The blood compatibility challenge. Part 4: Surface modification for hemocompatible materials: Passive and active approaches to guide blood-material interactions

Acta Biomaterialia 94 (2019) 33–43

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Review article

The blood compatibility challenge. Part 4: Surface modification for hemocompatible materials: Passive and active approaches to guide blood-material interactions Manfred F. Maitz a,g,⇑, M. Cristina L. Martins b,c,d, Niels Grabow e, Claudia Matschegewski e,f, Nan Huang g, Elliot L. Chaikof h,i,j, Mário A. Barbosa b,c,d, Carsten Werner a, Claudia Sperling a a

Institute Biofunctional Polymer Materials, Max Bergmann Center of Biomaterials, Leibniz-Institut für Polymerforschung Dresden e.V., Dresden, Germany i3S, Instituto de Investigação e Inovação em Saúde, Portugal c INEB, Instituto de Engenharia Biomédica, Rua Alfredo Allen, 208, 4200-135 Porto, Portugal d ICBAS, Instituto de Ciências Biomédicas Abel Salazar, Universidade do Porto, Porto, Portugal e Institut für Biomedizinische Technik, Universitätsmedizin Rostock, Friedrich-Barnewitz-Str. 4, 18119 Rostock, Germany f Institute for ImplantTechnology and Biomaterials (IIB) e.V., Friedrich-Barnewitz-Str. 4, 18119 Rostock, Germany g Key Laboratory of Advanced Technology for Materials of Education Ministry, School of Materials Science and Engineering, Southwest Jiaotong University, Chengdu 610031, China h Department of Surgery, Beth Israel Deaconess Medical Center, Harvard Medical School, 330 Brookline Avenue, Boston, MA 02115, United States i Wyss Institute for Biologically Inspired Engineering at Harvard University, 3 Blackfan Circle, Boston, MA 02115, United States j Harvard-MIT Division of Health Sciences and Technology, 77 Massachusetts Avenue, Cambridge, MA 02139, United States b

a r t i c l e

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Article history: Received 3 May 2019 Received in revised form 29 May 2019 Accepted 13 June 2019 Available online 19 June 2019 Keywords: Hemocompatibility Surface modification Coagulation Blood platelets Passivation Bioactive coating

a b s t r a c t Biomedical devices in the blood flow disturb the fine-tuned balance of pro- and anti-coagulant factors in blood and vessel wall. Numerous technologies have been suggested to reduce coagulant and inflammatory responses of the body towards the device material, ranging from camouflage effects to permanent activity and further to a responsive interaction with the host systems. However, not all types of modification are suitable for all types of medical products. This review has a focus on application-oriented considerations of hemocompatible surface fittings. Thus, passive versus bioactive modifications are discussed along with the control of protein adsorption, stability of the immobilization, and the type of bioactive substance, biological or synthetic. Further considerations are related to the target system, whether enzymes or cells should be addressed in arterial or venous system, or whether the blood vessel wall is addressed. Recent developments like feedback controlled or self-renewing systems for drug release or addressing cellular regulation pathways of blood platelets and endothelial cells are paradigms for a generation of blood contacting devices, which are hemocompatible by cooperation with the host system. Statement of significance This paper is part 4 of a series of 4 reviews discussing the problem of biomaterial associated thrombogenicity. The objective was to highlight features of broad agreement and provide commentary on those aspects of the problem that were subject to dispute. We hope that future investigators will update these reviews as new scholarship resolves the uncertainties of today. Ó 2019 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.

Contents 1. 2.

Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 34 Physico-chemical approaches versus bioactive modification to target the blood-material interaction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 35 2.1. Follow on comments/discussion. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 36

⇑ corresponding author. E-mail address: [email protected] (M.F. Maitz). https://doi.org/10.1016/j.actbio.2019.06.019 1742-7061/Ó 2019 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.

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3. 4. 5. 6. 7. 8. 9.

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Passivating technology: reduced versus controlled protein adsorption . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.1. Follow on comments/discussion. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Biofunctionalization technology: stable immobilization versus release systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.1. Follow on comments/discussion. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Type of bioactives: Biological vs. Synthetic inhibitors . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.1. Follow on comments/discussion. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Target: Enzymes and the coagulation cascade vs. whole cells . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.1. Follow on comments/discussion. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Target venous or arterial (low- or high shear stress) system . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7.1. Follow on comments/discussion. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Highlight on endothelialization as natural hemocompatible implant surfaces: material concepts vs. surface functionalization . . . . . . . . . . . . . . 8.1. Follow on comments. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Conclusion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

