Biocompatibility Pathways: Biomaterials-Induced Sterile Inflammation

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Biocompatibility Pathways: Biomaterials-Induced Sterile Inflammation, Mechanotransduction, and Principles of Biocompatibility Control David F. Williams* Wake Forest Institute of Regenerative Medicine, Richard H. Dean Biomedical Building, 391 Technology Way, Winston-Salem, North Carolina 27101, United States ABSTRACT: This paper addresses a significant paradox in biomaterials science; biocompatibility phenomena have been experienced and described for over 50 years but without an agreed understanding of the framework of mechanisms that control the events that occur when a biomaterial is exposed to the tissues of the human body. The need for such an understanding has become more urgent as biomaterials are now used in wide-ranging applications such as tissue engineering, drug and gene delivery, and imaging contrast agents. A detailed analysis of these phenomena, especially in terms of clinical outcomes rather than in vitro experiments, determines that two overarching mechanisms, mechanotransduction and sterile inflammation associated with damageassociated molecular patterns, are responsible for the vast majority of phenomena. In contrast, interfacial interactions, for so long being assumed to play pivotal roles in biocompatibility, especially relating to protein adsorption, are actually relatively unimportant unless, through conformational changes, they are able to participate in 3D ECM development. Critical to this new view of biocompatibility is the fact that the combination of mechanotransduction and sterile inflammation, especially focusing on inflammasome activation and the immunology of the balance between inflammation and fibrosis, allows biomaterials science to encompass mechanisms of innate and adaptive immunity without recourse to the traditional implications of pathogen induced responses of the immune system. In this way, a system of biocompatibility pathways can be generated; these are able to explain a wide range of clinical biocompatibility challenges, including nanoparticle translocation and internalization, intraocular lens opacification, leukocytedominated responses to metallic wear debris in joint replacement, stem cell differentiation of nanostructured hydrogels, tissue responses to incontinence meshes, and restenosis of intravascular stents. Perhaps even more importantly, the identification of these molecular pathways of biocompatibility offers prospects of the control of the host response by targeting specific points in these pathways, for example the inhibition of epithelial to mesenchymal transformation that can result in excessive fibrosis, and the inhibition of activation of the NLRP3 inflammasome following exposure to biomaterial-induced stresses; this should lead to a more effective translation of biocompatibility understanding into better clinical outcomes. KEYWORDS: inflammasome, fibrosis, damage-associated molecular patterns, medical devices, tissue engineering, nanotoxicology

1. INTRODUCTION

system, and that there is no such thing as a universally biocompatible material. At the time when the term was starting to be used seriously, the majority of relevant applications pertained to implantable devices, with some references to artificial organs, extracorporeal systems, wound healing products, and simple drug delivery systems. Essentially, biocompatibility was considered as a perturbation of the wound healing process that inevitably occurred following a surgical procedure to implant a device; the result was described as the foreign body reaction.3 The range of applications of biomaterials has opened up considerably in recent years,4 so it is necessary to consider the host response in terms of tissue engineering products, drug, gene and vaccine

Biocompatibility is a widely used, but poorly understood, term. Appearing intermittently in the literature for several years, it was first seriously defined in the 1980s at a consensus conference on definitions in biomaterials.1 The definition that emerged was that biocompatibility refers to “the ability of a material to perform with an appropriate host response in a specific application”. Although recognizing that this is a conceptual definition that does not have practical utility, it did firmly embed in the thinking of biomaterials scientists that biocompatibility has to be considered in terms of the precise situation in which a biomaterial is used. It follows that biocompatibility phenomenon associated with any one biomaterial will vary depending on the application. Two consequences of this, discussed recently,2 are that biocompatibility is not a property of a material but of a biomaterial - host © 2016 American Chemical Society

Received: October 1, 2016 Accepted: November 29, 2016 Published: November 29, 2016 2

DOI: 10.1021/acsbiomaterials.6b00607 ACS Biomater. Sci. Eng. 2017, 3, 2−35

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ACS Biomaterials Science & Engineering Table 3. Tissue Compartmentsa

delivery systems, and diagnostic systems in addition to implantable devices. Wound healing, therefore, is not a good starting point to discuss mechanisms.5 More importantly, the use of biomaterials in the majority of implantable devices has largely been predicated on the need to minimize interactions between them and the host, and most tests to determine “biological safety” of products is based on the need to approach chemical and biological inertness. However, this does not apply to all biomaterials applications, and certainly not to those used in tissue engineering where scaffold materials are required to facilitate molecular and mechanical signaling to the target cells. In view of these issues, attention has turned toward the identification of mechanisms by which biomaterials and hosts interact with each other within the multitude of biocompatibility scenarios.6−8 There are currently over 100 different classes of biomaterial in clinical use or in an advanced state of testing for eventual use.9 These demonstrate a remarkable array of chemical, mechanical, physical and biological properties (summarized in Table 1) and can be presented to host tissues with wide-ranging

solid tissue or organ in vivo cardiovascular system in vivo nervous system in vivo sensory organs in vivo tubular system in vivo skin and mucous membranes ex vivo bioreactor, cell culture system extracorporeal circulatory system a

This nonexhaustive list indicates the range of tissue compartments that may be exposed to biomaterials.

variations that are introduced by clinical skills and patient specific factors (age, gender, diabetes, medications, lifestyle, etc.), and also the immense difficulties we have in correlating in vitro and animal test data with clinical performance in humans, we could anticipate that a comprehensive, overarching description of biocompatibility mechanisms would be impossible to attain. Faced with this situation, there are several potential procedures that could be used to establish a framework of biocompatibility mechanisms. One such procedure could involve the identification of broadly based biocompatibility pathways, that is, the major sequences of events, grounded in the established processes of materials and biological sciences, that control the development of the host response in any given situation.10 In this paper, the essential features of biocompatibility pathways are set out and discussed. This is neither a matter of semantics nor an academic exercise. If an overarching framework of biocompatibility pathways can be identified, then mechanisms and procedures that could lead to the control of biocompatibility maybe identified. This is indeed where this analysis leads, which will be explored at the end of the paper.

Table 1. Characteristics and Properties of Biomaterialsa chemical nature

mechanical properties

physical properties

biological properties

metallic (pure metal, alloy, etc.) polymeric (synthetic, biopolymer etc.) ceramic (oxides, phosphates etc.) carbons (crystalline, amorphous, nanostructured etc.) composites engineered biological components elastic modulus elastic limit hardness ductility strength electrical electronic optoelectronic magnetic biostability biodegradability bioresorbability permeability cytotoxicity

2. FUNDAMENTAL BIOCOMPATIBILITY PARADIGM With very few exceptions, when man-made, engineering, or commodity materials are used as biomaterials, they are not intrinsically compatible with physiological systems, nor have they been designed to be so. Moreover, the tissues of the human body have not evolved in order to benignly accommodate these materials within their midst and they are treated as “potentially harmful”. The default position, therefore, is that there is inherent incompatibility between these two compartments, the biomaterial, and the tissue. It is, however, much more serious than that. The problem with this default position is that the human body has evolved in such a way as to have exquisite detection mechanisms that readily identify foreign objects, and there are powerful defensive mechanisms that deal with such objects once they have been detected; we are confronted, therefore, with an active incompatibility and not a passive one. These mechanisms evolved naturally to deal with bacteria and viruses, but they are often capable of diversion toward any synthetic material that might find its way into the body or any type of biological stress that may arise with this use. This becomes especially important when biomaterial components have some similarities with bacteria and viruses, both in size and chemistry, so that these reactions are invited to take place and means to avoid them have to de devised. The introduction of a biomaterial into the human body normally represents a physiologically stressful event and the body is expected to present some adaptive

a

This nonexhaustive list summarizes the principal material groups and property characteristics that may influence biocompatibility.

morphologies (Table 2). The tissue compartments or components with which they come into contact are also very varied (Table 3). Add to these multiple characteristics the Table 2. Morphologies of Biomaterial Components solid macroscopic object solid object with microscale topography solid object with nanoscale topography solid object with porous surface solid porous object macroscopic mesh homogeneous gel heterogeneous, nanostructured gel microparticulate collection nanoparticulate collection thin or hollow fiber membrane 3


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ACS Biomaterials Science & Engineering

contrast to most forms of the biomaterial, however, the tissue provides a dynamic environment in which the cells and their ECM change over time. Solid tissue is also permeated by circulating cells, such as blood and inflammatory cells, and by circulating molecules, such as proteins. In this simplest of conditions, the biocompatibility of the system is concerned with the effect that the physical presence of the material has on the dynamic response of the tissue components. This is the conceptual starting point, from which a series of positions emerge where the biomaterial is presented to the physiological environment in different forms or in different situations. Each of these takes the basic inert scene and adds, in turn, the complexities of chemical reactions with solid surfaces or soluble components, reactions with solid microscale entities, reactions with nanoscale entities, and those influenced by pharmacological agents.

response. The tissues of the body are aqueous-based, and have a collection of species, both cellular and molecular, that are mobile and aggressive, so that the already corrosive environment is powerfully enriched by these active agents. Aggressive host-biomaterial interaction mechanisms are readily available as soon as the physician or technologist exposes the former to the latter. A few general principles have to be borne in mind. The first concerns location; the consequences of the interactions between material and host may relate to the vicinity of the material, giving the local foreign body response. Alternatively, or additionally, the effects may be remote, affecting either the whole body (a systemic response), or affecting a specific discrete remote site, for example, the site where a corrosion or degradation product is eventually stored. Local foreign body responses are important with implanted devices. Systemic effects can occur with any biomaterial but have taken on greater significance as molecular or nanoscale biomaterials are used for drug and gene delivery or imaging contrast agents, usually delivered by injection. In addition, and especially with respect to tissue engineering processes, interactions may take place within in vitro bioreactor or microfluidics systems. Second, the mechanisms of biocompatibility should not be expected to show linear progression with time. In many situations, usually on the side of the host, one event may be triggered spontaneously, at any time, the effect of which can be powerfully amplified by one or more mechanisms, changing the whole nature of the response in a short space of time; it is often sensible to consider biocompatibility in terms of metastable systems. Third, as alluded to earlier, although biocompatibility is obviously controlled by the nature of the material, device or agent, it is clearly influenced by many other factors. Biocompatibility phenomena vary from patient to patient, and may vary with the techniques used to administer the biomaterial to the patient. Conceptually, it is best to consider the totality of biocompatibility by identifying those physiological processes that may be perturbed by the presence of biomaterials, and describing the mechanisms by which these perturbations may take place. For example, how do cells that normally interact with a specialized extracellular matrix respond to the surface of a nonphysiological biomaterial? How does a plasma protein respond to a metallic intravascular stent when it normally interacts with an intact endothelium? It has been common practice to separate out different materials, different applications and different phenomena. Thus, interactions with blood have been considered separately from interactions with solid tissues; bone and soft tissue biocompatibility have usually been considered separately and the biocompatibility of implant materials has been characterized separately from that of materials used in tissue engineering scaffolds, drug delivery, and imaging systems. Here, another approach is taken, based on the mechanisms by which physiological processes are perturbed by the presence of biomaterials. The biocompatibility paradigm presented here originates with the hypothesis that the biomaterial is a solid object that is immobile, chemically nonreactive with physiological components, and unchanging with time; mechanisms may then be added to this basic situation as more realistic characteristics are considered. In this model, the tissue is generic; it consists, to varying extents and with varying individual characteristics, of cells and their extracellular matrix. It could be solid tissue, blood, a collection of cells, or a tissue-engineering construct. In

