Fabrications of small diameter compliance bypass conduit using electrospinning of clinical grade polyurethane

Abstract

Objective

Compliance and viscoelastic mismatches of small diameter vascular conduits and host arteries have been the cause of conduit’s failure.

Methods

To reduce these mismatches, the aim of this study was to develop and characterize a polyurethane conduit, which mimics the viscoelastic behaviors of human arteries. Electrospinning technique was used to fabricate tubular polyurethane conduits with similar properties of the human common carotid artery. This was achieved by manipulating the fiber diameter by altering the syringe flow rate of the solution. The mechanical and viscoelastic properties of the fabricated electrospun polyurethane conduits were, then, compared with commercially available vascular conduits, expanded polytetrafluoroethylene, polyethylene terephthalate (Dacron®) and the healthy human common carotid arteries. In addition, a comprehensive constitutive model was proposed to capture the visco-hyperelastic behavior of the synthetic electrospun polyurethanes, commercial conduits and human common carotid arteries.

Results

Results showed that increasing the fiber diameter of electrospun polyurethanes from 114 to 190 nm reduced Young’s modulus from 8 to 2 MPa. Also, thicker fiber diameter yielded in higher conduits’ viscosity. Furthermore, the results revealed that proposed visco-hyperelastic model is strongly able to fit the experimental data with great precision which proofs the reliability of the proposed model to address both nonlinear elasticity and viscoelasticity of the electrospun polyurethanes, commercial conduits and human common carotid arteries.

Conclusions

In conclusion, statistical analysis revealed that the elastic and viscous properties of 190 nm fiber diameter conduit are very similar to that of human common carotid artery in comparison to the commercial expanded polytetrafluoroethylene and Dacron® that are up to nine and seven times stiffer than natural vessels. Therefore, based on our findings, from the mechanical point of view, by considering the amount of Young’s modulus, compliance, distensibility and viscoelastic behavior, the fabricated electrospun polyurethane with fiber diameter of 189.6 ± 52.89 nm is an optimum conduit with promising potential for substituting natural human vessels.

Keywords

Vascular conduit polyurethane electrospun viscoelasticity elasticity

Introduction

Commercially available prosthetic vascular conduits such as expanded polytetrafluoroethylene (ePTFE or Teflon®) and polyethylene terephthalate (PET, Dacron®) are manufactured under different conventional textile fabrication techniques (weaving, knitting and braiding).^1–3 Although commercially available conduits earn good success with approximately five years patency rate of 80% to 90% in replacing high-flow and large-diameter arteries, but none of them is appropriate for small diameter bypass conduit (<6 mm), mainly due to compliance and viscoelastic mismatches.^3–5

In fact, an ideal vascular conduit should closely mimic the biomechanical behaviors of a natural vessel, such as, compliance, distensibility, nonlinearity, viscoelasticity, and anisotropy to achieve proper transmission of propagation of pulsatile blood flow waveform across the vascular bypass conduit into the arterial system.^3,6 The mechanical behavior of a vascular conduit can be modified by altering the type of material and fabrication technique.⁷ A common material widely used in vascular conduit fabrication is medical-grade polyurethane.^7,8

Medical-grade polyurethanes exhibit more compliant as well as viscoelastic properties than other commercially available conduits because of the presence of hard and soft segments in their structures.^9,10 Polyurethanes have relatively good biocompatibility, a broad range of mechanical properties and are inherently more thromboresistant than most other polymers.^2,10–12 On the other hand, earlier studies demonstrated that electrospinning is a common and capable technique for fabrication of vascular conduits. Electrospinning as a versatile technique for creating nanofibers conduits at the scale of collagen and elastin fibers, can mimic the natural extracellular matrix (ECM) structure with controllable fiber’s diameter, fiber’s orientation, pore size distribution, porosity and microarchitecture.¹³ In addition, Khodadoust et al. revealed that the mechanical properties of electrospun polyurethane/PET are in the range of the native vessels, while commercial PET conduits are much stiffer than natural vessels.¹⁴ The mechanical behavior of electrospun polyurethane (EPU) conduits can be tailored by manipulating fiber diameter via alternation in the electrospinning parameters (i.e. syringe flow rate, collector speed, distance between syringe and collector, applied voltage, density of the polymer, etc.).^1,7,15 Also, the recent reports on small diameter polyurethane conduits showed that by altering the microstructure of the conduit using electrospinning method, a conduit with more compliance to the native vessels could be fabricated.^16,17

Furthermore, various mechanical properties of EPU conduits such as Young’s modulus, strength, strain to failure and burst pressure were investigated by tensile and biaxial burst pressure tests.^8,18 It has been known that stresses in arteries are induced mostly from the blood pressure that is in the range of 80 to 120 mmHg. Therefore, it is reasonable to calculate the tangent elastic modulus for vascular conduits in the same physiological range. To best of our knowledge, investigating the elastic modulus and viscoelastic behaviors of EPUs, ePTFE and Dacron® in the physiological range is not reported, deserving more attention and further investigation.⁷

