J Mater Sci: Mater Med (2015) 26:112 DOI 10.1007/s10856-015-5448-9

ENGINEERING AND NANO-ENGINEERING APPROACHES FOR MEDICAL DEVICES

Experimental study on the construction of small threedimensional tissue engineered grafts of electrospun poly-e-caprolactone Guang-Chang Zhu • Yong-Quan Gu • Xue Geng • Zeng-Guo Feng • Shu-Wen Zhang Lin Ye • Zhong-Gao Wang

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Received: 16 August 2014 / Accepted: 23 November 2014 Ó Springer Science+Business Media New York 2015

Abstract Studies on three-dimensional tissue engineered graft (3DTEG) have attracted great interest among researchers as they present a means to meet the pressing clinical demand for tissue engineering scaffolds. To explore the feasibility of 3DTEG, high porosity poly-ecaprolactone (PCL) was obtained via the co-electrospinning of polyethylene glycol and PCL, and used to construct small-diameter poly-e-caprolactone–lysine (PCL–LYS–H) scaffolds, whereby heparin was anchored to the scaffold surface by lysine groups. A variety of small-diameter 3DTEG models were constructed with different PCL layers and the mechanical properties of the resulting constructs were evaluated in order to select the best model for 3DTEGs. Bone marrow mononuclear cells were induced and differentiated to endothelial cells (ECs) and smooth muscle cells (SMCs). A 3DTEG (labeled ‘10-4 %’) was successfully produced by the dynamic co-culture of ECs on the PCL–LYS–H scaffolds and SMCs on PCL. The fluorescently labeled cells on the 3DTEG were subsequently observed by laser confocal microscopy, which showed that the ECs and SMCs were embedded in the 3DTEG. Nitric oxide and endothelial nitric oxide synthase assays showed that the ECs behaved normally in the 3DTEG. This study consequently provides a new thread to produce small-diameter tissue engineered grafts, with

G.-C. Zhu  Y.-Q. Gu  S.-W. Zhang  Z.-G. Wang (&) Department of Vascular Surgery, Xuanwu Hospital, Capital Medical University, No. 45 Changchun Street, Xicheng District, Beijing 100053, China e-mail: [email protected] X. Geng  Z.-G. Feng  L. Ye School of Materials Science & Engineering, Beijing Institute of Technology, Beijing 100081, China

excellent mechanical properties, that are perfusable to vasculature and functional cells.

1 Introduction Although the development of cardiovascular surgery has been rapid, it has been somewhat hindered by the poor effect of small-diameter blood vessels (inner diameter \6 mm), especially in the treatment of long-segment infrapopliteal stenotic/occlusive diseases [1, 2]. As cardiovascular disease remains one of the major causes of death and amputation throughout the world, the pressing clinical demand for prosthetic grafts has promoted research in tissue engineered grafts for replacement therapy [3–5]. To date, three technologies have been used internationally to produce tissue engineered grafts, which have been assessed in clinical trials [6]. Firstly, there is a biodegradable scaffold approach. Through this technique a functional artery was constructed by Niklason and colleagues [7] with a polyglycolic acid (PGA) scaffold and cultured smooth muscle cells (SMCs). The resulting graft was then transferred to a porcine saphenous artery, where it remained stable for at least 24 days. However, with this approach the tissue mechanics may be impaired by polymer remnants in the engineered grafts. The second technique is known as the cell self-assembly approach. In the investigation of this method, L’Heureux et al. [8] rolled co-cultured human SMCs, fibroblasts and vitamin C into a functional blood vessel, which was found to mimic a native artery. Although the engineered graft had a strong burst pressure, blood easily infiltrated the cell sheet layers [9]. The third and final approach is known as the ‘body as a bioreactor’ approach. In this way, Shin’Oka et al. [10] constructed a tissue engineered graft with a polycaprolactone–polylactic acid

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(PLA) copolymer reinforced with PGA. Cells from a patient’s peripheral vein were then cultured on the graft for 10 days. The tissue engineered graft was subsequently implanted into the patient’s pulmonary artery, where it remained patent for more than 7 months. Animal studies showed that the grafts transform into mature blood vessels after implantation and grow as the subjects by recruiting SMCs and endothelial cells (ECs) from autologous neighboring vessels [11]. However, without a mature extracellular matrix and mechanical strength, the grafts can only be implanted into blood vessels with low pressure. To improve mechanical strength of 3DTEGs, Yuan et al. [12] cocultured ECs, SMCs and fibroblast cells on three sub-regional areas of a polydimethylsiloxane (PDMS) membrane to fabricate 3DTEGs through a stress-induced rolling strategy. Despite good mechanical strength, partitions appeared between neighboring layers of the PDMS membrane under the impact of blood flow. Furthermore, the overlap of different cells resulted in a stress response. In recent years, small-diameter three-dimensional tissue engineered grafts have attracted research attention; however, their assembly and manipulation have proved challenging. In this study, we constructed nine small-diameter 3DTEG models with different poly-e-caprolactone (PCL) membranes and poly-e-caprolactone–lysine (PCL–LYS–H) scaffolds in order to identify the best model for 3DTEGs (Table 1). The identified 3DTEG was then constructed through the co-culture of ECs on the selected PCL–LYS–H scaffold and SMCs on PCL.

2 Materials and methods 2.1 Materials EBM-2 complete medium was purchased from Lonza Group Ltd (Basel, Switzerland). Low-glucose Dulbecco’s modified Eagle’s medium (LG-DMEM), fetal bovine serum (FBS), trypsin and ethylene diamine tetraacetic acid Table 1 3DTEG models fabricated with different PCL membranes and PCL–LYS–H scaffolds PCL–LYS–H scaffolds

PCL membranes 4%

6%

8%

10

10-4 %

10-6 %

10-8 %

20

20-4 %

20-6 %

20-8 %

30

30-4 %

30-6 %

30-8 %

Nine 3DTEG models, labeled ‘10-4 %’, ‘10-6 %’, ‘10-8 %’, ‘204 %’, ‘20-6 %’, ‘20-8 %’, ‘30-4 %’, ‘30-6 %’ and ‘30-8 %’, were fabricated with inner PCL-LYS-H scaffolds (labeled ‘10’, ‘20’ and ‘30’) and outer two-layer sheets of PCL membranes (labeled ‘4 %’, ‘6 %’ and‘8 %’)

