Design and Fabrication of a Biomimetic Vascular Scaffold Promoting

Aug 22, 2018 - Department of Biomedical Engineering, University of Memphis, Memphis , Tennessee 38017 , United States. ACS Appl. Bio Mater. , 2018, 1 ...
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Design and Fabrication of a Biomimetic Vascular Scaffold Promoting in Situ Endothelialization and Tunica Media Regeneration Tong Wu,†,# Jialing Zhang,‡,# Yuanfei Wang,§ Binbin Sun,† Meng Yin,*,∥ Gary L. Bowlin,⊥ and Xiumei Mo*,†

ACS Appl. Bio Mater. Downloaded from pubs.acs.org by UNIV OF SOUTH DAKOTA on 09/12/18. For personal use only.

†

State Key Lab for Modification of Chemical Fibers and Polymer Materials, College of Chemistry, Chemical Engineering and Biotechnology, Donghua University, Shanghai 201620, China ‡ Cardiovascular Center, Children’s Hospital of Fudan University, Shanghai 201102, China § State Key Laboratory of Bioreactor Engineering, School of Resources and Environmental Engineering, East China University of Science and Technology, Shanghai 200237, China ∥ Department of Cardiothoracic Surgery, Shanghai Children’s Medical Center, Shanghai Jiaotong University School of Medicine, Shanghai 200127, China ⊥ Department of Biomedical Engineering, University of Memphis, Memphis, Tennessee 38017, United States S Supporting Information *

ABSTRACT: Multilayered vascular scaffolds may be considered advantageous in regenerating vascular tissues due to the nature of mimicking the native structure of a blood vessel. However, there are currently limited small-diameter vascular scaffolds integrating the specific features of native tunica intima (anti-thrombus and rapid endothelialization) and tunica media (the alignment and ingrowth of smooth muscle cells (SMCs), structural elements capable of promoting vascular regeneration and function). To address this limitation, we developed a modified electrospinning method capable of fabricating a bilayer vascular scaffold with a 2-mm inner diameter and investigated the in vivo performance and regenerative capacity using a rat abdominal aorta, with a 2-month implantation period. The vascular scaffold was fabricated from poly(L-lactide-cocaprolactone)/collagen (PLCL/COL) nanofibers and nanofiber yarns, comprising the luminal and medial layers, respectively. Heparin and anti-CD133 antibody (HEP/CD133) were incorporated into the PLCL/COL nanofibers comprising the luminal layer. The mechanical characterization demonstrated compliance of the bilayer scaffold, which was comparable to the human saphenous vein and improved over commercially available e-PTFE grafts. The incorporated components (HEP/CD133) were released over a period of nearly 40 days, during which the nanofibers and nanofiber yarns maintained their structure. Moreover, the released heparin contributed to lumen anticoagulation functionality initially, and the incorporated anti-CD133 antibody promoted the development of a neo-intima. In addition, SMCs proliferated and penetrated throughout the entire nanofiber yarn outer structure. In vivo evaluations demonstrated that a monolayer of endothelial cells (CD31 positive), as well as the aligned and infiltrated smooth muscle tissues (α-SMA positive), were regenerated on the inner and outer layers of the fabricated scaffold, respectively, demonstrating the capacity to regenerate structures mimicking native blood vessels. In conclusion, the functionalized bilayer scaffold can be viewed as a promising candidate for in situ vascular tissue regeneration. KEYWORDS: electrospun nanofibers, nanofiber yarns, anti-CD133 antibody, endothelialization, smooth muscle, bilayer vascular scaffold

1. INTRODUCTION The application of tissue-engineered scaffolds has been extensively investigated to repair tissue defects by integrating biomaterials with cells and bioactive molecules.1 From a biomimicking perspective, a vascular scaffold should ideally replicate the three concentric layers in a native blood vessel, including tunica intima, tunica media, and tunica adventitia. More specifically, the intima and media are considered as the vital layers for vascular function, consisting of a continuous monolayer of endothelial cells (ECs) and circumferentially aligned smooth muscle cells (SMCs), respectively.2 Based on the natural design, it is hypothesized that multilayered scaffolds will be advantageous in regenerating vascular tissues relative to © XXXX American Chemical Society

a scaffold comprised of a single layer. De Valence et al. fabricated a bilayer vascular scaffold by electrospinning, consisting of a high-porosity nanofiber scaffold and a lowporosity layer of nanofibers on either the luminal or the adventitial side.1 The in situ results demonstrated that a lowporosity layer combined with a high-porosity scaffold was critical to reduce blood leakage immediately upon implantation and achieve long-term cell infiltration. By comparing the bilayer scaffolds with two different microarchitectures, they Received: June 25, 2018 Accepted: August 22, 2018 Published: August 22, 2018 A