Due to significant difference in reported mechanical properties of healthy human blood vessels, it would be hard to set a proper base of mechanical properties needed in the fabrication of artificial vascular conduits.^7,19,20 Hence, healthy common carotid artery (CCA) was chosen as a representative of human arteries to use as a guideline target of mechanical information in fabrication of vascular conduits. It should be noted that the mechanical behavior of the CCA plays an undeniable role in the formation and progression of atherosclerosis in the carotid bifurcation that is a leading cause of stroke.²¹

Therefore, the aim of this research was to fabricate proper vascular conduits with close mechanical properties as those of healthy human CCA arteries using electrospinning method and clinical grade polyurethane. To manipulate the elastic and viscoelastic properties of EPUs, conduits with three different ranges of fiber diameter were fabricated. The mechanical properties of fabricated EPUs were, then, compared with commercially available bypass conduits and human CCA. Thereafter, the optimum fabricated EPU conduit with promising potential for substituting natural human vessels was selected, by considering the amount of Young’s modulus, compliance, distensibility and viscoelastic behavior under physiological condition. At the end, a comprehensive visco-hyperelastic model was proposed to simulate both nonlinear elastic and time-dependent responses of samples.

Materials and methods

Materials

Polyurethane

A medical-grade thermoplastic PU (3485A-polyester based, $ρ =$ 1200 kg/m³) was purchased from Bayer Material Science AG, Germany. Dimethylformamide (DMF) was purchased from Sigma-Aldrich, UK.

Commercial vascular conduit

Clinical ePTFE bypass conduit with 5 mm internal diameter was supplied by GORE-TEX®, W. L. Gore & Associates, Flagstaff, USA. The Dacron bypass conduit with an internal diameter of 5 mm diameter was supplied by DuPont, Wilmington, DE, USA.

Human common carotid artery

Healthy human CCAs (n = 3) were removed from male cadaver upon permission from the donators under the ethical rules based on the 2013 Declaration of Helsinki. The age range was from 29 to 37 years and cause of the death was due to the accident. All the extracted samples were kept in normal saline and refrigerated until the tests were begun. All samples were tested within 12 h. A minimum of three samples was taken from each CCA and only the results of samples ruptured away from the grips were taken into consideration. Samples were sprayed with the saline solution to prevent drying of the tissue during the experiment.

Polyurethane vascular conduit fabrication

Polyurethane was dissolved in DMF 10% (w/v) concentration, and stirred for 24 h and finally, a homogeneous polyurethane solution containing 10 w/v %, a viscosity of 1.4 Pas and a conductivity of 12.5 µS/cm was obtained. In brief, the prepared 10% w/v solution of pure polyurethane was loaded into a plastic syringe immediately to prevent phase separation²² and were pumped at three different syringe flow rates of 0.4, 0.5 and 0.6 ml/h, in order to obtain different range of smooth fiber diameter vascular conduits based on polyurethane, named EPU_Min, EPU_Med and EPU_Max (Min, Med and Max refer to the fabricated conduits with the minimum, medium and maximum fiber diameters), respectively. The collector speed was set at 1500 r/min and a positive output lead of a high DC voltage was set at 15 KV, collected using a thin sheet of aluminum foil.⁷ The needle-collector distance of 10 cm, syringe gauge of 22 and collector diameter of 5 mm were kept constant for all three EPUs. The electrospinning duration for the samples was at least about 7 h at room temperature ( $\sim$ 24°C) and relative humidity of 40%. All test samples, including the commercial conduits, CCAs and fabricated EPUs, were ring shaped with a width of 5 mm.

Vascular conduit characterization

A representative sample from the EPUs was sputter coated with a thin layer of gold for 40 s to prevent charging of the surface and to promote the quality of imaging for the scanning electron microscope (SEM). The obtained images were analyzed with ImageJ software to determine the fiber diameter. The average fiber diameter was obtained from measuring the mean value of 40 individual fibers from each SEM. Furthermore, Fourier-transform infrared spectroscopy (FTIR; Bomen MD100) spectrum was used to identify the presence of functional groups of structure in the range of 400 to 4000 cm⁻¹.

Mechanical characterization

Mechanical properties of all samples were measured by carrying out a series of uniaxial tensile and stress relaxation tests. The mean ± standard deviation (SD) thickness and inner diameter of CCAs, Dacron®, ePTFE and EPUs were measured precisely using a micrometer with a resolution of 0.005 mm (Mitutoyo Corporation), as reported in Table 1. Ten cycles of preconditioning at a pre-load of 0.05 N were applied to each sample before any measurement.^23,24

Table 1.

Diameter, thickness and physiological stress of human CCA, ePTFE, Dacron® and fabricated EPUs.

