Co-Electrospun Blends of PLGA, Gelatin, and ... - ACS Publications

Biomacromolecules 2011, 12, 399–408

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Co-Electrospun Blends of PLGA, Gelatin, and Elastin as Potential Nonthrombogenic Scaffolds for Vascular Tissue Engineering Jingjia Han,† Philip Lazarovici,† Colin Pomerantz,† Xuesi Chen,§ Yen Wei,‡ and Peter I. Lelkes*,† Integrated Laboratory for Cellular Tissue Engineering and Regenerative Medicine, School of Biomedical Engineering, Science and Health Systems and Department of Chemistry, Drexel University, Philadelphia, Pennsylvania, 19104, United States, and Key Laboratory of Polymer Ecomaterials, Changchun Institute of Applied Chemistry, Chinese Academy of Science, Changchun, 130022, P.R. China Received September 24, 2010; Revised Manuscript Received November 16, 2010

In search for novel biomimetic scaffolds for application in vascular tissue engineering, we evaluated a series of fibrous scaffolds prepared by coelectrospinning tertiary blends of poly(lactide-co-glycolide) (PLGA), gelatin, and elastin (PGE). By systematically varying the ratios of PLGA and gelatin, we could fine-tune fiber size and swelling upon hydration as well as the mechanical properties of the scaffolds. Of all PGE blends tested, PGE321 (PLGA, gelatin, elastin v/v/v ratios of 3:2:1) produced the smallest fiber size (317 ( 46 nm, 446 ( 69 nm once hydrated) and exhibited the highest Young’s modulus (770 ( 131 kPa) and tensile strength (130 ( 7 kPa). All PGE scaffolds supported the attachment and metabolization of human endothelial cells (ECs) and bovine aortic smooth muscle cells (SMCs) with some variances in EC morphology and cytoskeletal spreading observed at 48 h postseeding, whereas no morphologic differences were observed at confluence (day 8). The rate of metabolization of ECs, but not of SMCs, was lower than that on tissue culture plastic and depended on the specific PGE composition. Importantly, PGE scaffolds were capable of guiding the organotypic distribution of ECs and SMCs on and within the scaffolds, respectively. Moreover, the EC monolayer generated on the PGE scaffold surface was nonthrombogenic and functional, as assessed by the basal and cytokine-inducible levels of mRNA expression and amidolytic activity of tissue factor, a key player in the extrinsic clotting cascade. Taken together, our data indicate the potential application of PGE scaffolds in vascular tissue engineering.

Introduction Cardiac bypass surgery is preferentially carried out using autologous arteries or veins; however, many patients have an inadequate supply, thus requiring alternative approaches.1 Although biopolymers, such as expanded poly(tetrafluoroethylene) (e-PTFE) and Dacron, have successfully been used as large diameter (>5 mm) vascular substitutes, they are thrombogenic, rendering them ineffective in small diameter (i.e., coronary bypass) applications.1,2 Grafts made of natural materials or cellular sheets, however, have limited mechanical strength, makingtheminadequateforapplicationinthearterialcirculation.3-6 Recent attempts to tissue engineer small caliber vascular grafts via autologously derived cells are currently in clinical trials.7,8 However, even if these trials prove to be successful, the rather long time needed for producing these grafts (up to 9 months) limits their potential application in acute cases.8 Therefore, further research is required for creating nonthrombogenic offthe-shelf available small-diameter grafts. We hypothesized that mechanically appropriate and nonthrombogenic scaffolds for application as off-the-shelf available vascular grafts could be produced by blending synthetic * To whom correspondence should be addressed. Tel: 215-762-2071 (Office), 215-762-8106 (Lab). Fax: 215-895-4983. E-mail: pilelkes@ drexel.edu. † Integrated Laboratory for Cellular Tissue Engineering and Regenerative Medicine, School of Biomedical Engineering, Science and Health Systems, Drexel University. ‡ Department of Chemistry, Drexel University. § Chinese Academy of Science.

polymers with natural extracellular matrix (ECM) proteins. Collagen, the most abundant matricellular protein in the arterial ECM network, provides the primary structural framework of the blood vessel wall and harbors signaling cues for vascular cells. In some tissue engineering applications, the denatured form of collagen, gelatin, has been used as an economical substitute for collagen.9 Specifically, by electrospinning collagen out of fluoroalcohols such as HFP, collagen may be denatured into gelatin.10 Elastin, which forms an integral part of the elastic fibers in the ECM, contains critical cues that modulate vascular smooth muscle cell (SMC) activity and phenotype as well as confer elasticity to the vessel wall.11-13 Indeed, uncontrolled proliferation of SMCs in elastin knockout mice has led to arterial stenosis.14 Therefore, in manufacturing scaffolds that mimic natural ECM compositions, collagen/gelatin and elastin are two indispensable natural components that provide proper structural support and biological cues. Because attempts to engineer scaffolds purely from natural materials have typically resulted in constructs that exhibit rather weak mechanical properties,15 we have incorporated a biodegradable polymer, poly(lactic-coglycolic acid) (PLGA, a synthetic polymer that is used in many FDA-approved biomedical devices), into our blends in an effort to mechanically strengthen the construct. To develop a nonthrombogenic neointimal lining, the endothelial monolayer on our scaffolds must express negligible levels of tissue factor (TF) under basal conditions so as to avoid activation of the extrinsic coagulation pathway. To be physiologically competent, the endothelial cells (ECs) must respond to an inflammatory insult by significantly elevating TF expression.16,17

10.1021/bm101149r  2011 American Chemical Society Published on Web 12/23/2010

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In this study, we adopted a parametric approach of designing a fibrous scaffold for engineering small caliber vascular grafts with improved mechanical properties. To achieve this goal, we first studied how the physical properties of PLGA/gelatin/elastin (PGE) composite electrospun fibers could be fine-tuned by adjusting the ratios of PLGA and gelatin in the blend. Second, we investigated the effect of the varying ratios of PLGA and gelatin on vascular cell morphology and growth on and penetration into the scaffolds. Lastly, we showed that the electrospun PGE scaffolds could support a functional, nonthrombogenic endothelium. Taken together, this finding supports our notion that PGE scaffolds could be a lead candidate material for the development of small caliber vascular grafts.

