Accepted Manuscript Hybrid small-diameter vascular grafts: anti-expansion effect of electrospun poly εcaprolactone on heparin-coated decellularized matrices Wenhui Gong, Dong Lei, Sen Li, Peng Huang, Quan Qi, Yijun Sun, Yijie Zhang, Zhe Wang, Zhengwei You, Xiaofeng Ye, Qiang Zhao PII:

S0142-9612(15)00876-5

DOI:

10.1016/j.biomaterials.2015.10.066

Reference:

JBMT 17163

To appear in:

Biomaterials

Received Date: 20 August 2015 Revised Date:

20 October 2015

Accepted Date: 26 October 2015

Please cite this article as: Gong W, Lei D, Li S, Huang P, Qi Q, Sun Y, Zhang Y, Wang Z, You Z, Ye X, Zhao Q, Hybrid small-diameter vascular grafts: anti-expansion effect of electrospun poly ε-caprolactone on heparin-coated decellularized matrices, Biomaterials (2015), doi: 10.1016/j.biomaterials.2015.10.066. This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting proof before it is published in its final form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.

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Hybrid small-diameter vascular grafts: anti-expansion effect of electrospun poly εcaprolactone on heparin-coated decellularized matrices

Zhe Wang1, Zhengwei You2*, Xiaofeng Ye1*, Qiang Zhao1*

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Wenhui Gong1#, Dong Lei2#, Sen Li1, Peng Huang2, Quan Qi1, Yijun Sun2, Yijie Zhang2,

1 Department of Cardiac Surgery, Ruijin Hospital, Shanghai Jiaotong University School of Medicine, Shanghai, P.R. China

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2 State Key Laboratory for Modification of Chemical Fibers and Polymer Materials, College of Materials Science and Engineering, Donghua University, Shanghai, P.R.

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China #These authors contributed equally to this work.

*Corresponding author: Qiang Zhao, Xiaofeng Ye and Zhengwei You

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Email: [email protected] or [email protected] or [email protected]

Address: Department of Cardiac Surgery, Ruijin Hospital, No.197, Ruijin Er Road,

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Shanghai, China Postal code: 200025

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Telephone: 086-021-64370045

Fax: 086-021-64671787

Conflict of Interest

The authors confirm that there are no known conflicts of interest associated with this publication and there has been no significant financial support for this work that could have influenced its outcome.

ACCEPTED MANUSCRIPT Hybrid small-diameter vascular grafts: anti-expansion effect of electrospun poly ε-caprolactone on heparin-coated decellularized matrices Wenhui Gong1#, Dong Lei2#, Sen Li1, Peng Huang2, Quan Qi1, Yijun Sun2, Yijie Zhang2, Zhe Wang1, Zhengwei You2*, Xiaofeng Ye1*, Qiang Zhao1*

Shanghai, P.R. China

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1 Department of Cardiac Surgery, Ruijin Hospital, Shanghai Jiaotong University School of Medicine, 2 State Key Laboratory for Modification of Chemical Fibers and Polymer Materials, College of Materials Science and Engineering, Donghua University, Shanghai, P.R. China #

These authors contributed equally to this work.

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Corresponding author: Qiang Zhao, Xiaofeng Ye and Zhengwei You Email: [email protected] or [email protected] or [email protected]

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Abstract:

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Small-diameter vascular grafts (SDVGs) (D < 6 mm) are increasingly needed in clinical settings for cardiovascular disease, including coronary artery and peripheral vascular pathologies. Vessels made from synthetic polymers have shortcomings such as thrombosis, intimal hyperplasia, calcification, chronic inflammation and no growth potential. Decellularized xenografts are commonly used as a tissue-engineering substitute for vascular reconstructive procedures. Although acellular allogeneic vascular grafts have good histocompatibility and antithrombotic properties, the decellularization process may damage the biomechanics and accelerate the elastin deformation and degradation, finally resulting in vascular graft expansion and even aneurysm formation. Here, to address these problems, we combine synthetic polymers with natural decellularized small-diameter vessels to fabricate hybrid tissue-engineered vascular grafts (HTEV). The donor aortic vessels were decellularized with a combination of different detergents and dehydrated under a vacuum freeze-drying process. Polycaprolactone (PCL) nanofibers were electrospun (ES) outside the acellular aortic vascular grafts to strengthen the decellularized matrix. The intimal surfaces of the hybrid small-diameter vascular grafts were coated with heparin before the allograft transplantation. Histopathology and scanning electron microscope revealed that the media of the decellularized vessels were severely injured. Mechanical testing of scaffolds showed that ES-PCL significantly enhanced the biomechanics of decellularized vessels. Vascular ultrasound and micro-CT angiography showed that all grafts after implantation in a rat model were satisfactorily patent for up to 12 weeks. ES-PCL successfully prevented the occurrence of vasodilation and aneurysm formation after transplantation and reduced the cell inflammatory infiltration. In conclusion, the HTEV with perfect histocompatibility and biomechanics provide a facile and useful technique for the development of SDVGs. Keywords: Vascular grafts, Electrospinning, PCL, Mechanical properties, Endothelialization