DOI: 10.1021/acsabm.8b00269 ACS Appl. Bio Mater. XXXX, XXX, XXX−XXX

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ACS Applied Bio Materials

were not aligned, and the infiltration of SMCs into the interior of the scaffold was insufficient. To integrate the specific features of anti-thrombus, rapid reendothelialization, and SMCs alignment and ingrowth into the same scaffold, we have designed a bilayer vascular scaffold to meet these design requirements. More specifically, the scaffold was comprised of PLCL/COL nanofibers incorporating heparin and anti-CD133 antibody in the intimal layer, and aligned/porous PLCL/COL nanofiber yarns in the outer layer. PLCL has been widely used in soft tissue engineering due to its high mechanical strength and elasticity. We have used PLCL to construct vascular scaffolds which have shown sufficient mechanical properties when in vivo implantation was performed. However, synthetic polymer alone, or the simple combination with a natural polymer, is insufficient for in situ vascular tissue engineering. As such, PLCL was selected as the synthetic part, combined with collagen, and further incorporated with biological agents in this study. Heparin is a commonly used anticoagulant16−18 which has been utilized extensively to prevent thrombus formation during the early period post scaffold implantation. CD133 is a glycoprotein, expressed on the circulating hematopoietic and putative endothelial regenerating cells.19−21 It has been demonstrated that scaffolds with surface immobilization of anti-CD133 antibody promote re-endothelialization post vascular injury.19,20,22 However, to date, there has been no study investigating the encapsulation of anti-CD133 antibodies within nanofibers by coaxial electrospinning and the antibody’s sustained release from the fibrous structure. Herein, the antiCD133 antibodies within the core of the nanofibers were used for the long-term recruitment of endothelial progenitor cells (EPCs), driving luminal re-endothelialization. Regarding the design of the outer layer of the vascular scaffold, this study utilized aligned and porous nanofiber yarns fabricated by dynamic liquid electrospinning, which have been shown to enhance cell orientation and penetration.23−26 As such, PLCL/ COL nanofiber yarns were investigated to support SMCs alignment and infiltration, promoting the regeneration of a functional media layer. The uniaxial mechanical properties and compliance of the fabricated bilayer scaffolds were then measured. In addition, the release kinetics of heparin and anti-CD133 antibody, blood biocompatibility, luminal EPCs adhesion, and SMCs ingrowth in vitro were evaluated. A preliminary in vivo evaluation of the bilayer scaffolds was conducted using an abdominal aortic rat model over a 2-month implantation period.

concluded that the majority of the cellular infiltration was derived from the marginal tissues rather than the bloodstream, providing a critical key insight into the design of nextgeneration vascular scaffolds. Electrospinning has gained popularity over the past several decades to produce fibers from nano- to microscale, as well as the capacity to generate scaffolds possessing a high porosity and large surface area-to-volume ratio.3 A larger selection of natural and synthetic polymers can be electrospun to fabricate fibrous scaffolds with different microstructures.4−6 In general, the natural biomaterials, such as collagen (COL), chitosan, gelatin, and silk fibroin (SF), have good biocompatibility and low thrombogenicity, whereas the synthetic biomaterials, such as polyurethanes (PU), poly(lactide-co-glycolide) (PLGA), and poly(L-lactide-co-caprolactone) (PLCL), can provide suitable mechanical strength and predictable/regulated biodegradability.7−9 As such, the composite or blended scaffolds consisting of natural and synthetic materials have been investigated for vascular tissue engineering.10−12 Zhang et al. fabricated a trilayer scaffold by combining electrospinning and braiding.2 Different ratios of SF/PLCL electrospun nanofibers were used to construct the intima (with heparinization) and the media, and SF yarns were braided as the adventitia to enhance the ability to withstand high pressures. The scaffolds exhibited satisfactory compliance and bursting strength, anticoagulant properties, and in vitro cytocompatibility with ECs and SMCs. However, no in vivo evaluation was presented in the study to investigate the re-endothelialization and the reconstruction of the media using the vascular scaffold in situ. In another study, a trilayer vascular scaffold, consisting of aligned PLCL/COL nanofibers forming the intima, PLGA/SF nanofiber yarns comprising the media, and random PLCL/ COL nanofibers forming the adventitia, was evaluated.13 The results of this study demonstrated that the trilayer vascular scaffold supported cell infiltration, scaffold biodegradation, and collagen production after a 10-week subcutaneous implantation. More importantly, the bare scaffolds, which contained no biological cues, were insufficient in supporting rapid reendothelialization and regeneration of the media comprised of smooth muscles in situ. This is critical, because any delayed regeneration may result in blood clotting and loss of scaffold structural integrity prior to functional vascular tissue regeneration. Coaxial and emulsion electrospinning are commonly used techniques to fabricate drug-loaded nanofibers, and the unique core−shell structure ensures the sustained release of the core solution.14 To promote and enhance vascular regeneration in vivo, Zhang et al. encapsulated vascular endothelial growth factor (VEGF) in fibers comprising the intima, and plateletderived growth factor-bb (PDGF) in fibers comprising the adventitia, to form a bilayer vascular scaffold.11 The in vivo results demonstrated that the vascular scaffold permitted ECs’ adhesion onto the lumen, and SMCs grew on the surface of the outer layer without thrombosis following implantation into a rabbit carotid artery after 4 weeks. Han et al. further designed a trilayer vascular scaffold by incorporating VEGF into intimal fibers and PDGF into medial fibers, and adding a third, outer layer of PCL/gelatin fibers for mechanical stability.15 Their results demonstrated that the dual release of VEGF and PDGF promoted re-endothelialization and inhibited SMCs’ hyperproliferation in the lumen, and the scaffold maintained patency in a rabbit carotid artery for 8 weeks. One limitation of both studies was that the SMCs which developed the adventitia

2. EXPERIMENTAL METHODS 2.1. Materials. The PLCL (50:50, Mn: 450 000) used in this study was purchased from Jinan Daigang Biomaterial Co., Ltd. (Jinan, China). Porcine-derived type I collagen (Mn: 100 000) was obtained from Chengdu Kele Bio-tech Co., Ltd. (Chengdu, China). The solvent, 1,1,1,3,3,3-hexafluoro-2-propanol (HFIP), was purchased from Shanghai Darui Fine Chemicals Co., Ltd. (Shanghai, China). The cross-linking agent, 25% aqueous glutaraldehyde (GA) solution, was purchased from Sinopharm Chemical Reagent Co., Ltd. (Shanghai, China). Bovine serum albumin (BSA) was purchased from Sigma-Aldrich (St. Louis, MO, USA). Mouse bone marrowderived EPCs were isolated and identified after 7 days of culture as previously described.27 The cryopreserved EPCs, passage 2 (P2), were reanimated and cultured, and P3 EPCs were utilized. The SMCs (human source) were obtained from the Institute of Biochemistry and Cell Biology (Chinese Academy of Sciences, China). All cell culture B

