation region to recreate the same effect. Thus, the mandrels and methods using the mandrels described herein provide precise control of device thickness, device dimensions, device shape, anisotropy, and bending modulus, allowing for curved fibers as opposed to linear or isotropic alignment.

[0087] In another aspect, the fibers are circumferentially aligned and attached at least partially to at least a portion of the heart valve, e.g., at least partially to the concave or abdomen portions of the leaflets. In yet another aspect, the fibers are aligned in a non-directional or isotropic pattern to at least a portion of the valve, e.g., in the portion between the abdomen and coaptation surfaces of the leaflets.

[0088] Prosthetic heart valves are fabricated by electrodeposition of one or more biodegradable, biocompatible polymer compositions. Examples of useful polymer compositions include poly(ester urethane) urea (PEUU), poly(ether ester urethane) urea (PEEUU), poly(ester carbonate) urethane urea (PECUU), poly(carbonate) urethane urea (PCUU), polymers derived from alpha-hydroxy acids, polylactide, poly(lactide-co-glycolide), poly(L-lactide-co-caprolactone), polyglycolic acid, poly(dl-lactide-co-glycolide), poly(l-lactide-co-dl-lactide), polymers comprising lactone monomers, polycaprolactone, and polymers comprising carbonate linkages. The polymers include one or more of: polymers having ester bonds, polycarbonate, polyglyconate, poly(trimethylene carbonate), poly(glycolide-co-trimethylene carbonate), poly(glycolide-co-trimethylene carbonate-co-dioxanone), polyurethane, polycarbonate urethane, polyester urethane, polymers with ester bonds, polyalkanoates, polyhydroxybutyrate, polyhydroxyvalerate, polydioxanone, polygalactin, natural polymers, chitosan, collagen, elastin, alginate, cellulose, hyaluronic acid, and gelatin.

[0089] In yet another aspect, there is provided a method of fabricating a prosthetic heart valve structure, together with a product of the method comprising electrodepositing a biodegradable, biocompatible polymer composition onto an electrodeposition target, e.g., a mandrel, as described herein. In another aspect of the method, the polymer composition comprises a synthetic polymer. In another aspect of the method, the synthetic polymer is selected from the group consisting of poly(ester urethane) urea (PEUU), poly(ether ester urethane) urea (PEEUU), poly(ester carbonate) urethane urea (PECUU), poly(carbonate) urethane urea (PCUU), alpha-hydroxy acid derived polymers, polylactide, poly(lactide-co-glycolide), poly(L-lactide-co-caprolactone), polyglycolic acid, poly(dl-lactide-co-glycolide), poly(L-lactide-co-caprolactone), poly(glycolic acid), poly(dl-lactide-co-glycolide ... glycolide), poly(l-lactide-co-dl-lactide), polymers with lactone monomers, polycaprolactone, polymers with carbonate bonds, polycarbonate, polyglyconate, poly(trimethylene carbonate), poly(glycolide-co-trimethylene carbonate), poly(glycolide-co-trimethylene carbonate-co-dioxanone), polyurethanes, polycarbonate urethanes, polyester urethanes, polymers with ester bonds The polymer is selected from the group consisting of one or more of polyalkanoates, polyhydroxybutyrates, polyhydroxyvalerates, polydioxanones, polygalactins, natural polymers, chitosan, collagen, elastin, alginates, cellulose, hyaluronic acid, and gelatin. In another embodiment of the method, the synthetic polymer is PEUU, PEEUU, PECUU, or PCUU. In another embodiment of the method, the anisotropy of the electrodeposited polymer composition is oriented in at least a portion of the structure, thereby creating an anisotropic portion of the structure. In another embodiment of the method, for example, more than 50% of the fibers of the electrodeposited polymer are oriented circumferentially within the concave central portion or abdomen of at least two leaflet portions, and / or for example, more than 50% of the fibers of the electrodeposited polymer are oriented longitudinally at or immediately adjacent to the coaptation interface between at least two concave leaflet portions.In another aspect of the method, the shape and size of the electrodeposition target mimic natural anatomy, shape, and dimensions to replicate healthy or pathological human or animal anatomy (e.g., valve anatomy). The electrodeposited valve structure is removed from the electrodeposition target, and optionally, coapted leaflet portions are separated, leaving a coaptation surface that coapts at least a portion of the leaflet portions. The valve structure may be rinsed or hydrated in a suitable solution, such as water, saline, or PBS. The valve structure may optionally be seeded with cells, and optionally, cells may be cultured on the valve structure so that the cells cover and / or infiltrate at least a portion of the valve structure. In another aspect, the method further comprises electrodepositing, spraying, or otherwise adding or incorporating a second polymer, which may be an ECM gel, a drug, water, saline, PBS, cell culture medium, cells, a biological agent, a salt, a buffer, a cytokine, a growth factor, or a combination thereof, onto the electrodeposition target.