Artery/conduit	Inner diameter (mm)	Thickness (mm)	$σ_{80}$ (KPa)	$σ_{120}$ (KPa)
CCA	4.89 ± 0.73	0.61 ± 0.37	34	52.6
ePTFE	5 ± 0.01	0.4 ± 0.0.1	66.625	99.93
Dacron®	5 ± 0.01	0.7 ± 0.01	38.07	57.10
EPU_Min	5 ± 0.01	0.32 ± 0.09	83.28	124.92
EPU_Med	5 ± 0.01	0.29 ± 0.07	91.89	137.84
EPU_Max	5 ± 0.01	0.34 ± 0.11	78.38	117.58

EPU: electrospun polyurethane.

Since the diameter of the samples is adequately larger than its wall thickness,²⁵ radial stress can be negligible compared to the circumferential and longitudinal stresses.²¹ While the circumferential stress is the dominant one, the stress of samples in physiologic range (80 to 120 ml Hg) can be determined according to the thin-walled pressure vessel equation below

σ_{θ} = \frac{\Pr}{t}

(1)

where P, r and t are internal pressure (Pa), radius (mm) and thickness (mm) of the samples, respectively. Therefore, the physiologic tangent elastic modulus (E_P) can be easily determined as the slope of the stress/strain curve at the range of physiologic stress.²⁵ The physiological stress of the samples is reported in Table 1.

Furthermore, considering thin-walled tube with a constant thickness, density and Young’s modulus at physiologic range (E_P) for all samples, the compliance and distensibility of the samples at the physiologic range can be measured using equations below

C_{P} = \frac{2 π r^{3} (1 - ϑ^{2})}{t E_{P}}

(2)

D_{P} = \frac{1}{A_{0}} C_{P} = \frac{2 r (1 - ϑ^{2})}{t E_{P}}

(3)

where r,

ϑ

and t are radius (mm), Poisson’s ratio and thickness (mm) of the samples, respectively.

Nonlinear elastic behavior

All tension tests were performed by a universal tension test device (SANTAM Corporation-STM-1) equipped with a 250 N load cell with an accuracy of ±0.01 N. The strain rate of 1%/min was employed to carry out the tension tests to obtain the nonlinear stress–strain curves of human CCAs, Dacron®, ePTFE and EPUs.

Viscoelastic properties assay

Stress relaxation test describes how the samples relieve stress with time under constant strain. In order to study the viscoelastic behaviors of human CCAs, Dacron®, ePTFE and EPUs, a series of stress relaxation tests at different strain levels (depending on the failure strain of the samples) ranging from 10% to 50% at a strain rate of 10%/min and holding time of 5 min was conducted following the preconditioning protocol.

Visco-hyperelastic model

Very few numbers of fabricated vascular conduits have been characterized by a constitutive model to capture their mechanical behavior. Therefore, a strain energy density function-based polynomial model was implemented for the nonlinear elastic behavior, and Prony series was used to characterize the viscoelastic response of the fabricated and commercial synthesis conduits and healthy human CCA. For this purpose, the experimental stress relaxation data were used to adjust the constitutive model parameters under uniaxial stress relaxation. Generally, the visco-hyperelastic constitutive model is utilized to describe materials that are able to undergo large, recoverable elastic strain and time-dependent materials such as biological soft tissues.

For isotropic hyperelastic materials, the strain energy density function (U) is a function of the right Cauchy–Green deformation tensor (C). Under the assumption that all samples are incompressible and isotropic materials, the general polynomial model is expressed as equation (4)

U = \sum_{i + j = 1}^{N} C_{ij} {(I_{1} - 3)}^{i} {(I_{2} - 3)}^{j}, i, j = 0, 1, 2, \dots, N

(4)

The special case of the general polynomial model (equation (4)) by considering N = 2 is expressed as equation (5)

\begin{array}{l} U = C_{10} (I_{1} - 3) + C_{01} (I_{2} - 3) \\ + C_{11} (I_{1} - 3) (I_{2} - 3) \\ + C_{20} {(I_{1} - 3)}^{2} + C_{02} {(I_{2} - 3)}^{2} \end{array}

(5)

where U is the function of two principal strain invariants of I₁ and I₂ as express as below

I_{1} = λ_{1}^{2} + λ_{2}^{2} + λ_{3}^{2}

(6)

I_{2} = λ_{1}^{2} λ_{2}^{2} + λ_{2}^{2} λ_{3}^{2} + λ_{1}^{2} λ_{3}^{2}

(7)

where

λ_{1}

λ_{2}

and

λ_{3}

are the principal stretch ratios. It is well known that

λ_{1} λ_{2} λ_{3} = 1

for incompressible material.²⁶ Under the assumption that all samples are incompressible and isotropic materials, the stretch ratios are as follows

λ_{1} = λ, λ_{2} = λ_{3} = \frac{1}{\sqrt{λ}}

(8)

where

λ

is the stretch ratio in the direction of the tension. So, equations (6) and (7) can be written as

I_{1} = λ^{2} + 2 λ^{- 1} and I_{2} = λ^{- 2} + 2 λ

(9)

The elongation (e) of the samples was obtained during the uniaxial tensile test. Therefore, the stretch ratio of $λ$ can be calculated according to the equation of $λ = 1 + e$ . Also, C_ij are the material constants.