Experimental Section Materials. PLGA 90/10 (w/w) was synthesized and characterized at the Changchun Institute of Applied Chemistry, Chinese Academy of Science, P.R. China.18 Gelatin (bovine skin, type B powder), R-elastin (soluble bovine, lyophilized powder), and 1,1,1,3,3,3-hexafluoro2-propanol (HFP) were purchased from Sigma (St. Louis, MO).19 Goat antihuman polyclonal VE-cadherin IgG was from Santa Cruz Biotechnology (Santa Cruz, CA). Alexa Fluoro 594 donkey antigoat IgG (H+L) was purchased from Invitrogen (Carlsbad, CA). Tumor necrosis factor-R (TNF-R, human recombinant, cat no: T0157) was from Sigma. TaqMan gene expression assays (human TF, assay ID: Hs00175225_m1 and human housekeeping gene peptidylprolyl isomerase A (PPIA), assay ID: Hs99999904_m1) were purchased from Applied Biosystems (Carlsbad, CA). Human factor VIIa (FVIIa, cat no: HCVIIA-0031) and human factor X (FX, cat no: HCX-0050) proteins were purchased from Haematologic Technologies (Essex Junction, VT). Chromogenic substrate S-2765 was purchased from Diapharma (West Chester, OH). Preparation of PGE Fibrous Matrices by Electrospinning. Electrospinning was carried out essentially as previously described.19 In brief, PLGA, gelatin, and elastin were dissolved in HFP at 10, 8, and 20%, respectively, then mixed as tertiary blends at volume ratios of 1:2:1, 2:2:1, 3:2:1, 1:3:1, 2:3:1, 3:3:1, 1:4:1, 2:4:1, and 3:4:1, correspondingly abbreviated as PGE121, PGE221, PGE321, PGE131, PGE231, PGE331, PGE141, PGE241, and PGE341. Depending on the experimental needs, three types of fibrous samples were prepared: (a) 20-50 µm thick fibrous mats on circular 12 mm glass coverslips (Fisher Scientific) for scanning electron microscopy (SEM) analysis, (b) 100-200 µm thick fibrous mats on circular 15 mm glass coverslips (Fisher Scientific) for in vitro cell culture studies, and (c) 0.5 to 1 mm thick mats on a grounded rectangular aluminum plate for mechanical testing and histology. For the latter tests, the mats were then peeled off the plate. All samples for cell culture studies were sterilized by exposure on both sides to a conventional ultraviolet (UV) source (30 W, 65 cm working distance) for 20 min each in a laminar flow hood and were secured in 24-well culture plates using Viton O-rings (ColeParmer, Vernon Hills, IL) before cell seeding.20 Scanning Electron Microscopy Analysis. Both dry and hydrated samples were sputter-coated with platinum/palladium (Pt/Pd) and then visualized with a SEM (XL-30 Environmental SEM-FEG or Zeiss Supra 50VP).19 For preparation of hydrated samples, electrospun PGE fibers on circular glass coverslips were soaked in distilled and deionized H2O (dd H2O) for 2 h, then briefly dried on a hot plate (70-80 °C, 2-5 min). Average fiber diameters were calculated using the SEM-internal dedicated software or UTHSCSA ImageTool 3.0 (http://ddsdx.uthscsa.edu/dig/download.html). Pore sizes of hydrated samples were calculated by image analysis using Image J 1.43 (http://rsbweb.nih.gov/ ij/download.html) following published procedures.21 Average and standard deviation were calculated from 20 random measurements per image and from at least 8 SEM images for each sample.21 Mechanical Tensile Test. The tensile properties of electrospun PGE mats upon hydration were characterized with an Instron System (model 5564, Norwood, MA) using routine tensile testing.22 The dimensions

