Surface Modification of Polytetrafluoroethylene (PTFE) with a Heparin-immobilized Extracellular Matrix (ECM) Coating for Small-diameter Vascular Grafts Applications

Abstract

Intimal hyperplasia, thrombosis formation, and delayed endothelium regeneration are the main causes that restrict the clinical applications of PTFE small-diameter vascular grafts (inner diameter < 6 mm). An ideal strategy to solve such problems is to facilitate in situ endothelialization. Since the natural vascular endothelium adheres onto the basement membrane, which is a specialized form of extracellular matrix (ECM) secreted by endothelial cells (ECs) and smooth muscle cells (SMCs), functionalizing PTFE with an ECM coating was proposed. However, besides ECs, the ECM-modified PTFE improved SMC growth as well, thereby increasing the risk of intimal hyperplasia. In the present study, heparin was immobilized on the ECM coating at different densities (4.89 ± 1.02 μg/cm², 7.24 ± 1.56 μg/cm², 15.63 ± 2.45 μg/cm², and 26.59 ± 3.48 μg/cm²), aiming to develop a bio-favorable environment that possessed excellent hemocompatibility and selectively inhibited SMC growth while promoting endothelialization. The results indicated that a low heparin density (4.89 ± 1.02 μg/cm²) was not enough to restrict platelet adhesion, whereas a high heparin density (26.59 ± 3.48 μg/cm²) resulted in decreased EC growth and enhanced SMC proliferation. Therefore, a heparin density at 7.24 ± 1.56 μg/cm² was the optimal level in terms of antithrombogenicity, endothelialization, and SMC inhibition. Collectively, this study proposed a heparin-immobilized ECM coating to modify PTFE, offering a promising means to functionalize biomaterials for developing small-diameter vascular grafts.

Graphical abstract

1. Introduction

Cardiovascular disease (CVD) is a leading cause of human deaths worldwide [1–4]. Vascular transplantation is an effective and common means to treat severe CVD, but the low availability of autologous blood vessels limits its clinical applications, thus stimulating the development of synthetic vascular grafts [1, 5]. Polyethylene terephthalate (PET) and expanded polytetrafluoroethylene (ePTFE) have been utilized for large-diameter vessel replacements (inner diameter > 6 mm) for several decades due to their excellent biocompatibility and robust mechanical properties, as well as commercial availability [2, 5, 6]. Unfortunately, the outcomes for small-diameter vascular grafts (inner diameter < 6 mm) remain unsatisfactory until now because of the occurrence of occlusion induced by the acute thrombogenicity, intimal hyperplasia, mechanical mismatch with native tissue, and the lack of endothelium regeneration [7–9].

In the vasculature, the interior surface is the endothelium, which is consisted of a monolayer of endothelial cells (ECs), acting as a non-thrombogenic interface and playing a crucial role in the prevention of occlusion [10–13]. More specifically, vascular ECs adhere on the vascular basement membranes (VBMs), which are thin, dense sheets of self-assembled extracellular matrix (ECM) composed of a mixture of constituents, including elastin, laminin, collagen IV, entactin/nidogen, heparan-sulfate proteoglycans, and so on [14–16]. Therefore, developing a biomimetic VBM on the luminal surface of vascular grafts seems an ideal strategy to achieve in situ endothelialization and long-term patency. Researches on biomimetic VBMs have been reported previously, such as the influences of fiber diameters, collagen IV, and laminin on EC behaviors [17–19]. Even so, it is challenging to obtain a fully biomimetic VBM, because of the complicated chemical composition as well as the complex nano- and sub-micron topography features of natural VBMs. Therefore, a complete ECM modification, which could serve a wide variety of biochemical and biophysical functions, seems to be a feasible approach, as it endows the intimal surface of small-diameter vascular grafts with improved biocompatibility. In this regard, various ECM-related materials have been explored. One of the materials was small intestine submucosa, which was an acellular matrix material with good biological properties, specifically related to limited immune reaction and rapid remodeling [20]. Another attractive material was based on the use of decellularized vessels. It had been reported that decellularized xenografts could be employed as potential scaffolds for small-diameter vascular substitutes because of their ability to facilitate the vascular-remodeling process after implantation [21, 22]. Considering the limited sources of human or animal tissues and the complex, time-consuming decellularization process, an ECM coating secreted by vascular cells was proposed as a functionalized means to modify vascular biomaterials. Huang’s group had explored introducing the ECM secreted by ECs onto Ti substrates [23, 24] and polydopamine (PDA) coated 316L stainless steel [25]. The VBM is composed of the ECMs secreted by the ECs and SMCs [10, 26], in consideration of this, Han et al. innovatively introduced a nature-inspired double-deck SMC/EC ECM onto the Ti substrate [27]. The results suggested that these ECM coatings enhanced the material’s hemo-, cyto-, and tissue compatibility significantly.

Polytetrafluoroethylene (PTFE) has been widely applied as a vascular graft material, but its bioinertness would hamper the EC attachment and endothelialization process, restricting the application for small-diameter vascular graft development [5, 7, 28]. In this study, a nature-inspired ECM coating was introduced onto the PDA/polyethyleneimine (PEI) coated PTFE via co-culture and decellularization of ECs and SMCs, followed by heparin immobilization. Heparin, a negatively charged natural polysaccharide, has been used in fabricating vascular grafts because of its thrombosis impeding property [7, 29, 30]. Moreover, it was reported that heparin could enhance EC proliferation while inhibiting SMC expansion. However, its antiproliferative property might also damage the endothelium [31, 32]. By providing an appropriate dosage of heparin when fabricating vascular grafts, the behaviors of ECs and SMCs could be selectively regulated [31, 32]. Therefore, heparin immobilization with varying concentrations (0.5 mg/mL, 2 mg/mL, 5 mg/mL, and 10 mg/mL) was evaluated in this study.