1. Introduction Biomedical devices contacting blood disturb the balance of proand anti-coagulant factors in the vasculature. This finely tuned balance is a prerequisite for stable blood circulation which ensures the supply of all tissues with oxygen, nutrients and removal of metabolites. Consequently, a disturbance of this balance may lead to thrombosis and infarctions or induce bleeding. Different from other implants, activation products of blood contacting devices are not restricted to the site of the device, but they may be distributed with the blood stream to remote and vital organs. Local thrombi or platelet aggregates may be released and cause embolization in the next perfused organ. As another example, activated leukocytes during hemodialysis tend to sequester in the lung. The reason why biomedical devices lead to an imbalance is that -unlike the endothelium- they cannot provide anti-coagulant properties and, additionally, they may exhibit pro-coagulant triggers on their surface. Clinically, this usually is counteracted by systemic application of anticoagulant drugs. However, such systemic treatment is associated with enhanced bleeding risks. Surfaces that minimize the activation of blood defense could significantly reduce the amount of systemic medication needed to maintain the balance. Several review articles on hemocompatible surface modification have been published recently, which summarize the existing concepts and technologies to decrease blood coagulation and inflammatory processes at the surface [1–8]. This article refers to

36 37 37 37 38 38 38 38 38 38 39 39 40 41

these review articles for a complete overview and only briefly introduces the molecules applied for hemocompatible surface modification. The main focus is on a discussion-like presentation of application-specific aspects that have to be considered individually (see Fig. 1). Generally, one can distinguish between passive surfaces that attempt to minimize the interaction with blood defense systems and active surfaces that specifically counteract activation processes or support repair. Passivating modifications attempt to suppress protein adsorption either by a stealth effect using uncharged (e.g. poly(ethylene glycol) (PEG)) or zwitterionic (betaines) hydrophilic brushes or hydrogels on the surface (see part 2 of this series) [9]. Alternatively, super-hydrophobic surfaces, which prevent wetting by water or blood, have been suggested to impede the adsorption of proteins. They are based on highly porous surfaces with entrapped air (lotus effect) or with entrapped hydrophobic liquid (SLIPS, ‘slippery, liquid-infused, porous surface’) [10–13]. Bioactive hemocompatible surfaces help to maintain the balance between material and blood processes, as they provide anticoagulant or fibrinolytic properties directly at the implant surface. The local supply of substances to counteract prothrombotic surface properties decreases the total drug load for the patient compared to systemic application. Many bioactive surface modifications attempt to mimic mechanisms of the blood vessel wall. Thus, the anticoagulant proteins tissue factor pathway inhibitor (TFPI) [7,14], thrombomodulin [15,16] or activated protein C [17] have been immobilized on the surface of

Fig. 1. Approaches of hemocompatible surface modification. Various molecules (arranged horizontally), as named in the introduction of this review, are applied to improve the performance of a surface in blood. In their application, different considerations have to be taken and different concepts are followed (arranged vertically), which present the main discussion in this review. The color scheme reflects the graphical abstract and is graded according to the rank of a substance between the different poles.

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biomaterials to be applied on vascular grafts or stents. Grafting of the fibrinolytic proteins tissue plasminogen activator (tPA) or urokinase prevents the deposition of blood clots [18–24]. Coating the indirect anticoagulant heparin is the only active anticoagulant surface treatment currently on the market (various providers) [25]. The linear polysaccharide heparin acts like the glycosaminoglycan heparan sulfate of the blood vessel wall as an indirect anticoagulant [26,27]. Direct peptide-based coagulation inhibitors, namely hirudin or the hirudin derivative bivalirudin, have been suggested as anticoagulant coatings [28,29]. Due to their better stability, synthetic inhibitors to coagulation factors thrombin or Factor Xa (FXa) have been suggested as an alternative [30– 33]. Contact system specific inhibitors, which address high molecular weight kininogen, kallikrein, FXIIa or FXIa, increasingly receive attention for clinical anticoagulation, but so far only corn trypsin inhibitor of this class of inhibitors has been suggested for immobilization on surfaces. Antibodies, peptide and synthetic inhibitors are probably next candidates for immobilization [34–39]. Blood platelets are also selected as targets of hemocompatible functionalized surfaces. The lipid prostaglandin E1 (PGE1), as stimulator of the adenyl cyclase, and the synthetic phosphodiesterase inhibitor dipyridamole, both have been immobilized to polymer surfaces in stable and in release systems for suppressing platelet aggregation [40–42]. As adenosine diphosphate (ADP) is among the strongest activators of blood platelets, immobilization of the ADP degrading enzyme apyrase can decrease the thrombogenicity of surfaces [43]. Nitric oxide has multiple functions in vivo, including prevention of platelet activation and aggregation on a cGMP dependent pathway. NO releasing coatings have been developed either as direct release system with immobilized NO donors (Ndiazeniumdiolates) or with immobilized catalysts (organoselenium- or copper based), which release NO from endogenous or pharmacologically substituted soluble donors (Snitrosothiol compounds) [44–48]. Finally, the fibrinogen-receptor (GP IIb/IIIa) has been blocked by the antibody abciximab using an elution system on vascular stents [49]. While many of the mentioned strategies of bioactive surface modification attempt to mimic endothelial properties by the device surface, several approaches directly take use of vital endothelium as ideal of a hemocompatible surface [50–53] and support endothelialization of the device surface. These cover specific surface texture, proteins of the extracellular matrix (fibronectin and laminin) and growth factors (VEGF) [54–56], or concepts for

