Macromolecular Research, Vol. 24, No. 4, pp 371-379 (2016) DOI 10.1007/s13233-016-4042-4
www.springer.com/13233 pISSN 1598-5032 eISSN 2092-7673
Characterization and Preparation of Bioinspired Resorbable Conduits for Vascular Reconstruction Soo-kyeong Yang1, Muhammad Shafiq2,3, Daeheum Kim1, Chulhwan Park1, Youngmee Jung2,3, and Soo Hyun Kim*,2,3,4 1
Graduate School, Department of Chemical Engineering, Kwangwoon University, Seoul 01899, Korea Korea University of Science and Technology, 176 Gajeong-dong, Yuseong-gu, Daejeon 34113, Korea 3 Center for Biomaterials, Biomedical Research Institute, Korea Institute of Science and Technology, Cheongryang, Seoul 02792, Korea 4 KU-KIST Graduate School of Converging Science and Technology, Korea University, Seoul 02857, Korea 2
Received November 5, 2015; Revised January 20, 2016; Accepted January 23, 2016 Abstract: Soft tissues such as blood vessels possess mechanical behavior characterized by a ‘J-shaped’ stress-strain curve with a low-stiffness and a highly elastic zone. These biomechanical characteristics result in rapid endothelialization, smooth muscle cell regeneration, and aneurysm inhibition. The objective of this study was to fabricate biodegradable vascular grafts mimicking the mechanical properties of native arteries. Vascular grafts (inner diameter = 5.0 mm, length = 2.0 cm) were fabricated by dip coating poly(L-lactide-co-ε-caprolactone) (PLCL) copolymers on polydioxanone (PDO) fibers. We used PDO fibers of different diameters to yield vascular grafts with a range of mechanical properties. Biomechanical properties, microstructure, and biocompatibility of the grafts were assessed using circumferential tensile testing, burst pressure measurement, scanning electron microscopy, and subcutaneous implantation, respectively. Vascular grafts possessed circumferential tensile strength and strain in the range of 4.07 to 5.98 MPa and 2.83 to 3.47 MPa, respectively, and were circumferentially stronger than expanded polytetrafluoroethylene (ePTFE) grafts. Burst pressure was physiologically relevant in the range of 1323.0 to 1736 kPa and water entry pressure was between 102.66 to 257.33 kPa. Mechanical properties of the grafts were also assessed in vivo after subcutaneous implantation in Sprague-Dawley rats for up to 8 weeks. Examination of the retrieved grafts indicated that the ‘J-shaped’ strain/stress curve was maintained for up to 3 week in PDO/PLCL vascular grafts. In contrast, ePTFE grafts did not maintain ‘J-shaped’ stress-strain behavior after in vivo implantation. Histological analysis demonstrated cellularization within PDO/PLCL grafts, whereas, PTFE grafts showed cellularization and neotissues mainly at the outer side. Our results suggest a new methodology for the fabrication of biodegradable vascular grafts with mechanical behaviour comparable to the native arteries that might avoid failure due to mechanical mismatch between the graft and native arteries. Keywords: vascular graft, poly(L-lactide-co-ε-caprolactone), polydioxanone, blood vessel, compliance, J-shaped stress/ strain curve.
