Novel nanofiber-based material for endovascular scaffolds Rui Wang,1,2 Nicole Levi-Polyanchenko,1,2 Michael Morykwas,1,2 Louis Argenta,1 William D. Wagner1,2 1
Department of Plastic and Reconstructive Surgery, Wake Forest University School of Medicine, Medical Center Blvd, Winston-Salem, North Carolina 27157 2 Virginia Tech - Wake Forest University School of Biomedical Engineering and Science, Medical Center Blvd, Winston-Salem, North Carolina 27157 Received 12 February 2014; revised 17 June 2014; accepted 25 June 2014 Published online 15 July 2014 in Wiley Online Library (wileyonlinelibrary.com). DOI: 10.1002/jbm.a.35267 Abstract: Conventional collagen-based heart valves eventually fail because of insufficient replacement of graft material by host tissue. In this study, type I collagen was blended with silk fibroin and the synthetic elastic polymer poly (glycerol–sebacate) (PGS) in varying proportions to create multifunctional electrospun nanofibrous materials tailored for use as endovascular scaffolds such as heart valve replacement. Depending on the blended material the elastic moduli ranged from 2.3 to 5.0 Mpa; tensile stresses ranged from 0.8 to 1.5 Mpa; and strains ranged from 30% to 70%. Electrospun materials with a weight ratio of 4.5:4.5:1 (collagen, fibroin, and PGS) (termed PFC mats) were the most similar to native heart valves. In vitro degradation of PFC mats was 0.01% per week.
Endothelial cells adhered to, proliferated, and formed cell– cell junctions on PFC mats. Compared with collagen hydrogels and electrospun collagen mats respectively 220–290% less platelet adhesion was observed for PFC mats. The study demonstrates that PFC material has superior mechanical properties, low degradation, and reduced thrombogenic potential and suggests that further investigation of this bioC 2014 material for cardiovascular applications is warranted. V Wiley Periodicals, Inc. J Biomed Mater Res Part A: 103A: 1150–1158, 2015.
Key Words: heart valve, heart valve scaffold, nanocomposite scaffold, polyglycerol sebacate, fibroin, silk, collagen
How to cite this article: Wang R, Levi-Polyanchenko N, Morykwas M, Argenta L, Wagner WD. 2015. Novel nanofiber-based material for endovascular scaffolds. J Biomed Mater Res Part A 2015:103A:1150–1158.
INTRODUCTION
Heart valve malfunction continues to be a major health problem, affecting one in forty adults in the United States.1 Replacement of native valves with animal derived valves or totally synthetic bioprostheses is the standard of care if the native valve cannot be repaired.2 Although numerous potential valve substitutes are available, none are optimal and all suffer from major potential complications such as degeneration, structural failure, thrombogenicity, or loosening of position.3–5 Typically, bioprosthetic collagen based valves have early valve degeneration and significant risks of postoperative failure. The failure of these materials and the thrombogenicity of mechanical devices necessitate the development of improved materials for heart valve repair and/or replacement. The purpose of this study was to fabricate an improved material composed of blended Type I collagen, silk fibroin, and poly (glycerol–sebacate) (PGS) which could serve as an extracellular matrix (ECM) mimetic and functional heart valve graft. In order to create an appropriate graft material that meets physiological requirements, a strategy of incorporating natural proteins with an elastic synthetic polymer was applied. Type I collagen was selected because it is the
most abundant and load-bearing component of valve tissue while elastin provides flexibility. Moreover, type I collagen motifs specifically bind to a2b1 and avb1 integrins through RGD, DGEA, or GFOGER peptide sequences to facilitate endothelial cell adhesion.6,7 Silk fibroin was selected to improve the strength of graft material. Interchain hydrogen bonds in silk fibroin protein assemble the polypeptide chains into a highly crystalline b-sheet conformation which imparts a slow degradation rate.8,9 Since elastin extracted from tissue is insoluble, PGS, an elastic polymer was blended with two other materials in order to substitute for native elastin.10 Moreover, PGS has been reported to support endothelial cell growth and has superior hemocompatibility over other commonly used synthetic polymers such as used poly (L-lactide-co-glycolide) (PLGA).11–13 To create a biomimetic valvular graft-blended material, electrospinning, a fabrication technique that applies high voltage power to a solution of polymer blends at a syringe tip, was applied. Fibers are generated when the electrostatic force overcomes the surface tension of the polymer solution.14 This technique was used to create an ECM analogous graft blended material which was tested physically,
Correspondence to: William D. Wagner; e-mail: [email protected] Contract grant sponsor: Department of Plastic and Reconstructive Surgery Research Funds
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mechanically, and in studies designed to investigate endothelial cell growth and thrombogenicity. MATERIALS AND METHODS
Materials Silk fibroin protein was extracted using an aqueous solvent processing method according to the published procedure with modifications.8 Raw silk was boiled in 0.02M Na2CO3 at 100 C for 30 min, rinsed twice with DDH2O, squeezed, and air dried. Fibroin was then dissolved in 5.0M CaCl2, and centrifuged at 20003g to remove contaminants. The fibroin solution was dialyzed and lyophilized. The identity and purity of fibroin was confirmed following amino acid analysis and molecular weight determination using sodium dodecyl sulphate-polyacrylamide gel electrophoresis.15–18 PGS prepolymer was synthesized from glycerol and sebacic acid following published methods.13,19 Type I Collagen from calf skin was purchased commercially (Elastin Products Corp, MO). Fabrication and characterization of the electrospun mats Electrospun solutions were prepared using type I collagen, silk fibroin, and PGS at different weight ratios of 9:0:1, 8:1:1, 4.5:4.5:1, 1:8:1, and 0:9:1 respectively, dissolved in 1, 1, 1, 3, 3, 3-hexafluoro-2-propanol (HFP). A syringe was fixed on a Baxter infusion pump (Model AS50) to eject the polymer solution at a rate of 3 mL/h. A 35 kV high voltage (Gamma High Voltage Research, Ormond Beach, FL) was applied at a distance of 20 cm between the metal collector plate (12 cm by 12 cm) and the syringe tip to generate a sufficient electrostatic force for electrospun mat formation. The electrospun fabricated mats were removed from the collector plates and heated at 120 C for 48 h followed by treatment with glutaraldehyde vapor for 24 h to crosslink the material.19,20 Electrospun mats were washed with 0.2M glycine before use.21 Images of mats taken by scanning electron microscopy (JEOL JSM-6330F) were used to measure fiber diameters with NIH Image J software. Sixteen random measurements of fiber diameters were obtained from each mat. Chemical functional groups were characterized using a Perkin-Elmer FTIR spectrophotometer. The thermal transition temperatures were measured using differential scanning calorimetry (DSC) (TA instruments, New Castle, DE) from 260 C to 300 C at an increment rate of 20 C/min (n 5 3 per sample). To measure mat porosity, the dry weight of samples was measured and samples were hydrated in 5 mL DDH2O for 30 min at room temperature on a mechanical shaker. After blotting dry, the hydrated weight was determined.22 The volume of each electrospun mat sample was obtained by dividing the weight by the density of the mat. Then, the porosity was calculated using the following equation: porosity 5 volume of water/(volume of water 1 volume of electrospun mat). The density and weight ratios of type I collagen (1.40 g/cm3), silk fibroin (1.31 g/cm3), and PGS (1.13 g/cm3) were used to calculate the densities of the composites.10,22–24
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Mechanical testing An Instron 5500R mechanical tester (Instron Corporation, Norwood, MA) with a 500 N load cell and BlueHill software was used to perform uniaxial tensile tests of the mats at an elongation rate of 10 mm/min. A dumbbell stamp (ASTM D638-IV cutting die, Pioneer-Dietecs Corporation, Norwood, MA) was used to prepare samples of each electrospun mat. Prior to analysis, all mats were hydrated in 100 mL DDH2O for 10 min. Three measurements from each type of electrospun mat were used to calculate the elastic modulus, tensile strength, and percentage elongation. A suture retention test was performed according to modifications of a published protocol.25,26 Electrospun mats were prepared 2 cm in length and 0.5 cm in width (n 5 5) and monofilament prolene (3-0 monofilament; Ethicon, Somerville, NJ) was placed 0.5 cm from the edge of each mat to form a single loop. Using a BOSE-Electroforce mechanical tester (Bose Electroforce, Eden Prairie, MN), samples were stretched at a rate of 10 mm/min until the suture completely tore through the material. Material degradation For degradation, mats (n 5 5) were air dried and stamped into 1 3 10 mm strips and conditioned in degradation solution (10 mL phosphate buffered saline (PBS) containing 0.1% sodium azide at 37 C) during the first week. At specific time intervals the material was removed, washed three times with deionized water, and weighed to determine weight loss. Starting material weight loss during 1 week of conditioning represented water weight in the sample thus weights at 1 week were used as baseline values for the degradation study. Samples from three batches of mats were tested in separate experiments. The weights of samples were measured every week over a 30-week period. The percentage weight loss was calculated from the ratio of mass change after degradation to the original mass of the scaffold according to the formula: Weight loss (%) 5 Wl/Wi 3 100%. (Wi, initial weight of the electrospun mats and Wl, weight loss of the same electrospun mats after exposure in degradation solution.) One extra sample was prepared at each time point for morphological assessment using scanning electron microscopy (SEM). Cell culture study To test cell-material interaction, PFC mats (at 4.5:4.5:1:Collagen: Fibroin: PGS weight ratio) were sterilized with ethanol for 30 min followed by ultraviolet light exposure of each side for 1 h followed by hydration in sterile PBS. To mimic the in vivo condition where blood proteins would interact with implanted material, fibronectin (Sigma, St Louis, MO) was used at a concentration of 100 mg/mL to coat mats and culture dishes. HUVECs (ATCC: CRL-1730) were seeded on the substrates in a 16-well culture plate at a cell density of 50,000 cells/well. On Days 3 and 7, rhodamine-phalloidin and sytox green dyes (Invitrogen, Eugene, OR) were used to stain F-actin and nuclei, respectively. Image J software (NIH) was used to quantitate cell numbers from representative confocal micrographs.