[0090] In use, the valve prostheses described herein are implanted in a patient at the site of a native valve, e.g., a valve annulus. In the case of a heart valve, the device is sewn into place at the heart valve annulus and, in some cases, connected, for example, via cusps on the prosthetic valve, to the papillary muscles of the atrio-ventricular valves or the coaptation surfaces of the ventriculo-arterial valves. The prosthesis is sewn and connected to a frame, such as a stent or similar framework as is commonly known in the art, and then placed and implanted in the native valve annulus. Suitable frames, fabricated from shape-memory metals such as, for example, but not limited to, nitinol or polymers, are commonly known, and the appropriate frame configuration can be determined.

[0091] [Example] [Example 1] A three-leaflet version of the mandrel design described herein is essentially shown in Figures 1A-1F, and its use in preparing a tricuspid valve is illustrated in Figures 5(A)-5(D). Figure 5(A) is a photographic image of the dual-component mandrel before polymer fiber deposition. As shown in Figure 5(B), the polymer, in this case PEUU, is electrodeposited for 3 hours around the conductive portion of the target, selectively depositing the polymer fibers on the conductive target. The processing conditions for this fabrication were: polymer voltage 11 kV, second stream (PBS) voltage 8 kV, mandrel voltage -5 kV, polymer flow rate 1.5 ml / hr, second stream flow rate 1.2 ml / hr, polymer-mandrel gap 15.5 cm, second stream-mandrel gap 4.5 cm, PEUU solvent weight / volume 12%, humidity <40%, raster speed 0 cm / s, and mandrel speed 372 rpm. As shown in Figure 5(D), the mandrel is placed in a chuck, rotated, and moved longitudinally. In a typical embodiment, the mandrel is rotated and the electrodeposition nozzle is not rotated around the mandrel; the spatial position and relative orientation of the polymer nozzle and mandrel can be controlled manually or, more typically, by computer using a standard robot and stage. The resulting tricuspid valve is shown in Figure 5(C). The dual-component design (shield + target) aims to focus fiber deposition only in the concave zone. The part shape and dimensions can be varied based on the patient's anatomy. Similar concepts can be applied to non-biomedical applications requiring fiber deposition in concave areas.

[0092] [Example 2] A bicuspid (bicuspid) version of the mandrel design described herein, as shown essentially in Figures 2A-2D and 2F, and its use in creating a bicuspid valve with fiber deposition essentially as described in Example 1, and a prosthetic bicuspid valve are shown in Figures 4A-4E. Polymer fibers, such as PEUU, are deposited essentially as described in Example 1. Similarly, control of anisotropy and modulus is achieved by varying the mandrel speed and raster speed, respectively.

[0093] [Example 3] Example 3 provides a qualitative test of leaflet coaptation at rest for a tricuspid valve fabricated using a dual-component mandrel, as described in Example 1. The tricuspid valve was removed from the mandrel with a three-leaflet design ( FIG. 5A ), and the valve was immersed in PBS. Qualitative examination of the valve construct fabricated as described in Example 1 demonstrated leaflet coaptation at rest when the valve construct was immersed in liquid ( FIG. 5B ). In contrast, the leaflets of conventional valves obtained by electrospinning a regular shape are flat or cylindrical. In the specific case of a flat or cylindrical mandrel, the lack of leaflet indentation and the need to structurally connect the different leaflets do not allow for proper coaptation at rest.

[0094] [Example 4] Example 4 provides further measurements and analysis of the three-dimensional shape of a tricuspid valve made using the dual-component mandrel described in Example 1. Photographs of a tricuspid valve made as described in Example 1, shown in Figures 6A and 6B, demonstrate how the valves and methods described herein for making these valves produce tricuspid valves with a physiological curvature that is dictated by the shape of the mandrel used to make them.

[0095] [Example 5] Example 5 provides data comparing the thickness of leaflet valves fabricated using the dual-component mandrel described herein with a native porcine tricuspid valve. This example also demonstrates how thickness is linearly affected by deposition time. Material processing variables were the same as those used in Example 1. A specific thickness of interest can be achieved based on deposition time. Figure 7A shows a graph of thickness versus deposition time for engineered heart valves and native porcine tricuspid valves (n = 3 (3 hours), n = 4 (4 hours), n = 4 (native porcine tricuspid valve)). The engineered heart valves were fabricated by independent fabrication, where a predetermined fabrication time can be set to achieve a prosthetic leaflet thickness comparable to that of the native leaflet, demonstrating that a specific thickness of interest can be achieved based on deposition time. Figures 7B and 7C show the thickness distribution over the leaflet region of the native porcine tricuspid valve and the engineered tricuspid valve, respectively, 3 hours after fabrication. The thickness map comparison (natural vs. prosthetic) shows comparable leaflet thickness values across the entire surface.