2. Materials and Methods

2.1. Materials

Medical-grade PTFE sheets with a thickness of 1 mm were purchased from Scientific Commodities Inc. (USA). Branched PEI (average Mn ~ 10000), heparin sodium salt from porcine intestinal mucosa (CAS Number: 9041081, Grade I-A, ≥ 180 USP units/mg), 2-(N-morpholino) ethanesulfonic acid (MES), 1-ethyl-3-(3-(dimethylamino)propyl)-carbodiimide (EDC), N-hydroxysuccinimide (NHS), Toluidine Blue O (TBO), and dopamine hydrochloride were all purchased from Sigma-Aldrich (USA). Human umbilical artery smooth muscle cell (HUASMC), human umbilical vein endothelial cell (HUVEC), and Human Endothelial Cell Medium kit (Catalog#: 1001) were obtained from ScienCell Research Laboratories (USA). Smooth Muscle Cell Growth Medium-2 BulletKit (Catalog#: CC-3182) was purchased from Lonza (USA). Porcine whole blood with Na citrate (Catalog#: 50414268) was bought from Fisher Scientific (USA). The 4’,6-diamidino-2-phenylindole (DAPI, Catalog#: 40043), rhodamine phalloidin (Catalog#: 00027), and live/dead viability kit (Catalog#: 30002) were purchased from Biotium (USA). Cell counting kit-8 (CCK-8, Catalog#: HY-K0301) was obtained from MedChemExpress (USA). Nitrite assay kit (Catalog#: K544) was bought from BioVision (USA), and prostacyclin ELISA kit (Catalog#: EKU08421) was obtained from Biomatik (USA). The water used in this study was purified by a Milli-Q system (Millipore, USA). All the chemicals were of analytical grade and were used without further treatment if not specially mentioned.

2.2. Surface modification of PTFE

Detailed procedures are presented in Figure 1. Briefly, dopamine and branched PEI were dissolved in the Tris buffer (pH 8.5) with concentrations of 2 and 0.5 mg/mL, respectively. The PEI concentration was controlled at a low level (0.5 mg/mL) because an appropriate dosage of PEI was beneficial for cell adhesion, while an overdose might cause cell death [7, 33]. Upon dissolution, the dopamine immediately started undergoing self-polymerization into polydopamine (PDA), reacting with PEI to form PDA/PEI conjugates [34, 35]. After ultrasonication cleaning in 20% (v/v) ethanol solution for 30 min, the PTFE sheets were immersed in a freshly prepared PDA/PEI coating solution for 24 h at 37 °C. Subsequently, the obtained samples were removed from the solution and washed three times with deionized water and named PTFE_pp. Afterward, HUVECs and HUASMCs (mixed in a 1:1 ratio) were seeded on the PTFE_PP (sterilized with 75% (v/v) ethanol) with a seeding density of 5 × 10⁴ cells/cm², and co-cultured in the medium that contained a 1:1 (v/v) mixture of the Human Endothelial Cell Medium and Smooth Muscle Cell Growth Medium-2 for 3 days under standard conditions (37 °C, 5% CO₂ under a moist environment), followed by a typical decellularization operation as reported previously [24, 27] to obtain the ECM-coated samples (ECM/PTFE_PP). Thereafter, heparin (0.5 mg/mL, 2 mg/mL, 5 mg/mL, and 10 mg/mL in MES buffer, pH 6.5) was activated in the presence of 120 mM EDC and 60 mM NHS for 4 h at 25 °C. The ECM/PTFE_PP samples were then incubated in the activated heparin sodium solutions at 25 °C overnight, obtaining the 0.5He-ECM/PTFE_PP, 2He-ECM/PTFE_PP, 5He-ECM/PTFE_PP, and 10He-ECM/PTFE_PP samples. All biological procedures were performed in a sterile environment.

Figure 1.

Schematic illustration to show the stepwise surface modification procedures of PTFE.

2.3. General characterization

The surface topographies of the samples were characterized using a scanning electron microscope (SEM, Jeol, NeoScope JCM-5000) and an atomic force microscope (AFM, Bruker BioScope Catalyst, Multimode 8) in the tapping mold. A contact angle instrument (Dataphysics, OCA15EC) was utilized to evaluate the wettability of various samples, and a volume of 5 μl of phosphate buffered saline (PBS) rather than water was dropped on the sample surfaces for the wettability test. Fourier transform infrared spectroscopy (FTIR, Bruker, Tensor 27) was used to identify the functional groups on the sample surfaces. The scanning range was set from 400 to 4000 cm⁻¹, with a resolution of 4 cm⁻¹. The surface chemical compositions were measured using the X-ray photoelectron spectroscopy (XPS, Thermo Scientific, K-Alpha), and an XPS peak-fitting software (XPS PEAK41) was applied to analyze the spectra.

2.4. Heparin density characterization–TBO assay

The immobilized amounts of heparin were characterized by the TBO assay as described in previous studies [31, 36]. In specific, the TBO solution (0.04 wt%) was prepared by dissolving the TBO powder in aqueous 0.01 M HC1/0.2 wt% NaCl, and then reacted with heparin of known concentrations to form the heparin-TBO blue complex. Afterward, the mixture was centrifuged at 3500 rpm for 10 min, and the complex was carefully rinsed with PBS and then dissolved in a 5 ml ethanol/0.1 M NaOH mixture (4:1, v/v) solution. Finally, the absorbance of the resultant solution was measured at 530 nm using a microplate reader (MULTISKAN FC, Thermo Fisher Scientific), and the standard curve was thus established. To determine the amounts of immobilized heparin, different samples were incubated in 5 ml TBO solution under static conditions at 37 °C for 4 h. Subsequently, the formed heparin–TBO complex was eluted and dissolved in 5 ml ethanol/0.1 M NaOH mixture (4:1, v/v) solution. After completing decolorization of the samples, the absorbance of the supernatants was measured at 530 nm, and the heparin amounts on sample surfaces were evaluated using the obtained calibration standard curve.

2.5. Hemocompatibility evaluation

2.5.1. Platelet adhesion test

Briefly, platelet rich plasma (PRP) was obtained by centrifuging porcine whole blood (with Na citrate) at 1500 rpm for 15 min, then a quantity of 500 μL/cm² PRP was added to the samples, which had been sterilized and rinsed with the PBS solution before testing, followed by incubation at 37 °C for 2 h. After incubation, the samples were rinsed with PBS to wash away the unattached platelets and fixed with 4% paraformaldehyde at 4 °C overnight. Next, samples were dehydrated with a series of ethanol solution (50%, 70%, 80%, 90%, and 100%, v/v) and sufficiently dried in a desiccator, then imaged using SEM.