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(EDTA) were purchased from Gibco-BRL (Gaithersburg, MD, USA). Gelatin, phosphate-buffered saline (PBS), goat anti-human CD31 polyclonal antibody, fluorescein isothiocyanate (FITC), conjugated mouse anti-goat IgG, rhodamine (TRITC), AffiniPure goat anti-mouse IgG and FicollÒ were purchased from Sigma Chemical Co. (St Louis, MO, USA). Monoclonal mouse anti–human von Willebrand factor (vWF) antibody, monoclonal mouse anti-human smooth muscle actin (a-actin), monoclonal mouse anti-human calponin antibody and monoclonal mouse anti-human smooth muscle myosin heavy chain (SMMHC) antibody were purchased from Dako Corporation (Santa Barbara, CA, USA). Texas red-conjugated goat anti-rabbit IgG was purchased from Jackson ImmunoResearch Laboratories Inc. (West Grove, Pennsylvania, USA). Dioctadecylindocarbocyanine-labeled acetylated low-density lipoprotein (Dil-Ac-LDL) was purchased from Molecular Probes (Junction City, NY, USA). CellTrackerTM Orange (CMTMR) and TRIzolÒ reagent were purchased from Invitrogen (Carlsbad, CA, USA). Calcein-AM was purchased from Dojindo (Kumamoto, Japan). The nitric oxide (NO) assay kit (enzymic method) and the endothelial nitric oxide synthase (eNOS) assay kit were purchased from Nanjing Jiancheng Bioengineering Institute (Nanjing, China). The activated partial thromboplastin time (APTT) kit was purchased from Diagnostic Stago Inc. (Parsippany, NJ, USA). The RevertAidTM First Strand cDNA Synthesis kit was purchased from Fermentas Life Sciences (Ontario, USA). SYBRÒ Green PCR Master Mix was purchased from Applied Biosystems (Foster City, CA, USA). Agarose was purchased from Biowest (Seville, Spain). Fibronectin (Fn) was purchased from Calbiochem (Schwalbach, Germany). SurgiSealÒ was purchased from Adhezion Biomedical, LLC (Wyomissing, PA, USA). PCL (average Mw = 76500) and PCL–LYS–H scaffolds were provided by Beijing Institute of Technology. Beagle canines were provided by Beijing Keyu Experimental Animal Cultivation Center [licensure: SCXK (Jing) 2012-0004]. All procedures were approved by the Animal Care and Use Committee. 2.2 PCL–LYS–H scaffold fabrication and detection of the anticoagulant effect A solution of PCL (10 % m/v) in dichloromethane was heated up to 30 °C for electrospinning at 18 kV direct current power supply. The receiver with a diameter of 4 mm, was a metal rod with a -4 kV voltage, which was 15 cm from the syringe pump. The solution was infused from the syringe pump at a rate of 2 ml/h. The syringe pump moved horizontally very slowly. Meanwhile, the receiver revolved very slowly to get even PCL scaffolds. PCL scaffolds (inner diameter = 4 mm) were electrospun

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for 10 min, 20 min and 30 min (labeled: ‘10’, ‘20’, and ‘30’, respectively) and dried in a vacuum at room temperature for 24 h. The scaffolds (length: 5–6 cm) were then immersed in a saturated solution of lysine (0.5 g/ml) and shaken at 37 °C for 5 days. They were then rinsed in deionized water to remove unreacted lysine. PCL–LYS scaffolds were subsequently obtained after drying in a vacuum at room temperature for 24 h. PCL–LYS–H scaffolds were synthesized from the PCL–LYS scaffolds and a solution of heparin sodium (1:2) through dehydration, condensation, rinsing and drying. Heparin was anchored onto the PCL scaffold surface by the lysine groups via an esterification reaction. Heparin is known to have an effect on the coagulation of blood. APTT is a common means to detect the anticoagulant effect of a material’s surface. Consequently, the APTT of our scaffolds was assessed with an APTT kit (used according to the manufacturer’s instructions) and a coagulation analyzer (TECHROMIV PLUS, Nrufahm N.B., Germany). Blood samples were obtained from healthy subjects and dosed with sodium citrate (1.09 mol/l) prior to centrifugation to obtain the plasma. PCL–LYS–H and PCL samples were cut into 1 cm 9 1 cm blocks and attached to the inner side of a quadruple colorimetric cup. The APTT was then assessed for plasma, PCL–LYS–H and PCL (n = 4). 2.3 Co-spinning of PCL and polyethylene glycol (PEG) Solutions of PEG (Mw = 20,000; 4, 6, 8 % m/v) in dichloromethane were labeled ‘4 %’, ‘6 %’ and ‘8 %’, respectively, and co-electrospun with 10 % PCL. Two syringe pumps were connected to a high-voltage (20 kV) direct current power supply. The collection screen, which was 18 cm from the syringe pumps, was a metal plate with a -4 kV voltage. PCL and PEG were extruded from the needle tip at a constant rate by the two syringe pumps, which were parallel to each other with a 4 cm horizontal distance. After 2 h, the membrane was taken out and dried in a vacuum at room temperature for 24 h. It was then rinsed repeatedly in deionized water to remove excess PEG. The membrane was then freeze-dried to obtain a PCL membrane with increased porosity. 2.3.1 Detection of PCL membrane porosity PCL, PCL co-electrospun with PEG and a portion of porcine carotid artery were fixed with 2.5 % glutaraldehyde. The samples were cut into 3 mm 9 3 mm blocks, rinsed with PBS, fixed with 1 % osmic acid, dehydrated gradually with graded ethanol and vacuum-dried prior to being

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sputter-coated with gold. The morphology of the samples was observed by scanning electron microscopy (SEM; Hitachi, TM-1000, Japan) at a voltage of 20 kV. The porosities of PCL and PCL co-electrospun with PEG were determined using a mercury intrusion porosimeter (AutoPoreIV 9510, Micromeritics, USA). Samples were subjected to a pressure of 0.5–60,000 psia in predefined steps. The porcine carotid artery was used as a control. 2.3.2 Analysis of the mechanical properties The mechanical properties of PCL and PCL co-electrospun with PEG were measured using an electronic tension tester (DLL-5000, Caldecott Shanghai Machinery Equipment Co. Ltd, Shanghai, China) at room temperature. The materials were cut into 12 mm 9 20 mm strips, and then stretched at a rate of 100 mm/min. A load–elongation curve was recorded to give the breaking tenacity, elongation at break and maximum force of PCL and PCL co-electrospun with PEG. 2.4 Fabrication of the 3DTEG model PCL membranes, labeled ‘4’, ‘6’ and ‘8 %’, were cut into 5 cm 9 3 cm blocks and rolled into two-layer tubular sheets along an axis diameter of 4 mm. Notably, the sheets remained tubular after withdrawal of the axis. PCL-LYS-H scaffolds (the inner tubes), labeled ‘10’, ‘20’ and ‘30’, were then enclosed by these PCL two-layer tubular sheets. The edges of the outer tubular sheets were bound with SurgiSealÒ glue. Then, nine 3DTEG models (length: 5 cm) were fabricated with different labels: ‘10-4 %’, ‘10-6 %’, ‘10-8 %’, ‘20-4 %’, ‘20-6 %’, ‘20-8 %’, ‘30-4 %’, ‘306 %’ and ‘30-8 %’, according to the composition of the model. 2.4.1 Analysis of the suture strength The nine 3DTEG models were divided and stitched to pieces of cloth (1.2 cm 9 5 cm). The edge distance of the stitch and the distance between the stitches were all 2 mm. The suture strength was measured using an electronic tension tester (DLL-5000) at room temperature. The experimental modality, edge distance of stitch, distance between stitches, and width and thickness of the samples were all recorded electronically. The samples were stretched slowly at a rate of 5 mm/min and a load–elongation curve was recorded to obtain the respective strengths at the point that the samples broke. The suture strength was recorded as the ratio of the breaking tenacity and the stitch number. The suture strength of a porcine carotid artery was measured as a control.