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1. Introduction

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Currently, small diameter vascular grafts (SDVGs, D < 6 mm) are increasingly needed in the clinic for coronary artery bypass grafting and BT shunts for congenital heart diseases and some peripheral vascular diseases. The conventional surgical treatment for these diseases, using saphenous veins or internal mammary arteries, has proven effective. However, autologous vascular grafts may be delimited by the shortage of donors for patient conditions and anatomic variations, and the harvest procedure will cause trauma. Therefore, it is particularly important to fabricate biologically responsive vascular alternatives [1]. Prosthetic vascular conduits made of synthetic materials, such as Dacron and polytetrafluoroethylene (PTFE), have been developed as substitutes for large-diameter vascular grafts (≥6-mm inner diameter) successfully. However, unfortunately, the acute thrombogenicity, anastomotic intimal hyperplasia, aneurysm formation, infection, and progression of atherosclerotic disease usually lead to significantly reduced durability and availability of the artificial vascular grafts [2]. To address these challenges, tissue engineering has emerged as a new technique to make biologically active SDVGs [3]. The purpose of tissue engineering vascular grafts is to develop alternative grafts that integrate with the patient’s native tissue to restore physiologic function [4]. The latest and more attractive engineering approach is based on the use of decellularized extracellular matrix (ECM) [5]. Decellularized xenografts have been identified as potential scaffolds for small-diameter vascular substitutes [6] because they facilitate the remodeling process and fast endothelialization after implantation [7-10]. It is crucial that these grafts possess mechanical properties that allow them to withstand physiological flow and pressure [11]. However, the decellularization process might damage the structure of the vessel and accelerate the elastin deformation and degradation, resulting in vascular graft expansion and even aneurysm formation [12-14]. Cell proliferation after vascular grafting may strengthen the mechanical properties of blood vessels, but the complete cellularization process is very time-consuming, taking approximately 6 months [15]. Therefore, the decellularized vascular grafts need to be strengthened before implantation. Synthetic polymers have also been widely investigated in vascular tissue engineering [16]. Among them, polycaprolactone (PCL) is representative. PCL is a biodegradable polymer material with excellent biocompatibility and good mechanical properties. It has been approved by the Food and Drug Administration (FDA) as a carrier into the body due to its characteristic of easily degradability and is widely used in tissue engineering graft design [17,18]. Fabricating polymers into suitable 3D structures is important for their applications in tissue engineering. Recently, electrospinning (ES) has emerged as an important technology to make tissue-engineering scaffolds with biomimetic nanofibrous structures [19]. ES technology has fabricated various polymers, including polyvinylidene fluoride-co-hexafluoropropylene (PVDF-HFP) [20], polyurethane (PU), polycaprolactone (PCL), and poly- (lactic-co-glycolic acid) (PLGA) into nanofibrous

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2. Materials and methods

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3D vascular grafts [21]. Electrospun nanofibrous scaffolds offer precise porosity and pore size distribution and architecture, and some show similar mechanical properties, such as tensile strength and elastic modulus, as those of native vessels [20, 21]. Furthermore, electrospun nanofibrous scaffolds are beneficial to biological function as well, e.g., compared to solvent cast PVDF-HFP film, electrospun PVDF-HFP nanofibers showed enhanced interaction with endothelial cells [20], less platelet adhesion and aggregation [21]. Next-generation vascular grafts may utilize hybrid materials, combining both synthetic and natural polymers to improve the graft property [22, 23]. In the current study, we generated a type of hybrid tissue-engineered vessel (HTEV) containing a decellularized rat aortic vessel and poly ε-caprolactone (PCL) electrospun fibers to avoid the poor histocompatibility of synthetic materials and enhance the biomechanics of the natural acellular matrix. We used exogenous heparin to modify the intima of acellular vessels to resist platelet aggregation and inhibit thrombosis [24]. The mechanical characterization, pathology and scanning electron microscope (SEM) images of the in vitro grafts were studied. Then, the grafts were studied in vivo using a rat model for long-term applications in direct comparison to no-ES decellularized controls. In this study, we demonstrate that the new designed hybrid grafts with electrospun PCL reinforced the decellularized rat aortic vessels, can resist biomechanical failures and has a much better long-term performance compared to the decellularized rat aortic vessels alone.