2.5.2. Whole blood clotting test

All samples were punched into round pieces and placed in 24-well plates. 100 μl anticoagulated whole blood was dispensed onto the surfaces of pre-warmed samples, followed by adding 20 μL of 0.2 M CaCl₂ solution into the blood to start the blood coagulation cascade. After incubation for 60 min at 37 °C, 2.5 mL of distilled water was carefully added into each well, and then incubated for another 5 min at 37 °C. 100 μL of anticoagulated whole blood mingling with 2.5 mL distilled water was regarded as the control group. The absorbance of the resultant hemoglobin solution was measured at 540 nm (OD₅₄₀) with a microplate reader (MULTISKAN FC, Thermo Scientific). Finally, the blood-clotting index (BCI) was calculated using the following formula (1) [18].

$$\mathrm{BCI}=\frac{D_s}{D_0}$$ (1)

where D_s is the OD₅₄₀ value of the test sample, and D₀ is the OD₅₄₀ value of the control group.

2.5.3. Hemolysis test

The hemolysis test was conducted according to the previous literature [37, 38]. Briefly, samples (10 mg) were incubated in 10 mL PBS for 72 h at 37 °C. Then 0.2 mL diluted blood, which was prepared by diluting porcine whole blood (with Na citrate) with PBS solution at a volume ratio of 4:5, was added to each sample and incubated at 37 °C for 1 h, followed by centrifugation at 3000 rpm for 5 min. The absorbance of supernatant was measured at 545 nm with an ultraviolet spectrophotometer, and the hemolysis ratio (HR) was defined as the following formula (2). PBS solution and deionized water were applied as the negative and positive controls, respectively.

$$\mathrm{HR}=\frac{O{D}_{SA}-O{D}_{NE}}{O{D}_{PO}-O{D}_{NE}}\times 100\%$$ (2)

where OD_SA represents the absorbance of experimental samples, OD_PO represents the absorbance of positive control, and OD_NE stands for the absorbance of negative control.

2.6. Cell culture and in vitro cytocompatibility evaluation

All samples were punched and placed in 24-well plates, followed by sterilization using 75% (v/v) ethanol. Two types of cells, namely, HUVECs and HUASMCs, were separately seeded onto various samples with the same concentration of 6 × 10⁴ cells per well for cell adhesion analysis, and 1 × 10⁴ cells per well for all other characterizations. The Human Endothelial Cell Medium and Smooth Muscle Cell Growth Medium-2 were used for HUVEC and HUASMC cultures respectively. The culture medium was replaced every two days, and at least three replicates were performed on all tests for statistical analysis.

The cell adhesion test was conducted at 6 h after cell seeding. Fixed cells (in the 4% paraformaldehyde solution at 4 °C for 24 h) were stained with rhodamine and DAPI successively after sufficient PBS rinse to wash away the non-adhered cells. After imaging with a fluorescence microscope (DMI3000, Leica), the adhered cells on various samples were counted to obtain the cell adhesion densities. For the cell proliferation test, a CCK-8 assay kit was applied to evaluate cell proliferation at 1 d, 3 d, and 5 d. Cell viability was measured by the live/dead viability kit after 5 d of culture according to the manufacturer’s instructions. For the cell shape and cytoskeleton characterization, the cells cultured on different samples were stained and imaged at 3 d, and the relevant structure parameters were measured using the Image J software. A centrifugation cell adhesion assay was utilized for in situ relative cell adhesion force measurement, and the detailed procedures were performed as reported in previous studies [17, 39, 40]. For evaluating the nitric oxide (NO) and prostacyclin (PGI2) production properties of HUVECs, related ELISA kits were utilized according to the manufactures’ recommended protocols after 3 d of culture, and the cell culture medium was not replaced until testing. By doing so, the cumulative release amounts of NO and PGI2 after 3 days of culture could be estimated.

2.7. Statistical analysis

All data were presented as means ± standard deviation. Statistical analysis was analyzed using the one-way analysis of variance (ANOVA). The Tukey’s test was then applied to evaluate the specific differences between the data, and the significant differences were presented as *, **, and ***, which meant p < 0.05, p < 0.01, and p < 0.001, respectively.

3. Results and Discussion

3.1. Surface characterization

3.1.1. Surface chemistry

The surface chemistry properties of various samples were first characterized using FTIR. In Figure 2, all samples showed peaks at 1202 cm⁻¹ and 1148 cm⁻¹, and the two characteristic peaks were attributed to the asymmetrical and symmetrical C-F stretching of PTFE. For the PTFE_PP, a new peak could be found at 1655 cm⁻¹, which was ascribed to the carbonyl groups derived from PDA [41]. The broad band near 3308 cm⁻¹ resulted from the stretching vibrations of the O-H and N-H groups from PDA and PEI. The new peaks indicated the successful co-deposition of PDA/PEI on the PTFE. After the introduction of ECM coating, the intensities of both of the peaks at 1655 cm⁻¹ and 3308cm⁻¹ increased. The former was attributed to the carbonyl groups in amide I, and the latter was due to the introductions of the O-H and N-H groups from ECM proteins.

Figure 2.

FTIR spectra of different samples. The square dashed regions were enlarged for clarity.

Moreover, two new peaks appearing at 1456 cm⁻¹ and 1554 cm⁻¹ were observed, which were assigned to the bending vibrations of N-H of amide III, and C-N and N-H of amide II, respectively [42–44]. ECM is a complex system composed of proteins and polysaccharides (laminin, collagen, elastin, dextran, and so on) [10]. Therefore, the vibration peaks of amide I, amide II, and amide III suggested the successful incorporation of ECM coating. When heparin was further grafted onto the ECM/PTFE_PP, a new shoulder peak representing S=O vibration at 1040 cm⁻¹ appeared. Additionally, the immobilization of heparin introduced plenty of C=O, O-H, C-N, and N-H groups onto the ECM/PTFE_PP, thus leading to the obviously enhanced absorption at 1456 cm⁻¹, 1554 cm⁻¹, 1655 cm⁻¹, and the broad band near 3308 cm⁻¹. Meanwhile, the intensities of these peaks gradually increased with the increase of heparin density.

