An extracellular matrix-mimetic coating with dual bionics for cardiovascular stents

Abstract Anti-inflammation and anti-coagulation are the primary requirements for cardiovascular stents and also the widely accepted trajectory for multi-functional modification. In this work, we proposed an extracellular matrix (ECM)-mimetic coating for cardiovascular stents with the amplified functionalization of recombinant humanized collagen type III (rhCOL III), where the biomimetics were driven by structure mimicry and component/function mimicry. Briefly, the structure-mimic was constructed by the formation of a nanofiber (NF) structure via the polymerization of polysiloxane with a further introduction of amine groups as the nanofibrous layer. The fiber network could function as a three-dimensional reservoir to support the amplified immobilization of rhCoL III. The rhCOL III was tailored for anti-coagulant, anti-inflammatory and endothelialization promotion properties, which endows the ECM-mimetic coating with desired surface functionalities. Stent implantation in the abdominal aorta of rabbits was conducted to validate the in vivo re-endothelialization of the ECM-mimetic coating. The mild inflammatory responses, anti-thrombotic property, promotion of endothelialization and suppression of excessive neointimal hyperplasia confirmed that the ECM-mimetic coating provided a promising approach for the modification of vascular implants.


Introduction
Percutaneous coronary intervention (PCI) with a vascular stent is the most successful treatment for coronary heart diseases [1,2]. Although the risk of in-stent restenosis (ISR) has decreased after the advent of drug-eluting stent (DES) [3,4], anti-proliferative drugs on the stent [such as paclitaxel (PTX) and sirolimus] for suppressing the proliferation of smooth muscle cell (SMC) also inhibit the endothelialization process after stent implantation. The delayed endothelialization caused by these drugs poses a risk for the incomplete healing of neointima and thus leading to late thrombosis (LST), threatening the life of patients [5,6]. Currently, bioresorbable vascular stents (BRS) are designed to eliminate foreign body reactions (FBR) that arise from non-degradable metallic stents [7], since surrounding tissues absorb it completely and leave only healthy vessels [8,9]. However, currently applied coatings on BRS in the clinic follows the strategy of the coatings on non-degradable metallic DES, which means LST may also occur on BRS. Therefore, it is essential that desired BRS coating should induce a mild inflammatory response, suppress the excessive hyperplasia of SMCs and promote re-endothelialization [10].
Resting endothelial cells (ECs) are known to inhibit coagulation, but during PCI surgery, injury and stimulation can contribute to a decrease in anti-coagulation and leukocyte recruitment [9]. In addition, rapid inflammatory responses occur, accompanied by monocyte aggregation around the stent and their differentiation into macrophages [10,11]. The damage to endothelial layer during PCI is unavoidable, as both acute inflammatory responses and thrombus deposition can lead to delayed endothelialization and excessive hyperplasia, eventually resulting in the failure of stent implantation [11,12]. Since many complicated biological responses around the material/vessel interface have a vital impact on the efficiency of implantation, multi-functional modification, including anti-inflammation, anti-coagulation, and rapid endothelialization functions, is a promising trajectory for the design of an ideal coating for vascular stent [13][14][15][16]. Many researchers have constructed a multi-functional surface via multiple substances. Yang et al. reported a phenolic-amine chemistry-mediated coating, which suppressed the inflammation via tannic acid (TA) and enhanced the anti-coagulation ability via thrombin inhibitor bivalirudin (BVLD) [17]. In addition, the vascular endothelial growth factor (VEGF) is widely applied to vascular stent modification. Sun et al. designed a stent coated with bilayered PLGA nanoparticles containing VEGF and PTX [18]. PTX suppressed SMC proliferation, and VEGF promoted rapid reendothelialization. Biomaterials, such as TA, hyaluronic acid (HA), and epigallocatechin gallate, have been widely applied in the modification of cardiovascular stents. These materials are often combined with other substances (i.e. heparin, BVLD and PTX) to create stent coatings [19][20][21]. The application of most multifunctional coatings is limited by the high price and complicated assembly process. In addition, the introduction of multiple biomaterials may impact the functions of these substances. Hence, introducing a single biomacromolecule to meet the multifunctional requirements of stents could be a novel and promising method for modifying cardiovascular materials.
Recombinant humanized collagen is a designable, water-soluble and biocompatible material [22][23][24]. However, its application for blood-contacting modification is impeded by the activation and adhesion of platelets. Considering this, we designed a tailored recombinant protein containing cell adhesive motives, Gly-Glu-Arg (GER) and Gly-Glu-Lys (GEK) triplets [25], referred to as recombinant humanized collagen type III (rhCol III). Compared to most recombinant collagens, rhCol III shows great anti-coagulation properties, due to the deletion of hydroxyproline-containing fragments, which is known to bind surface integrin a2b1 to activate platelets [26]. In addition, rhCol III also exhibits anti-inflammatory capacities and selectively promotes EC proliferation [22,24,26]. As previously reported, the layer-by-layer (LBL) coating of HA and rhCol III is a promising blood-contact material [26]. However, it is still far from ideal as a vascular stent coating due to the tedious assembly process of LBL, the lack of stability, and the limited introduction of rhCol III. Therefore, it is important to construct a stable coating using single rhCol III and explore the efficiency of its functions in response to complicated biological environments.
Mimicking the extracellular matrix (ECM) is a widely accepted approach for modifying scaffolds. The inherent fiber structure and functional matrix components (i.e. collagen, laminin and fibronectin) could help to support cell/material interactions [26][27][28][29][30]. To design an ECM-mimetic coating with both structuremimic and function-mimetic could benefit from topographical surface and multiple functions of components. The polymerization of polysiloxane has been used to construct fibrous or porous network [31][32][33]. Furthermore, the polymerization of polysiloxane to form a nanofiber (NF) structure can enable the loading of functional components onto a three-dimensional reservoir and then the immobilization of more functional components. Based on this, we designed an ECM-mimetic coating with amplified functionalization of rhCOL III, where the biomimetics were driven by both structure mimicry (a fibrous network layer) and component/function mimicry (an rhCOL III-based matrix that supports cell growth and anti-coagulant ability).
As shown in Fig. 1, the ECM-mimetic coating was prepared in three steps, including basic NF coating formation, aminofunctionalized substrate construction, and rhCol III covalent immobilization. It was designed to inhibit inflammation, reduce thrombosis, suppress intimal hyperplasia, and promote endothelialization. The NFs were formed on a Fenton-treated surface, where oxidative functional groups were introduced via the hydrolyzation and cross-linking of methyl trichlorosilane (MTS), followed by further introduction of amino groups by (3-aminopropyl) triethoxysilane (ATPES), which hydrolyzed and bound to the Fenton-treated NFs. The rhCol III was immobilized via the condensation reaction (carboxyl and amino groups) to construct the dual ECM-mimetic coating. Subsequently, the hemocompatibility and cytocompatibility were evaluated in vitro. Animal experiments were performed to test the anti-thrombosis and anti-inflammatory capacities, and in vivo vascular implantation was conducted to evaluate the potential of endothelialization and anti-hyperplasia. For further comparison, rhCol III was directly immobilized on the PLA substrate, to demonstrate the positive effect of the nanofibrous layer (NF layer) on rhCol III immobilization.