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2.1. Preparation of decellularized scaffolds Thoracic aortas were obtained from 8-month-old male Sprague-Dawley (SD) rats (180 g – 200 g, Shanghai Laboratory Animal Center, Shanghai, China) after the rats were killed with pentobarbital sodium. Phosphate buffered saline (PBS) with 1 % penicillin/streptomycin (Gibco-BRL, Grand Island, NY, USA) was used to wash the grafts. The arteries were 30 to 40 mm in length with interior lumen diameters from 1.5 to 2.0 mm. Excess connective tissue was removed and the intercostal artery branches were ligated with 7-0 prolene sutures (Ethicon, Shanghai China). Decellularized rats aortas (DRA) were produced as we previously described with slight modifications [24-26]. PBS with 0.5 % sodium dodecyl sulfate (436143, Sigma-Aldrich, St. Louis, MO, USA), 0.5 % Triton X-100 (0694, Amresco, Solon, OH，USA) and 0.5 % sodium deoxycholate (D6750, Sigma-Aldrich, St. Louis, MO, USA) were prepared for decellularization. Then, the aorta grafts were placed on an orbital shaker (120 cycles per minute) for a total of 96 h, and the solvent was replaced at 6, 12, 24, 48 and 72 h (4 °C). The tissue sections were incubated with DNase (0.2 mg/mL, DN25, Sigma-Aldrich, St. Louis, MO, USA) and RNase ribonuclease A (20 mg/mL, 9001994, Sigma-Aldrich, St. Louis, MO, USA) at 37 °C for 12 h to remove the cellular components. The conduits were thoroughly rinsed with PBS for 8 cycles of 12 h (4 °C) under continuous shaking (100 cycles per minute) and finally stored in PBS at 4 °C and used within 1 week. All animal experiments were performed

ACCEPTED MANUSCRIPT according to protocols approved by the Ethical Committee of Shanghai Jiaotong University.

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2.2. Microscopic study and DNA quantification The DRA and native rat aortas (NRA) were fixed in 10% formaldehyde, dehydrated in graded ethanol and xylene, and subsequently embedded in paraffin. Sections of 5 mm were cut and were stained with hematoxylin and eosin (H&E) by the Masson and Verhoff-von Gieson methods. The specimens were fixed in 2.5% glutaraldehyde and then dehydrated with a graded ethanol series. After gradient dehydration, the dried specimens were sputter-coated with gold before being examined under scanning electron microscopy (SEM). The cross-section and adventitial and intimal surfaces of the DRA and NRA were analyzed with SEM (Hitachi S2520, Japan). The cellular DNA in the aorta was measured using Quant-iT PicoGreen dsDNA Assay Kit (P11496, Invitrogen Inc., Carlsbad, CA, USA).

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2.3. Electrospinning on decellularized scaffolds Polycaprolactone was electrospun (ES-PCL) outside of the decellularized rat aorta to produce the bilayered HTEV. The customized electrospinning apparatus with a rotating cylindrical collector and a doughnut-shape copper electrode (Fig. S1A-B) was applied. First, the decellularized rat aorta was immersed and deployed in PBS solution. Then, a suitable steel mandrel was inserted into the vascular lumen (Fig. S1C). After vacuum freeze-drying (Scientz-10N, Ningbo, China), the decellularized vessel was gently attached onto a steel mandrel while keeping a good tubular structure (Fig. S1D). Then, the steel mandrel was fixed in the collector, which was placed at a distance of 15 cm from the needle tip and rotated at 500 rpm. Next, a 12.5 % (w/v) polycaprolactone (PCL, Mn 80,000 g/mol, Sigma-Aldrich, St. Louis, MO, USA) solution in hexafluoroisopropanol (DaruiFine Chemical Co., Ltd, China) was pumped out via a 21-gauge needle at a flow rate of 0.8 mL/h with a high positive voltage of 10 kV applied. Compared with the conventional electrospinning technology using a needle or tablet as an electrode, this method can greatly reduce the formation time of the external layer [27-29]. The electrospinning process was carried out for 30 minutes to produce an approximately 100-µm thick nanofibrous PCL layer outside of the decellularized vessel. The resultant HTEV (Fig. S1E) was dried in vacuum. The morphology of the surfaces and cross-section of the HTEV was observed through SEM (Hitachi S2520, Japan) as described earlier, and silver sputter-coated samples were used [30]. 2.4. Mechanical testing of scaffolds 2.4.1. Scaffolds thickness measurement The scaffolds were separated into 3 groups: native rat aortas (NRA), decellularized rat aortas (DRA) and hybrid tissue engineering vessels (HTEV). A high-precision slide caliper (Everpower, 557115, Foshan, China) was used to measure the thickness of the scaffolds. Six samples were measured for each group and averaged. 2.4.2. Suture retention strength testing