[0096] [Example 6] Example 6 provides data related to the leaflet mechanics of valves fabricated using the dual-component mandrel described herein. To demonstrate control over the in-plane and out-of-plane mechanics of the valve leaflets, the mandrel design described herein was tested for nine different conditions covering the operating range of interest for valve applications (mandrel tangential velocity: 0.3-3 m / s, raster (longitudinal) linear velocity: 0-2.5 cm / s) (see Figure 8). A PEUU was used for these experiments. The results shown in Figure 8 (A-I) demonstrate that mandrel velocity directly controls mechanical anisotropy (increasing differences in response to mandrel longitudinal direction for VI, V2, and V3). In contrast, raster velocity did not significantly affect the level of anisotropy (no significant difference between circumferential and longitudinal directions for R0, R1, and R2). Representative images of valve constructs demonstrate the feasibility of employing the presented mandrel design for different fabrication configurations.

[0097] [Example 7] Example 7 provides test results on the mechanical response of leaflet valves fabricated using the dual-component mandrel described herein compared to a native porcine tricuspid valve when mandrel tangential and raster velocities were varied. The processing conditions for these fabrications were: polymer voltage 11 kV, secondary stream (PBS) voltage 8 kV, mandrel voltage -5 kV, polymer flow rate 1.5 ml / hr, secondary flow rate 1.2 ml / hr, polymer-mandrel gap 15.5 cm, secondary stream-mandrel gap 4.5 cm, PEUU solvent wt / v %; humidity less than 40%; raster velocities 0, 0.16, and 2.5 cm / s, while mandrel tangential velocities were 0.3, 1.5, and 3 m / s. Figure 9A is a graph showing the in-plane mechanical response of engineered valves tested in biaxial tensile tests in isostress mode for nine configurations obtained by varying the mandrel tangential velocity (ω, controlling for anisotropy) and raster velocity (υ, controlling for bending modulus), n = 3 mean ± standard error (n = 3 mean ± standard error). Figure 9B is a graph showing the anisotropy ratio (AR), defined when the mechanical strain ratio between the longitudinal and circumferential directions is used as a metric of anisotropy. AR vs. rotational velocity summarizes the results in Figure 9A, showing the AR for the valve configurations in Figure 9A as well as values for the native porcine tricuspid valve. The appropriate mandrel velocity required to fabricate a tricuspid valve with native mechanics (approximately 1.5 m / s) was identified by linear interpolation of the AR at 0.3, 1.5, and 3 m / s.Figure 9C shows a graph of the biaxial response (valve radial = mandrel longitudinal; valve leaflet circumferential = mandrel circumferential) of an engineered tricuspid valve replicating the AR of a native porcine tricuspid valve, using data generated with a customized biological tissue biaxial testing apparatus and protocol previously described in "Biaxial Mechanical Evaluation of Planar Biological Materials" by M. Sacks in Journal of Elasticity and the Physical Science of Solids, July 2000, Vol. 61, Issues 1-3, pp. 199-246. Testing was performed at room temperature under quasi-static conditions, employing an equiaxial biaxial protocol with a peak of 400 kPa after the specimen was continuously immersed in PBS during testing. A preconditioning free-float condition was utilized as the reference configuration.

[0098] [Example 8] Example 8 provides further test results on the mechanical response of leaflet valves fabricated using the dual-component mandrel described herein compared to native porcine tricuspid valves when mandrel tangential and raster speeds were varied. Specifically, the relationship between raster speed and elastic modulus (out-of-plane behavior) was investigated. Material processing variables were the same as those utilized in Example 7. Figure 10 is a graph showing leaflet bending modulus for different values of mandrel tangential speed (0.3, 1.5, 3 m / s) and raster speed (0, 0.25, 2.5 cm / s), n = 3 mean ± standard error. Comparison with porcine tricuspid valve values (n = 5 mean ± standard error) demonstrates the ability of the mandrel design to replicate the bending modulus of the native valve. While the elastic modulus is largely insensitive to changes in mandrel speed (Figure 10), the raster speed determines the bending stiffness, demonstrating the ability of this new design to achieve physiologically relevant values of bending stiffness for stentless composite geometries. Data were generated using a custom-made biological tissue bending device previously developed and validated in

[12] (Mirnajafi A et al., "The flexural rigidity of the aortic valve leaflet in the commissural region," Journal of Biomechanics, Volume 39, Issue 16, 2006, Pages 2966-2973). Testing was performed at room temperature under quasi-static conditions, with the specimen continuously immersed in PBS during testing. The curvature range was ±0.12°, and the Eulero-Bernoulli theory was employed for moment-curvature characterization. The combined biaxial testing and bending stiffness evaluation demonstrated the method / prototype's ability to isolate and control in-plane and out-of-plane engineered valve mechanisms.