In a uniaxial relaxation test, the test sample is subjected to the sudden strain and then kept constant over the duration of the relaxation test while stress is relaxed and measured over time. The initial stress is due to the hyperelastic response of the material while the stress relaxes due to its viscous response. Dimensionless relaxation modulus of the viscoelastic material, defined by Prony series expansion, defined as

K_{R} (t) = 1 - \sum_{k = 1}^{N} {\bar{K}}_{k}^{P} (1 - e^{\frac{- t}{τ_{k}}})

(10)

For hyperelastic material behavior, the stress relaxation of the samples is expressed as

C_{ij}^{R} = C_{ij}^{0} (1 - \sum_{k = 1}^{N} {\bar{K}}_{k}^{P} (1 - e^{\frac{- t}{τ_{k}}}))

(11)

where

{\bar{K}}_{k}^{P}

and

τ_{k}

are the material constants. Also,

C_{ij}^{\infty} = C_{ij}^{0} (1 - \sum_{k = 1}^{N} {\bar{K}}_{k}^{P})

, whereas

C_{ij}^{0}

and

C_{ij}^{\infty}

are the instantaneous and long-term modulus. The experimental stress relaxation data were used to adjust an isotropic visco-hyperelastic constitutive model to define the nonlinear mechanical behavior of samples under uniaxial stress relaxation.

Data and statistics

Statistical analysis was performed using Statistical Toolbox of MATLAB software. Analysis of variance (ANOVA) was carried out to study the significant effect of the independent variable on measured variables, namely, the E_P and the amount of stress relaxation in 5 min. The independent variable was the type of arteries, i.e. Commercial ones, fabricated EPUs and CCAs. Furthermore, to find the difference within the groups, a Tukey–Kramer honestly significant difference (HSD) post hoc test was carried out. A P-value of less than 0.05 was considered as significant.

Results

Vascular conduits characterization

Figure 1 presents the morphology as well as fiber diameter histogram of the fabricated conduits EPUs. The mean ± SD fiber diameter of EPU_Min, EPU_Med and EPU_Max were 114.1 ± 50.01, 142.2 ± 54.6 and 189.6 ± 52.89 nm, respectively.

Figure 1.

SEM and fiber diameter histogram of (a) EPU_Min, (b) EPU_Med and (c) EPU_Max (SEM magnification: 50 kx).

Also, the FTIR of the fabricated conduit is depicted in Figure 2 and the main peaks of used polyurethane (i.e. absorption bands of C=O (amide I) stretching and N–H bending vibrations (amide II) in the wave numbers of 1756 and 1530 cm⁻¹) were observed and supported the suggested structure.

Figure 2.

The FTIR of fabricated EPUs.

Mechanical properties studies

Human common carotid artery

The stress–strain and normalized stress relaxation-time curves of human CCAs are illustrated in Figure 3(a) and (b). According to Figure 3(a), the ultimate stress and strain of human CCAs were about 2.8 MPa and 58%, respectively. Also, Figure 3(b) depicts that the amount of relaxation of normalized stress at strain levels of 20% to 50% varied from 21% to 28%.

Figure 3.

(a) Stress–strain curve of human CCAs at a strain rate of 1%/min and (b) normalized stress relaxation–time curve of human CCAs at four different strain levels of 20%, 30%, 40% and 50%.

Commercial ePTFE and Dacron®

Figure 4(a) presents the stress–strain curve of ePTFE and Dacron®. Figure 4(a) clearly shows that ePTFE was stiffer than Dacron®. Also, normalized stress relaxation-time curves of Dacron® and ePTFE at strain levels of 10%, 20%, 30% and 40% are illustrated in Figure 4(b) and (c). According to Figure 4(b) and (c), ePTFE exhibits a more viscous behavior than Dacron®, where the amount of stress relaxation in ePTFE was about 31% to 38%, while for the Dacron® it was from 8% to 22% (Figure 4(c)).

Figure 4.

(a) Stress–strain curves of ePTFE and Dacron®, (b) normalized stress relaxation-time of ePTFE at four strain levels of 10%, 20%, 30% and 40% and (c) normalized stress relaxation-time of Dacron® at four strain levels of 10%, 20%, 30% and 40%.