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A 20 mm-long scaffold was cut into two segments. The 7-0 polypropylene sutures (Ethilon, Ethicon, Inc., USA) were placed on each side of the two segments 1 mm from the edge, separated by 120 degrees. The suture strength was measured on an Instron5969 mechanical tester (Instron Corporation, USA). A constant deflection speed (2 mm/min) was used along the longitudinal axis of the scaffold until the sutures were pulled through the vessel edge, and the maximum suture strength was recorded. Five specimens were tested for each sample and averaged. 2.4.3. Radial and axial tensile testing Tensile tests were performed at room temperature with a deflection speed of 2 mm/min. A special wire hook was manufactured to test the ultimate radial tensile stress. A tubular specimen (length 4 mm) was passed through two hooks, which were fixed at the two test-clamps. The distance between the two clamps was set to 10 mm. Each test was repeated on five specimens. Tubular specimens (length 1 cm) were clamped at their cut ends for the ultimate axial tensile test. The specimens were extended until rupture and the stress-strain curves were recorded. 2.4.4. Burst pressure testing Burst pressure testing was performed by balloon pump pressure (pressure syringe, Allwell Medical Inc., USA). PBS was filled into the 20-ml syringe, which was inserted at the end of the sample tube and the other end was also firmly secured and sealed with 5-0 sutures ligature to prevent leakage. The pressure was increased by 300 mmHg/s until failure or leakage occurred [31]. The burst pressure was recorded as the maximum pressure prior to construct failure or leakage.

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2.5. Heparin immobilization and evaluation Heparin solution (HEP, H4784, Sigma-Aldrich, St. Louis, MO, USA) was dissolved in 0.2 M acetate buffer at pH 4.6 to attain a concentration of 0.5 mg/ml as described previously [22]. The DRA (n = 6) and HTEV (n = 6) grafts were incubated in the above heparin solution to modify the vascular intimal surfaces before implantation. In accordance with this procedure [25, 26], the heparin was adsorbed alternatively for 5 min, with two consecutive adsorption steps being separated by 3 rinsing steps of 15 min with PBS. After the completion of the heparin modifications, the grafts were stored at 4 °C and used immediately for the in vivo study. The immunofluorescence study of the HEP-HTEV was performed as described but with a slight modification [25]. The HTEVs were embedded in an optimum cutting temperature compound (Sakura, 4583, Tokyo, Japan). For the visualization of HEP binding, the sections were incubated with 1 mg/ml FITC-HEP (H-7482, Invitrogen Inc., Carlsbad, CA, USA) in PBS solution. Then, the sections were rinsed with PBS to remove non-adsorbed FITC-HEP and examined in the red channel of the laser confocal scanning microscope (LCSM; Zeiss LSM710). 2.6. Platelet adhesion test Human platelet-rich plasma was obtained from healthy adult volunteers and prepared by centrifugation of 10 ml citrated whole blood. The native rat aorta (NRA group, n = 6), non-heparinized-DRA (DRA group, n = 6) and heparin-modified DRA

ACCEPTED MANUSCRIPT (HEP-DRA group, n = 6) were incubated with 2 mL platelet-rich plasma (1×108 cells/mL) for 1 hour at 37 °C under static conditions. The result of blood platelet adhesion in vitro was also observed through SEM (Hitachi S2520, Japan) as described earlier [24]. The number of adherent platelets was determined by detecting the activity of lactate dehydrogenase (LDH Release Assay Kit, Beyotime, Nantong, China) present after cell lysis as previously described [24].

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2.7. Implantation in a rat model The heparinized DRAs (n = 6) and HTEVs (n = 6) were implanted as new abdominal aorta interposition grafts in 10-week-old SD rats by two end-side anastomoses. The native aorta between the two anastomoses was ligated with 3-0 silk sutures to avoid shunting. The surgical procedure was as previously described with a slight modification [24]. The rats were anesthetized by pentobarbital sodium at 30 mg/kg via intraperitoneal injection. A ventral midline incision was made and the rat intestines were pushed aside with warm wet gauze, wrapped, finally the infrarenal aorta was exposed. After systemic administration of 300 IU/kg heparin, both the proximal and distal sides were gently cross-clamped by vascular clamps. The graft was anastomosed with the infrarenal aorta in an end-to-side continuous suturing manner with 8-0 prolene sutures (Ethicon, Shanghai, China) under operating loupes. The surgical wound was closed in layers with absorbable sutures. Rats were monitored during recovery from anesthesia. DRAs and HTEVs were explanted 12 weeks after implantation.