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TABLE I. Fiber Diameters in Mats of Different Compositions Sample Type
Fiber Diameters (nm)
Collagen : PGS (9 : 1) Collagen : Fibroin: PGS (8:1:1) Collagen : Fibroin: PGS (4.5:4.5:1) Collagen : Fibroin: PGS (1:8:1) Fibroin : PGS (9:1)
2067 6 168 4577 6 697 2952 6 240 784 6 77 694 6 43
All values represent means 6 SEM. The fiber diameters were measured from 16 randomly selected fibers of two representative scanning electron micrographs. Diameters ranged from 600 to 3000 nm. FIGURE 1. FTIR spectra of collagen, PGS, fibroin, and composites.
Platelet adhesion to mats and mats with cells R Human whole blood was drawn into a 2.7 mL BD VacutainerV Coagulation Tube (BD, Franklin Lakes, NJ) and centrifuged at 800 rpm for 15 min at 25 C with Harrier 18/80 centrifuge (Sanyo Gallenkamp, Loughborough, UK) to obtain platelet rich plasma (PRP). Three materials: PFC mats, (at weight ratio of 4.5:4.5:1 Collagen: Fibroin:PGS) Type I electrospun collagen mats and collagen hydrogels (rat tail, type I collagen, BD Biosciences, Bedford, MA) were prepared to compare platelet adhesion. PRP was diluted with PBS to 2.16 3 108 platelets/mL, and 100 mL was applied to the center of the substrates. After 15 min materials were washed twice with PBS followed by fixation in 2.5% glutaraldehyde. Factin, the contractile protein expressed in platelets, was stained with rhodamine-phalloidin to identify platelet interaction with materials. Adherent platelets on various substrates were assessed following confocal microscopy. The morphology of platelets was evaluated by SEM. Numbers of adherent platelets were quantified using NIH Image J Software. In a separate experiment, platelet adhesion was examined using materials that were preseeded with HUVECs for 24 h. The rationale for this experiment was to closely simulate an in vivo state where endothelial cells would produce a nonthromobenic glycocalyx and condition the material.
composites at high collagen blends (Collagen: Fibroin: PGS; 0:9:1, 1:8:1) closely mimicked the FTIR spectrum of collagen. This is seen through the broad peak at 3275 cm21, which is assigned to the OAH stretch and NAH stretch of collagen (Fig. 1—region 1). Amide I and II of collagen were present at 1622–1630 cm21 (Fig. 1—region 4) and 1546– 1515 cm21 (Fig. 1—region 5), respectively, while amide III can be seen at 1228–1237 cm21 (Fig. 1—region 6). At high silk fibroin blends (Collagen: Fibroin: PGS 1:8:1, 0:9:1), the composite spectra closely mimicked the FTIR spectrum of silk fibroin. Silk fibroin was found in the b sheet conformation within the composite blend based on the carbonyl stretching frequencies for amides I, II, and III at 1620 cm21, 1512 cm21, and 1230 cm21 respectively.27,28 The stretching frequency at 3280 cm21 is indicative of the ANAH stretching vibration shown as a broad peak for amide A. The peak at 1700 cm21 was assigned to be the C@O stretch in amide I b sheets, and 1225–1233 cm21 referred to the CAN stretch and CANAH bend in amide III b sheets structure.30 The stretching frequencies for the PGS polymer can be seen in the spectra of all of the composite blends. The C@O stretch for the ester can be seen at 1734 cm21 (Fig. 1—region 3) and the broad OAH stretch at 3500 cm21 represents the hydrogen bonded hydroxyl group of PGS.13 The peaks shown at
Statistical analysis Data are presented as mean 6 standard error of the mean (SEM) unless otherwise noted. Statistical significance was determined using one-way analysis of variance (ANOVA) following post hoc testing for multiple groups when appropriate. A p value of p < 0.05 was considered statistically significant. StatView 5.0 (SAS Institute, Cary, NC) was used to perform the statistical analysis. RESULTS
Composite characterization FTIR spectroscopy was used to determine the structural characteristics of the composite materials. The vibrational stretching frequencies of the pure collagen, silk fibroin and PGS polymers were found to be comparable to literature values (Fig. 1).9,13,27–31 The characteristic stretching frequencies of Type I collagen, silk fibroin, and PGS were evident in the composite materials. The FTIR spectra of the
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FIGURE 2. DSC scans of electrospun mats.
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TABLE II. Uniaxial Tensile Testing of Electrospun Mats Sample Type Collagen Collagen: PGS (9:1) Collagen: Fibroin: PGS (8:1:1) Collagen: Fibroin: PGS (4.5:4.5:1) Collagen: Fibroin: PGS (1:8:1) Fibroin: PGS (9:1)
Elastic Modulus (Mpa) a
3.67 6 0.12 2.25 6 0.16b 2.76 6 0.20b 4.11 6 0.13a,c 4.97 6 0.27d
Stress (Mpa) a,d
0.69 6 0.06 1.30 6 0.09b,c 1.10 6 0.09b 1.45 6 0.05c N/A: too brittle to be determined 0.82 6 0.09d
Strain (mm/mm) 0.23 6 0.03a d 0.62 6 0.02b 0.44 6 0.03b,c 0.41 6 0.01a,b,c 0.33 6 0.12a,c
All values represent means 6 SEM. The tensile stress, strain and elastic modulus of electrospun mats at different composite ratios were measured. Means having a similar superscript letter are not significantly different. Means not having a similar superscript letter (a,b,c,d) are significantly different (p < 0.05).
2928 cm21 and 2855 cm21 (Fig. 1—region 2) represent the sp3 CAH symmetric and antisymmetric stretches of PGS. Fiber morphology and diameter Fiber diameters ranged from 694 to 4577 nm (Table I) and showed high interconnectivity. In general, thinner and more rounded fibers were observed for the electrospun mats with higher fibroin content (0:9:1, 1:8:1 for Collagen: Fibroin: PGS weight ratios) as compared with the thicker and more flat fibers of electrospun mats with high proportions of collagen (9:0:1, 8:1:1, 4.5:4.5:1). Thermal transition analysis of electrospun mats DSC scans were analyzed to determine thermal transition temperatures of electrospun mats. By incorporating increasing amount of fibroin, a shift of thermal transition temperatures to a higher range was observed in comparison to PGS alone (226 C) (Fig. 2). The results suggest that the electrospun mats made from collagen, silk fibroin, and PGS composites would be thermally stable for in vivo application. Porosity measurements The porosity of the electrospun mats ranged from 67 to 80% which would enable diffusion of nutrients and oxygen transport.
FIGURE 3. Forces required to tear sutures from electrospun mats and procine heart valves. Duplicate samples shown by the two bars were tested for each material. Collagen: Fibroin: PGS (4.5:4.5:1) showed the highest suture pull-out strength with an average maximum load of 0.32 N as compared with a maximum average load of 0.64 N for fresh porcine heart valve.
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Mechanical tensile testing The elastic moduli of the electrospun mats (Table II) ranged from 2.25 Mpa (Collagen: PGS at a 9:1 ratio) to 4.97 MPa (Fibroin: PGS at a 9:1 ratio). Electrospun mats with collagen alone had a similar elastic modulus of 3.67 Mpa compared with collagen-based porcine valvular grafts: 3.68 MPa for fresh valves and 3.95 MPa for glutaraldehyde fixed valves.31 Addition of PGS to primarily collagen containing mats created a material with significantly higher percentage of elongation but less ultimate tensile strength. All mats produced had ultimate tensile stress values that ranged between 0.69 MPa and 1.45 MPa. Among all electrospun mats produced, the composites at 4.5:4.5:1 of collagen: silk fibroin: PGS weight ratio had the best overall mechanical strength and elasticity. The suture retention test was used to measure the maximum force required to disrupt sutures from the material. For this test, material thickness was similar to the thickness of fresh aortic valve. Collagen: silk fibroin: PGS (4.5:4.5:1) mats required the greatest force to tear the tissue with maximum suture retention force of 0.32 N. Fresh porcine heart valve had a maximum force of 0.64 N (Fig. 3). On the basis of the results of all materials and composites studied, electrospun mats containing collagen-fibroinPGS (4.5:4.5:1 weight ratio) were selected for further study. This material, termed PFC, was used to evaluate degradation, cellular compatibility, and thrombogenicity. The primary selection criteria at this point in the study was based
FIGURE 4. In vitro degradation of PFC mat during a 30-week period. Each data point is presented as the mean 6 SEM (n 5 4). [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
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FIGURE 5. Quantitation of HUVEC grown on culture dishes or PFC mats. Each bar represents the mean 6 SEM for measurements taken from representative photomicrographs (n 5 4). A significant (p < 0.05) increase in the number of HUVECs was observed from day 3 to 7 of cells on both substrates.
on mechanical properties and similarity to native fresh heart valves. Degradation of PFC mats Degradation of the PFC mats after 1 week of preconditioning was found to be only 0.3% weight loss per week over a 30-week incubation period (Fig. 4). To the unaided eye, PFC mats remained intact during the entire experiment. By SEM, nanofibers maintained a consistent morphology within the meshwork during the course of study. No significant differ-
FIGURE 7. Confocal images illustrating platelet adhesion to different substrates. PRP (1.08 3 107 platelets) was added to each well of a 48well plate. The images were acquired after 15 min of incubation on various substrate surfaces at 37 C. A: PFC mat, B: collagen mat, C: culture dish, D: collagen gel. More adhered platelets and the formation of microthrombi on the collagen gel and collagen mat are observed as compared with the culture dish and the PFC mat. Platelets were visualized by rhodamine–phalloidin staining (red) for F-actin protein (Scale bar: 50mm).
ences in fiber diameters were found over the course of the experiment (data not shown). While not part of this current study design samples were maintained under conditions of degradation for long term and at 108 weeks the change in total weight was 4% (detailed data not presented).
