Electrospun Tecophilic/gelatin nanofibers with potential for small diameter blood vessel tissue engineering
Elham Vatankhah1,2, Molamma P Prabhakaran2,*, Dariush Semnani1, Shahnaz Razavi3, Mohammad Morshed1, Seeram Ramakrishna2,4
1
Department of Textile Engineering, Isfahan University of Technology, Isfahan 84156-83111,
Iran 2,*
Center for Nanofibers and Nanotechnology, E3-05-14, Nanoscience and Nanotechnology
Initiative, Faculty of Engineering, National University of Singapore, 2 Engineering Drive 3, Singapore 117576. Email:[email protected] 3
Department of Anatomical Sciences and Molecular Biology, School of Medicine, Isfahan
University of Medical Sciences, Isfahan 81744-176, Iran. 4
Department of Mechanical Engineering, Faculty of Engineering, National University of
Singapore, 2 Engineering Drive 3, Singapore 117576.
This article has been accepted for publication and undergone full peer review but has not been through the copyediting, typesetting, pagination and proofreading process which may lead to differences between this version and the Version of Record. Please cite this article as an ‘Accepted Article’, doi: 10.1002/bip.22524 © 2014 Wiley Periodicals, Inc.
Biopolymers
Abstract Tissue engineering techniques particularly using electrospun scaffolds have been intensively utilized in recent years for the development of small diameter vascular grafts. However, the development of a completely successful scaffold that fulfills multiple requirements to guarantee complete vascular regeneration remains challenging. In this study, a hydrophilic and compliant polyurethane namely Tecophilic (TP) blended with gelatin (gel) at a weight ratio of 70:30 (TP(70)/gel(30)) was electrospun to fabricate a tubular composite scaffold with biomechanical properties closely simulating those of native blood vessels. Hydrophilic properties of the composite scaffold induced non-thrombogenicity while the incorporation of gelatin molecules within the scaffold greatly improved the capacity of the scaffold to serve as an adhesive substrate for vascular smooth muscle cells (SMCs), in comparison to pure TP. Preservation of the contractile phenotype of SMCs seeded on electrospun TP(70)/gel(30) was yet another promising feature of this scaffold. The nanostructured TP(70)/gel(30) demonstrated potential feasibility towards functioning as a vascular graft. Keywords: blood vessel, nanofibrous tubular scaffold, biomechanical properties, blood compatibility, smooth muscle cells
1. Introduction Coronary artery disease is the leading cause of mortality worldwide. Limited availability of healthy autologous vessels along with the failure of synthetic grafts such as the woven Dacron and expanded polytetraflouroethylene (ePTFE), as small diameter vessels (< 6 mm) have forced bioengineers to develop alternative grafts for the regeneration of vascular tissues. Tissue engineering is a promising approach for creating small diameter vessels by combining autologous vascular cells with a natural and/or synthetic scaffold 1.
2 John Wiley & Sons, Inc.
Page 2 of 43
Page 3 of 43
Biopolymers
The rational design of a scaffold depends on its ability to mimic the native extracellular matrix (ECM) as much as possible in terms of physical, chemical and mechanical features. Development of an optimal tubular scaffold plays a significant role in vascular tissue engineering. Bioengineered scaffolds serve as the temporary matrix and provide structural, mechanical and biological support throughout the tissue formation, assisting the process of tissue remodeling2. Among the numerous techniques used to fabricate scaffolds, electrospinning has been widely used because of the high porosity, large surface area and nanofibrous structure of electrospun scaffolds, mimicking the physical nano features of native ECM 3. Furthermore, the simplicity of electrospinning technique to fabricate a tubular scaffold makes it more popular for vascular tissue regeneration applications. In vascular tissue engineering of small diameter arteries, a recurring problem is their failure owing to higher possibility of thrumbus formation which is due to experiencing lower flow rate and having higher ratio of surface area to blood volume compared to larger diameter vascular grafts which increase the time interaction with blood
4, 5
. Other common causes of long-term
failure in small grafts are compromised visco-elastic mechanical behavior of blood vessels after long-term implantation of vascular scaffolds and intimal hyperplasia at the anastomotic site, with proliferation and migration of vascular smooth muscle cells from the media to intima and subsequent synthesis of matrix proteins. Compliance mismatch between a native vessel and the implanted vascular graft leads to excessive strain on anastomosis, resulting in suture failure and consequent development of false aneurysm 6. Thus, a vascular scaffold should be made of a durable biomaterial allowing cell infiltration, adhesion and proliferation while having strength, compliance and suture retention strength comparable with those of native blood vessels
1, 2
. Blood compatibility is also clearly a crucial
3 John Wiley & Sons, Inc.
Biopolymers
Page 4 of 43
factor to raise success of the engineered constructs since the vascular scaffolds come in contact with blood. Recently several attempts have been made towards the engineering of vascular constructs that closely mimic the strength and compliance to native blood vessels
7, 8
. In addition various
approaches were carried out to design non-thrombogenic interface 9. In this study, we sought to develop a durable electrospun scaffold to mimic the fibrous structural and J-shaped mechanical properties of native arteries while blood compatibility is prioritized and considered during this study. Due to the importance of the compositional ratio of the polymer constituting an electrospun scaffold in determining mechanical properties of the scaffold 10, it is anticipated that electrospinning of a blend of compliant and stiff polymers might be a promising approach towards this goal. Polyurethane is often used in blood-contacting devices due to its thrombo-resistance, elasticity, good compatibility with tissue and blood, and resistance to mechanical degradation
11
.