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2.8. Imaging of the grafts in vivo Six weeks after implantation, the rats were anesthetized by pentobarbital sodium as previously described. The velocity of blood flow and the pulse wave were assessed by a Vevo 2100 ultrasound platform (VisualSonics, Canada) equipped with a 21 MHz and a 16 MHz probe to evaluate the function of the implant. Nonsurgical rats were tested as the control. A three-dimensional CT reconstruction of DRAs (n = 3) and HTEVs (n = 3) was also performed at 6 weeks under anesthesia. Iopamidol (Niopam370, Bracco, Italy) was administered via the tail vein at 1 mg/kg•min by micro-pump. Images were acquired using a Siemens micro-CT scanner (Inveon Hybrid Micro-PET/CT, Siemens, Germany), and the patency of the implanted grafts was evaluated. 2.9. Explanation and tissue analysis Twelve weeks after implantation, the recipient rats were killed and the DRAs and HTEVs were explanted for tissue analysis. The tissue samples were fixed in 10% formalin buffer, dehydrated in graded ethanol and xylene, and subsequently embedded in paraffin. Sections of 5 µm were cut and stained with hematoxylin and eosin (H&E). The specimens were also stained immunohistochemically for the presence of endothelial cells with von Willebrand factor (vWF) (ab6994, Abcam, Cambridge, UK), for presence of interstitial cells with α-SMA (ab5694, Abcam, Cambridge, UK), and for inflammatory cells with CD68 (ab125212, Abcam, Cambridge, UK), followed by

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2.10. Statistical analysis All data were expressed as the mean ± SD. The statistical analysis for the determination of the differences in the measured properties between groups was accomplished using Student t test, performed with SPSS version 19.0 (SPSS, Chicago, IL), and p values < 0.05 were considered statistically significant.

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3. Results

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3.1. Histological analysis of NRAs and DRAs Decellularized scaffolds were produced from rat aortas through a decellularization process as previously described [24] by means of osmotic shock, enzyme digestion and mechanical processes. Compared with NRAs (Fig. 1A), the cellular components were not detected by hematoxylin-eosin staining in DRAs, and the intima was also significantly injured (Fig. 1B). The images by the ’s trichrome and Verhoeff-van Gieson staining showed the disorders of vascular media and the destroyed structure of the elastic fiber after decellularization (Fig. 1D & F), compared with fresh rat aorta vessels (Fig. 1C & E). The total DNA content (301.7±15.6 ng/ml) decreased significantly after the decellularization treatment when compared to NRAs (89.7±11.6 ng/ml) (Fig. 1G; p< 0.05). The SEM images revealed that the histoarchitecture of decellularized grafts was maintained the same as NRA (Fig. 2A), but the cross section showed the middle fiberboard was disconnected (Fig. 2B). The adventitia of the DRAs lost its fine fibers after decellularization compared with that of the NRAs (Fig. 2C, D), and the intimal surface was also rougher than that of the NRAs (Fig. 2E, F).

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3.2. HTEVs fabrication and characterization HTEVs were prepared by ES-PCL out of the DRAs. The morphology and dimensions of the HTEVs were characterized by SEM (Fig. 3A,B). The external PCL layer showed a homogeneous fibrous morphology (Fig. 3A-C) similar to that of the native vessels (Fig. 2C), while its structure was more compact (Fig. 3C-E). Furthermore, the two layers were tightly attached to each other without apparent delamination (Fig. 3D, E). The inner diameter was approximately 1.5 mm, and the wall thickness of the PCL nanofibers was approximately 100-150 µm (Fig. 3D, E). A 2 cm long HTEV was then dried under vacuum and sterilized by ethylene oxide before allogeneic implantation (Fig. S1E & Fig. 3F). 3.3. Assessment of heparin immobilization The surface morphology of heparin-modified graft was observed through fluorescent immunohistochemistry by confocal microscopy (Zeiss LSM710). The

ACCEPTED MANUSCRIPT grafts under white laser showed the outline of the HTEVs and indicated the stratification of the vascular elastic fibers, but densified the PCL nanofiber outside the decellularized matrix (Fig. 4A). The red fluorescence distributed on the surface of the tissue cross-section indicates the presence of HEP bound to the decellularized rat aorta (Fig. 4B). The HEP-FITC did not stain the outer layer of the PCL nanofiber (Fig. 4C).

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3.4. Platelet adhesion test Human platelet-rich plasma was prepared and used to test the adherence of platelets as previously described [24]. Compared with the NRA group and the HEP-DRA group, more platelets adhered to non-heparin-modified DRAs (Fig. 4). Platelets that adhered to non-heparin DRA demonstrated aggregate and changed morphology (Fig. 5B). In contrast, the platelets maintained a spheroid shape in the other two groups, indicating that they were less activated. The number of platelets that adhered to graft intimal surface was further quantified by measuring the LDH activity. Heparin modification significantly reduced the number of platelets that adhered to the HEP-DRAs (OD value = 0.30±0.05, p0.05) (Fig. 4D).