FIGURE 6. Confocal microscopic images illustrating HUVEC morphology and density. A: PFC mat at day 3, B: PFC mat at day 7, C: culture dish at day 3, D: culture dish at day 7. Cells were seeded at 50,000 per dish on a 48-well plate. Representative photomicrographs depict increased cell numbers from day 3 to day 7 on both substrates. Cells were stained for F-actin protein (red) using rhodamine-phalloidin and for nuclei (green) using sytox green (Scale bar: 50 mm).
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FIGURE 8. Platelet adhesion to different substrates. Adherent platelets were counted from representative photomicrographs taken from the culture dish, PFC mat, collagen mat, and collagen gel after 15 min incubation at 37 C. Each bar represents the mean 6 SEM (n 5 3). Means having a similar superscript letter are not significantly different. Means not having a similar superscript letter are significantly different (p < 0.05).
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FIGURE 9. Scanning electron micrographs of collagen gel (A and D), collagen mat (B and E) and PFC mat (C and F) after 15 min of incubation with PRP. The images demonstrate the presence of a single layer of platelets on the PFC mat and minor platelet activation in comparison to activated platelets on the collagen mat and collagen gel (Magnification: 31500 for A, B, and C, scale bar: 10 mm; 35000 for D, E, and F, scale bar: 1 mm).
Cell adhesion and proliferation Biocompatibility and nonthrombogenic property is essential for materials in contact with blood. When HUVECs were cultured on PFC mats, cell numbers increased significantly from days 3 to 7 (p < 0.05) (Fig. 5). Morphologically, an organized endothelial cell monolayer formed on PFC mats. The intense staining pattern of F-actin at the cell borders for cells cultured on PFC mats suggested the formation of tight endothelial cell junctions [Fig. 6(B)]; however, future studies will be needed to prove the formation of tight junctions. In contrast, HUVECs on culture dishes did not form uniform monolayer (Fig. 6).
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Platelet adhesion to PFC mats and mats preseeded with HUVECs The thrombogenic potential of PFC mats precultured with and without endothelial cells was accessed by examining platelet adhesions. Confocal images of platelets stained with rhodamine–phalloidin demonstrated a low level of platelet adhesion on polysterene culture dishes (Fig. 7). Single platelets or small clumps comprised of 2–3 platelets were present on PFC mats. In sharp contrast, platelets, and platelet aggregates adhered extensively to the collagen gel (positive control). Images demonstrate increased platelets numbers as well as increased sizes of platelet aggregates (Fig. 7). Platelet numbers on collagen mats, and collagen gels were
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FIGURE 10. Confocal images illustrating the association and interaction of platelets with HUVECs on different substrate. PRP (1.08 3 107 platelets) was added to each well of a 48-well plate. The images were acquired after 15 min of incubation of platelets on various substrate surfaces at 37 C. A: PFC mat, B: collagen mat, C: culture dish, D: collagen gel. Formation of microthrombi on the collagen gel and increased size of platelet aggregates was observed on the collagen gel and collagen mat as compared with the culture dish and PFC mat. F-actin (red) of cells and platelets were stained using rhodamine–phalloidin and nuclei (green) of cells were stained using sytox green (Scale bar: 50 mm).
respectively 150 and 790% higher (p < 0.05) than on the PFC mats (Fig. 8). The morphology of platelets on various substrates was examined by SEM. Extensive activation and fusion of degranulated platelets was observed on collagen gels [Fig. 9(A,D)]. Platelets on electrospun collagen mats also were activated and stellate shaped [Fig. 9(B,E)]. Platelets on PFC mats appeared to be distinctly spherical [Fig. 9(C,F)] and morphologically at minimal activation state in comparison to the appearance of platelets on collagen electrospun mats or the collagen gels. In a second experiment, platelets were applied to mats cultured for 24 h with HUVECs (50,000 cells/well in a 48well plate) (Fig. 10). Adherent platelets were observed on collagen electrospun mats as well as the formation of microthrombi of larger sizes on collagen gels. In clear contrast, platelets on PFC mats were all spherical in shape and no microthrombi were observed (Fig. 10). A 60% reduction in total adherent platelets was observed on the PFC mat as
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compared with the culture dish (Fig. 11) (p < 0.05). The number of platelets decreased 290% on PFC mats compared with electrospun collagen mats (p < 0.05), and 220% compared with collagen gels (p < 0.05).
DISCUSSION
In this study, an electrospinning fabrication technique was used to produce PFC mats for potential use as prosthetic heart valves. The electrospinning fabrication procedure is less time consuming compared with decellularized tissue processing procedures.32 As a feasible fabrication technique, electrospinning on molds could be used to recreate the geometry of heart valve tissue or for use in transcatheter heart valves procedures.33 PFC mats demonstrated superior strength and flexibility compared to primarily collagen rich biomaterials. Blending silk fibroin with type I collagen and PGS created a novel and strong material that still possessed elastic properties.
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FIGURE 11. Platelet interaction with cells cultured on different substrates. Numbers of adhered platelets were counted from representative photomicrographs taken from culture dish, PFC mat, collagen mat, and collagen gel substrates after 15 min incubation at 37 C. White bars (w) indicate areas of the material without endothelial cells gray bars (䊏) indicate platelet counts for areas with cells. Each bar represents the mean 6 SEM (n 5 3). Means not having a similar superscript letter are significantly different (p < 0.05).
Various ratios of type I collagen and silk fibroin were tested to mimic the mechanical and physiological properties desired for heart valve construction. Results showed electrospun mats at a 4.5:4.5:1 weight ratio of collagen: silk fibroin: PGS (PFC mat) had properties closely matching heart valve mechanical properties, an interconnective network, cell compatibility, and low thrombogenecity. In the suture retention tests, PFC mats showed sufficient suture retention, suggesting that PFC mats could be used in surgical implant procedures. A major limitation of the study includes the absence of in vivo data in an appropriate animal model on the performance of PFC as a valvular graft. Studies prior to implantation should include assessment of PFC in a tissue bioreactor where performance under increasing pressures can be evaluated. In addition, the PFC material should be tested in conjunction with devices such as catheters to be used in clinical deployment. This study suggested a new strategy in improving the strength of the graft material by composite selection in order to prevent cell-mediated buckling, a major challenge in maintaining the mechanical integrity of valvular graft.32,33 This was accomplished by incorporating silk fibroin with helical collagen and the elastic polymer PGS to obtain both strength and flexibility.34–36 The physical properties and molecular interactions demonstrated by FTIR suggest that PFC fibers have a property imparting strength and viscoelasticity to the composite material.29,37,38 PFC mats showed limited degradation with minimal weight loss during a 30-week degradation period. The ultrastructure of the nanofibers showed little change and appeared superior to the reported degradation of fully cross-linked collagen or polylactic acid electrospun constructs.9,39–41 The results of the studies on PFC mats are consistent with published in vivo degradation studies indicating that silk fibroin degrades slower than most collagen-based materials, such as polycaprolactone-collagen scaffolds.42,43
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In this study, HUVEC attached and proliferated on PFC mats. This is the first step to achieve tissue remodeling and maturation of implanted heart valves in vivo.6,7,44,45 Published studies have shown elevated thrombogenic risk associated with decellularized collagen-based heart valve grafts.46 In the present study significant reductions in platelet adhesion and platelet activation were observed on PFC mats. The observation that the PFC material had reduced platelet adherence and low activation was a surprising finding since enhanced endothelial cells affinity was observed for the material. The potential exists that the PFC composite had a molecular interaction of components that result in selective patterns of masked or exposed binding sites for endothelial cells and platelet cell surface ligands. Future studies will be required to identify the exact molecular mechanism. CONCLUSIONS
A major accomplishment of the current investigation was the development and testing of a new composite material that can be used as material for valvular surgery. Material comprising PFC mats with collagen: fibroin: PGS at a ratio of 4.5:4.5:1 demonstrated comparable mechanical strength as fresh heart valve tissue. Functional tests showed minimal weight loss and sustained nanofiber structural integrity over a 30-week degradation study. Organized endothelial monolayers formed on PFC mats and based on platelet adhesion studies were less thrombogenic compared to collagen nanofiber mats and collagen gels. This study demonstrated that the PFC composite is a promising material, with potentially superior hemocompatibility compared with structurally similar collagen materials. ACKNOWLEDGMENTS
The authors gratefully acknowledge the use of resources of the Wake Forest University Center of Nanotechnology and Molecular Materials; Dr. Roy Hantgan and Dr. Mark Lively, Wake Forest University Biochemistry Department and Protein Analysis Core Laboratory; Dr. Jessica Sparks, Miami University Chemical and Paper Engineering. The authors acknowledge Dr. Christopher MacNeill’s assistance in the scientific comments and review of the manuscript. They thank Robin Simmons for the assistance in manuscript preparation and Victoria Jiang for editorial comments REFERENCES 1. Nkomo VT, Gardin JM, Skelton TN, Gottdiener JS, Scott CG, Enriquez-Sarano M. Burden of valvular heart diseases: A population-based study. Lancet 2006;368:1005–1011. 2. Apte SS. Current developments in the tissue engineering of autologous heart valves: Moving towards clinical use. Future Cardiol 2011;7:77–97. 3. Mol A. Review article: Tissue engineering of semilunar heart valves: Current status and future developments. J Heart Valve Dis 2004;13:272–280. 4. Jordan JE, Williams JK, Lee SJ, Raghavan D, Atala A, Yoo JJ. Bioengineered self-seeding heart valves. J Thorac Cardiovasc Surg 2012;143:201–208. 5. Breuer CK. Application of tissue-engineering principles toward the development of a semilunar heart valve substitute. Tissue Eng 2004;10(11-12):1725–1736. 6. Gu X, Masters KS. Regulation of valvular interstitial cell calcification by adhesive peptide sequences. J Biomed Mater Res A 2010; 93:1620–1630.