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2.4.2 Analysis of the burst pressure To assess the burst pressure, one end of a 3DTEG model was tied and a long balloon was inserted into the model with a metal rod through the other end. The 3DTEG model and balloon were then connected to a three-armed tube. The other two ends of the three-armed tube were connected, respectively, with a pressure gauge and a mini air compressor (ZB-0.11/7, Zhejiang Wenling Juba Machinery Co. Ltd, China). Air was infused into the balloon at a rate of 3 ml/min. The burst pressure was obtained when the 3DTEG model broke suddenly. Four samples of each model were tested to obtain an average burst pressure and the burst pressure of a porcine carotid artery was measured as a control. 2.5 Isolation and differentiation of bone marrow mononuclear cells to vascular ECs and SMCs 2.5.1 Isolation of bone marrow mononuclear cells (BM MNCs) Bone marrow was obtained from the bilateral posterior superior iliac spines of four beagle canines. The procedure was approved by the Animal Care and Use Committee. Bone marrow was drawn into a 20 ml syringe containing 5 ml heparin sodium solution (100 l/ml). A total of 60 ml bone marrow was obtained from beagle canines. It was subsequently added into four sterile 50 ml centrifuge tubes containing an equal volume of PBS. The resulting bone marrow solutions were transferred to the surface of FicollÒ (1.077 g/ml) in 50 ml centrifuge tubes. MNCs were isolated by density gradient centrifugation at 1,500 rpm for 15 min. The top layer, a dull red liquid, was removed in order to obtain the reddish, cloudy middle layer, which is known to contain the mononuclear cells. This layer was then rinsed with PBS and centrifuged twice at 1,500 rpm for 15 min. The supernatant was discarded to harvest the mononuclear cell pellet.

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with 15 ml EBM-2 every 2–3 days. As much as 80–90 % of the cells fusions, the cells of each flask were sub-cultured into next passage. After discarding the culture fluid and rinsing with PBS, the cells in each flask were digested with 3 ml trypsin (0.25 %)/EDTA (0.02 %) for about 5 min. Then, 10 % FBS (2 ml) was added into each flask to stop digestion. After centrifugation and rinsing, the cells of this first passage were seeded at a density of 1 9 104/cm2 into another flask and incubated at 37 °C, 5 % CO2 and saturated humidity. The growth conditions and morphologic features were observed by inverted microscopy (IX71, Olympus, Japan) every day. Cells following two or three passages were used in this study. 2.5.2.1 Characterization of ECs Cells from the second passage were seeded at the density of 1 9 104/cm2 in 24-well plates pre-coated with gelatin (0.2 %) and incubated at 37 °C for 36 h. The adherent cells were fixed with 4 % paraformaldehyde at room temperature for 15 min and washed with PBS. The fixed cells were then permeated with 0.3 % Triton for 15 min and incubated with non-immune goat serum (10 %) to block nonspecific binding. The primary antibodies (goat anti-human CD31 polyclonal antibody and rabbit anti-human vWF polyclonal antibody) were added to the 24-well plates and the plates then kept at 4 °C overnight. After washing with PBS, the cells were incubated with secondary antibody (FITC-conjugated mouse anti-goat IgG and Texas red-conjugated goat antirabbit IgG) at 37 °C for 2 h. The stained cells were observed by fluorescence microscopy (IX-71, Olympus) with PBS as a negative control. Ten visual fields were randomly chosen to calculate the percentage of double-stained cells. For the DiI-Ac-LDL uptake assay, ECs were seeded in a 35 mm petri dish and cultured at 37 °C for 24 h. According to the manufacturer’s instructions, the cells were incubated with 10 lg/mL of Dil-Ac-LDL at 37 °C for 4 h. After washing twice with PBS, DiI-Ac- LDL uptake was detected by fluorescence microscopy.

2.5.2 Differentiation of BM MNCs into vascular ECs

2.5.3 Differentiation of BM MNCs into vascular SMCs

Culture flasks (75 ml) were pre-coated with 3 ml gelatin (0.2 %) for 2 h. Then, 15 ml EBM-2 [including vascular endothelial growth factor (VEGF), human fibroblastic growth factor b (hFGF-b), insulin-like growth factor 1 (IGF-1), human epidermal growth factor (hEGF), hydrocortisone and ascorbic acid] supplemented with 5 % FBS and 1 % penicillin–streptomycin was added into the flasks. BM MNCs were seeded at a density of 1 9 105/cm2 into the flasks and then incubated at 37 °C, 5 % CO2 and saturated humidity in a HeracellTM 150 incubator (Thermo Electron, Osterode, Germany). Each flask was replenished

BM MNCs were seeded at the density of 1 9 105/cm2 into flasks (75 ml), which were pre-coated with 2 ml Fn solution (20 lg/ml) for 2 h. Then, 15 ml DMEM (including platelet-derived growth factor BB (PDGF-BB), transforming growth factor b1 (TGF-b1) and IGF-1) supplemented with 10 % FBS and 1 % penicillin–streptomycin was added into the flasks. The cells were incubated at 37 °C, 5 % CO2 and saturated humidity in a HeracellTM 150 incubator. Each flask was replenished with 15 ml DMEM every 2–3 days. The sub-culture was performed in a similar manner to that of the ECs.