Han et al. of rectangular PGE mats (15 mm × 5 mm × (0.5 to 1) mm, L × W × H) were measured with a digital caliper after soaking in 1× phosphatebuffered saline (PBS, Mediatech, Herndon, VA) for 2 h. A 10 Newton load cell and strain speed of 1 mm/min were used. Young’s modulus was calculated from the slope of the initial linear segment of the stress-strain curve by segmental linear regression. Cell Culture. Human EA.hy926 ECs and bovine aortic SMCs (passages 6-11) were both maintained in Dulbecco’s modified Eagle’s medium (DMEM) with 4.5 g/L glucose supplemented with 10% fetal bovine serum (FBS, Hyclone, Logan, UT), 0.5% PenStrep (10 000 IU/ mL penicillin, 10 000 µg/mL streptomycin solution, Mediatech, Manassas, VA) under standard culture conditions (37 °C, 5% CO2).23,24 The culture medium was changed every 3 days. PGE Scaffold-Vascular Cell Interactions. Immunofluorescence. For each cell type, cells were seeded onto the electrospun scaffolds at a density of 5000 cells/cm2 and maintained in complete DMEM for 48 h or until confluence at day 8. Constructs were fixed in 10% buffered formalin for 1 h at room temperature (RT) and permeabilized with 0.25% Triton-X100 (Sigma) in 1X PBS. Constructs were then incubated for 15 min with 2 µg/mL Hoechst 33258 (Bisbenzimide, Sigma) and 1 µg/mL rhodamine-phalloidin (Phalloidin-tetramethylrhodamine B isothiocyanate, Sigma), staining nuclei and actin microfilaments, respectively. VE-cadherin staining was carried out on EC-seeded constructs at day 8. Following permeabilization, nonspecific staining was blocked with 1% bovine serum albumin (BSA, g96%, Sigma) in 1X PBS for 2 h. The samples were incubated overnight with goat antihuman polyclonal VE-cadherin IgG (2 µg/mL) at 4 °C followed by Alexa Fluoro 594 donkey antigoat IgG (H+L) (1 ug/mL) for 1 h at RT as the secondary antibody. Constructs were then mounted in VECTASHIELD mounting medium with DAPI as nuclear counterstain (Vector Laboratories Inc., Burlingame, CA). Negative controls were carried out in parallel omitting the primary antibody. All constructs were visualized with a Leica DMRX upright microscope equipped with appropriate fluorescence filters. Digital images were acquired with a Leica DC 300FX camera (Leica, Wetzlar, Germany).25 For immunofluorescence quantification, 100-200 cells in random fields were analyzed from the fluorescent micrographs using Image J software. AlamarBlue Assay. The alamarBlue (AB) assay measures metabolic activity of cells in culture.26 Cell metabolization can then be translated into cell numbers, following proper calibration and adequate controls.26,27 For this study, cell metabolization was evaluated continually using the AB assay (Biosource, Alameda, CA), as previously described.19 ECs/ SMCs were seeded at 5000 cells/cm2 separately onto PGE fiber-coated glass coverslips and tissue-culture-treated polystyrene (TCP, control). Following overnight incubation, supernatants were removed and replaced with 1 mL of fresh complete medium containing 5% (v/v) AB in each well. After a 4 h incubation, triplicate 200 µL aliquots of AB-containing medium were removed from each well for fluorescence readings. The remaining AB-containing medium was replaced with 1 mL of fresh complete DMEM, and the cells were returned to the incubator. Continual assessment of AB fluorescence was performed on the same cell population every other day for up to 8 days until the cells reached confluence. All data were normalized to AB “fluorescence blank” readings of AB-containing medium in the absence of cells for each time point. The initial AB fluorescence readings taken after overnight incubation were defined as day 0 for each substrate. Histology. The PGE scaffolds (5 mm × 5 mm × (0.5 to 1) mm) were placed in a six-well culture plate on an orbital shaker and seeded overnight with a suspension of either ECs at 500 000 cells/mL or SMCs at 2 000 000 cells/mL, respectively. The samples were harvested at day 8. Constructs were fixed in 10% formalin for 1 h at RT and kept overnight at 4 °C. After several washes with 1× PBS, samples were embedded in paraffin and sectioned transversely using a Leica RM2255 microtome at 5 µm thickness, followed by hematoxylin and eosin (H&E) staining. Tissue Factor Expression and Amidolytic Activity. TF mRNA expression and a chromogenic assay for TF activity were used to assess

Co-Electrospun Blends of PLGA, Gelatin, and Elastin


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Table 1. Fiber Diameters, Percent of Swelling and Pore Area of the Electrospun PGE Fibersa components

relative concentration by weights PLGA (%)

gelatin (%)

fiber diameter (nm) dry

hydrated

percent of swelling (%)

pore area (µm2)

384 ( 22 384 ( 51 317 ( 46b

825 ( 35 873 ( 50 446 ( 69c

115 127 41

2.57 ( 0.22 2.03 ( 0.14 0.57 ( 0.04c

903 ( 24 1196 ( 79 756 ( 58c

1001 ( 30 1735 ( 103 822 ( 74c

11 45 8

1.99 ( 0.11 2.78 ( 0.26 3.46 ( 0.46c

939 ( 15 825 ( 47 792 ( 41

1174 ( 22 1364 ( 34 775 ( 30c

25 65 -2

1.92 ( 0.23 2.13 ( 0.18 4.74 ( 0.43c

elastin (%)

PGE121 PGE221 PGE321

22 36 45

35 28 25

43 36 30


18.5 31.25 41

44.5 37.5 32

37 31.25 27


16 28 37

52 44 39

32 28 24

a Note: All data are expressed as mean ( SE, n ) 30 for fiber diameter, n ) 160 for pore area. 121, PGE221, PGE321).

