The effect of thick fibers and large pores of electrospun poly(ε-caprolactone) vascular grafts on macrophage polarization and arterial regeneration.

Biomaterials xxx (2014) 1e11

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The effect of thick fibers and large pores of electrospun poly(ε-caprolactone) vascular grafts on macrophage polarization and arterial regeneration Zhihong Wang a,1, Yun Cui a,1, Jianing Wang a, Xiaohu Yang b, Yifan Wu a, Kai Wang a, Xuan Gao c, Dong Li a, Yuejie Li c, Xi-Long Zheng d, Yan Zhu b, Deling Kong a, c, **, Qiang Zhao a, * a State Key Laboratory of Medicinal Chemical Biology, Key Laboratory of Bioactive Materials, Ministry of Education, Collaborative Innovation Center of Chemical Science and Engineering (Tianjin), Nankai University, Tianjin 300071, China b Tianjin State Key Laboratory of Modern Chinese Medicine, Tianjin University of Traditional Chinese Medicine, Tianjin, China c Institute of Biomedical Engineering, Chinese Academy of Medical Science, Tianjin 300192, China d Department of Biochemistry & Molecular Biology, Libin Cardiovascular Institute of Alberta, Faculty of Medicine, University of Calgary, Calgary, Alberta, Canada T2N 4N1

a r t i c l e i n f o

a b s t r a c t

Article history: Received 1 March 2014 Accepted 27 March 2014 Available online xxx

The vascular grafts prepared by electrospinning often have relatively small pores, which limit cell infiltration into the grafts and hinder the regeneration and remodeling of the grafts into neoarteries. To overcome this problem, macroporous electrospun polycaprolactone (PCL) scaffolds with thicker fibers (5 e6 mm) and larger pores (w30 mm) were fabricated in the present study. In vitro cell culture indicated that macrophages cultured on thicker-fiber scaffolds tended to polarize into the immunomodulatory and tissue remodeling (M2) phenotype, while those cultured on thinner-fiber scaffolds expressed proinflammatory (M1) phenotype. In vivo implantation by replacing rat abdominal aorta was performed and followed up for 7, 14, 28 and 100 d. The results demonstrated that the macroporous grafts markedly enhanced cell infiltration and extracellular matrix (ECM) secretion. All grafts showed satisfactory patency for up to 100 days. At day 100, the endothelium coverage was complete, and the regenerated smooth muscle layer was correctly organized with abundant ECM similar to those in the native arteries. More importantly, the regenerated arteries demonstrated contractile response to adrenaline and acetylcholine-induced relaxation. Analysis of the cellularization process revealed that the thicker-fiber scaffolds induced a large number of M2 macrophages to infiltrate into the graft wall. These macrophages further promoted cellular infiltration and vascularization. In conclusion, the present study confirmed that the scaffold structure can regulate macrophage phenotype. Our thicker-fiber electrospun PCL vascular grafts could enhance the vascular regeneration and remodeling process by mediating macrophage polarization into M2 phenotype, suggesting that our constructs may be a promising cell-free vascular graft candidate and are worthy for further in vivo evaluation. Ó 2014 Elsevier Ltd. All rights reserved.

Keywords: Vascular grafts Macrophage polarization Cellularization Electrospinning PCL

1. Introduction

* Corresponding author. Institute of Molecular Biology, Nankai University, Tianjin 300071, China. Tel.: þ86 22 23501229; fax: þ86 22 23498775. ** Corresponding author. Institute of Biomedical Engineering, Baidi Road 236, Nankai District, Tianjin 300192, China. Tel.: þ86 22 87893696; fax: þ86 22 87893696. E-mail addresses: [email protected], [email protected] (D. Kong), [email protected] (Q. Zhao). 1 These authors contributed equally to this work.

Small-diameter vascular grafts (SMVGs) (D < 6 mm) are increasingly needed in the clinic for coronary disease and hemodialysis. Unfortunately, the acute thrombosis and subsequent occlusion often lead to the failure of the transplantation when commercialized poly(ethylene terephthalate) (Dacron) or expanded poly(tetrafluoroethylene)(e-PTFE) vascular grafts were used. In this regard, development of vascular grafts with slow degradation and controlled regenerative process has become a new concept and

http://dx.doi.org/10.1016/j.biomaterials.2014.03.078 0142-9612/Ó 2014 Elsevier Ltd. All rights reserved.

Please cite this article in press as: Wang Z, et al., The effect of thick fibers and large pores of electrospun poly(ε-caprolactone) vascular grafts on macrophage polarization and arterial regeneration, Biomaterials (2014), http://dx.doi.org/10.1016/j.biomaterials.2014.03.078

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Z. Wang et al. / Biomaterials xxx (2014) 1e11

Scheme 1. This schematic illustrates the hypothesis of the present study. Thicker-fiber grafts with larger pores may modulate the polarization of macrophages into M2 phenotype, which secret wound-healing cytokines, enhance cell infiltration, vascularization and tissue remodeling. However, the thinner-fiber grafts with smaller pores maintain the macrophages in M1 phenotype, which secret pro-inflammatory factors and play a negative role in tissue regeneration.

direction. These grafts provide a favorable environment for the recruitment of autogenous vascular cells. After a full degradation of the polymer scaffold, “neo-artery” could be generated [1]. The degradable electrospun polycaprolactone (PCL) vascular grafts have been evaluated in rat aorta implantation model [2e4] and the patency of the grafts was followed for up to 18 months [5]. The PCL grafts showed excellent structural integrity and patency. Cell invasion and neovascularization of the graft wall rapidly increased until 6 months; however, at 12 and 18 months, regression of cell number and capillary density and severe calcification were observed within the graft wall. It was probably due to the slow degradation, lower compliance and dense fibrous structure of electrospun PCL that failed to regenerate a viable arterial wall with vasomotricity and elasticity. Recently, it was reported that a fast degradable vascular graft prepared from elastic poly (glycerol sebacate) (PGS) successfully regenerated into functional new artery within 90 d [6]. The fast degradation, elastic property and large pores had led to rapid host remodeling and integration. Cellularization of the vascular graft, also known as engraftment, is a key factor for the tissue regeneration and remolding [7]. In most cases, the vascular grafts are fabricated by electrospinning [8e10], which is a versatile processing technique to provide non-woven fibrous structure. However, one major problem associated with electrospinning is the small pores, which often result in insufficient cell infiltration. To solve this problem, various techniques have been tested to increase pore size and overall porosity, including the salt/ polymer leaching [11,12], modification of collector apparatus, posttreatment by laser radiation [13]. In addition, adjusting the electrospinning parameters can control the scaffold porosity. It has been reported that the pore size is closely related to the fiber diameter of electrospun mats, which means that the pore size increases with the increase of the fiber diameter [14,15]. During electrospinning, the fiber diameter can be readily controlled by changing the parameters, such as the concentration of the polymer solution, voltage, solvent, etc. [16,17].