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7. Shekaran A, Garcia AJ. Nanoscale engineering of extracellular matrix-mimetic bioadhesive surfaces and implants for tissue engineering. BBA- Gen Subjects 201;1810:350–360. 8. Rockwood DN, Preda RC, Yucel T, Wang X, Lovett ML, Kaplan DL. Materials fabrication from Bombyx mori silk fibroin. Nat Protocols 2011;6:1612–1631. 9. Horan RL, Antle K, Collette AL, Wang Y, Huang J, Moreau JE, Volloch V, Kaplan DL, Altman GH. In vitro degradation of silk fibroin. Biomaterials 2005;26:3385–3393. 10. Pomerantseva I, Krebs N, Hart A, Neville CM, Huang AY, Sundback CA. Degradation behavior of poly(glycerol sebacate). J Biomed Mater Res A 2009;91:1038–1047. 11. Motlagh D, Yang J, Lui KY, Webb AR, Ameer GA. Hemocompatibility evaluation of poly(glycerol-sebacate) in vitro for vascular tissue engineering. Biomaterials 2006;27:4315–4324. 12. Lee, KW, Stolz DB, Wang Y. Substantial expression of mature elastin in arterial constructs. Proc Natl Acad Sci USA 2011;108: 2705–2710. 13. Wang Y, Ameer GA, Sheppard BJ, Langer R. A tough biodegradable elastomer. Nat Biotech 2002;20:602–606. 14. Sell SA, Wolfe PS, Garg K, McCool JM, Rodriguez IA, Bowlin GL. The use of natural polymers in tissue engineering: a focus on electrospun extracellular matrix analogues. Polymers 2010;2:522– 553. 15. Matsumoto A, Lindsay A, Abedian B, Kaplan DL. Silk fibroin solution properties related to assembly and structure. Macromol Biosci 2008;8:1006–1018. 16. Matsumoto A, Chen J, Collette AL, Kim UJ, Altman GH, Cebe P, Kaplan DL. Mechanisms of silk fibroin sol-gel transitions. J Phys Chem 2006;110:21630–21638. 17. Lucas F, Shaw JT, Smith SG. The amino acid sequence in a fraction of the fibroin of Bombyx mori. Biochem J 1957;66:468–470. 18. Schroeder WA, Kay L, Lewis B, Munger N. The amino acid composition of bombyx mori silk fibroin and of tussah silk fibroin. J Am Chem Soc 1955;77:3908–3913. 19. Sant S, Khademhosseini A. Fabrication and characterization of tough elastomeric fibrous scaffolds for tissue engineering applications. In: Engineering in Medicine and Biology Society (EMBC), 2010 Annual International Conference of the IEEE. 2010 20. Yi F, LaVan DA. Poly(glycerol sebacate) nanofiber scaffolds by core/shell electrospinning. Macromol Biosci 2008;8:803–806. 21. Sung HW, Chen CN, Huang RN, Jsu JC, Chang WH. In vitro surface characterization of a biological patch fixed with a naturally occurring crosslinking agent. Biomaterials 2000;21:1353–1362. 22. Soliman S, Sant S, Nichol JW, Khabiry M,Traversa E, Khademhosseini A. Controlling the porosity of fibrous scaffolds by modulating the fiber diameter and packing density. J Biomed Mater Res A 2011;96:566–574. 23. Minoura N, Tsukada M, Nagura M. Fine structure and oxygen permeability of silk fibroin membrane treated with methanol. Polymer 1990;31:265–269. 24. De Cupere VM, Van Wetter J, Rouxhet PG. Nanoscale organization of collagen and mixed collagen2pluronic adsorbed layers. Langmuir 2003;19:6957–6967. 25. Trowbridge EA, Lawford PV, Crofts CE. Pericardial heterografts: A comparative study of suture pull-out and tissue strength. J Biomed Eng 1989;11:311–314. 26. Edwards MB, Draper ERC, Hand JW, Taylor KM, Young IR. Mechanical testing of human cardiac tissue: Some implications for MRI safety. J Cardiovasc Magn Reson 2005;7:835–840. 27. Um IC, Kweon HY, Park YH, Hudson S. Structural characteristics and properties of the regenerated silk fibroin prepared from formic acid. Int J Biol Macromol 2001;29:91–97.
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28. Cilurzo F, Gennari CG, Selmin F, Marotta LA, Minghetti P, Montanari L. An investigation into silk fibroin conformation in composite materials intended for drug delivery. Int J Pharm 2011; 414(1–2):218–224. 29. Renugopalakrishnan V, Chandrakasan G, Moore S, Huston TB, Berney CV, Rajendra S, Bhatnagar S. Bound water in collagen. Evidence from Fourier Transform Infrared and Fourier transform infrared photoacoustic spectroscopic study. Macromolecules 1989;22:4121–4124. 30. Hayashi T, Mukamel S. Vibrational2Exciton couplings for the amide I, II, III, and A modes of peptides. J Phys Chem B 2007;111: 11032–11046. 31. Vesely I, Noseworthy R. Micromechanics of the fibrosa and the ventricularis in aortic valve leaflets. J Biomec 1992;25:101–113. 32. Cebotari S. Use of fresh decellularized allografts for pulmonary valve replacement may reduce the reoperation rate in children and young adults: Early report. Circulation 2011;124(11S):S115– S123. 33. Dijkman PE, Driessen-Mol A, Frese L, Hoerstrup SP, Baaijens FP. Decellularized homologous tissue-engineered heart valves as offthe-shelf alternatives to xeno- and homografts. Biomaterials 2012; 33:4545–4554. 34. Billiar KL, Sacks MS. Biaxial mechanical properties of the native and glutaraldehyde-treated aortic valve cusp. II. A structural constitutive model. J Biomecl Eng 2000;122:327–335. 35. Sacks MS, Schoen FJ, Mayer JE. Bioengineering challenges for heart valve tissue engineering. Annu Rev Biomed Eng 2009;11: 289–313. 36. Beun LH, Beaudoux XJ, Klein JM, de Wolf FA, Stuart MA. Selfassembly of silk-collagen-like triblock copolymers resembles a supramolecular living polymerization. ACS Nano 2011;6:133–140. 37. Jiang C, Wang X, Gunawidjaja R, Lin YH, Gupta MK, Kaplan DL, Naik RR, Tsukruk VV. Mechanical properties of robust ultrathin silk fibroin films. Adv Funct Mater 2007;17:2229–2237. 38. Zoccola M, Aluigi A, Vineis C, Tonin C, Ferrero F, Piacentino MG. Study on cast membranes and electrospun nanofibers made from keratin/fibroin blends. Biomacromolecules 2008;9:2819–2825. 39. Kim K, Yu M, Zong X, Chiu J, Fang D, Seo YS, Hsiao BS, Chu B, Hadjiargyrou M. Control of degradation rate and hydrophilicity in electrospun non-woven poly(d,l-lactide) nanofiber scaffolds for biomedical applications. Biomaterials 2003;24:4977–4985. 40. Zong X, Ran S, Kim KS, Fang D, Hsiao BS, Chu B. Structure and morphology changes during in vitro degradation of electrospun poly(glycolide-co-lactide) nanofiber membrane. Biomacromolecules 2003;4:416–423. 41. Liu T, Teng WK, Chan BP, Chew SY. Photochemical crosslinked electrospun collagen nanofibers: Synthesis, characterization and neural stem cell interactions. J Biomed Mater Res A 2010;95:276–282. 42. Zhou J, Cao C, Ma X, Hu L, Chen L, Wang C. In vitro and in vivo degradation behavior of aqueous-derived electrospun silk fibroin scaffolds. Polym Degrad Stab 2010;95:1679–1685. 43. Tedder M.E. Stabilized collagen scaffolds for heart valve tissue engineering. Tissue Eng Part A 2009;15:1257–1268. 44. Butcher JT, Mahler GJ, Hockaday LA. Aortic valve disease and treatment: The need for naturally engineered solutions. Adv Drug Deliv Rev 2011;63:242–268. 45. Zou L, Cao S, Kang N, Huebert RC, Shah VH. Fibronectin induces endothelial cell migration through beta1-integrin and Src dependent phosphorylation of fibroblast growth factor receptor-1 at tyrosines 653/654 and 766. J Biol Chem 2012;287:7190–7202. 46. Schopka S. Recellularization of biological heart valves with human vascular cells: In vitro hemocompatibility assessment. J Biomed Mater Res B Appl Biomater 2009;88:130–138.