the thrombogenicity and functionality of EC monolayers on various scaffolds. To compare the role of PGE blend composition versus the three-dimensionality of the fibrous PGE scaffolds on TF expression, we tested ECs on both 15 mm glass coverslips covered with various electrospun fibrous PGE scaffolds (as above) and similar circular glass coverslips that were dip-coated with identical PGE solutions (diluted at 15% v/v). Following dip-coating, the coverslips were dried overnight on a hot plate at 37 °C in a chemical hood to ensure the complete evaporation of HFP solvent. ECs were seeded on both systems at 100000/cm2 to obtain a confluent monolayer overnight. At 24 h postseeding, the medium was exchanged for basal DMEM medium without FBS and with or without10 ng/mL TNF-R to assess the effect of inflammatory cytokines on TF mRNA expression and TF activity in cultured ECs. Cells were harvested after 4 h for assaying TF mRNA levels or after 5 h for measuring TF activity. Real-Time QuantitatiVe PCR. Total RNA was isolated with TRIzol reagent (Invitrogen) after the samples were washed with ice-cold 1× PBS twice; the extracted RNA was analyzed spectrophotometrically with a Nanodrop 1000 (Thermo Fisher Scientific). Total RNA (600 ng) was reverse transcribed with a high capacity cDNA reverse transcription kit (Applied Biosystems). Real-time PCR was performed using a Stratagene Mx3000P QPCR system (Agilent Technologies, Santa Clara, CA). A 10 µL reaction mixture containing 1 µL of cDNA from each sample was mixed with 3.5 µL of nuclease-free water (Applied Biosystems), 0.5 µL of target gene/housekeeping gene, and 5 µL of TaqMan universal PCR master mix (Applied Biosystems). A comparative Ct method was used for analyzing the expression level of human TF, which was normalized to the Ct value of human housekeeping gene PPIA. Fold of change was then calculated as 2-∆∆Ct. Tissue Factor ActiVity Assay. The activity of TF on the surface of ECs was assessed by a two-step chromogenic, amidolytic assay.28,29 In brief, as part of the extrinsic coagulation pathway, TF combines with FVIIa and forms the FVIIa-TF complex that is capable of activating FX to FXa, which in this assay then cleaves a specific chromogenic substrate, S-2765. Upon termination of the experiment, the EC monolayers were washed twice with ice-cold 1× PBS and incubated at 37 °C for 1 h on an orbital shaker with 150 µL of a mixture of FVIIa (10 nM) and FX (175 nM) in HEPES buffer (10 mM HEPES, 5 mM CaCl2, 1 mg/mL BSA). Subsequently, 50 µL of the incubation mixture was transferred in duplicate to a 96-well plate together with 50 µL of TBS/EDTA buffer (1 mg/mL BSA, 20 mmol/L EDTA, pH 8.5) to terminate generation of FXa. Then, 100 µL of S-2765 substrate (1 mM stock solution) was added to each well, and the optical density was measured at 405 nm by kinetic reading in a microplate reader at 37 °C for 30 min. Negative controls were carried out in parallel in the absence of FVIIa. In our study, the rates of changes in optical readings were linear within the initial 30 min for all samples. TF activities were expressed as changes in optical density (∆O.D.405nm) per minute. Statistical Analysis. All experiments were repeated at least three times in triplicate unless otherwise mentioned. Data are expressed as mean ( SE when applicable. Linear regression analysis, Student’s t

b

P < 0.05. c P < 0.01 within each group (e.g., PGE

test, and single factor ANOVA were used for parameter estimation and hypothesis testing. Data were considered to be statistically significant when *: p < 0.05 and **: p < 0.01.

Results and Discussion Co-electrospining blends of synthetic PLGA and natural collagen/gelatin and elastin out of an “environment-friendly” and inexpensive solvent, which is capable of dissolving both types of components, remains a challenge. As a strong organic solvent with high polarity, HFP is still being regarded as an ideal solvent for this application because of its volatility, miscibility with water and many organic solvents, as well as its low surface tension.30 In our previous study, we described novel electrospun scaffolds made of tertiary blends of PGE focusing on morphology, fiber diameter, and mechanical properties under both dry and hydrated conditions as well as cell adhesion and proliferation of H9c2 myoblasts and bone marrow stromal cells.19 Here we expanded the PGE series and tested how, by parametrically adjusting the ratios of two of the PGE components (PLGA and gelatin) and, hence, by fine-tuning the physical properties of the ensuing blends, we could modulate the interaction of such scaffolds with vascular cells, with the ultimate aim to generate a nonthrombogenic endothelium. Morphology, Diameter, and Pore Size of PGE Fibers. Whereas fiber sizes for each individual blend were rather uniform (standard errors generally did not exceed (10%, Table 1), we were able to achieve a wide range of fiber diameters depending on the various PGE compositions ranging from 320 to 1200 nm in the dry state and from 450 to 1750 nm when hydrated (Figure 1, Table 1). Whereas the size of the fibers varied for each PGE composition (Table 1), we also observed some differences in the pore size of various hydrated PGE blends, which were in the range of 0.6-4.7 µm2. Such small pores would normally not permit penetration of cells into fibrous scaffolds, made of stiff synthetic polymers, such as PLGA.19 In contrast, the elastic nature of PGE fibers/scaffolds will permit highly migratory cells, such as smooth muscle cells, to penetrate through the pores and enter the scaffolds, as seen in Figure 5 for PGE 341 and also previously described for the interaction of rat cardiac myoblasts and rat bone marrow stromal cells with PGE 312.19 In extending our previous studies,19 we found that for all of these tertiary blends, no chemical cross-linking was required to obtain stable fibrous constructs upon hydration (Figure 1D-F). In contrast, binary blends of natural and synthetic polymers, such as gelatin and polyaniline, did not yield stable structures upon hydration and require cross-linking.20 As previously reported,19 PGE fibers are homogeneous rather than