DOI: 10.1021/acsabm.8b00269 ACS Appl. Bio Mater. XXXX, XXX, XXX−XXX

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where M1 and M0 were the actual and theoretical weights, respectively, of heparin or CD133/BSA, in the bilayer scaffold. The release kinetics of heparin and CD133/BSA were examined by first taking a 50 mg bilayer scaffold sample and immersing the sample into a 15 mL centrifuge tube with 5 mL PBS solution, which was incubated on a continuous horizontal shaker at 37 °C and 120 rpm. At the pre-scheduled time points, 1 mL of the release medium was sampled for analysis, and an equal volume of fresh PBS was replenished for the subsequent incubation period. The amount of heparin and CD133/BSA in the release medium was analyzed using the Toluidine blue method and the BCA protein assay kit as previously noted. To quantify concentration in the releasate, known concentrations of heparin and CD133/BSA were used to obtain standard curves. The results were expressed as the cumulative percent of released heparin or CD133/BSA. All testing was conducted in triplicate at each time point. After 40 days incubation, the inner and outer surfaces of the vascular scaffold were observed by SEM. 2.5. In Vitro Evaluation of Blood Compatibility and EPCs Recruitment. The anticoagulation performance of PLCL/COLHEP/CD133, PLCL/COL-HEP, and PLCL/COL nanofibers were evaluated by standard platelet adhesion and hemolysis tests. Fresh blood samples were collected from healthy New Zealand White Rabbits in compliance with the institutional guidelines of Shanghai Jiaotong University School of Medicine. The drawn blood samples were first injected into BD Blood Collection Vacutainer containing 3.5 mg of K2EDTA as the anticoagulant followed by inversions to ensure thorough mixing. Platelet-rich plasma was obtained by centrifugation at 1200 rpm for 10 min at 25 °C using a plastic centrifuge tube for use in the platelet adhesion analysis. Normal saline was used to dilute the blood at a ratio of 4:5 (v/v) for the hemolysis test. Disk scaffold samples which were 14 mm in diameter were disinfected with 75% ethanol immersion and rinsed with PBS solution three times. Then, 0.5 mL of platelet-rich plasma was overlaid onto each sample. After a 2-h incubation at 37 °C with mild shaking, the samples were gently rinsed with PBS to remove non-adhered platelets and then fixed in 4% paraformaldehyde, dehydrated with gradient alcohol (30%, 50%, 70%, 80%, 90%, 95%, and 100%), and freezedried. Finally, the scaffolds with platelets adhered on the nanofibers were sputter-coated with gold for SEM observation.27,32 The hemolysis analysis was performed using the previously reported procedure.27,33 To begin, 15 mL centrifuge tubes containing 10 mL of normal saline were incubated in a water bath at 37 °C for 30 min. Then, the samples were immersed into the centrifuge tubes, and 0.2 mL diluted blood was added. For this study, 10 mL of normal saline solution and double distilled water combined with 0.2 mL of diluted blood was separately served as the negative and positive control. After incubation at 37 °C for 1 h, the samples were removed, and the tubes were centrifuged at 3000 rpm for 10 min. After centrifugation, 100 μL of the supernatant was transferred into a 96well plate, and absorbance was measured with a microplate reader (BioTek, USA) at 545 nm. The hemolysis rates were calculated using eq 3:

media and reagents were purchased from Gibco Life Technologies Co. (USA), unless otherwise specified. 2.2. Fabrication of Bilayer Vascular Scaffold. To create the inner layer of the bilayer scaffold, PLCL/COL (w/w, 3:1) was dissolved in HFIP at a concentration of 10% (w/v) as the shell fiberforming solution. For the fiber core solution, 0.15 g/mL heparin and 20 μg/mL anti-CD 133 antibodies with 1 mg/mL BSA were dissolved in a phosphate-buffered saline (PBS) solution. To obtain the PLCL/ COL nanofibers incorporating heparin and anti-CD 133 antibody (PLCL/COL-HEP/CD133), coaxial electrospinning was conducted using an applied voltage of +14 kV, flow rates of 1.0 mL/h (shell) and 0.1 mL/h (core), a coaxial nozzle with 0.34 and 1.12 mm inner diameter, respectively, of the inner and outer nozzle, and a rotating (500 rpm) stainless-steel rod (2 mm diameter and 10 cm length) as the collector. After electrospinning, the nanofibers (on the mandrel) were cross-linked by GA vapor for 20 min. A dynamic liquid system, which was previously described in detail, was used to fabricate electrospun PLCL/COL nanofiber yarns as the outer layer of the bilayer scaffold.28,29 Briefly, an applied voltage of +12 kV and a flow rate of 1.0 mL/h were used during electrospinning. A dynamic liquid vortex was generated to twist the nanofibers into nanofiber yarns as the PLCL/COL nanofibers reached the water surface. Then, the nanofiber yarns were drawn through the drain hole in the center of the water basin and directly collected on the rotating mandrel with the cross-linked PLCL/COL-HEP/CD133 nanofibers on it. Finally, the bilayer vascular scaffolds were frozen at −80 °C for 24 h and lyophilized using a Labconco freeze-dryer (LIGHT ACE HK Ltd.) overnight. The devices used for coaxial electrospinning and dynamic liquid electrospinning are presented in Figure S1. Heparin-loaded PLCL/COL (PLCL/COL-HEP) and PLCL/COL nanofibers were electrospun for comparison purposes. Before testing ensued, samples were stored in a vacuum oven at room temperature. 2.3. Characterization. The overall morphology of the bilayer vascular scaffold was observed by a digital camera (Nikon D5600) and a scanning electron microscope (SEM, JSM-5600, Japan) at an accelerating voltage of 15 kV and a working distance of 8 mm. The PLCL/COL-HEP/CD133, PLCL/COL-HEP, and PLCL/COL nanofibers along with the PLCL/COL nanofiber yarns were imaged for analysis using the SEM. The core−shell structure of a single PLCL/COL-HEP/CD133 nanofiber was observed by a transmission electron microscope (HT7700, Hitachi). A universal material testing machine (H5K-S, Hounsfield, UK) with a 50 N load cell was used to evaluate the tensile mechanical properties of the bilayer vascular scaffold at a temperature of 20 °C, a relative humidity of 65%, and an extension rate of 10 mm/min. The compliance of the hydrated, bilayer vascular scaffold was tested using TM 1 Test Bench System (Bose Electro force, USA) employing a fluid flow rate of 100 mL/min and a testing frequency of 1 Hz. Compliance under the pressure range of 80−120 mmHg was calculated using eq 1: % compliance =