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in situ capture of endothelial cells or their progenitors [57–63]. Besides the support of endothelialization, suppression of excessive proliferation of vascular smooth muscle cells on vascular stents, clinically apparent as restenosis, is a target of surface treatment. Coatings with anti-proliferative drugs, mainly paclitaxel, and drugs of the –limus family (sirolimus, everolimus, zotarolimus) or statins are commonly applied on these devices to prevent restenosis [64– 66]. All types of surface modification provide certain advantages and disadvantages, and hemocompatible coating strategies have to be selected according to requirements of the specific applications. The following discussion in this text should provide a guideline for the selection of modification strategies, indicated in Fig. 1, but also highlight the persisting limitations and provoke new ideas for this exciting and widely applied field of research. Approaches of bioactive surface modification are compared with passive modifications. Consideration of arguments in favor of or against the incorporation of active players on biomaterials surfaces is presented. Several strategies, how these substances are provided, either by permanent immobilization on the surface or in release systems, are discussed. Additionally, pro-thrombogenic reactions that are preferential targets are considered. Several possibilities to interfere with the blood coagulation cascade and platelets are presented and possibilities to support endothelialization of surfaces are highlighted. 2. Physico-chemical approaches versus bioactive modification to target the blood-material interaction In contrast to passive modifications, which mainly act via a stealth effect, bioactive hemocompatible surfaces act by specific inhibition of coagulation factors, blood platelets or other players in coagulation and hemostasis (an overview on the different strategies can be found in Fig. 2). Anti-inflammatory modifications may act by inhibition of the complement system or leukocytes. Alternatively, biofunctionalization strategies may be applied to support endothelialization of the surface to make use of the superior hemocompatible properties of healthy endothelium. These specific interactions typically provide more effective hemocompatibility than passive coatings. However, as a disadvantage, bioactive modifications frequently suffer more from degradation, saturation or exhaustion in vivo than passive modifications [6]. Bioactive surfaces therefore have their main domain for temporary application,

Fig. 2. Overview of passive versus bioactive strategies for surface modification.

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such as catheters or extracorporeal artificial organs. They are also useful when endothelialization, as a desired physiological effect, takes over the anticoagulant properties of the coating after a certain time, such as on vascular stents. 2.1. Follow on comments/discussion Passive strategies for hemocompatible surface modification have the longest tradition and still are applied for long-term blood contacting devices or as a complementary strategy, when bioactive modifications are exhausted. They frequently provide a persistent and substantially better hemocompatibility than the untreated surface [1–4,6,7,67]. Control of surface properties such as wettability, charge, morphology, crystal structure or electronic surface properties using chemical and physical methods, have been applied to influence composition and conformation of the adsorbing protein layer in a passivating way. Low conformational change and an accumulation of passivating proteins like albumin in the protein adsorbate is intended [3,9]. Low adsorption, however, may also pose the risk of ‘non-adhesive encounters’. Surfaces may lead to the activation of blood platelets or the coagulation cascade, but resist the formation of a blood clot. Activated cells or thrombotic material spall from the surface and cause harm in downstream organs, as demonstrated for Poly(hydroxyethyl methacrylate) (pHEMA) based hydrogels [68]. Bioactive coating strategies frequently also apply passivating molecules like poly(ethylene glycol) (PEG) as non-adhesive spacer molecules. These have various functions: they prevent non-specific protein adsorption, provide mobility to the biomolecule and shield the biomolecule and its reaction partner from influences of the surface [69]. Thus, concepts of passive coatings are also a basis for the active coatings. Passive coatings usually intend to modulate protein adsorption and to decrease the degree of denaturation of adsorbed proteins. The immobilization of small molecules with selective affinity for plasma proteins can support protein adsorption in a physiological conformation. For instance, aliphatic chains, resembling fatty acids, have been immobilized. Bilirubin or fatty acids use albumin as transport molecule. The such decorated surface exposing bound albumin without conformational change exhibits good hemocompatibility [70–74]. Although this modification is not bioactive, it applies highly selective molecular interactions to obtain a high blood compatibility. The clinical success of bioactive coatings is typically limited because they lose their activity due to saturation and consumption. In addition, immobilized biomolecules may undergo degradation during storage or due to physiological turn over in vivo. The lifetime of release systems may be increased via the reservoir capacity of the device. Saturation and consumptions are not general drawbacks of bioactive surfaces. The indirect coagulation inhibitor heparin is frequently applied as hemocompatible coating. It acts as a catalyst and is not used up [75]. However, loss of anticoagulant activity of a heparinized surface due to heparin degradation or nonspecific protein adsorption is a matter of debate. There are reports that explants of heparinized surfaces after 12 weeks in vivo still showed antithrombin-binding activity [76]. But the heparin degrading enzyme heparanase has elevated concentrations in the plasma of typical recipients of stents and vascular grafts, such as diabetic patients and patients with kidney disease [77,78]. This enzyme may limit the lifetime of the coating in vivo [78]. Furthermore, a heparin coated surface lost its anticoagulant property after one hour incubation in whole blood in vitro, observed in subsequent exposure to fresh blood [79]. Authors attributed this loss of anticoagulant activity mainly to shielding effects of adsorbed proteins, whereas the heparin activity assay of the resorted 12-