XPS was utilized to further analyze the surface chemical composition of various samples. The XPS spectra and elemental compositions were shown in Figure 3 and Table 1, respectively. In Figure 3A, only C1s and F1s were detected on the PTFE, while all other samples presented additional characteristic peaks of O1s and N1s. Furthermore, for the ECM-PTFE_PP and heparin-immobilized samples, new peaks corresponding to S2s and S2p could be found, indicating that the sulfur contents from the ECM coating (such as chondroitin sulfate, dermatan sulfate, heparin, heparin sulfate, and keratin sulfate) and heparin were successfully introduced. The C1s core level scans exhibiting the detailed information of carbon-containing bonds were shown in Figure 3B. As expected, when compared to the PTFE, each surface modification step obviously enhanced the C–C/C=C band, and introduced the new bands of C=O, C–O, and C–N. These new carbon peaks came from the PDA and PEI (the PTFE_PP), as well as ECM proteins (the ECM/ PTFE_PP) and heparin (heparin-immobilized samples). From the statistical results in Table 1, it could be found that N content increased from 0 for the PTFE to 5.77% for the PTFE_PP, and 9.90% for the ECM/PTFE_PP, which suggested the successful co-deposition of PDA/PEI and the subsequent ECM coating preparation. In addition, S was detected at a rate of 0.97%, 1.15%, 2.14%, and 3.01% on the 0.5He-ECM/PTFE_PP, 2He-ECM/PTFE_PP, 5He-ECM/PTFE_PP, and 10He-ECM/PTFE_PP, respectively, higher than that of the ECM-PTFE_PP (0.62%), exhibiting an uptrend. Moreover, in contrast with the ECM/PTFE_PP, the O contents of heparin-immobilized samples increased obviously, due to the fact that heparin possessed high O and S contents.

Figure 3.

XPS analysis of different samples. (A) XPS survey spectra; (B) Gauss-fitted high-resolution spectra of the carbon peak (C1s).

Table 1.

Elemental composition of samples obtained via XPS

Samples	C (%)	F (%)	N (%)	O (%)	S (%)
PTFE	31.96	68.04	0	0	0
PTFEPP	31.22	48.03	5.77	14.98	0
ECM/PTFEPP	46.97	26.57	9.90	15.94	0.62
0.5He-ECM/PTFEPP	36.96	25.79	10.19	26.09	0.97
2He-ECM/PTFEPP	37.76	22.86	10.99	27.24	1.15
5He-ECM/PTFEPP	41.24	16.01	11.61	29.00	2.14
10He-ECM/PTFEPP	42.93	11.89	16.39	25.78	3.01

3.1.2. Surface morphology

As can be seen from Figure 4, the PTFE showed a relatively smooth surface (SEM images), while nano- and micropores could be found on its surface from the high-resolution AFM characterization results. After the co-deposition of PDA/PEI, the PTFE_PP showed a rougher surface with many apparent aggregates, and the distribution of these aggregates was inhomogeneous (SEM images), in agreement with the observation of Luo et al. [45]. Its Ra value increased to 32.24 ± 8.33 nm, 2.34 times of that of the PTFE (13.76 ± 3.21 nm). It had been reported that self-polymerized PDA formed nanoparticles in an aqueous solution, and the addition of PEI could suppress the precipitation of PDA aggregates [7, 34, 35]. However, some relatively large PDA/PEI aggregates with different sizes (AFM images) were still observable in the present study. From the SEM image of the ECM/PTFE_PP, markedly less aggregates were found when compared with the PTFE_PP because most of the aggregates were washed away during the ECM preparation process. The Ra value of the ECM/PTFE_PP was 4.69 ± 1.26 nm, much smaller than that of the PTFE_PP. It could be summarized that the ECM covered the surface of the PTFE_PP, thus creating a smoother surface. Similar speculation had been reported before [24, 27]. After further heparin immobilization, more aggregates anchored onto the substrate surfaces (SEM images) when compared to the ECM/PTFE_PP. It was because heparin molecules tended to aggregate when dried from the aqueous solution [46]. Furthermore, noticeably less aggregates could be seen on the 0.5He-ECM/PTFE_PP due to its lower heparin density. From the AFM 2D and 3D images of heparin-immobilized samples, a higher heparin density led to a rougher surface, and the height differences from the cross-section profile lines confirmed this trend. Moreover, Ra values of the 0.5He-ECM/PTFE_PP, 2He-ECM/PTFE_PP, 5He-ECM/PTFE_PP, and 10He-ECM/PTFE_PP were 6.53 ± 2.25 nm, 26.93 ± 5.84 nm, 43.89 ± 8.62 nm, and 61.70 ± 11.28 nm, respectively, increasing with the increase of heparin density.

Figure 4.

Surface topography characterization by SEM and AFM (including 2D and optimized 2D images, cross-section profile lines corresponding to the red lines shown in 2D and optimized 2D images, Ra values, and 3D images).

3.1.3. Assessment of heparin immobilization and surface wettability

The heparin amounts were calculated via the TBO assay as reported in previous studies [31, 36]. Heparin could be covalently grafted to the bioactive groups provided by PEI, PDA, and ECM proteins on the ECM/PTFE_PP, including -NH₂ and -COO-. As illustrated in Figure 5A, the heparin densities on the 0.5He-ECM/PTFE_PP, 2He-ECM/PTFE_PP, 5He-ECM/PTFE_PP, and 10He-ECM/PTFE_PP were 4.89 ± 1.02 μg/cm², 7.24 ± 1.56 μg/cm², 15.63 ± 2.45 μg/cm², and 26.59 ± 3.48 μg/cm², respectively. With the increase of heparin concentration, the heparin amounts immobilized onto the samples gradually increased, consistent with the FTIR and XPS results.

Figure 5.

Heparin densities (A) and PBS contact angles (B) of different samples.

The PBS contact angle was measured to characterize the surface hydrophilicity of various samples. In Figure 5B, the virgin PTFE exhibited strong hydrophobicity, having a PBS contact angle of 121.24 ± 4.68°. After the co-deposition of PDA/PEI, the PBS contact angle dramatically decreased to 84.13 ± 2.84° because of the introduction of plenty of hydrophilic groups [7, 25, 41]. For the ECM/PTFE_PP, more hydrophilic biomolecules from the ECM proteins were immobilized onto the PTFE_PP, so it presented a further decreased PBS contact angle of 57.86 ± 5.18°, similar with the results from previous studies [25, 27]. Obviously, compared to the ECM/PTFE_PP, the 0.5He-ECM/PTFE_PP and 2He-ECM/PTFE_PP showed significantly higher PBS contact angles of 64.61 ± 4.22° and 62.19 ± 3.27°, respectively. Even though heparin was rich in hydrophilic groups (including hydroxyl, amine, carboxyl, and sulfo groups), the EDC/NHS coupling chemistry might have changed the original molecular conformation and orientation of ECM proteins, thus affecting the hydrophilicity of the ECM coating. However, with the further increase of heparin concentration, many more hydrophilic groups from heparin were introduced, thus PBS contact angles of the 5He-ECM/PTFE_PP and 10He-ECM/PTFE_PP significantly decreased to 45.72 ± 4.89° and 40.25 ± 5.96°, respectively.