(R P2 − R P1) R P1

×

104 R 2 − R1

hemolysis rate = (Ab1 − Ab3)/(Ab2 − Ab3) × 100%

(1)

where Ab1, Ab2, and Ab3 refer to the absorbance of the supernatant from the sample, positive control, and negative control, respectively. EPCs were reanimated and cultured using an EGM-2 MV Bullet kit (Lonza, Switzerland) consisting of endothelial basal medium, 5% fetal bovine serum, human epidermal growth factor (hEGF), vascular endothelial growth factor (VEGF), fibroblast growth factor-basic (hFGF-B), insulin-like growth factor 1 (IGF-1), ascorbic acid, and heparin at an atmosphere of 5% CO2 and 37 °C. In addition, 10% fetal bovine serum and 1% antibiotic−antimycotic were added to the culture medium. The nanofibers were positioned into 24-well plates, disinfected by 75% ethanol vapor for 24 h, and washed with sterile PBS three times at room temperature. Then, 2.0 × 104 EPCs were seeded per well, and the media was refreshed every 2 days. At 1, 4, and 8 h and 1, 4, and 7 days post culture, cell adhesion and proliferation were evaluated using Cell Counting Kit-8 (Dojindo Lab., Japan) as previously reported.5 The morphology of the EPCs were

where R is the internal radius, P1 is the lower internal pressure, and P2 is the higher internal pressure.30 2.4. Encapsulation Efficiency and Release Kinetics of Heparin and Anti-CD133 Antibody. The encapsulation efficiency was determined according to a previously published method.17 Briefly, 50 mg of the bilayer vascular scaffold was dissolved in 4-mL dichloromethane/PBS (1:1, v/v) solution, and then centrifuged at 4000 rpm for 5 min. The aqueous supernatant was pipetted, and the concentration of heparin and anti-CD133 antibody/BSA (CD133/ BSA) were separately determined by the Toluidine blue method and a BCA protein assay kit (K812-1000, BioVision, Inc., San Francisco, USA), respectively.31 The encapsulation efficiency (EE) was calculated using eq 2: %EE = M1/M 0 × 100%

(3)

(2) C

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ACS Applied Bio Materials observed by SEM and an inverted fluorescence microscope (IFM, Olympus IX71, Japan) after 4 days of culture. Specifically, the samples were rinsed twice with PBS and fixed in 4% paraformaldehyde solution at 4 °C for 4 h. Half of the fixed samples were rinsed with PBS three times and then dehydrated in gradient concentrations of ethanol. After drying in a vacuum oven at room temperature overnight, the samples were observed under SEM as described earlier. The other half of the fixed samples were rinsed with PBS and permeabilized with 0.1% Triton X-100 (Sigma, USA) at room temperature for 5 min. Then, the samples were rinsed again with PBS and stained with 4′,6′-diamidino-2-phenylindole hydrochloride (DAPI, Invitrogen, USA) and rhodamine-conjugated phalloidin (Invitrogen, USA) to observe the nuclei and cytoskeletons of cells under the IFM. 2.6. In Vitro Evaluation of SMCs Ingrowth into the Nanofiber Yarns. SMCs were incubated in Dulbecco’s modified Eagle’s medium (DMEM) with 10% fetal bovine serum and 1% antibiotic−antimycotic in an atmosphere of 5% CO2 and 37 °C. After disinfection with 75% ethanol vapor and rinsing with PBS solution, the PLCL/COL nanofibers and nanofiber yarns were seeded with SMCs at a density of 2.0 × 104 cells per well. Proliferation of SMCs was evaluated by the Cell Counting Kit-8 after culturing for 1, 4, and 7 days. Cell morphology on the samples was observed by SEM after culturing for 4 days as previously described. To examine SMCs migration into the nanoyarns, the samples pre-seeded with SMCs for 4 days were harvested, embedded in freezing medium, and frozen at −80 °C. The frozen samples were sectioned into 10-μm thick slices and stained with DAPI and α-SMA (Invitrogen, USA). Specifically, the samples were fixed in 4% paraformaldehyde solution at 4 °C for 4 h, permeabilized with 0.1% Triton X-100 for 5 min, and blocked by 3% BSA/PBS solution for 1 h. Then, the samples were treated by αSMA primary antibody (1:100) at 4 °C overnight, followed by staining using Alexa Fluor 555 second antibody for 1 h at room temperature. Afterward, DAPI was used to stain the nuclei of cells for 5 min at room temperature. Between each procedure, the samples were washed with PBS solution three times. The fluorescence micrographs were acquired using the IFM. 2.7. In Vivo Evaluation. All experimental procedures for in vivo evaluation were performed under institutional guidelines for animal care and approved by the Animal Ethics Committee of Shanghai Children’s Medical Center, Shanghai Jiaotong University School of Medicine (Shanghai, China). Seven male SD rats with weight of 200− 300 g (Shanghai Slaccas Experimental Animal Ltd., China) were used for the abdominal aorta replacement models. The bilayer vascular scaffolds and control scaffolds made of an inner layer only (PLCL/ COL-HEP/CD133 fibers) with a 2 mm inner diameter were trimmed to 15 mm length and sterilized with ethylene oxide for 3 h. Afterward, all materials and scaffolds were handled aseptically. The rats were anesthetized with 30 mg/kg pentobarbital sodium (Merck KGaA, Darmstadt, Germany) by intraperitoneal injection. In each rat, the abdominal aorta with a length of 12 mm was removed. Then, the vascular scaffolds were implanted into the section of resected abdominal aorta and sutured to the proximal and distal ends of the native vessel using 8-0 monofilament nylon sutures. Afterward, the skin wound was closed using 3-0 nylon sutures under stringent aseptic conditions. Each type of scaffold was implanted into three rats, respectively. After implantation for 2 months, the animals were euthanized, and the scaffolds were harvested for histological analysis. The explanted scaffolds were fixed in 4% paraformaldehyde at 4 °C for 4 h, embedded in paraffin, and sectioned into 10-μm thick slices. Then, the slices were stained with hematoxylin and eosin (H&E) and Masson’s trichrome, respectively, for evaluation. Immunohistochemical staining of CD31 and α-SMA were also performed on the fixed slices to examine re-endothelialization and smooth muscle regeneration. DAPI was used to stain the nuclei of cells for 5 min. For comparison to a native blood vessel, an autologous vessel in a rat was also extracted and evaluated for histology. 2.8. Statistics Analysis. Statistical analyses were performed in Origin 9.0 (Origin Lab Inc., USA). Averages and standard deviations were calculated from at least three replicates and are presented as

