Cardiovascular medicine currently relies heavily on biomaterial implants or other blood-contacting apparatuses such as vascular grafts, stents, pacemakers, guidewires, artificial kidney, extracorporeal oxygenation, heart valves, and ventricular assist devices. A concerning complication of these devices relates to the formation of surface-induced thrombi. The thrombi pause the risk of occlusions or downstream emboli resulting in debilitating strokes, pulmonary thrombosis, organ infarction, complete circulatory collapse, and even death. To this day, anticoagulants are the only clinically available response to implants thrombosis, but they are not without risks, such as bleeding, particularly leading to intracerebral/extracerebral hemorrhage. Many other approaches to minimize clotting have been investigated, ranging from mechanical alteration, biochemical modifications, endothelial cell (EC) reseeding, and stem cells to build a nonthrombogenic surface.
Since Dr. Gott’s original publication more than 5 decades ago on the preparation of nonthrombogenic graphite surface with heparin, considerable attention has been given to the use of artificial materials and nanotechnology to modify blood-contacting surfaces. Clotting is a multistep process that involves primary hemostasis (proteins and platelets adhesion, platelet activation and degranulation, and platelet aggregation) and secondary hemostasis with the action of thrombin on a fibrin plug. The production of “thromboresistant” surfaces targets each of those steps above. A quick synopsis of the processes that govern thrombosis from blood–material interaction will be discussed below as well as the surface modifications to counteract them. Ideally, the nonthrombogenic coating should be biocompatible, non-biodegradable, antiproliferative, and durable. Although these criteria are straightforward, they represent one of the biggest challenges in modern medicine.
In preparation for this review, a systematic and comprehensive search of peer-reviewed publications was performed using keywords including hemocompatibility, surface-induced thrombosis, inhibition of proteins adsorption, and surface modification. Each type of thromboresistant material was entered in the Pubmed and Embase search engines. More than 30 publications spanning several decades were examined on each of the chemical and biological modifications for blood-contacting surfaces described herein.
Mechanism of Surface-Induced Thrombosis
Clot formation on blood-contacting surfaces is complex and mediated by primary hemostasis and the intrinsic pathway of secondary hemostasis. Primary hemostasis is initiated by the adsorption of coagulation proteins such as fibrinogen (mostly), fibronectin, vitronectin, and von Willebrand factors (vWF) as seen in Figure 1. These plasma proteins interact with their milieu including the artificial surfaces through hydrogen bonding, van der Waals interactions, π–π stacking, and electrostatic interactions to form layers of about 100–200 Å thick.1 This layer forms within seconds of blood contact and is evolutive in nature; there is considerable evidence that overtime, different polypeptides displace other proteins already anchored on the surface, a phenomenon first observed by Vroman et al.2,3 and described as the “Vroman effect.” After a rapid initial adsorption, fibrinogen is displaced from the surface within minutes by other plasma proteins by unknown mechanisms.2–42–42–4 Proteins with higher affinity for the foreign surface but of low concentration in plasma such as kininogen may exhibit a delayed adsorption when compared with fibrinogen and cause this Vroman effect.
Platelets later bind to adsorbed proteins via platelets surface receptors such as glycoprotein GPIb (CD42) and GPIIb/IIIa (αIIbβ3 integrin). In particular, GPIb adheres to adsorbed vWF, whereas GPIIb/IIIa adheres to adsorbed fibrinogen, fibronectin, vitronectin, and vWF.5,6 The most abundant plasma protein albumin also adsorbs onto material surfaces; however, it is generally thought of as an anticoagulant. The reason for this is because platelets are less adhesive to albumin, as it lacks known amino acid sequences that bind platelets, namely the Arg-Gly-Asp (RDG) sequence found on fibrinogen, vWF, fibronectin, and vitronectin.7–97–97–9 Nevertheless, platelets also bind to little albumin via nonspecific interactions. Ultimately, physical and chemical properties of blood-contacting surfaces determine the amount of proteins adsorbed at the blood–polymer interface.