Inflammatory response has received increasing attention in the tissue regeneration especially the vasculogenesis. Breuer group has demonstrated that the main role of mesenchymal stem cells (MSCs) pre-seeded on the vascular graft is the secretion of monocyte chemoattractant protein-1 (MCP-1) to recruit the monocyte/ macrophage via a paracrine mechanism [18,19]. In addition to biological signals, physical cues including the structure and morphology of the scaffolds have been accepted as an important factor influencing the phenotypes of the macrophages (M1, proinflammatory phenotype; M2, the immunomodulatory and tissue remodeling phenotype), which further regulate the remodeling of wound healing and regeneration [20]. In general, transition of macrophage polarization into M2 is a promising strategy to promote the remodeling of vascular grafts. Ratner group has demonstrated that the pore size around 30e40 mm increased vascularization and reduced fibrotic response due to the shift in macrophage phenotype toward the M2 state [21]. Bowlin group investigated the effects of fiber and pore size of the electrospun mats on the macrophage polarization in cell culture, and found that the fiber with larger diameter favored the polarization of macrophages into the M2 phenotype [22]. We hypothesized that increasing the fiber diameter of electrospun vascular grafts can increase the pore size and improve the cell infiltration and migration. Meanwhile, the thicker fibers and larger pores may help modulate the phenotype and functions of macrophages, and mediate the regeneration process of vascular grafts. Scheme 1 illustrates how these two factors work cooperatively to promote tissue regeneration and remolding.

2. Materials and methods 2.1. Materials Poly(ε-caprolactone) (PCL) pellets (Mn 80,000) were purchased from Sigma (St Louis, USA). Methanol, chloroform and alcohol were obtained from Tianjin Chemical Reagent Company (Tianjin, China). Triton X-100 was purchased from Alfa Aesar (London, England). Wistar rats (male, weight 280e320 g) were purchased from the


Z. Wang et al. / Biomaterials xxx (2014) 1e11 Laboratory Animal Center of the Academy of Military Medical Sciences (Beijing, China). 2.2. Graft fabrication The scaffolds were made by electrospinning using a setup previously described in detail [23]. In brief, a polymer solution was ejected at a continuous rate using a syringe pump through a stainless steel needle (2 mm i.d.). A high voltage was applied to the needle with a variable high-voltage power supplier. Polymer fibers were collected on a grounded rotating stainless steel mandrel mounted on a home-made stand. The thicker-fiber tubular scaffolds were fabricated with the following conditions: 25% w/v PCL in CHCl3/MeOH (5:1, v/v), needleecollector distance 17 cm, flow rate 8 ml/h and voltage 11 kV with a 21-G needle. These conditions were modified as follows to make the thinner-fiber scaffolds: 10% w/v PCL in CHCl3/MeOH (5:1, v/v), needleecollector distance 20 cm, flow rate 2 ml/h and voltage 18 kV. The obtained grafts were placed in vacuum overnight to remove the residual solvent. Prior to implantation, the grafts were sterilized by immersing in 75% ethanol for 30 min, and then exposed to UV light overnight. 2.3. Fiber and pore size measurements The average fiber diameter was measured based on scanning electron microscopy (SEM) images. For each sample, five SEM images were analyzed, and at least 70 fibers were manually measured on each image and analyzed using ImageJ software (NIH USA, 2008). Results are expressed as mean standard deviation. Scaffold porosity was measured with a liquid intrusion method [22]. Dry grafts were weighed and then immersed in 100% ethanol overnight for complete wetting. Grafts were then gently wiped to remove excess ethanol and weighed again. Graft porosity was calculated with the following equations: VEtOH ¼ (MwetMdry)/QEtOH, VPCL ¼ Mdry/QPCL and Scaffold porosity ¼ VEtOH/(VPCL þ VEtOH) 100% (VEtOH is the volume of ethanol entrapped in the graft pores, Mdry and Mwet are the dry and wet weights of the grafts, QPCL is the density of PCL (1.145 g mm3) and QEtOH is the density of ethanol (0.789 mg mm3)). Pore sizes were analyzed by both manual measurement and theoretical calculations based on average fiber diameter and scaffold porosity. From SEM images (magnification 1500), pore sizes were measured by manually fitting an ellipse in representative pores formed by fibers in the same plane. The size of each pore was the average between the long and short diameters of the fitted ellipse. At least 30 pores per SEM image were measured and five images were included per scaffold sample. Results are given as mean standard deviation. Theoretical estimation of pore sizes was done according to the model established by Tomadakis and Sotirchos [24]. In the simplified form of the model, a direct relationship among total porosity (i), average fiber diameter (u) and characteristic pore size for randomly oriented overlapping fiber structures is established: pore diameter ¼ u/ln(i). 2.4. Mechanical testing Longitudinal mechanical properties were measured on a tensile-testing machine with a load capacity of 100 N (M1600, Schenk AG, Germany). Grafts with 1 cm in length were clamped with a 1 cm inter-clamp distance and pulled longitudinally at a rate of 10 mm/min until rupture. Maximum stress and strain at rupture were measured. The Young’s modulus representing the elasticity was obtained by measuring the slope of the stressestrain curve in the elastic region. Transverse tensile mechanical test of the grafts before and after implantation was performed on a Testometric tensile tester (M350, UK) (n ¼ 3) at a crosshead speed of 0.1 mm/s. A custom-made mounting frame was designed to mount the tubular structures on the tensile tester. 2.5. Culture of RAW264.7 cells on the scaffolds RAW264.7 cells were cultured in DMEM with 10% fetal bovine serum (FBS). For cell infiltration assay, RAW264.7 cells (1.25 105 cells) were seeded on ethanol disinfected scaffolds (10 mm discs) with either large or small pores. After 72 h, fluorescein diacetate (FDA) was used to stain the cells on the grafts, followed by observation under fluorescence microscope. The morphologies of the cells on the grafts were analyzed by SEM. 2.6. Real-time PCR for gene expression analysis 5 105 cells were seeded on the surface of scaffolds (30 mm discs). After 72 h, RNA was isolated from the cells using TRIzol reagent (Invitrogen). The concentration of RNA was determined by spectrophotometry (NanoDrop ND-1000, U.S). Reverse transcription was performed using the Reverse Transcription System (Promega, U.S). Quantitative PCR (Q-PCR) assay was done using LightCycler 480 SYBR green (Roche) and analyzed using Q-PCR instrument (Bio-Rad, U.S). Cells that were cultured on tissue culture plates were used as control. Gene expression levels were calculated using the 2DDCt method. Experiments were done at least three times with at least three biological replicates in each. Primers used in Q-PCR are listed in Table S3 in Supporting Information, including proinflammatory mediators (iNOs, TNF-a, IL-6),