mean ± standard deviation. Statistical differences were determined by one-way ANOVA analysis, and the differences were considered statistically significant at p < 0.05.

3. RESULTS AND DISCUSSION 3.1. Structure and Morphology of the Bilayer Vascular Scaffold. The bilayer scaffold was fabricated via coaxial electrospinning and dynamic liquid electrospinning. As shown in Figure S1, we first fabricated the PLCL/COL-HEP/ CD133 nanofibers as the inner layer of the vascular scaffold by coaxial electrospinning. After a 20 min cross-linking of the nanofiber layer in GA vapor, the rotating rod along with the asobtained inner layer was directly transferred to collect the PLCL/COL nanofiber yarns made by dynamic liquid electrospinning. In this way, the cross-linked PLCL/COL-HEP/ CD133 nanofibers can reduce the loss of bioactive agents during the fabrication of nanofiber yarns in water vortex. Figure 1 shows a photograph and representative SEM images

Figure 1. (A) Photograph and (B) SEM micrograph of the cross section of the bilayer vascular scaffold. (C,D) SEM micrographs of the surface of the inner and outer layer of the bilayer vascular scaffold; TEM micrograph inset in (C) was a single PLCL/COL-HEP/CD133 nanofiber; SEM micrograph inset in (D) was the cross section of the PLCL/COL nanofiber yarns.

of the bilayer vascular scaffold, which consisted of PLCL/ COL-HEP/CD133 nanofibers in the inner layer (Figure 1C) and PLCL/COL nanofiber yarns in the outer layer (Figure 1D). The dense PLCL/COL nanofibers encapsulated with heparin and anti-CD133 antibody in the inner layer possessed a core−shell structure, as shown in the transmission electron microscopy (TEM) image (inset in Figure 1C). The porous PLCL/COL nanofiber yarns in the outer layer showed porous and aligned organization (Figure 1D). Several random nanofibers were interconnected with the neighboring yarns. 3.2. Mechanical Properties. The tensile mechanical properties of the PLCL/COL-HEP/CD133 and PLCL/COL nanofibers, and PLCL/COL nanofiber yarns are separately shown in Figures S2 and S3. The ultimate stress and elongation at break of the PLCL/COL-HEP/CD133 nanofibers were 11.8 ± 0.62 MPa and 133.13 ± 14.49%, respectively, showing no significant difference compared with the PLCL/COL nanofibers (15.73 ± 2.55 MPa and 104.8 ± 11.11%, respectively). The mechanical strength of the nanofiber yarns in the parallel direction to their alignment (16.00 ± 1.30 MPa) was significantly larger than that in the perpendicular direction to their alignment (2.21 ± 0.37 MPa). The elongation at break of PLCL/COL nanofiber yarns in the D

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Figure 2. Tensile mechanical properties of the bilayer vascular scaffold: (A) Stress−strain curves at parallel and perpendicular directions; (B) ultimate tensile stress, (C) elongation at break, and (D) Young’s modulus of the bilayer vascular scaffold and a native porcine coronary artery.

parallel direction (257.68 ± 22.73%) and perpendicular direction (319.01 ± 13.73%) were both larger than that of random PLCL/COL nanofibers (104.8 ± 11.11%). The tensile mechanical properties and compliance of the bilayer vascular scaffolds were investigated to determine their adaptability as a vascular graft. Figure 2A shows the tensile stress−strain curves of the bilayer vascular scaffold at the parallel and perpendicular directions. The ultimate tensile stress (Figure 2B), elongation at break (Figure 2C), and Young’s modulus (Figure 2D) were separately compared with values for a natural porcine coronary artery.15 As shown in Figure 2A, the stress−strain curves at both parallel and perpendicular directions display two stages because of the bilayer structure. The ultimate tensile stress of the bilayer scaffold was 14.73 ± 3.11 MPa in the parallel direction and 7.38 ± 0.16 MPa in the perpendicular direction, both of which were greater than that of the porcine coronary artery (2.6 MPa) (Figure 2B). 15,34 Additionally, the elongations at break at parallel and perpendicular directions were 238.92 ± 38.97% and 220.62 ± 25.43%, respectively, both of which were larger than that of the porcine coronary artery (100%) (Figure 2C).15,34 Similarly, Young’s modulus of the bilayer scaffold at parallel and perpendicular directions (34.48 ± 1.22 and 13.23 ± 0.89 MPa, respectively) were greater than that of the porcine coronary artery (1 MPa) (Figure 2D).15,34 Even in wet condition, the tensile mechanical property is comparable to that of the porcine coronary artery. As shown in Figure S4, the ultimate tensile stress of the bilayer scaffold in wet condition was 13.43 ± 2.78 MPa in the parallel direction and 6.65 ± 1.08 MPa in the perpendicular direction. The elongations at break at parallel and perpendicular directions were 207.37 ± 20.45% and 200.06 ± 17.56%, respectively. The compliances of the bilayer vascular scaffold, natural blood vessels, and commercial e-PTFE are presented in Table 1. The compliance of the bilayer vascular scaffold was 1.43 ± 0.046%/100 mmHg, which is comparable to the value of human saphenous vein (0.7−1.5%/100 mmHg) and greater than the commercial e-PTFE (0.1%/100 mmHg).8,35,36 As the