Upon adhesion, platelets become activated, undergo a change from a discoid to a dendritic shape, and release α and dense granules. Dense granules containing adenosine diphosphate, serotonin, platelet factor 4, and calcium further recruit circulating platelets to expand the aggregate, which is stabilized by fibrinogen. The phospholipid bilayer of platelets also catalyzes the activation of prothrombin to thrombin, which exerts a positive feedback on platelets activation and converts fibrinogen to fibrin, whereas calcium and phospholipids also participate in the activation of coagulation factors. Furthermore, blood contact with materials induces the expression of tissue factor (TF) by monocytes and thus activates the extrinsic pathway, a phenomenon that has been observed in vivo with the use of cardiopulmonary bypass.10,11 Although there is evidence of activation of the extrinsic pathway, the intrinsic pathway exerts the dominant role in surface-mediated coagulation with interplay between factor XI, factor XII, prekallikrein, and high molecular weight kininogen (HMWK) as illustrated in Figure 2.
Passivation of Surfaces
It includes the long list of processes whereby layers of proteins-repellent molecules are added onto blood-contacting surfaces. Table 1 provides a description of the chemical make-up of some well-studied passive coatings as well as a summary of their applications with supporting and countering evidences (Figure 3).
Pyrolytic carbon thromboresistant properties were discovered somewhat serendipitously in the laboratory of Dr. Gott after its development as a coating for nuclear fuel rods in the 1960s. In 1965, Bokros et al.43 developed pyrolytic carbon, a durable material and with Dr. Gott investigated the possibility of binding heparin to these durable carbons. Although unsuccessful, the pyrolytic carbon proved to be the least thrombogenic materials tested at the time and the best rigid material for prosthetic valve reconstruction.43,44 By 1970s, many prosthetics valves were made of pyrolytic carbons from Bokros et al.43 original invention. Another carbon-based coating that applies the same mechanism for hemocompatibility (i.e., albumin adhesion) is the diamond-like carbon (DLC) coating. Initial experiments revealed that mouse fibroblasts and peritoneal macrophages culture on plates coated with DLC showed no cellular damages compared with control plates. This preliminary result demonstrated that they were biocompatible and could be used satisfactorily as coating for biomedical implants, namely orthopedic implants to serve as a protective film to prevent the release of metal ions and wear debris into tissues.45
Although antithrombotic moieties can be added to blood-contacting surfaces, the clinical application of this approach is limited by erosion or desorption. A novel undertaking consists of mimicking the biologically inert cell surface. The thrombogenicity of blood cellular elements is a feature of their intracellular content, whereas the extracellular surface is an antithrombogenic electrically neutral membrane made of phosphorylcholine that prevents fouling.46,47 Both phosphorylcholine head groups and other zwitterionic materials have been exploited in the development of antifouling surfaces. These zwitterionic molecules bind water from their surrounding milieu to form a hydration layer that is bioadhesion resistant.29–3129–3129–31 Their molecular design incorporates a quaternary ammonium cation and an anion (phosphate, sulfonate, or carboxylate) to make up phosphobetaine, sulfobetaine, or carboxylbetaine, respectively (Figure 4). These groups have been grafted onto modified polymer surfaces and resisted protein adsorption and platelet adhesion. Stents coated with phosphorylcholine such as Biodiv Ysio exhibit enhanced hemocompatibility.48,49
Biological macromolecules such as albumin also play an important role in clotting: it is the most abundant peptide in plasma, the first protein that adsorbs onto implants, and it lacks the platelet-recognized RDG sequence. Furthermore, it was observed that patients suffering from nephrotic syndrome present with hypoalbuminemia and an increased tendency to thrombosis. The degree of hypoalbuminemia correlates with the extent of thrombosis, and hyperaggregable platelets have been found.50,51 Additionally, the anticoagulant action of heparin on antithrombin III (ATIII) is because of its negatively charged sulfate groups binding to the positively charged epsilon-amino lysyl group of ATIII. Because albumin also contains negatively charged group, it was hypothesized to simulate the same effect on ATIII.33 These observations led to thromboresistance testing on albumin-based coats. Although other proteins can displace albumin on polymer surfaces, covalent immobilization strengthens the binding of albumin.