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Department of Plastic and Reconstructive Surgery, Wake Forest University School of Medicine, Medical Center Blvd, Winston-Salem, North Carolina 27157 2 Virginia Tech - Wake Forest University School of Biomedical Engineering and Science, Medical Center Blvd, Winston-Salem, North Carolina 27157 Received 12 February 2014; revised 17 June 2014; accepted 25 June 2014 Published online 15 July 2014 in Wiley Online Library (wileyonlinelibrary.com). DOI: 10.1002/jbm.a.35267 Abstract: Conventional collagen-based heart valves eventually fail because of insufficient replacement of graft material by host tissue. In this study, type I collagen was blended with silk fibroin and the synthetic elastic polymer poly (glycerol–sebacate) (PGS) in varying proportions to create multifunctional electrospun nanofibrous materials tailored for use as endovascular scaffolds such as heart valve replacement. Depending on the blended material the elastic moduli ranged from 2.3 to 5.0 Mpa; tensile stresses ranged from 0.8 to 1.5 Mpa; and strains ranged from 30% to 70%. Electrospun materials with a weight ratio of 4.5:4.5:1 (collagen, fibroin, and PGS) (termed PFC mats) were the most similar to native heart valves. In vitro degradation of PFC mats was 0.01% per week.
Endothelial cells adhered to, proliferated, and formed cell– cell junctions on PFC mats. Compared with collagen hydrogels and electrospun collagen mats respectively 220–290% less platelet adhesion was observed for PFC mats. The study demonstrates that PFC material has superior mechanical properties, low degradation, and reduced thrombogenic potential and suggests that further investigation of this bioC 2014 material for cardiovascular applications is warranted. V Wiley Periodicals, Inc. J Biomed Mater Res Part A: 103A: 1150–1158, 2015.
Key Words: heart valve, heart valve scaffold, nanocomposite scaffold, polyglycerol sebacate, fibroin, silk, collagen
How to cite this article: Wang R, Levi-Polyanchenko N, Morykwas M, Argenta L, Wagner WD. 2015. Novel nanofiber-based material for endovascular scaffolds. J Biomed Mater Res Part A 2015:103A:1150–1158.
INTRODUCTION
Heart valve malfunction continues to be a major health problem, affecting one in forty adults in the United States.1 Replacement of native valves with animal derived valves or totally synthetic bioprostheses is the standard of care if the native valve cannot be repaired.2 Although numerous potential valve substitutes are available, none are optimal and all suffer from major potential complications such as degeneration, structural failure, thrombogenicity, or loosening of position.3–5 Typically, bioprosthetic collagen based valves have early valve degeneration and significant risks of postoperative failure. The failure of these materials and the thrombogenicity of mechanical devices necessitate the development of improved materials for heart valve repair and/or replacement. The purpose of this study was to fabricate an improved material composed of blended Type I collagen, silk fibroin, and poly (glycerol–sebacate) (PGS) which could serve as an extracellular matrix (ECM) mimetic and functional heart valve graft. In order to create an appropriate graft material that meets physiological requirements, a strategy of incorporating natural proteins with an elastic synthetic polymer was applied. Type I collagen was selected because it is the
most abundant and load-bearing component of valve tissue while elastin provides flexibility. Moreover, type I collagen motifs specifically bind to a2b1 and avb1 integrins through RGD, DGEA, or GFOGER peptide sequences to facilitate endothelial cell adhesion.6,7 Silk fibroin was selected to improve the strength of graft material. Interchain hydrogen bonds in silk fibroin protein assemble the polypeptide chains into a highly crystalline b-sheet conformation which imparts a slow degradation rate.8,9 Since elastin extracted from tissue is insoluble, PGS, an elastic polymer was blended with two other materials in order to substitute for native elastin.10 Moreover, PGS has been reported to support endothelial cell growth and has superior hemocompatibility over other commonly used synthetic polymers such as used poly (L-lactide-co-glycolide) (PLGA).11–13 To create a biomimetic valvular graft-blended material, electrospinning, a fabrication technique that applies high voltage power to a solution of polymer blends at a syringe tip, was applied. Fibers are generated when the electrostatic force overcomes the surface tension of the polymer solution.14 This technique was used to create an ECM analogous graft blended material which was tested physically,
Correspondence to: William D. Wagner; e-mail: [email protected] Contract grant sponsor: Department of Plastic and Reconstructive Surgery Research Funds
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mechanically, and in studies designed to investigate endothelial cell growth and thrombogenicity. MATERIALS AND METHODS
Materials Silk fibroin protein was extracted using an aqueous solvent processing method according to the published procedure with modifications.8 Raw silk was boiled in 0.02M Na2CO3 at 100 C for 30 min, rinsed twice with DDH2O, squeezed, and air dried. Fibroin was then dissolved in 5.0M CaCl2, and centrifuged at 20003g to remove contaminants. The fibroin solution was dialyzed and lyophilized. The identity and purity of fibroin was confirmed following amino acid analysis and molecular weight determination using sodium dodecyl sulphate-polyacrylamide gel electrophoresis.15–18 PGS prepolymer was synthesized from glycerol and sebacic acid following published methods.13,19 Type I Collagen from calf skin was purchased commercially (Elastin Products Corp, MO). Fabrication and characterization of the electrospun mats Electrospun solutions were prepared using type I collagen, silk fibroin, and PGS at different weight ratios of 9:0:1, 8:1:1, 4.5:4.5:1, 1:8:1, and 0:9:1 respectively, dissolved in 1, 1, 1, 3, 3, 3-hexafluoro-2-propanol (HFP). A syringe was fixed on a Baxter infusion pump (Model AS50) to eject the polymer solution at a rate of 3 mL/h. A 35 kV high voltage (Gamma High Voltage Research, Ormond Beach, FL) was applied at a distance of 20 cm between the metal collector plate (12 cm by 12 cm) and the syringe tip to generate a sufficient electrostatic force for electrospun mat formation. The electrospun fabricated mats were removed from the collector plates and heated at 120 C for 48 h followed by treatment with glutaraldehyde vapor for 24 h to crosslink the material.19,20 Electrospun mats were washed with 0.2M glycine before use.21 Images of mats taken by scanning electron microscopy (JEOL JSM-6330F) were used to measure fiber diameters with NIH Image J software. Sixteen random measurements of fiber diameters were obtained from each mat. Chemical functional groups were characterized using a Perkin-Elmer FTIR spectrophotometer. The thermal transition temperatures were measured using differential scanning calorimetry (DSC) (TA instruments, New Castle, DE) from 260 C to 300 C at an increment rate of 20 C/min (n 5 3 per sample). To measure mat porosity, the dry weight of samples was measured and samples were hydrated in 5 mL DDH2O for 30 min at room temperature on a mechanical shaker. After blotting dry, the hydrated weight was determined.22 The volume of each electrospun mat sample was obtained by dividing the weight by the density of the mat. Then, the porosity was calculated using the following equation: porosity 5 volume of water/(volume of water 1 volume of electrospun mat). The density and weight ratios of type I collagen (1.40 g/cm3), silk fibroin (1.31 g/cm3), and PGS (1.13 g/cm3) were used to calculate the densities of the composites.10,22–24
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Mechanical testing An Instron 5500R mechanical tester (Instron Corporation, Norwood, MA) with a 500 N load cell and BlueHill software was used to perform uniaxial tensile tests of the mats at an elongation rate of 10 mm/min. A dumbbell stamp (ASTM D638-IV cutting die, Pioneer-Dietecs Corporation, Norwood, MA) was used to prepare samples of each electrospun mat. Prior to analysis, all mats were hydrated in 100 mL DDH2O for 10 min. Three measurements from each type of electrospun mat were used to calculate the elastic modulus, tensile strength, and percentage elongation. A suture retention test was performed according to modifications of a published protocol.25,26 Electrospun mats were prepared 2 cm in length and 0.5 cm in width (n 5 5) and monofilament prolene (3-0 monofilament; Ethicon, Somerville, NJ) was placed 0.5 cm from the edge of each mat to form a single loop. Using a BOSE-Electroforce mechanical tester (Bose Electroforce, Eden Prairie, MN), samples were stretched at a rate of 10 mm/min until the suture completely tore through the material. Material degradation For degradation, mats (n 5 5) were air dried and stamped into 1 3 10 mm strips and conditioned in degradation solution (10 mL phosphate buffered saline (PBS) containing 0.1% sodium azide at 37 C) during the first week. At specific time intervals the material was removed, washed three times with deionized water, and weighed to determine weight loss. Starting material weight loss during 1 week of conditioning represented water weight in the sample thus weights at 1 week were used as baseline values for the degradation study. Samples from three batches of mats were tested in separate experiments. The weights of samples were measured every week over a 30-week period. The percentage weight loss was calculated from the ratio of mass change after degradation to the original mass of the scaffold according to the formula: Weight loss (%) 5 Wl/Wi 3 100%. (Wi, initial weight of the electrospun mats and Wl, weight loss of the same electrospun mats after exposure in degradation solution.) One extra sample was prepared at each time point for morphological assessment using scanning electron microscopy (SEM). Cell culture study To test cell-material interaction, PFC mats (at 4.5:4.5:1:Collagen: Fibroin: PGS weight ratio) were sterilized with ethanol for 30 min followed by ultraviolet light exposure of each side for 1 h followed by hydration in sterile PBS. To mimic the in vivo condition where blood proteins would interact with implanted material, fibronectin (Sigma, St Louis, MO) was used at a concentration of 100 mg/mL to coat mats and culture dishes. HUVECs (ATCC: CRL-1730) were seeded on the substrates in a 16-well culture plate at a cell density of 50,000 cells/well. On Days 3 and 7, rhodamine-phalloidin and sytox green dyes (Invitrogen, Eugene, OR) were used to stain F-actin and nuclei, respectively. Image J software (NIH) was used to quantitate cell numbers from representative confocal micrographs.