3. MECHANOTRANSDUCTION: ROLE OF MECHANICAL STIMULI IN BIOCOMPATIBILITY Mechanotransduction is the collective term that describes the molecular and cellular processes that are involved with the conversion of mechanical stimuli into biochemical signals. These phenomena have been receiving increasing attention within biological sciences in recent years,11 having dominant roles in determining cell shape, proliferation, migration, apoptosis and other parameters, such that developmental biology,12 stem cell lineage specification,13 cancer biology,14 disease progression,15 and regenerative medicine16 are all powerfully controlled by mechanical stimulating events. As described by Iskratsch et al.,11 one crucial aspect of organ formation, tissue repair, regeneration and aging is the dynamic interaction between cells and their microenvironment and the forces that are applied, where the same biochemical components will have differing effects on cells when the mechanical system is altered. Myosin motors that exert forces on actin filaments anchored to cell−cell or cell−matrix adhesions, and mechanosensors that are responsive to counter forces from matrices and other cells, play important roles here. The actomyosin contractility mechanosensors exist in a quasisteady state of tension, allowing cells to continuously determine organ shape. When forces are applied, or more importantly when forces are changed, mechanotransduction pathways, involving sensing and signaling processes, lead to changes in gene and protein expression profiles. Most signaling pathways eventually culminate with nuclear proteins binding to specific genomic elements that modulate transcription.17 The time scale for these events may be milliseconds/seconds for the stretching of mechanosensors, hours for altered gene expression, and days or weeks for altered cell function and tissue development. It would seem intuitively obvious that because all biomaterials applications involve the perturbation of mechanical environments and, quite often, the deliberate application of forces that are unlikely to be of normal physiological character, mechanotransduction pathways should play a prominent role in biomaterial−host interactions. Indeed it is proposed here that mechanotransduction is the primary, inescapable, baseline phenomenon in biocompatibility. It is emphasized that there are several parameters that define a mechanical environment, including stress, strain, strain rate, and strain energy, and that it is unclear exactly which metrics control mechanobiological effects;18 these issues have to be resolved in due course but do not negate the argument presented here. 4


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Figure 1. Schematic representation of responses of endothelial cells to laminar shear stress. Shear stress-induced mechanosensing and intracellular signaling lead to the modulation of gene expression and cellular functions. Reproduced with permission from ref 46. Copyright 2005 Elsevier.

3.1. Musculoskeletal Tissues. Reference to a series of diverse biomaterials applications indicates just how comprehensive and widespread are mechanotransduction effects. The musculoskeletal system provides a good basis for discussion. Placing implantable devices within bones and joints has for long been associated with perturbations of stress fields and alterations in bone structure, often referred to as stress shielding,19 although largely described phenomenologically rather than mechanistically. Far more information has been recently obtained with reference to bone remodeling and regeneration in the context of disease progression and tissue engineering.20 Several types of cell, stress systems and specific signaling pathways are involved in bone mechanotransduction. Osteoblasts, osteoclasts, and osteocytes are clearly important, but bone lining cells, mesenchymal stem cells (MSCs), macrophages and lymphocytes are also affected. There is interplay between structural stresses (tensile and compressive) and fluid shear stresses, particularly within interstitial bone fluid. Of considerable relevance is the fact that many molecules involved in mechanotransduction signaling pathways may also be involved in responses of bone to biomaterials. For example, the Wnt/β-catenin pathway is primarily involved with the mechanosensitory function of osteocytes,21 and is also associated with the host response to synthetic hydroxyapatite, β-catenin being up-regulated in peri-implant bone cells.22 The elastic modulus of tissue engineering scaffolds regulates regenerative responses through Wnt/β-catenin signaling.23 In the context of Wnt/β-catenin, sclerostin, a Wnt inhibitor, is down-regulated by mechanical forces.24,25 At the same time, hyaluronan can improve bone regeneration through hostbiomaterial interactions that involve selective binding of sclerostin and enhanced osteoblast function,26 while sclerostin antibody treatment can improve implant fixation under some circumstances.27 There has been much speculation of the role of microRNAs in mechanotransduction in bone,20 which is reasonable in view of their involvement in post-transcriptional regulation of skeletal development;28 the evidence so far is sparse but mechanical force-induced specific microRNA expression has been observed in human periodontal ligament stem cells.29 There is, however, extensive evidence, albeit

mostly in vitro, of the regulation of microRNA expression by biomaterials surfaces.30,31 The clinical emphasis on therapies for bone repair and regeneration in recent years has resulted in a reappraisal of the regulation of bone remodeling,32−34 many studies emphasizing the critical role of mechanical stress.35,36 Just as mechanotransduction is now being considered as a factor in the pathogenesis of osteoarthritis37 and bone cancer,38 it is inevitable that mechanotransduction is a critical factor in the development of the response of bone or bone cells to biomaterials, either in the form of implantable devices or in the context of bone tissue engineering. Still within the musculoskeletal system, the biocompatibility of dental implants is dependent on applied mechanical stresses. A number of recent studies emphasize the multiscale nature of the mechanisms of the tissue response, specifically involving the interaction between topographical and biomechanical factors that control the development of osseointegration,39,40 which defines the manner in which bone and materials such as titanium develop functional attachment. Mechanical forces are also known to significantly influence tendon healing.41 The use of hydrogels to facilitate the repair of cartilage defects strongly depends on the mechanical interactions between the hydrogels and chondrocytes, especially on the dynamic nature of force application.42 The situation is complex because there are variable, often contrary, effects of dynamic loading on proteoglycan and collagen synthesis. 3.2. Cardiovascular System. Biomaterials interact with the cardiovascular system under many circumstances, but the most prominent with respect to biocompatibility relate to vascular grafts, intravascular stents and prosthetic heart valves, where endothelial responses and thrombogenicity are the main features. Vascular endothelial cells control many homeostatic functions in normal healthy blood vessels and in response to chemical and mechanical stimuli. They influence vascular remodeling through the production of both growth promoting and inhibiting substances; they also modulate hemostasis and thrombosis through the secretion of procoagulant, anticoagulant, and fibrinolytic molecules, mediate inflammatory responses by chemotactic agents, chemokines, and cytokines, and regulate vascular smooth muscle cells via vasodilators and 5


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ACS Biomaterials Science & Engineering vasoconstrictors. As reviewed by Chu and Chien,43 hemodynamic forces are essential for these physiological functions under normal conditions, but perturbation to these forces can induce endothelial dysfunction by adversely modulating the cell signaling and gene expression patterns. With respect to the biocompatibility of devices used within blood vessels, interventional procedures themselves will cause such perturbations and inevitably will be associated with vascular changes. As discussed by Goel et al., postinterventional changes such as restenosis in stents represent the sum of two processes, intimal hyperplasia, the thickening of the tunica intima, and arterial remodeling, resulting in changes to the overall vessel morphology,44 and these take place in the absence of prolonged contact with a device. It is abundantly clear that disturbed flow patterns, which include recirculation eddies, flow separation, and reciprocating flow, and which have spatiotemporal variations, strongly influence the activation of a number of atherogenic genes. In straight segments of arteries, there is usually laminar flow and high unidirectional shear stresses with down-regulation of genes such as monocyte chemotactic protein-1 and platelet-derived growth factors.45 However, in disturbed flow there may be low and reciprocating shear stresses that induce a sustained activation of these genes, shown in Figure 1.46 Compared to laminar flow, the perturbed shear stresses result in higher cell turnover, DNA synthesis, adhesion molecule expression, sustained oxidative stress, vascular smooth muscle cell activation, and the promotion of both atherosclerosis and thrombosis.43 As with bone mechanotransduction, microRNAs also have roles in flow-dependent vascular remodeling, where, for example, a group of microRNAs such as miR-10a, mir-19a, and miR-23b are induced by high shear stress and have an atheroprotective effect.47 With prosthetic vascular grafts in humans, it is well-known that the grafts remain largely without an endothelium,48 the majority of the graft length being covered with compacted fibrin. This arises because there is very limited passage of any cells through the vessel walls. Although this may not be a significant problem with large diameter grafts, it is a major limitation to the performance (i.e., clinical patency) of smalland medium-sized grafts due to thickening of the fibrin layer and to the intimal hyperplasia in the region of anastomoses, especially the distal anastomoses. It may well be that the mechanical properties of the grafts influence fibrovascular infiltration along the graft length because of compliance mismatch,49 although there is limited evidence for this. On the other hand, it is very clear that flow disturbances at anastomoses, for example, distal end-to-side anastomoses,50 strongly influence the hyperplasia seen at these locations.51 It also appears that mechanotransduction is a limiting factor in the performance of saphenous vein grafts in coronary bypass since the mechanosensitive transcriptional factor Egr-1 has been shown to regulate insulin-like growth factor-1 receptor expression in vascular smooth muscle cells, contributing to neointima formation.52 In a similar manner, the use of synthetic arteriovenous grafts in long-term hemodialysis vascular access, which have a primary patency rate at 2 years of only 25%, is adversely affected by disturbances in hemodynamic shear stresses within the graft; neointimal hyperplasia is frequently seen at the graft venous anastomosis in association with aberrant wall shear stresses on either side of the junction.53 With coronary interventions involving stents, two separate but inter-related mechanical phenomena, involving both fluid

and solid mechanics, are experienced. Immediately after stent deployment, struts compress the vessel wall and cause flow disturbances at the intima.54 Compression of the endothelium leads to immediate endothelial denudation and subintimal hemorrhage, both causing inflammation and the onset of proliferative processes, including vascular smooth muscle cell proliferation, ECM formation and intimal hyperplasia.55 These inflammation-induced effects are simultaneously influenced by the flow disturbances associated with the interfacial geometry determined by the stent strut patterns embedded in the endothelium. Koskinas et al. have given detailed consideration to the combined influences of these two mechanical effects with respect to in-stent restenosis and thrombosis:56 it is clear that mechanotransduction is a profound determinant of the biocompatibility of stents. It is also relevant to point out the basic role of hemodynamics in thrombus formation, and especially platelet activation and aggregation.57,58 The situation is complex because there are at least three distinct shear-dependent platelet aggregation mechanisms, involving processes mediated by integrin αIIbβ3 under low shear, by VWF − GPIb−integrin αIIbβ3 at high shear and solely by VWF − GPIb adhesion without the involvement of integrins at pathologically high shear. During thrombus development, flow separation, temporal shear gradients and turbulent flow become more prevalent with hemodynamic perturbations and the degree of perturbation correlates directly with the magnitude of platelet aggregation.59 Although mechanisms are different, red blood cells are also affected by hemodynamic conditions; upon deformation, the cells release chemicals such as ATP, which participate in vascular signaling.60 Hemolysis is a mechanically induced component of the biocompatibility of devices such as mechanical heart valves. 3.3. Tissue Engineering Substrates. Mechanotransduction profoundly affects the behavior of stem cells, both under natural circumstances and within tissue engineering systems, for example in in vitro bioreactors. The force-dependent cell signaling processes in stem cell differentiation have been reviewed by Yim and Sheetz,61 with special emphasis on focal adhesions, mechanosensitive ion channels, cytoskeletal contractivity, Rho GTPase signaling, calcium signaling and nuclear regulation. There are many individual components of the various pathways in these systems that are clearly forcedependent, including the binding of vinculin to talin during initial stages of focal adhesion assembly62 and the activation of RhoA and Cdc42 in neurogenesis in neural stem cells.63 Dealing first with in vitro bioreactor-based tissue engineering, two separate types of mechanical cue influence stem cell behavior, only one of which is concerned with a biomaterial property. These aspects have been discussed by Steward and Kelly with respect to the mechanical regulation of MSC differentiation.64 The first type refers to the shear stress system imposed by the mechanics of the bioreactor, which, as described by Yeatts et al.,65 include spinner flasks, rotating wall bioreactors and perfusion bioreactors. Each of these provides different stress systems and dynamic variations in shear stresses. A primary shear stress driven signaling pathway in the differentiation of MSCs in both osteogenesis and chrondrogenesis is that of mitogen activated protein kinases (MAPKs). Mechanical stresses are involved in pathway activation and in the up-regulation of the proteins on which the pathways depend, shown in Figure 2. Although the physical characteristics of any biomaterial scaffold or template, including 6