Fabricated vascular conduit

Figure 5(a) and (b) demonstrates the stress–strain and normalized stress relaxation-time curves of the fabricated conduits of EPUs at three different range of fiber diameter obtained by altering the syringe flow rate of the solution from 0.4 to 0.6 ml/h, mandrel (collector) diameter and speed of 5 mm and 1500 r/min, the voltage of 15 kV and needle-collector distance of 10 cm. Although stress relaxation tests of each conduit at three different range of fiber diameter were done at five strain levels of 10%, 20%, 30%, 40% and 50%, the results obtained for the strain level of 30% (to mimic the maximum physiological strain²⁷) are only presented for three fabricated conduits due to similarity of results and also to limit the number of graphs (Figure 5). According to Figure 5(a), the ultimate stress and strain of EPU_Min, EPU_Med and EPU_Max were about 10.72 MPa and 133%, 6.13 MPa and 116% and 3.24 MPa and 130%, respectively. Figure 5(b) demonstrates that EPU_Max was more viscous than EPU_Min and EPU_Med, where the amount of stress relaxation in 5 min at strain level of 30% in EPU_Max was about 28% while for the EPU_Min and EPU_Med were 27% and 21%, respectively.

Figure 5.

(a) Stress–strain curve of fabricated EPUs at strain rate of 1%/min and (b) normalized stress relaxation-time curve of fabricated EPUs at strain level 30%.

The mean ± SD E_P, compliance and distensibility of the samples at the physiologic blood pressure range are calculated and illustrated in Figure 6(a), (b) and (c), respectively. The mean ± SD E_P of 1.51 ± 0.51, 14.23 ± 0.32, 10.7 ± 0.24, 7.51 ± 1.2, 4.61 ± 0.81 and 2.18 ± 0.42 MPa were for CCA, ePTFE, Dacron®, EPU_Min, EPU_Med and EPU_Max, respectively (Figure 6(c)). According to Figure 6(a), ePTFE was the stiffest sample, while CCA was the most compliant one among others. Also, the mean ± SD compliance of 15.19 ± 1.61, 1.39 ± 0.04, 1.06 ± 0.03, 3.31 ± 0.34, 5.94 ± 0.61 and 10.72 ± 1.09 $% / mmHg 10^{- 2}$ were for CCA, ePTFE, Dacron®, EPU_Min, EPU_Med and EPU_Max, respectively (Figure 6(b)). Figure 6(b) reveals that CCA and Dacron® were more and less compliant than others, while the compliance of EPU_Max was clearly closest to CCA among the samples. Furthermore, the mean ± SD distensibility of 1.081 ± 0.11, 0.177 ± 0.006, 0.135 ± 0.003, 0.421 ± 0.039, 0.75 ± 0.079 and 1.366 ± 0.102 $1 / mmHg 10^{- 2}$ were for CCA, ePTFE, Dacron®, EPU_Min, EPU_Med and EPU_Max, respectively. Also, EPU_Max and Dacron® were more and less distensible than other samples (Figure 6(c)).

Figure 6.

(a) E_P, (b) compliance and (c) distensibility of human CCA, ePTFE, Dacron® and fabricated EPUs.

Visco-hyperelastic model

The experimental data were used to adjust and obtain the visco-hyperelastic constitutive equation constants. Material constants representing the hyperelastic and viscoelastic response of the samples are listed in Tables 2 and 3, respectively. Also, Figure 7 presents stress–strain curves from the proposed visco-hyperelastic constitutive model fitted to experimental data of EPUs, ePTFE, Dacron® and human CCAs at a strain rate of 1%/min. Moreover, the mentioned physiologic stress range of the natural human vascular conduit is depicted in Figure 7.²⁵ Furthermore, the capability of the model to capture the hyperelastic and viscoelastic behavior of the samples is demonstrated in Figures 7 and 8.

Table 2.

Material constants of hyperelastic response section of the constitutive equation.

	Material constants: Hyperelastic response (MPa)
Artery/conduit	c₁₀	c₀₁	c₂₀	c₁₁	c₀₂
CCA	−2.978	3.275	28.09	−70.59	46.35
ePTFE	−667.98	691.67	152843.70	−343984	194788.43
Dacron®	−7.64	12.92	3612.42	−8562.95	5189.37
EPU_Min	−5.67	7.86	2.51	−10.56	14.29
EPU_Med	−5.98	7.44	0.981583	−4.90	8.73
EPU_Max	−1.23	5.89	20.18	−56.98	48.80

EPU: electrospun polyurethane; CCA: common carotid artery.

Table 3.

Material constants of viscoelastic response section of the constitutive equation.