Surface Topography and Energy Toward Passivation
Although extensive human clinical data are still missing on many of the above modifications, newer concepts have surfaced to target proteins adhesion, namely by using surface topography and chemistry to imitate nature. Bioinspired coatings, the newest and promising development, have recently gathered much interest. One method known as slippery, liquid-infused, porous surface (SLIPS) mimics the surface of the Nepenthes pitcher plant where a film of slippery aqueous liquid prevents attachment of insects by repelling the oil on their legs. Insects are thus able to slide down to the bottom to reach the plant juices.52 This technology combines a slippery liquid trapped within the microstructures of a porous substrate biomaterial.53 The porosity of the surface is merely the glue that holds the lubricant in place. The idea that surface topography could affect wettability first appeared in the literature in 1996,54 but their blood-compatibility potential is just now being appreciated.
Surface energy is an equally important aspect of these biomimetic surfaces. It determines how a surface interacts with its milieu. Water-repellent surfaces are able to do so because the dry surface has a lower surface energy than molecules of water. To disturb the chemical bonds that maintain the spherical shape of a water droplet, energy has to be applied. Thus on surface with lower energy, water retains its droplet form and does not wet.54 The lubricant in the SLIPS technology has an omniphobic character, repelling oils and water, a remarkable achievement in light of the very low surface energy of oil molecules and their nonpolar nature.55 Different models exit to explain how surface energy determines the wetting capability of a surface. In the Cassie–Baxter model, which describes a solid–air–liquid interface, a liquid droplet float over air pockets formed in the papillae of the surface minimizing contact as seen in Figure 5. This model is most relevant to these bioinspired surfaces.
Wong et al.53 recreated a SLIPS that withstands mechanical damage and exhibit high water repellency even under high pressure (approximately 676 atm). Under high-speed camera, they also observe self-healing characteristics where the fluid moved to damaged areas in 150 msec and recovered their liquid-repellent function. Their self-healing property means that the lifespan of coated implants may be prolonged by countering wear and tear from the deterioration of the material. SLIPS also effectively repel complex fluid such as blood.53
This crude observation was further characterized by Leslie et al.56in vitro by comparing fluorescently labeled fibrin deposed on SLIPS-coated surface that was exposed to whole human blood with low dose heparin (0.25 U/ml). They observed a 27 fold reduction in platelet adhesion compared with the control surface. Flow studies were also performed with blood circulating at shear strain rate of 1000 per second for 16 hours at the most with consistent low sliding angle (<3°).56 The potential for this technology still remains untapped, but it is showing great promise for surface modifications. The “long-term” evidence should be further explored beyond the sliding angle, looking at actual platelet activity, cytotoxicity, cellular damages/hemolysis, and inflammatory markers. The strongest evidence in favor of this technology came from in vivo experiment where a modified perfusion tubing assembled into a porcine femoral arteriovenous shunt for a duration of 8 hours remained free of thrombus even in the absence of any heparin compared with expectedly occluded conduits in the control group.56
Activation of Surfaces
They involve surface changes that inhibit processes or agents of coagulation. The oldest attempt is heparinized surfaces and dates back to the 1960s when -Gott et al. demonstrated that clotting time could be prolonged by attaching heparin to graphite-coated surfaces via electrostatic interactions. Along with heparin, some other well-known modifications are succinctly presented in Table 2 (Figure 6).