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2.5.3.1 Characterization of SMCs Cells from the third passage were seeded at a density of 1 9 105/cm2 into a flask (75 ml) pre-coated with gelatin (0.2 %) and incubated at 37 °C, 5 % CO2. The flask was replenished with 15 ml DMEM every 2–3 days. On day 14, cells were seeded at a density of 1 9 104/cm2 in 24-well plates and incubated at 37 °C. The adherent cells were fixed with 4 % paraformaldehyde for 5 min and washed with PBS. The fixed cells were permeated with 0.3 % Triton for 15 min and incubated with non-immune goat serum (10 %). After removing the goat serum, monoclonal mouse anti-human smooth muscle actin (1:400), monoclonal mouse anti-human calponin antibody (1:250) and monoclonal mouse anti-human SMMHC antibody (1:250) were added separately to the 24-well plates, which were kept at 4 °C overnight. After removing the primary antibodies and washing with PBS, the cells were incubated with secondary antibody at 37 °C for 2 h. Cell nuclei were counter-stained with 4,6-diamino-2-phenylindole (DAPI) after removal of the secondary antibodies and washing with PBS. The stained cells were observed by fluorescence microscopy (IX-71, Olympus) with PBS as a negative control. The cytoskeletal protein smooth muscle 22 alpha (SM22a) is expressed specifically in vascular SMCs. The expressionof SM22a was detected by real-time quantitative PCR. SMCs were seeded into a flask (75 ml) with DMEM supplemented with 10 % FBS and incubated at 37 °C, 5 % CO2 and saturated humidity in a HeracellTM 150 incubator (Thermo Electron) for 8 h. Then, SMCs were cultured with serum-free DMEM for 24 h to starve the cells. About 1 9 107 SMCs were collected for extraction of their RNA using TRIzolÒ reagent (Invitrogen), which was used according to the manufacturer’s instructions. cDNA synthesis was performed with a RevertAidTM First Strand cDNA Synthesis kit (Fermentas Life Sciences) primed with oligo (dT). For the real-time PCR, 2 ll template cDNA was mixed with 10 ll 2 9 SYBRÒ Green PCR Master Mix and 1 ll each of the forward and reverse primers to obtain final volumes of 20 ll. The primer sequences for SM22a were forward: 50 -GAAGTTTCTCCAGTTGTC-30 and reverse: 50 -ATACTTCTTTCGCGTCCTCG-30 . Amplifications were performed using ABI Prism 7700 (Applied Biosystems, Foster City, CA, USA) with the following temperature profile: a pre-denaturation step of 3 min at 95 °C followed by 40 cycles of denaturation (30 s at 94 °C), annealing (30 s at 60 °C) and extension (45 s at 72 °C), and a last extension at 72 °C for 5 min. The thermal denaturing step was performed to generate the dissociation curves to verify amplification specificity followed by 1.5 % agarose gel electrophoresis. A glyceraldehyde 3-phosphate dehydrogenase (GAPDH) control primer probe was used for normalization as follows: forward primer 50 -GGTCACCAGGGCTGCT TT-30 , reverse primer 50 -ATTTGATGTTGGCGG GAT-30 .

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2.6 Design, construction and analysis of the 3DTEG 2.6.1 Design of the 3DTEG For construction of the 3DTEG, ECs and SMCs derived from BM MNCs were labeled with CellTrackerTM Orange (CMTR) and Calcein-AM. Then, the ECs were seeded on PCL–LYS–H scaffolds labeled ‘10’ and SMCs were seeded on tubular PCL sheets labeled ‘4 %’. The PCL–LYS–H scaffolds treated with ECs were then enclosed by tubular PCL sheets containing SMCs. The edges of the outer tubular sheets were bound with SurgiSealÒ glue to construct the final 3DTEG. After dynamic co-culture for a week, the fluorescently labeled cells on the 3DTEG were observed by laser confocal microscopy to explore the growth of the ECs and SMCs. Assays of the DNA index (DI), NO and eNOS activity of the ECs were performed to evaluate the feasibility of the 3DTEG construct (Fig. 1).

2.6.2 Construction of the 3DTEG Sixteen PCL–LYS–H scaffolds, labeled ‘10’, which were sterilized with epoxyethane, were coated with Fn (20 lg/ ml) for 2 h. They were then all enclosed by sterilized tubular PCL sheets and placed in four culture flasks (25 cm2) that were not pre-coated with gelatin. Eight sterilized tubular PCL sheets, labeled ‘4 %’, which were pre-coated with Fn, were placed in another four culture flasks (25 cm2) without gelatin pre-coating. ECs from a second passage were added into a sterile 50 ml centrifuge tube containing 1 ml EBM-2. Then, 1 ml CellTrackerTM Orange (5 lM) was added into the centrifuge tube, which was incubated at 37 °C, 5 % CO2 and saturated humidity for 60 min in a HeracellTM 150 incubator. The dyed cells were washed with PBS and centrifuged twice at 1,500 rpm for 15 min. The supernatant was discarded to harvest the cell pellet, which was then treated with 1 ml EBM-2. Then, 2 ll of the suspension of dyed cells was spread on a glass slide and observed by fluorescence microscopy (IX-71, Olympus; excitation wavelength: 561 nm, emission wavelength: 595 nm). The dyed ECs (2.5 9 105) were seeded into the above culture flasks (25 cm2) with PCL–LYS–H scaffolds, followed by the addition of 15 ml EBM-2 (including VEGF, hFGF-b, IGF-1, hEGF, hydrocortisone and ascorbic acid) supplemented with 5 % FBS and 1 % penicillin–streptomycin. They were incubated at 37 °C, 5 % CO2 and saturated humidity for 24 h. To evenly spread the ECs on the surface of the PCL–LYS–H scaffolds, the flasks were turned 90° clockwise every 2 h during the first 12 h.

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Fig. 1 Pictorial representation of 3DTEG construction. a The left image demonstrates that ECs labeled with CellTrackerTM Orange can be seeded on PCL–LYS–H scaffolds; the right image demonstrates that SMCs labeled with Calcein-AM can be seeded on tubular PCL sheets. b The left shows a PCL–LYS–H scaffold with ECs; the right shows a tubular PCL sheet with SMCs. c The left image demonstrates

that the PCL–LYS–H scaffold with ECs can be enclosed by the tubular PCL sheet with SMCs; the right image shows a PCL–LYS–H scaffold with ECs enclosed by a tubular PCL sheet with SMCs. d The right image shows the constructed 3DTEG; the left image demonstrates a transect of 3DTEG (Color figure online)

SMCs from the second passage were added into another 50 ml centrifuge tube containing 1 ml DMEM. Calcein-AM (1 ml, 15 lM) was added into the centrifuge tube, which was incubated at 37 °C, 5 % CO2 and saturated humidity for 90 min. The dyed cells were washed twice with 45 ml PBS. The supernatant was discarded to harvest the dyed cells, which were treated with 1 ml DMEM. Then, 2 ll of the suspension of dyed SMCs was spread on a glass slide and observed by fluorescence microscopy (IX-71; excitation wavelength: 490 nm, emission wavelength: 515 nm). The dyed ECs (5 9 105) were seeded into the above culture flasks with tubular PCL sheets, followed by addition of 15 ml DMEM (including TGF-b, PDGF and IGF-1) supplemented with 10 % FBS and 1 % penicillin–streptomycin. They were incubated at 37 °C, 5 % CO2 and saturated humidity for 24 h, and turned 90° clockwise periodically to spread the SMCs evenly on the surface of the PCL sheets. The PCL–LYS–H scaffolds seeded with dyed ECs and the tubular PCL sheets seeded with dyed SMCs were taken out separately from the culture flasks (25 cm2) using tweezers. The tubular PCL sheets that enclosed the PCL– LYS–H scaffolds were stripped off to remove the ECs on the outer surface of the scaffolds. The tubular PCL sheets, which were seeded with dyed SMCs, were unwound to double enclose the PCL–LYS–H scaffolds, which were seeded with dyed ECs. The edges of the outer tubular sheets were bound with SurgiSealÒ glue to construct 3DTEGs. Eight tubular PCL sheets, which were not seeded with dyed SMCs, were unwound to double enclose the PCL–LYS–H scaffolds seeded with dyed ECs as a control.