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TABLE I. Fiber Diameters in Mats of Different Compositions Sample Type
Fiber Diameters (nm)
Collagen : PGS (9 : 1) Collagen : Fibroin: PGS (8:1:1) Collagen : Fibroin: PGS (4.5:4.5:1) Collagen : Fibroin: PGS (1:8:1) Fibroin : PGS (9:1)
2067 6 168 4577 6 697 2952 6 240 784 6 77 694 6 43
All values represent means 6 SEM. The fiber diameters were measured from 16 randomly selected fibers of two representative scanning electron micrographs. Diameters ranged from 600 to 3000 nm. FIGURE 1. FTIR spectra of collagen, PGS, fibroin, and composites.
Platelet adhesion to mats and mats with cells R Human whole blood was drawn into a 2.7 mL BD VacutainerV Coagulation Tube (BD, Franklin Lakes, NJ) and centrifuged at 800 rpm for 15 min at 25 C with Harrier 18/80 centrifuge (Sanyo Gallenkamp, Loughborough, UK) to obtain platelet rich plasma (PRP). Three materials: PFC mats, (at weight ratio of 4.5:4.5:1 Collagen: Fibroin:PGS) Type I electrospun collagen mats and collagen hydrogels (rat tail, type I collagen, BD Biosciences, Bedford, MA) were prepared to compare platelet adhesion. PRP was diluted with PBS to 2.16 3 108 platelets/mL, and 100 mL was applied to the center of the substrates. After 15 min materials were washed twice with PBS followed by fixation in 2.5% glutaraldehyde. Factin, the contractile protein expressed in platelets, was stained with rhodamine-phalloidin to identify platelet interaction with materials. Adherent platelets on various substrates were assessed following confocal microscopy. The morphology of platelets was evaluated by SEM. Numbers of adherent platelets were quantified using NIH Image J Software. In a separate experiment, platelet adhesion was examined using materials that were preseeded with HUVECs for 24 h. The rationale for this experiment was to closely simulate an in vivo state where endothelial cells would produce a nonthromobenic glycocalyx and condition the material.
composites at high collagen blends (Collagen: Fibroin: PGS; 0:9:1, 1:8:1) closely mimicked the FTIR spectrum of collagen. This is seen through the broad peak at 3275 cm21, which is assigned to the OAH stretch and NAH stretch of collagen (Fig. 1—region 1). Amide I and II of collagen were present at 1622–1630 cm21 (Fig. 1—region 4) and 1546– 1515 cm21 (Fig. 1—region 5), respectively, while amide III can be seen at 1228–1237 cm21 (Fig. 1—region 6). At high silk fibroin blends (Collagen: Fibroin: PGS 1:8:1, 0:9:1), the composite spectra closely mimicked the FTIR spectrum of silk fibroin. Silk fibroin was found in the b sheet conformation within the composite blend based on the carbonyl stretching frequencies for amides I, II, and III at 1620 cm21, 1512 cm21, and 1230 cm21 respectively.27,28 The stretching frequency at 3280 cm21 is indicative of the ANAH stretching vibration shown as a broad peak for amide A. The peak at 1700 cm21 was assigned to be the C@O stretch in amide I b sheets, and 1225–1233 cm21 referred to the CAN stretch and CANAH bend in amide III b sheets structure.30 The stretching frequencies for the PGS polymer can be seen in the spectra of all of the composite blends. The C@O stretch for the ester can be seen at 1734 cm21 (Fig. 1—region 3) and the broad OAH stretch at 3500 cm21 represents the hydrogen bonded hydroxyl group of PGS.13 The peaks shown at
Statistical analysis Data are presented as mean 6 standard error of the mean (SEM) unless otherwise noted. Statistical significance was determined using one-way analysis of variance (ANOVA) following post hoc testing for multiple groups when appropriate. A p value of p < 0.05 was considered statistically significant. StatView 5.0 (SAS Institute, Cary, NC) was used to perform the statistical analysis. RESULTS
Composite characterization FTIR spectroscopy was used to determine the structural characteristics of the composite materials. The vibrational stretching frequencies of the pure collagen, silk fibroin and PGS polymers were found to be comparable to literature values (Fig. 1).9,13,27–31 The characteristic stretching frequencies of Type I collagen, silk fibroin, and PGS were evident in the composite materials. The FTIR spectra of the
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FIGURE 2. DSC scans of electrospun mats.
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TABLE II. Uniaxial Tensile Testing of Electrospun Mats Sample Type Collagen Collagen: PGS (9:1) Collagen: Fibroin: PGS (8:1:1) Collagen: Fibroin: PGS (4.5:4.5:1) Collagen: Fibroin: PGS (1:8:1) Fibroin: PGS (9:1)
Elastic Modulus (Mpa) a
3.67 6 0.12 2.25 6 0.16b 2.76 6 0.20b 4.11 6 0.13a,c 4.97 6 0.27d
Stress (Mpa) a,d
0.69 6 0.06 1.30 6 0.09b,c 1.10 6 0.09b 1.45 6 0.05c N/A: too brittle to be determined 0.82 6 0.09d
Strain (mm/mm) 0.23 6 0.03a d 0.62 6 0.02b 0.44 6 0.03b,c 0.41 6 0.01a,b,c 0.33 6 0.12a,c
All values represent means 6 SEM. The tensile stress, strain and elastic modulus of electrospun mats at different composite ratios were measured. Means having a similar superscript letter are not significantly different. Means not having a similar superscript letter (a,b,c,d) are significantly different (p < 0.05).
2928 cm21 and 2855 cm21 (Fig. 1—region 2) represent the sp3 CAH symmetric and antisymmetric stretches of PGS. Fiber morphology and diameter Fiber diameters ranged from 694 to 4577 nm (Table I) and showed high interconnectivity. In general, thinner and more rounded fibers were observed for the electrospun mats with higher fibroin content (0:9:1, 1:8:1 for Collagen: Fibroin: PGS weight ratios) as compared with the thicker and more flat fibers of electrospun mats with high proportions of collagen (9:0:1, 8:1:1, 4.5:4.5:1). Thermal transition analysis of electrospun mats DSC scans were analyzed to determine thermal transition temperatures of electrospun mats. By incorporating increasing amount of fibroin, a shift of thermal transition temperatures to a higher range was observed in comparison to PGS alone (226 C) (Fig. 2). The results suggest that the electrospun mats made from collagen, silk fibroin, and PGS composites would be thermally stable for in vivo application. Porosity measurements The porosity of the electrospun mats ranged from 67 to 80% which would enable diffusion of nutrients and oxygen transport.
FIGURE 3. Forces required to tear sutures from electrospun mats and procine heart valves. Duplicate samples shown by the two bars were tested for each material. Collagen: Fibroin: PGS (4.5:4.5:1) showed the highest suture pull-out strength with an average maximum load of 0.32 N as compared with a maximum average load of 0.64 N for fresh porcine heart valve.
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Mechanical tensile testing The elastic moduli of the electrospun mats (Table II) ranged from 2.25 Mpa (Collagen: PGS at a 9:1 ratio) to 4.97 MPa (Fibroin: PGS at a 9:1 ratio). Electrospun mats with collagen alone had a similar elastic modulus of 3.67 Mpa compared with collagen-based porcine valvular grafts: 3.68 MPa for fresh valves and 3.95 MPa for glutaraldehyde fixed valves.31 Addition of PGS to primarily collagen containing mats created a material with significantly higher percentage of elongation but less ultimate tensile strength. All mats produced had ultimate tensile stress values that ranged between 0.69 MPa and 1.45 MPa. Among all electrospun mats produced, the composites at 4.5:4.5:1 of collagen: silk fibroin: PGS weight ratio had the best overall mechanical strength and elasticity. The suture retention test was used to measure the maximum force required to disrupt sutures from the material. For this test, material thickness was similar to the thickness of fresh aortic valve. Collagen: silk fibroin: PGS (4.5:4.5:1) mats required the greatest force to tear the tissue with maximum suture retention force of 0.32 N. Fresh porcine heart valve had a maximum force of 0.64 N (Fig. 3). On the basis of the results of all materials and composites studied, electrospun mats containing collagen-fibroinPGS (4.5:4.5:1 weight ratio) were selected for further study. This material, termed PFC, was used to evaluate degradation, cellular compatibility, and thrombogenicity. The primary selection criteria at this point in the study was based
FIGURE 4. In vitro degradation of PFC mat during a 30-week period. Each data point is presented as the mean 6 SEM (n 5 4). [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
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FIGURE 5. Quantitation of HUVEC grown on culture dishes or PFC mats. Each bar represents the mean 6 SEM for measurements taken from representative photomicrographs (n 5 4). A significant (p < 0.05) increase in the number of HUVECs was observed from day 3 to 7 of cells on both substrates.
on mechanical properties and similarity to native fresh heart valves. Degradation of PFC mats Degradation of the PFC mats after 1 week of preconditioning was found to be only 0.3% weight loss per week over a 30-week incubation period (Fig. 4). To the unaided eye, PFC mats remained intact during the entire experiment. By SEM, nanofibers maintained a consistent morphology within the meshwork during the course of study. No significant differ-
FIGURE 7. Confocal images illustrating platelet adhesion to different substrates. PRP (1.08 3 107 platelets) was added to each well of a 48well plate. The images were acquired after 15 min of incubation on various substrate surfaces at 37 C. A: PFC mat, B: collagen mat, C: culture dish, D: collagen gel. More adhered platelets and the formation of microthrombi on the collagen gel and collagen mat are observed as compared with the culture dish and the PFC mat. Platelets were visualized by rhodamine–phalloidin staining (red) for F-actin protein (Scale bar: 50mm).
ences in fiber diameters were found over the course of the experiment (data not shown). While not part of this current study design samples were maintained under conditions of degradation for long term and at 108 weeks the change in total weight was 4% (detailed data not presented).