weeks in vivo samples had been determined after clearing the adsorbed proteins [76]. Other examples of long-term activity due to a non-exhaustive catalytic effect are nitric oxide (NO) generating surfaces, based on organoselenium or copper compounds, with glutathione peroxidase-like activity [44–48], immobilized ADP degrading apyrase [43], thrombomodulin [15,16] or activated protein C [17]. Shielding by protein adsorption and proteolytic degradation, respectively, appear as the main routes of fading activity with time. In addition, in situ renewal of surface functionalization is possible: enzyme ligation offers an opportunity for catalyzing reversible bond forming reactions that could enable the molecular regeneration of bioactive thin films in situ. Staphylococcus aureus sortase A (SrtA) catalyzes the covalent transpeptidation of a C-terminal ‘‘sorting motif” LPXTG to N-terminal oligoglycine (e.g. GGG) nucleophiles through an acyl-enzyme complex forming an LPXT-GGG bond [80,81]. Due to its synthetic simplicity and the very limited occurrence of the LPXTG motif in native proteins, SrtA-catalyzed transpeptidation has been broadly applied in protein purification, labeling, and immobilization onto solid supports [82–86]. The low catalytic activity and substrate affinity of wild type (WT) SrtA necessitates high molar excess of the enzyme and long incubation times to approach reaction completion, limiting its effectiveness and applicability. An evolved SrtA mutant (eSrtA) has recently been generated that exhibits 120-fold higher LPETG-ligation activity than the wild-type enzyme [85,87]. This enzyme suggested the possibility of multiple rapidly catalyzed cycles of removal and reversible assembly of bioactive molecular films onto oligoglycine modified surfaces both in vitro and in vivo. Indeed, eSrtA transpeptidation is reversible, enabling multiple cycles of peptide bond formation and cleavage, and thereby, facilitating film regeneration in the presence of whole blood in vitro and in vivo [88]. These studies have established a rapid, orthogonal, and reversible scheme, which can be used to regenerate selective molecular constituents with the potential to extend the lifetime of bioactive films, or likewise, as a method to load and release any of a number of materialbound constituents for controlled drug loading and delivery, or as a strategy for material dissolution [88]. This advances the concept of ‘‘in situ regeneration” of bioactive surfaces at will. The same principle would apply to regenerating a device with immobilized drug containing micro- or nano-particles, whether the drug is synthetic or biologic, once the reservoir is depleted. While this area is evolving, it is likely that additional schemes, which are suitable for in vivo application, will be developed in the future. This principle could also be applied to regenerating a ‘‘non-adsorptive” surface coating, such as PEG, if that surface fouls over time. Loss of thromboprotective property is not restricted to bioactive modifications, since protein resistant modifications also have only a limited lifetime. This can be caused through an oxidative degradation of the coating and the resulting loss of its passivating properties. Therefore, the clinical application of PEGylated surfaces is limited, despite their success in vitro [89,90].

3. Passivating technology: reduced versus controlled protein adsorption Passivating approaches for biomaterials hemocompatibility suppress blood activation processes by reduced interaction of blood components with the surface. Highly hydrophilic, water binding molecules provide a stealth effect on the surface, which should prevent the interaction with cells and proteins in blood [9,74]. Superhydrophobic surfaces (self-cleaning coatings) show a specific nanostructured geometry that creates an energy barrier to prevent contact of liquid with dissolved proteins and cells with surfaces [12,13]. Besides these extreme concepts, surface