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effects do vary between adult and neonatal cell sources and with other relevant variables. The potential role of biomaterials as stem cell regulators has been extensively analyzed by Murphy et al.13 It is relevant to repeat here a major part of their conclusions: “Although there are many mechanisms at play at the cell/material interface, the fundamental interaction that all cells must have is a link between the cytoskeleton and the material. The consequences of this interaction include a cascade of events in the cell, all of which are initiated by the cytoskeleton or by structures that link it to the material...the cytoskeletal protein actin and its molecular motor myosin II bind and slide past one another to contract the cell. This mechanism is highly organized in muscle, yet it is present in all adherent cell types and in stem cells it enables them to “feel” the stiffness and topography of the environment, as well as to control their size, shape and polarity. Although such inherent properties of the material may seem disparate, they are united by a common contractilitybased mechanism that directs stem cells towards specific lineages based on the degree of activation”. 3.4. Nanoparticle Internalization. If the strategy to develop an overarching theory of biocompatibility discussed in section 2.1 is valid, it has to accommodate the evidence associated with the fate of injected nanoparticles, for example imaging contrast agents. There are significant toxicological concerns about such nanoparticles,71 with much discussion in the literature about mechanisms such as reactive oxygen species (ROS) generation. Rarely have such discussions involved consideration of the mechanical characteristics of the nanoparticles, but data now indicate that the dynamics of the physical interactions between nanoparticles and cell membranes, involving deformation of the latter, are important during the internalization of particles, a critical event in the overall toxicological pathway. With reference to drug-containing nanoparticles, Li et al. recognized that current views on nanoparticle uptake in cells suggest that the properties of nanoparticles that influence internalization are size, shape, surface chemistry, and ligand arrangement, but they went on to consider the potential influence of nanoparticle hardness.72 Using particle dynamics simulation, they showed that rigid nanoparticles can enter cells by endocytosis, whereas for soft nanoparticles such as liposomes, the endocytosis process can be frustrated due to wrapping-induced shape deformation and nonuniform ligand distribution. Gonzalez-Rodriguez and Barakat have also used computational modeling to analyze the receptor-mediated nanoparticle internalization into endothelial cells, in which the kinetics of the internalization process, the dynamics of binding, and the stress distribution between nanoparticle and cell membrane were computed.73 They demonstrated the existence of an optimal radius, of around 50 nm, for receptormediated nanoparticle internalization. Below this radiu,s the process becomes rapidly impaired by membrane bending rigidity, whereas above this, cytoplasmic rigidity reduces the efficiency of internalization. They also showed that internalization is most efficient at a receptor-particle bond elastic constant of around 5 nM/μm. 3.5. Case for Mechanotransduction. The four areas discussed in the previous sections provide good evidence of the role, or potential role, of mechanotransduction in biocompatibility phenomena. Although these are quite diverse examples, they do not conclusively prove the argument proposed in this paper that mechanotransduction is the primary, frontline

Figure 2. Influence of culture conditions in a perfusion system on stem cell signaling. Shear stress and controlled oxygen tension provide stimulus to cells growing on 3D scaffolds, which influences HIF (hypoxia inducible factor) and MAPK (mitogen-activated protein kinase pathways). Reproduced with permission from ref 65. Copyright 2013 Elsevier.

porosity, have some influence on fluid flow, they are not the primary determinant of the shear stresses that impact on the cells. The second type of mechanical cue is that of structural stresses, perhaps best seen in cell-seeded scaffolds in static culture where hydrostatic pressure results in stress transfer between biomaterial surfaces and cell membranes. The precise nature of the stresses at these interfaces, including magnitude and type (especially tensile or compressive), has a strong influence on the gene expression of the cells and the differentiation pathway down which they are directed.66 The mechanisms here are likely to reflect the normal processes of stem cell−matrix interactions within the microenvironment of cell niche67 and the material property most likely to influence the cell fate is substrate stiffness, or elasticity. In particular, MSCs clearly respond to 3D hydrogel stiffness, being modulated by integrin binding through reorganization of ligand presentation at the nanoscale;68 matrices of 11−30 kPa stiffness induce MSC osteogenic differentiation whereas those of 2.5−5 kPa show adipogenesis. At this stage, there is a lack of consistency in the details of the causal relationship between stiffness and cell fate when considering all types of cell and all practical conditions, largely because of the interactivity between different mechanisms, but it is clear that mechanotransduction is a primary controlling factor in the phenomena of biomaterial−bioreactor induced stem cell differentiation. The situation is similar to in vivo tissue engineering, where much evidence points to a role of mechanical stress in tissue regeneration associated with injectable scaffolds. Myocardial tissue engineering provides a good example. The disparity between the stiffness of myocardium and injectable hydrogels, and the importance of associated stress fields, has been addressed by Reis et al.69 When cardiovascular progenitor cells are contained in cardiac ECM − fibrin hybrid scaffolds, their differentiation is affected by the stiffness as well as the composition of the hydrogel.70 For example, VWF gene expression is up-regulated with increasing gel stiffness, although it is fair to say that such 7


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mediated by the nature of the adsorbed layer, which is likely to vary over time. 4.2. Protein Behavior at Biomaterials Surfaces. Haynes and Norde published an authoritative review of the behavior of globular proteins at solid−liquid interfaces in 199487 and their analysis has formed the basis for much subsequent work, including an important discussion on the role of adsorbed proteins on cell behavior by Wilson et al. in 2005.88 According to these reviews, adsorption behavior is predicated on four processes, the structural rearrangement of the protein molecule, the dehydration of the material surface, the redistribution of charged proteins at the interface and protein surface polarity. Although there is some connectivity between these factors, there is clearly no single force or effect that dominates protein adsorption, which explains the complexity of the overall phenomenon, and the relative lack of uniformity of facts and opinions on the consequences of the adsorption. One of the problems with identifying these consequences is the difference in behavior when materials are exposed to single protein solutions and to complex physiological fluids such as blood. In single solutions, the amount of adsorption tends to increase with time and protein concentration up to the point when a monolayer is established. The rate of adsorption then tends to decrease on the basis of the number of available binding sites. In more complex biological fluids, including serum, plasma, and blood, competitive adsorption characteristics on surfaces that generally have limited number of binding sites depends on relative concentrations and surface affinities, as first discussed by Vroman.89 It was initially found that fibrinogen adsorbed to materials from blood with an increasing surface concentration up to a maximum at an intermediate contact time, but is then replaced by other plasma proteins; this is seen with many other proteins, details of which depend on surface charge and specific surface chemistry. Although this occurs with both hydrophilic and hydrophobic surfaces, and although it may appear that more proteins adsorb to hydrophilic materials but the stronger binding occurs with hydrophobic materials, no universality can be seen with this phenomenon.90 The competitive adsorption process is generally discussed in terms of what is called “the Vroman effect”;91 the displacement of one adsorbed protein (A) by a second protein (B) is considered to take place by a process whereby B embeds itself in A, the resulting transient complex forming a taller and less dense structure exposing A to the solution, when it then desorbs.92 Over time, the composition of the adsorbed layer changes as faster-diffusing molecules such as albumin are displaced by proteins with a higher affinity for the surface, such as vitronectin. The variation in adsorption with the wettability of the surface is an important, but confusing, issue. The events are related to the structure of water at the interface93 where there is a relatively less dense water region against hydrophobic surfaces with an open hydrogen-bonded network and relatively more dense water region against hydrophilic surfaces with a collapsed hydrogen bonded network. The interphase region in which water structure varies (including degrees of dehydration) can extend tens of nanometers from the solid surface and should control subsequent events. Hydrophobic surfaces tend to support adsorption of proteins because expulsion of solute from solution into this region is energetically favorable. Hydrophilic surfaces tend not to support adsorption because this mechanism is energetically unfavorable. These comments, however, only reflect general trends, and are not universally applicable. It may be, as suggested by Vogler,93 that

pathway. Supporting this argument, however, is the fact that there are many other areas of biology/clinical medicine in which, in recent years, mechanobiology has been gaining ascendancy as a controlling factor in important phenomena. For example, as Tyler has recently discussed, an understanding of the mechanobiology of brain function is critical.74 On the basis of this knowledge, the influence of glial cell mechanosensitivity on foreign body reactions to biomaterials in the central nervous system has been identified,75 as has the influence of cell−substrate mechanical interactions on neural cells in neural tissue engineering processes.76 The present paper is not alone in drawing attention to the primacy of mechanical aspects of biocompatibility, as indicated by the review of Mazza and Ehret.77 The field is far too young to identify all key mechanotransduction pathways that are involved in biocompatibility. Moreover, it is not yet possible to insert extensive quantitative data into these pathways in order to demonstrate how the pathways compare when discussing different anatomical and clinical systems. It is not necessary, of course, to claim that there are biomaterial-specific mechanotransduction pathways; it is relevant, however, to note that recent discoveries of pathways in mechanotransduction in general are already being implicated in biocompatibility-related mechanotransduction phenomena. One of these is the Hippo pathway, shown to be involved in the promotion of cell death and differentiation and the inhibition of cell proliferation, which is now recognized as a strong factor in mechanotransduction78 and has been implicated in the control of cardiac progenitor cell fate through the dynamic sensing of substrate mechanics.79 Naturally the focal points of mechanistic discussions are the mechanical properties of actin networks, which determine the dynamics of cell stiffness and transmit forces during cytokinesis, cell motility, and cell shape changes.80

4. INTERFACIAL PHENOMENA IN BIOCOMPATIBILITY 4.1. Potential Significance on the Biomaterial−Host Interface. It was noted earlier that mechanical forces are experienced within the biomaterial−host system as soon as a material is placed within or on the relevant tissue components. Equally rapidly, molecules from those tissue components will be attracted to the materials surface through the influence of interface energetics. This has been recognized for decades, and early commentaries on biocompatibility focused on these initial events, especially with respect to protein adsorption on surfaces.81−84 Because many proteins are highly surface active, they have a strong affinity for some surfaces, resulting in a rapid initial adsorption phase,85 concentrating them in far higher levels than found in solutions.86 Several questions arise concerning the relevance of this process and its variability with different biomaterials. It is inevitable that from the first few seconds, a solid macroscale biomaterial will not directly contact any tissue components apart from the adsorbed proteins. There are two important consequences. First, while the cumulative effect of establishing noncovalent bonds between the initially arriving proteins may not be of great significance in relation to the behavior of macromolecules in the tissues in general, it is possible for the adsorbed layer to have a powerful effect on those proteins that are involved with the initiation of cascade processes, with special relevance to the blood coagulation cascade and the complement activation cascade. Second, the interaction between cells (of hard or soft tissue and blood within in vivo environments, or in ex vivo culture) and biomaterials is 8