Table 1. Compliances of the Bilayer Vascular Scaffold, Natural Blood Vessels, and Commercial e-PTFE sample

% compliance/100 mmHg

bilayer vascular scaffold [8, 35] human saphenous vein [8, 35] human artery [35, 36] e-PTFE

1.43 ± 0.046 0.7−1.5 4.5−6.2 0.1

scaffold is biodegradable after in vivo implantation, the initial mechanical strength and elongation of the scaffold should be greater than that of a native blood vessel to account for loss of mechanical integrity. Afterward, with the regeneration of SMCs in the vascular scaffold, the regenerated vascular wall can further support the required mechanical strength to counterpoise the hemodynamic forces encountered in the vascular circuit. 3.3. Encapsulation Efficiency and Release Behavior of Heparin and Anti-CD133 Antibody. Heparin and CD133/ BSA were encapsulated into the PLCL/COL nanofibers to prevent thrombosis and modulate re-endothelization of the lumen, respectively. The encapsulation efficiencies of heparin and CD133/BSA in the bilayer vascular scaffold were 40.39 ± 3.71% and 43.23 ± 5.63%, respectively. The loss of encapsulation efficiency is attributed to the electrospinning process. Some proteins may be distributed on the surface of the nanofibers instead of within the core.37−39 The release curves of heparin and CD133/BSA of the bilayer scaffold over a 40-day period are presented in Figure 3A,B. The initial burst release of heparin (around 15%) and CD133/BSA (10%) was separately observed, and then, the cumulative release was continuous at a relatively slow rate. At 40 days, the cumulative release amount of heparin and CD133/BSA was 69.89 ± 2.03% and 72.98 ± 1.19%, respectively. It has been reported that several factors influence the release behaviors of proteins in the core−shell fibers, such as the property of the shell fibers (hydrophilicity, degradability, etc.) and the cooperation with E

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Figure 3. (A) Heparin and (B) CD133/BSA release curves after immersing the bilayer scaffolds in PBS solution at 37 °C and 120 rpm for 40 days. SEM micrographs showing the morphologies of (C) the nanofibers in the inner layer and (D) the nanofiber yarns in the outer layer after heparin and CD133/BSA released for 40 days.

Figure 4. (A) SEM micrographs and (B) fiber diameter distribution of PLCL/COL, PLCL/COL-HEP, and PLCL/COL-HEP/CD133 nanofibers.

stabilizer (BSA).37,39−41 In this study, the hydrophilicity and degradability of the PLCL/COL nanofibers, and the existence of BSA contributed to the sustained and stable release of heparin and the anti-CD133 antibody. Figure 3C,D shows the morphologies of PLCL/COL nanofibers and nanofiber yarns after immersing them in the degradation media for 40 days. With the release of HEP/CD133, the average diameter of the PLCL/COL-HEP/CD133 nanofibers increased to 641.80 ± 190.13 nm, mainly owing to the swelling of the nanofibers in the degradation medium. Although the nanofibers and yarns swelled and crimped, they maintained the fiber and yarn organizations. In our previous study, we have demonstrated that the nanofibers or nanofiber yarns made of PLCL/COL maintained their structure and mechanical strength during biodegradation for more than 4 months.29,30,42 Thus, the bilayer vascular scaffold is able to support the structural

integrity and mechanical strength until fresh tissue regeneration. 3.4. Blood Compatibility and EPCs Growth on the Lumen Surface. We separately fabricated PLCL/COL, PLCL/COL-HEP, and PLCL/COL-HEP/CD133 nanofibers, with an average fiber diameter of 347.84 ± 160.12, 372.04 ± 180.52, and 431.38 ± 185.93 nm, respectively (Figure 4). The SEM images in Figure 5A−C show platelet adhesion on the different types of nanofibers, respectively. The PLCL/COL nanofibers were blood-compatible, with only a few adhered platelets. By inducing heparin into the nanofibers, the platelet adhesion was further inhibited on the PLCL/COL-HEP and PLCL/COL-HEP/CD133 nanofibers. Figure 5D displays the hemolysis rate of different nanofibers. The hemolysis test is based on the degree of erythrolysis that occurs when the material contacts erythrocytes in vitro. It is widely used to evaluate the destruction potential that implants present to F

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Figure 5. SEM micrographs of platelet adhesion on (A) PLCL/COL nanofibers, (B) PLCL/COL-HEP nanofibers, and (C) PLCL/COL-HEP/ CD133 nanofibers. (D) Hemolysis rates of PLCL/COL, PLCL/COL-HEP, and PLCL/COL-HEP/CD133 nanofibers. White arrows indicate the platelet adhesion. The insets in panels A−C are the SEM images magnifying the area outlined in a square in the original SEM images.

erythrocytes.27 The results show that the hemolysis rates of all the PLCL/COL, PLCL/COL-HEP, and PLCL/COL-HEP/ CD133 nanofibers are less than 5%. According to previous studies, equal to or less than 5% is the criterion for excellent blood compatibility.43 A lower hemolysis rate usually indicates better hemocompatibility. Thus, the PLCL/COL-HEP/ CD133 nanofibers may maintain better lumen patency relative to the PLCL/COL nanofibers. Figure 6A,B shows EPCs adhesion and proliferation on the different types of nanofibers and tissue culture plate (TCP). After separately culturing EPCs for 8 h and 7 days, EPCs adhesion and proliferation on PLCL/COL-HEP/CD133 nanofibers were significantly greater (p < 0.05) than the PLCL/COL-HEP and PLCL/COL nanofibers. This enhanced adhesion is mainly because CD133 is a cell surface antigen, which is specifically expressed on EPCs.21 As such, the released anti-CD133 antibodies from the PLCL/COL-HEP/CD133 nanofibers enhanced EPCs adhesion to the engineered surface. The fluorescence and SEM micrographs in Figure 6C,D display the EPCs morphology on different nanofibers. EPCs showed superior intercellular interaction and formed a continuous cell monolayer on the nanofibers loaded with anti-CD133 antibodies, in comparison to the PLCL/COLHEP and PLCL/COL nanofibers. The results indicate that the PLCL/COL-HEP/CD133 nanofibers in the inner layer are advantageous in promoting EPCs recruitment and maturation on the lumen surface. 3.5. SMCs Ingrowth on the Nanofiber Yarns in the Outer Layer. Figure 7A shows SMCs proliferation on PLCL/ COL nanofibers, PLCL/COL nanofiber yarns, and TCP. At the initial 24 h, the proliferation of SMCs showed no