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3.5. Mechanical properties The wall thicknesses, suture retention strength, radial and axial tensile testing, and burst pressure test of the vessels were measured to characterize the mechanical properties of different types of grafts (Fig. 6), and the typical stress-strain curves were shown in Fig. S2. The rat aorta was significantly affected by the decellularization in terms of wall thicknesses, suture retention strength and ultimate tensile strength. PCL is a linear semi-crystalline polymer with a high toughness [22]. Accordingly, the external ES-PCL layer significantly enhanced the mechanical strength of HTEV graft compared to the DRA. The wall thicknesses of decellularized vessels (137±7 µm) were significantly lower than those of NRAs (170±8 µm) and HTEV grafts (192±12 µm) (Fig. 6A). The suture strength of the HTEV was approximately 1.5 times higher than that of the decellularized vessels (Fig. 6B). The decellularization significantly decreased the tensile stress, and the ES-PCL improved the mechanical strength of the HTEV graft (Fig. 6C, D). A typical stress-strain curve was shown in Fig. S2 A and B. There was no significance difference in the ultimate radial strain among the three groups (Fig. S2C). However, the electrospinning evidently enhanced the graft axial strain (Fig. S2D). The tensile strength properties ensure that the PCL/DRA hybrid vascular grafts can satisfy the requirement for artificial replacement. The burst pressure of DRAs was slightly lower than that of NRAs without being significantly different. The HTEVs could resist an average of 2060 ± 127 mmHg of burst pressure, which was significantly higher than that of the DRAs (1465 ± 65 mmHg) and NRAs (1688 ± 114 mmHg) (Fig. 6E). Overall, the hybrid grafts exhibited superior mechanical properties and could be practically sutured to allograft rat aortic models without rupture.

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3.6. Operative results and patency evaluation The DRAs were grafted into10-week-old SD rat abdominal aortas as new abdominal interposition aortas. The native aorta between the two anastomoses was ligated to avoid shunting. The HTEVs were also implanted into rat abdominal aortas and showed good performance of suturing without bleeding. The mean operation time was 78 ± 15 min. All of the operations were successfully accomplished with no postoperative complications. At 6 weeks post-implantation, the patency of the implanted grafts was evaluated by a Vevo 2100 ultrasound platform. Doppler echography demonstrated that all of the grafts remained patent (Fig. 7A). There was no blood flow in the native aorta between the two anastomoses and the anastomosis was patent but with slow blood flow in the DRA group (Fig. 7B). The maximum flow rate of the DRA grafts was not significantly lower than that of the HTEV grafts, but it was significantly lower than that of the autologous abdominal aorta (AAA) (Fig. 7F). At the same time, the grafts were examined by micro-CT. The results indicated 100% patency and no evidence of stenosis. CT images show a larger diameter of DRA grafts compared with HTEV and AAA (Fig. 7G). One case of DRA grafts displayed local vascular bulging with pseudo-aneurysm (Fig. 7C, D). However, the HTEV graft was patent with no damage both in terms of CT image (Fig. 7E) and pathological anatomy after implantation (Fig. 8).

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3.7. Histology and immunohistochemistry after implantation Twelve weeks after implantation, the grafts were retrieved for histology and analyses. The H&E staining of both groups showed no evidence of stenosis and thrombosis. The fiber structure of the DRA was obviously destroyed 12 weeks after implantation, while that of HTEV was intact (Fig. 8A, B). Immunohistochemistry staining of the DRAs reveals slightly more vWF+ cells in the intima compare to the HTEV group (Fig. 8C, D). There was a small number of α-SMA+ cells located in the vascular graft media in both group (Fig. 8E, F). In contrast, a large amount of cells resident in the adventitia of both groups were positive for α-SMA staining. However, the α-SMA+ on the DRA adventitia was haphazard in arrangement, while it was neatly lined in fusiform in HTEV group (Fig. 8F), which indicated that PCL may induce muscle fiber growth. To identify the inflammatory response of grafts, staining against CD68 was performed. The CD68+ cells out of the DRA graft largely showed disorder. The results also indicated that there were numbers of CD68+ cells outside the HTEV grafts, but a clear separation zone was found between the vessel adventitia and the CD68+ cells (Fig. 8H), showing that ES-PCL reduced extravascular inflammatory cell infiltration.