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Jones et al. found that although it appears intuitive that there should be a link between lens comfort, as a surrogate for lens biocompatibility, and surface deposition of proteins and other molecules, this has been remarkably difficult to demonstrate in clinical studies.101 It has to be concluded that although it is inevitable that proteins interact with biomaterial surfaces, and these interactions may have some indirect influence on biocompatibility and device performance, there is little evidence that protein adsorption is a reproducible universal mediator of biocompatibility performance. As alluded to above, there are some exceptions to this general position, largely derived from observations on the performance of long-term implantable devices. These are now discussed in relation to the activation of certain cascade processes, the potential influence of adsorbed proteins on specific cellular behavior, and the special case of protein adsorption on nanoparticles. 4.3. Cascade Processes. It is necessary here to consider whether interfacial reactions are able to trigger or activate cascade processes within host tissues. Two important possibilities are the clotting cascade and complement activation. 4.3.1. Blood Coagulation on Foreign Surfaces. It has been known for a very long time that blood tends to clot on contact with foreign surfaces, some early detailed discussions of the phenomenon being given by Lozner and Taylor102 and Margolis.103 It was postulated that a different pathway to clotting than that seen in normal physiological situations was in play with foreign surface induced thrombogenicity, being referred to as the intrinsic pathway to differentiate it from the normal extrinsic pathway. Ratnoff identified a key molecule in the initiation of this intrinsic pathway, being termed factor XII, or the Hageman Factor.104 For many years, as recently described by Markiewski et al.,105 the conventional view was that a proteolytic cascade system is activated on adsorption of plasma proteins on the foreign surface and the conformational changes that then take place. The binding of factor XII to the surface triggers contact activation, which generates factor XIIa, and sequentially other factors, including XI, with HMWK (high molecular weight kininogen) acting as one of the cofactors. The intrinsic and extrinsic pathways, the latter usually being assumed to be activated by the expression of tissue factor (TF) on cells, converge on a common pathway at factor X, which proceeds via prothrombin to thrombin, fibrinogen, and fibrin. It was also assumed that different materials had different abilities to activate this process because it was a common experience that blood in paraffin-coated tubes took much longer to clot than blood in glass. The general consensus was that clotting was favored by anionic surfaces. However, in spite of very many attempts to classify materials in terms of their thrombogenicity, and to identify those characteristics that promote the process, very little success has been achieved and questions have been asked about the validity of adsorption-driven mechanisms. Sefton et al. in 2001 attempted to develop surface modified polymers to give improved blood compatibility on the basis of assumptions of the putative influence of surface properties, but found many of the modifications made the material worse not better.106 Sefton and colleagues returned to this question a few years later.107 Some critical observations were that although factor XII has been observed on some blood-contacting medical devices, it is not found there in activated form; only minute amounts of thrombin or thrombin-antithrombin III complex (TAT) are generated when biomaterials are incubated with undiluted

inappropriate parameters are used to characterize wettability so that true correlations will not emerge; he prefers water adhesion tension, τo, to the more normally quoted surface free energy or Zisman’s critical surface tension. The relationship will also be influenced by many other factors. The charge on the surface should play a role; however, this charge is likely to be shielded by hydrating water94 and may be counterbalanced by small ions in the interphase region, and is influenced by the changing dielectric constant in the structured water. This is potentially important with respect to the consequences of the adsorption since there may be discrimination between biologically important ions that preferentially solubilize divalent ions in more-dense water regions relative to less-dense water regions, where there is enrichment of monovalent ions.93 The entropy of the system may be increased by conformational changes in the adsorbed proteins, which may allow strengthening of protein attachment over time and influence desorption behavior. Thus, when wettability or charge suggest that protein adsorption to a surface should be minimal, conformational change may actually drive the adsorption process. Many reviews of the development of the foreign body response to biomaterials, and specifically to implanted devices, indicate that protein adsorption is the first event to take place and processes of adsorption, desorption and conformational change play significant roles in subsequent events, including inflammation, the immune response and fibrosis. With certain exceptions alluded to above and explained below, the reality is that after decades of experimental work and clinical evaluation, there is very little evidence that this is universally the case, and very little evidence that ultimate biocompatibility-related clinical outcomes are directly dependent on the surface properties of the materials. The complexity of these processes and the multiple parameters that influence the interfacial reactions mask our understanding of the mechanisms that are involved. It is possible that some indirect protein effects may play some role in overall clinical performance. For example there is some evidence that boundary lubrication in joint replacements may be influenced by protein adsorption, affecting wear rates and prosthesis loosening.95 It has been known for several decades that proteins, adsorbed or in solution, can influence the corrosion of metals,96 including the only intentional degradable implantable metal magnesium;97 this implies that the overall biocompatibility will be affected, although the clinical relevance has not been demonstrated. The same situation arises with metallic electrodes, where interactions inevitably take place between surface ions (e.g., gold) and proteins such as albumin under the influence of applied potential, but the interactions are complex, with uncertain consequences.98 Contact lenses provide an interesting model for protein adsorption and clinical relevance. Ionic, high-water poly(hydroxyethyl methacrylate) (pHEMA) lenses attract a large amount of tear film protein, especially the positively charged lysozyme, compared to nonionic pHEMA, whereas silicone hydrogels attract even less. However, the percentage of denatured protein on the silicone is higher, and silicone hydrogels have not reduced the prevalence of ocular complications such as inflammation and microbial keratitis.99 The silicone hydrogel can be coated with poly(ethylene oxide), which is highly protein resistant in vitro, but this property does not appear to be replicated in vivo, and clinical performance is not improved.100 In a major review of contact lens discomfort, 9


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mediated coagulation pathway is the main, if not exclusive, pathway that initiates fibrin formation in vivo. The observations discussed in the above paragraphs suggest that although it is an incontrovertible fact that blood tends to clot on contact with foreign objects, both in vitro and in vivo, the role of the actual material surface is far from clear and indeed the conventional view that plasma protein adsorption on the surface is the controlling process in biomaterials−related thrombogenicity has to be questioned. There have been many attempts to reduce thrombogenicity risk with surface modifications intended to minimize or eliminate the adsorption; these often show promise under in vitro conditions but rarely do they make any difference in terms of clinical outcomes.115 With medical devices, patients with prosthetic mechanical heart valves are certainly at risk of thromboembolic events, which has to be managed by antithrombotic therapy, but the main risk factors are associated with altered blood flow and hemostatic activation caused by vessel wall disruption and not the nature of the biomaterials that are used.116,117 Similarly, ventricular assist devices and artificial hearts, now an acceptable treatment option for end-stage heart failure, have a number of potential causes of complications but thromboembolic events are well-managed and relate to hemodynamics and not material characteristics.118,119 Extracorporeal circulation, for example in hemodialysis and membrane oxygenation, is associated with risks of hemostatic complications, with both clotting and bleeding needing management.120 Obviously the hemodynamic conditions are very unphysiological in these circuits and are the fundamental causes of disturbed hemostasis. Central venous catheters are very widely used, especially for the infusion of drugs into cancer patients, and are associated with several complications, including catheter-related thrombosis.121,122 The risk factors include those that are insertionrelated, catheter-related and patient-related, the most important being the damage to the vessel wall on insertion, the location of the catheter tip and the type of catheter, with peripherally inserted central catheters having higher rates of thrombus. The stiffness of the polymer is relevant as stiff catheters are more likely to damage the endothelium, but polymer surfaces themselves are rarely considered risk factors. Infection is also common and thrombosis and infection often occur together. There have been attempts to reduce both phenomena by surface coatings; one recent report suggests that a nonleaching poly sulfobetaine coating may be successful with this objective, although the mechanisms are not clear and the results have yet to be confirmed in clinical studies.123 Thrombus formation has become a significant and controversial issue with coronary stents. The initial high rates of thrombosis seen with bare metal stents was markedly reduced through the use of antiplatelet agents but appeared to increase with the introduction of drug-eluting stents, with a significant number of patients presenting with thrombosis several years after implantation.124,125 It has become clear that stent thrombosis has very little to do with the stent materials or any interfacial phenomena, but is due to a variety of procedurerelated issues that result in inflammation of the vessel wall, particularly stent under-expansion and malapposition.126,127 As noted later, it is obviously difficult to establish, by direct observation in the clinical setting, the chronology and mechanisms involved in the short period of protein adsorption on medical device surfaces. The experiences with those medical devices discussed above that come into significant contact with

plasma alone; higher levels of adsorbed kallikrein and factor XII do not correlate with TAT formation. The inescapable conclusion is the contact phase proteins, by themselves, play little role in the activation of coagulation. The question arises now as to whether, and if so, how, the biocompatibility of biomaterials influences thrombogenicity through plasma protein adsorption and activation. The work of Sefton referred to above107 addressed the broader implication through a consideration of the role of blood cells, particularly platelets and leukocytes, and complement (discussed below) in thrombus formation. They draw a very useful analogy by considering thrombogenicity as a special case of inflammation, also discussed later. Inflammation is characterized as a leukocytic response to a stimulus and may itself contribute to thrombin generation through expression of TF on monocytes, the activation of platelets by released inflammatory mediators and the blocking of inhibitors of coagulation. Foreign-surface thrombosis (and indeed thrombosis in general) may then be viewed as a multiprotein and multicellular process rather than simply being involved with activation of a cascade and of platelets.108 With reference to contact activation, Vogler and Siedlecki also concluded that existing paradigms based on the biophysics of contact activation do not provide explanations of experimental and clinical observations,109 that surface activation of factor XII to XIIa is not specific to anionic/hydrophilic surfaces, and very importantly, this autoactivation is moderated by proteins in the solution phase. This hypothesis is consistent with the position that proteins rapidly diffuse from concentrated solution and inflate the interphase, which then undergoes a slow decrease in volume due to the efflux of interfacial water, which in turn increases interphase protein concentration. Nie et al. examined the parameters of protein adsorption, clotting times, platelet adhesion and activation and complement activation on silicon surfaces modified by a range of single molecule layers of different charge and chemistry, whereas hydrophilic surfaces in general had lower protein adsorption and platelet activation, the results were very variable, with some surfaces of extremely high protein and platelet affinity having little platelet activation and TAT generation.110 Returning to the questions of multiparameter processes, Sperling et al. found that neither platelet adhesion on hydrophobic surfaces without concurrent contact activation, nor contact activation on negatively charged surface without the presence of activated platelets led to significant thrombus.111 However, the presence of few platelets was enough to propagate coagulation significantly if contact activation was simultaneously initiated; they proposed the coexistence of different activation processes that leads to an on−off switch for coagulation activation on biomaterials as opposed to the direct scaling of platelet adhesion and contact activation with certain physicochemical surface properties. Long et al. have also compared coagulation with inflammation and have questioned the role of factor XII in apparent contact activation, noting that factor XII deficient patients are not physiologically associated with an increased bleeding risk.112 This latter point was raised by Maas et al.113 and Schmaier,114 who query the physiological role of factor XII, the former noting that activation may occur through the activity of mis-folded proteins, not necessarily at the point of foreign surface contact. This is consistent with the hypothesis of Vogler109 concerning the interphase region. Long et al.112 go further to speculate that this absence of bleeding in factor XII deficient states leads to the conclusion that TF 10


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ACS Biomaterials Science & Engineering blood show that, provided clinical techniques are sufficiently good to avoid undesirable tissue trauma and appropriate anticoagulation regimes are used in order to counteract the effects of disturbed blood flow, biomaterial surface-induced thrombogenicity is not a major clinical problem. Although recognizing the controversial nature of associated mechanisms, it is suggested here that this is not a significant, fundamental biocompatibility pathway. 4.3.2. Complement Activation by Foreign Surfaces. Brief mention must be made of complement at this stage. The complement system consists of more than 20 plasma proteins; these function as either enzymes or binding proteins, and the activation of this system is instrumental in the body’s defense mechanisms against infection. The significance is that one of the mechanisms by which the system is activated, the so-called alternative pathway, involves interactions with surfaces.107 These may be fungal and bacterial polysaccharides, lipopolysaccharides, or biomaterials. As with the more generalized classical pathway, the alternative pathway starts with an initial enzyme that catalyzes the formation of the C3 convertase, which in turn generates the C5 convertase and the subsequent assembly of the terminal complement complex. Specifically with the alternative pathway it is the hydrolysis of the internal thioester group of the plasma protein C3 in the fluid phase that initiates the activation. The resulting hydrolyzed C3 binds and activates Factor B, which cleaves another C3 molecule into C3a and C3b, forming the C3bBb convertase. Of critical importance here, C3b needs to bind to a surface for the process to continue. The presence of a carbonyl group in the C3b thioester binding site allows covalent bonding to hydroxyl or amine groups, which suggests that certain polymers or ceramics may be able to facilitate this process; this is, however, slow, and clinical experience indicates that such complement activation is unlikely to be an important mechanism in the host response to medical device. There is one situation where surface-mediated complement activation may be more important and this is when there is a very large contact area where, if the material does have relevant charged groups, the rate of production could be sufficiently high to have clinical consequences. This is especially true for chronic hemodialysis.128 Although extracorporeal circulation is utilized in a several procedures, the contact between blood and synthetic materials in the majority of cases is transient, measured usually over a number of hours and in the majority of cases this occurs only once. In hemodialysis, however, a typical recipient receives treatment for 5 h at a time, three times a week, over many months and possibly years. In most dialysis machines, the surface area of contact is very large, on the order of several meters squared, giving ample opportunity for prolonged blood−material interactions. Not all membrane materials have the same effect on blood, and specifically on complement, and on subsequent clinical consequences. In particular, cellulosic materials were shown to have the greatest effect, primarily because the repeating polysaccharide units of cellulose are similar to bacterial wall polysaccharides. It was appreciated that the alternative pathway was the most significant since this does not require antigen− antibody binding but was dependent on molecular binding to hydroxyl and/or amine groups on the membrane surface. As a result of these observations, the use of membranes moved away from cellulose to modified cellulose, in which the hydroxyl groups were replaced by groups such as acetate, and to totally synthetic polymers.129