Introduction Cardiovascular diseases are leading cause of morbidity and mortality word wide.1 The biologic grafts, like autologus tissues, allografts, and xenografts are gold standard treatments for the replacement of diseased vessels and arteries because of the decreased rates of thromboembolic events. However, shortcomings such as, donor-site associated toxicity risks, calcific degradation, and secondary graft failure hamper their full utilization.2-3 On the other hand, artificial vas*Corresponding Author. E-mails:
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cular prosthesis such as, polyethylene terephthalate (PET, Dacron) or expanded polytetrafluoroethylene (ePTFE, GoreTex), which have been widely used clinically for bypass grafting or for the replacement of occluded or aneurismal arteries tissues are so stiff that fabricated tubes show little inflation over full pressure regions. In addition, long-term complications including, stenosis, early stage thromboembolization, late stage neointimal hyperplasia, calcification, and infection render these materials marginally suitable for grafting especially for small-diameter vascular grafts.4,5 Thrombus formation and neointimal hyperplasia on the surface of the small-diameter artificial grafts become more critical factors for graft failure compared to medium to large diameter grafts, because the
© The Polymer Society of Korea and Springer 2016
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effective flow area of the small-diameter grafts must be smaller when the same amount of thrombosis or neointimal hyperplasia is generated.6,7 In vascular tissue engineering, a recurring problem has been disparity in mechanical properties between a native vessel and an implanted vascular graft. These mechanical property differences (i.e. compliance mismatch) are thought to be responsible, in part, for graft failure due to intimal hyperplasia. In addition, the difference in mechanical properties between a native artery and an artificial graft induces hemodynamic flow disturbance and stress concentration near the anastomosis, causing thrombosis and intimal hyperplasia.8-10 Therefore, it is crucial to develop vascular grafts (for the vascular milieu and other organ systems as well) that closely mimic the mechanical properties of the tissues that they are meant to replace or augment. The biomechanical properties of native blood vessels have been extensively characterized and it has been realized that their response to stress and strain is an important feature.11 Native tissue tubular structures, including blood vessels, exhibit characteristic nonlinear ‘J-shaped’ mechanical behavior indicated on stress-strain traces through an initial low-stiffness region at low strain attributable to elastin content that gives rise to generally an order-of-magnitude increase in stiffness through a curvilinear transition at high strain attributable to collagen content.12-16 This ‘J-shaped’ stress-strain curve combines soft, compliant mechanics and large levels of stretchability, with a natural ‘strain-limiting’ that protects biological tissues from excessive strain.17,18 It is the elastic nature of elastin that includes recoil after each cycle of vessel inflation and prevents arterial dilatation in vivo.13-15 Furthermore, strong elastin and collagen networks in the wall of native blood vessels provide good surgical properties and prevent rupture during in vivo experiments. Researchers have identified the complex nature of the stress-strain behavior of vessels and are actively seeking to recapitulate those properties through a variety of natural and synthetic materials, construction modalities (e.g. electrospinning and casting), and various composite designs via different approaches.19-25 In addition, the mechanical properties of various fabricated scaffolds made from blends of elastin, collagen, and synthetic polymers have been reported.24,25 However, there remains a need for tissue-engineered scaffolds that are capable of recapitulating the ‘J-shaped’ stress-strain behavior, and methods for making such scaffolds. The objective of this study was therefore to develop bioinspired resorbable conduits mimicking the mechanical response of native arteries for vascular reconstruction. To this end, we used elastomeric resorbable polymeric materials because they can also transduce mechanical stimulation to cells.26,27 We used mechano-elastic poly(L-lactide-co-ε-caprolactone) (PLCL, molar ratio 50:50) copolymer and polydioxanone (PDO) to recreate complex structure on a macroscale in a tubular scaffold. Copolymers of L-lactide (LA) and ε-caprolactone (CL) (PLCL) exhibit various mechanical properties, depending upon their 372
composition. For example, the high-molecular-weight equimolar copolymer PLCL (50:50) is a biodegradable elastomer. Owning to its tunable biodegradability and mechano-active nature, it has been widely investigated as a candidate biomaterial for cardiovascular regeneration.28-30 Ideally, a PLCL mesh or vascular graft will degrade and sorbed with the passage of time in vivo. Concomitantly, cells from the surrounding tissues will invade, proliferate, and produce ECM to form neo-vessels.30 Previously, we and others have reported the satisfactory performance of small-diameter tissue-engineered vascular grafts fabricated from PLCL copolymers in vitro and in vivo.28 In those reports, vascular grafts were developed using various processing techniques, such as electrospinning, salt-leaching, and scaffold-membrane approach. PLCL vascular graft induced autologus tissue regeneration by the proliferation and differentiation of endothelial and smooth muscle cells (SMCs) after in vivo implantation avoiding risks such as, intimal hyperplasia, occlusion, and aneurysm.31,32 PDO was chosen to be the second component of our tissue-engineered vascular grafts. PDO is a colorless, crystalline, bioabsorbable polymer that is currently in use as a commercially available wound closure suture. These polymers degrade by hydrolytic process, generally resulting in low molecular weight species, which can be metabolized or bioabsorbed by the body. As a suture material, PDO is highly flexible with excellent strength retention and shape memory and offers several advantages over the more traditional bioresorbable polymers like poly(glycolic acid) (PGA), poly(lactic acid) (PLA) and poly(lactic-co-glycolic acid) (PLGA), because it has a slower resorption rate and induces a lower inflammatory response.33 PDO has received approval of the US Food and Drug Administration (FDA) to be used as a suture material in gynecology. Vascular grafts consisted of electrospun PDO fibers showed less thrombosis compared to PGA and Dacron synthetic grafts. However, PDO lacks the potential for cell infiltration in vivo and shows slight tendency for the formation of aneurysm.34,35 We fabricated tubular grafts by gel-spinning PLCL solution onto non-woven PDO fibers of different microstructure and assessed the morphology and biomechanical behavior of these grafts in vitro and in vivo.