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anti-inflammatory mediator (Arg, FiZZ1), and wound-healing mediators including vascular endothelial growth factor (VEGF), transforming growth factor-b1 (TGF-b1), basic fibroblast growth factor (bFGF), macrophage inflammatory protein-1 alpha (MIP-1a) and MCP-1. 2.7. In vivo implantation in rats The use of experimental animals was approved by the Animal Experiments Ethical Committee of Nankai University and carried out in conformity with the Guide for Care and Use of Laboratory Animals. In brief, rats were anesthetized with chloral hydrate (300 mg/kg) by an intraperitoneal injection. Heparin (100 units/kg) was administered for anticoagulation by tail vein injection before surgery. A midline laparotomy incision was then performed and the abdominal aorta was isolated, clamped, and transected. The tubular PCL grafts (2.0 mm in inner diameter and 1.0 cm in length) were sewed in an end-to-end fashion with 8e10 interrupted stitches using 9-0 monofilament nylon sutures (Yuan Hong, Shanghai, China). No anticoagulation drug was administered to the rats after surgery. Before sacrifice, rats were anesthetized with a chloral hydrate. The patency of the grafts was tested by Skin Ultrasound (MD-300, MEDA, Tianjin, China) at the predetermined time points (7, 14, 28 and 100 d). Then animals were sacrificed by injection of overdose chloral hydrate. Grafts were explanted, rinsed with saline and cut into two parts from the middle. One part was fixed with 4% paraformaldehyde and embedded in OCT for frozen cross-section. The other part was longitudinally cut into two pieces. One piece was first observed by stereomicroscope, and then fixed with paraformaldehyde for longitudinal section. The other piece was dehydrated with gradient ethanol for SEM examination. 2.8. Assessment of vascular function Aortic ring bioassay was performed as previously described [25]. After the grafts were dissected and cleaned from connective and fat tissues, the rings of 3 mm in length were obtained and bathed in the standard Krebs buffer (composition in mM: NaCl, 118.4; KCl, 4.7; CaCl2, 2.5; MgSO4, 1.2; KH2PO4, 1.2; NaHCO3, 25; dextrose, 11.1; Na2Ca EDTA, 0.029; pH 7.4) at 37 C and gassed with carbogenic mixture (95% O2 and 5% CO2). All preparations were stabilized under a resting tension of 2.00 g for 1 h with the buffer changed every 15 min. The presence of functional smooth muscle cells was indicated by the contractile responses induced by adding KCL (60 mM). The functionality of the neo-endothelium was confirmed by the relaxation by acetylcholine (Ach, 10 mM) in pre-constricted segments by adrenaline (AD, 1 mM). Isometric forces were recorded with force transducers connected to a PowerLab/870 Eight-channel 100 kHz A/D converter (AD Instruments, Sydney, Australia). Results were obtained from 6 individual rings derived from the explants from 3 animals. 2.9. Histological analysis of the explanted vascular grafts To prepare samples for SEM, the explants were rinsed with PBS and fixed with 2.5% glutaraldehyde overnight, and dehydrated in ascending series of ethanol. Samples were affixed onto aluminum stubs with carbon tape, sputter-coated with gold, and observed by SEM (Quanta 200, Czech). For sectioning and staining, the explants were fixed with 4% paraformaldehyde, dehydrated by 30% sucrose solution until the grafts sank to the bottom. After embedded in OCT, the explants were cryosectioned to 6 mm in thickness. Subsequently, the sections were stained with hematoxylin and eosin (H&E), Masson’s trichrome, and Verhoeffevan Gieson (VVG). Images were observed under an inverted microscope (Nikon Eclipse TE2000-U Kanagawa, Japan) and analyzed by Nikon NIS Elements software. To perform immunofluorescent staining, the frozen sections were fixed in cold acetone for 10 min, air-dried, and rinsed once with 0.01 mM PBS. Then slides were incubated in 5% normal goat serum (Zhongshan Golden bridge Biotechnology, China) for 45 min at 4 C. For intracellular antigen staining, 0.1% Triton-PBS was used to permeate the membrane before incubation with serum. Then the sections were incubated with primary antibodies in PBS overnight at 4 C, followed by incubation with secondary antibody in PBS for 2 h at room temperature. The nuclei were counterstained with 4,6-diamidino-2-phenylindole (DAPI) containing mounting solution (Dapi Fluoromount G, Southern Biotech, England). Endothelial cell staining was performed using rabbit anti-von Willebrand factor (vWF, 1:200, Dako, USA) primary antibody. The mural cells were stained using mouse anti-a-SMA (a-SMA, 1:100, Boster, China) and mouse anti-smooth muscle myosin heavy chain I (MYH, 1:200, Abcam, USA) primary antibody. To observe inflammatory cells in the explanted grafts, anti-CD68 antibody (CD68, 1:100, Abcam, USA) was used. To visualize M2 macrophages, anti-Mannose Receptor antibody (CD206, 1:200, Abcam, USA) was used. Elastin was stained using rabbit polyclonal anti-rat elastin (1:2000, Abcam, USA) primary antibody. Alexa Fluor 488 goat anti-mouse IgG (1:200, Invitrogen) and goat anti-rabbit IgG (1:200, Invitrogen) were used as the secondary antibodies, respectively. The sections without incubation with primary antibodies were used as negative controls. Slides were observed under a fluorescence microscope (Zeiss Axio Imager Z1, Germany).