Problems with chronic inflammation associated with the socalled dialysis syndrome have not gone away through this change of preferred membrane, however, and there are still several immunologically based risks associated with dialysis that contribute to the high mortality and morbidity rates in chronically dialyzed patients. The most significant complications are chronic inflammatory conditions such as vasculopathy. In addition, uremic toxins cause disruption of several granulocyte and monocyte functions. Complement activation is still considered as the most significant phenomenon here, but in addition, the activation of mononuclear cells, cytokine synthesis and release, and the production of ROS and nitric oxide are all possible. With initiation of the terminal pathway of the alternative complement system, some of the molecules are anaphylatoxins (C3a and C5a), whereas others, such as Ba, C3b, and the terminal complement complex induce cytotoxic effects on monocytes, neutrophils and platelets. Most of these factors are detected in increased amounts in dialysis patients; cellulosic membranes cause a 70% increase in C3 turnover. At the extreme end of the spectrum of immunological responses to hemodialysis is the possibility of anaphylactic or anaphylactoid reactions.130 This possibility has been recognized for several decades and the prevalence may be in the region of 1% of patients. 4.4. Protein Adsorption and Cell Behavior. As mentioned above, conventional wisdom has determined that the initial adsorption of proteins on a biomaterial surface controls all subsequent events and is the main determinant of the characteristics of the foreign body response. Major reviews of this response repeatedly start with this mantra. Wilson et al. state “ the rapid adsorption of proteins from blood effectively translates the structure and composition of the foreign surface into a biological language...it is to this language that the cells respond, contributing to the ultimate outcomes in both implantation and tissue culture situation”.131 Gentleman and Gentleman state “These early interactions with the surface play a fundamental role in determining cell adhesion, differentiation, and ultimate tissue formation at the interface”.132 Wang et al. introduce a review on proteins and bioceramics with statements such as “ proteins from the surrounding body fluids will be spontaneously adsorbed onto their surfaces, and then cellular attachment, proliferation and migration occurs” and “ the protein adsorption behavior plays a vital role during bone tissue regeneration”.133 Any thorough analysis of biocompatibility pathways has to confront this apparently pivotal phenomenon in the biomaterial−host interaction and determine whether it is backed by solid evidence and whether it can be explained mechanistically. Many papers on this subject, while accepting this mantra, also refer to contradictory and confusing data. Wei et al., in discussing the influence of surface wettability on the protein adsorption process and subsequent osteoblast behavior, state that it is difficult to obtain consistent correlations between the surface characteristics and these biological processes, with totally contradictory data on cell attachment and protein adsorption on hydrophilic and hydrophobic surfaces.134 Brevig et al. studied the role of adsorbed proteins in leukocyte adhesion and activation and noted very contradictory evidence in the literature, where adsorbed albumin had been demonstrated to have either significant effects or no effects on leukocyte activation in different experiments.135 11


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but not caused by, the IOL implant into important clinical outcomes. Nibourg et al. did not need to implicate biomaterial surfaces in these mechanisms. It is true that they did state in passing that the IOL materials and design do influence the extent of PCO. One of the citations used to support this statement137 actually specifically states that it is the IOL design and not the material that influences PCO. Other studies support the importance of design rather than materials,138,139 whereas others are equivocal.140 It should be noted, as pointed out by Bozukova et al.,141 the design issue largely concerns the role of the IOL optic acting as a barrier to cell migration rather than being involved in causation of migration and EMT. Evidence concerning surface properties, protein adsorption and cell behavior on IOL surfaces is very confusing. In the paper of Bozukova et al. mentioned above, it is stated on one page that “implants exhibiting hydrophilicity associated with strong cell adhesion”, while on the next page report “biomaterials with a hydrophilic surface are known for effective reduction of protein deposition and cell adhesion”. They do refer to published data that suggest that low aqueous contact angles (75°) have low cell attachment, whereas maximum attachment is seen at intermediate angles. The work of Awasthi et al. confirm this confusing situation142 through the comments: “Comparison of hydrophobic and hydrophilic materials showed that the type might influence PCO development...although it is wellrecognized that a hydrophilic acrylic material is more biocompatible, IOLs made of this material have been shown to support LEC adhesion, migration, and proliferation and thus PCO development compared with an IOL made of PMMA or hydrophobic acrylic materials”. This example has shown that it is possible to explain a biocompatibility phenomenon in an implanted device using well-established biological pathways, in this case associated with intervention-related tissue damage, without invoking protein adsorption mechanisms. This does not imply that adsorption processes are always unimportant and it is necessary to look at a few other clinical scenarios. Osseointegration and Dental Implants. Endosseous dental implants represent another class of highly successful implantable medical devices that has a history of controversy over biocompatibility mechanisms. This is an area that is confused by the positions adopted by implant manufacturers with claims over the putative benefits of their surface features; these will be set aside in this discussion, as will the various definitions of osseointegration and related terms about bioinert and bioactive materials. The scientific team behind the development of osseointegrated dental implants preferred to avoid the majority of phenomenologically based definitions, reiterating that the definition should have a clinical basis, related to the development and maintenance of clinically asymptomatic rigid fixation of materials in bone during functional loading.143 The question is how this functional adaptation to bone is achieved and what are the roles of materials and protein adsorption. The dominant material is titanium, with various types of surface treatment and coatings, including hydroxyapatite. There is considerable uncertainty about the role of surface properties on protein and cell behavior and bone formation; Goriainov et al. discussed these inconsistencies.144 For example, hydrophilic rough Ti surfaces are reported to induce more osteoblast differentiation, although cell proliferation was disturbed. Increasing wettability of surfaces resulted in reduced

This confusing situation may be substantially due to the complexities of the systems involved (as shown in Tables 1−3) and the wide range of potential variables, especially those with spatiotemporal character. The evidence is presented here with respect to phenomena with implanted devices and tissue engineering culture systems. 4.4.1. Behavior of Fully Differentiated Cells Relevant to in Vivo Applications. The issue to be discussed in this section is whether the adsorption of proteins and their subsequent rearrangement influences the cells of the host response and subsequent clinical outcomes of the biomaterials and devices. There is a major barrier to addressing this issue; it is very difficult, if not impossible, to examine directly the sequence of events that take place at the surface of a device under clinical conditions. It is possible to determine certain end points of clinical biocompatibility phenomena, such as loosening of joint prostheses and in-stent restenosis, but very difficult to interrogate in real time the processes that lead up to these events. Posterior Capsular Opacification and Intraocular Lenses. There is one situation where the questions can be examined with some objectivity, and that is with intraocular lenses. These are highly successful implanted devices, but some patients do suffer from one significant complication, which is posterior capsular opacification (PCO). This is a multifactorial physiological process involving the central posterior capsule that has a significant impact on high- and low-contrast acuity and low-contrast sensitivity. It arises from the proliferation, migration and abnormal differentiation of residual lens epithelial cells and fibers in the capsular bag. The relevance here is that the effects are obvious to the patient and can be readily monitored by physicians; moreover the process may take place slowly over a few years, and take place in the absence of significant mechanical forces and in the absence of significant release of chemical substances from the biomaterials. There have been references in the literature to the role of the biomaterials surfaces and protein adsorption in the pathogenesis of PCO; it is worth examining the evidence. Nibourg et al. have recently discussed the pathogenesis and prevention of PCO.136 Their main observations are as follows. Lens epithelial cells (LECs) are normally situated in a single layer on the inner side of the anterior lens capsule. They are mostly removed during cataract surgery but some remain in the capsular bag afterward. The LECs are able to proliferate and migrate to the posterior capsule. Moreover, LECs are able to transdifferentiate, especially to myofibroblasts, which are primarily responsible for the PCO. This epithelial to mesenchymal transformation, EMT, can cause the lens capsule surface to become wrinkled, because the myofibroblasts contain α-smooth muscle actin and therefore have contractile properties; when in the visual axis, these wrinkles give rise to visual disturbances. EMT is triggered within inflammatory responses, induced in this case by the surgery itself, where the damaged ocular tissue releases chemokines. Also the family of TGF-β growth factors has been implicated in the EMT process. A latent form of TGFβ is present in aqueous humor and is activated by trauma. Signaling by TGF-β starts with its binding to serine/threonine kinases on the cell surface, eventually resulting in cell signaling by phosphorylation of Smad proteins. The purpose of this brief summary of mechanisms of EMT and the consequential PCO is to demonstrate that there are well-established signaling pathways that translate ocular tissue damage in the region of, 12


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conformational change, depending to a large extent on their rigidity. FN does undergo such changes; under in vitro conditions, hydrophobic surfaces can cause significant change and unfolding, induced by interfacial hydration.152 Liao et al. have carried out computer simulations of FN adsorption and conformation on hydroxyapatite surfaces, demonstrating variable binding and unfolding through several types of interaction, for example electrostatic interactions and hydrogen bonding between the guanidine group and phosphate groups.153 The significance here is that the conformational changes on FN are now known to be involved with the promotion of osteogenic differentiation of MSCs.154 Conformational remodeling of FN regulates VEGF signaling and angiogenesis.155,156 Two related phenomena then control the development of the interfacial region, concerning cells and the ECM. The adsorbed and modified FN layer is able to influence the initial ECM development arising from the reorganization of the blood clot. As discussed by Rico et al., the surface bound FN is able to induce 3D FN fibrillation,157 so that FN is not confined to the surface. This is consistent with the known mechanisms of fibronectin ECM assembly in developmental biology,158 where matrix assembly is usually initiated by ECM glycoproteins binding to cell surface receptors, such as FN dimers binding to α5β1 integrins. In this case, the relevant cells are osteoblasts or their precursors, which are known to adhere to FN through these integrins, such binding directing osteoblast cell survival, proliferation, bone-specific gene expression and matrix mineralization.159,160 This process is potentially autocatalytic because the osteoblasts, once adherent to surface-bound or matrix FN, are able to secrete their own FN.161 The osteoblasts may originate locally, or be derived from circulating osteogenic cells162 or from bone marrow stem cells. Several factors, both materials- and biology-related, can influence the above process, and some subsidiary mechanisms make contributions. For example, the development of functional osseointegration requires mineralization of the matrix, and it has been suggested that rapid mineralization at the interface is regulated by expression of RANK, RANKL and osteoprotegerin.163 Osteoblast-derived VEGF also regulates differentiation and bone formation during bone repair.164 Surface modifications have been used in attempts to enhance or accelerate bone formation, including topographical guidance of osteoblast recruitment165 and up-regulation of TGFβ/BMP signaling, 166 enhancing ligand density (especially RGD peptides) on surfaces,167,168 precoating with FN169 or VN,170 and incorporating slow release of BMP-2.171 The mechanisms outlined above adequately explain the host response to contemporary dental implants, which, for the vast majority of patients, is entirely adequate. Clinically the performance varies very little with the minor surface variations that have been introduced, but slight variations, for example with nanotopography and hydroxyapatite coatings, are easily explained by these mechanisms. Should implants be made of less corrosion resistant metals than titanium, then an ongoing inflammatory response would ensue, as discussed in a later section, resulting in less bone and more soft tissue. Interestingly, because collagen is involved in matrix formation as well as FN, and because collagen I also binds to cells through integrins, it is the relative roles of FN and collagen that controls how much bone rather than fibrous tissue forms at this interface.