Fabrication of rhCol III coatings
The PLA sheets were successively cleaned with acetone, ethanol, and ultrapure (UP) water. Hydroxyl groups on the PLA substrate were introduced via Fenton's reagent (pH 2.5-3.0), in which FeSO 4 and hydrogen peroxide (H 2 O 2 ) reacted with the substrate. The plates were immersed in hexane containing 0.2% vol MTS for 12 h, prior to being covered by NFs. Before coating of rhCol III, the PLA plates and stents were treated by Fenton reagent and then immersed in the ethanol of 2% vol ATPES for 2 h to form amino groups. The substrates were washed with ethanol and UP water, and immersed in the MES buffer solution (pH 5.0) containing EDC (0.08 M) and rhCol III (2 mg/ml). The sheets only covered by rhCol III were omitted in the preparation of NF.

Characterization of rhCol III coatings
The elemental composition was analyzed by X-ray photoelectron spectroscopy (XPS, Thermo Fischer, ESCALAB Xiþ, USA). The chemical bonds were evaluated by attenuated total reflectance-Fourier-transform infrared spectrometry (ATR-FTIR, Thermo Fisher Nicolet Is5, USA). The morphology of the samples and the diameter of NF were observed using a scanning electron microscope (SEM; FEI, USA) and atomic force microscopy (Bruker Dimension Icon AFM, USA). Water contact angles (WCA) were estimated using ThetaFlex (Biolin Scientific, Finland).

