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.