FIGURE 6. Confocal microscopic images illustrating HUVEC morphology and density. A: PFC mat at day 3, B: PFC mat at day 7, C: culture dish at day 3, D: culture dish at day 7. Cells were seeded at 50,000 per dish on a 48-well plate. Representative photomicrographs depict increased cell numbers from day 3 to day 7 on both substrates. Cells were stained for F-actin protein (red) using rhodamine-phalloidin and for nuclei (green) using sytox green (Scale bar: 50 mm).
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FIGURE 8. Platelet adhesion to different substrates. Adherent platelets were counted from representative photomicrographs taken from the culture dish, PFC mat, collagen mat, and collagen gel after 15 min incubation at 37 C. Each bar represents the mean 6 SEM (n 5 3). Means having a similar superscript letter are not significantly different. Means not having a similar superscript letter are significantly different (p < 0.05).
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FIGURE 9. Scanning electron micrographs of collagen gel (A and D), collagen mat (B and E) and PFC mat (C and F) after 15 min of incubation with PRP. The images demonstrate the presence of a single layer of platelets on the PFC mat and minor platelet activation in comparison to activated platelets on the collagen mat and collagen gel (Magnification: 31500 for A, B, and C, scale bar: 10 mm; 35000 for D, E, and F, scale bar: 1 mm).
Cell adhesion and proliferation Biocompatibility and nonthrombogenic property is essential for materials in contact with blood. When HUVECs were cultured on PFC mats, cell numbers increased significantly from days 3 to 7 (p < 0.05) (Fig. 5). Morphologically, an organized endothelial cell monolayer formed on PFC mats. The intense staining pattern of F-actin at the cell borders for cells cultured on PFC mats suggested the formation of tight endothelial cell junctions [Fig. 6(B)]; however, future studies will be needed to prove the formation of tight junctions. In contrast, HUVECs on culture dishes did not form uniform monolayer (Fig. 6).
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Platelet adhesion to PFC mats and mats preseeded with HUVECs The thrombogenic potential of PFC mats precultured with and without endothelial cells was accessed by examining platelet adhesions. Confocal images of platelets stained with rhodamine–phalloidin demonstrated a low level of platelet adhesion on polysterene culture dishes (Fig. 7). Single platelets or small clumps comprised of 2–3 platelets were present on PFC mats. In sharp contrast, platelets, and platelet aggregates adhered extensively to the collagen gel (positive control). Images demonstrate increased platelets numbers as well as increased sizes of platelet aggregates (Fig. 7). Platelet numbers on collagen mats, and collagen gels were
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FIGURE 10. Confocal images illustrating the association and interaction of platelets with HUVECs on different substrate. PRP (1.08 3 107 platelets) was added to each well of a 48-well plate. The images were acquired after 15 min of incubation of platelets on various substrate surfaces at 37 C. A: PFC mat, B: collagen mat, C: culture dish, D: collagen gel. Formation of microthrombi on the collagen gel and increased size of platelet aggregates was observed on the collagen gel and collagen mat as compared with the culture dish and PFC mat. F-actin (red) of cells and platelets were stained using rhodamine–phalloidin and nuclei (green) of cells were stained using sytox green (Scale bar: 50 mm).
respectively 150 and 790% higher (p < 0.05) than on the PFC mats (Fig. 8). The morphology of platelets on various substrates was examined by SEM. Extensive activation and fusion of degranulated platelets was observed on collagen gels [Fig. 9(A,D)]. Platelets on electrospun collagen mats also were activated and stellate shaped [Fig. 9(B,E)]. Platelets on PFC mats appeared to be distinctly spherical [Fig. 9(C,F)] and morphologically at minimal activation state in comparison to the appearance of platelets on collagen electrospun mats or the collagen gels. In a second experiment, platelets were applied to mats cultured for 24 h with HUVECs (50,000 cells/well in a 48well plate) (Fig. 10). Adherent platelets were observed on collagen electrospun mats as well as the formation of microthrombi of larger sizes on collagen gels. In clear contrast, platelets on PFC mats were all spherical in shape and no microthrombi were observed (Fig. 10). A 60% reduction in total adherent platelets was observed on the PFC mat as
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compared with the culture dish (Fig. 11) (p < 0.05). The number of platelets decreased 290% on PFC mats compared with electrospun collagen mats (p < 0.05), and 220% compared with collagen gels (p < 0.05).
DISCUSSION
In this study, an electrospinning fabrication technique was used to produce PFC mats for potential use as prosthetic heart valves. The electrospinning fabrication procedure is less time consuming compared with decellularized tissue processing procedures.32 As a feasible fabrication technique, electrospinning on molds could be used to recreate the geometry of heart valve tissue or for use in transcatheter heart valves procedures.33 PFC mats demonstrated superior strength and flexibility compared to primarily collagen rich biomaterials. Blending silk fibroin with type I collagen and PGS created a novel and strong material that still possessed elastic properties.
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ORIGINAL ARTICLE
FIGURE 11. Platelet interaction with cells cultured on different substrates. Numbers of adhered platelets were counted from representative photomicrographs taken from culture dish, PFC mat, collagen mat, and collagen gel substrates after 15 min incubation at 37 C. White bars (w) indicate areas of the material without endothelial cells gray bars (䊏) indicate platelet counts for areas with cells. Each bar represents the mean 6 SEM (n 5 3). Means not having a similar superscript letter are significantly different (p < 0.05).
Various ratios of type I collagen and silk fibroin were tested to mimic the mechanical and physiological properties desired for heart valve construction. Results showed electrospun mats at a 4.5:4.5:1 weight ratio of collagen: silk fibroin: PGS (PFC mat) had properties closely matching heart valve mechanical properties, an interconnective network, cell compatibility, and low thrombogenecity. In the suture retention tests, PFC mats showed sufficient suture retention, suggesting that PFC mats could be used in surgical implant procedures. A major limitation of the study includes the absence of in vivo data in an appropriate animal model on the performance of PFC as a valvular graft. Studies prior to implantation should include assessment of PFC in a tissue bioreactor where performance under increasing pressures can be evaluated. In addition, the PFC material should be tested in conjunction with devices such as catheters to be used in clinical deployment. This study suggested a new strategy in improving the strength of the graft material by composite selection in order to prevent cell-mediated buckling, a major challenge in maintaining the mechanical integrity of valvular graft.32,33 This was accomplished by incorporating silk fibroin with helical collagen and the elastic polymer PGS to obtain both strength and flexibility.34–36 The physical properties and molecular interactions demonstrated by FTIR suggest that PFC fibers have a property imparting strength and viscoelasticity to the composite material.29,37,38 PFC mats showed limited degradation with minimal weight loss during a 30-week degradation period. The ultrastructure of the nanofibers showed little change and appeared superior to the reported degradation of fully cross-linked collagen or polylactic acid electrospun constructs.9,39–41 The results of the studies on PFC mats are consistent with published in vivo degradation studies indicating that silk fibroin degrades slower than most collagen-based materials, such as polycaprolactone-collagen scaffolds.42,43
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In this study, HUVEC attached and proliferated on PFC mats. This is the first step to achieve tissue remodeling and maturation of implanted heart valves in vivo.6,7,44,45 Published studies have shown elevated thrombogenic risk associated with decellularized collagen-based heart valve grafts.46 In the present study significant reductions in platelet adhesion and platelet activation were observed on PFC mats. The observation that the PFC material had reduced platelet adherence and low activation was a surprising finding since enhanced endothelial cells affinity was observed for the material. The potential exists that the PFC composite had a molecular interaction of components that result in selective patterns of masked or exposed binding sites for endothelial cells and platelet cell surface ligands. Future studies will be required to identify the exact molecular mechanism. CONCLUSIONS
A major accomplishment of the current investigation was the development and testing of a new composite material that can be used as material for valvular surgery. Material comprising PFC mats with collagen: fibroin: PGS at a ratio of 4.5:4.5:1 demonstrated comparable mechanical strength as fresh heart valve tissue. Functional tests showed minimal weight loss and sustained nanofiber structural integrity over a 30-week degradation study. Organized endothelial monolayers formed on PFC mats and based on platelet adhesion studies were less thrombogenic compared to collagen nanofiber mats and collagen gels. This study demonstrated that the PFC composite is a promising material, with potentially superior hemocompatibility compared with structurally similar collagen materials. ACKNOWLEDGMENTS
The authors gratefully acknowledge the use of resources of the Wake Forest University Center of Nanotechnology and Molecular Materials; Dr. Roy Hantgan and Dr. Mark Lively, Wake Forest University Biochemistry Department and Protein Analysis Core Laboratory; Dr. Jessica Sparks, Miami University Chemical and Paper Engineering. The authors acknowledge Dr. Christopher MacNeill’s assistance in the scientific comments and review of the manuscript. They thank Robin Simmons for the assistance in manuscript preparation and Victoria Jiang for editorial comments REFERENCES 1. Nkomo VT, Gardin JM, Skelton TN, Gottdiener JS, Scott CG, Enriquez-Sarano M. Burden of valvular heart diseases: A population-based study. Lancet 2006;368:1005–1011. 2. Apte SS. Current developments in the tissue engineering of autologous heart valves: Moving towards clinical use. Future Cardiol 2011;7:77–97. 3. Mol A. Review article: Tissue engineering of semilunar heart valves: Current status and future developments. J Heart Valve Dis 2004;13:272–280. 4. Jordan JE, Williams JK, Lee SJ, Raghavan D, Atala A, Yoo JJ. Bioengineered self-seeding heart valves. J Thorac Cardiovasc Surg 2012;143:201–208. 5. Breuer CK. Application of tissue-engineering principles toward the development of a semilunar heart valve substitute. Tissue Eng 2004;10(11-12):1725–1736. 6. Gu X, Masters KS. Regulation of valvular interstitial cell calcification by adhesive peptide sequences. J Biomed Mater Res A 2010; 93:1620–1630.