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2.10. Statistical analysis All quantitative results were obtained from at least three samples for analysis. Data were expressed as the mean standard deviation. A two-tailed paired Student’s t-test was used to compare the differences. A value of p < 0.05 was considered to be statistically significant.

3. Results

Table 1 Quantitative measurement of the structural parameters of the electrospun PCL grafts. Measurement

Thinner-fiber grafts

Fiber size Pore size Porosity Radius

0.69 4.66 65.86 1.65

0.54 1.53 12.89% 1.30

Thicker-fiber grafts 5.59 40.88 83.22 30.46

0.67 13.67 1.76% 3.67

3.1. Graft fabrication and characterization Two types of electrospun mats with different fiber diameters were prepared by changing the electrospinning conditions. The fiber morphology and dimension were characterized by SEM. Results showed that the fibers had a smooth surface with welldefined fiber morphology (Fig. 1A,B). The averaged diameter was about 0.69 0.54 mm for the thinner-fiber electrospun mat; whereas it increased to 5.59 0.67 mm for the thicker-fiber mat. Due to the increase of fiber diameter, the pore size was markedly increased by almost 8-fold, and the porosity also showed marked increase in the thicker-fiber mat (Table 1). The tubular grafts with thinner or thicker fibers were fabricated under the conditions similar to electrospun mats fabrication. The inner-diameter of the grafts was 2.0 mm and the wall thickness is about 400e500 mm. The cross-section of the tubular grafts demonstrated homogeneous fiber distribution, and the thickerfiber grafts showed much larger pores than the thinner-fiber grafts (Fig. 1DeI).

Table 2 The mechanical properties of the electrospun PCL vascular grafts. Measurement

Thinner-fiber grafts

Young’s modulus (MPa) Maximum stress (MPa) Maximum load (N) Strain at rupture (%)

17.44 13.35 9.35 168.40

0.91 1.47 1.70 8.76

Thicker-fiber grafts 21.00 8.72 5.11 639.20

1.39 0.84 0.95 24.15

3.2. Mechanical properties The mechanical properties of two types of tubular grafts were characterized by tensile testing, and the typical stressestrain curves were shown in Fig. 1C. The mechanical performances of two grafts were significantly different in terms of elongation. The corresponding value of thicker-fiber grafts was about 3 times higher than the thinner-fiber grafts, suggesting that the toughness of

Fig. 1. Structure and mechanical property of electrospun PCL. SEM images of electrospun PCL mats with thicker fibers (A) and thinner fibers (B). Cross-sections of tubular thickerfiber grafts (DeF) and thinner-fiber grafts (GeI). The representative tensile strain-stress curve was shown in C.



thicker-fiber grafts was much greater than the thinner-fiber grafts. Besides, the increased fiber diameter resulted in a moderate increase in Young’s modulus, but a decrease in tensile stress (Table 2). 3.3. In vitro macrophage polarization on the grafts RAW264.7 cells were seeded on the surfaces of the two types of electrospun PCL mats. SEM images showed that cells plainly spreading on the surface of the thinner-fiber PCL mat, few of them infiltrated into the depth of the mat after culture for 3 d (Fig. 2A). In contrast, in the thicker-fiber PCL mat cells attached on individual fibers in 3 dimensions (Fig. 2B). Fluorescent images from FDA labeling confirmed the SEM observation (Fig. 2C,D). To investigate the effects of the structure of electrospun PCL mats on the polarization of macrophage, the related gene expression was detected using real-time PCR. Data showed that the mRNA expression of IL-6, TNF-a and MIF-1, which are related to M1 phenotype, was evidently elevated in cells cultured on the thinnerfiber mat (10% PCL) at day 3 (Fig. 2E). Whereas, the genes associated with M2 phenotype, including ArgI and FiZZ1, were highly expressed in cells on the thicker-fiber mat (25% PCL) (Fig. 2F). These findings suggest that thicker-fiber scaffolds could stimulate the macrophage polarization into M2 phenotype, while, thinner-fiber scaffolds directed the macrophage into M1 phenotype. Similar results were reported by Bowlin group [22]. In addition, the expression of MCP-1 and matrix metalloproteinase 2 (MMP2) genes was increased on both types of mats. The mRNA levels of pro-angiogenic growth factors, including VEGF and bFGF, were not significantly increased. 3.4. Patency of the implanted vascular grafts Grafts were implanted into rat abdominal aorta to replace a segment of native aorta. At specified times (day 7, 14, 28, and 100), the patency of the implanted grafts was examined by ultrasound. Results indicated that most of the grafts remained patent (Fig. 3A and Fig. S1) with only one (one out of thirteen) occluded.

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Stereomicroscopic images showed that the implanted grafts were integrated well with the native vessels. The luminal surface was clean and free of platelet aggregates or thrombi (Fig. 3B). The inner diameters calculated from the ultrasound test showed no significant change at each time points, indicating that no obvious restenosis occurred (Fig. 3C).