Each 3DTEG was implanted in a sterile pipe, in which 15 ml DMEM (including TGF-b, PDGF and IGF-1) supplemented with 10 % FBS and 1 % penicillin–streptomycin was added. The DMEM medium in the pipe was driven by a peristaltic pump (BT100-1L-A, Baoding Longer Precision Pump Co. Ltd, Baoding, China). The flow rate of the DMEM medium was increased by 0.5 ml/ min every 4 h from 0.5 ml/min to 7 ml/min (15 dyn/cm2). The pipe and peristaltic pump were incubated at 37 °C, 5 % CO2 and saturated humidity for 1 week. The DMEM medium was replenished with 15 ml EBM-2 every 2 days and closely observed for color variation to prevent contamination by bacteria. Four 3DTEGs were taken from the pipes to harvest ECs via trypsination [trypsin (0.25 %)/EDTA (0.02 %)] on day three and seven and stored at -80 °C. Before harvesting the ECs, 0.5 cm sample segments of each 3DTEG were cut out and stored at -80 °C to make frozen sections. Controls were prepared in the same way.

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2.6.3 Observation of the 3DTEG by laser confocal microscopy Frozen Sections (20 lm thick) of the 3DTEGs were cut with a freezing microtome (CM1900, Leica, Germany) at -20 °C and mounted on 10 % glycerin-coated glass slides. The sections were store at -20 °C before use. The sections were examined using laser confocal microscopy (LSM780,ZEISS, Germany; excitation wavelength: 561, 488 nm) to explore the growth of ECs and SMCs.

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2.6.4 NO, eNOS and DNA index assays NO and eNOS activity, and the DI of the ECs were evaluated to explore the function of the ECs in the 3DTEG before and on days three and seven after construction of the 3DTEGs. For biochemical analysis, an ECs suspension was prepared with 0.9 % sodium chloride at a density of 1 9 106/ ml. Then, the ECs were broken up by ultrasonication and centrifuged at 3,000 rpm/min for 15 min. The supernatant was used to measure the levels of NO and eNOS activity, according to the manufacturer’s protocol. The NO content of the ECs was measured by a NO assay kit. The eNOS assay was performed using a NO synthase assay kit. The absorbance of the supernatant was determined using an ultraviolet spectrophotometer (Hitachi U-2800, Tokyo, Japan) at 530 nm. A suspension of ECs (100 ll) was permeabilized with 70 % ethanol for 60 min and deposited on 300-mesh copper grids. The grids were rinsed with PBS and centrifuged repeatedly at 1,000 rpm for 5 min to isolate the nuclei. Then, the DNA was stained with 1 ml propidium iodide (PI; 0.5 %) for 30 min. The DI and DNA histograms were generated using a fluorescence-activated cell sorter420 (FACS-420, Becton–Dickinson, Mountain View, CA, USA). The coefficient of variation (CV; \2.0 %) was set according to microspheres from a standard sample before analysis. Samples with a dispersion index (DI) ranging from 0.89 to 1.14 were classified as diploid [13].

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and 8 % (m/v) solutions of PEG to fabricate three different PCL membranes, labeled ‘4 %’ (thickness: 0.11–0.25 mm), ‘6 %’ (thickness: 0.16–0.3 mm) and ‘8 %’ (thickness: 0.15–0.25 mm), as shown in Fig. 2. The assessment of the APTT showed that the APTTs for plasma from healthy subjects, PCL and the PCL–LYS–H scaffolds were 43.2 ± 0.3, 40.1 ± 0.5 and 51.45 ± 0.3 s, respectively. This indicated that PCL itself cannot prolong the APTT, while PCL–LYS–H can. It also showed that heparin on the surface of PCL–LYS–H scaffolds has the expected anticoagulant effect. The porosities of PCL co-electrospun with 4, 6 and 8 % solutions of PEG were 91.18, 88.73 and 89.25 %, respectively. All the samples had greater porosity than the porcine carotid artery control (64.2 %) and PCL alone (83 %). The greater porosity may be beneficial to the attachment and embedding of cells. The mechanical properties of PCL co-electrospun with PEG had some degree change as compared with plain PCL. The tensile behavior of PCL co-electrospun with 4 and 6 % solutions of PEG fell slightly (P [ 0.05), while PCL coelectrospun with an 8 % solution of PEG had an increased breaking tenacity, elongation at break and maximum force (Table 2). 3.2 Assessment of the 3DTEG model Triple-layered 3DTEG models were successfully fabricated and all the models displayed similar softness, ductility and a smooth lumen.

2.7 Statistical analysis 3.2.1 Suture strength All the analyses were performed with SPSS version 19.01 software (SPSS Inc., Chicago, IL, USA). The data are expressed as mean ± standard deviation (SD). A comparison of two groups was analyzed by an independent-sample test. The mean comparison in the groups was analyzed by one-way analysis of variance (ANOVA). Differences were considered significant when the p value was \0.05.

The analysis of the suture strength showed that the 3DTEG models had higher suture strengths than the porcine carotid artery control (0.75 ± 0.04 N). The suture strengths of the 3DTEGs labeled ‘10-6 %’, ‘20-4 %’ and ‘20-6 %’ were significantly higher than that of the porcine carotid artery control (P \ 0.05; Table 3). 3.2.2 Burst pressure

3 Results 3.1 PCL–LYS–H scaffolds and PCL co-electrospun with PEG Three different PCL–LYS–H scaffolds, which were labeled ‘10’ (wall thickness: 0.3–0.5 mm), ‘20’ (wall thickness: 0.6–0.7 mm) and ‘30’ (wall thickness: 0.78–0.9 mm), were successfully fabricated. PCL was co-electrospun with 4, 6

The analysis of the burst pressure showed that the burst pressure of the 3DTEG model labeled ‘10-4 %’ was significantly higher (548.00 ± 9.97 kPa) than that of the porcine carotid artery control (443.50 ± 22.25 kPa) and the other 3DTEG models (P = 0.001). There was no significant difference between the porcine carotid artery control and the 3DTEG models with the exception of ‘104 %’(P [ 0.05; Table 4; Fig. 3).