Stability tests
The FITC-rhCol III was synthesized by mixing of rhCol III and FITC in a saturated sodium bicarbonate solution for 24 h. The solution was dialyzed in the dialysis tubing (3.5 kD) for 7 days and then freeze-dried. The preparation of FITC-rhCol III coatings was the same, but rhCol III was replaced with FITC-rhCol III. The samples were immersed in PBS on the shaker at 37 C and removed after 7, 14 and 28 days to observe the distribution of FITC-rhCol III under a laser-scanning confocal microscope (LSM880 Airyscan with STEDYCON, Carl Zeiss, Germany).

Platelet adhesion test
Platelet-rich plasma (PRP) was collected from the supernatant of New Zealand white rabbit blood centrifugated at 1,500 r/min for 15 min. The PRP was incubated on the samples for 1 h at 37 C. Next, the samples were washed with PBS and immersed in 2.5% glutaraldehyde solution overnight for fixation [14,34]. Then, the platelet morphologies were observed by SEM after dehydration by gradient alcohol [35].

Ex vivo blood circulation thrombogenicity test
The ex vivo blood circulation test was carried out on New Zealand white rabbits (2.5-3 kg), as the arteriovenous shunt model is a common approach to evaluate blood compatibility. The coatings of rhCOL III with and without NF were prepared on the polyvinyl chloride tubes. The tubes were connected from the carotid artery to the jugular vein to form a blood circuit for 2 h. Then, the blood clots and the cross-section of the tubes were photographed and recorded for further evaluation of thrombosis.

Smooth muscle cells compatibility test
Primary human umbilical artery smooth muscle cells (HUASMCs) were grown in DMEM containing 10% (v/v) FBS and 1% (v/v) penicillin-streptomycin. The subsequent processes were similar to those mentioned above in the HUVEC compatibility test.
Macrophage cell compatibility test RAW 264.7 mouse leukemia cells were cultured as the similar method mentioned in the HUVECs compatibility test, and the media after incubation were collected to determine the inflammatory factors. The expression of tumor necrosis factor-alpha (TNF-A) and interleukin-10 (IL-10) in cell supernatant was measured by ELISA Kit (Jiangsu Meimian Industrial Co., Ltd, Yancheng, China).

Subcutaneous implantation test
The animal experiments were carried out in accordance with the Animal Care and Use Committee of Sichuan University. Male adult Sprague-Dawley rats (Chengdu Dashuo Experimental Animal Co., Ltd, Chengdu, China) under anesthesia were implanted with 6 mm Â 6 mm samples subcutaneously. The samples with surrounding tissue were removed and fixed in 4% paraformaldehyde for at least 24 h after 15 and 30 days. The surrounding tissue was stained with hematoxylin and eosin (H&E), as well as immunofluorescence (anti-CD68), to evaluate the thickness of the fibrous capsule and inflammation.

Stent implantation evaluation
The rhCol III (NF) and bare stents were crimped before the experiment. Then, the sterilized stent was implanted in the abdominal aorta of the rabbit. After 1 and 3 months of implantation, the stents and surrounding vascular tissues were excised and washed with 800 U heparin solution to avoid clot deposition in the separate vessels. Then, the vessels containing stents were immersed in 4% paraformaldehyde for 1 week and stained with H&E and immunohistochemistry antibodies [anti-CD68, -CD31, -eNOS and -A-smooth muscle actin (a-SMA)] for further evaluation.

Statistical analysis
The stained slices were analyzed by Image-Pro Plus 6.0, and statistical analyses (more than three parallel samples) were conducted using GraphPad Prism 8.0, including one-way analysis of variance (ANOVA) and t-test to determine the statistical significance (*P < 0.05, **P < 0.01, ***P < 0.001 and ****P < 0.0001). All of the results are expressed as mean 6 standard deviation (SD).