3.5. Cellular infiltration Cell infiltration was detected by H&E and DAPI staining. In the thinner-fiber grafts, cells only slightly infiltrated into the grafts from both the luminal and abluminal sides of the grafts, no cells were observed within the central graft wall at day 28 (Fig. 3D,E). In contrast, the macroporous structure remarkably enhanced cell penetration into the depth of the grafts (Fig. 3F,G). The nucleated cells were homogeneously distributed within the graft wall at all time points in the thicker-fiber grafts (Fig. 3H). In order to identify the type of infiltrated cells, cross-sections of explanted vascular grafts were analyzed by immunofluorescent staining using a-SMA and MYH antibodies. a-SMAþ cells mainly distributed in the graft wall near the surrounding connective tissue at day 7 (Fig. 4A), but migrated into the luminal side at day 14. While at day 28, a-SMAþ cells were organized into a thin media layer. This layer progressively increased its thickness at day 100. MYHþ cells were not observed at day 7, and there were a few at day 14. MYHþ cells became abundant and formed a media layer at day 28 and 100, indicating the regeneration of contractile smooth muscle layer (Fig. 4B). CD68þ macrophages were identified in the graft wall with uniform distribution. The number of CD68þ cells continuously increased over time (Fig. 4C). The phenotype of macrophages was discerned by CD206 antibody, which detects the M2 macrophage. CD206þ cells were observed within the surrounding tissue at day 7 and 14 with few cells found within the graft wall (Fig. 4D). At day 28 and 100, the number of CD206þ cells markedly increased. These cells were uniformly located within the entire graft wall.

Fig. 2. Evaluation of cell infiltration and macrophage polarization by in vitro culture of Raw264 cells. SEM images show the attachment on the thinner-fiber PCL mat (A) and the thicker-fiber PCL mat (B). Fluorescent images show the infiltration and distribution of the cells within the thinner-fiber mat (C) and the thicker-fiber mat (D). The gene expression of macrophages was detected using real-time PCR. E, the expression of M1 macrophage related genes; (F), the expression of M2 macrophage related genes.


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Fig. 3. Evaluation of the tubular PCL grafts in vivo. A, The ultrasound image of the thicker-fiber graft at day 100; B, the luminal morphology of the thicker-fiber graft at day 100 was observed under stereomicroscope; C, the inner diameter of the grafts was calculated based on ultrasound measurement (n ¼ 4). H&E staining show cellularization of the thinnerfiber (D) and thicker-fiber grafts (F) at day 28; cellularization of the thinner-fiber (E) and thicker-fiber grafts (G) was also confirmed by DAPI staining. The density of cells within the thicker-fiber grafts was calculated based on DAPI staining (n ¼ 4) (H).

3.6. Vascular tissue regeneration and remodeling Rate of endothelialization of the graft was examined by SEM. Endothelial cells with cobblestone-like morphology were nicely organized along the blood stream. At day 14, the cobblestone-like endothelial cells were only observed near the suture area (Fig. 5B). Although a large number of cells could also be observed from sample sections that were approximately a quarter of the graft length upstream from the suture area, almost all of these cells remained undifferentiated as no specific morphology could be observed, which may be indicative of an early stage of endothelialization (Fig. 5C). Further along the graft, hardly any cell could be found from samples that were near the halfway point of the implant, where the fibrous structure of the graft material could still be clearly visualized (Fig. 5D). At day 28 however, the endothelial coverage has already extended significantly compared to day 14, reaching approximately a half of the graft length (Fig. 5E). After 100 days, as expected, complete endothelialization was achieved, covering the entire lumenal side of the graft (Fig. 5E). Immunofluorescent staining with vWF antibody further confirmed the formation of endothelium and its similarity to that of the native artery (Fig. 6A,B) MYH-positive cells were well organized into circumferential layers, indicating the formation of arterial media at day 100 (Fig. 6C,D). 3.7. Extracellular matrix secretion Extracellular matrix is an important aspect of vascular remodeling process. In order to detect the histology structure of the

explanted grafts and the ECM secretion, histological staining including H&E, Masson, and VVG staining were performed. Data showed that a neo-tissue layer in the lumen of the graft was formed at day 100 (Fig. 6G, H). In addition, abundant collagen was observed in the graft (Fig. 6I, J). At the same time, some fibrous elastin could also be identified, which was organized between the smooth muscle cell layers (Fig 6K, L). Immunofluorescent staining showed that the elastin aligned circumferentially in a pattern like that in native arteries (Fig. 6E, F). These observations indicated that the graft was trending to remodel into the native vessels. 3.8. Biomechanical properties and physiological function The mechanical properties of explanted grafts were evaluated by the transverse tensile test, and the representative stressestrain curves were shown in Fig. S3. The mechanical properties (tensile strength and elongation) did not show significant difference before and after implantation for 100 d. This was because the slow degradation of PCL matrix, although ECM was secreted during the vascular regeneration, was not enough to evidently alter the mechanical properties. Standard dual pin wire myography was performed to examine the physiological functions of the regenerated endothelium and smooth muscle layer. Among the three tested explants, two displayed apparent vascular function of contraction and relaxation (Fig. 7 and Fig. S2A), indicating that these explants were responsive to various vasoactivators. The grafts constricted in response to the vasomotor agonists, KCL and AD. In addition, the grafts displayed vasodilation to acetylcholine (Ach), suggesting the presence of



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Fig. 4. Cellular distribution and changes during the remodeling process. A, Distribution and organization of a-SMAþ cells (green) at each time-point; B, MYHþ cells; C, CD68þ cells; D, CD206þ cells. Cell nuclei in all immunofluorescent micrographs were counterstained by DAPI (blue). Scale bar: 200 mm. (For interpretation of the references to colour in this figure legend, the reader is referred to the web version of this article.)