Tecophilic® (TP) is a family of hydrophilic polyether-based thermoplastic aliphatic polyurethanes. It is a medical grade polymer capable of absorbing water content of up to 150% of the weight of its dry resin 12. Gelatin (gel), is a stiff and brittle biopolymer collagen, possesses many integrin binding sites
14
13
extracted from
, which makes it an appropriate candidate to
blend with TP for fabricating a composite scaffold for vascular tissue regeneration. Detta et al. used electrospun scaffolds of Tecoflex and gelatin from separate spinnerets for application in vascular tissue engineering, but reports on the blood compatibility of the scaffolds were not available
15
. In another study, Wang et al. fabricated a bilayered tubular electrospun
scaffold of PU (outer layer) and gelatin-heparin (inner layer) as an artificial blood vessel, whereby heparin provided blood compatibility of the scaffolds16. However, the lack of cell
4 John Wiley & Sons, Inc.
Page 5 of 43
Biopolymers
binding sites on the outer layer of the scaffold limits its application as a substrate for vascular tissue engineering. Therefore we aimed to investigate all aspects required for a successful tissue engineered vascular graft, during this study. It is noteworthy that a blend ratio of 70:30 for TP:gel was found in our previous study as an optimum composition to develop a suitable electrospun substrate that enhanced the contractility of smooth muscle cells (SMCs)
17
. Therefore, this composition was chosen in this study to
further examine the biomechanical and blood compatibility properties in detail.
2. Materials and Methods 2.1. Materials Tecophilic® (TP) with a product name of (SP-80A-150) was a kind gift from Lubrizol. Gelatin (gel) type A (300 Bloom) from porcine skin, bovine serum albomin (BSA), anti CD41 and 1,1,1,3,3,3-hexafluoro-2-propanol (HFP) were purchased from Sigma-Aldrich (Singapore). Human aortic smooth muscle cells (HASMC) and smooth muscle cell medium (SMCM) were obtained from ScienCell Research Laboratories (Singapore). 4′, 6-diamidino-2-phenylindole (DAPI) was purchased from Invitrogen. Anti-alpha smooth muscle actin antibody and antismooth muscle myosin heavy chain were obtained from Abcam (HongKong). 2.2. Fabrication of scaffolds Polymeric solutions were prepared by dissolving TP and gel with weight ratios of 100:0, and 70:30 in HFP to obtain a total concentration of 8% (w/v), followed by 24 h stirring at room temperature. Each solution was loaded to a 5 ml syringe capped with a 27 G blunted stainless steel needle, connected to the positive terminal of a high voltage power supply (10 kV) and a grounded flat collector wrapped in aluminum foil was placed at a distance of 12 cm from the needle tip to collect fibers. The flow rate was set on 1 ml/h using a programmable syringe pump. 5 John Wiley & Sons, Inc.
Biopolymers
Electrospun fibers were collected on 15 mm cover slips for SEM and in vitro studies while electrospun sheets were used for other analysis. Pure gelatin solution was also electrospun to provide control samples for chemical evaluation. For fabricating tubular scaffold of TP(70)/gel(30), a cylindrical mandrel of 5 mm diameter rotating at a speed of 150 rpm was used. The collector was traversed at 20 mm/s to obtain uniform thickness of the scaffold along the length of the mandrel. Electrospun scaffolds were transferred to a vacuum desiccator for at least 48 h to ensure no residual solvent on the scaffold surfaces. 2.3. Morphological and chemical Characterization Scanning electron microscopy (SEM) (JSM5600, JEOL, Japan) was used to study the morphological features of electrospun scaffolds, operated at an accelerating voltage of 15 kV. Scaffolds were coated with platinum (JEOL JFC-1300 Auto fine coater, Japan) before SEM analysis. The fiber diameter was measured from SEM micrographs using image analysis software (Image J, National Institutes of Health, USA). Results were presented in the form of mean ± standard deviation of 100 random points per image. Structural pore properties were investigated by a CFP-1200-A capillary flow porometer (PMI, New York, NY). Galwick with a defined surface tension of 15.9 dynes/cm (PMI, New York, NY) was used as the wetting liquid for porometry measurements. The pore size was measured from three samples of each type of electrospun membrane with the similar thickness of 0.35 µm and a dimension of 3 × 3 cm2. The total porosity of the electrospun scaffolds was estimated by measuring the apparent density of the electrospun scaffolds using the weights of precisely cut samples of defined area and thickness. The overall porosity was calculated according to the following equation:
6 John Wiley & Sons, Inc.
Page 6 of 43
Page 7 of 43
Biopolymers
Porosity (%) = (1 −
(1)
ρ0 ) × 100 ρ
where ρ0 and ρ are apparent density of the scaffold and density of the bulk polymer, respectively. Water contact angle as a wettability index of the electrospun scaffolds was determined using sessile drop method-based video contact angle system mounted with a CCD camera (VCA Optima, AST Products, Billerica, MA). Distilled water droplet size was set at 1.0 µl and each measurement was recorded after 20 s. For chemical analysis of the electrospun scaffolds, attenuated total reflectance Fourier transform infrared (ATR-FTIR) spectra were measured using an Avatar 380 (Thermo Nicolet, Waltham, MA) spectrometer over the range 400- 4000 cm-1 at a resolution of 4 cm-1. 2.4. Biodegradation assessment Cover slips containing electrospun nanofibrous scaffolds were placed in 24-well plate containing 1 ml of phosphate buffer solution (PBS; pH 7.4) in each well and were incubated in vitro at 37° C. PBS was refreshed after every alternate days. After a period of 7, 14, and 28 days, samples were washed and subsequently dried in a vacuum desiccator at room temperature for 48 h. SEM of the scaffolds were performed to understand the change in nanofibers morphology during these periods. For assessment of swelling ratio (water-uptake capacity) and mass loss of the scaffolds, three square-shaped sheets of each composition were immersed in PBS at 37°C for 7, 14, and 28 days. Samples were weighed before immersion (W0), after tap drying with a filter paper (W) and completely drying in a vacuum desiccator (Wd). The swelling ratio and mass loss were calculated from following formula:
7 John Wiley & Sons, Inc.
Biopolymers
Swelling ratio (%) =
Mass loss (%) =
Page 8 of 43
(2)
W − Wd × 100 Wd
(3)
W0 − Wd ×100 W0
2.5. Biomechanical properties of the conduit Mechanical properties of the tubular scaffold were determined by uniaxial tensile testing of scaffold specimens equilibrated with PBS overnight at 37°C using a uniaxial testing system (Instron 5943) with a 50 N load cell and an extension rate of 10 mm/min. At least five specimens were tested for each mechanical evaluation. Tensile tests of ring samples with 3 mm width (W) and certain diameter (2R0) were conducted after measuring the thickness (t) on the mentioned tensile apparatus with two stainless steel hooks subsequently fixed in its clamps for gripping the ring. All ring specimens were preconditioned prior to testing with 10 cycles at 13% circumferential stretch where stretch, was defined as the sample length (L) normalized to initial length (L0). Stress-strain curves in terms of the circumferential stress against circumferential strain were computed according to current length and cross section area, shown schematically in Figure 1, with the incompressibility assumption for the tubular scaffold (equations 4-7);
A0 = 2 . W . t
(4)
F A0
(5)
L0 = π . R0
(6)
σ=
8 John Wiley & Sons, Inc.
Page 9 of 43
Biopolymers
(7)
L R A0 = = L0 R0 A
where σ is corresponding stress of applied force (F), and A is cross sectional area. Maximum stress value before failure and its corresponding value of stretch were taken from the stress-strain curve, respectively as the ultimate tensile strength (UTS) and stretch to failure (STF). Linear regression of the slope in stress-strain plots was used to calculate the elastic modulus. Elastic modulus (EM) was measured between 10 and 30% strain, termed as EM1, while EM2 was measured between the maximum strain minus 20% and the ‘maximum’ strain, with linear curve fits to these regions of the stress-strain curve 7. The results of the tensile test of ring samples were used to calculate the theoretical compliance (C) (equation 8) and burst strength (BS) of the tubular scaffold regarding thin-walled cylinder approximation relating pressure to stress scaled by wall thickness and initial radius (equation 9) and Barlow equation (equation 10) 7, 18. (Figure 1) (8)
R2 − R1 %compliance R1 = × 10 4 100 mmHg P2 − P1
σ=
Pb =
(9)
P R0 t
(10)
σ UTS t R0
where, R1 and R2 are internal radii at pressure P1 and P2, respectively. P, Pb, σ, σUTS, and R0, are transmural pressure, burst strength, tensile stress, ultimate tensile strength, and initial internal radius, respectively. Compliance was measured as the change in radius relative to the initial
9 John Wiley & Sons, Inc.
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Page 10 of 43
radius over the change in tensile stress at 200 and 300 kPa which are equivalent to physiological arterial blood pressure of 120/80 mmHg (systolic/diastolic) according to equation (9). Rectangular specimens were also prepared from longitudinally opened tubular scaffolds in longitudinal direction with dimensions of 20 × 5 mm2. Longitudinal samples were used to determine the suture retention force. One end of the specimen was fixed to the stage clamp of the uniaxial load test machine and a single commercial 5-0 vicryl loop (Ethicon Inc) was placed at 3 mm from the short edge of the opposite end of each sample and secured to a hook connected to the upper clamp of the testing device. The specimen was pulled at a crosshead speed of 10 mm/min until either the suture ripped or the scaffold fractured. Suture retention force (SRF) was considered as the maximum force recorded prior to failure. 2.6. Hemolysis test Hemolysis test provides information on whether the interaction of erythrocytes with biomaterial leads to hemolytic process. By providing a culture medium enriched with red blood cells, it is possible to determine whether biomaterials used in this research can destroy cells and digest the hemoglobin inside. In this regard, electrospun scaffolds were cut into small bars (about 5 × 25 mm2) and put into a vial. After rinsing bars three times with normal saline (0.9% (w/v) sodium chloride solution in distilled water), they were immersed in 1 ml of normal saline for 3 h at 37°C. Human whole blood was collected in accordance with regulations of the NUS Institutional Review Board (NUS-IRB) into blood collection tubes with 3.2% buffered sodium citrate solution. 1 ml of anticoagulant blood diluted three times with normal saline was added to each vial allowing samples to remain soaked in the solution at 37°C for 60 min. Diluted anticoagulant blood was added to distilled water and normal saline, which served as positive and negative controls, respectively. The hemolytic percentage was determined by photometric analysis of