[0099] [Example 9] The microstructure of the engineered valve leaflets and a comparison with a native porcine tricuspid valve are shown in Figures 11A and 11B. Material processing parameters were the same as those used in Example 1. Figure 11A is a multiphoton microscopy image of a native porcine tricuspid valve microstructure showing the collagen fiber network identified by acquiring collagen second-harmonic generation. A 500 μm × 500 μm × 100 μm volume was analyzed in five different valves and five different locations within the valve leaflets, including the coaptation plane (upper left TL and upper right TR), the ventral region (center C), and the transition zone between two locations (the left ML of the middle leaflet and the right MR of the middle leaflet). The shape of the collagen fibers was identified by digital image analysis as utilized in "Fiber Microarchitecture in the Longitudinal Radial and Circumferential Radial Planes of the Ascending Thoracic Aortic Aneurysm Media" by A. Tsamis et al., Journal of biomechanics 46(16), 2787-2794. The analytical method quantifies the predominant angle of fiber orientation by the mean of the fiber angle distribution θ and the level of fiber alignment with the orientation index (OI). This widely adopted metric (see, for example, A. D'Amore et al., "Characterization of the complete fiber network topology of planar fibrous tissues and scaffolds," Biomaterials 31(20), 5345-5354) is 0.5 for a set of randomly oriented fibers and equals 1 for a set of parallel fibers. Values for native tissue are reported in Table 2. The predominant direction of alignment (n = 5 independent leaflets) is also indicated by white arrows. The rotation of the fibers is from the coaptation plane to the abdominal region. Figure 11B is a multiphoton microscopy image of an engineered valve leaflet using the same imaging and digital analysis techniques as Figure 11A. Figures 11A and 11B use the same imaging and digital analysis techniques. Not only did the leaflets report physiological levels of fiber alignment (Table 2: 0I = 0.57–0.62), but the primary angles of alignment also tended to be comparable to those of native valve leaflets.This result cannot be achieved with conventional electrospinning electrodes (eg, flat mats or rotating drums) where the direction of primary alignment remains the same within the same structure.

[0100] [Table 2]

[0101] [Example 10] This example shows the results of in vitro testing of leaflet coaptation and suture retention of engineered valves (see Figures 12(A) through 12(F)). Material processing parameters were the same as those used in Example 1. Electrospun valves were implanted ex vivo into the tricuspid position of native porcine hearts, the right ventricle was slowly filled with saline, and pressure values were monitored with a Millar pressure transducer (mikro-Cath®, Millar, Houston, TX). Simultaneously, images of the coapted leaflets were acquired for healthy native porcine valves (A, B, C) and engineered valves (D, E, F), with Δρ > 30 mmHg, indicating adequate leaflet coaptation and adequate suture retention.