	Material constants: Viscoelastic response
Artery/conduit	g₁	τ₁(s)	g₂	τ₂(s)	g₃	τ₃(s)
CCA	0.11459	8.8844	0.14549	148.41	−0.26008	6.48E+06
ePTFE	0.2049	6.221	1.73E-01	98.317	−0.3779	6.41E+06
Dacron®	0.17273	2.9953	5.97E-02	80.947	−0.23238	95500
EPU_Min	0.133796	12.471	9.11E-02	152.29	−0.22907	2.40E+05
EPU_Med	0.16584	11.035	9.41E-02	78.378	−0.25991	2.72E+05
EPU_Max	0.19426	7.2064	1.04E-01	117.11	−0.29861	2.09E+05

EPU: electrospun polyurethane; CCA: common carotid artery.

Figure 7.

Stress–strain curves of human CCA, ePTFE, Dacron® and fabricated EPUs at a strain rate of 1%/min from experimental data and constitutive model.

Figure 8.

Normalized stress relaxation-time curves of human CCA, ePTFE, Dacron® and fabricated EPUs at a strain level of 30%, from experimental data and constitutive model.

Statistical analysis

Statistical analysis results revealed that the type of arteries, i.e. commercial ones, fabricated EPUs and CCAs, has a significant effect on the E_P and the amount of stress relaxation in 5 min both with P < 0.0001. In addition, ANOVA post hoc Tukey–Kramer HSD multiple range tests were conducted to investigate significances among the six groups of commercial vascular conduits, EPUs and CCA, where E_P and the amount of stress relaxation in 5 min were set as the dependent variables. Post hoc comparisons indicated that the E_P of both commercial samples, i.e. ePTFE and Dacron, and that of natural arteries (CCAs) were significantly different (both with P < 0.0001). Furthermore, within EPUs, change in the fiber diameter by altering the syringe flow rate of electrospinning exhibited a strongly significant effect on E_P with P < 0.0016. Also, despite the increase in the mean amount of stress relaxation in 5 min of fabricated grafts by enhancing the fiber diameter, but there is no significant effect on the amount of stress relaxation in 5 min (P > 0.05).

However, within all groups, the difference of E_P of EPU_Max and CCA samples was not significant (P > 0.05). Also, no significant difference was found in the amount of stress relaxation of EPU_Max and that of CCA samples as well.

Discussion

It has been known that compliance and viscoelastic mismatches between the conduit and the host natural artery, particularly the anastomotic line stress concentration, lead to changes in wall shear stress tensor, pulse propagation velocity, blood flow patterns, reflection in pressure pulse and subsequently the occurrence of intimal hyperplasia, all reported as the main reason behind conduit failure.^6,16,28–31 Thus, it would be ideal to fabricate biocompatible vascular conduits with the near structural and mechanical properties as those of healthy biological arteries. To achieve this objective, the mechanical properties of the three individual human healthy CCA samples were measured and used as guidelines to tailor the properties of the EPU conduits. Furthermore, the mechanical behavior of commercial ePTFE and Dacron® conduits was also analyzed. For all groups (i.e. healthy CCA, fabricated EPUs and commercial conduits), 10 cycles of preconditioning at a pre-load of 0.05 N were applied to each sample before any measurement began. Preconditioning of the blood vessels is an important step in the mechanical testing of soft biological tissues. Earlier studies have shown that the conditioning of the tissue can be fully reached in about 10 loading–unloading cycles and increasing the number of cycles beyond that does not change the hysteresis.^23,24

One of the advantages of the fabricated EPU conduits over the commercial ePTFE and Dacron® is their nanofibrous structure which can closely mimic the structure of natural ECM, increasing the possibility of matching the biological properties of human healthy arteries, in addition to their mechanical properties.^1,32,33

The mean ultimate stress and failure strain of human CCAs in this study were measured to be 2.8 MPa and 58%, respectively, in agreement to that of about 2 MPa and 61% reported in the earlier study.³⁴ Moreover, there was reasonable consistency between the mean E_P of human CCAs (1.5 MPa) obtained in the current study to that of 0.4–1.5 MPa range.²¹ The same mechanical test procedures used for human CCAs were utilized to examine the fabricated EPUs and commercial conduits and the results were compared together. To find the optimum sample for substituting natural blood arteries, both stress–strain curves, especially E_P as well as viscoelastic properties were considered.

The obtained results for the mean E_P of ePTFE and Dacron® (14.23 and 10.7 MPa) in the present study is in good agreement to that of 13.24 and 9.4 MPa, as reported by Pietrabissa.³⁵ By exploring Figures 6(a) and 7, the results clearly revealed that ePTFE was the stiffest, while the human CCAs were the most compliant ones among the others. E_P of Dacron® was also close to that of ePTFE and still much stiffer than natural human CCAs, while the average E_P of fabricated EPUs was in the range of 2.18 to 7.51 MPa where the lowest number was obtained for EPU_Max, the closest to that of human CCAs (P > 0.05).²⁵

Based on the mechanical tests conducted on the fabricated EPUs, and taking into account the electrospinning procedure, we arrived at the conclusion that by decreasing the syringe flow rate from 0.6 to 0.4 ml/h at a constant needle-collector distance of 10 cm, more time for the solvent of EPUs was required to reach the collector, which leads to the reduction of the fiber diameter as suggested by other investigators,³⁶ as presented in Figure 1. In accordance with the results of earlier works,^7,37–39 it was shown that the reduction of fiber diameter enhances the stiffness and decreases the compliance and distensibility of the conduits, as illustrated in Figures 5, 6(a) and 7.