Experimental Materials. Polydioxanone (PDO) absorbable monofilaments were purchased from Meta BioMed (MEPFIL-D Series, Seoul, South Korea). The retention of PDO is more than 65% of the initial strength after 28 days at 37 oC. We used 4-0, 3-0, and 2-0 PDO fibers whose characteristics are given in Table 1. PDO fibers make the structures to tubes (ID=5.0 mm) (Fig. 1(A)-(B)). Synthesis of Poly(L-lactide-co-ε-caprolactone). Poly(Llactide-co-ε-caprolactone) (50:50) copolymer was synthesized using our previous published method described in detail elsewhere.36,37 Briefly, polymerization was carried out in a 100 mL Macromol. Res., Vol. 24, No. 4, 2016
Bioinspired Resorbable Conduits for Vascular Reconstruction
Table I. Dimensions of PDO Fibers Used for the Fabrication of PDO/PLCL Vascular Grafts and Identification Codes of the Prepared Grafts Size USP
Diameters by EP or USP Methods (mm) EP
Average Value
Individual Value
Gauge
Minimum
Maximum
Minimum
Maximum
4-0
1.5
0.200
0.249
1.75
0.275
3-0
2
0.250
0.339
0.225
0.375
2-0
3
0.340
0.399
0.325
0.450
Identification Code
Diameter of PDO Fibers (mm)
Graft Wall Thickness (mm)
Inner Diameter of the Graft (mm)
4-0
0.369
2.0
5.0
3-0
0.295
1.8
5.0
2-0
0.225
1.5
5.0
glass ampoule containing L-lactide (100 mmol), ε-caprolactone (100 mml), 1,6-hexanediol (0.5 mmol), and stannous octoate (1 mmol). The ampoule was sealed under vacuum after purging three times with nitrogen and reaction was carried out in an oil bath at 150 oC for 24 h. The obtained polymer was dissolved in chloroform and filtered through a 4.5 μm pore membrane filter. The polymer was precipitated into an excess of methanol, filtered and dried under vacuum. Preparation of Vascular Grafts. PDO/PLCL vascular grafts were fabricated using extrusion-particulate leaching technique. PLCL was dissolved in chloroform (1% w/v) and sodium chloride particles (50 μm) (1:2) were added. PLCL/salt solution was extruded on PDO fibers by a custom-designed piston extrusion set up. The solvent was evaporated for 48 h at room temperature and grafts were dried in vacuum oven for 24 h. The resulting PLCL/salt composite tubular scaffolds were leached in deionized water with shaking for 3 days, freeze-dried for 24 h, and sterilized with ethylene oxide gas. Evaluation of Mechanical Properties. Tensile properties of vascular grafts (5 mm × 10 mm) were examined using a tensile testing equipment (Model 5567, Iinstron, TestResoruces, Shakopee, MN, USA) with a 10 N load cell at a cross-head speed of 10 mm/min (n = 8). Stress-strain curves were plotted using load versus elongation data obtained from the experiments. Burst Pressure Measurement. Burst pressure of the vascular grafts was measured using a custom-designed set up. Samples were mounted on instrument and pressurized with water. The pressure was increased by 5 kPa to a maximum of 30 kPa for 5 minutes and maintained at 30 kPa for 30 min. Afterwards, the pressure was gradually increased until the graft failure occurred.38 Water Entry Pressure. Water entry pressure was measured using a custom-designed leak-free set up. A section of graft (length=20 mm) was attached to the system via leak proof connectors. Deionized water was used to pressurize the graft to a value less than the expected water entry pressure, approximately 2.5 pounds per square inch (psi). The pressure was Macromol. Res., Vol. 24, No. 4, 2016