[0102] [Example 11] Figure 13 provides a graphical representation of valve function during in vitro testing of (A) the dynamics of a state-of-the-art commercially available prosthetic valve (n = 5, Carpentier-Edwards® Duraflex®) and (B) the dynamics of an engineered valve (n = 3). Pulsatile flow across the valve was generated by the Thoratec percutaneous VAD system, a commercially available ventricular assist device operating at a constant frequency of 70 beats per minute. Two cameras continuously detected valve motion, and pressure and flow were recorded by sensors. Digital image processing was performed to detect orifice area using a dedicated Matlab code (Mathworks®, Natick, MA). Figures 13(A) and 13(B) show the detected orifice area (white) for the Carpentier-Edwards and engineered valves, respectively, during systole. Figure 13(C) shows a bar graph representation of the bending deformation index (BDI) for the Carpentier-Edwards bioprosthetic valve and the engineered valves in Figures 13(A) and 13(B). BDI, a widely adopted metric for bending stiffness (see Non-Patent Document 14, "In vitro hydrodynamics, cusp-bending deformation, and root distensibility for different types of aortic valve-sparing operations: Remodeling, sinus prosthesis, and reimplantation" by A. Erasmi et al. in The Journal of Thoracic and Cardiovascular Surgery, Volume 130, Issue 4, October 2005, pp. 1044-1049), was calculated at mid-diastole. Figure 13(D) is a bar graph representation of the geometric orifice area (GEO) comparison of the Carpentier-Edwards bioprosthetic valve and the engineered valves of Figures 13(A) and 13(B), where GEO was calculated from the images processed at peak systole shown in (A) and (B).Figure 13(E) is a bar graph representation of the peak systolic pressure for the Carpentier-Edwards bioprosthetic valve and the engineered valves of Figures 13(A) and 13(B). Figure 13(F) is a bar graph representation of the mean systolic pressure for the Carpentier-Edwards bioprosthetic valve and the engineered valves of Figures 13(A) and 13(B). Figure 13(G) is a bar graph representation of the mean pressure drop for the Carpentier-Edwards bioprosthetic valve and the engineered valves of Figures 13(A) and 13(B) during a complete cycle including systole and diastole. Figure 13(H) is a bar graph representation of the mean flow rate for the Carpentier-Edwards bioprosthetic valve and the engineered valves of Figures 13(A) and 13(B) during a complete cycle including systole and diastole. None of the comparisons presented in Figures 13(C) through (H) showed statistically significant differences, indicating that the engineered valves have dynamic functional characteristics comparable to commercially available bioprosthetic valves. This superior dynamic performance was determined by the ability to control the valve mechanics and anatomy discussed herein.

[0103] [Example 12] Figure 14 shows a photographic representation of (A) a flow duplicator in which a 40% glycerol solution was used to mimic blood viscosity for evaluation of valve function under physiological flow conditions, (B) an engineered valve holder with a white arrow pointing to the engineered valve, and (C) a schematic diagram of the flow duplicator. Here, 1) is a desktop computer; 2) to 4) are the pressure and flow signal acquisition system, 5) is the preload sensor, 6) is the flow meter, 7) is the capacitor, 8) is the pressure sensor, 9) is the flanged valve holder, 10) is the side camera, 11) is the Thoratec percutaneous VAD system (Thoratec, Pleasanton, CA), 12) is the front camera, and 13) is the VAD controller.

[0104] The invention further includes the subject matter of the following clauses.

[0105] (Article 1) 1. An electrodeposition target having a surface with a pattern of conductive and non-conductive surface portions, the target being mounted on a mandrel having an axis of rotation and comprising a spindle electrically connected to the conductive portions of the target.

[0106] (Article 2) 10. The electrodeposition target of claim 1, wherein the mandrel comprises a non-conductive sheath insulating at least a portion of the conductive portion.

[0107] (Article 3) The electrodeposition target described in clause 2, characterized in that the target comprises a support portion arranged around the rotational axis of the mandrel, a conductive insert having a plurality of ridges extending longitudinally from the support portion and a plurality of recessed portions between the ridges, and a non-conductive layer on at least a portion of the support portion and at least a portion of the ridges.

[0108] (Article 4) 4. The electrodeposition target of clause 3, wherein the insert has two concave portions, the two concave portions being symmetrical or asymmetrical about the rotation axis of the mandrel.

[0109] (Article 5) 5. The electrodeposition target of claim 4, wherein a cross section of the target perpendicular to the rotation axis at the concave portion is "U" shaped.

[0110] (Article 6) 6. The electrodeposition target of any one of clauses 3 to 5, wherein the non-conductive layer is continuous around the plurality of recessed portions.

[0111] (Article 7) 7. An electrodeposited target according to any one of clauses 3 to 6, characterized in that the concave portion has the shape of a valve cusp, such as the shape and size of the cusps (leaflets) of a normal or pathological human or animal mitral, tricuspid, aortic or pulmonary valve, for example, a normal or pathological valve cusp.

[0112] (Article 8) The target is: a. a support portion having a non-conductive surface and a radius disposed about an axis of rotation of the mandrel; b. a leaflet portion attached to the support portion and extending longitudinally from the support portion along the axis of rotation, the leaflet portion comprising three concave conductive portions defined by three conductive ridges extending radially from the axis of rotation and having peaks, the leaflet portion comprising a first portion adjacent to and extending from the support portion and an optional second portion extending longitudinally from the first portion opposite the support portion, the radius of the ridges of the first portion decreasing by less than 10% from the support portion to the second portion, and, if present, the radius of the ridges of the second portion decreasing by at least 50%, and optionally at least 60%, 70%, 75%, 80%, 90%, 95%, or 99% in the second portion, the ridges further comprising a non-conductive layer extending from the support portion over at least a portion of the peaks of the ridges of the first portion. 8. The electrodeposition target according to any one of clauses 1 to 7, characterized in that