Results and discussion
Characterization of the NF substrates The synthesis of NFs was conducted through the hydrolysis of MTS, which reacted with the hydroxy groups of the substrate treated with Fenton's reagent and induced the polymerization of polysiloxane on the surface. As shown in Fig. 2A, the surface of PLA was smooth and flat. When the substrate was covered by NFs, the roughness increased from 1.73 to 211 nm (Fig. 2B), and the mean diameter of NF was about 119.4 nm. After undergoing the Fenton and ATPES processes, NFs were linked with hydroxyl groups and amino groups, and the mean diameters were 232.6 and 343.1 nm, respectively ( Fig. 2A). The roughness was significantly different between amino-silanized bare PLA (RMS ¼ 4.41 nm) and NF substrate (RMS ¼ 337 nm) on account of the surface-to-volume ratio (Fig. 2B). This indicates that more NH 2 groups could be immobilized on the PLA surface. The increase in the nitrogen (N) content in Fig. 2D also confirmed that more amino groups were connected to the NF substrate. In Fig. 2C, the FTIR spectra of the NF substrate had a new peak at 1070 cm À1 , compared with the bare PLA, and this peak was assigned to the stretching vibration of Si-O-Si [32]. After APTES modification, the peak near 1270 cm À1 was enhanced due to the Si-C bond [33]. In addition, the peak near 1600 cm À1 was attributed to carbonyl stretching, and its intensity weakened because of the complete coverage of NFs. In contrast, the spectra of NH 2 (PLA) were similar to the bare PLA because the coating thickness was not sufficient to measure, while the thickness of NH 2 (NF) coatings was 1.42 mm according to Fig. 2E.
In summary, the fibrous network layer was successfully prepared on the PLA substrate to mimic ECM structure and provided a three-dimensional reservoir for more amino immobilization. In addition, the layer was hydrophobic because the WCA is about 162.6 ( Supplementary Fig. S1). The NFs on the surface presented overhanging morphology and low surface energy. Thus, the water drops rested on the protrusions, entrapping air underneath and eventually rolling off the substrate [31,32].

Characterization of the rhCOL III coatings
FITC-rhCOL III and rhCOL III were immobilized on the aminosilanized NF substrates as mentioned above. The rhCOL III particles and NH 2 groups were closely integrated via amide bonds, due to a chemical reaction between rhCOL III carboxyl and amino on the substrate (in MES buffer and EDC activation). Figure 2F presents the morphology after the immobilization of rhCOL III, and the conglutination of spherical particles was observed around NFs. In Fig. 2G, the fluorescence signals on FITC-rhCOL III (NF) were much stronger than that on FITC-rhCOL III (PLA) coatings, which confirmed that more rhCOL III was immobilized on the NF substrate. This phenomenon indicates that the fiber network interface increased the amount of rhCOL III immobilization.
The stability of the coatings has a significant impact on the effectiveness of implantation, and the function of the coatings should be maintained for at least 7 days, as a modified bloodcontact device. As shown in Fig. 2H, the PLA stent coated with FITC-rhCOL III (NF) presented strong fluorescence signals before immersion. To test the coating stability, the FITC-rhCOL III (NF) stent was immersed in PBS solution, and the loss of rhCOL III after 7, 14, and 28 days was observed. There was no major attenuation in the fluorescence on the immersed stent after 7 days. After 14 days of immersion, a slight decrease in fluorescence indicated that the rhCOL III content of the coating decreased. Fluorescence signals were detected after 28 days of immersion, implying that the coating had good stability and could maintain long-term function after implantation.
The above results indicate that the introduction of a fibrous network layer increased the space for rhCOL III immobilization and the coating was stable due to the covalent binding between rhCOL III and NF layer. With the immobilization of hydrophilic rhCOL III, the rhCOL III (NF) coating was hydrophilic ( Supplementary Fig. S1) on account of a 2D capillary effect [38]. The dual ECM-mimetic coating was a biomimetics driven by a fibrous network layer and a rhCOL III-based matrix. For further investigation of the advantages of dual bionics, the rhCOL III (PLA) coating (single ECM component mimetic) was prepared and evaluated.

