Figure 3. Schematic representation of intrinsic blood-coagulation mechanism. Upon the implantation of biomaterials, proteins from plasma and ECM will be adsorbed on the surface rapidly resulting in a thin protein film deposition and subsequent restructure of the blood-material interface[26]. Plasma protein in high concentration will first be adsorbed onto the surface such as fibrinogen (FGN) and fibronectin (FN). Those proteins will eventually be replaced by trace proteins with high affinity such as high molecular weight kininogen (HK) and clotting Factor XII (FXII), known as “Vroman effect”[27]. The complement system is also triggered resulting in immune response to the biomaterial. Once adsorbed on the surface, FXII undergoes conformational change and is activated to FXIIa. Activated FXIIa will sequentially activate other clotting factors: FXI and FIX. FIXa complexed with cofactor FVIIIa will further activate FX to FXa leading to thrombin generation and conversion of fibrinogen to insoluble fibrin. Platelets will be activated and adhere to the surface through protein adsorption. Activated platelets release other clotting factors to facilitate the thrombin formation and further platelets activation and aggregation. Thrombin further promotes the polymerization of fibrin. Together with platelets aggregation, an insoluble thrombus is formed.
Designing surfaces of blood-contacting biomaterials should consider the protein adsorption, thrombin generation, platelet adhesion and cellular behavior at the interface especially ECs and SMCs to improve grafts patency and thrombogenicity reduction. The intact vascular endothelium is responsible for anticoagulant properties and vascular protective functions. Endothelium contains prostacyclin and nitric oxide (NO) that exhibit signal-inhibit effects to inhibit platelet aggregation and activation as well as the proliferation SMCs[27]. Overgrowth of SMCs is an early stage of intimal hyperplasia formation. Inspired by the thromboresistive nature of the vascular endothelium, achieving fully endothelialization on the luminal surface either throughin vitro or in situ approaches has been highlighted as the ultimate solution[26].
Topography on the micron- and nanometer scale of the surface plays a crucial role in antithrombogenicity. For example, picosecond laser ablation technology was adopted to micropattern PEG-functionalized PLA vascular grafts with parallel microgrooves with varying geometries[28]. It was found that all microstructured surfaces were non-toxic and non-hemolytic. A specific feature with 20 to 25 μm wide and 6 to 7 μm deep favored the adhesion of EC. The hydrophobicity of patterned surface was significantly increased with the water contact angel changed from 71.1 ± 0.2° to 112 ± 1° after laser ablation. Since PEG element was homogenously incorporated in the substrate, the topographic change would contribute to the increased hydrophobicity instead of the removal of PEG from the top layer. However, higher platelet adhesion on patterned surface may be attributed to the increased surface roughness due to the presence of nanopores after micropatterning. It was believed by the authors that underin vivo conditions, the platelet adhesion on microstructured surface would be mitigated due to the micro shear gradient produced by hemodynamics around the patterns.
It was suggested that the heparin-like molecule (heparan sulfate) residing on vascular endothelium plays a key role in thromboresistance[29]. Heparinization of biomaterials has been widely used in clinical practice to improve hemocompatibility. Various heparinized blood-contacting devices are currently in market[30] such as Palindrome™ Precision H-heparin coated dialysis catheter (Medtronic) and Affinity Pixie™ Arterial Filter (Medtronic).
A biodegradable PLA vascular stent was fabricated by 3D printing and heparinized through PDA/PEI intermediates to improve hemocompatibility and anticoagulation property[31]. The surface of PLA is lack of functional groups that limits its heparinization potential. Mussel-inspired natural PDA can bind to substrates under mild aqueous conditions instead of organic solvents. Since the amine groups provided by PDA are insufficient, amine-rich PEI was introduced onto the surface to effectively conjugate with heparin. Heparinization resulted in significant increase in stent flexibility as evaluated by a three-point bending test (1.00 ± 0.11 N of heparin-coated stents v.s. 1.39 ± 0.24 N of bare PLA stents). It was confirmed that those heparinized stents suppressed SMCs proliferation while promoted ECs proliferation. The in vitro adhesion tests showed that fewer fibrinogen and platelets attached to heparinized stents compared to PDA/PEI coated ones, which reveals their anti-thrombogenic properties. When implanted those stents in porcine models, the heparinized stents showed the most promising lumen patency with inhibited neointima hyperplasia and lowest area restenosis.