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hydrophilicity and charge may be also tuned to support preferential adsorption of passivating proteins, like albumin, and adsorption in native, non-activated conformation. 3.1. Follow on comments/discussion Highly hydrophilic, water binding, or very hydrophobic, water repellent surfaces are the leading concepts to suppress protein interaction with the surface. Poly(hydroxyethyl methacrylate) (pHEMA) and poly(ethylene glycol) (PEG) bind water molecules through hydrogen bonds to the hydroxyl- or ethylene oxide (– CH2–CH2–O)– chains, creating a strongly bound hydration layer. Steric exclusion due to densely packed and flexible extended PEG chains (PEG brushes) contributes further to the antifouling property of PEGylated surfaces. However, although all these surfaces were very successful in vitro, their performance in vivo did not fulfill expectations, perhaps due to oxidative and enzymatic degradation of PEG when used for long periods of time [89,90]. External cell membranes are mostly composed of phospholipids that contain an electrically neutral zwitterionic head group (phosphorylcholine). Synthetic zwitterionic surfaces based on phosphorylcholine, sulfobetaines or carboxybetaines are thought to mimic the physical properties of the external cellular membrane and therefore promise high hemocompatibility. Besides mimicking cell membrane properties, they are highly anti-adhesive due to their strong hydration capacity. The charged groups can bind water molecules tightly in a structured way by electrostatic interactions. In addition, their non-fouling capacity is improved when surface hydration is combined with chain flexibility (steric repulsion) [91], whereas soluble or nanoparticle-shaped betaines may also lead to platelet aggregation and activation [92]. Superhydrophobic surfaces depend on their specific nanostructured geometry that creates an energy barrier to prevent liquid from entering the cavities of the nanostructure, resulting in water droplets rolling off from the surface [10–13]. However, in physiological conditions, these surfaces lose their repelling capacity, since the high protein concentration in blood decreases the surface tension [9]. It also has to be considered that a liquid at a superhydrophobic surface actually forms a liquid-air interface with high protein denaturating characteristics [93]. Nevertheless, a superhydrophobic membrane coating, consisting of a complex disordered, fractal-like network of fluorinated silicon oxide nanospheres, was developed for blood oxygenation. In this coating gas is trapped between the nanospheres, thus avoiding blood to contact the membrane for at least 24 h [12,13]. Fluorinated materials, such as expanded polytetrafluoroethylene (ePTFE; TeflonÒ) have been described as bio-inert and non-thrombogenic and have been used clinically for blood contacting medical devices, such as large diameter vascular grafts. Although PTFE frequently is rapidly covered densely with blood platelets [94], it does not show high thrombogenicity in clinical application and surface fluorination even is recommended to decrease platelet activation [95]. A passivating form of monolayer platelet adhesion has been suggested [96]. The nonthrombogenicity may be restricted to the high blood flow rates at the luminal surface and thus the materials cannot be used in small-caliber artificial grafts (<5 mm), where the flow rate is lower [97]. The ratio of albumin to fibrinogen adsorption of surfaces is frequently regarded as an indicator of blood compatibility, as albumin is a passivating protein, whereas fibrinogen supports platelet adhesion (see part 2 of this series) [95,98,99]. However, the conformation of adsorbed fibrinogen is more relevant for platelet adsorption than the total amount, because only denatured fibrinogen gamma chains expose platelet binding sites [100,101]. A soft fibrinogen layer, which is adsorbed via the aC domains does not expose the

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binding sites or cannot transduce the mechanical forces for integrin signaling to adherent platelets and therefore does not get covered by platelets [102]. Multiple ways of surface modification by chemical or physical methods have been applied to tune the composition and conformation of the protein adsorbate in a favorable way [103,104]. Although general rationales for surface properties with favorable protein adsorption have been suggested, it is not yet possible to predict the leading parameters of hemocompatibility from a given set of physical-chemical surface properties. Various strategies have been developed to create albuminbinding surfaces as a rational strategy for albumin adsorption in its native conformation to prevent cell adhesion and the adsorption of other proteins involved in coagulation and complement activation. Coatings rely on hydrophobic patches within hydrophilic surfaces (e.g. CH3/OH) [105,106] or, as mentioned before, on the immobilization of albumin-binding compounds, such as antibodies, warfarin, bilirubin- or fatty acid-like compounds, such as alkyl chains with 16–18 carbons [70–74]. However, despite promising results in vitro, none of these strategies were translated into clinical setting. 4. Biofunctionalization technology: stable immobilization versus release systems Covalent immobilization of bioactive molecules is the most stable form of biofunctionalization. However, covalent immobilization may be challenging to establish, because molecules have to be immobilized in correct orientation and the immobilization technology must not interfere with their interaction with the target molecule [16]. As a matter of principle, surface immobilized molecules only act locally and do not influence coagulation or inflammatory processes in the blood volume. Surfaces with stably immobilized inhibitors and antagonists generally are not renewable as soon as they are occupied by their target molecules. In this context, planar surfaces of biomedical devices allow only low amounts of target molecules to bind [31]. Release systems overcome most of these problems providing a much higher capacity for the bioactive substance, but eventually all of the bioactive component is consumed. 4.1. Follow on comments/discussion Direct immobilization of bioactive molecules is frequently associated with a loss of activity because of insufficient flexibility of the molecules to reach the binding sites of the target molecules [69]. Therefore, the linear molecule heparin has been end-point immobilized to provide access to binding sites inside the chain for coagulation factors and for antithrombin [107,108]. Other molecules are usually immobilized via flexible spacer molecules, like PEG. These spacers also help to overcome some problems of unsuitable orientation of the bioactive molecules. However, many molecules of interest, like the clinically applied direct oral anticoagulants (DOACS) dabigatran, rivaroxaban, apixaban, or the direct thrombin inhibitor Argatroban do not contain suitable reactive groups for immobilization in active form. In these cases, modification of the inhibitor molecule is necessary for a conjugation to a surface [32]. The low surface density of irreversibly immobilized molecules can be compensated by presentation of the bioactives in hydrogels with a suitable mesh size to allow the penetration of the target molecules [109]. While immobilized synthetic inhibitors usually show only local activity, most of the physiological inhibitory systems shift their activity from the (cell-)surface into the fluid phase. Cells release tissue factor pathway inhibitor (TFPI), or surface-associated thrombo-