3.2. Hemocompatibility evaluation

Satisfying hemocompatibility is considered to be of the most important factors that determines the long-range patency of implanted vascular grafts [1, 2, 5]. In the present study, the platelet adhesion, whole blood clotting test, and hemolysis test were conducted to evaluate the hemocompatibility.

As shown in Figure 6A, massive platelets attached onto the PTFE, obviously more than all other samples, and relatively more platelets distributed on the PTFE_PP and 0.5He-ECM/PTFE_PP as well. The amounts of adhered platelets on the ECM/PTFE_PP, 2He-ECM/PTFE_PP, 5He-ECM/PTFE_PP, and 10He-ECM/PTFE_PP were less, with no stark differences found among them. Quantitative characterization of the adherent platelets on each surface was consistent with the SEM images that the numbers of adhered platelets (platelets/mm²) on the PTFE, PTFE_PP, ECM/PTFE_PP, 0.5He-ECM/PTFE_PP, 2He-ECM/PTFE_PP, 5He-ECM/PTFE_PP, and 10He-ECM/PTFE_PP were 4511, 2340, 1591, 3150, 1805, 1652, and 1178, respectively (Figure 6B).

Figure 6.

SEM images (A) and statistical results (B) of the platelet adhesion test. BCI values (C) and hemolysis rates (D) of different samples.

Compared to the PTFE, the co-deposition of PDA/PEI significantly increased the hydrophilicity of the PTFE_PP, leading to less platelets adhered. This was in agreement with the previous finding that the increase of hydrophilicity corresponded to the decrease of platelet adhesion [7, 47]. In the presence of foreign biomaterials, albumin, fibrinogen, and other plasma proteins would be absorbed quickly, followed by attachment of platelets. The denaturation of fibrinogen is generally considered as a significant cause of platelet activation and aggregation [48–50]. Previous studies suggested that the ECMs secreted by ECs or SMCs possessed excellent ability of reducing fibrinogen adhesion and inhibiting its denaturation. Moreover, most of the platelets adhered on the ECM surfaces through accidental physical adsorption and not fibrinogen denaturation pathway [24, 25, 27]. Therefore, for the ECM/PTFE_PP, the number of adhered platelets on it would decrease. Heparin, as an anticoagulant with frequent clinical use, could prevent the protein fouling on the surfaces and inactivate proteins involved in thrombus development, thus impeding blood coagulation [32, 47, 48]. Strangely, compared to the ECM/PTFE_PP, the amount of adhered platelet on the 0.5He-ECM/PTFE_PP was significantly higher (p < 0.01). The reason was speculated that the heparin immobilization process might have damaged the integrity and biological activities of the ECM. As a result, the negative effects of heparin immobilization offset the benefit of heparin at a low density (0.5He-ECM/PTFE_PP), thus resulting in a poor antithrombotic property. With the increase of heparin density, the numbers of adhered platelets on the 2He-ECM/PTFE_PP, 5He-ECM/PTFE_PP, and 10He-ECM/PTFE_PP decreased gradually, comparable with that of the ECM/PTFE_PP, while no significant differences existed among them.

A higher BCI value indicated better anticoagulation capability. BCI values of various samples were shown in Figure 6C. Compared to the PTFE, the BCI value of the PTFE_PP significantly decreased. The reason might be that the PTFE_PP attracted much more plasma proteins adsorption (like fibrinogen) because of the presence of PDA, resulting in the blood cells attachment. After the introduction of ECM coating, the improved hydrophilicity of the ECM/PTFE_PP restricted the adsorption of plasma proteins [51, 52], such as fibrinogen [27]. However, the denaturation of absorbed proteins would be inhibited due to the favorable biocompatibility of the ECM coating. The positive effects of the ECM coating counteracted its negative effects on blood cell attachment, thus no obvious difference was found between the ECM/PTFE_PP and PTFE_PP. For heparin-immobilized samples, the 0.5He-ECM/PTFE_PP exhibited a close BCI value to the ECM/PTFE_PP because of its relatively low heparin density. With increasing heparin density, the BCI values of the 2He-ECM/PTFE_PP, 5He-ECM/PTFE_PP, and 10He-ECM/PTFE_PP presented a general uptrend, indicating their improved anticoagulation properties. The difference between the 2He-ECM/PTFE_PP and 5He-ECM/PTFE_PP was statistically insignificant, yet the 10He-ECM/PTFE_PP possessed the highest BCI value, significantly higher than all other samples. The reason might be that heparin could prevent the intrinsic blood coagulation by inactivating proteins involved in thrombus development [32, 53].

As shown in Figure 6D, all hemolysis rates were less than 2%, and no obvious hemolysis phenomenon could be observed from the inset images, suggesting that these samples could be regarded safe for direct contact with blood [38, 54]. Moreover, no significant differences could be found between the ECM/PTFE_PP and the heparin-immobilized samples, though heparin played a significant role in the management of complement mediated hemolysis [32].

3.3. HUVEC performances

3.3.1. HUVEC adhesion, proliferation, and viability

For a vascular graft, behaviors and functions of ECs that adhere on the luminal surface decide its endothelialization process, antithrombosis, and anti-hyperplasia properties to a great extent [55, 56]. Thus, the HUVEC growth and functions were evaluated to investigate the endothelialization effectiveness of various samples.

Figure 7A exhibited the fluorescence images of HUVECs adhered on various samples after a 6 h culture. It could be found that HUVECs on the PTFE did not spread actively. Their round and dwindled shapes suggested that the PTFE surface lacked biological conditions and was not suitable for cell adhesion and spreading. After surface modification, the strong expression of F-actin stress fibers suggested the improved HUVEC spreading, demonstrating the relatively high activation level of HUVECs on these samples. In Figure 7B, the HUVEC number on the PTFE was 33 cells/mm², significantly lower than that on the PTFE_PP of 139 cells/mm². It was reported that PTFE vascular grafts could not be endothelialized normally, resulting in a poor patency and a high risk of thrombosis [2]. For the PTFE_PP, native serum proteins that could interact with the adhesion receptors on the cell membrane were more likely to adsorb onto it and maintain their native structures and activities [57], thus leading to its improved cell adhesion. In addition, PEI has been reported to effectively enhance cell attachment and adhesion [33, 58].