then increased by 0.5 psi increments and held at each pressure for approximately 15 seconds. At the first sign of water on the outer surface, the water entry pressure was recorded on the data sheet. Pore Size Measurement. The morphology of scaffolds was studied with a scanning electron microscope (SEM, Hitachi, Tokyo, Japan) using a voltage of 15 kV. The specimens were coated with gold using a sputter-coater (Eiko IB3, Tokyo, Japan). The pore diameter on the inner and outer surfaces of the graft was measured from the obtained micrographs. Subcutaneous Implantation of Vascular Grafts. In order to access the mechanical properties and evaluate the biocompatibility of the developed grafts in vivo, vascular grafts were subcutaneously implanted in Sprague-Dawley rats (n = 32, age = 10 weeks, weight = ~250 to 300 g, Japan SLC Inc, Japan). Animal were divided into four groups and four grafts were implanted per group for each time point. Under general anesthesia by intravenous injection of urethane (0.5 g/kg), the dorsal skin was shaved and sterilized using 70% ethanol. PDO/PLCL grafts (ID = 5.0 mm, length = 2.0 cm) were implanted into the subcutaneous dorsum. In addition, PTFE grafts (ID = 4.5 mm, length = 2.0 cm) were implanted as a positive control. The skin was closed with 3-0 monofilament nylon sutures. After 1, 3, 5 or 8 weeks of implantation, animals were sacrificed and grafts were retrieved. The grafts were rinsed with saline and then cut into two parts from the middle. One part was fixed with 10% neutral-buffered formalin for 24 h and embedded in paraffin for sectioning. Parrafin-embeded samples were sectioned into slices (thickness = 6.0 μm) using a microtome (Leica, Germany). The other part of the graft was immersed in PBS (1%) and characterized using tensile testing.
Results PLCL (50:50) was prepared using ring opening polymerization (Figure 1(D)). Number average molecular weight (Mn) of the copolymer was found to be 167 kDa as evaluated using gel permeation chromatography.37 373
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Figure 1. (A-H) Schematic illustration of the experimental setup for the fabrication of vascular grafts. (A) Photomicrograph of PDO fibers. (B) Scanning electron micrograph of PDO fibers. Scale bar =1 mm. Fibers with three different diameters (identification codes =4-0, 3-0, and 2-0) were used. (C) Photomicrograph of commercial ePTFE vascular graft. (D) Synthesis of PLCL (50:50) by ringopening polymerization. L-lactide (LA) and ε-caprolactone (CL) were reacted at 150 oC for 24 h. (E) Fabrication of PDO/PLCL vascular grafts. PDO fibers were coated with PLCL/salt solution. The resulting PDO/PLCL scaffolds were subjected to salt leaching to remove salt, freeze-dried, and stored at -20 oC for subsequent use. (F) photographs of 4-0, 3-0, and 2-0 vascular grafts. (G) PDO/ PLCL or PTFE grafts were cut into desired dimensions, sterilized using ethylene oxide gas, and (H) implanted subcutaneously in SD rats for 1W, 3W, 5W, and 8W.