functional endothelium and the responsiveness of smooth muscle. Note that the magnitude of contractile responses in explanted grafts were less than those in the native aorta, which could be due to the slow degradation of PCL matrix so as to restrict the deformation of regenerated vascular tissue. However, one of the three explants failed to display detectable contraction and relaxation (Fig. S2B). MYH immunofluorescent staining of the cross-section showed that there was much less smooth muscle regeneration compared to the other two explants (Fig. S2C, D). 4. Discussion Cell infiltration is the initial step with a critical role in subsequent regeneration of the vascular grafts, such as vascularization, ECM secretion and deposition [26]. Limited cell infiltration often prevents the regeneration and remodeling progress of vascular grafts in the long term [5]. To solve this problem, two strategies have been attempted to date. One is to focus on rapid degradation of the graft. The space generated during the scaffold degradation promotes cell infiltration and vascular tissue regeneration. Wang group utilized a fast-degradable poly (glycerol sebacate) (PGS) to fabricate the vascular grafts. The grafts could be completely degraded within 2 weeks in vivo, leading to sufficient infiltration, almost complete host remodeling and good integration within 3 months [6]. Although the results were a very promising in rat aorta implantation, there have been still several issues to be considered if such grafts are considered for clinical use. It was reported that the regeneration capability of humans is much slower than that of rodent animals [27]. If cell migration and synthesis

of ECM are not substantial enough to withstand the arterial pressure, rapid degradation of the vascular grafts may result in loss of the mechanical strength and subsequent aneurysms, rupture, and probably the death [28]. For this reason, we choose PCL to prepare the grafts because of its relatively slow degradation rate. There are several methods to create macropores within the grafts by using porogen [29,30] or post-modification [31]. The present study has used a simple method to increase the pose size by optimizing the fiber thickness. Our data demonstrated that cells could easily infiltrate into the grafts and then uniformly distributed within the graft wall right after implantation. Unlike the regular electrospun PCL grafts with small pores [5], there was no cellular regression occurring after 100 d. Inflammation is also an important physiological response during the implantation of foreign materials or devices. It has been widely accepted that macrophages are an important source of both inflammatory and anti-inflammatory signals. These signals arrive in mammalian wounds 48e96 h after injury, where they clear dead cells, release proinflammatory cytokines, and subsequently produce the factors that reduce inflammation and stimulate angiogenesis and fibroblast migration and replication [32]. Rapid early monocyte recruitment has been shown to be intricately involved in postnatal blood vessel formation [33]. In particular, successful arteriogenesis is dependent on MCP-1induced monocyte recruitment to preexistent collateral arterioles [34]. A previous report has demonstrated that inflammation mediates the transformation process of tissue-engineered vascular grafts into functional neovessels [18]. Inflammation, especially by the macrophages, plays a critical role in neovessel formation and the development of stenosis in tissue-engineered vascular grafts [19].


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Fig. 5. SEM images show the endothelialization process of the thicker-fiber PCL grafts over time in rat aorta implantation. The explants were longitudinally cut into two pieces. A, graft of day 14; E, graft of day 28; F, graft of day 100. B, C and D are three selected areas in A with higher magnification.

Additionally, macrophages demonstrate various phenotypes, including M1, M2a, M2b and M2c, of which M2 phenotype plays a positive role in tissue repairing and wound-healing processes [35]. Bowlin group reported the effects of the structure of electrospun scaffolds (including fiber and pore dimensions) on modulating macrophage polarization as shown in their in vitro experiment, assuming that naïve macrophages preferably acquire a functional M2 phenotype when in contact with scaffolds having larger pores and (or) thicker fibers [22]. Our results have confirmed this viewpoint from the in vivo aspect. The pore size of our thicker-fiber mat was around 30 mm, which is in agreement with the previous report that pore size between 30 and 40 mm is most favorable for the angiogenesis and reduces fibrotic response [21]. We observed the change of M2 macrophage (CD206þ cells) during the regeneration process of the vascular grafts. These cells were observed in the surround tissue at day 7, and then progressively migrated into the graft wall with increase in number and accompanied with the formation of new capillaries (Fig. S4). Artery is typically composed of three concentric layers, endothelium, tunica media and adventitia. The endothelium plays critical roles in preventing thrombogenesis and keeping the artery patent [36]. The tunica media, composed by multiple layers of smooth muscle cells, contributes to the tensile strength, compliance, and vasoactive response of the artery [37]. Up to date, most of the reported work paid much attention to rapid endothelialization of the vascular grafts, but not enough

attention to the regeneration of functional tunica media. The previous reports by us and others have shown that the electropsun PCL grafts with thinner fibers and smaller pores could not facilitate the regeneration of functional smooth muscle layer [23,5]. However, our present data showed that the regenerated smooth muscle layer is structurally close to the native tunica media. The regenerated smooth muscle tissue consists of several layers of MYHþ cells which are well organized within the elastin network, indicating that the smooth muscle cells have contractile phenotype and the vascular grafts is in the tissue remodeling stage after implantation for 100 d (Fig. 7). More importantly, we used tissue bioassay to demonstrated that the explanted grafts have contractile response to vasoactivators and relaxation response to acetylcholine, further confirming the successful regeneration of functional tunica media and endothelium in the thicker-fiber grafts. The origin of the infiltrated cells within the graft wall is another issue which has been debated for decades. Previously, it was believed that the new cells were mainly those migrating from the connected native arteries. Recently, several groups have reported that there are stem cell niches existing in the adventitia of blood vessels, and these resident progenitor cells are important cell source for the vascular growth, repairing, and the occurrence and progression of vascular diseases [38e40]. We also observed c-kitþ cells within the graft wall at day 14 and 28. However, they were not detected at day 7 and 100 (Fig. S5).



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Fig. 6. Histological analysis and deposition of extracellular matrix in the regenerated grafts at day 100 in comparison with native aorta. Cross-sectional images of the regenerated grafts (A,C and E) and native artery (B, D and F) were immunochemically stained to identify the endothelial cells, smooth muscle cells and elastin. H&E staining show the structure of the explanted grafts (G) and native aorta (H). Masson’s trichrome staining show the presence of collagen (green) in the explanted grafts (I) and native aorta (J). Verhoeff’s staining show the presence of elastin (black) in the explanted grafts (K) and native aorta (L). Scale bar: 100 mm.