Figure 7.

Fluorescence images (A) and numbers (B) of adhered HUVECs on different samples at 6 h.

A significant improvement in cell adhesion was achieved after the introduction of ECM coating. Compared to the PTFE_PP, the HUVEC density on the ECM/PTFE_PP significantly increased by 1.88 times to 262 cells/mm². Surface modification of vascular biomaterials with ECM components, such as some ECM proteins (like fibronectin, laminin, and collagen IV) or functional peptides (like GFPGER, YIGSR, and cyclic RGD), could improve the biological functions remarkably [17, 59, 60]. Therefore, the complete ECM would endow the vascular biomaterials with more physiological and biological functions. Relevant studies had showed that the nature-inspired ECM coating could significantly improve the material’s hemocompatibility, cytocompatibility, and tissue compatibility [24, 25, 27]. In the present study, the ECM coating provided plenty of binding sites, which could bind to the integrin proteins on the HUVEC membranes, thus enhancing HUVEC adhesion. However, compared with the ECM/PTFE_PP, the numbers of adhered HUVECs significantly decreased after heparin immobilization. HUVEC densities on the 0.5He-ECM/PTFE_PP, 2He-ECM/PTFE_PP, and 5He-ECM/PTFE_PP samples were 105 cells/mm², 68 cells/mm², and 35 cells/mm², respectively, decreasing gradually with the increase of heparin density. For the 10He-ECM/PTFE_PP, the HUVEC number slightly increased to 58 cells/mm², significantly lower than that of the 0.5He-ECM/PTFE_PP, while no significant differences could be found among the 2He-ECM/PTFE_PP, 5He-ECM/PTFE_PP, and 10He-ECM/PTFE_PP. It was speculated that the decrease of the HUVEC adhesion was due to the decline of the functional groups of ECM proteins because a number of the bioactive groups had been reacted and occupied by heparin. Furthermore, the EDC/NHS would also cross-link ECM proteins with other ECM proteins, affecting the original biological activities of the ECM coating.

Figure 8A showed the statistical results of HUVEC proliferation from the CCK-8 assay. As it can be seen, the PTFE showed the lowest values through the culture period, significantly lower than all other samples, indicating that HUVECs proliferated slowly or were not able to proliferate on PTFE materials. After the PTFE was coated with PDA/PEI, besides HUVEC proliferation, the HUVEC viability significantly increased as well, from 85.15 ± 5.06% to 92.26 ± 2.01% (Figure 8B). For the ECM-PTFE_PP, collagen IV, laminin, and other ECM proteins were introduced. They provided signaling biomacromolecules similar to the in vivo environment for the cells and abundant binding sites for cell proliferation, thus the HUVEC proliferation improved further. Compared to the PTFE_PP, the ECM/PTFE_PP showed a similar cell population at day 1, while the values increased to 2.02 times and 2.68 times as many on days 3 and 5, respectively. And its viability increased to 97.82 ± 1.85%, which was the highest value among all the samples (Figure 8B). When heparin was immobilized, these samples showed similar absorbance values to the ECM/PTFE_PP at day 1, with no significant differences. Extending the culture period, the statistical results indicated that HUVEC proliferation was negatively correlated with the heparin density. At day 3, the ECM-PTFE_PP, 0.5He-ECM/PTFE_PP, and 2He-ECM/PTFE_PP exhibited the similar HUVEC proliferation, significantly better than the 5He-ECM/PTFE_PP and 10He-ECM/PTFE_PP. At day 5, the ECM/PTFE_PP and 2He-ECM/PTFE_PP maintained excellent HUVEC proliferation. Their absorbance values were significantly higher than those of the 0.5He-ECM/PTFE_PP, 5He-ECM/PTFE_PP, and 10He-ECM/PTFE_PP. The reason was speculated that as more heparin was introduced, more functional groups of the ECM coating would be consumed. However, as a bioactive molecule with good biocompatibility, heparin could help to increase the EC proliferation by binding and stabilizing angiogenic growth factors, such as vascular endothelial growth factor (VEGF) and fibroblast growth factor 2 (FGF-2) [32, 36, 61]. Therefore, introducing an appropriate dosage of heparin could yield the best overall benefit. This was especially true for the 2He-ECM/PTFE_PP, which always exhibited similar HUVEC adhesion and proliferation results with the ECM/PTFE_PP in this study. In Figure 8B, after heparin immobilization, the HUVEC viability values of the 0.5He-ECM/PTFE_PP, 2He-ECM/PTFE_PP, 5He-ECM/PTFE_PP, and 10He-ECM/PTFE_PP were 94.45 ± 1.24%, 95.19 ± 0.88%, 93.64 ± 1.33%, and 92.96 ± 1.14%, respectively, without significant differences among them, indicating the limited effects of heparin on HUVEC viability.

Figure 8.

(A) Statistical results of HUVEC proliferation from the CCK-8 assay. (B) Statistical results of HUVEC viability from the live/dead assay at day 5.

3.3.2. HUVEC morphology

As can be seen in Figure 9A, HUVECs on the PTFE exhibited a typical non-adherent state with restricted-spread shapes. However, following a series of surface modifications, the cellular-substrate interaction was greatly improved, mainly owing to the introduction of biological functional groups from PDA, PEI, ECM components, and heparin. Thus, HUVECs on these modified samples presented ellipse, polygon, or circular morphologies and possessed significantly larger average coverage areas than the PTFE of 2211.27 μm². HUVECs on the PTFE_PP showed the largest size of 7172.46 μm², while the spread of HUVECs was restricted after the introduction of ECM coating and heparin, especially for the ECM/PTFE_PP of 4483.16 μm² (p < 0.001), 2He-ECM/PTFE_PP of 5356.43 μm² (p < 0.05), and 5He-ECM/PTFE_PP of 3852.91 μm² (p < 0.001) (Figure 9B). This was likely attributed to the significant increase of the cell population over the limited sample area, though both the ECM coating and heparin were beneficial for improving HUVEC spreading [10, 32, 62, 63]. Moreover, the HUVEC spreading area of the 5He-ECM/PTFE_PP was significantly smaller than those of other heparin-immobilized samples (0.5He-ECM/PTFE_PP of 6599.72 μm² and 10He-ECM/PTFE_PP of 5902.36 μm²). The reason might be that as more heparin was immobilized, the integrity of the ECM coating was negatively affected, subsequently leading to weakened HUVEC affinity and restricted HUVEC spreading. For the 10He-ECM/PTFE_PP, relatively low HUVEC proliferation on it contributed to more space for HUVEC spreading, thus the HUVEC coverage area value increased.