[0113] (Article 9) The target is: a. a support portion having a non-conductive surface and a radius disposed about an axis of rotation of the mandrel; b. a leaflet portion attached to the support portion and extending longitudinally from the support portion along the axis of rotation, the leaflet portion comprising two concave conductive portions defined by two conductive ridges extending radially from the axis of rotation and having peaks, the leaflet portion comprising a first portion adjacent to and extending from the support portion, the ridges of the first portion decreasing in radius by less than 10% from the support portion to a distal end of the first portion, the ridges further comprising a non-conductive layer extending from the support portion over at least a portion of the peaks of the ridges of the first portion. 9. The electrodeposition target according to any one of clauses 1 to 8, characterized in that

[0114] (Article 10) 10. The electrodeposition target of either clause 8 or 9, wherein the non-conductive layer is continuous around the periphery of the recessed conductive portion.

[0115] (Article 11) 10. The electrodeposition target of any of clauses 8 or 9, wherein the circumferential width of the ridges decreases with longitudinal distance from the cylindrical portion.

[0116] (Article 12) 10. The electrodeposition target of any of clauses 8 or 9, wherein the first portion has substantially the same radius as the cylindrical portion.

[0117] (Article 13) an insulating sheath; and a removable conductive insert having a cylindrical portion and including conductive ridges, peaks of the ridges, and a recessed conductive portion defined by the ridges; 13. The electrodeposition target of any one of clauses 1 to 12, wherein the insulating sheath covers and insulates at least a portion of the non-conductive portion of the insert and at least a portion of the raised peaks of the raised portions of the conductive insert.

[0118] (Article 14) 14. The electrodeposition target of any one of clauses 1 to 13, wherein the support portion is cylindrical.

[0119] (Article 15) 15. The electrodeposition target of any one of clauses 1 to 14, further comprising a flange extending radially around at least a portion of the support.

[0120] (Article 16) 1. A prosthetic valve formed from a matrix of polymer fibers, comprising: a. a tubular (not meant to be cylindrical, but can have a circular, elliptical, or any closed shape in cross section perpendicular to the longitudinal axis) support portion defining an opening and having a longitudinal axis; and b. at least two concave leaflets extending longitudinally from said support portion; Each leaflet includes a leaflet having a concave central portion, a peripheral portion about the concave central portion, a proximal end connected to the support portion, and a distal end longitudinally distal to the support portion. An artificial valve, characterized in that the peripheral portions of adjacent valve leaflets are adjacently coapted at the support portions to form a coaptation surface between the adjacent valve leaflets.

[0121] (Article 17) 17. The prosthetic valve of clause 16, wherein the leaflets have a bending modulus in the range of 500 kPa to 500,000 kPa, a mechanical strain in the range of 0 to 100, and / or a stress in the range of 0 to 5,000 kPa.

[0122] (Article 18) 17. A valve according to clause 16, characterized in that the matrix comprises an anisotropic portion with an orientation index ranging from 0.5 to 0.8.

[0123] (Article 19) 19. A valve according to any one of clauses 16 to 18, characterized in that the matrix at the interface and / or peripheral portion is anisotropic with the matrix fibres being biased in the longitudinal direction and / or the matrix in the concave central portion is anisotropic with the matrix fibres being biased in the circumferential direction.

[0124] (Article 20) 19. A valve according to any one of claims 16 to 18, characterized in that it has two cusps which are optionally asymmetric.

[0125] (Article 21) 19. A valve according to any one of claims 16 to 18, characterized in that it has three cusps which are optionally asymmetric.

[0126] (Article 22) 19. A valve according to any one of clauses 16 to 18, characterized in that it has two cusps formed as mitral valve cusps or three cusps formed as tricuspid valve cusps.

[0127] (Article 23) 23. A valve according to any one of clauses 16 to 22, characterized in that the matrix is formed by electrospinning.

[0128] (Article 24) The matrix may be poly(ester urethane) urea (PEUU), poly(ether ester urethane) urea (PEEUU), poly(ester carbonate) urethane urea (PECUU), poly(carbonate) urethane urea (PCUU), polymers derived from alpha-hydroxy acids, polylactide, poly(lactide-co-glycolide), poly(L-lactide-co-caprolactone), polyglycolic acid, poly(dl-lactide-co-glycolide), poly(l-lactide-co-dl-lactide), polymers comprising lactone monomers, polycaprolactone, polymers comprising carbonate linkages, polycarbonate, polyglyconate, poly 24. The prosthetic valve device of any one of clauses 16-23, comprising a polymer composition selected from the group consisting of one or more of poly(trimethylene carbonate), poly(glycolide-co-trimethylene carbonate), poly(glycolide-co-trimethylene carbonate-co-dioxanone), polyurethane, polycarbonate urethane, polyester urethane, polymers with ester bonds, polyalkanoates, polyhydroxybutyrate, polyhydroxyvalerate, polydioxanone, polygalactin, natural polymers, chitosan, collagen, elastin, alginate, cellulose, hyaluronic acid, and gelatin.