Pore Size Measurement. Morphological analysis of the grafts was performed using SEM and the pore sizes on the inner and outer surfaces of the grafts were determined from the micrographs. SEM micrographs of grafts are shown in Figure 2 whereas the pore sizes on the inner and outer surfaces of the grafts are tabulated in Table II. The average pore size on the inner and outer surfaces of PTFE grafts was found to be 16.83±2.54 μm and 15.50±3.04 μm, respectively (Figure 2(D), (H); Table II). On the other hand, the average pore sizes on the inner surface of the PDO/PLCL grafts were found to be 44.72±5.33 μm, 47.95±2.26 μm, and 46.83±2.54 μm for 374
Figure 2. SEM micrographs of the inner and outer surface of PTFE and PDO/PLCL vascular grafts. (A) 4-0 inner, (B) 3-0 inner, (C) 2-0 inner, (D) PTFE inner, (E) 4-0 outer, (F) 3-0 outer, (G) 2-0 outer, and (H) PTFE outer. Grafts were sputter coated with gold prior to SEM microscopy. Scale bars, 100 µm.
4-0, 3-0, and 2-0 grafts, respectively (Figure 2(A)-(C); Table II). The average pore sizes on the outer surface of 4-0, 3-0, and 2-0 PDO/PLCL grafts were about 42.25±2.51 μm, 47.19±3.93 μm and 45.50±3.04 μm, respectively (Figure 2(E)-(G); Table II). Mechanical Properties. Tensile properties, such as tensile stress, Young’s modulus, and strain at failure were determined from the stress-strain curves. Figure 3 shows stress-strain traces of PTFE and PDO/PLCL grafts. PDO/PLCL grafts demonstrated ‘J-shaped’ curves on stress-strain traces similar to native Macromol. Res., Vol. 24, No. 4, 2016
Bioinspired Resorbable Conduits for Vascular Reconstruction
Table II. Pore Size on the Inner and Outer Surface of PTFE and PDO/PLCL Vascular Grafts Scaffold Type
Inner Surface (µm)
Outer Surface (µm)
4-0
44.72 ± 5.33
42.25 ± 2.51
3-0
47.95 ± 2.26
47.19 ± 3.93
2-0
46.83 ± 2.54
45.50 ± 3.04
PTFE
16.83 ± 2.54
15.50 ± 3.04
Figure 3. Stress-strain traces of PTFE and PDO/PLCL vascular grafts. PDO/PLCL vascular grafts showed ‘J-shaped’ curves on stress-strain traces and exhibited high tensile strength compared with PTFE grafts. Of PDO/PLCL grafts, grafts fabricated from 2-0 and 3-0 PDO fibers showed high tensile strength.
arteries.13 By contrast, PTFE grafts did not exhibit ‘J-shaped’ stress-strain curves. It can also be seen that the PDO/PLCL grafts exhibit better elastic region than PTFE grafts under the strain value of 100%. The tensile strain of PDO/PLCL grafts
was comparable to PTFE grafts. Of all grafts, the PTFE graft showed the highest strain value, indicating that it was soft in comparison. Since PLCL is elastin like rubbery material, on stress-strain curve, low strain behavior is consistent with PLCL, whereas, the high strain behavior is dominated by the PDO fibers. PDO fibers enhance the tensile strength of the graft. The vascular graft fabricated from 2-0 PDO fibers exhibited the highest value of tensile strength because 2-0 PDO fibers have the largest diameter among all fibers used in this study. According to the manufacturer’s data sheet, PDO fibers with larger diameter have higher strength. Vascular grafts fabricated from 3-0 and 4-0 PDO fibers documented similar value of tensile stress. Elastic modulus values were closely related to the compliance because the inverse of modulus is compliance in elastic materials like rubbers.39 Burst Pressure and Water Entry Pressure. Arterial substitutes need to withstand arterial pressure immediately upon implantation. Therefore, burst pressure was measured to identify the maximum pressure that grafts may endure before failure and to access whether they could exhibit adequate strength to endure the physiologic hemodynamic force. The burst pressure values of 4-0, 3-0, 2-0, and PTFE grafts were 1323.0±303.0 kPa, 1755.3±516.0 kPa, 1736.0±426.0 kPa, and 656.6±85.0 kPa, respectively (Figure 4(A)). The water entry pressure (kPa) values of 4-0, 3-0, 2-0 and PTFE grafts were 102.66±3.0, 211.66±58, 257.33±24, and 81.66±17 kPa, respectively (Figure 4(B)). As can be noted from these values and observed from the figure, PDO/PLCL possessed significantly high burst strength compared to PTFE grafts (*p