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modulin acts by release of activated protein C (APC) into the blood phase. This suggests that synthetic concepts, which are based on release systems, may simulate physiological conditions better. Suitable release profiles without over- and under-dosage and without being influenced by local pH or other metabolic influences are crucial for release systems. Physiological release systems act as feedback control systems, which release bioactive substances only upon a special trigger. Recently, such adaptive systems, which respond on the current coagulation activity, have been developed by different groups and show an important step towards smart bioactive materials [79,110–113]. These systems may consist of nanocomplexes of the anticoagulant heparin and a peptide, which is cleaved by activated coagulation factors. Active coagulation cleaves the peptide and releases the anticoagulant heparin at a self-titrating concentration [110]. Alternatively, the coagulation cleavable peptides can be used as covalent cross-linkers in hydrogels of 4-armed poly(ethylene glycol) (starPEG) and heparin. Activated coagulation enzymes degrade the hydrogel and release heparin at an adjusted amount to stop the coagulation. Upstream coagulation factors in the coagulation cascade proved to have better timing of the anticoagulant release and can induce the same anticoagulant effect at a lower heparin release [79,111]. Instead of heparin, also the fibrinolytic tissue plasminogen activator (tPA) has been released from coagulation-responsive hydrogels or nanocapsules to degrade a clot as it forms [112,113]. 5. Type of bioactives: Biological vs. Synthetic inhibitors Bioactive surfaces containing biomolecules with peptide/protein or polysaccharide structure are closest to the biological system. However, biomolecules have the limitation that they undergo inactivation and degradation in vivo. They typically also require expensive preparation and purification or production by recombinant organisms and have limited stability at sterilization or shelf storage. Synthetic inhibitors are frequently inspired and optimized from their biological model. They are optimized for higher affinity and show favorable stability at sterilization, storage and in vivo. 5.1. Follow on comments/discussion Due to their mode of action, immobilized inhibitors are prone to saturation and consumption. This is much less the case for natural indirect inhibitors such as heparin, thrombomodulin, or tissue plasminogen activator. They act as catalysts and avoid the problem of saturation. However, they depend on additional molecules (AT III, Protein C, or plasminogen, respectively), which may be limited in cases of severe illness [114]. Heparin-antithrombin conjugates have been suggested to overcome such limitation [115]. The limited stability of the biomolecules in vivo, which also leads to a loss of activity has been discussed before. In order to avoid loss of activity of biomolecules during processing of the final device, sterilization methods with less protein denaturation than steam sterilization may be required. Biomolecules may be more harmful than synthetic analogues. Examples are immuno-reactions against non-human proteins and peptides, as described for hirudin [116]. Life-threatening heparininduced thrombocytopenia type II (HIT II) has been ascribed also to immobilized heparin, similar to systemic heparin [117,118]. 6. Target: Enzymes and the coagulation cascade vs. whole cells Coagulation factors are the main target of hemocompatible bioactive surface modifications, because of their well-described structure and activity. Alternatively, the fibrinolytic pathway has been a target of surface modifications [2]. In contrast to the

enzyme systems, blood platelets show a complex response and are rarely a focus of bioactive surface modification. In contrast to surface modification aiming at the inhibition of blood platelet activation, the support of endothelialization of vascular stents via immobilized antibodies or aptamers for endothelial progenitor markers is a frequently applied approach [61]. 6.1. Follow on comments/discussion The predominant target molecules for current clinically applied bioactive coatings are coagulation factors, mainly FXa and thrombin via heparin-catalyzed inhibition by antithrombin. Experimental strategies aim at direct inhibition of thrombin, Factor Xa, or Factor XIIa [28–33]. Many biomaterials induce coagulation by the contact phase system and the following intrinsic pathway (discussed in part 3 of this series). Specific inhibitors of contact activation, such as C1 inhibitor, corn trypsin inhibitor, infestin or specific antibodies 15H8 and 3F7 find interest, because they do not affect physiological, tissue-factor mediated, coagulation [34–39]. A different strategy is to support the fibrinolytic system. Tissue type plasminogen activator (tPA), urokinase, streptokinase and active plasmin have been immobilized directly to surfaces [18– 24]. In addition, surfaces with enhanced affinity for tPA and plasmin, mimicking structural properties of a fibrin clot have been formed to accumulate these enzymes from plasma and enhance local fibrinolysis [119,120]. All concepts of enhanced fibrinolysis, however, pose a risk of fragile thrombus formation and embolization. While there is a plethora of concepts to target individual enzymes, there are much less anti-platelet and, to the best of our knowledge, no anti-leukocyte surface modifications [40–43,49]. Besides the complex molecules, also nitric oxide (NO) generating and releasing systems target blood platelets [44–46]. Surface modifications directed against blood cells typically require drug release instead of surface immobilization to prevent cell adhesion to the surface via receptor/inhibitor interaction. Many drugs bind to intracellular ligands. This excludes their application for surface immobilization and allows only drug release systems. Antibiotics or antiseptic substances are applied on central venous catheters to prevent bacterial colonization with the risk of catheter related thrombosis and sepsis [121]. However, many antiseptics as such show pro-coagulant properties and require additional anticoagulant surface treatments [122]. Intact endothelium presents the best hemocompatible surface, which is desired for all long-term implants like vascular stents or blood vessel grafts [50–53]. Therefore, there are many coating strategies to promote the growth of endothelial cells and to suppress the growth of vascular smooth muscle cells, which are summarized and discussed below. 7. Target venous or arterial (low- or high shear stress) system It is accepted clinical knowledge that arterial thrombosis is mainly driven by shear-activated blood platelets. Venous thrombosis, however, is driven by the plasmatic coagulation cascade. Consequently, patients with stents or arterial grafts have more benefit from systemic anti-platelet drugs (aspirin, P2Y12 inhibitors e.g. clopidogrel) for prophylaxis of stent/graft thrombosis than from anticoagulants. Prophylaxis of venous thrombosis or thrombi by atrial fibrillation, however, is based on systemic anticoagulants (heparin, warfarin or DOACs) [123–126]. 7.1. Follow on comments/discussion Such site-specific behavior is not yet well considered in the design of hemocompatible surfaces. More focus should be put on