significant difference among the different groups. However, the proliferation of SMCs on PLCL/COL nanofiber yarns was significantly greater (p < 0.05) than PLCL/COL nanofibers and TCP when cultured for 4 and 7 days. The enhanced porosity of the scaffold is mainly due to the enlarged pore size and porosity of nanofiber yarns relative to the dense nanofibers.28 The SEM micrographs in Figure 7B,C show the SMCs morphology on PLCL/COL nanofibers and nanofiber yarns. The SMCs grew with a random pattern and polygonal morphology on PLCL/COL nanofibers, whereas they spread along the direction of PLCL/COL nanofiber yarns, as marked by the white arrows. The fluorescence micrographs in Figure 7D,E display the cross section of PLCL/COL nanofibers and nanofiber yarns after SMCs were cultured for 4 days. The SMCs remained on the surface of PLCL/COL nanofibers (Figure 7D), whereas they infiltrated into the PLCL/COL nanofiber yarns, due to the larger porosity and pore size of the three-dimensional yarn structure relative to the dense nanofibers (Figure 7E). The results indicate that the PLCL/COL nanofiber yarns can provide sufficient spaces for SMCs proliferation and transmural ingrowth. 3.6. Analyses of the Explanted Scaffolds. After implanting the scaffolds in the rat abdominal aorta for 2 months, we conducted the histological analysis of transverse sections at the middle sites by H&E and Masson’s trichrome staining (Figure 8). Figure 8A−D shows the typical H&E images of the bilayer vascular scaffold. We demonstrated that the tubular structure was well-kept, and cells infiltrated into the interior of the scaffold. Also, the biodegradation of the scaffold was observed, where the fragments of the scaffold were wrapped by cells and fibrous tissues. As such, the porous and G

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Figure 6. (A) EPCs adhesion after culturing for 1, 4, and 8 h and (B) proliferation after culturing for 1, 4, and 7 days on PLCL/COL, PLCL/COLHEP, and PLCL/COL-HEP/CD133 nanofibers; *indicates significant difference at p < 0.05 level. (C) Fluorescence micrographs and (D) SEM micrographs of EPCs morphology on different types of nanofibers after culturing for 4 days.

Figure 7. (A) SMCs proliferation on PLCL/COL nanofibers, PLCL/COL nanofiber yarns, and TCP after culturing for 1, 4, and 7 days; *P < 0.05 when compared with that on PLCL/COL nanofibers and TCP. (B,C) SEM micrographs and (D,E) fluorescence micrographs of SMCs morphology on (B,D) PLCL/COL nanofibers and (C,E) PLCL/COL nanofiber yarns after culturing for 4 days.

(Figure 8E−H). As expected, the regeneration of muscle fibers (red) and collagen fibers (blue) was induced after implanting the scaffold in vivo for 2 months. The regenerated muscle and collagen fibers enclosed the scaffold fragments and contributed

swelled structure of nanofiber yarns, as well as the biodegradation of the scaffold, contributed to the inward growth of cells.25,29 The micrographs of Masson’s trichrome staining indicate the production of muscle and collagen fibers H

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Figure 8. Representative (A−D) H&E and (E−H) Masson’s trichrome staining micrographs of the bilayer vascular scaffold after implantation for 2 months. The white arrows in panels A−D indicate the scaffold fragments, and the arrows in panels E−H indicate the regenerated collagen fibers (blue).

Figure 9. Fluorescence micrographs of (A−D) the bilayer vascular scaffold after implantation for 2 months and (E−H) the autologous vessel both after immunohistochemical staining of (A,E) DAPI, (B,F) CD31, and (C,G) α-SMA. (D,H) The correspondingly merged fluoresence micrographs.

to maintaining the integral scaffold structure and mechanical support. We then examined the regeneration of the endothelium and smooth muscle using the specific label of CD31 and α-SMA. The fluorescence micrographs in Figure 9A−D shows the regeneration of the endothelium (green) and smooth muscle (red) on the bilayer vascular scaffold, and the cell nucleus was stained by DAPI (blue). We also extracted an autologous vessel as control and stained it under identical conditions (Figure 9E−H). The cell distribution in the entire bilayer scaffold was similar to an autologous vessel, which includes a monolayer of endothelial cells forming the lumen (green) and aligned smooth muscles (red) in the vessel wall. In summary, the encapsulated heparin in the inner layer contributed to the initial anticoagulation in the early period of implantation, and the anti-CD133 antibody simultaneously promoted EPCs recruitment and rapid re-endothelialization in the subsequent stages. Additionally, the aligned and porous nanofiber yarns in the outer layer contributed to the regeneration and organization of the SMCs. In Figure 9D, the gap between endothelium and smooth muscles is attributed to the insufficient infiltration of SMCs derived from the marginal tissues during such a short period and the large thickness of the inner layer. During the in situ regeneration in a period of 2 months, the continuous