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Figure 1. SEM micrographs of PGE fibrous scaffolds: (A) dry PGE321, (B) hydrated PGE321, (C) macroscopic image of hydrated PGE321 fibrous mat, (D) hydrated PGE321, (E) hydrated PGE331, and (F) hydrated PGE341. Scale bar ) 5 µm in parts A, B, and D-F and 1 cm in part C.

a mixture of individual PLGA, gelatin, and elastin ones, which is in contrast with the findings by Boland et al.,15 who report the heterogeneous deposition of coelectrospun collagen and elastin fibers. Hydrated PGE mats from all blends exhibited a macroscopic appearance resembling opaque hydrogels. (A representative image for PGE 321 is shown in Figure 1C.) When examined by SEM, we noticed an increase in fiber diameter and an increasingly flattened fiber morphology as the volume fraction of gelatin increased (e.g., PGE321, PGE331, and PGE341 in Figure 1). Individual fiber diameters changed differentially upon hydration. Specifically, by comparison with their dry counterparts, the diameter of hydrated PGE221 fibers more than doubled in size, whereas the diameter of hydrated PGE341 fibers remained essentially unchanged (Table 1). Within each group, the lowest degree of hydration-induced swelling was observed in fibers with the highest PLGA content, for example, in the group of PGE121, PGE221 and PGE321, PGE321; that is, fibers with the highest content of PLGA (45%) and the lowest content of gelatin (25%) swelled the least upon hydration (by ∼41%). Furthermore, the same PGE321 blend yielded the smallest average fiber diameters (317 ( 80 nm when dry and 446 ( 121 nm upon hydration, see Table 1). PGE141, which has the highest content of gelatin (52%) and the lowest content of PLGA (16%), yielded fibers of intermediate diameter in both dry and hydrated conditions. In contrast, PGE121 (with an elastin content of ∼43%) and PGE341 (with an elastin content of ∼24%), which yield significantly different fiber diameters in the dry state (PGE121 with small diameter and PGE341 with intermediate diameter), produced fibers with similar intermediate diameters (∼800 nm) upon hydration (Table 1). When correlating the content of each component material with the percent of swelling or the resultant fiber diameter under both dry and hydrated conditions, no strong correlation was found (R2 < 0.5 by linear regression, data not shown), suggesting that the individual fiber diameter and the percent of swelling of resultant PGE scaffolds may not be directly correlated with

the relative content of individual materials. Ju et al.21 reported that increasing the fiber diameters of their electrospun PCL-collagen scaffolds significantly increased pore sizes in the scaffolds. However, in this study, we did not find a statistically significant correlation between these two parameters when evaluating the entire set of our nine blends (R2 < 0.5 by linear regression, data not shown). We believe that this discrepancy might stem from differences in our experimental settings, that is, varied blend compositions. In the study by Ju et al.,21 the ratios of their various binary PCL-collagen blends were consistently kept at 1:1, whereas we varied the relative ratio of PLGA and gelatin from 1 to 3 in our tertiary PGE blends. In addition, Ju et al.21 measured the pore areas in the scaffolds in the dry state, which does not represent the real situation for cell infiltration because of the swelling and junction fusion of fibers when the scaffolds are hydrated. In support of our results, Powell et al.31 also reported that the swelling as well as partial degradation of hydrated gelatin fibers have led to significant loss of porosity. Whereas the porosity of most electrospun scaffolds is high and usually exceeds 80%,32 the nature of the electrospinning process, that is, the sequential layerTable 2. Tensile Properties of the Electrospun PGE Matsa components

Young’s modulus (kPa)

tensile strength (kPa)


134 ( 51 427 ( 41 770 ( 131b

16 ( 7 102 ( 26 130 ( 7b


85 ( 47 130 ( 15 358 ( 19b

9(5 40 ( 5 90 ( 14b


3(1 129 ( 5 150 ( 15b

5(3 43 ( 8 65 ( 10b

a Note: All data are expressed as mean ( SE (n ) 3). b P < 0.01 with each group.



403

Figure 2. Correlation of Young’s modulus and tensile strength of the hydrated PGE fibrous mats with relative concentrations (w/v) of PLGA and gelatin.

by layer deposition of the fibers, dictates that the porosity of electrospun scaffolds will decrease with increasing thickness of the fibrous mats. Taken together, these data suggest that in addition to the fiber diameter and composition, other factors such as the percent of swelling, the (random) orientation of fibers, and the thickness of the fibrous mats might affect both pore size and porosity of electrospun scaffolds. Tensile Properties of PGE Mats. Of the nine blends, PGE321 (with the highest content of PLGA) produced the highest Young’s modulus (770 ( 131 kPa) and the highest tensile strength (130 ( 7 kPa), whereas PGE141 (with the highest content of gelatin) yielded the lowest values for both Young’s modulus (3 ( 1 kPa) and tensile strength (5 ( 3 kPa) (Table 2). The Young’s moduli of most PGE scaffolds are well in the range of those of natural arteries,33,34 whereas the tensile strengths are significantly lower than those of natural coronary arteries (1.4-11.14 MPa),35 which may be due to the absence of vascular cells in our scaffolds at the time of mechanical testing and also explained by the fact that we tested isotropic sheets and not anisotropic tubes. Lee et al.36 reported that their electrospun fibers of collagen, elastin, and PLGA blends are characterized by a Young’s modulus of 0.85 MPa (our PGE321 is 0.77 MPa) and a tensile strength of 0.37 MPa. (Our PGE321 is 0.13 MPa.) The differences between our and their data can be explained by differences in the morphology of the scaffolds (we tested mats, they tested tubes), the blend composition (45% collagen, 15% elastin, and 40% PLGA for Lee et al., whereas our PGE321 was composed of 25% gelatin, 30% elastin, and 45% PLGA), and the composition of PLGA that was used by Lee et al. (50/50 w/w) and by us (90/10 w/w). We noticed that the mechanical properties of various PGE scaffolds in the current study were lower than those in our previous study,19 in which we used PGE at ratios of 312 and 222. We suggest that this difference might be due to different blend compositions, which favor PLGA’s major contribution to the mechanical properties. Figure 2 presents a correlation of the relative contents of PLGA and gelatin with the Young’s moduli and tensile strengths of resultant composite fibrous mats. As seen in Figure 2A,C, a reasonable positive correlation exists between the Young’s modulus/tensile strength of PGE scaffolds and the relative PLGA content. Furthermore, we found a statistically more significant negative correlation between gelatin contents and mechanical properties (R2 ) 0.6625 and R2 ) 0.8962 for Young’s modulus