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7. Shekaran A, Garcia AJ. Nanoscale engineering of extracellular matrix-mimetic bioadhesive surfaces and implants for tissue engineering. BBA- Gen Subjects 201;1810:350–360. 8. Rockwood DN, Preda RC, Yucel T, Wang X, Lovett ML, Kaplan DL. Materials fabrication from Bombyx mori silk fibroin. Nat Protocols 2011;6:1612–1631. 9. Horan RL, Antle K, Collette AL, Wang Y, Huang J, Moreau JE, Volloch V, Kaplan DL, Altman GH. In vitro degradation of silk fibroin. Biomaterials 2005;26:3385–3393. 10. Pomerantseva I, Krebs N, Hart A, Neville CM, Huang AY, Sundback CA. Degradation behavior of poly(glycerol sebacate). J Biomed Mater Res A 2009;91:1038–1047. 11. Motlagh D, Yang J, Lui KY, Webb AR, Ameer GA. Hemocompatibility evaluation of poly(glycerol-sebacate) in vitro for vascular tissue engineering. Biomaterials 2006;27:4315–4324. 12. Lee, KW, Stolz DB, Wang Y. Substantial expression of mature elastin in arterial constructs. Proc Natl Acad Sci USA 2011;108: 2705–2710. 13. Wang Y, Ameer GA, Sheppard BJ, Langer R. A tough biodegradable elastomer. Nat Biotech 2002;20:602–606. 14. Sell SA, Wolfe PS, Garg K, McCool JM, Rodriguez IA, Bowlin GL. The use of natural polymers in tissue engineering: a focus on electrospun extracellular matrix analogues. Polymers 2010;2:522– 553. 15. Matsumoto A, Lindsay A, Abedian B, Kaplan DL. Silk fibroin solution properties related to assembly and structure. Macromol Biosci 2008;8:1006–1018. 16. Matsumoto A, Chen J, Collette AL, Kim UJ, Altman GH, Cebe P, Kaplan DL. Mechanisms of silk fibroin sol-gel transitions. J Phys Chem 2006;110:21630–21638. 17. Lucas F, Shaw JT, Smith SG. The amino acid sequence in a fraction of the fibroin of Bombyx mori. Biochem J 1957;66:468–470. 18. Schroeder WA, Kay L, Lewis B, Munger N. The amino acid composition of bombyx mori silk fibroin and of tussah silk fibroin. J Am Chem Soc 1955;77:3908–3913. 19. Sant S, Khademhosseini A. Fabrication and characterization of tough elastomeric fibrous scaffolds for tissue engineering applications. In: Engineering in Medicine and Biology Society (EMBC), 2010 Annual International Conference of the IEEE. 2010 20. Yi F, LaVan DA. Poly(glycerol sebacate) nanofiber scaffolds by core/shell electrospinning. Macromol Biosci 2008;8:803–806. 21. Sung HW, Chen CN, Huang RN, Jsu JC, Chang WH. In vitro surface characterization of a biological patch fixed with a naturally occurring crosslinking agent. Biomaterials 2000;21:1353–1362. 22. Soliman S, Sant S, Nichol JW, Khabiry M,Traversa E, Khademhosseini A. Controlling the porosity of fibrous scaffolds by modulating the fiber diameter and packing density. J Biomed Mater Res A 2011;96:566–574. 23. Minoura N, Tsukada M, Nagura M. Fine structure and oxygen permeability of silk fibroin membrane treated with methanol. Polymer 1990;31:265–269. 24. De Cupere VM, Van Wetter J, Rouxhet PG. Nanoscale organization of collagen and mixed collagen2pluronic adsorbed layers. Langmuir 2003;19:6957–6967. 25. Trowbridge EA, Lawford PV, Crofts CE. Pericardial heterografts: A comparative study of suture pull-out and tissue strength. J Biomed Eng 1989;11:311–314. 26. Edwards MB, Draper ERC, Hand JW, Taylor KM, Young IR. Mechanical testing of human cardiac tissue: Some implications for MRI safety. J Cardiovasc Magn Reson 2005;7:835–840. 27. Um IC, Kweon HY, Park YH, Hudson S. Structural characteristics and properties of the regenerated silk fibroin prepared from formic acid. Int J Biol Macromol 2001;29:91–97.
1158
WANG ET AL.
28. Cilurzo F, Gennari CG, Selmin F, Marotta LA, Minghetti P, Montanari L. An investigation into silk fibroin conformation in composite materials intended for drug delivery. Int J Pharm 2011; 414(1–2):218–224. 29. Renugopalakrishnan V, Chandrakasan G, Moore S, Huston TB, Berney CV, Rajendra S, Bhatnagar S. Bound water in collagen. Evidence from Fourier Transform Infrared and Fourier transform infrared photoacoustic spectroscopic study. Macromolecules 1989;22:4121–4124. 30. Hayashi T, Mukamel S. Vibrational2Exciton couplings for the amide I, II, III, and A modes of peptides. J Phys Chem B 2007;111: 11032–11046. 31. Vesely I, Noseworthy R. Micromechanics of the fibrosa and the ventricularis in aortic valve leaflets. J Biomec 1992;25:101–113. 32. Cebotari S. Use of fresh decellularized allografts for pulmonary valve replacement may reduce the reoperation rate in children and young adults: Early report. Circulation 2011;124(11S):S115– S123. 33. Dijkman PE, Driessen-Mol A, Frese L, Hoerstrup SP, Baaijens FP. Decellularized homologous tissue-engineered heart valves as offthe-shelf alternatives to xeno- and homografts. Biomaterials 2012; 33:4545–4554. 34. Billiar KL, Sacks MS. Biaxial mechanical properties of the native and glutaraldehyde-treated aortic valve cusp. II. A structural constitutive model. J Biomecl Eng 2000;122:327–335. 35. Sacks MS, Schoen FJ, Mayer JE. Bioengineering challenges for heart valve tissue engineering. Annu Rev Biomed Eng 2009;11: 289–313. 36. Beun LH, Beaudoux XJ, Klein JM, de Wolf FA, Stuart MA. Selfassembly of silk-collagen-like triblock copolymers resembles a supramolecular living polymerization. ACS Nano 2011;6:133–140. 37. Jiang C, Wang X, Gunawidjaja R, Lin YH, Gupta MK, Kaplan DL, Naik RR, Tsukruk VV. Mechanical properties of robust ultrathin silk fibroin films. Adv Funct Mater 2007;17:2229–2237. 38. Zoccola M, Aluigi A, Vineis C, Tonin C, Ferrero F, Piacentino MG. Study on cast membranes and electrospun nanofibers made from keratin/fibroin blends. Biomacromolecules 2008;9:2819–2825. 39. Kim K, Yu M, Zong X, Chiu J, Fang D, Seo YS, Hsiao BS, Chu B, Hadjiargyrou M. Control of degradation rate and hydrophilicity in electrospun non-woven poly(d,l-lactide) nanofiber scaffolds for biomedical applications. Biomaterials 2003;24:4977–4985. 40. Zong X, Ran S, Kim KS, Fang D, Hsiao BS, Chu B. Structure and morphology changes during in vitro degradation of electrospun poly(glycolide-co-lactide) nanofiber membrane. Biomacromolecules 2003;4:416–423. 41. Liu T, Teng WK, Chan BP, Chew SY. Photochemical crosslinked electrospun collagen nanofibers: Synthesis, characterization and neural stem cell interactions. J Biomed Mater Res A 2010;95:276–282. 42. Zhou J, Cao C, Ma X, Hu L, Chen L, Wang C. In vitro and in vivo degradation behavior of aqueous-derived electrospun silk fibroin scaffolds. Polym Degrad Stab 2010;95:1679–1685. 43. Tedder M.E. Stabilized collagen scaffolds for heart valve tissue engineering. Tissue Eng Part A 2009;15:1257–1268. 44. Butcher JT, Mahler GJ, Hockaday LA. Aortic valve disease and treatment: The need for naturally engineered solutions. Adv Drug Deliv Rev 2011;63:242–268. 45. Zou L, Cao S, Kang N, Huebert RC, Shah VH. Fibronectin induces endothelial cell migration through beta1-integrin and Src dependent phosphorylation of fibroblast growth factor receptor-1 at tyrosines 653/654 and 766. J Biol Chem 2012;287:7190–7202. 46. Schopka S. Recellularization of biological heart valves with human vascular cells: In vitro hemocompatibility assessment. J Biomed Mater Res B Appl Biomater 2009;88:130–138.
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