In addition, our data have suggested a potential mechanism underlying the regeneration of vascular grafts in rat aorta implantation. Macrophage infiltration was the initial and critical step during the process of regeneration. CD68þ macrophages were observed within the graft wall with uniform distribution and their number increased over time. In contrast, CD206þ cells (M2 macrophages) were observed within the surrounding tissue at day 7 and 14, and migrated into the graft wall with marked increase in number at day 28 and 100. These macrophages may stimulate cell migration into the grafts, and modulate their proliferation, differentiation, secretion and tissue remodeling. a-SMAþ cells had apparent migration trend. It appeared that loosely distributed a-SMAþ cells within the graft at day 7 moved

to the luminal side at day 14, formed a dense layer at day 28, and finally generated new tunica media at day 100. MYHþ cells were first observed at day 14 with low number, but became abundant with well-organized multiple-layered structure at day 28 and 100, indicating the remodeling process toward mature and functional media. Endothelium formation was observed at day 14, and became complete at day 28. Staining of the longitudinal sections with vWF antibody demonstrated that the endothelial cells might migrate from adjacent native arteries. Capillaries were observed within the graft at day 14, and showed a trend of increase in number over time without regression at day 100. The above processes have been summarized in Fig. 8.

Fig. 7. The physiological functions of the explanted grafts were assessed by wire myography. The rings of explanted grafts displayed contractile responses in response to KCL and adrenaline. Acetylcholine induced relaxation in regenerated grafts pre-constricted with adrenaline, indicating the presence of functional endothelium.


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Fig. 8. Proposed changes and migration of cells during the remodeling process of vascular grafts. Macrophage infiltration is the initial and critical step during the process. CD68þ macrophages were observed within the graft wall with uniform distribution, while, CD206þ cells migrated from the surrounding tissue into the graft wall. The macrophages stimulate the adventitial cells for activation, proliferation and migration from the outer into the graft. It appeared that loosely distributed a-SMAþ cells within the graft at early period moved to the luminal side and formed a dense layer at day 28. This a-SMAþ cell layer transformed into MYHþ cells and enhanced its thickness after 100 days. These a-SMAþ and MYHþ cells will synthesize extracellular matrix and are organized into the artery wall. Endothelialization and capillary formation were observed at early period and increased over time. At the long term, macrophages exist in the grafts and play a role in graft degradation and regulation of tissue remodeling.

5. Conclusions The present study has demonstrated that the thicker-fiber electropsun PCL grafts with larger pores markedly enhanced cell infiltration, vascularization and efficient regeneration of functional tunica media in comparison with the regular thinner-fiber PCL grafts. The scaffold structure was particularly important to stimulate the macrophage polarization into M2 phenotype, which consequently mediated the regeneration of cell-free PCL grafts into neo-arteries in vivo. Our findings suggest that the combined biomaterial-immunomodulatory approach should help with better design of next generation of vascular grafts. Acknowledgment The work was financially supported by NSFC projects (No. 81000680, 81171478 and 81371699), National Basic Research Program (973, No. 2011CB964903 and 2011CB606202), and Doctoral Fund of Ministry of Education of China (20111106110039). We thank Xiaohua Jia from Institute of Automation, Chinese Academy of Sciences, Beijing, China, for providing PCR primers. Appendix A. Supplementary data Supplementary data related to this article can be found at http:// dx.doi.org/10.1016/j.biomaterials.2014.03.078. References [1] Garg K, Sell SA, Madurantakam P, Bowlin GL. Angiogenic potential of human macrophages on electrospun bioresorbable vascular grafts. Biomed Mater 2009;4:031001. [2] Pektok E, Nottelet B, Tille JC, Gurny R, Kalangos A, Moeller M, et al. Degradation and healing characteristics of small-diameter poly(ε-caprolactone) vascular grafts in the rat systemic arterial circulation. Circulation 2008;118: 2563e70. [3] Innocente F, Mandracchia D, Pektok E, Nottelet B, Tille JT, de Valence S, et al. Paclitaxel-eluting biodegradable synthetic vascular prostheses: a step towards reduction of neointima formation? Circulation 2009;120(suppl. 1):S37e45. [4] de Valence S, Tille JC, Giliberto JP, Mrowczynski W, Gurny R, Walpoth BH, et al. Advantages of bilayered vascular grafts for surgical applicability and tissue regeneration. Acta Biomater 2012;8:3914e20. [5] de Valence S, Tille JC, Mugnai D, Mrowczynski W, Gurny R, Möller M, et al. Long term performance of polycaprolactone vascular grafts in a rat abdominal aorta replacement model. Biomaterials 2012;33:38e47.