monolayer of endothelium was almost remolded, whereas the SMCs need longer time to infiltrate throughout the entire scaffold. Besides, the inner layer of the vascular scaffold is not as thin as the intima of a normal blood vessel. The SMCs are expected to infiltrate into the space left by the biodegradation of the dense nanofibers in a longer time. In addition, the original thickness of the inner and outer layer of the bilayer vascular scaffold is 96.72 ± 3.66 and 543.94 ± 5.43 μm, respectively, as measured from the SEM micrograph in Figure 1. However, as shown in Figure 9, the thickness of the regenerative intima and media is 27.15 ± 4.9 and 196.06 ± 3.75 μm, respectively, after implanting the bilayer vascular scaffold in a rat abdominal aorta for 2 months. In this case, the thickness of the regenerative intima and media is comparable to that of the native rat abdominal aorta (36.96 ± 8.70 μm for intima and 105.14 ± 5.87 μm for media, respectively). It should be noted that the decreased thickness of the regenerative intima and media, as compared to the layer thickness of the original bilayer scaffold, is mainly due to three factors: (i) the large thickness of the original inner layer of dense nanofibers, as sufficient nanofibers are required to incorporate biological agents; (ii) the insufficient infiltration of SMCs from the marginal tissues in such a short period; and (iii) the biodegradation of nanofibers and nanofiber yarns with the regeneration of new tissues. I

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and ingrowth in the outer layer of the vascular scaffold. The regenerated tissues showed similar organization to the native blood vessel, indicating that the bilayer vascular scaffold is a potential candidate for early-stage vascular tissue regeneration. Long-term investigation of the scaffolds in large animals will be necessary to be performed for further improving the design and enhancing its functionality.

To better demonstrate the superiority of PLCL/COL nanofiber yarns as the outer layer of a vascular scaffold, we implanted a scaffold that only consisted of PLCL/COL-HEP/ CD133 nanofibers for in vivo evaluation under matching condition. Before the explanation of the scaffold, we investigated the patency of the single-layer vascular scaffold using color Doppler flow imaging (CDFI), which indicated the lumen patency after implantation for 2 months (Figure S5A). No thrombus was observed, which can be attributed to the release of heparin and anti-CD133 antibody, and the degraded scaffold fragments were enclosed in the regenerated tissues as well. However, the single-layer scaffold collapsed when explanted, mainly due to the missing support of the dynamic flowing of blood, as indicated in the images of H&E (Figure S5B,C). The significant differences between the bilayer and single-layer scaffold were visible regarding in vivo vascular tissue regeneration. The results indicate that the PLCL/COL nanofiber yarns in the outer layer are necessary to support the infiltration of abundant cells into the interior of the scaffold. In this way, tissues can be regenerated along with the scaffold fragments, as the vascular wall, to provide sufficient mechanical support. We showed that the bilayer scaffold mimicked the structure and function of a native vessel, which improved the outcome of early-stage vascular tissue regeneration. Relative to the adventitia that mainly consists of collagen and fibroblasts, intima and media play critical roles in maintaining vascular function and structure. While the intima aims for lumen patency, thrombus inhibition, and endothelium maturation, the media endows the blood vessel diastolic and systolic function, and maintains the structure of vascular wall. To this end, the remolding of intima and media is much more needed than the adventitia when designing a vascular scaffold. As the SMCs are mainly derived from the marginal tissues rather than the bloodstream, the tunica media will be more difficult to be regenerated if we first construct an adventitia before the SMCs penetrating into the media. As such, we design a scaffold that can remold the functional intima and media first. Then, through the secretion from the intrinsic cells in situ, the tissues in adventitia are expected to be further regenerated. It will be necessary to investigate the potential of the bilayer scaffold to induce vascular tissue regeneration over a long period of time in larger animals and determine the risk for aneurysm.

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ASSOCIATED CONTENT

S Supporting Information *

The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acsabm.8b00269. Figure S1, schematic diagram of coaxial electrospinning and dynamic liquid electrospinning for the fabrication of the bilayer vascular scaffold; Figured S2 and S3, mechanical properties of the PLCL/COL-HEP/CD133 and PLCL/COL nanofibers, and PLCL/COL nanofiber yarns; Figure S4, mechanical properties of the bilayer vascular scaffold in wet condition; Figure S5, color Doppler flow image showing the patency of the singlelayer vascular scaffold after implantation for 2 months, and H&E staining images of the single-layer vascular scaffold after implantation for 2 months (PDF)

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AUTHOR INFORMATION

Corresponding Authors

*E-mail: [email protected]. *E-mail: [email protected]. ORCID

Tong Wu: 0000-0001-8822-1868 Xiumei Mo: 0000-0001-9238-6171 Author Contributions #

T.W. and J.Z. contributed equally to this work.

Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS This research was supported by National Major Research Program of China (2016YFC1100202), National Natural Science Foundation of China (31470941, 81671833, 31771023), Science and Technology Commission of Shanghai Municipality (15JC1490100, 15441905100, 16CR3078B), Fundamental Research Funds for the Central Universities (CUSF-DH-D-2017047), Light of textile project (J201404), Collaborative Innovation Center for Translational Medicine (TM201504), Ai You Foundation (2017SCMC-AY002), and Pudong New Area Science and Technology Development Fund Minsheng Scientific Research (Medical and Health) Project (PKJ2016-Y33).

4. CONCLUSION Regarding biomimicking the structure and function of a native blood vessel, we designed a bilayer vascular scaffold based on PLCL/COL electrospun nanofibers and nanofiber yarns. Heparin and anti-CD133 antibody were encapsulated into the PLCL/COL nanofibers as the inner layer, whereas the porous and puffed PLCL/COL nanofiber yarns were explored as the outer layer. The bilayer vascular scaffold showed better compliance performance than the commercial e-PTFE and matched compliance with the human saphenous vein. In vitro studies indicated that the heparin and anti-CD133 antibody were sustained released for almost 40 days, contributing to reducing platelet adhesion and enhancing re-endothelization of the lumen. Moreover, the porous nanofiber yarns promoted the proliferation and ingrowth of SMCs. In vivo evaluation demonstrated that the bilayer vascular scaffold supported lumen patency and promoted rapid endothelialization after implantation in the abdominal artery of a rat for 2 months. Furthermore, the nanofiber yarns promoted SMCs alignment

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