cell numbers and enhanced cell differentiation in vitro, whereas increased wettability resulted in better bone apposition in vivo. Hydrophilic surfaces were shown to adsorb proteins in a more flexible conformation in vitro with improved adhesion and spread of cells, whereas another study showed that protein binding was stronger on hydrophobic surfaces. Many studies have attempted to show the value of various calcium phosphate surfaces. Surmenev et al. had to conclude that the mechanisms of osteogenesis on CaP coatings in vivo was very unclear, with conflicting reports varying from highly effective to no effect and even negative effects.145 The work of Velasco-Ortega et al. demonstrates the scale of the problem;146 they examined four different surface treatments for Ti, noting there were no significant differences in cell responses in a subcutaneous rat model, but claiming that the trends observed suggested a better performance with sandblasted surfaces. Although there are many of the usual comments in the literature that the first phase of the bone−material interaction involves the adsorption of plasma proteins, and that these events control everything else, there is no evidence that the major proteins, albumin, fibrinogen, and the globulins, play any role in bone formation adjacent to biomaterials. The situation is quite different, however, with some glycoproteins, especially fibronectin (FN) and vitronectin (VN), which can be adsorbed onto implant materials very rapidly; there is also some evidence that glycoproteins may adsorb to a greater extent to CaP ceramics than to Ti.147 FN is important because its two forms, the soluble plasma protein and the insoluble cellular form that exists in the ECM, both play a role in the bone response. As reviewed by Chatakun et al., FN promotes cell adhesion and migration and also promotes the osteoblastic differentiation of vascular smooth muscle cells.148 The FN adsorption appears to be correlated to osteoblast adhesion through both actin cytoskeleton formation and morphology related to FN fibrillogenesis. Notwithstanding the practical uncertainties mentioned above, the clinical situation is quite clear. A number of commercial implants, mostly made of Ti but with a variety of surface finishes, including CaP coatings, give very good performance with fairly uniform biological and biomechanical characteristics, where differences are largely associated with clinician and patient variables. It is proposed that the biocompatibility pathways here are as follows. However, careful the clinician is, the trauma exerted on the alveolar bone during surgery initiates an inflammatory response.149 This may be associated with effects on osteocytes through heating effects and mechanical disruption.150 The implant is placed within this incipient inflammatory environment, with a rapid infiltration of inflammatory cells, which release a cocktail of pro-inflammatory cytokines and chemokines. Clot formation rapidly takes place at the interface and blood cells are trapped in this region. It is possible that major plasma proteins are adsorbed on the implant surface but these are irrelevant to subsequent events. Glycoproteins, and especially FN, take the place of the initial interfacial layer. It is possible that leukocytes adhere to, and may be activated by, this layer, as proposed by Brevig et al.135 but this is unlikely to be important in most situations. The influence of the FN is most likely dependent on conformational changes that can take place and the interactions between adsorbed FN molecules and the integrins on cells, especially osteoblasts. As described by Fenoglio et al.,151 adsorbed proteins on solid surfaces have variable resistance to 13


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ACS Biomaterials Science & Engineering Biomaterials in Blood Vessels and Endothelialization. Devices within the cardiovascular system were discussed in section 3 in relation to hemodynamic factors and mechanotransduction and in section 4.3 above. It is necessary to return briefly here because of the long-standing view that surface properties and protein adsorption are precursors to the response of the vascular endothelium to biomaterials. In devices intended to remain within the circulatory system, it is usually desirable that an endothelial lining is established on the surface of the material; this would apply to vascular grafts48 and coronary artery stents.172 Once again, although it is ubiquitous, the dominant plasma protein albumin plays little role in the host response, especially in the context of the effects of wall shear stresses already discussed and inflammation and the immune response to be discussed later. On the other hand, proteins of the clotting cascade and FN may play significant roles. The pathophysiology of restenosis after stent placement has been discussed by Jukema et al., who noted the interrelation between inflammation and ECM development within the zones of endothelium and vascular smooth muscle cells.55 It is obviously difficult to establish by direct observation the chronology and mechanisms involved in the short period of protein adsorption and initial endothelialization, and so indirect evidence has to be considered. Most attention has recently been focused on FN. In an extensive in vitro study aimed at the molecular mechanisms by which adsorbed proteins influence endothelial cells, Yang et al. reported four components of these potential mechanism.173 First some adsorbed proteins bind to cell surface receptors generating signal transduction, which activates cell surface integrins through increasing intracellular calcium levels. Integrin expression may also be enhanced through the effect of molecules such as thrombospondin in the adsorbed layer on TGF-β signaling pathway. Then those adsorbed proteins that contain the RGD sequence, including thrombospondin, VN and FN, activate focal adhesion pathways, resulting in increased focal adhesion formation, actin cytoskeleton organization and cell adhesion and spreading. Finally fibrin and FN interact within the layer to further regulate the actin cytoskeleton. It is important to note with respect to the last point that FN contains two major fibrinbinding sites;174 this binding is important for cell migration and adhesion in several situations, including fibrin clot reorganization, and could be very relevant here in establishing the cell supportive ECM adjacent to the biomaterial. It is also possible for other ECM proteins, for example collagen I, to assist FN in supporting endothelialization.175 This critical role of FN is also consistent with the recently discussed ideas of the activation of endothelial cells by FN in angiogenesis, where there is a transition from quiescent to active state with enhanced migration and proliferation.176 It is also supported by recent evidence of a signaling pathway by which FN achieves endothelial cell chemotaxis through transactivation of FGF receptor-1.177 Kim et al. have been able to model endothelial cell migration and spreading on FN substrates and explored mechanisms involving ligand density, cell integrin expression level and integrin-ligand binding affinity.178 The critical role of FN in the regulation of endothelialization, assisted by fibrin, thrombomodulin and collagen, is supported by attempts, usually in vitro, to improve endothelial cell attachment by surface modification. Sun et al. heparinized polycaprolactone, resulting in enhanced FN adsorption and endothelial cell attachment, while, critically, reducing smooth muscle cell attachment.179 Traub et al. observed the promotion

of endothelial cell attachment through the use of a bifunctional protein consisting of a FN domain and VEGF-A.180 Teng Goh et al., in a review of cardiovascular bypass devices, noted that FN combined with stem cell homing factor SDF-1α, when adsorbed onto polyester grafts, improved endothelial cell coverage and reduced intimal hyperplasia181 and Chlupac et al. reported on the beneficial effect of combined fibrin and FN, as well as laminin, on attachment of endothelial cells under conditions of shear stress.182 It is also known that endothelial cells adherent on FN-rich surfaces, in association with platelet endothelial cell adhesion molecule-1, respond to shear stresses with an adaptive stiffening response.183 This discussion demonstrates that the process of protein adsorption is a potentially important factor in the development of the host response to implanted devices, but has to be seen in the context of intervention-induced trauma and inflammation and mechanotransduction. The mere adsorption of major plasma proteins, to give a layer most likely of monomolecular dimension, is of no consequence. If, however, the adsorbed protein undergoes conformational change such that domains are exposed that are able to bind to receptors of relevant cells, and if they are able to participate in 3D ECM development, then they may well regulate the host response. FN appears to be the dominant glycoprotein here, possibly assisted by VN, collagen, thrombomodulin, fibrinogen and fibrin depending on the circumstances. Within a stent, the balance between endothelialization and smooth muscle cell controlled neoatherosclerosis will, to some extent, be controlled by the manner in which FN can be organized on the biomaterial surface. 4.4.2. Proteins and the Differentiation of Stem Cells on Biomaterials Surfaces. As discussed by the present author recently184 one of the major difficulties in developing our understanding of biocompatibility phenomena, and one of the major hurdles in identifying appropriate scaffolds for tissue engineering, has been the assumption that knowledge about biocompatibility in traditional medical device technology can be transferred to the performance of these scaffolds. Nothing could be further from the truth; an objective of the present paper is to identify overarching paradigms in biocompatibility mechanisms but this has to reflect the quite different circumstances that exist in different modes of application of biomaterials. As noted in section 2, the most important specification for a tissue engineering scaffold or template is that it should recapitulate the microenvironment of the niche of the target cell; this is not achieved by designing materials that avoid stimulation of protein and cellular activation. This situation was expressed, and discussed in detail, by Lutolf and Hubbell in 2005185 who specifically stated; “Biomaterials play central roles···as designable biophysical and biochemical milieus that direct cellular behavior and function. The guidance provided by biomaterials may facilitate restoration of structure and function of damaged and dysfunctional tissues... providing a 3D support to interact biomolecularly with cells to control their function, guiding the spatially and temporally complex multicellular processes of tissue formation”. With reference to ex vivo bioreactor based tissue engineering, the interaction between biomaterials and proteins is not a matter of how a biomaterial surface engages with the complex dynamic in vivo physiological environment, where the latter imposes itself on the former, but how the situation is reversed such that the material, including its surface, imposes itself on an artificial, cell-containing, physiological-mimicking fluid environ14


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of cell-adhesive motifs. Clearly the choice of specific chemistries is critical; a hybrid of two biopolymers will not work if neither has suitable cell adhesive domains, as revealed by Ziv et al. with silk−alginate hydrogels.196 The final factor here is nanostructure. It is clear that some forms of nanostructure influence cell adhesion and phenotype197 and although mechanisms have not been determined, mechanostransduction is most likely the dominant factor. The role of proteins is not at all clear. It has been suggested that cells in contact with nanofibers produce quite different protein and glycoprotein expression than those on planar surfaces.198 The most likely consequences of the nanoscale here is the greater contact between substrate and target cell so that mechanotransduction factors are enhanced; also, through the selection of medium, for example with different differentiation factors199 and the presence of appropriate peptide motifs on the surfaces, including those of self-assembling peptide nanostructured scaffolds,200 the binding to cell adhesive domains is optimized. This is compatible with the earlier comment that a biomaterial has to impose itself on the cell-containing medium, where the template architecture, the motifs on the surface and the biomolecules present in the medium are more important than the material chemistry and protein adsorption profile. 4.5. Protein Coronas on Nanoparticles. Protein adsorption has become a major factor in the clinical use of nanoparticles in drug and gene delivery and imaging systems.201 Knowledge about the mechanisms and consequences of this phenomenon is only just entering the literature, but several good reviews have recently been published.202,203 It is not the purpose here to consider the effects of the adsorbed proteins on the function of the relevant nanoparticles, but it is necessary to evaluate the effects on biocompatibility, and specifically on toxicity and effects on the immune system, and on translocation and migration of the particles; these aspects are discussed in more detail in section 6. Once a nanoparticle (NP) comes into contact with tissues and body fluids, proteins and other biomolecules are rapidly adsorbed onto the NPs, forming a surface layer, known as the protein corona (PC), forming what may be described as the NP-PC complex. The nature of the PC will depend on the physicochemical parameters of the NP, including size, shape, potential, surface chemistry, and the surrounding fluid, whether it is blood, saliva, pulmonary fluid, and so on. Prevailing opinions are that the PC is established quickly, in less than a minute, and that it does not change too much, at least qualitatively, over time and during passage through the body.204 The PC is considered to have two parts, the first a tightly bound layer which has a long exchange time and is described as the hard corona, and the second being the soft corona, which has a fast exchange time with weak binding affinity. The presence of the PC can be either advantageous or disadvantageous as far as biocompatibility and toxicity are concerned. A major factor in NP toxicity is the effect of positive surface charges on cell plasma membranes because of their interaction with cell surface negative charge.205 The PC may mask these factors. On the other hand, proteins in the corona, especially the soft corona, can undergo conformational change, with potentially increased immunogenicity and greater activation of complement. The significant questions here concerns whether consistent pathways linking NP and NP-PC characteristics can be identified and then exploited in order to maximize the beneficial effects of NPs, especially targeting and internalization,