and tensile strength, respectively). Surprisingly, we found little or no statistically significant correlation between elastin contents and mechanical properties of the scaffolds (not shown). We surmise that the low stiffness of elastin led to its negligible effect on the mechanical properties of the composite fibers and that PLGA and gelatin may cojointly control the tensile properties of the resultant fibrous scaffolds.37 Therefore, by adjusting the relative ratio of these two components, we can effectively modulate the mechanical properties of electrospun PGE scaffolds. Biocompatibility of PGE Scaffolds. To confirm the biocompatibility of PGE scaffolds, we investigated the interactions of ECs and SMCs with our scaffolds. In line with William’s Dictionary of Biomaterials,38 we define biocompatibility in the context of our studies as the ability of our scaffolds to support appropriate vascular cell adhesion, migration, metabolization, and differentiation in conjunction with promoting histotypic organization, that is, the formation of an EC monolayer on the scaffold surface and of a multilayered assembly of the SMCs in the scaffold interior. To assess the effects of the scaffold composition on EC and SMC morphology, we seeded the cells onto the various PGE scaffolds. Histotypic morphology was evaluated by fluorescence microscopy 48 h and 8 days post seeding. Differences in EC morphology were observed 48 h after seeding. As shown in Figure 3A, ECs cultured on PGE131 (Young’s modulus ) ∼85 kPa, relative concentration of PLGA ) 0.185, hydrated fiber diameter ∼1000 nm, Tables 1 and 2) appeared more elongated, whereas on PGE121 (Young’s modulus ) ∼134 kPa, relative concentration of PLGA ) 0.22, hydrated fiber diameter ∼800 nm), that is, a slightly “stiffer” scaffold with smaller fibers containing less of the natural proteins, the cells remained more rounded (Figure 3A). ECs cultured on rigid glass surfaces (Young’s modulus >1 GPa) appeared more “typical” EC: parsley-shaped with a significantly bigger cell size (Figure 3). In quantifying these visual observations, we found that ECs cultured for 48 h on PGE121, PGE131, and glass were significantly different from each other in terms of cell area and shape factor (Figure 3C,D). However, upon confluence at day 8, the EC monolayers displayed similar morphology, regardless of which scaffold they were cultured on (Figure 3A, lower row). This histotypic EC monolayer configuration on various PGE scaffolds, characterized by the nonoverlapping appearance of nuclei and perimembranal assembly of actin microfilaments

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Figure 3. (A) Morphology of EA.hy926 endothelial cells on PGE fibrous scaffolds or glass coverslips at 48 h postseeding and at confluence (day 8). Staining for nuclei-hoechst (blue) and actin cytoskeleton-phalloidin (red). (B) Immuofluorescence staining for monolayer formation on PGE121. Staining for nuclei-DAPI and intercellular junctions-VE-cadherin. (C) Cell area analysis 48 h postseeding (cell area expressed as number of pixels). (D) Shape factor analysis 48 h postseeding (shape factor: 0 ) line, 1 ) circle). (E) Morphology of bovine aortic smooth muscle cells on PGE fibrous scaffolds. Staining as in panel A. (C,D) Data are expressed as mean ( SE, representative from three experiments, n ) 90. **: P < 0.01. Scale bar ) 50 µm.

(Figure 3A, lower row), was also confirmed by VE-cadherin staining (Figure 3B). Similar to ECs, SMCs attached and spread on all PGE scaffolds by 48 h postseeding. However, in contrast with the ECs, there was no morphological difference on SMCs after 48 h or 8 days (Figure 4 and data not shown). Moreover, 8 days postseeding, the SMCs had invaded the scaffolds and formed multiple cell layers, as evidenced by their overlapping appearance of nuclei and microfilaments (Figure 3E). In summary, the differences in blend composition, fiber diameters, and mechanical properties of PGE scaffolds are mainly manifested in terms of the initial adhesion and spreading of the ECs. Importantly, all PGE scaffolds are “biocompatible” and facilitate the histotypic culture of ECs and SMCs.