10 John Wiley & Sons, Inc.
Page 11 of 43
Biopolymers
supernatant that was pre-centrifuged at 1000 rpm and read at 540 nm using a spectrophotometric plate reader. The hemolysis ratio (HR) was calculated as follows:
HR % =
(11)
Ats − Anc ×100 Apc − Anc
where Ats is the absorbency of test sample and Apc, Anc are the absorbance values of positive and negative controls, respectively. All hemolysis tests were done in triplicates. 2.7. Platelet adhesion assay The capability of electrospun scaffolds to prevent thrombosis was investigated by evaluating the degree of platelet adhesion on them based on two techniques (i) SEM (ii) immunofluorescent staining. Briefly, anticoagulant blood was centrifuged at 2000 rpm for 10 min at room temperature to softly sediment white and red blood cells in the lower fraction, while platelet rich plasma (PRP) remained suspended. The upper layer of the PRP was carefully collected using a pipette to limit contamination by other cellular elements and placed in a sterile tube. Scaffolds placed in 24 well plates were washed three times with distilled water and equilibrated with PBS for 30 min at 37°C, then incubated with a sample of 200 µl of PRP for 1 h at 37°C. Subsequently, scaffolds were washed repeatedly with PBS to remove un-adhered platelets. Samples were fixed using glutaraldehyde (2.5 wt-%), dehydrated with graded ethanol from 50% to 100%, dried overnight and sputter coated with platinum and were observed under SEM. In addition, the adherent platelets were stained by mouse anti-human CD41 and goat anti-mouse IgG Alex Fluor-488 conjugate (Invitrogen) and viewed using a laser scanning confocal microscope (LSCM; Olympus, FluoView FV1000). 2.8. In vitro Cell Culture HASMCs were cultured in smooth muscle cell medium (SMCM) containing basal medium complemented with 2% fetal bovine serum (FBS), 1% smooth muscle cell growth supplement 11 John Wiley & Sons, Inc.
Biopolymers
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(SMCGS), and 1% penicillin/streptomycin solution (P/S) in a 150 cm2 cell culture flask. Cells were incubated under standard incubator (37° C, 5% CO2) for 6 days and the culture medium was changed every 3 days. Electrospun fibers collected on glass cover slips and placed in a 24 well plate were sterilized under ultraviolet (UV) radiation for 2 h, washed thrice with PBS and subsequently immersed in SMCM overnight before cell seeding. Confluent HASMCs were trypsinized with 0.25 % trypsin containing 0.1% EDTA. Deattached cells were centrifuged at 1500 rpm for 5 min, counted by trypan blue assay using a hemocytometer, and seeded on electrospun scaffolds and tissue culture polystyrene (TCP) as the control at a density of 10,000 cells per well. 2.9. Cytocompatibility and proliferation assay The viability of HASMCs on electrospun scaffolds was examined qualitatively and quantitatively using Live-Dead Cell Staining Kit (Invitrogen) and cell metabolic activity assay, respectively. Cell viability/cytotoxicity assay based on the simultaneous determination of live and dead cells was employed after 1, 4, 7, and 10 days of cell culture according to the manufacture’s protocol. Briefly, staining solution was prepared by mixing two probes of calcein AM and ethidium homodimer-1 (EthD-1) in PBS for detection of live and dead cells, respectively. Cell-seeded scaffolds were washed with PBS and incubated with 200 µl of staining solution for 30 min at 37° C. Further, samples were washed with PBS and viewed under Leica DM IRB fluorescent microscope. The proliferation of HASMCs was investigated by 3-(4, 5-dimethylthiazol-2-yl)-5-(3carboxymethoxyphenyl)-2-(4-sulfophenyl)-2H-tetrazolium (MTS) assay (CellTiter 96 AQueous One solution; Promega, Madison, WI). Briefly, after 1, 4, 7, and 10 days, cell-seeded scaffolds were rinsed with PBS, incubated with 1 ml of serum free medium containing 20 % of MTS
12 John Wiley & Sons, Inc.
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Biopolymers
reagent at 37° C for 3 h, aliquoted into a 96-well plate (100 µl/well) and absorbance of the obtained dye was measured at 490 nm using a spectrophotometric plate reader (FLUOstar Optima, BMG Lab Technologies, Offenburg, Germany). Samples with culture medium but without cells were set to determine the background absorbance to be subtracted. The intensity of obtained color is directly proportional to the metabolic activity of the cell population. 2.10. HASMC morphology on electrospun scaffolds After 4 and 10 days of culture, cell seeded constructs were harvested to observe morphological characteristics of HASMCs by SEM. Briefly, cell-cultured samples were washed with PBS, fixed with 3 % glutaraldehyde for 3 h, rinsed with distilled water and dehydrated through a graded series of ethanol solutions (50 to100 % (v/v)). Samples were then treated with hexamethyldisilazane and air-dried in a fume hood. Completely dried specimens were sputter coated with platinum and analyzed by SEM. 2.11. HASMC cellular phenotype analysis To identify whether the interaction of cells with nanofibers might affect their phenotype, the presence of two specific smooth muscle cell markers (smooth muscle alpha actin (α-SMA), and smooth muscle myosin heavy chain (SM-MHC) of cells seeded for 5 days on electrospun scaffolds were analyzed by immunofluorescent staining. Briefly, formalin-fixed cells were incubated with 250 µl of 0.1% Triton-X100 for 3min at room temperature to permeabilize the cell membrane. Triton-X100 was removed by washing samples thrice with PBS and non-specific binding sites were blocked by incubating cells with 150 µl of 3 % BSA for 90 min at room temperature. Samples were then incubated with 150 µl of either primary rabbit anti- human αSMA or mouse anti- human SM-MHC (1:200 dilution) for 2 h at room temperature under shaking condition. After three washes with PBS, either a goat anti-rabbit IgG Alexa Fluor-488