Anti-coagulant performance
Incubation with PRP to observe platelet adhesion and activation on the substrate is a common way to test the anti-coagulant performance ex vivo. As shown in Fig. 3A, a large number of activated platelets were adhered to the PLA sheet, as observed by SEM. In contrast, platelets were fewer and exhibited low activation on the rhCOL III (PLA) and rhCOL III (NF) substrates. After analyzing the density of platelets on the sheets in Fig. 3B, rhCOL III could obviously suppress platelet adhesion and activation and the anticoagulation of rhCOL III (NF) coating was better. The rhCOL III (NF) substrate in Fig. 3A displays a larger image of the platelets on the rhCOL III (NF) coating, and the platelets observed were not active. A previous study demonstrated the efficiency of rhCOL III in repressing platelet adhesion [26], while the NF layer exhibited a micro/nanostructure and decreased the contact point for platelet adhesion [32]. In brief, the rhCOL III (NF) coating amplified the anti-coagulation of rhCOL III and influence platelet/material interaction, which had a positive effect on the resistance of platelet adhesion.
To further investigate the anti-thrombotic performance of the coatings, an ex vivo circulation from the carotid artery to the jugular vein of the rabbit was constructed to evaluate the samples in flowing blood. As shown in Fig. 3D, the bare tube was covered by severe thrombus deposition after 1 h of circulation, while there was little thrombus deposition on the rhCOL III (PLA) and rhCOL III (NF) samples from macroscopical observation. Moreover, the inner side of the tubes was observed by SEM, indicating the adhesion of fibrous protein and the activation of platelets. In Fig. 3E, the bare tube was filled with severe thrombosis, which consisted of a large amount of activated platelets and cross-linked fibrin. The increased fibrin network density promoted the aggregation of platelets and erythrocytes, which gave rise to the formation and deposition of thrombosis [39]. Due to the immobilization of rhCOL III, fewer platelets adhered to the rhCOL III (PLA) groups, compared with bare PLA. There were few platelets adhered on the rhCOL III (NF) surface. The coating construction in the tube was different from that on the PLA substrate, and the compact nanoparticles constituted the surface, instead of NFs. This was due to the hydrolyzation of the flowing MTS ( Supplementary Fig.  S2). The results of the PRP incubation and the ex vivo perfusion experiment proved the rhCOL III (NF) coating had good antithrombotic properties.

Cell viability
Stent implantation destroys the endothelial layer of blood vessels, leading to inflammation, hyperplasia and thrombosis. Therefore, re-endothelialization on the inner surface of the stent is essential to reduce inflammation and promote blood compatibility after implantation. To evaluate the proliferation of ECs and SMCs on the rhCOL III (NF) coating, CCK-8 assay and TRITCphalloidin staining were performed to monitor the cell viability and morphology. As shown in Fig. 4A and B, the proliferation of HUVECs had no significant difference among PLA, rhCOL III (PLA) and rhCOL III (NF) samples after 24 h. However, the cell morphology on the PLA was different. The stretching of HUVECs on the rhCOL III (NF) coating increased cell adhesiveness and spreading areas, which was beneficial to multiplication. The number of adhered HUVECs was significantly higher on the rhCOL III (NF) substrate than on the others after 72 h. The rhCOL III (NF) coating vastly enhanced the proliferation of HUVECs compared to rhCOL III (PLA), and most cells were spindle shaped, which indicated healthy cell morphology.
In contrast, the rhCOL III (NF) coating showed essential inhibition of SMC proliferation (Fig. 4D). The viability of SMC on the rhCOL III (PLA) was as high as that on the bare PLA substrate, which exhibited that rhCOL III had no inhibition or promotion of the proliferation of SMCs. The resistance of SMC proliferation on the rhCOL III (NF) coating might be attributed to the NF layer. The ECM structure mimetic NF layer provided a 3D nanofibrous substrate and supported cell/material interaction. As a result, the fibrous substrate increased the expression of contractile markers [40] and then tended to suppress SMC proliferation.
These results indicate that the rhCOL III (NF) coating fulfills the requirements of the intravascular stent for endothelial healing and inhibitory SMC proliferation.