4. Discussion An ideal SDVG, which is a major clinical essential in blood vessel replacement surgery, should possess at least three properties: a confluent endothelium, differentiated quiescent smooth muscle cells (SMCs), and sufficient mechanical

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integrity and elasticity [1, 32, 33]. Many different approaches, including decellularized xenografts, natural scaffolds containing one or more ECM proteins and degradable polymeric scaffolds, have been examined [34]. However, decellularization damaged the extracellular matrix to some extent (Fig. 1,2) and decreased the graft mechanical properties (Fig. 6). The optimal graft should also allow cells to proliferate, migrate and synthesize their own ECM [35]. PCL, which is a type of inexpensive, biocompatible and biodegradable polymer materials, has been widely used in biomedical engineering and is already FDA approved for clinical applications (e.g., drug delivery system) [18,19]. PCL has superior mechanical, rheological and viscoelastic properties and could be readily fabricated into various implants. PCL has also been shown to be a promising substitute for native grafts [36]. Together with a wide range of 3D structure designs and surface modifications [22, 35], PCL grafts could create an instructive microenvironment to promote the cell functionalization and tissue organization of the vasculature [36]. In view of all of these advantages, we chose PCL to fabricate the external layer of HTEV grafts. Electrospinning (ES) is an efficient technique to produce nanofibrous scaffolds, which mimic the structure of native arteries. The mechanical properties of electrospun scaffolds are usually higher than other porous structures. It could be easily adapted to produce PCL nanofibrous scaffolds, the tensile properties of which were comparable to those of natural arteries [37]. We designed and customized an electrospinning device (Fig. S1A) for this study. First, we manufactured a series of cylindrical receivers (steel mandrels) with various diameters to match diverse decellularized vessels. The receivers could be rotated at a wide range of rates (300-1500 rpm) (Fig. S1A). Second, the doughnut-shaped copper electrode connected to the injection part (Fig. S1B) was designed to improve the efficiency. When a conventional needle or tablet was used as the electrode, the PCL nanofibers that were produced hit various fields of the collectors due to the electrostatic repulsion between charged fibers, and only a small fraction of the PCL nanofibers could collected by the decellularized vessel. Thus, it took a long time to produce the external PCL layer with a sufficient thickness (typically 2 h for 100 µm) [27-29]. The doughnut-shape copper electrode yielded a confined electric field, which constrained the fibers to ensure that the vast majority of them deposited and coated around the decellularized vessel placed in the center of the electric field. Thus, the utilization of the produced PCL fibers was greatly enhanced. We could produce the desired PCL layer with a thickness of 100 µm within 30 min. For the design of the HTEVs, the main purpose is to obtain a vascular graft with both sufficient mechanical properties and optimal hemocompatibility. We fabricated a hybrid TEVG using the synthetic polymer PCL and DRAs by the electrospinning technique and subsequently illustrated the advantages of its mechanical properties (Fig. 6 and Fig. S2). We found that decellularization significantly decreased the radial and axial tensile stress and that the ES-PCL improved the mechanical strength of the HTEV (Fig. 6C, D). The electrospinning of PCL evidently enhanced the axial strain and burst pressure of the graft. The HTEV made of ES-PCL/DRA had good tensile properties in vitro, which can satisfy the

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requirement for artificial replacement. It is an important to incorporate the natural anticoagulant heparin into the scaffold. Heparin has been immobilized on the tissue engineering scaffolds by means of covalent bonding or layer-by-layer assembly [38-40]. We used exogenous heparin to modify the endometrial stromal of the decellularized vessels to resist platelet aggregation, such as by layer-by-layer assembly (Fig. 4) [22]. The effects of heparin functionalization on the platelet adhesion were investigated by SEM (Fig. 5A-C) and an LDH Release Assay (Fig. 5D). Immobilized heparin can support and promote the growth of endothelial cells by the stabilization effect for VEGF secreted from the cells by heparin and the further stimulation for its secretion [41]. The ideal vascular graft should be biocompatible, resistant to thrombosis, elastic and should match the compliance with the native host vessel [42]. Different material behavior between the native and the graft vessels and a mismatch at the anastomosis could in fact induce a local flow disturbance and may facilitate thrombus formation and endothelial hyperplasia [43]. The decellularization process itself might damage the structure of vasculature and accelerate the elastin deformation, so we paid more attention to strengthening the vascular wall tension strength to prevent the dilation of the blood vessels. In fact, postoperative ultrasound and CT angiography confirmed that our hypothesis was right. The 8-week results of Doppler echography demonstrated that there was no intimal hyperplasia and thrombosis in either the DRA group or the HTEV group. However, CT images 12 weeks after implantation showed larger diameter DRA grafts, even pseudo-aneurysm formation, compared with HTEV grafts (Fig. 7). Grafts were pathologically examined 12 weeks after implantation. The fiber structure of the DRA graft was obviously destroyed (Fig. 8A), but it was intact in the HTEV group (Fig. 8B), showing the advantage of ES-PCL outside the decellularized matrix. The Walpoth and Moeller groups indicated that electrospun PCL grafts were degraded to 20% of original molecular weight at 18 months post-implantation, but the grafts neither dilated nor had a significant increase in compliance [44]. Furthermore, α-SMA+ cells, which might consist of SMCs and myofibroblasts, were positive on the HTEV wall with a regular arrangement, suggesting an active vascular remodeling process that involves the synthesis and deposition of ECM proteins (Fig 8F). The rapid and adequate endothelialization on the graft lumen is the major key to achieving long-term TEVG patency [45]. The complete endothelialization at 12 weeks was further confirmed by immunostaining using a vWF antibody. Interestingly, the DRA graft intima was thicker than that of HTEV graft, and also there were more vWF+ cells (Fig. 8C & D). The regular arrangement of α-SMA+ cells out of the HTEV graft indicated that the PCL may promote SMCs and myofibroblasts to synthesize ECM. The mechanism of HTEV mitigating graft intimal hyperplasia may be due to the decreased extravascular inflammatory infiltration, which would lessen intimal injury protected by the restricted vasodilation, which will be our next research point. Decellularized scaffolds seem to be an appealing strategy. However, there is a debate on the optimal level of decellularization as under-decellularization might leave traces of immunogenic cells, while over-decellularization may compromise