Traditional drug-eluting stents rely on the incorporation of cytotoxic or cytostatic drugs such as paclitaxel for inhibiting the migration and proliferation of SMCs[32]. However, those drug-eluting stents are always associated with delayed re-endothelialization due to the suppressed growth of ECs. Co-immobilization of two or more biomolecules into the vascular grafts is developed to obtain complementary or synergistic functions in SMCs suppression while ECs promotion. However, bioactive molecules with relatively distinct therapeutic effects will impair the combined efficacy due to the absent interactions between those molecules, which further hinders their practical use[33].
An endothelium mimicking coating was developed through the sequential conjugation of heparin and nitride oxide (NO)-releasing substance on 316L stainless steel stents[34]. There are other studies that investigated the combined effects of NO and heparin on healing outcomes of vascular grafts based on employment of NO donors[35]. Whereas the safe therapeutic dose of NO remains uncertain and the half-lives of those NO donors are unsatisfactory, limiting their applications in long-term devices[34]. The NO-releasing compound used in this study was selenocystamine (SeCA) to realize in situcatalytic generation of NO. The bioactivity of both biomolecules was retained and not affected by each other. The heparin/SeCA treated stents combined the anticoagulant function by heparin and anti-platelet adhesion by NO-releasing. The migration and growth of SMCs were effectively suppressed, whereas the growth of ECs was promoted. When implanted the heparin/SeCA coated stents in iliac arteries of rabbits, the enhanced re-endothelilization and suppressed restenosis were achieved.
Nitinol (NiTi), known as shape memory alloys, is widely designed for self-expanding vascular stents to prevent the possible plastic deformation in vessels due to the balloon expandable stents. However, excessive nickel ion releasing from the nitinol can lead to cellular inflammation[36]. A nanocomposite coating composed of TiO2 nanotubes and chitosan-heparin particles, was developed to obtain improved hemocompatibility as well as enhanced corrosion resistance. The TiO2 nanotubes were deposited on NiTi alloy by electrochemical anodization followed by chitosan-heparin NPs coating via an intermediate dip-coated PEI layer. Those nanoparticles can act as drug carriers for sustained release of heparin. There was a continuous release of heparin for 2 weeks after the initial release. It was reported that the anodization of highly ordered nanotubular structure to nitinol surface would improve its corrosion resistance and reduce nickel ions releasing[37]. The TiO2 nanotubes layer effectively reduced the release of nickel ions, while the nanoparticles coating also inhibited those ions releasing. Compared to bare metallic and anodized stents, the chitosan-heparin incorporation resulted in significant reduced hemolysis ratio and platelet adhesion as well as enhanced the attachment, spreading and proliferation of ECs. Whereas the effects of this nanocomposite coated nitinol stents on SMCs were not determined.
Many researchers aimed to promote the adhesion and growth of ECs, yet obtained limited re-endothelialization. Possible reason can be the ignorance of the competitive growth between ECs and SMCs[34,38]. Therefore, the efficient surface engineering techniques on ECs proliferation are suggested to perform a co-culture assay of ECs and SMCs. For the aforementioned heparin/SeCA treated stents, they exhibited a synergistic effect on ECs over SMCs. ECM peptides can be incorporated to vascular grafts to influence cellular behavior and output specific interactions to surrounding. Several biomolecules incorporated in vascular grafts such as RGD peptides are not cell-specific, that raises the concern about competition between ECs and SMCs. The REDV polypeptide is specifically recognized by ECs making it an ECs-specific biomolecule. Xue et al. covalently immobilized REDV peptides on nitinol reinforced PET microfibrous grafts through PDA NPs[38]. Such surface modification on microfilaments produced hierarchical micro/nanostructures that benefit cell attachment and proliferation. REDV immobilization grafts improved the hemocompatibility with untraceable hemolysis rate as well as ECs proliferation and increased release of NO. Besides single peptide, Peng et al. studied the effects of multiple-peptides (YIGSR, RGD, and REDV) immobilization of silk fibroin scaffolds on ECs[39]. YIGSR-modified scaffolds showed the highest cell migration rate compared to RGD- and REDV-modified scaffolds. Whereas dual-peptides (YIGSR+RGD) significantly enhanced the proliferation of ECs compared to other dual-peptides combination.