Figure 9.

Fluorescence images (A) and the coverage areas (B) of HUVECs cultured on different samples after 3 days incubation.

3.3.3. NO and PGI2 release

NO is an important endothelial-derived regulatory molecule. It possesses extensive metabolic, vascular, and cellular effects on maintaining vascular homeostasis, such as preventing the adhesion, aggregation, and activation of platelets, modulating EC functions, inhibiting the SMC proliferation, and accelerating vascular remodeling [64–67]. Like NO, PGI2 could also exert important platelet inhibitory effects and inhibit SMC proliferation, acting synergistically with NO in this regard [65, 68, 69]. Moreover, NO release of ECs could be potentiated by PGI2 [70].

Figure 10 shows the NO and PGI2 release amounts of HUVECs cultured on various samples. Obviously, the PTFE exhibited the lowest values of NO and PGI2 production. Compared to the PTFE, the PTFE_PP promoted the release of NO (p < 0.001) and PGI2 (p < 0.001). In addition, the ECM/PTFE_PP further significantly enhanced the production of NO (p < 0.001) and PGI2 (p < 0.05) in comparison with the PTFE_PP. This was consistent with the results of previous studies where the introduction of functional groups from ECM was beneficial for the NO and PGI2 productions of HUVECs [27, 69]. After the heparin immobilization, the 0.5He-ECM/PTFE_PP showed a similar NO level with the ECM/PTFE_PP, while significantly higher than those of the 2He-ECM/PTFE_PP, 5He-ECM/PTFE_PP, and 10He-ECM/PTFE_PP (Figure 10A). Similar results had also been corroborated by Liu et al.[31], who investigated the influence of heparin density on the NO release activity of HUVECs. Their study suggested that, in comparison with unmodified samples, NO release amounts significantly decreased on the samples with high heparin densities (30.9 ± 3.4 μg/cm² and 43.1 ± 5.1 μg/cm²). When the heparin densities decreased to 3–5 μg/cm², the NO release level tended to rise, and a significant improvement was achieved. Furthermore, as discussed earlier, the introduction of heparin would damage the integrity of the ECM coating and thus offset its positive effects on HUVEC behaviors, including the NO and PGI2 production. However, in Figure 10B, no significant differences could be found among the ECM/PTFE_PP, 2He-ECM/PTFE_PP, 5He-ECM/PTFE_PP, and 10He-ECM/PTFE_PP, thus indicating that heparin played a positive role in facilitating the PGI2 release of HUVECs. For the 0.5He-ECM/PTFE_PP, its PGI2 level was significantly lower than those of the ECM/PTFE_PP and 5He-ECM/PTFE_PP (p < 0.05), due to its low heparin density that was not sufficient to compensate for the negative effects of heparin immobilization process on the integrity and biological activities of ECM coating.

Figure 10.

The production levels of the NO (A) and the PGI2 (B) from HUVECs cultured on different samples.

3.3.4. HUVEC adhesion force

In vivo, ECs of blood vessels are always exposed to physiological flow conditions, so they might be sloughed off if the adhesion strength is not strong enough [71, 72]. The EC-matrix adhesion, as a crucial mechanical phenomenon in biology, affects various EC behaviors, including growth, differentiation, and migration [39, 40, 73, 74]. Therefore, the excellent EC-substrate adhesion should be considered when designing vascular grafts. Quantitative results of HUVEC adhesion forces on different samples were shown in Figure 11. The forces values of the PTFE, PTFE_PP, and ECM/PTFE_PP were 108.15 ± 28.34 pN, 1306.67 ± 303.16 pN, and 3968.12 ± 673.24 pN, respectively. ECs adhere to the substrates primarily via integrins that could bind to external ligands in the ECM, such as collagen, fibronectin, vitronectin, and so forth. Integrin clusters normally provide the primary mechanical junctions between the actin cytoskeleton and the substrates via forming focal adhesions [71, 75]. For the PTFE_PP, the existence of the PDA enhanced the adsorption of soluble adhesion proteins in the culture medium (e.g., vitronectin and fibronectin), thus its adhesion force value increased greatly when compared to the PTFE. After further introduction of the ECM coating, much more integrin-protein bonds brought about the highest adhesion force value of the ECM/PTFE_PP, significantly higher than all other samples. Since the heparin would react with the functional groups of ECM coating, and EDC/NHS coupling chemistry could impact the original biological activities of ECM proteins, HUVEC adhesion forces of heparin-immobilized samples decreased. Correspondingly, the force values decreased along with the increase of heparin density from 3012.68 ± 300.36 pN for the 0.5He-ECM/PTFE_PP to 2003.26 ± 400.19 pN for the 10He-ECM/PTFE_PP. Moreover, it should be noted that the cell adhesion force was also influenced directly by the cell spreading, as well as the topography and stiffness of the substrates [17, 73, 76].

Figure 11.

HUVEC adhesion forces on different samples.

3.4. HUASMC performances

3.4.1. HUASMC adhesion, proliferation, and viability

In natural blood vessels, healthy SMCs contribute to maintaining a normal structure and functions of the ECs, while excessive proliferation of SMCs results in late hyperplasia and increases the lumen loss, thereby narrowing the arteries [77]. Hence, in addition to ECs, the performances of HUASMCs on various samples were examined as well.

After 6 h incubation, the fluorescence images of adhered HUASMCs on various samples were taken and shown in Figure 12A. It could clearly be seen that most of adhered cells on the PTFE, PTFE_PP 0.5He-ECM/PTFE_PP, 2He-ECM/PTFE_PP, and 5He-ECM/PTFE_PP exhibited a round shape, indicating their weak and unstable adhesion status [78]. In contrast, plenty of HUASMCs on the ECM/PTFE_PP and 10He-ECM/PTFE_PP spread noticeably and presented in the shape of an elongated spindle. From Figure 12B, the PTFE presented the lowest value of 22 cells/mm². After surface modification, the HUASMC density significantly increased to 79 cells/mm² for the PTFE_PP (p < 0.05), further up to 229 cells/mm² for the ECM/PTFE_PP (p < 0.001), primarily because of the abundant binding sites provided by PDA, PEI, absorbed plasma proteins, and ECM proteins.