AB metabolic activities of ECs and SMCs cultured on various PGE fibrous matrices are presented in Figure 4. As assessed by AB fluorescence, the metabolic activities of both EC and SMC significantly increased over an 8 day time course. For all PGE compositions, the metabolization of EC was approximately 50-70% lower on the scaffolds than on TCP. The EC metabolization on the scaffolds seemed to be independent of the PGE blend composition, suggesting that it might be the three-dimensionality of the fibrous substrates rather than the blend composition that might play a key role in affecting EC metabolization in vitro. As previously noted, after 8 days, ECs formed confluent monolayers on all PGE scaffolds. (See Figure 3B.) In contrast, SMCs displayed a similar rate of metabolization on all PGE scaffolds, which was essentially identical to that on



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Figure 4. (A) EA.hy926 endothelial cell and (B) bovine aortic smooth muscle cell metabolic activity on various substrates over an 8 day time course by alamarBlue (AB) assay. *: P < 0.05, **: P < 0.01.

Figure 5. Histological analysis (H&E staining) of cell-seeded PGE scaffolds: (A) lined (arrow) with a confluent monolayer of endothelial cells (EA.hy926) and (B) scaffold interior populated with bovine aortic smooth muscle cells. Scale bar ) 50 µm.

TCP. We surmise that the 3-D microenvironment of the fibrous scaffolds in concert with the PGE blend composition might facilitate SMC metabolization and penetration into the scaffold, similar to the organotypic multi-SMC-layers in the tunica media of blood vessel. Our data suggest that during the 8 day study, PGE scaffolds support the metabolization and organotypic arrangement of both ECs and SMCs. In line with the above microscopic visualization, histological analysis of cell-seeded 3-D PGE mats indicated that after 8 days postseeding, ECs formed a monolayer on the surface of electrospun PGE scaffolds without migrating into the fibrous meshwork (Figure 5A), whereas the SMCs penetrated rather uniformly into the scaffold, forming a multilayered tissue-like structure reminiscent of SMC organization in the tunic media (Figure 5B). Ju et al.21 recently designed bilayered poly(εcaprolactone) (PCL)-collagen scaffolds in which they expanded pore size in the outer layer via increased fiber diameter in an effort to aid in the infiltration of SMCs. In comparison, all PGE scaffolds, composed of submicrometer-sized fibers, facilitate the penetration of SMCs. This indicates that PGE scaffolds are innately suitable in terms of guiding the organotypic distribution and sorting out of vascular cells. Tissue Factor Expression and TF Amidolytic Activity of EC monolayers. For these studies, we focused on comparing TF expression and activity in EC monolayers seeded after 24 h on PGE121, PGE321, and PGE141 because of the representative compositions as well as distinct mechanical properties of those scaffolds (e.g., PGE141 has the highest content of gelatin, the lowest Young’s modulus, and the lowest tensile strength; see Table 1). Basal levels of TF mRNA in ECs cultured on PGE scaffolds were somewhat higher than those on TCP (Figure 6A). However, this increase did not lead to higher basal amidolytic activities, as reflected in the rates of optical reading changes

(Figure 6B). In contrast, basal TF mRNA and activity levels on dip-coated coverslips were identical to those on TCP (Figure 6A,B). Whereas previous studies already observed that some electrospun scaffolds exhibited low thrombogenic potential with low TF release from EC cultured on them,39 this study is, to the best of our knowledge, the first one to investigate differential TF expression on 3-D electrospun fibrous scaffolds and to compare it with TF expression on 2-D coated surfaces. On the basis of our results, we expect that in an in vivo setting, endothelialized PGE grafts will not exhibit elevated TF activity and hence will not activate the extrinsic coagulation cascade. To test the functionality of the EC monolayers, we stimulated the cells with a paradigmatic inflammatory cytokine, TNF-R. Upon stimulation with 10 ng/mL TNF-R, the TF levels of EC monolayers cultured on various substrates significantly increased and were comparable to those on TCPs (Figure 6C,D), suggesting a physiological behavior of EC monolayer on PGE. Notably, upon stimulation with TNF-R, TF mRNA levels on all substrates increased by ∼50 fold, whereas TF protein activity was elevated ∼2 fold. This observation is in line with data reported by Xie et al.,40 who also found a ∼2 to 3 fold increase in TF activity in EA.hy926 cells upon stimulation with 10 ng/ mL TNF-R. We believe that this discrepancy between elevated TF mRNA levels and protein activity may be a result of the reported stability of TF mRNA in ECs stimulated with TNF-R, suggesting differential control of TF protein at the transcriptional and translational levels.41 Meng et al.42 noticed reduced TF secretion from ECs cultured on electrospun fibrous meshes of multiwalled carbon naotubes/ polyurethane composites in comparison with those on solutioncast films, implying a preferred preservation of the anticoagulant ability of ECs cultured on fibrous scaffolds. Zhang et al.43 recently reported low platelet adhesion in vitro to ECs cultured

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Figure 6. Effect of PGE blend composition on endothelial cell tissue factor expression and activity: (A) basal tissue factor mRNA expression, (B) basal tissue factor activity,(C) TNF-R-induced (10 ng/mL, 4 h) tissue factor mRNA expression, and (D) TNF-R-induced (10 ng/mL, 5 h) tissue factor activity in EA.hy926 monolayers. Data are normalized to the basal expression of TF mRNA on corresponding substrates in part C. Tissue factor activities are expressed in rates of changes in optical readings in parts B and D. *: P < 0.05 versus TCP, n.s.: not significant. Table 3. Comparison of Electrospun Hybrid Fibrous Matrices for Vascular Tissue Engineering components

PGE321 (this study)

PCL/collagen/elastin21,36,46

silk/PEO tube47,48,50

fiber size (µm) Young’s modulus (MPa) tensile strength (MPa) vascular cell adhesion and growth vascular cell distribution low TF activity