13 John Wiley & Sons, Inc.
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Page 14 of 43
conjugate (1:400 dilution), or a goat anti-mouse IgG Alexa Fluor-546 conjugate (1:400 dilution) (Invitrogen) was used as appropriately. Nuclei were stained with 4', 6-diamidino-2-phenylindole, Dihydrochloride (DAPI; Invitrogen) at a dilution of 1:1000. Further, cell-scaffold constructs were washed with PBS and visualized under laser scanning confocal microscopy (Olympus, FluoView FV1000) after mounting on glass sides. Cells seeded on TCP with the same treatment were set as the control. In order to investigate the cellular penetration after 5 and 10 days of cell culture, towards the interior of the composite TP(70)/gel(30) scaffolds, the immunofluorescent staining of α-SMA was performed and cell infiltration was analyzed using Imaris software. 2.12. Statistical analysis All data presented are expressed as mean±standard deviation (SD). Statistical analysis was carried out by performing one-way ANOVA followed by Tukey post hoc tests for multiplecomparison of different samples using S-PLUS (TIBCO Software Inc., USA). A value of p
Elham Vatankhah1,2, Molamma P Prabhakaran2,*, Dariush Semnani1, Shahnaz Razavi3, Mohammad Morshed1, Seeram Ramakrishna2,4
1
Department of Textile Engineering, Isfahan University of Technology, Isfahan 84156-83111,
Iran 2,*
Center for Nanofibers and Nanotechnology, E3-05-14, Nanoscience and Nanotechnology
Initiative, Faculty of Engineering, National University of Singapore, 2 Engineering Drive 3, Singapore 117576. Email:[email protected] 3
Department of Anatomical Sciences and Molecular Biology, School of Medicine, Isfahan
University of Medical Sciences, Isfahan 81744-176, Iran. 4
Department of Mechanical Engineering, Faculty of Engineering, National University of
Singapore, 2 Engineering Drive 3, Singapore 117576.
This article has been accepted for publication and undergone full peer review but has not been through the copyediting, typesetting, pagination and proofreading process which may lead to differences between this version and the Version of Record. Please cite this article as an ‘Accepted Article’, doi: 10.1002/bip.22524 © 2014 Wiley Periodicals, Inc.
Biopolymers
Abstract Tissue engineering techniques particularly using electrospun scaffolds have been intensively utilized in recent years for the development of small diameter vascular grafts. However, the development of a completely successful scaffold that fulfills multiple requirements to guarantee complete vascular regeneration remains challenging. In this study, a hydrophilic and compliant polyurethane namely Tecophilic (TP) blended with gelatin (gel) at a weight ratio of 70:30 (TP(70)/gel(30)) was electrospun to fabricate a tubular composite scaffold with biomechanical properties closely simulating those of native blood vessels. Hydrophilic properties of the composite scaffold induced non-thrombogenicity while the incorporation of gelatin molecules within the scaffold greatly improved the capacity of the scaffold to serve as an adhesive substrate for vascular smooth muscle cells (SMCs), in comparison to pure TP. Preservation of the contractile phenotype of SMCs seeded on electrospun TP(70)/gel(30) was yet another promising feature of this scaffold. The nanostructured TP(70)/gel(30) demonstrated potential feasibility towards functioning as a vascular graft. Keywords: blood vessel, nanofibrous tubular scaffold, biomechanical properties, blood compatibility, smooth muscle cells
1. Introduction Coronary artery disease is the leading cause of mortality worldwide. Limited availability of healthy autologous vessels along with the failure of synthetic grafts such as the woven Dacron and expanded polytetraflouroethylene (ePTFE), as small diameter vessels (< 6 mm) have forced bioengineers to develop alternative grafts for the regeneration of vascular tissues. Tissue engineering is a promising approach for creating small diameter vessels by combining autologous vascular cells with a natural and/or synthetic scaffold 1.
2 John Wiley & Sons, Inc.
Page 2 of 43
Page 3 of 43
Biopolymers
The rational design of a scaffold depends on its ability to mimic the native extracellular matrix (ECM) as much as possible in terms of physical, chemical and mechanical features. Development of an optimal tubular scaffold plays a significant role in vascular tissue engineering. Bioengineered scaffolds serve as the temporary matrix and provide structural, mechanical and biological support throughout the tissue formation, assisting the process of tissue remodeling2. Among the numerous techniques used to fabricate scaffolds, electrospinning has been widely used because of the high porosity, large surface area and nanofibrous structure of electrospun scaffolds, mimicking the physical nano features of native ECM 3. Furthermore, the simplicity of electrospinning technique to fabricate a tubular scaffold makes it more popular for vascular tissue regeneration applications. In vascular tissue engineering of small diameter arteries, a recurring problem is their failure owing to higher possibility of thrumbus formation which is due to experiencing lower flow rate and having higher ratio of surface area to blood volume compared to larger diameter vascular grafts which increase the time interaction with blood
4, 5
. Other common causes of long-term
failure in small grafts are compromised visco-elastic mechanical behavior of blood vessels after long-term implantation of vascular scaffolds and intimal hyperplasia at the anastomotic site, with proliferation and migration of vascular smooth muscle cells from the media to intima and subsequent synthesis of matrix proteins. Compliance mismatch between a native vessel and the implanted vascular graft leads to excessive strain on anastomosis, resulting in suture failure and consequent development of false aneurysm 6. Thus, a vascular scaffold should be made of a durable biomaterial allowing cell infiltration, adhesion and proliferation while having strength, compliance and suture retention strength comparable with those of native blood vessels
1, 2
. Blood compatibility is also clearly a crucial
3 John Wiley & Sons, Inc.
Biopolymers
Page 4 of 43
factor to raise success of the engineered constructs since the vascular scaffolds come in contact with blood. Recently several attempts have been made towards the engineering of vascular constructs that closely mimic the strength and compliance to native blood vessels
7, 8
. In addition various
approaches were carried out to design non-thrombogenic interface 9. In this study, we sought to develop a durable electrospun scaffold to mimic the fibrous structural and J-shaped mechanical properties of native arteries while blood compatibility is prioritized and considered during this study. Due to the importance of the compositional ratio of the polymer constituting an electrospun scaffold in determining mechanical properties of the scaffold 10, it is anticipated that electrospinning of a blend of compliant and stiff polymers might be a promising approach towards this goal. Polyurethane is often used in blood-contacting devices due to its thrombo-resistance, elasticity, good compatibility with tissue and blood, and resistance to mechanical degradation
11
.