ment in order to guide the tissue generation described by Lutolf and Hubbell. It follows that protein-related biocompatibility in tissue engineering is controlled by the nature of the culture medium and the manner in which the biomaterial surface, whether natural or modified, can influence the target cells. Of considerable relevance here is the architecture of the biomaterial template. If the template is a microporous polymer or ceramic, even though the surface area may be relatively large, the proportion of the cells seeded into the construct that actually come into direct contact with the material surface is small; cell behavior is therefore governed by cell−cell interactions and the mechanotransduction effects previously discussed. The reality is that microporous structures (i.e., porosities >50% and pore sizes measured in hundreds of microns) made of synthetic materials rarely produce effective tissue regeneration in reasonable volumes, and the precise characteristics of the material are not very relevant. Because culture media usually contain serum proteins, the nonspecific adsorption of these proteins onto surfaces is usually considered as a deterrent to cell adhesion and proliferation. An important goal here is to minimize the nonspecific adsorption while encouraging various bioactive signaling processes. As discussed by Grafahrend et al., although this can be accomplished on simple 2D surfaces, this is rare in 3D scaffolds.186 There have been attempts to modify surfaces by physicochemical means, for example by plasma treatment,187 but although they may alter surface characteristics including wettability, the effects on protein adsorption and cell adhesion and proliferation are variable and generally irreproducible. The problem, as discussed by Pashuck and Stevens188 is that synthetic polymers typically lack cell adhesion sites, and nonspecific protein adsorption makes it worse. Several approaches are being used to address this issue, the majority involving either the use of biopolymers instead of synthetic polymers, or hydrogels, nanostructures, protein/ peptide functionalized materials, or, more usually, a combination of these. Biopolymers offer many advantages, but although some, such as collagen I, can be reconstituted into a fibrillar matrix form where the polypeptide chains support cell adhesion and spreading,189 not all have this capability; silk, considered by many to be an attractive option lacks specific domains for cell attachment in most of its forms.190 Proteins may be coated onto synthetic polymers. Chatakun et al. have reviewed the effects of FN, osteopontin, tenascin, bone sialoprotein and BMP-2, showing variable effects; interestingly pretreatment with FN enhanced early stage osteogenic differentiation of MSCs but the proteins had mixed effects in later stages of proliferation.191 It should not be forgotten, of course, that cells in culture may secrete their own ECM proteins which can affect subsequent behavior. In a review of the design of polymeric materials for culturing human pluripotent stem cells, Higuchi et al. noted that even the better materials require the secretion of ECMs from these cells or the adsorption of proteins from a conditioned medium to support the cell lines.192 Of greater relevance is the development of hydrogels as templates, which could be either synthetic or natural, and especially protein-conjugated hydrogels.193 With synthetic polymers, PEG has been conjugated to a variety of proteins, including fibrinogen194 and collagen,195 and there are many examples of collagen/hyaluronic acid/chitosan hybrids and similar structures. One significant issue here is that the composition of the conjugate can be adjusted in order to attempt optimization of mechanical properties and presentation 15


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biocompatibility; Schutte et al. studied the release of cytokines, chemokines and growth factors from monocytes and macrophages cultured on common biomaterials surfaces and although they found some differences, these were not significant (apart from a PVC polymer with known cytotoxic additives) and not reflective of a shift in in vivo inflammation or wound healing.209 The authors concluded that these polymers were similar in being nonreactive through native signaling pathways. What are of relevance here are moieties that may be released from the surfaces, over time, through leaching, diffusion, erosion, or degradation processes or those situations in which the biomaterial is itself presented to the physiological environment in a mobile or labile form. With polymeric systems, such moieties include monomers, oligomers, catalysts, antioxidants, processing additives and degradation products. With metallic systems, they are metal ions, impurities, corrosion products, and retained ancillary manufacturing products such as those used in surface treatments. Mobile components include those molecules used in biomaterials-based therapeutics (such as dendrimers) and metallic-based imaging agents, for example, those involving gadolinium and iron oxide. It is necessary to temporarily separate out biostable manufactured implantable devices from other systems in order to assess the potential role of material chemistry. Over the last few decades, as noted in the introduction, the choices of clinically proven implantable devices has led to the use of materials that are essentially inert, and there are very few examples where the chemistry is a determinant of biocompatibility. The present author was involved with clinical studies of alloy systems that were being introduced into orthopedic and dental practice before their corrosion resistance was optimized, and there certainly was an influence of the alloy chemistry on the inflammatory host response.210 This does not happen with contemporary devices and materials; as noted below, the development of the local host response is largely mediated by factors other than chemistry. The situation may be different with devices that do degrade, intentionally or otherwise, or are injected in a form with incomplete structure. With nonimplantable devices, the host response may be quite different as the chemical entity may be deliberately exposed to the physiological system in a fluid or suspension form. These products could be in the form of polymer therapeutics211 or contrast agents,212 with far greater opportunity to interact with the defensive mechanisms on the body, with either beneficial or adverse effects. 5.2. Essence of Inflammation, Immunity, and Fibrosis. The classical view of the host response to an implanted material involves acute inflammation, chronic inflammation, and fibrosis, the extent of each phase depending on a number of factors. In recent years, there has been a trend to consider these events as a continuum within the mechanisms of the immune response. More importantly, it has become possible to consider these events in terms of the evolution of theories about inflammasomes, damage-associated molecular patterns (DAMPs), sterile inflammation, and the immunology of fibrosis. The biomaterials community has struggled with the implications of the involvement of the immune systems in biocompatibility since the former has traditionally been associated with the interactions between hosts and pathogens, while the latter is associated with interactions between host and nonpathogens. The key to a better understanding of this situation originates with the work of Matzinger who developed ideas of the so-called danger model and different concepts to

and minimize adverse effects in relation to immunotoxicology, or better still, to combine these effects in relation to biomaterial controlled immunomodulation.206 Considerable progress is being made, for example, with the modulation of the glycans component of the PC207 and the modification of quantum dots with a small molecule that induces a protein-misfolding event in the PC that produces cell-specific, receptor-mediated endocytosis.208 4.6. Case against Interfacial Phenomena. The analysis above indicates that in the majority of situations where biomaterials are used clinically, protein adsorption is of minor importance in biocompatibility pathways. In particular, biomaterial-induced thrombogenicity is not a major clinical problem provided appropriate clinical techniques are used, and even when thromboembolic events do occur it is flow disturbances in the blood that provides the main driver. Flow patterns and hemodynamic forces within the vascular system are nonuniform, where flow ranges from laminar with high and directed wall shear stresses to irregular distributions of low wall shear stresses, these variations being associated with different endothelial cell and platelet responses. Complement activation in blood-contacting devices is only important in situations where very large surface membranes are involved over protracted periods of time, or in a very small number of cases where idiosyncratic anaphylactic reactions are experienced. For tissue-contacting devices, even though surfaces do become covered with proteins, interactions between these proteins and cells are also of minor importance, where mechanical stresses and surgery-induced trauma dominate the pathways and where phenomena such as EMT and 3D organization of glycoproteins in the interfacial zone play important roles. With synthetic materials used as tissue engineering scaffolds, the relationships between protein adsorption and cell adhesion, proliferation and differentiation are variable and irreproducible; however the design of surfaces that resist nonspecific adsorption and the incorporation of certain protein motifs in hydrogel structures are of potential importance. The one situation where protein adsorption does have an impact is that of nanoparticles, where the NP-PC complex can influence toxicity.

5. ROLE OF BIOMATERIAL CHEMISTRY IN INFLAMMATION AND IMMUNE RESPONSES The third major potential mediator of the development of the host response, after mechanotransduction and adsorbed proteins, is the influence of biomaterial chemistry on the cells and the ECM in the vicinity of a biomaterial and on the physiological system in general. On the one hand this concerns the chemistry of the surface and, especially, of any components that may be released from that surface. On the other hand, the effects have to be placed into the context of inflammation and the immune response. Both of these are immense subjects in themselves and it is only possible to deal with the principles here insofar as they relate to biocompatibility pathways. 5.1. Essence of Biomaterials Chemistry. With respect to the chemical structure of the biomaterial, the surface chemistry of the material will not be considered here since, as discussed in section 4, any effects arising from the surface are likely to be mediated by the rapidly forming interfacial layer or by topographical effects discussed later. It should also be noted that it is always possible to demonstrate effects of different surface chemistries of the release of biomolecules from cells in vitro but that this is not usually relevant to clinical 16


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Figure 3. NLRP3 inflammasome activation within macrophages induced by engineered nanomaterials. Reproduced with permission from ref 228. Copyright 2011 American Chemical Society.

surfaces, are able to sense endogenous molecules as well as microbial PAMPS, although this is by no means the only, or indeed most important, process. Second, the release of intracellular cytokines and chemokines activate common pathways downstream of PRRs. IL-1, including both IL-1α and IL-1β, is likely to be a key mediator here, acting through the IL-1 receptor. IL-1β appears to be highly significant, its secretion by inflammatory cells being dependent on a multiprotein complex termed an inflammasome, which operates through the activation of caspase1. Inflammasomes are innate immune system receptors and sensors that induce inflammation in response to pathogens and molecules derived from host proteins.219 Several families of PRRs are important components of the inflammasome complex; upon sensing certain stimuli, a relevant receptor can oligomerize to become a caspase −1-activating scaffold, which subsequently cleaves the proinflammatory IL-1 family of cytokines, which may result in inflammatory cell death. Inflammasomes are named after the protein that forms the scaffold. Most are formed with NOD-like receptors (NLRs) while others contain the absent in melanoma 2 (AIM) − like receptors (ALRs). The NLRP3 inflammasome is activated in response to a wide range of stimuli, and has been implicated in BISI events, as discussed below. Certain adapter proteins are involved, especially including ASC (apoptosis-associated specklike protein containing a CARD), shown in Figure 3. It is also noted that ROS constitutes an additional danger signal with NLRP3 stimulation and caspase 1 activation,220 with the production of ROS by neutrophils in the acute phases of exposure to biomaterials and nanoparticles. Third, there is direct activation by receptors that are not normally associated with microbial recognition. Importantly, several endogenous molecules released from necrotic cells, such as heat shock proteins and nucleic acids, or which are present in the ECM such as hyaluronan and heparan sulfate, have been reported to act as DAMPs and activate TLRs.218 The significant issue here with BISI is the fact that the degree of inflammation in response to any challenge, and the temporal profile of the process, will determine the resulting host response. If inflammation is insufficient, whether caused by pathogens, dead cells, or exogenous irritants, then the response

replace the standard self-and nonself paradigm in the 1990s,213,214 although these ideas were not translated to biomaterials until some time later.215−217 This concept is consistent with recently expressed views on sterile inflammation,218 described as inflammation that is the result of trauma, ischemia-reperfusion injury, or chemically induced injury, typically in the absence of any microorganism. Just as with microbially produced inflammation, sterile inflammation is associated with the recruitment of neutrophils and macrophages and the production of pro-inflammatory chemokines and cytokines, especially TNF and IL-1. In relation to biomaterials, and especially those medical products that have a long, usually intended, residence time in the body, it is necessary to consider the progress of sterile inflammation from the moment of initial contact through to ultimate, clinical acceptance or elimination. The recent discussions about the immunology of sterile inflammation and fibrosis now allow such an analysis. It is important to note that the mechanisms of biomaterial-induced sterile inflammation (BISI) have to be consistent with those that are implicated in similar conditions, especially sterile inflammatory diseases, including both those associated with endogenous molecules and those with exogenous substances. Of relevance are diseases associated with chronic inhalation of irritants such as asbestos, crystal deposition, for example of monosodium urate in joints leading to gout, and, possibly, atherosclerosis and endothelial cell dysfunction following engulfment of cholesterol and Alzheimer’s disease, where neurotoxicity is associated with activated microglial cells adjacent to β-amyloid plaques. Although the ideas about DAMPs and sterile inflammation are only now being developed, the proposals of Chen and Nuňez218 are worth consideration. They proposed that there are three, not necessarily mutually exclusive, mechanisms by which sterile stimuli trigger inflammation. The first involves pattern recognition receptors (PRRs). PRRs normally sense conserved structural moieties within microorganisms, sometimes called PAMPS, pathogen-associated molecular patterns; it is suggested that in this first phase PRRs are activated by mechanisms similar to those used by microorganisms and PAMPs. This is partly based on the observations that Toll-like receptors (TLRs), the transmembrane proteins located at cell 17