Inflammatory response of the coatings in vivo and in vitro
Inflammatory cells adhere to medical devices after implantation, leading to inflammatory responses around the material/tissue interface. This causes morphology differentiation and the secretion of inflammatory cytokines [41,42]. To evaluate the level of macrophage inflammation in vitro, RAW 264.7 mouse leukemia cells (RAW cells) were cultured on the PLA sheets, and the morphology and inflammatory cytokines were monitored. TNF-A is a proinflammatory cytokine produced by activated macrophages [43], and IL-10 is an anti-inflammatory cytokine secreted by regulatory macrophages [44]. The expression of TNF-A and IL-10 in the cell supernatant was detected by ELISA. According to the images of FDA-stained RAW cells, more spindle-shaped activated cells were observed on the bare PLA sheets, and the pseudopodia had elongated (Fig. 5E). The number of differentiated macrophages on the PLA substrate is higher than rhCOL III (PLA) and rhCOL III (NF). Comparatively, the rhCOL III (PLA) and rhCOL III (NF) groups inhibited the differentiation of RAW cells. The expression of TNF-A was higher in the supernatant of the PLA group, while the expression of IL-10 was lower than that of rhCOL III (PLA) and rhCOL III (NF). Moreover, there was no significant difference between the expression of rhCOL III groups, which proved that rhCOL III effectively suppressed inflammatory responses.
For further evaluation of the anti-inflammation capacity, the PLA sheets with coatings were subcutaneously implanted in Sprague-Dawley rats. The sheets attract inflammatory cells and result in the formation of a fibrous capsule around the material. The thickness of the fibrous capsule is an essential indication of the inflammation level, and severe inflammation would bring about a thicker fibrous coating [45,46]. Figure 5F shows the H&Estained tissue sections 15 and 30 days after subcutaneous implantation. The capsule fibrous thickness of the bare PLA, rhCOL III (PLA), and rhCOL III (NF) groups was 62.947 6 8.716 and 48.126 6 3.190 mm, 40.879 6 4.168 and 30.751 6 7.108 mm and 26.092 6 3.700 and 14.357 6 4.781 mm after 15 and 30 days, respectively. The encapsulation of rhCOL III (NF) was significantly thinner than rhCOL III (PLA) and bare PLA sheets (Fig. 5F). At these times points, the dual ECM-mimetic coating exhibited good antiinflammatory capacities and decreased the thickness of the fibrous capsule via increased immobilization of rhCOL III and fibrous structure.
Inflammatory cell infiltration can reveal the degree of inflammatory response. For further inflammatory evaluation, CD68 fluorescence staining was performed to observe macrophage distribution [47]. As shown in Fig. 5G, fewer fluorescent spots were observed in the newly formed fibrous capsules of rhCOL III (NF) sheets, compared with bare PLA and rhCOL III (PLA). This demonstrated much more inflammatory cells aggregated and infiltrated the capsules of PLA samples at 15 and 30 days. Moreover, there were few macrophages spotted on the rhCOL III (NF) substrates, which could indicate the material caused mild inflammation during implantation. As a result, the thickness and the macrophage distribution in the fibrous capsules showed a similar trend, in which dual ECM-mimetic coating resulted in lower inflammatory responses and better biocompatibility, compared with single ECM component mimetic coating.