In light of the success of heparin, other thrombin-inhibitors such as hirudin have been tested on a variety of substrates with optimistic results. Phaneuf et al.68 have functionalized Dacron stents with recombinant hirudin (r-hirudin) using a cross-linked modified albumin base. Li and Henry69 used a polyethylene glycol (PEG) spacer arm to covalently attach the C-terminus of r-hirudin to a biodegradable microfibrous vascular graft. Although less thoroughly investigated than hirudin, immobilization of thrombomodulin onto surfaces that interface with blood has had some merit. The outcome in preventing graft thrombosis proved successful in thrombomodulin covalently immobilized on films via a spacer PEG arm. Decreased complement activation was also reported.70In vivo studies have complemented these findings with recombinant thrombomodulin (rTM) coated onto the surface of expanded polytetrafluoroethylene stents implanted in porcine model showing decreased neointimal hyperplasia.71
Activation of the clotting cascade by inflammation is a less well-characterized and less targeted aspect of surface-induced thrombosis but recently, Yoshitomi et al.72 provided some evidence that free-radicals scavengers can reduce inflammation and thrombosis. Their work stemmed from the observation that nitric oxide-releasing polymer generated reactive oxygen species (ROS) that amplify inflammatory cytokines, the extent of which contributed to blood coagulation. A polymer containing nitroxide radicals, which act as radical scavengers, was synthesized and coated onto glass beads. Control beads were coated with poly(chloromethylstyrene) (PCMS), which generates ROS. When exposed to whole blood, they showed a statistically significant decrease in levels of ROS, tumor necrosis factors-α, and interleukin-1β on the nitroxide-coated beads. They also measured a reduced consumption of leukocytes and platelets, and scanning electron microscopy imaging showed persistently smooth nitroxide-coated beads surface as opposed to aggregates visible on control beads.72
These nitroxide radicals are functional groups containing an N–O bond attached to a side chain such as R3N–O. They are free radicals themselves, but they are stable because of resonance structure involving the unpaired electron and steric hindrance from side chains.73 The nitroxide radicals-containing polymer (NRP) have yet to perform in long-term studies but provided compelling evidence that targeting inflammatory processes can be a source of blood-compatible biomaterials.
The endothelium is a natural antithrombogenic barrier, and the issue of clotting with blood-contacting medical devices is inherent to the absence of this layer. Synthetic grafts made of polytetrafluoroethylene (PTFE) are the most widely used grafts. Although successful in large-caliber vessels (>5 mm) with high flow rate, they have proven disappointing in small diameter because of rapid thrombosis and fibrinous thickening that leads to occlusion of the vessels.74 Reduced blood flow in small-sized vessels also increases the contact time between the blood and the luminal surfaces of the graft, increasing the likelihood of a thrombogenic reaction. Although tissue-engineered vascular grafts were developed to overcome the limitations of small diameter prostheses, they too are susceptible to thromboembolic complications without (re)-endothelialization with ECs or stem cells.
The concept of EC seeding first appeared in the literature in 1978 from Herring et al.75 and refined with in vitro studies in the early 1980s. Three decades later, EC transplantation by seeding (cells mixed with blood or plasma and added to graft) or sodding (cells placed directly onto graft) of autologous EC onto the small-caliber grafts has improved the patency rate to level comparable with that of autologous vein grafts.76 A 6-year multicenter study of Veith et al.77 reported saphenous vein grafts patency of 75% at 3 years, findings that were upheld by Taylor et al.78 A clinical trial led by Meinhart et al.76 demonstrated that endothelialized ePTFE grafts showed a 3-year patency rate of 72.9%.
The initial seeding protocol involved scraping EC from the lumen of saphenous veins with steels wool pledges and mixing them with blood used to preclot the grafts.75 The presence of an EC layer was confirmed by the identification of EC-specific Wiebel-Palade bodies and by staining with fluorescent-labeled von Willebrand antibody.79 Because of their immunogenicity, ECs are harvested from autologous sources, namely progenitors in the blood or bone marrow, or from adipose tissue. Upon seeding, EC adhesion is enhanced in vitro through exposure to physiological flow rates and shear stress. The protocols have become more sophisticated but still require maturation of the seeded layer before implantation, which makes this technology impractical for emergency surgeries. In situ endothelialization was considered to circumvent this requirement by allowing transanastomotic EC migration, but it does not provide complete graft coverage.80 Other issues to consider with EC transplantation include methods to promote stronger and more specific EC adhesions. For that purpose, studies have investigated precoating grafts with RGD peptide, a fibronectin-derived peptide recognized by both EC and platelets, which could thus potentially lead to thrombus formation.81 More selective EC-adhesion proteins exist such as laminin-derived YIGSR.82