The recorded stress relaxation of human CCA samples at a strain level of 30% was found to be about 26% slightly lower than that of 30% reported in the investigation of porcine CCA.⁴⁰ Also, the measured stress relaxation of ePTFE (38%) in this study is well in the range, as reported previously.⁴¹

Normalized stress relaxation-time curves of all samples from examined constitutive model compared with experimental data at the strain level of 30% are depicted in Figure 8. According to this figure, at a strain level of 30%, ePTFE was more viscous than other samples with a stress relaxation of 38%, while Dacron® and EPU_Min were less viscous than the rest with about 21% relaxation. The stress relaxation of human CCAs and EPU_Med and EPU_Max were about 26%, 27% and 28%, respectively, i.e. almost similar (P > 0.05).

Taken together, the results suggest that using commercial vascular conduit, EPU_Min and EPU_Med have a remarkable effect on the formation of intimal hyperplasia due to the compliance and viscoelastic mismatches, especially in small diameter conduits. However, using the fabricated EPU_Max could minimize the consequences of compliance and viscoelastic mismatches of the conduit and the host natural artery.

Furthermore, the results revealed that proposed visco-hyperelastic model is strongly able to fit experimental data (R² of 0.99) with great precision which proofs the reliability of the proposed model to predict both nonlinear elasticity and viscoelasticity of the EPUs, commercial conduits and human CCAs as observed in Figures 7 and 8.

In conclusion, the elastic properties of the fabricated EPU_Max are very similar to that of human CCA in comparison to the commercial ePTFE and Dacron® that are up to nine and seven times stiffer than natural vessels. Furthermore, the viscoelastic behavior of the EPU_Med and EPU_Max matches well with that of human CCA. Therefore, based on our findings, from the mechanical point of view, by considering the amount of Young’s modulus, compliance, distensibility and viscoelastic behavior, the fabricated EPU_Max is an optimum conduit with promising potential for substituting natural human vessels. This study overlooked the examination of endothelial cell behavior on the prepared EPUs which will be carried out in future studies.

Footnotes

Declaration of conflicting interests

The author(s) declared no potential conflicts of interest with respect to the research, authorship, and/or publication of this article.

Funding

The author(s) received no financial support for the research, authorship, and/or publication of this article.

References

Singh

Wong

Wang

Medical textiles as vascular implants and their success to mimic natural arteries. J Funct Biomater 2015; 6: 500–525.

Seifalian

, et al. In vivo biostability of a poly (carbonate-urea) urethane graft. Biomaterials 2003; 24: 2549–2557.

Kannan

, et al. Current status of prosthetic bypass grafts: a review. J Biomed Mater Res 2005; 74: 570–581.

Ravi

Chaikof

EL.

Polymeric materials for tissue engineering of arterial substitutes. Vascular 2009; 17: 45–54.

Teebken

Haverich

Tissue engineering of small diameter vascular grafts. European J Vasc Endovasc Surg 2002; 23: 475–485.

Sarkar

, et al. The mechanical properties of infrainguinal vascular bypass grafts: their role in influencing patency. Eur J Vasc Endovasc Surg 2006; 31: 627–636.

Hasan

, et al. Electrospun scaffolds for tissue engineering of vascular grafts. Acta Biomater 2014; 10: 11–25.

, et al. The preparation and performance of a new polyurethane vascular prosthesis. Cell Biochem Biophys 2013; 66: 855–866.

Gadley

Andrade

Maia

JM.

Effect of soft‐to‐hard segment ratio on viscoelastic behavior of model thermoplastic polyurethanes during phase transitions. Macromol Mater Eng 2016; 301: 953–963.

10.

Tiwari

, et al. New prostheses for use in bypass grafts with special emphasis on polyurethanes. Cardiovasc Surg 2002; 10: 191–197.

11.

Zhang

, et al. Pore size, tissue ingrowth, and endothelialization of small-diameter microporous polyurethane vascular prostheses. Biomaterials 2004; 25: 177–187.

12.

Salacinski

, et al. Performance of a polyurethane vascular prosthesis carrying a dipyridamole (Persantin) coating on its lumenal surface. J Biomed Mater Res 2002; 61: 337–338.

13.

Khorshidi

, et al. A review of key challenges of electrospun scaffolds for tissue‐engineering applications. J Tissue Eng Regen Med 2016; 10: 715–738.

14.