ACCEPTED MANUSCRIPT mechanical properties [45]. In the future, greater effort should be made into incorporating an anticoagulant or antiplatelet drug delivery and nanotechnology, which could enhance the biocompatibility and hemocompatibility of synthetic materials.

5. Conclusion

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In the present study, we applied an electrospun nano-PCL coating to a decellularized rat aorta and used exogenous heparin to modify the acellular vascular intima, making a type of HTEV with an excellent performance, including good mechanical properties, physical stability, anticoagulation properties and biological compatibility. Mechanical testing of scaffolds in vitro demonstrated that the ES-PCL out of the DRA provided sufficient and suitable mechanical properties for artery replacement. Implantation in rat abdominal aorta for 12 weeks confirmed freedom from thrombus, better vascular morphology and no expansion or aneurism for the hybrid grafts in comparison with the control. This study provides a new idea for artificial tissue engineering of small vascular grafts.

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The work was financially supported by National Natural Science Foundation of China (81571826) (31330029), the Fundamental Research Funds for the Central Universities (2232014A3-01) and DHU Distinguished Young Professor Program (B201303). We thank Zhai Nan and Yang Jie for help with SEM testing, and Hang Qin for help with fluorescent immunohistochemistry.

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Fig. 1. Light micrograph of a rat aorta before (left) and after (right) decellularization. (A, B) H&E stain, (C, D) Masson and (E, F) Verhoff-von Gieson stain. The endothelial cells and vascular smooth muscle cells within the vessels were completely removed using the decellularization protocol, leaving a partially destroyed and disordered fiber structure, and the intima was also significantly impaired. (G) DNA quantification analysis of native rat aortas (NRAs) showed a significantly lower value of OD compare with that of the decellularized rat aortas (DRAs). *Corresponds to a p< 0.05 of DRAs in comparison with NRAs. Scale bars = 100 µm (A~F).

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Fig. 2. Scanning electron microscopy (SEM) before (left) and after (right) decellularization. (A, B) SEM images of the cross section, (C, D) adventitial and (E, F) intima. Upper right corner images show the whole cross section, adventitial and intima of the vessel, respectively. Decellularization made the vascular fiber rough and the gap between the fibers large, and the intima became less smooth. Scale bars = 50 µm (A~F); Scale bars = 500 µm (Upper right corner of A and B); Scale bars = 100 µm (Upper right corner of C-F). Fig. 3. Structure characterization of electrospun PCL hybrid vascular grafts. (A, B) SEM images of nano-PCL fibers and (C) cross section of nano-PCL fibers. (D, E) SEM images of cross section of HTEV grafts. (F) Digital optical image of PCL/DRA hybrid vascular grafts.

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Fig. 4. The fluorescence image of HEP on the surface of the DRA. (A) Fiber skeleton of DRA under white laser. (B) The red fluorescence indicated the presence of HEP bound to the DRA. (C) A merged image of DRA and florescent HEP. Scale bars = 100 µm. Fig. 5. Effect of immobilization on anti-platelet adhesion heparin. (A) SEM images show the adhesion on fresh rat aorta, (B) un-heparin-modified DRA and (C) HEP-DRA. (D) Quantification of lactate dehydrogenase activity among three groups. *Corresponds to a p