It is a technique of consolidating powder compacts with the use of thermal energy. It involves a heat treatment applied to a powder coating to about one half or more of the alloy’s melting temperature. Upon heat exposure, the powder compact becomes densified, and the grain size increases, resulting in reduction of the total interfacial energy. This is an old technique that dates to the prehistoric era with pottery and continues to be used to date to fabricate bulk ceramic components and powder metallurgical parts. There are two main types of sintering: solid state and liquid phase.83 It creates a textured surface that entraps blood cellular components, forming a smooth biological lining called pseudointima.84,85
Early left ventricular assist device (LVAD) were made of smooth blood-contacting surfaces such as polyurethane elastomers, which were very thrombogenic because of surface irregularities that created flow-eddies and promoted clotting.86,87 With the development of textured surfaces, thromboembolic events decreased. The research done on the HeartMate 1000 IP (Thermo Cardiosystems, Inc, Woburn, MA), whose blood-contacting surface consists of sintered titanium microspheres, has highlighted the role of sintering in preventing thrombus formation. The reduced thrombogenicity resulted from the formation of a neointima from hematopoietic cells.84,88 Electron microscopy and molecular studies determined that the neointimal layer was made up of spindle-like cells resembling fibroblasts or myofibroblasts and hematopoietic cells namely CD45+ and CD34+ cells that proliferate on the protein background and differentiate within this milieu.89 One hypothesis that was proposed to explain the formation of neointima on the sintered surface is the presence of a three-dimensional surface topography. Zapanta et al.90 casted cavities into a polyurethane polymer applied to the surface of a blood pump before implantation and noted the formation of a stable neointima. Additionally, adherence to the polymer correlated with the size of the cavity: the larger the cavities, the stronger the neointimal adherence.90
Patients implanted with the HeartMate 1000 IP at 11 clinical centers in the United States were evaluated for thromboembolic complication at 62 ± 76 days in the setting of minimal systemic anticoagulation. Rose et al.84 found these patients to be at low risk for thromboembolic complications owing to the textured surface of their LVADs. Nevertheless, the use of sintering has been plagued with some controversial results. Previous designs of the HeartMate II axial flow pumps with sintered titanium exhibited severe thrombosis in areas of the pumps with small clearances. The pumps were redesigned with smooth titanium stators and sintered conduits with improved blood compatibility.91–9391–9391–93
With the emergence of molecular biology and the development of scanning probe microscopy, nanotechnology has become a key player in material sciences and surface chemistry with the next generation of coatings under investigation that incorporates nanotechnology. Its role in material sciences is rapidly expanding because of the small scale of its features that guarantees a more accurate interaction between materials and environment: cells and proteins truly interact with particles of similar size; there is a larger fraction of surface atoms, and flaws and imperfection are scaled down. Because of their large specific surface areas, their interactions with the environment are also faster.94
A successful application of this technology was originally developed by the U. S. Air Force for aerospace engineering and known as polyhedral oligomeric silsesquioxane (POSS). Its attractive physical and mechanical properties were augmented when combined with a poly(carbonate-urea) urethane (PCU) polymer. POSS-PCU has been functionalized by covalent linkage to endothelial progenitor cell (EPC)-specific antibody (anti-CD34) that attracts EPC leading to endothelialization, which prevents thrombosis (Figure 7).95 There is some evidence that it is a biocompatible material because it was used as a scaffold for a tracheobronchial transplant in a 36 year old male patient with tracheal cancer.96 Further data on thrombogenicity come from in vitro studies where polyvinyl chloride (PVC) spray coated with POSS-PCU then functionalized with anti-CD34 that showed hemocompatibility.95
The relative passive nature of many carbon-based materials such as pyrolytic carbons and DLC and their success as blood-contacting surfaces suggested an application for carbon nanotubes (CNTs) in blood–material interface. CNTs are single-walled or multiwalled cylinders of graphene.97 Their inert surface has been shown to be antiadhesive when exposed to platelets. Koh et al.98 prepared composite poly(lactic-co-glycolic-acid)–multiwalled CNT (PLGA-MWCNT) films onto which fibrinogen was immobilized before incubation with platelet-rich plasma. When compared with the pristine PLGA films, these CNT-functionalized films exhibited reduced platelets adhesion with no activation.98 More recently, Gaffney et al.99 found conflicting results when they exposed MWCNT-modified PVC to platelet-rich plasma and measured more platelet aggregation compared with bare PVC surfaces. In vivo, PVC perfusion tubing functionalized with MWNCTs exhibited massive thrombosis in the absence of anticoagulant.99 Interestingly, it has been shown that CNTs induce extracellular Ca2+ influx and Ca2+ release from intracellular stores.100,101 Reducing intraplatelets Ca2+ with prostacyclin analogs reduced MWCNT-induced platelet aggregation. It is likely, as Gaffney et al.99 pointed that the low platelets adhesion from Koh et al.98 is an indirect effect of plasma protein adsorption such as seen with the use of albumin.