Tecophilic® (TP) is a family of hydrophilic polyether-based thermoplastic aliphatic polyurethanes. It is a medical grade polymer capable of absorbing water content of up to 150% of the weight of its dry resin 12. Gelatin (gel), is a stiff and brittle biopolymer collagen, possesses many integrin binding sites
14
13
extracted from
, which makes it an appropriate candidate to
blend with TP for fabricating a composite scaffold for vascular tissue regeneration. Detta et al. used electrospun scaffolds of Tecoflex and gelatin from separate spinnerets for application in vascular tissue engineering, but reports on the blood compatibility of the scaffolds were not available
15
. In another study, Wang et al. fabricated a bilayered tubular electrospun
scaffold of PU (outer layer) and gelatin-heparin (inner layer) as an artificial blood vessel, whereby heparin provided blood compatibility of the scaffolds16. However, the lack of cell
4 John Wiley & Sons, Inc.
Page 5 of 43
Biopolymers
binding sites on the outer layer of the scaffold limits its application as a substrate for vascular tissue engineering. Therefore we aimed to investigate all aspects required for a successful tissue engineered vascular graft, during this study. It is noteworthy that a blend ratio of 70:30 for TP:gel was found in our previous study as an optimum composition to develop a suitable electrospun substrate that enhanced the contractility of smooth muscle cells (SMCs)
17
. Therefore, this composition was chosen in this study to
further examine the biomechanical and blood compatibility properties in detail.
2. Materials and Methods 2.1. Materials Tecophilic® (TP) with a product name of (SP-80A-150) was a kind gift from Lubrizol. Gelatin (gel) type A (300 Bloom) from porcine skin, bovine serum albomin (BSA), anti CD41 and 1,1,1,3,3,3-hexafluoro-2-propanol (HFP) were purchased from Sigma-Aldrich (Singapore). Human aortic smooth muscle cells (HASMC) and smooth muscle cell medium (SMCM) were obtained from ScienCell Research Laboratories (Singapore). 4′, 6-diamidino-2-phenylindole (DAPI) was purchased from Invitrogen. Anti-alpha smooth muscle actin antibody and antismooth muscle myosin heavy chain were obtained from Abcam (HongKong). 2.2. Fabrication of scaffolds Polymeric solutions were prepared by dissolving TP and gel with weight ratios of 100:0, and 70:30 in HFP to obtain a total concentration of 8% (w/v), followed by 24 h stirring at room temperature. Each solution was loaded to a 5 ml syringe capped with a 27 G blunted stainless steel needle, connected to the positive terminal of a high voltage power supply (10 kV) and a grounded flat collector wrapped in aluminum foil was placed at a distance of 12 cm from the needle tip to collect fibers. The flow rate was set on 1 ml/h using a programmable syringe pump. 5 John Wiley & Sons, Inc.
Biopolymers
Electrospun fibers were collected on 15 mm cover slips for SEM and in vitro studies while electrospun sheets were used for other analysis. Pure gelatin solution was also electrospun to provide control samples for chemical evaluation. For fabricating tubular scaffold of TP(70)/gel(30), a cylindrical mandrel of 5 mm diameter rotating at a speed of 150 rpm was used. The collector was traversed at 20 mm/s to obtain uniform thickness of the scaffold along the length of the mandrel. Electrospun scaffolds were transferred to a vacuum desiccator for at least 48 h to ensure no residual solvent on the scaffold surfaces. 2.3. Morphological and chemical Characterization Scanning electron microscopy (SEM) (JSM5600, JEOL, Japan) was used to study the morphological features of electrospun scaffolds, operated at an accelerating voltage of 15 kV. Scaffolds were coated with platinum (JEOL JFC-1300 Auto fine coater, Japan) before SEM analysis. The fiber diameter was measured from SEM micrographs using image analysis software (Image J, National Institutes of Health, USA). Results were presented in the form of mean ± standard deviation of 100 random points per image. Structural pore properties were investigated by a CFP-1200-A capillary flow porometer (PMI, New York, NY). Galwick with a defined surface tension of 15.9 dynes/cm (PMI, New York, NY) was used as the wetting liquid for porometry measurements. The pore size was measured from three samples of each type of electrospun membrane with the similar thickness of 0.35 µm and a dimension of 3 × 3 cm2. The total porosity of the electrospun scaffolds was estimated by measuring the apparent density of the electrospun scaffolds using the weights of precisely cut samples of defined area and thickness. The overall porosity was calculated according to the following equation:
6 John Wiley & Sons, Inc.
Page 6 of 43
Page 7 of 43
Biopolymers
Porosity (%) = (1 −
(1)
ρ0 ) × 100 ρ
where ρ0 and ρ are apparent density of the scaffold and density of the bulk polymer, respectively. Water contact angle as a wettability index of the electrospun scaffolds was determined using sessile drop method-based video contact angle system mounted with a CCD camera (VCA Optima, AST Products, Billerica, MA). Distilled water droplet size was set at 1.0 µl and each measurement was recorded after 20 s. For chemical analysis of the electrospun scaffolds, attenuated total reflectance Fourier transform infrared (ATR-FTIR) spectra were measured using an Avatar 380 (Thermo Nicolet, Waltham, MA) spectrometer over the range 400- 4000 cm-1 at a resolution of 4 cm-1. 2.4. Biodegradation assessment Cover slips containing electrospun nanofibrous scaffolds were placed in 24-well plate containing 1 ml of phosphate buffer solution (PBS; pH 7.4) in each well and were incubated in vitro at 37° C. PBS was refreshed after every alternate days. After a period of 7, 14, and 28 days, samples were washed and subsequently dried in a vacuum desiccator at room temperature for 48 h. SEM of the scaffolds were performed to understand the change in nanofibers morphology during these periods. For assessment of swelling ratio (water-uptake capacity) and mass loss of the scaffolds, three square-shaped sheets of each composition were immersed in PBS at 37°C for 7, 14, and 28 days. Samples were weighed before immersion (W0), after tap drying with a filter paper (W) and completely drying in a vacuum desiccator (Wd). The swelling ratio and mass loss were calculated from following formula:
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Swelling ratio (%) =
Mass loss (%) =
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(2)
W − Wd × 100 Wd
(3)
W0 − Wd ×100 W0
2.5. Biomechanical properties of the conduit Mechanical properties of the tubular scaffold were determined by uniaxial tensile testing of scaffold specimens equilibrated with PBS overnight at 37°C using a uniaxial testing system (Instron 5943) with a 50 N load cell and an extension rate of 10 mm/min. At least five specimens were tested for each mechanical evaluation. Tensile tests of ring samples with 3 mm width (W) and certain diameter (2R0) were conducted after measuring the thickness (t) on the mentioned tensile apparatus with two stainless steel hooks subsequently fixed in its clamps for gripping the ring. All ring specimens were preconditioned prior to testing with 10 cycles at 13% circumferential stretch where stretch, was defined as the sample length (L) normalized to initial length (L0). Stress-strain curves in terms of the circumferential stress against circumferential strain were computed according to current length and cross section area, shown schematically in Figure 1, with the incompressibility assumption for the tubular scaffold (equations 4-7);
A0 = 2 . W . t
(4)
F A0
(5)
L0 = π . R0
(6)
σ=
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(7)
L R A0 = = L0 R0 A
where σ is corresponding stress of applied force (F), and A is cross sectional area. Maximum stress value before failure and its corresponding value of stretch were taken from the stress-strain curve, respectively as the ultimate tensile strength (UTS) and stretch to failure (STF). Linear regression of the slope in stress-strain plots was used to calculate the elastic modulus. Elastic modulus (EM) was measured between 10 and 30% strain, termed as EM1, while EM2 was measured between the maximum strain minus 20% and the ‘maximum’ strain, with linear curve fits to these regions of the stress-strain curve 7. The results of the tensile test of ring samples were used to calculate the theoretical compliance (C) (equation 8) and burst strength (BS) of the tubular scaffold regarding thin-walled cylinder approximation relating pressure to stress scaled by wall thickness and initial radius (equation 9) and Barlow equation (equation 10) 7, 18. (Figure 1) (8)
R2 − R1 %compliance R1 = × 10 4 100 mmHg P2 − P1
σ=
Pb =
(9)
P R0 t
(10)
σ UTS t R0
where, R1 and R2 are internal radii at pressure P1 and P2, respectively. P, Pb, σ, σUTS, and R0, are transmural pressure, burst strength, tensile stress, ultimate tensile strength, and initial internal radius, respectively. Compliance was measured as the change in radius relative to the initial
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radius over the change in tensile stress at 200 and 300 kPa which are equivalent to physiological arterial blood pressure of 120/80 mmHg (systolic/diastolic) according to equation (9). Rectangular specimens were also prepared from longitudinally opened tubular scaffolds in longitudinal direction with dimensions of 20 × 5 mm2. Longitudinal samples were used to determine the suture retention force. One end of the specimen was fixed to the stage clamp of the uniaxial load test machine and a single commercial 5-0 vicryl loop (Ethicon Inc) was placed at 3 mm from the short edge of the opposite end of each sample and secured to a hook connected to the upper clamp of the testing device. The specimen was pulled at a crosshead speed of 10 mm/min until either the suture ripped or the scaffold fractured. Suture retention force (SRF) was considered as the maximum force recorded prior to failure. 2.6. Hemolysis test Hemolysis test provides information on whether the interaction of erythrocytes with biomaterial leads to hemolytic process. By providing a culture medium enriched with red blood cells, it is possible to determine whether biomaterials used in this research can destroy cells and digest the hemoglobin inside. In this regard, electrospun scaffolds were cut into small bars (about 5 × 25 mm2) and put into a vial. After rinsing bars three times with normal saline (0.9% (w/v) sodium chloride solution in distilled water), they were immersed in 1 ml of normal saline for 3 h at 37°C. Human whole blood was collected in accordance with regulations of the NUS Institutional Review Board (NUS-IRB) into blood collection tubes with 3.2% buffered sodium citrate solution. 1 ml of anticoagulant blood diluted three times with normal saline was added to each vial allowing samples to remain soaked in the solution at 37°C for 60 min. Diluted anticoagulant blood was added to distilled water and normal saline, which served as positive and negative controls, respectively. The hemolytic percentage was determined by photometric analysis of
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supernatant that was pre-centrifuged at 1000 rpm and read at 540 nm using a spectrophotometric plate reader. The hemolysis ratio (HR) was calculated as follows:
HR % =
(11)
Ats − Anc ×100 Apc − Anc
where Ats is the absorbency of test sample and Apc, Anc are the absorbance values of positive and negative controls, respectively. All hemolysis tests were done in triplicates. 2.7. Platelet adhesion assay The capability of electrospun scaffolds to prevent thrombosis was investigated by evaluating the degree of platelet adhesion on them based on two techniques (i) SEM (ii) immunofluorescent staining. Briefly, anticoagulant blood was centrifuged at 2000 rpm for 10 min at room temperature to softly sediment white and red blood cells in the lower fraction, while platelet rich plasma (PRP) remained suspended. The upper layer of the PRP was carefully collected using a pipette to limit contamination by other cellular elements and placed in a sterile tube. Scaffolds placed in 24 well plates were washed three times with distilled water and equilibrated with PBS for 30 min at 37°C, then incubated with a sample of 200 µl of PRP for 1 h at 37°C. Subsequently, scaffolds were washed repeatedly with PBS to remove un-adhered platelets. Samples were fixed using glutaraldehyde (2.5 wt-%), dehydrated with graded ethanol from 50% to 100%, dried overnight and sputter coated with platinum and were observed under SEM. In addition, the adherent platelets were stained by mouse anti-human CD41 and goat anti-mouse IgG Alex Fluor-488 conjugate (Invitrogen) and viewed using a laser scanning confocal microscope (LSCM; Olympus, FluoView FV1000). 2.8. In vitro Cell Culture HASMCs were cultured in smooth muscle cell medium (SMCM) containing basal medium complemented with 2% fetal bovine serum (FBS), 1% smooth muscle cell growth supplement 11 John Wiley & Sons, Inc.