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the fact that surface modification for venous catheters or venous valves may benefit more from addressing the coagulation cascade, whereas stents, arterial grafts or heart valves might benefit more from addressing blood platelets. Nitric oxide generating [44–48] or abciximab eluting [49] coatings for vascular stents, which prevent platelet activation, follow this concept and await further validation in the future. 8. Highlight on endothelialization as natural hemocompatible implant surfaces: material concepts vs. surface functionalization The endothelial cell layer provides the natural antithrombogenic surface in the cardiovascular system. Therefore, creating implant surfaces that support endothelialization at the implant site is one of the major goals in the design of bioactive endovascular devices [50–53] (see also Fig. 3). There is more than one decade of clinical experience with drug eluting systems to suppress excessive proliferation of vascular smooth muscle cells. More tailored approaches recently support adhesion and proliferation of endothelial cells and endothelial progenitor cells from blood. 8.1. Follow on comments Support of endothelial cell growth (as pictured in Fig. 4) can already be achieved without specific pharmacology by selection of appropriate substrate materials. Poly(L-lactic acid)/poly(4hydroxy butyrate) (PLLA/P(4HB)) blends and magnesium-based stent materials show better endothelialization than corresponding reference materials in vitro and in clinical studies [127–133].

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However, for these biodegradable materials, the distinct biological actions in response to the degradation products are still the subject of intensive scientific investigation [130]. Furthermore, enhanced adhesion of cells alone is not sufficient for appropriate endothelial expression of antithrombogenic markers. An endothelial layer in an activated state loses its protective effect by expression of a pro-coagulant and pro-inflammatory phenotype with downregulated thrombomodulin, tissue plasminogen activator (tPA), tissue factor pathway inhibitor (TFPI), NO synthesis and reduced heparan-sulfate containing anticoagulant glycocalyx, while up-regulated ligands for platelet and leukocytes adhesion (P/Eselectin, VCAM-1, ICAM-1, vWF) or activating coagulation by tissue factor expression [50,53]. The design of a specific surface topography can support the performance of endothelial cells (see also Fig. 4). Parallel grooves, stripes of fibronectin, hyaluronic acid, or heparin with width of about 25 mm tune morphology and orientation of adherent endothelial cells similar to the flow condition in vivo. In addition, on some topographies, cell functions like the the secretion of extracellular matrix proteins, NO, Prostaglandin I2, and tissue factor pathway inhibitor better resemble the physiological condition [134–136]. Beyond the scaffold material itself, surface coatings comprising chemical and/or biological functionalization with biomolecules are frequently applied approaches aiming to support vascular healing and re-endothelialization. Surface immobilization with biomolecules, such as endothelial specific antibodies (e.g. anti-CD34, anti-CD133), anti-cadherin, immobilization of DNA or peptide aptamers, has been shown to promote re-endothelialization in animal studies. This was achieved

Fig. 3. Scheme of challenges and approaches in the development of endovascular implants. Successful implant integration including prevention of thrombosis, inflammation and fibrosis desires for stimulation of re-endothelialization at the implant site (green arrows) as well as inhibition of smooth muscle cell hyperplasia in order to prevent instent-restenosis (red arrows). (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)

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Fig. 4. Endothelialization of implant surfaces to prevent clot formation and thrombosis after implantation by forming an inherent hemocompatible layer: cell attachment and spreading of human endothelial EA.hy926 cells grown on electrospun nanofibrous polymeric scaffold (A) and plain polymeric surface (B) for 48 h showing dependencies of cell attachment and phenotype maintenance from surface topography (scanning electron micrographs; bar = 20 lm). (Image source: C. Matschegewski, N. Grabow).