[0129] (Article 25) 10. A method for producing a valve structure, comprising electrodepositing a matrix of a biodegradable, biocompatible polymer composition onto an electrodeposition target according to any one of clauses 1 to 9.

[0130] (Article 26) 26. The method of clause 25, wherein the polymer composition comprises a synthetic polymer.

[0131] (Article 27) The synthetic polymers include poly(ester urethane) urea (PEUU), poly(ether ester urethane) urea (PEEUU), poly(ester carbonate) urethane urea (PECUU), poly(carbonate) urethane urea (PCUU), polymers derived from α-hydroxy acids, polylactide, poly(lactide-co-glycolide), poly(L-lactide-co-caprolactone), polyglycolic acid, poly(dl-lactide-co-glycolide), poly(l-lactide-co-dl-lactide), polymers comprising lactone monomers, polycaprolactone, polymers comprising carbonate linkages, and polycarbonate. 27. The method of claim 26, wherein the polymer is selected from the group consisting of one or more of polyglyconate, poly(trimethylene carbonate), poly(glycolide-co-trimethylene carbonate), poly(glycolide-co-trimethylene carbonate-co-dioxanone), polyurethane, polycarbonate urethane, polyester urethane, polymers with ester bonds, polyalkanoate, polyhydroxybutyrate, polyhydroxyvalerate, polydioxanone, polygalactin, natural polymers, chitosan, collagen, elastin, alginate, cellulose, hyaluronic acid, and gelatin.

[0132] (Article 28) 26. The method according to clause 25, wherein the synthetic polymer is PEUU, PEEUU, PECUU or PCUU.

[0133] (Article 29) 29. The method of any one of clauses 25 to 28, wherein the polymer matrix is directionally biased and deposited at one or more locations on the target to generate one or more anisotropic moieties.

[0134] (Article 30) 30. The method of clause 29, wherein more than 50% of the orientation of the electrodeposited polymer is circumferentially biased within the central portion of the concave portion and more than 50% of the orientation of the electrodeposited polymer is longitudinally biased at or near the coaptation surfaces of the leaflet portions.

[0135] (Article 31) 31. The method of any one of clauses 25 to 30, wherein the shape and size of the concave portion of the electrodeposition target mimics the shape and size of the cusps of a human or animal valve.

[0136] (Article 32) 32. The method according to clause 31, wherein said concave portion has the shape and size of a cusp (leaflet) of a normal or pathological human or animal mitral, tricuspid, aortic or pulmonary valve.

[0137] (Article 33) 33. The method of any one of clauses 25-32, further comprising removing the valve structure from the electrodeposition target and trimming the valve structure to separate the distal ends of the valve leaflets.

[0138] (Article 34) 34. The method of any one of clauses 25 to 33, further comprising seeding the valve structure with cells and optionally culturing the cells on the valve structure such that the cells cover and / or infiltrate at least a portion of the valve structure.

[0139] (Article 35) 35. The method of any one of clauses 25 to 34, further comprising electrodepositing, spraying, or otherwise adding or incorporating a second polymer composition to the electrodeposition target, the second polymer composition being ECM gel, a drug, water, saline, PBS, cell culture media, cells, a biological product, a salt, a buffer, a cytokine, a growth factor, or a combination thereof.

[0140] (Article 36) 25. A method of repairing or replacing a heart valve in a patient, comprising implanting in the patient a valve prosthesis according to any one of clauses 16 to 24.

[0141] (Article 37) 37. The method of clause 36, wherein the valve prosthesis is a cardiac valve prosthesis.

[0142] (Article 38) 38. The method of clause 36 or 37, wherein said implanting is performed by the percutaneous route.

[0143] (Article 39) 39. The method of any one of claims 36 to 38, wherein the implanting is a stentless replacement of a native heart valve.

[0144] (Article 40) 40. The method of claim 39, wherein the heart valve prosthesis is sewn to one or more of the patient's heart valve annulus and, optionally, to one or more of the papillary muscles for an atrioventricular valve or the coaptation surfaces for a ventricular-aortic valve.