Khodadoust

Mohebbi-Kalhori

Jirofti

Fabrication and characterization of electrospun Bi-hybrid PU/PET scaffolds for small-diameter vascular grafts applications. Cardiovasc Eng Tech 2018; 9: 73–83.

15.

Inoguchi

, et al. Mechanical responses of a compliant electrospun poly (L-lactide-co-ε-caprolactone) small-diameter vascular graft. Biomaterials 2006; 27: 1470–1478.

16.

Nezarati

, et al. Electrospun vascular grafts with improved compliance matching to native vessels. J Biomed Mater Res 2015; 103: 313–323.

17.

Montini-Ballarin

, et al. Mechanical behavior of bilayered small-diameter nanofibrous structures as biomimetic vascular grafts. J Mech Behav Biomed Mater 2016; 60: 220–233.

18.

Theron

, et al. Modification, crosslinking and reactive electrospinning of a thermoplastic medical polyurethane for vascular graft applications. Acta Biomaterialia 2010; 6: 2434–2447.

19.

Stekelenburg

, et al. Dynamic straining combined with fibrin gel cell seeding improves strength of tissue-engineered small-diameter vascular grafts. Tissue Eng Part A 2008; 15: 1081–1089.

20.

Soletti

, et al. A bilayered elastomeric scaffold for tissue engineering of small diameter vascular grafts. Acta Biomaterialia 2010; 6: 110–122.

21.

Kamenskiy

, et al. Nonlinear mechanical behavior of the human common, external, and internal carotid arteries in vivo. J Surg Res 2012; 176: 329–336.

22.

Jing

, et al. Electrospinning thermoplastic polyurethane/graphene oxide scaffolds for small diameter vascular graft applications. Mater Sci Eng C 2015; 49: 40–50.

23.

Ferruzzi

Bersi

Humphrey

Biomechanical phenotyping of central arteries in health and disease: advantages of and methods for murine models. Ann Biomed Eng 2013; 41: 1311–1330.

24.

Fung

Y-C.

Biomechanics: mechanical properties of living tissues. Berlin: Springer Science & Business Media, 2013.

25.

Kamenskiy

, et al. Biaxial mechanical properties of the human thoracic and abdominal aorta, common carotid, subclavian, renal and common iliac arteries. Biomech Model Mechanobiol 2014; 13: 1341–1359.

26.

Karimi

, et al. A nonlinear hyperelastic behavior to identify the mechanical properties of rat skin under uniaxial loading. J Mech Med Biol 2014; 14: 1450075.

27.

Bluestein D and Einav S. Techniques in the analysis of stability of pulsatile flow through heart valves, In: Leondes CT, Biomechanical Systems: Techniques and Applications: Cardiovascular Techniques, Vol. II, ch. 4, CRC Press, 2001, pp. 14–39.

28.

Haruguchi

Teraoka

Intimal hyperplasia and hemodynamic factors in arterial bypass and arteriovenous grafts: a review. J Artif Organs 2003; 6: 227–235.

29.

Ostdiek

, et al. An in vivo study of a gold nanocomposite biomaterial for vascular repair. Biomaterials 2015; 65: 175–183.

30.

Chlupac

Filova

Bacakova

Blood vessel replacement: 50 years of development and tissue engineering paradigms in vascular surgery. Physiol Res 2009; 58: S119.

31.

Zahedmanesh

Lally

A multiscale mechanobiological modelling framework using agent-based models and finite element analysis: application to vascular tissue engineering. Biomech Model Mechanobiol 2012; 11: 363–377.

32.

, et al. Aligned biodegradable nanofibrous structure: a potential scaffold for blood vessel engineering. Biomaterials 2004; 25: 877–886.

33.

Stitzel

, et al. Controlled fabrication of a biological vascular substitute. Biomaterials 2006; 27: 1088–1094.

34.

Tseders

Purinya

The mechanical properties of human blood vessels relative to their location. Polymer Mechanics 1975; 11: 271–275.

35.

Pietrabissa

Biomateriali per protesi e organi artificiali. Bologna: Patron Editore, 1996.

36.

Sill

von Recum

HA.

Electrospinning: applications in drug delivery and tissue engineering. Biomaterials 2008; 29: 1989–2006.

37.

Bhardwaj

Kundu

SC.

Electrospinning: a fascinating fiber fabrication technique. Biotechnology Advances 2010; 28: 325–347.

38.

McClure

, et al. The use of air-flow impedance to control fiber deposition patterns during electrospinning. Biomaterials 2012; 33: 771–779.

39.

Yao

Bastiaansen

Peijs

High strength and high modulus electrospun nanofibers. Fibers 2014; 2: 158–186.

40.

García

Martínez

Pena

Viscoelastic properties of the passive mechanical behavior of the porcine carotid artery: influence of proximal and distal positions. Biorheology 2012; 49: 271–288.

41.

Sarkar

Development of a synthetic small calibre vascular bypass graft. London: University College London, 2011.