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Fig. 2 The fabricated PCL membrane and PCL–LYS–H scaffold. a PCL membrane ‘4 %’. b PCL–LYS–H scaffold ‘10’ Table 2 Mechanical property of PCL and PCL coelectrospun with PEG Elongation at break (%) (M ± SD)

Breaking tenacity (MPa) P

(M ± SD)

Maximum force (MPa) P

(M ± SD)

P

0 % PEG

626.16 ± 98.07

–

0.86 ± 0.02

–

2.38 ± 0.15

–

4 % PEG

600.00 ± 34.43

0.654

1.27 ± 0.72

0.484

1.88 ± 0.40

0.207

6 % PEG 8 % PEG

546.1 ± 105.58 780.80 ± 35.70

0.166

1.06 ± 0.34

0.712

2.14 ± 0.41

0.524

0.014

2.43 ± 1.13

0.012

3.31 ± 0.71

0.024

0 % PEG is PCL that was not coelectrospun with PEG. 4, 6 and 8 % PEG represent PCL that were coelectrospun with 4, 6 and 8 % PEG. MPa is megapascals. The p value is given by comparison with 0 % PEG

Table 3 Suture strength of 3DTEG models and PCA

Table 4 Burst pressure of 3DTEG models and PCA

Group

Suture strength (N) (M ± SD)

P

Group

Burst pressure (kPa) (Mean ± SD)

P

PCA

0.75 ± 0.04

–

PCA

443.50 ± 22.25

–

10-4 %

0.82 ± 0.23

0.853

10- 4 %

548.00 ± 9.97

0.001

10-6 %

2.12 ± 0.87

0.001

10-6 %

455.75 ± 47.27

0.673

10-8 %

1.37 ± 0.92

0.104

10-8 %

444.00 ± 59.34

0.986

20-4 %

1.68 ± 0.66

0.018

20-4 %

476.75 ± 24.92

0.256

20-6 %

1.69 ± 0.43

0.017

20-6 %

474.68 ± 23.42

0.286

20-8 %

1.28 ± 0.25

0.161

20-8 %

420.00 ± 28.48

0.420

30-4 %

1.21 ± 0.49

0.222

30-4 %

474.00 ± 18.99

0.297

30-6 %

1.48 ± 0.40

0.058

30-6 %

452.00 ± 76.70

0.769

30-8 %

1.28 ± 0.13

0.161

30-8 %

406.00 ± 43.86

0.202

PCA is porcine carotid artery. N is newton. 10-4 %, 10-6 %, 10-8 %, 20-4 %, 20-6 %, 20-8 %, 30-4 %, 30-6 % and 30-8 % represent 3DTEG models with labels ‘‘10-4 %, 10-6 %, 10-8 %, 20-4 %, 20-6 %, 20-8 %, 30-4 %, 30-6 % and 30-8 %’’. The p value is given by comparison with PCA. Mean comparison in groups was analyzed by ANOVA (P = 0.033)

3.3 Identification of ECs and SMCs 3.3.1 Identification of ECs After induction and culture of BM MNCs with EBM-2 for 6–7 days, the cells grew rapidly and most cells proliferated adhesively to form more cell colonies. After approximately

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The p value is given by comparison with PCA. kPa is kilopascal. Mean comparison in groups was analyzed by ANOVA (P = 0.003)

1 week, the cell confluency increased to 80–90 % and a cobblestone morphology was noted; most cells also had a polygonal appearance, as observed by microscopy (Fig. 4). Immunofluorescence staining showed that the cells displayed the characteristic presence of platelet EC adhesion molecule (CD31, or PECAM 1) and von Willebrand factor (vWF), which suggested the phenotypic characteristics of ECs were intact. The percentages of CD31?/vWF? cells were more than 97.5 % in ten visual fields. The cells with

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Fig. 3 Burst pressures of the 3DTEG models and porcine carotid artery (PCA). The burst pressure of the 3DTEG model labeled ‘10-4 %’ was significantly higher than that of the porcine carotid artery control (P = 0.001) and other 3DTEG models (P \ 0.05)

uptake of DiI-Ac-LDL were more than 98 %, which suggested the normal function of ECs (Fig. 5).

3.3.2 Identification of SMCs After induction and culture of BM MNCs with DMEM for 5–7 days, the cells had abundant cytoplasm and stretched to increase their projections; the cells also divided and proliferated rapidly at this point. After 9–11 days of culture, the sparse and dense cell populations interlaced to

Fig. 4 Differentiation of BM MNCs to vascular ECs. a After induction and culture for 3 days; a few cells proliferated adhesively to form cell colonies and most cells were in a floating state. b After

form a ‘‘peaks and valleys’’ type morphology, which was observed by microscopy (Fig. 6). Immunofluorescence staining showed that the cells displayed the characteristic presence of a-actin, calponin and SMMHC. Myofilament-like structures, which were parallel to the longitudinal axes of the cells, were prominent in the cytoplasm at higher magnification. These findings suggested the phenotypic characteristics of SMCs were intact (Fig. 7). After 24 h of serum starvation, SM22a was detected by real-time quantitative PCR. The dissociation curves, which were generated on the basis of amplification and cycle

induction and culture for 9 days; the cells are confluent up to 80–90 % and have a cobblestone morphology

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Fig. 5 Identification of ECs with immunofluorescence staining. The images show that the cells are positive for vWF (a), CD31 (b) and DiI-AcLDL (c)

Fig. 6 Differentiation of BM MNCs to vascular SMCs. a After induction and culture for 3 days a few cells grew to triangular or spindle shapes and proliferated adhesively. b After induction and

culture for 9–11 days; the sparse and dense cell populations interlaced to form a ‘‘peaks and valleys’’ appearance

threshold (CT), showed that the melting temperature (TM) of SM22 was 77.3 °C (GAPDH: 81.7 °C). Agarose gel electrophoresis revealed the appearance of a clear band at 187 bp, which is characteristic of SM22a and a marker gene of contractile SMCs (Fig. 8).

culture for 24 h, ECs labeled with the red fluorescent dye and SMCs labeled with the green fluorescent dye grew well and displayed spindle, triangular and polygon shapes with bright fluorescence, as found by fluorescence microscopy. The labeled ECs and SMCs were seeded on PCL–LYS–H scaffolds labeled ‘10’ and PCL sheets labeled ‘4 %, respectively, and used to successfully construct 3DTEGs. NO and eNOS activity, and the DI of ECs were measured before seeding and on day three and seven after construction of the 3DTEGs.