The above references demonstrate a keen interest in developing biocompatible hemostable surfaces for blood-contacting implants. Although the scientific literature continues to produce new nonthrombogenic materials, very few have successfully transitioned from laboratory to clinical practice. Many options are offered but fewer in vivo studies are done.
Many experiments have demonstrated reduced platelet adhesion and proteins adsorption, but no surface has been completely thromboresistant. Although investigating newer surfaces is scientifically relevant, unless we can completely eliminate the need for anticoagulation or antiplatelets, these data are not clinically relevant yet. We must agree that the end goal for hemocompatibility research is to find an alternative to the EC layer or something very close to it by well-established parameters (i.e., clear cutoffs for platelets activation, protein deposition, and how these translate into the requirements for anticoagulation).
The biggest success has been limited to endothelialization of vascular grafts. The challenge with other antithrombogenic material is to make them as similar as possible to the natural endothelial lining of blood vessels, a daunting task in light of the fact that the endothelial layer has secretory functions (prostacyclin), recycling potential, and active anticoagulant moietes (heparan sulfate, glycocalyx, etc.) and antiplatelet activity. Many artificial surfaces replicate an isolated function of the endothelium where nonthrombogenicity is a pluridimensional effect. Future studies could look at combining the individual actions of these surfaces into one.
Moreover, the biology of surface-induced coagulation is complex involving several players as seen in Figure 8, and there is much that remains to be elucidated such as the role of white blood cells in thrombogenesis. Receptors and surface proteins on platelets are numerous with convoluted roles. Studies on shelf life and an assessment of the ageing of these coatings have yet to be conducted. Other considerations are often excluded while evaluating these hemocompatible materials. It is well known that primary hemostasis is more active in the high shear arterial flow, whereas secondary hemostasis is more at play in the low shear venous flow.102 There is little indication that current studies have made this distinction toward the development of nonthrombogenic surfaces.
The road toward the development of the ideal blood-contacting coating is uncertain because the defining criteria for such coat have not been agreed upon, and there exist no common standards for in vitro testing of biomaterials. This point was first presented by Ratner103,104 in 1993 as the “biocompatibility catastrophe” and revisited in 2007 with little progress made in defining hemocompatibility. Recently, Braune et al.105 offered a perspective on standards for in vitro blood-compatibility testing. Currently, the minimum required testing is outlined in the ISO 10993–4 catalogue; the five test categories indicated from the ISO standards for blood-compatibility evaluation are thrombosis, coagulation, platelets activity, hematology, and immunology, but as argued by Braune et al.,105 supplemental tests ought to be added. Extensive investigation of the intrinsic physical properties of biomaterials including porosity, roughness, charge density, elasticity, and wettability should be conducted owing to their procoagulant contribution. Standardized guidelines for the use of anticoagulant for in vitro testing are also lacking. Such guidelines will establish recommendations for the choice and concentration of anticoagulant among the formulations of heparin (low molecular weight heparin or unfractionated heparin) and sodium citrate to ensure the functionality of blood components for testing. Moreover, in vitro miniaturized testing systems, whether static or dynamic, must be designed ideally to incorporate physiological shear stress, flow rate, temperature, and hematocrit and adapted for arterial or venous circulation.105 Lastly, pyrogenicity testing should be reinforced given that pyrogens such as lipopolysaccharides and lipoteichoic acids activate hemostasis and can be easily introduced into testing systems during manufacturing.106,107
This article illustrates the long and tremendous history in developing hemocompatible materials. Given the successes accomplished in medicine because of the progress made, it is imperative that research continues to challenge paradigms. Only with relentless efforts will a truly hemocompatible material evolves out of some of the formula herein.
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