[0145] (Article 41) 38. The method of clause 36 or 37, wherein the valve is attached to a structured frame before implantation, and the structured frame containing the valve is attached, e.g., sutured, to the patient's valve annulus.

[0146] (Article 42) 42. The method according to any one of clauses 36 to 41, using the prosthetic valve according to any one of clauses 16 to 24.

Claims

1. 1. A method of creating a valve structure, comprising: electrodepositing a synthetic polymer composition onto an electrodeposition target to form the valve structure; and electrodepositing an extracellular matrix (ECM) gel onto the electrodeposition target to incorporate the ECM gel into the valve structure.

2. 10. The method of claim 1, further comprising removing the valve structure from the electrodeposition target and trimming the valve structure to form leaflets having distinct distal ends.

3. 10. The method of claim 1, further comprising seeding the valve structure with cells, and optionally culturing the cells on the valve structure such that the cells cover and / or infiltrate at least a portion of the valve structure.

4. 10. The method of claim 1, wherein the valve structure comprises about 70% to 85% synthetic polymer composition and about 15% to 30% ECM gel.

5. 10. The method of claim 1, wherein the synthetic polymer composition and the ECM gel are simultaneously deposited onto the electrodeposition target by electrospinning.

6. 6. The method of claim 5, wherein the synthetic polymer composition is electrospun onto the electrodeposition target and the ECM gel is sprayed onto the electrodeposition target.

7. 10. The method of claim 1, further comprising rotating the electrodeposition target while electrodepositing the synthetic polymer composition and the ECM gel onto the electrodeposition target.

8. 8. The method of claim 7, further comprising varying the rotational speed of the electrodeposition target while electrodepositing the synthetic polymer composition onto the electrodeposition target to create different degrees of circumferential bias for the fibers of the synthetic polymer composition electrodeposited on the electrodeposition target.

9. 9. The method of claim 8, wherein the electrodeposition target is rotated at a tangential velocity of about 0.3 m / sec to about 3 m / sec.

10. 9. The method of claim 8, wherein the electrodeposition target is rotated at a tangential velocity of about 1.5 m / s to produce fibers having an anisotropy ratio comparable to the natural anisotropy.

11. 10. The method of claim 1, further comprising adjusting a longitudinal distance between a source of the synthetic polymer composition and the electrodeposition target during electrodeposition to produce varying degrees of longitudinal bias in the synthetic polymer fibers deposited on the electrodeposition target.

12. 12. The method of claim 11, wherein the electrodeposition target is moved relative to the synthetic polymer composition source at a scanning speed of about 0.25 cm / sec to about 2.5 cm / sec to adjust the bending stiffness of a portion of the valve structure.

13. 10. The method of claim 1, further comprising varying the magnitude of a voltage applied to the electrodeposition target while electrodepositing the synthetic polymer composition and the ECM gel onto the electrodeposition target.

14. 10. The method of claim 1, wherein the electrodeposition target has an outer surface including at least one conductive surface portion and at least one non-conductive surface portion.

15. 10. The method of claim 1, wherein the synthetic polymer composition is deposited at one or more locations on the exterior surface of the electrodeposition target with a directional bias to produce one or more anisotropic portions of the valve structure.

16. 10. The method of claim 1, wherein the step of electrodepositing the synthetic polymer composition onto the electrodeposition target comprises depositing the synthetic polymer composition such that at least 50% of the deposited synthetic polymer composition is circumferentially biased around the periphery of the electrodeposition target and less than 50% of the deposited synthetic polymer composition is longitudinally biased toward a distal end of the electrodeposition target.

17. The electrodeposition target is a support disposed around the rotation axis of the electrodeposition target; a conductive body forming at least one conductive surface portion of the electrodeposition target connected to the support, the conductive body including a plurality of longitudinally extending ridges and recessed portions between the plurality of ridges; a non-conductive coating covering at least a portion of the plurality of protrusions forming at least one non-conductive surface portion of the electrodeposition target, the concave portion of the conductive body being uncovered.

18. A method for producing an artificial valve formed from a valve structure obtained by the method according to any one of claims 1 to 17, the artificial valve comprising: a tubular support portion defining an opening and having a longitudinal axis; at least two leaflets extending longitudinally from the tubular support portion, each leaflet having a concave central portion, a peripheral portion around the concave central portion, a proximal end connected to the tubular support portion, and a distal end longitudinally distal to the tubular support portion, the peripheral portions of adjacent leaflets being partially coapted at the tubular support portion and adjacent the tubular support portion to form a coaptation surface between adjacent leaflets; A manufacturing method comprising:

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