Figure 12.

Fluorescence images (A) and numbers (B) of adhered HUASMCs on different samples at 6 h.

Besides the anticoagulation and anti-inflammatory functions, heparin has a crucial advantage of inhibiting intimal hyperplasia when applied with an appropriate dosage. It was likely attributed to that heparin could inhibit the production of proteinases, which was related to vascular SMC matrix metabolism, thus leading to poor vascular SMC adhesion [32, 53]. Compared with the ECM/PTFE_PP, 0.5He-ECM/PTFE_PP, 5He-ECM/PTFE_PP, and 10He-ECM/PTFE_PP exhibited weakened HUASMC adhesion densities of 197, 210, and 218 cells/mm², respectively. The 2He-ECM/PTFE_PP showed a significantly lower cell density of 134 cells/mm². For the 0.5He-ECM/PTFE_PP, the fewest functional groups of the ECM coating were occupied due to its lowest heparin density, leading to no statistical difference between it and the ECM/PTFE_PP. In the meanwhile, the high heparin densities of the 5He-ECM/PTFE_PP and 10He-ECM/PTFE_PP enhanced the surface hydrophilicity and provided biological sites for the HUASMC recognition and adhesion. Therefore, the results suggested that the 2He-ECM/PTFE_PP (with a relatively low heparin density) possessed better inhibition effects on SMC adhesion, in agreement with the conclusions from previous studies [31, 62, 79].

After culture for 3 days, the morphologies of adhered HUASMCs on different samples were shown in Figure 13A. On the PTFE, HUASMCs presented an apoptotic shrinkage morphology and appeared more isolated at a low density, whereas great changes were observed on other samples where HUVSMCs grew well, and most cells were spindle shaped. It’s noteworthy that relatively higher proportions of cells on the 0.5He-ECM/PTFE_PP and 2He-ECM/PTFE_PP were in a round or “contractile” shape, indicating the inhibited biological activity of HUASMCs. From Figure 13B, it could be easily found that absorbance values of the PTFE were always the lowest throughout the entire culture period. After the introduction of PDA/PEI and subsequent ECM coating, the absorbance values of the PTFE_PP and ECM/PTFE_PP increased significantly, as well as the HUASMC viability (Figure 13C). The ECM/PTFE_PP exhibited the highest absorbance values among all the samples, suggesting the best HUASMC proliferation on it. Similar to the HUVEC, heparin inhibited the HUASMC proliferation to some degree, but showed no significant effects on the HUASMC viability (Figure 13C). According to previous studies, the inhibition effects of heparin on SMC proliferation were likely attributed to the neutralization of angiogenic growth factors (like bFGF) and the expression of contractile HUVSMC phenotype markers [47, 79]. Meanwhile, heparin could also enhance the recruitment of SMCs after binding to the stromal cell-derived factor-1a (SDF-1a) [32]. Therefore, besides its molecular weight, sulfation degree and disaccharide composition of the heparin [62], the heparin dosage that determines both stimulatory and inhibitory effects on SMC proliferation should be carefully selected [31, 79]. From the statistical results, compared with the high heparin density (10He-ECM/PTFE_PP), the low heparin densities exhibited a more significant inhibitory effect on the proliferation of HUASMCs. The absorbance values of the 0.5He-ECM/PTFE_PP, 2He-ECM/PTFE_PP, and 5He-ECM/PTFE_PP were 84.81%, 87.85%, and 85.28% at day 1; 84.93%, 87.23%, and 99.11% at day 3; and 98.31%, 89.06%, and 88.24% at day 5 respectively, of those of the 10He-ECM/PTFE_PP. Collectively, except for the PTFE, the 2He-ECM/PTFE_PP showed relatively low absorbance values during the culture period, demonstrating the weakened biological activities of the HUASMCs.

Figure 13.

(A) Fluorescence images of HUASMCs cultured on different samples after 3 days incubation. (B) Statistical results of HUASMC proliferation from the CCK-8 assay. (C) Statistical results of HUASMC viability from the live/dead assay at day 5.

3.4.2. HUASMC adhesion force

As shown in Figure 14, the HUASMC adhesion force results exhibited a similar trend with those of the HUVEC in Figure 11. The PTFE possessed the lowest value of 130.56 ± 20.34 pN, while the ECM-PTFE_PP exhibited the highest value of 1389.23 ± 498.67 pN. For heparin-immobilized samples, the HUASMC adhesion forces were 1034.89 ± 269.09 pN, 910.34 ± 190.45 pN, 820.59 ± 200.42 pN, and 861.26 ± 212.78 pN for the 0.5He-ECM/PTFE_PP, 2He-ECM/PTFE_PP, 5He-ECM/PTFE_PP, and 10He-ECM/PTFE_PP, respectively, without significant differences among these samples, indicating that the HUASMC adhesion force was not sensitive to the heparin density. Furthermore, compared to the HUVEC, the HUASMC exhibited statistically lower force values on the same samples in the present study, in agreement with the observation of a previous study [40].

Figure 14.

HUASMC adhesion forces on different samples.

4. Conclusions

PTFE has been employed in the field of cardiovascular grafts for decades, but the poor long-term patency when it is applied for small-diameter vascular grafts limits its clinical applications. In this study, the co-deposition of PDA/PEI was applied to overcome the bio-inertness of PTFE, a nature-inspired ECM coating prepared by co-culture/decellularization of ECs and SMCs was then introduced to provide a bio-favorable environment. Moreover, heparin was immobilized to the ECM coating with an optimized density to selectively promote HUVEC proliferation but inhibit HUASMC growth and obtain the satisfactory hemocompatibility. In summary, a facile surface modification method with properties of inhibiting thrombosis formation, preventing intimal hyperplasia, and promoting endothelialization was proposed in the present study. This method has great potential for functionalizing super-hydrophobic and other kinds of biomaterials for the development of small-diameter vascular grafts.

Highlights.

• The co-deposition of PDA/PEI was applied for PTFE functionalization.

• A suitable dosage of heparin could optimize the functions of extracellular matrix.

• The heparin-immobilized extracellular matrix exhibited favorable hemocompatibility.

• Behaviors of endothelial cell and smooth muscle cell were selectively regulated.

• The proposed method has great potential to modify various vascular biomaterials.