3.4 Constructed 3DTEGs ECs and SMCs were labeled with CellTrackerTM Orange (5 lM) and Calcein-AM (15 lM), respectively. After

Fig. 7 Identification of SMCs with immunofluorescence staining. The images show that the cells are positive for SMMHC (a), calponin (b) and a-actin (c)

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Fig. 8 The expression of SM22a detected by real-time quantitative PCR. a Amplification plot of SM22a. b Amplification plot of GAPDH. c Dissociation plot of SM22a. d Dissociation plot of GAPDH. e Agarose gel electrophoresis; the left run is the DNA

marker, the middle run is the band of SM22a and the right run is the band of GAPDH. The imaging shows that a clear band appears at 187 bp, which is characteristic of SM22a

3.4.1 Analysis of the 3DTEGs by laser confocal microscopy

dye, which grew on the PCL–LYS–H scaffold. The two outer layers represent the SMCs, tagged with a green fluorescent dye, which grew on the PCL sheets. The ECs with red fluorescence and the SMCs with green fluorescence are clearly visible in Fig. 9b, c. It was possible to view the varyingly labeled cells in the different areas appear by adjustment of the fine focus during microscopy.

The confocal images are shown in Fig. 9. Figure 9a shows a clear outline of the 3DTEG, in which two colors form three concentric circles. The innermost ring of the three circles represents the ECs, tagged with a red fluorescent

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J Mater Sci: Mater Med (2015) 26:112 Table 5 NO assay of ECs before and after construction of 3DTEGs Group

Assay of NO (lmol/L) (M ± SD)

P Pre–Post

EG–COG

18.06 ± 2.83

–

–

EG

21.51 ± 3.20

0.337

0.565

COG

20.43 ± 2.46

0.378

EG

29.10 ± 4.77

0.001

COG

23.19 ± 2.76

0.027

Pre Day 3

Day 7 0.043

EG is experimental group. COG is control group. Pre is before construction of 3DTEGs. Post is after construction of 3DTEGs

Fig. 9 Confocal microscopyof the 3DTEG. a The innermost circle shows ECs on the PCL–LYS–H scaffold. The two outer circles show the SMCs on PCL sheets. By adjustment of the fine focus, red-labeled ECs are clearly visible on the innermost circle (b), and green-labeled SMCs are obvious on the two outer circles (c) (Color figure online)

3.4.2 NO assay The level of NO (29.10 ± 4.77 lmol/l) on day seven after construction of the 3DTEGs was significantly higher than prior to seeding (18.06 ± 2.83 lmol/l) and on day three post-construction (21.51 ± 3.20 lmol/l) of the 3DTEGs (P = 0.001 and P = 0.018, respectively). However, the levels of NO did not change significantly before seeding and on day three after 3DTEG construction (P = 0.337). The NO level of the experimental group was significantly higher than that of the control group (23.19 ± 2.76 lmol/l) on day seven after construction of the 3DTEGs (P = 0.043). While the difference was not significant between day three and day seven in the control group (P = 0.275), as shown in Table 5 and Fig. 10. 3.4.3 eNOS activity The eNOS activity on day seven after construction of the 3DTEGs (2.081 ± 0.346 U/ml) was significantly higher than before construction (1.124 ± 0.238 U/ml; P = 0.001). However, it was not significantly higher than on day three after construction of the 3DTEGs (1.641 ± 0.271 U/ml; P = 0.079) and the eNOS activity on day three after construction was significantly higher than before construction

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Fig. 10 NO assay of ECs before and after construction of the 3DTEGs. Before construction of the 3DTEGs, ECs were derived from the same batch of cells for both the experimental and control groups. Consequently, the mean (M) and standard deviation (SD) of the NO content were equal in the experimental and control groups before construction

(P = 0.037). Indeed, there were no significant difference between day three and day seven after construction of the 3DTEGs in both the experimental and the control groups (P = 0.079 and P = 0.498, respectively). However, the difference was significant between the experimental group and control group (1.648 ± 0.201 U/ml) on day seven after 3DTEG construction (P = 0.042; Table 6). 3.4.4 DNA index (DI) assay Flow cytometry was used to assess the diploid apex, which showed that the ECs had two complete sets of

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chromosomes before and after construction of the 3DTEGs. There were no significant differences in the DI before construction (1.006 ± 0.078) and on day three (1.048 ± 0.051) or seven (1.063 ± 0.071) after construction of the 3DTEGs in the experimental group (P [ 0.05). There were also no significant difference between the experimental and control groups (1.035 ± 0.028 and 1.063 ± 0.071, respectively) on day three and seven postconstruction (P [ 0.05). The DI of the ECs ranged from 0.9 to 1.1. These findings suggested that the ECs were in the G0/G1 phase (Table 7; Fig. 11).

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Table 7 DI of ECs before and after construction of 3DTEGs Group

DNA Index (M ± SD)

P Pre–Post

EG–COG

1.006 ± 0.078

–

–

EG

1.048 ± 0.051

0.886

0.998

COG

1.035 ± 0.028

0.969

Pre Day 3

Day 7 EG

1.063 ± 0.071

0.726

COG

1.086 ± 0.078

0.432

0.985

There were no significant difference in DI between day 3 and day 7 after construction of 3DTEGs in both experimental and control group (P = 0.997, P = 0.788)

4 Discussion It has been the aim of several studies to construct 3DTEGs as tissue-engineered materials with excellent properties and superior performance. In this regard, PCL, PGA, PLA and polyhydroxybutyrate (PHB) have been popularly used as biodegradable materials with satisfactory tissue compatibility and plasticity that can be readily degraded biochemically into smaller molecules. Moreover, the degradation rate and three dimensional structures of these polymers can be easily controlled [14]. PCL and polyethylene glycol (PEG) are also known to be biologically safe and both have achieved FDA approval in the USA, which has facilitated their clinical application [15]. Electrospun PCL has a reasonable three-dimensional structure and pore geometry, affording constructs with good mechanical properties. In our study, PCL–LYS–H scaffolds were fabricated with heparin anchored to the scaffold surface by lysine groups via an esterification reaction. After surface modification, PCL–LYS–H scaffolds not only exhibited an anticoagulant effect but also had better biocompatibility [16]. However, because heparin inhibits the proliferation of SMCs [17], the functionalized PCL–LYS–H scaffolds were used as an inner layer of the 3DTEGs adopted in this study. Table 6 eNOS activity of ECs before and after construction of 3DTEGs Group

eNOS Activity (U/ml) (M ± SD)

P Pre–Post

EG–COG

1.124 ± 0.238

–

–

EG

1.641 ± 0.271

0.037

0.358

COG

1.453 ± 0.340

0.169

EG

2.081 ± 0.346

0.001

COG

1.648 ± 0.201

0.023

Pre Day 3

Day 7 0.042

Fig. 11 DNA index (DI) histograms generated by flow cytometry. a DI before construction of the 3DTEGs. b DI for the experimental group on day 3 after construction of the 3DTEGs. c DI for the experimental group on day 7 after construction. d DI for the control group on day 3 after construction. e DI for the control group on day 7 after construction