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(SMCGS), and 1% penicillin/streptomycin solution (P/S) in a 150 cm2 cell culture flask. Cells were incubated under standard incubator (37° C, 5% CO2) for 6 days and the culture medium was changed every 3 days. Electrospun fibers collected on glass cover slips and placed in a 24 well plate were sterilized under ultraviolet (UV) radiation for 2 h, washed thrice with PBS and subsequently immersed in SMCM overnight before cell seeding. Confluent HASMCs were trypsinized with 0.25 % trypsin containing 0.1% EDTA. Deattached cells were centrifuged at 1500 rpm for 5 min, counted by trypan blue assay using a hemocytometer, and seeded on electrospun scaffolds and tissue culture polystyrene (TCP) as the control at a density of 10,000 cells per well. 2.9. Cytocompatibility and proliferation assay The viability of HASMCs on electrospun scaffolds was examined qualitatively and quantitatively using Live-Dead Cell Staining Kit (Invitrogen) and cell metabolic activity assay, respectively. Cell viability/cytotoxicity assay based on the simultaneous determination of live and dead cells was employed after 1, 4, 7, and 10 days of cell culture according to the manufacture’s protocol. Briefly, staining solution was prepared by mixing two probes of calcein AM and ethidium homodimer-1 (EthD-1) in PBS for detection of live and dead cells, respectively. Cell-seeded scaffolds were washed with PBS and incubated with 200 µl of staining solution for 30 min at 37° C. Further, samples were washed with PBS and viewed under Leica DM IRB fluorescent microscope. The proliferation of HASMCs was investigated by 3-(4, 5-dimethylthiazol-2-yl)-5-(3carboxymethoxyphenyl)-2-(4-sulfophenyl)-2H-tetrazolium (MTS) assay (CellTiter 96 AQueous One solution; Promega, Madison, WI). Briefly, after 1, 4, 7, and 10 days, cell-seeded scaffolds were rinsed with PBS, incubated with 1 ml of serum free medium containing 20 % of MTS
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reagent at 37° C for 3 h, aliquoted into a 96-well plate (100 µl/well) and absorbance of the obtained dye was measured at 490 nm using a spectrophotometric plate reader (FLUOstar Optima, BMG Lab Technologies, Offenburg, Germany). Samples with culture medium but without cells were set to determine the background absorbance to be subtracted. The intensity of obtained color is directly proportional to the metabolic activity of the cell population. 2.10. HASMC morphology on electrospun scaffolds After 4 and 10 days of culture, cell seeded constructs were harvested to observe morphological characteristics of HASMCs by SEM. Briefly, cell-cultured samples were washed with PBS, fixed with 3 % glutaraldehyde for 3 h, rinsed with distilled water and dehydrated through a graded series of ethanol solutions (50 to100 % (v/v)). Samples were then treated with hexamethyldisilazane and air-dried in a fume hood. Completely dried specimens were sputter coated with platinum and analyzed by SEM. 2.11. HASMC cellular phenotype analysis To identify whether the interaction of cells with nanofibers might affect their phenotype, the presence of two specific smooth muscle cell markers (smooth muscle alpha actin (α-SMA), and smooth muscle myosin heavy chain (SM-MHC) of cells seeded for 5 days on electrospun scaffolds were analyzed by immunofluorescent staining. Briefly, formalin-fixed cells were incubated with 250 µl of 0.1% Triton-X100 for 3min at room temperature to permeabilize the cell membrane. Triton-X100 was removed by washing samples thrice with PBS and non-specific binding sites were blocked by incubating cells with 150 µl of 3 % BSA for 90 min at room temperature. Samples were then incubated with 150 µl of either primary rabbit anti- human αSMA or mouse anti- human SM-MHC (1:200 dilution) for 2 h at room temperature under shaking condition. After three washes with PBS, either a goat anti-rabbit IgG Alexa Fluor-488
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conjugate (1:400 dilution), or a goat anti-mouse IgG Alexa Fluor-546 conjugate (1:400 dilution) (Invitrogen) was used as appropriately. Nuclei were stained with 4', 6-diamidino-2-phenylindole, Dihydrochloride (DAPI; Invitrogen) at a dilution of 1:1000. Further, cell-scaffold constructs were washed with PBS and visualized under laser scanning confocal microscopy (Olympus, FluoView FV1000) after mounting on glass sides. Cells seeded on TCP with the same treatment were set as the control. In order to investigate the cellular penetration after 5 and 10 days of cell culture, towards the interior of the composite TP(70)/gel(30) scaffolds, the immunofluorescent staining of α-SMA was performed and cell infiltration was analyzed using Imaris software. 2.12. Statistical analysis All data presented are expressed as mean±standard deviation (SD). Statistical analysis was carried out by performing one-way ANOVA followed by Tukey post hoc tests for multiplecomparison of different samples using S-PLUS (TIBCO Software Inc., USA). A value of p