0.32 0.77 0.13 yes EC monolayer; SMC penetration yes

0.52 0.34 0.31 yes EC monolayer; SMC infiltration N/Ta

0.38 2.45 2.42 yes EC monolayer; SMC infiltration N/T

a

N/T: not tested.

on gelatin, elastin, PCL, and poliglecaprone blended electrospun fibrous scaffolds, further suggesting the usefulness of hybrid fibers as nonthrombogenic scaffolds in vascular tissue engineering. We conclude that PGE scaffolds might be suitable candidates for supporting a nonthrombogenic and physiologically competent endothelial monolayer. We want to stress the importance of elastin in our blends for enhancing endothelial adhesion and proliferation as well as for the low thrombogenicity of the composite biomaterial. Recently, Wise et al.44 reported that a multilayered elastin/PCL hybrid vascular graft, which possesses mechanical properties comparable to those of the internal mammary artery, has significantly reduced plasma clotting time and platelet adhesion. The same group also described the ability of tropoelastin/elastin to increase the expansion of hematopoietic stem and progenitor cells via proper elastic signaling.45 This latter finding might be useful for future use of PGE as base material for tubular vascular grafts in vivo and their endothelialization in situ via hematopoietic stem and progenitor cells. A number of recent studies suggest that electrospun hybrid biomaterials, made by blending natural and synthetic components, might be well-suited for applications in vascular engi-

neering based on their mechanical properties as well as in terms of the adhesion, metabolization, and organotypic distribution of vascular cells.15,21,36,46,47 Our PGE scaffolds belong to this category. As summarized in Table 3, the tensile properties of PGE scaffolds are in the same order of magnitude as the PCL-collagen-elastin scaffolds described by Lee et al.36 Using a blend of PLGA-collagen-elastin,36 the scaffolds developed by these authors yielded a Young’s modulus (0.85 MPa), which is similar to our PGE321 (0.77 MPa). Both of these values are in good agreement with the Young’s moduli of natural arteries (e.g., 0.7 to 0.8 MPa in the left common carotid artery).33 These findings support the notion that PLGA is a useful synthetic “base-material” of choice, which will improve the mechanical properties of fibrous constructs comprising natural ECM proteins. According to the data by Lee et al.,36,46 Soffer et al.,48 and Zhang et al.,47 both the Young’s modulus and tensile strength of tubular fibrous scaffolds are 10-20 time higher than those of fibrous mats made from the same materials. In our future work, we will electrospin optimized PGE blends into tubular structures, similar to the vascular grafts recently described by our group.49 We anticipate that our PGE vascular grafts will exhibit improved tensile strengths while preserving


suitable Young’s moduli. A comparison of various similar constructs (Table 3) indicates that PGE scaffolds, like scaffolds made of PCL-collagen-elastin and silk-PEO,21,47 support vascular cell growth and organotypic distribution. While providing sufficient strength, we are cognizant of the fact that silk as a base material will need significant modifications for enhanced long-term tissue biocompatibility.50,51 Uniquely among this group, ours is the first study to evaluate TF expression and activity, that is, demonstrating that distinct PGE blends promote the formation of a functional, nonthrombogenic endothelium. Taken together, we suggest that electrospun scaffolds made of PGE blends might hold promise for application in vascular tissue engineering. For our future graft design, we might need to come up with a multilayered PGE tube to better mimic that of natural coronary artery and to analyze such PGE tubes both in vitro and in vivo, as also discussed in the recent study by Chung et al.52 For example, McClure et al.53 designed a trilayered PCL/collagen/ elastin graft using three different compositions to achieve the best combination of materials. Similarly, Caves et al.54 incorporated synthetic collagen microfibers to enhance a multilamellar graft of elastin-like network, by which they could fine-tune fiber orientation and mechanical properties of the grafts. Finally, our laboratory recently designed a complex vascular graft composed of a micropatterned lumenal surface for enhanced endothelial adhesion and shear resistance and an electrospun outer layer to endow the graft with compliance very similar to that of natural arteries.49 In this context, we envisage designing a multilayered PGE graft comprised of different PGE blends with proper alignment of PGE fibers.

Conclusions In this study, we tested the hypothesis that the mechanical properties of PGE scaffolds composed of composite electrospun fibers could be fine-tuned by varying the relative ratios of the individual ingredients. Among the nine PGE blends tested, the elasticity and Young’s modulus of PGE321 make this particular blend most suitable for future applications as vascular grafts. Hence, we plan to focus our future studies primarily on this particular blend when designing and testing tubular PGE grafts under an in vivo setting. Other notable outcomes of our study are the innate capability of PGE fibrous scaffolds to guide the organotypic distribution of vascular cells and the ability to support the establishment of a nonthrombogenic endothelial lining. Our studies emphasize the usefulness of electrospinning as a platform technology for engineering scaffolds. The versatility of this fabrication process enhances its potential for use in vascular tissue engineering and provides the tools for generating off-the-shelf available vascular grafts for acute implantation. Acknowledgment. This work is partially supported by a Calhoun Fellowship and a Provost’s Fellowship (J.H.) from Drexel University and a translational grant by the Wallace H. Coulter Foundation (P.I.L.). We are grateful to Dr. Karen I. Winey, Department of Materials Science and Engineering, University of Pennsylvania for the use of the Instron system and Dr. Cora J. Edgell, University of North Carolina, Chapel Hill for the generous gift of EA.hy926 EC line.

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