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PEG is a nonionic, water-soluble polymer with abundant ethoxyl groups, which can form hydrogen bonds with water to play a role in hydrophilia. The hydrophilic surface results in improved cellular affinity and growth on the scaffold. Previously, Bhattarai et al. [18] fabricated four different types of polylactide (PLLA)/PEG electrospun matrices. The hydrophilicity of the PLLA fibers and the biological activity of the fibroblast cells were notably improved by adding a small fraction of PEG into the electrospinning solution. In this present study, the coelectrospinning of PCL with 4 % PEG resulted in the construct labeled ‘4 %’, which had greater porosity than the porcine carotid artery control, PCL alone and PCL coelectrospun with other concentrations of PEG. Meanwhile, the tensile properties of the 4 % construct (i.e., elongation at break, breaking tenacity and maximum force) did not significantly change. From the mechanical properties of the 3DTEG labeled ‘10-4 %’, it was clear that this construct exhibited the thinnest wall and had the greatest burst pressure among the 3DTEG models. It also had the greatest suture strength, as compared to the porcine carotid artery control (P \ 0.05). In accordance with these findings, we chose the ‘10-4 %’ 3DTEG as the best model for tissueengineered materials for further study. Mononuclear cells can be easily isolated from bone marrow and do not require culture and proliferation in vitro. BM MNCs can provide pluripotent primitive cells, including bone marrow mesenchymal stem cells (BMSCs) [19, 20], which exert an autocrine effect. BM MNCs have no associated immunological rejection problems and provide a solution for overcoming the shortage of ethically sourced embryonic stem cells [21]. In the present study, most cells proliferated adhesively to form additional cell colonies with a cobblestone morphology after induction and culture of the BM MNCs with EBM-2 after 9–11 days. Immunofluorescence staining showed that the cells were positive for vWF and CD31, which are the phenotypic characteristics of ECs that can be used to identify ECs [22– 24]. The ECs were also positive for the uptake of DiI-AcLDL, which suggested normal function of the ECs. After culture of the BM MNCs in DMEM, the sparse and dense cell populations interlaced to form a ‘‘peaks and valleys’’ type morphology, which was observed by microscopy. The cells were positive for a-actin, calponin and SMMHC, which are, respectively, the early, mid and late markers of SMC differentiation [25] that can be used to identify SMCs [23]. In this study we prepared a triple-layered 3DTEG. This layered morphology can help prevent the overlap of different cell types, which can help prevent a cell stress response. To this end, ECs and SMCs were seeded on the PCL–LYS–H scaffold or tubular PCL sheet, respectively. Importantly, the PCL–LYS–H scaffolds, which were used

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as the inner layer, can prevent blood infiltration between the cell sheet layers. A clear outline of the 3DTEG was observed by laser confocal microscopy and labeled ECs and SMCs were clearly visible in the confocal images. Furthermore, NO and eNOS activity, and the DI of the ECs were measured before and after construction of the 3DTEG. Although there was no significant difference in the level of NO prior to construction and 3 days after construction in the experimental group, the eNOS activity on day three post-construction was significantly higher than before construction. The level of NO on day seven after construction was significantly higher than before construction and 3 days after construction (P = 0.001 and P = 0.018, respectively). This finding indicated that the activity of the ECs increased gradually after construction of the 3DTEG. It is likely that many factors contributed to the increase in activity of the ECs. Firstly, shear stress is an important factor affecting the activity of ECs [23]. Previous reports have shown that the change in shear stress was marked in dynamic cell culture, varying from 10 to 30 dynes/cm2 [26, 27]. Quinta [28] reported that functional ECs expressed more eNOS protein after exposure to laminar shear stress of 15 dyne/cm2 for 24 h. In this study, the flow rate of the DMEM medium was maintained at 7 ml/min for a week in the 3DTEG (inner diameter: 4 mm). Using Poiseuille’s equation (s = 4 lQ/pr3) [29–31], the shear stress was found to be 15 dynes/cm2, which is helpful to facilitate ECs growth on the 3DTEG. Appropriate shear stress can up-regulate the expression of VEGF, PDGF, eNOS and PGI2 to regulate cell proliferation, and the constriction and dilation of blood vessels [32]. eNOS is the only rate-limiting enzyme catalyzing the production of NO from Larginine and NO production is directly dependent on the expression of eNOS. Secondly, an appropriate culture medium provides sufficient nutrient substances for cell proliferation. DMEM (including PDGF-BB, TGF-b1 and IGF-1) supplemented with 10 % FBS was used for the coculture of ECs and SMCs. FBS has abundant mitogenic factors that are of great benefit to cell proliferation. Niklason [27] co-cultured ECs and SMCs in DMEM, and the cells grew well for 8 weeks. Thirdly, it was very important that ECs and SMCs were co-cultured to determine the cell proliferation, phenotype and protein synthesis [33, 34]. In our study, the NO production and eNOS activity of the experimental group were significantly higher than that of the control group on day seven after construction of the 3DTEGs (P \ 0.05). Di Luozzo et al. [35] previously reported that SMCs increased the production of NO by ECsand inhibited endothelin-1 (ET-1) gene transcription, which is mediated by the up-regulation of eNOS activity in ECs. Some reports have shown that 6-keto-prostaglandin F1a (6-keto-PGF1a, the stable metabolite of PGI2) and

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endothelin concentration were elevated in ECs and SMCs during co-culture, more so than when ECs are cultured alone [36, 37]. Meanwhile, increased quantities of adhesion molecules (vWF and integrin-b) and extracellular matrix proteins (Fn and laminin) were expressed by co-cultured ECs, which may be beneficial for the function of co-cultured ECs [38]. Overall, BM MNCs-derived ECs grew and performed well in the 3DTEG and the findings presented here indicated that the ‘10-4 %’ 3DTEG, which displayed active porosity, could provide a suitable perfusable vasculature with sufficient nutrients for living cells, and may advance the field of vascular tissue engineering [6].

5 Conclusion In this study we developed a new small-diameter 3DTEG, which was constructed with BM MNCs-derived ECs and SMCs that were seeded on a PCL–LYS–H scaffold and PCL co-electrospun with PEG, respectively. This provides a new thread by which to produce small-diameter tissue engineered grafts with excellent mechanical properties, perfusable vasculature and a suitable surface for the attachment and proliferation of functional cells. Acknowledgments This work was supported by the National High Technology Research and Development Program of China (2011AA020507). We declare no bias toward any institute who supported this study. Conflict of interest The authors declare that they have no potential conflicts of interest.

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