from local endothelial cells and by capturing circulating endothelial progenitor cells [57–63]. Unfortunately, such functionalized surfaces are not selective enough and the growth of other cells is possible [51,52]. Non-specific adsorption of plasma proteins on the surface supports this adhesion of other blood cells, which obviates the impact of the intended surface modification. Surface functionalization with vascular endothelial growth factor (VEGF) has been shown to promote some measures of endothelialization [54–56]. Rapid exhaustion of the growth factor may be avoided by local gene therapy, where the surface is functionalized with the gene and a suitable carrier for non-viral transfection [137,138]. Anti-inflammatory and anti-proliferative effects are obtained by drug-eluting stents applying Limus drugs (e.g. sirolimus, everolimus, zotarolimus) or statins to effectively prevent restenosis, the occlusion of the stented vessel lumen by proliferating smooth muscle cells [64–66]. Nevertheless, a negative impact on reendothelialization due to the generalized anti-proliferative effect was reported. The non-endothelialized stent struts are the origin of late stent thrombosis. Novel technology concepts like dual drug-eluting stents with differential effects on the vascular endothelium and smooth muscle cells have been used to sustain biocompatibility and endothelialization while preventing restenosis [139]. Studies on dual drug eluting stent systems that are abluminally covered with sirolimus and luminally with the antiproliferative atorvastatin showed a stabilizing effect on endothelial cell function without disabling sirolimus’ drug efficacy in terms of inhibiting smooth muscle cell growth [140]. Recent multi-molecule immobilization approaches are intended to create multi-functional microenvironments to promote endothelialization and consequently to prevent thrombosis. Studies on co-immobilization of biomolecules, for example with heparin, fibronectin and VEGF as multi-functional coating on titanium substrates were found to enhance proliferation of endothelial cells as well as of endothelial progenitor cells [141]. In the context of promoting recruitment and proliferation of endothelial cells, advantageous techniques such as layer-by-layer coating is of rising interest in order to build multilayer films of biomolecules on material surfaces. Layer-by-layer fabricated gelatinchitosan-VEGF film on memory shape nickel-titanium alloy substrates were shown to support attachment and proliferation traits of endothelial cells in vitro [142]. There is ongoing research to explore the inhibition of other regulatory pathways in endothelial- and smooth muscle cells. For instance, the phosphatidylinositol 3-kinase PI3K/p110a pathway is involved in inflammation and thrombosis and plays a crucial role in the induction of pro-thrombotic proteins, such as tissue factor

(TF) and plasminogen activator inhibitor-1 (PAI-1). Its inhibition by the PI3K inhibitor PIK75 was shown to selectively inhibit neointima formation while maintaining endothelial function and reendothelialization following vascular injury. In particular, application of PIK75 resulted in delayed arterial thrombus formation in vivo in a carotid artery thrombosis mouse model, as well as in decreased expression of the pro-thrombotic proteins TF and PAI-1 [143]. Pioneering approaches in the context of targeting the PI3K pathway focus on the usage of the biomolecule fucoidan, a sulfated polysaccharide from brown seaweed that was demonstrated to exhibit heparin-like anticoagulant efficiency and antithrombotic activity [144]. In vivo experiments in a rabbit model showed reduced restenosis in fucoidan-coated stented animals accompanied with decreased smooth muscle cell proliferation in vitro [145,146]. Detailed studies on the effects of fucoidan on cell signaling and gene expression in endothelial cells detected an activation of PI3K pathway resulting in the induction of endothelial cell migration [147]. Thus, selective targeting of the PI3K/p110a pathway might provide a promising strategy to design hemocompatible surfaces and preventing device-related thrombosis and restenosis. Beyond those in vitro functionalization strategies aiming at capturing endothelial cells, ex vivo seeding of vascular grafts with endothelial progenitor cells has raised interest in tissue engineering approaches. In particular, umbilical cord blood-derived endothelial cells are a highly promising source of endothelial precursor cells that are assumed to survive in vivo more likely than for example human umbilical vein cells. This advantage seems to be founded in their higher proliferative capacity and higher sensitivity to angiogenic factors [148]. These cells seem to be the ideal source and are therefore increasingly used in tissue engineered grafts.

9. Conclusion Activation processes in blood and vessel wall upon contact with a biomedical device are an ongoing clinical problem, as highlighted in the first part of this series. Although the device is localized, complications, such as embolism, may be remote or even systemic. For this localized focus, clinically usually systemic anti-coagulant or anti-platelet prophylaxis usually is applied, which also poses systemic risks of bleeding and other side effects. Multiple strategies have been developed to reduce these activation processes at biomedical device surfaces, in order to reduce side effects of the device and in consequence also of the prophylactic treatment. They target the whole range from the primary response of proteins in contact with the foreign surface (see part

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