[6] Wu W, Allen R, Wang Y. Fast degrading elastomer enables rapid remodeling of a cell-free synthetic graft into a neo-artery. Nat Med 2012;18:1148e53. [7] Assmann A, Delfs C, Munakata H, Schiffer F, Horstkötter K, Huynh K, et al. Acceleration of autologous in vivo recellularization of decellularized aortic conduits by fibronectin surface coating. Biomaterials 2013;34:6015e26. [8] Hasan A, Memic A, Annabi N, Hossain M, Paul A, Dokmeci MR, et al. Electrospun scaffolds for tissue engineering of vascular grafts. Acta Biomater 2014;10:11e25. [9] Bashur CA, Eagleton MJ, Ramamurthi A. Impact of electrospun conduit fiber diameter and enclosing pouch pore size on vascular constructs grown within rat peritoneal cavities. Tissue Eng Part A 2013;19:809e23. [10] Rayatpisheh S, Heath DE, Shakouri A, Rujitanaroj PO, Chew SY, Chan-Park MB. Combining cell sheet technology and electrospun scaffolding for engineered tubular, aligned, and contractile blood vessels. Biomaterials 2014;35:2713e9. [11] Baker BM, Gee AO, Metter RB, Nathan AS, Marklein RA, Burdick JA, et al. The potential to improve cell infiltration in composite fiber-aligned electrospun scaffolds by the selective removal of sacrificial fibers. Biomaterials 2008;29:2348e58. [12] Wang K, Xu M, Zhu MF, Hong Su, Wang HJ, Kong DL, et al. Creation of macropores in electrospun silk fibroin scaffolds using sacrificial PEOmicroparticles to enhance cellular infiltration. J Biomed Mater Res A 2013;101:3474e81. [13] Lee BLP, Jeon H, Wang A, Yan Z, Yu J, Grigoropoulos C, et al. Femtosecond laser ablation enhances cell infiltration into three-dimensional electrospun scaffolds. Acta Biomater 2012;8:2648e58. [14] Rnjak-Kovacina J, Weiss AS. Increasing the pore size of electrospun scaffolds. Tissue Eng Part B 2011;17:365e72. [15] Zhong S, Zhang Y, Lim CT. Fabrication of large pores in electrospun nanofibrous scaffolds for cellular infiltration: a review. Tissue Eng Part B 2012;18: 77e87. [16] Pham QP, Sharma U, Mikos AG. Electrospun poly(ε-caprolactone) microfiber and multilayer nanofiber/microfiber scaffolds: characterization of scaffolds and measurement of cellular infiltration. Biomacromolecules 2006;7:2796e805. [17] Rnjak-Kovacina J, Wise SG, Li Z, Maitz PK, Young CJ, Wang Y, et al. Tailoring the porosity and pore size of electrospun synthetic human elastin scaffolds for dermal tissue engineering. Biomaterials 2011;32:6729e36. [18] Roh JD, Sawh-Martinez R, Brennan MP, Jay SM, Devine L, Rao DA, et al. Tissueengineered vascular grafts transform into mature blood vessels via an inflammation-mediated process of vascular remodeling. Proc Natl Acad Sci 2010;107:4669e74. [19] Hibino N, Yi T, Duncan DR, Rathore A, Dean E, Naito Y, et al. A critical role for macrophages in neovessel formation and the development of stenosis in tissue-engineered vascular grafts. FASEB J 2011;25:4253e63. [20] Mokarram N, Merchant A, Mukhatyar V, Patel G, Bellamkonda RV. Effect of modulating macrophage phenotype on peripheral nerve repair. Biomaterials 2012;33:8793e801. [21] Madden LR, Mortisen DJ, Sussman EM, Dupras SK, Fugate JA, Cuy JL, et al. Proangiogenic scaffolds as functional templates for cardiac tissue engineering. Proc Natl Acad Sci 2010;107(34):15211e6. [22] Garg K, Pullen NA, Oskeritzian CA, Ryan JJ, Bowlin GL. Macrophage functional polarization (M1/M2) in response to varying fiber and pore dimensions of electrospun scaffolds. Biomaterials 2013;34:4439e51. [23] Zheng W, Wang Z, Song L, Zhao Q, Zhang J, Li D, et al. Endothelialization and patency of RGD-functionalized vascular grafts in a rabbit carotid artery model. Biomaterials 2012;33:2880e91.


Z. Wang et al. / Biomaterials xxx (2014) 1e11 [24] Tomadakis MM, Sotirchos SV. Transport-properties of random arrays of freely overlapping cylinders with various orientation distributions. J Chem Phys 1993;98:616e26. [25] Zhu Y, Bian Z, Lu P, Karas RH, Bao L, Cox D, et al. Abnormal vascular function and hypertension in mice deficient in estrogen receptor beta. Science 2002;295:505e8. [26] Zilla P, Bezuidenhuit D, Human P. Prosthetic vascular grafts: wrong models, wrong questions and no healing. Biomaterials 2007;28:5009e27. [27] Swartz DD, Andreadis ST. Animal models for vascular tissue-engineering. Curr Opin Biotechnol 2013;24:916e25. [28] Salacinski HJ, Goldner S, Giudiceandrea A, Hamilton G, Seifalian AM, Edwards A, et al. The mechanical behavior of vascular grafts: a review. J Biomater Appl 2001;15:241e77. [29] Ki CS, Park SY, Kim HJ, Jung HM, Woo KM, Lee JW, et al. Development of 3-D nanofibrous fibroin scaffold with high porosity by electrospinning: implications for bone regeneration. Biotechnol Lett 2008;30:405e10. [30] Wang K, Zhu MF, Li T, Zheng WT, Li L, Xu M, et al. Co-electrospraying of poly(ethylene oxide) increased the pore size of electrospun polycaprolactone and enhanced cell infiltration. J Biomed Nanotechnol 2014;10:1588e98. [31] Yixiang D, Yong T, Liao S, Chan CK, Ramakrishna S. Degradation of electrospun nanofiber scaffold by short wave length ultraviolet radiation treatment and its potential applications in tissue engineering. Tissue Eng Part A 2008;14:1321e9.

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[32] Park JE, Barbul A. Understanding the role of immune regulation in wound healing. Am J Surg 2004;187:11Se6S. [33] Arras M, Ito WD, Scholz D, Winkler B, Schaper J, Schaper W. Monocyte activation in angiogenesis and collateral growth in the rabbit hind limb. J Clin Invest 1998;101:40e50. [34] Heil M, Ziegelhoeffer T, Wagner S, Fernández B, Helisch A, Martin S, et al. Collateral artery growth (arteriogenesis) after experimental arterial occlusion is impaired in mice lacking CC-chemokine receptor-2. Circ Res 2004;94:671e7. [35] Brown BN, Badylak SF. Expanded applications, shifting paradigms and an improved under standing of hostebiomaterial interactions. Acta Biomater 2013;9:4948e55. [36] Bonetti PO, Lerman LO, Lerman A. Endothelial dysfunction: a marker of atherosclerotic risk. Arterioscler Thromb Vasc Biol 2003;23:168e75. [37] Hu J, Sun X, Ma HY, Xie CQ, Chen YE, Ma PX. Porous nanofibrous PLLA scaffolds for vascular tissue engineering. Biomaterials 2010;31:7971e7. [38] Hu YH, Zhang ZY, Torsney E, Afzal AR, Davison F, Metzler B, et al. Abundant progenitor cells in the adventitia contribute to atherosclerosis of vein graft sin ApoE-deficient mice. J Clin Invest 2004;113:1258e65. [39] Hu YH, Xu QB. Adventitial biology: differentiation and function. Arterioscler Thromb Vasc Biol 2011;31:1523e9. [40] Majesky MW, Dong XR, Hoglund V, Daum G, Mahoney Jr WM. The adventitia: a progenitor cell niche for the vessel wall. Cells Tissues Organs 2012;195:73e81.


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