In vivo stent implantation experiment
The PLA vascular stent is a promising absorbable bloodcontacting device. When the stent has been implanted for a long time, before complete degradation, chronic inflammation contributes to late stent restenosis and disordered endothelium layer [48]. To evaluate intimal hyperplasia and situ endothelialization in vivo, bare PLA and rhCOL III (NF)-coated stents were implanted in the rabbit abdominal aorta model and harvested after implantation for three months.
The cell morphology on the neointima of the inner surface of the lumen is shown in Fig. 6A. From the SEM result, there were few cells on the bare PLA stent after 3 months, and the growth of cells was disordered. The irregular state of the cells indicated different cell classes [26] and a low level of endothelialization. However, compared with the bare PLA stent, the number of cells increased, and the cells proliferated along the same direction on the rhCOL III (NF) coated stent. CD31 is a membrane receptor expressed by ECs to regulate the response to biomechanical stimuli [49] and was expressed around the EC edge. Besides, high CD31 expression indicated healthy and functioning ECs [13]. ECs secrete endothelial nitric oxide synthase (eNOS) to circulate nitric oxide metabolites and regulate blood pressure [50,51]. The eNOS expression demonstrates the physiology of blood vessels and endothelialization degree [52]. For further assessment of the cell type, morphology, and distribution of the neointima, the inner surface of the lumen was stained by immunofluorescence (anti-CD31 and anti-eNOS) and observed by LSCM in Fig. 6B. The immunofluorescent image is in agreement with the SEM results. From the quantitative analysis in Fig. 6D, the CD31 expression of the inner cell layer of rhCOL III (NF) was higher than that of the bare PLA stent after 3 months. A similar trend was also observed in the eNOS expression, compared with CD31 expression in the tissue. The high expression of eNOS in the rhCOL III (NF) group indicated that the regenerated endothelium was healthy and intact. As shown in Fig. 6B, ECs did not cover the inner surface of the bare PLA stent completely, while the rhCOL III (NF) surface was covered by ECs arranged closely and in the same direction. Hence, the above result demonstrates the rhCOL III (NF) coating enhanced EC proliferation and accelerated the recovery of the endothelium layer.
For further evaluation, the stent cross-section with surrounding tissue was stained by H&E and immunohistochemistry antibodies (anti-CD68 and anti-A-SMA) to investigate the restenosis rate, inflammation and endothelialization. Inflammatory responses could not be avoided after implantation and endothelial dysfunction is linked to inflammatory responses [53,54]. The propped stent injured the endothelium cells of the blocked blood The analysis of CD68 fluorescence signals of the capsule. *P < 0.05, **P < 0.01 and ***P < 0.001, ****P < 0.0001 (oneway ANOVA, error bars represent the SD). vessels and FBR was maintained until the complete degradation of the stent. Serious inflammatory responses lead to excessive thickening of vascular intima and restenosis. In Fig. 6E, numerous CD68-positive macrophages aggregated around the bare PLA stent struts after 3 months. Moreover, there were few macrophages in the neointima, and the activated macrophages mainly aggregated along the bare PLA struts, which indicated acute inflammation occurred and was sustained prior to collection from the New Zealand rabbits. In contrast, the rhCOL III (NF) coating suppressed inflammation remarkably after implantation, and few CD68-positive macrophages were observed around the coated struts, which suggests that the rhCOL III (NF) stent caused mild inflammatory responses.
Neointimal hyperplasia is a considerable threat to stent failure. At present, excessive proliferation of SMCs is considered the major contributor to the abnormal thickening of intimal and ISR [55,56]. SMCs in a contractile phenotype have a low proliferative rate, but after vascular injury, SMCs turn to a synthetic phenotype and undergo a proliferative response [57][58][59]. The switch of the SMC phenotype from contractile to proliferative leads to the deterioration of blood vessels and excessive hyperplasia. The A-SMA was specifically expressed in SMCs with contractile phenotype [26]. The expression of A-SMA around the rhCOL III (NF) stent was much higher than that of the bare PLA stent (Fig. 6H), indicating that most SMCs had low proliferative capacity and the neointima of rhCOL III (NF) groups was more mature. The maintenance of the contractile phenotype might arise from mild inflammatory responses, and a low level of inflammation would not stimulate SMCs and could induce the contractile phenotype of SMC. As a result, decreasing the acute proliferation of SMCs contributed thicker neointima and less lumen loss.
As shown in Fig. 6E, the restenosis rate of the bare PLA stent was as high as 20.68 6 1.05%, and that of the rhCOL III (NF) coating was about 16.10 6 0.64% after 3 months of implantation. The rhCOL III (NF) coating inhibited intimal hyperplasia and reduced restenosis by suppressing the aggregation of inflammatory cells and excessive proliferation of SMCs. Inflammation and coagulation were closely linked, due to the damage of EC function after implantation, and then persistent inflammation gave rise to abnormal hyperplasia and delayed re-endothelialization. The fibrous ad-layer and rhCOL III mimicked ECM dual structure and cross-section of the vascular structure, including stents. *P < 0.05, **P < 0.01 and ***P < 0.001, ****P < 0.0001 (t-test, error bars are defined as SD). component to enable the injured blood vessel in a low level of inflammation, and endothelialization accelerated without excessive proliferation of SMCs.
From the in vivo results, rhCOL III (NF) is an ideal vascular stent coating, benefiting from the reduction of restenosis on account of a mild inflammation response, contractile SMC induction, and endothelialization acceleration.

Conclusion
In this work, an ECM-mimetic coating with dual bionics was prepared successfully on PLA stents, where the biomimetics were driven by structure mimicry and component/function mimicry. The NFs on the PLA constructed an ECM fibrous structure as well as a three-dimensional reservoir to support the amplified immobilization of rhCol III. Then, rhCol III, a multi-functional collagen, provided reliable anti-coagulation and anti-inflammatory properties and promoted EC proliferation. As a result, this ECM-mimetic rhCOL III (NF) coating possessed superior anti-thrombosis and anti-inflammatory capacities, promoted re-endothelialization and inhibited SMC proliferation by inducing contractile SMC phenotype in vivo. In view of the stable interface and multiple functions endowed by the fibrous ad-layer and tailored rhCOL III, the rhCOL III (NF) coating is a promising material for the efficient modification of cardiovascular stents.

Supplementary data
Supplementary data are available at Regenerative Biomaterials online.