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Figure 4. Mechanisms driving major macrophage activation phenotypes in tissue repair, regeneration, and fibrosis. Recruited and resident macrophages undergo marked phenotypic and functional changes in response to DAMPs, PAMPs, growth factors, cytokines, and other mediators released in the local tissue microenvironment. Macrophages produce a variety of factors that stimulate the proliferation, differentiation, and activation of fibroblasts, epithelial cells, endothelial cells, and stem and progenitor cells that facilitate tissue repair. During the later stages of the repair process, they assume a regulatory pro-resolving phenotype that ensures that the tissue-damaging inflammatory response is suppressed and normal tissue architecture is restored. If the process is not controlled effectively, persistent inflammation and/or maladaptive repair processes can lead to tissuedestructive fibrosis. DAMP, damage-associated molecular pattern; PAMP, pathogen-associated molecular pattern; Treg cell, regulatory T cell; IRF5, interferon regulatory factor 5; NOS2, nitric oxide synthase 2; LXR, liver X receptor; AREG, amphiregulin; Arg1, arginase-1; IRF4, interferon regulatory factor 4; PPARγ, peroxisome proliferator-activated receptor γ; FGF, fibroblast growth factor; GAL-3, galectin-3; TGF, transforming growth factor; IC, immune complex; GR, glucocorticoid receptor; ATF3, activating transcription factor 3; and SOCS, silencer of cytokine signaling. Reproduced with permission from ref 222. Copyright 2016 Elsevier.

by macrophage-colony stimulating factor (M-CSF), TNF-α or IFNγ. They may also regulate the proliferation of adjacent parenchymal or stromal cells or activate stem cell and local progenitor cells, as illustrated in Figure 4. These cells then mostly exhibit an anti-inflammatory phenotype, M2, responding to IL-10 and other inhibitory molecules. At this stage there is much discussion in the literature about the real distinction between these phenotypes; the purpose of the discussion here is to determine whether biomaterials are able to influence this polarization. With respect to the role of adaptive immune responses in fibrosis, activation of T and B lymphocytes leads to the production of both proinflammatory and profibrotic cytokines.222 The processes are strongly influenced by effector cells. IL-12 drives Th1 differentiation, which produce proinflammatory cytokines, whereas IL-4 drives Th2 differentiation, which secrete anti-inflammatory cytokines. Both Th17 and T regulatory cells (Tregs) play important roles: an imbalance between Th1/Th2/Th17 subsets appears to drive inflammation in the early stages of disease, where Th1 and Th17 dominate, and then drive fibrosis later, when Th2 dominates. 5.3. Role of Biomaterials in Sterile Inflammation and Fibrosis. 5.3.1. Experimental Considerations. It is now necessary to translate the recent advances in understanding sterile inflammation and the immunology of fibrosis into potential pathways in the host response to biomaterials. At this stage there is little direct evidence; the review here includes biomaterials in all current forms. Christo et al. have recently reported on investigations on the early responses to the

is persistent. If it is excessive, then it can lead to chronic or systemic inflammatory disease. The situation is influenced by the role of sterile inflammation in fibrosis and tissue regeneration. Wick et al. have discussed the immunology of fibrosis;221 the guiding principles are that in all forms of fibrosis, inflammatory-immunologic reactions take place in the earliest stages of a response, promoting subsequent pro-fibrotic processes, and that elements of both the innate and adaptive immune systems are involved. In the former case, neutrophils, eosinophils, and mast cells all play roles, where pro-fibrotic molecules are released, with stimulation of fibroblast proliferation through synthesis of cyclooxygenase and prostaglandins, expression of collagens and the orchestration of fibrosis. It is likely that the balance between matrix metalloproteinases (MMPs) and the counteracting tissue inhibitors of matrix metalloproteinases (TIMPs), controls the balance between ECM deposition and breakdown. Macrophages, and especially the phenomenon of macrophage polarization, have received most attention here.222 It is evident that monocytes and macrophages are recruited and activated by a number of different mechanisms and that their functional characteristics control tissue repair and fibrosis. In wound healing, the initial inflammatory response is considerably diminished if there is a macrophage deficiency but their removal at later stages results in decreased debridement and less efficient repair since these macrophages are characterized by the production of many different growth factors and other soluble mediators. The early stage pro-inflammatory phenotype is usually referred to as the M1 macrophage,221 being activated 18


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implanted with a degree of tissue damage. It is inevitable that the mechanical environment is disturbed, affecting both fibroblasts and especially myofibroblasts that are the cause of the contracture. Breast tissue is subjected to movement during normal body function so that the process is likely to be progressive; interestingly, subglandular implants have a 2-fold increase in risk compared to submuscular implants,232 which are better protected from movement. It is also relevant that the TLR4 stimulation is involved in this ECM remodeling, which is affected by estrogen (specifically 17β-Estradiol) that increases transition to myofibroblast phenotype and contraction;233 mechanical forces are also known to stimulate stromal estrogen production in the breast.234 Consistent with the discussion in 5.2 above about the combined effects of innate and adaptive immunity in BISI, Wolfram et al.231 have demonstrated that the balance between Tregs and Th1/Th17 activity is influential in the progression of capsular contraction. Surgical Meshes. Surgical meshes provide another controversial example of implant-related biocompatibility problems. Meshes are used for hernia repair235 and for the treatment of stress urinary incontinence (SUI)236 and pelvic organ prolapse (POP).237 This is a difficult area of patient management because the need for therapy arises because soft connective tissues have failed to maintain function and/or integrity and require support while “healing” or “regeneration” takes place. Although minor cases can be treated very successfully by conservative means, more complex lesions rely on implanted devices, almost always in the form of a mesh, to hold tissues in place. Patients are often compromised in terms of obesity, poor nutrition, diabetes, and other lifestyle factors, and the site of implantation, primarily in SUI and POP, may be classified as clean-contaminated so that infection may be a risk factor. Clinical complications include recurrence and adhesions and, in the case of SUI tapes, erosion through the urethra or vagina. Meshes may be synthetic, usually polypropylene, or acellular xenograft-derived products, usually porcine or bovine, derived from the dermis or small intestine.238 The discussion here is confined to synthetic meshes because variable immune responses to mammalian ECM devices, through both anti-Gal and antinon Gal antibodies239 complicates this issue in as yet an unclear manner. Histopathological characterization of the tissue that forms within and around the meshes does vary considerably, with both areas of chronic inflammation and fibrosis. There have been many attempts to implicate the chemical nature of the materials used, and their putative degradation products, in the development of these responses, but there is no significant evidence of this. Of greater importance are the design and mechanical properties, especially stiffness, of the mesh. On the assumption that the site is indeed sterile, the development of this response is again well-explained by the combination of mechanotransduction and sterile inflammation. Although clinical techniques attempt to minimize mechanical stresses, for example, transvaginal SUI tapes are placed “tension-free”, forces are inevitable through normal body movement. Recent studies of soft tissue healing make it abundantly clear that inflammation and fibrosis occur side-by-side and that mechanical forces stimulate the responses of all relevant cells.240,241 Most pathological observations of tissue within meshes reveal mixed cell populations predominantly of leukocytes (mainly CD45+), T and B lymphocytes, mast cells, M1 and M2 macrophages, fibroblasts and myofibroblasts.242−244 Neither the nature of the polymer nor of any

peritoneal injection of poly(methyl methacrylate) (PMMA) beads, with special emphasis on the role of inflammasomes.223 The experiments did not reveal any involvement of NLRP3 in the development of the foreign body response; on the contrary AIM2 expression influenced leukocyte infiltration and controlled collagen deposition. The authors noted the general plasticity of inflammasome triggers, indicating that inflammasome activation may be associated with cell damage occurring during implantation. These results tend to support an earlier study, also with the peritoneal injection of PMMA particles, which showed that the full development of the response was not dependent on NLRP3 but was dependent on the generation of active caspase 1 and the secretion of IL-1β.224 Samelko et al., in attempting to identify the mechanisms of adverse responses to metal−on−metal hip replacements217 provide in vitro and in vivo data that indicate cobalt alloy particles induce macrophage dominated inflammation, which was abrogated if danger signaling was blocked. However, blocking TLR4 had no effect. It was concluded that excessive innate immune reactivity to cobalt debris was likely to be inflammasome driven, possibly the result of direct and secondary DAMPs and maybe other specific toxicity responses, but not dominated by TLR4 activation. On the other hand, several studies have shown the involvement of NLRP3 activation in the inflammatory response to nanoparticles, especially biomedically applied nanoparticles. This was reviewed by Sun et al.220 Ramadi et al. demonstrated the critical role of the NLRP3 inflammasome in the response to injected silver-based nanoparticles in mice,225 similar results being obtained for CdSe/ZnS quantum dots.226 A number of in vitro and in vivo studies have revealed the importance of surface chemistry on immune cell activation and inflammasome involvement. Li et al. modified mesoporous silica scaffolds with −COOH, −NH2, −SO3H, and other functional groups and found that inflammation was dependent on the surface chemistry and that the NLRP3 inflammasome was necessary for IL-1β production and immune cell infiltration in vivo.227 Similarly, Lunov et al. found that amino-functionalized polystyrene nanoparticles activated the NLRP3 inflammasome in macrophages,228 while Yang et al. found that benzoic acidfunctionalized carbon nanoparticles modulate inflammatory cell recruitment and NLRP3 inflammasome activation.229 5.3.2. Implications for Clinical Biocompatibility. Most of the above observations refer to injections of microparticles or nanoparticles. It is necessary to consider the range of clinically used devices in order to interpret the influence of sterileinflammation related phenomena. There have been several high profile challenges with the biocompatibility of implantable devices. Silicone Breast Implants. Silicone gel-filled breast implants provide one main example; it is not necessary to consider the controversy over putative silicone-induced autoimmunity because this mechanism was determined to not be relevant,230 but it is appropriate to consider capsular contraction because this does affect a significant number of patients.231 This is the generation of excessive fibrous tissue around the implants; for a long time, this was associated with adverse reactions to either the silicone elastomer or to silicone gel that diffuses through the shell, but with very little evidence. Although most publications on this subject state that the pathogenesis of capsular contraction is unknown, the condition appears to result from a combination of mechanotransduction and sterile inflammation. These implants are relatively large and are 19


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Figure 5. General schematic showing the effects of cobalt alloy particulate on macrophages acting directly and indirectly on three major proinflammatory innate immune pathways: (1) general toxins (such as inducing hypoxia like cell responses), (2) as danger associated molecular patterns, DAMPs (inflammasome induced activation), and (3) interacting with the pathogen associated molecular pattern (PAMP) pathway of TLR4. Reproduced with permission from ref 217.

induced in periprosthetic tissues by surgical trauma, including necrosis, ischemia and synovitis,248 these danger signals then being amplified by the micrometer-sized wear particles; this size is important because the average particle resembles the size of bacteria. Macrophages, and possibly foreign body giant cells dominate, with a strong emphasis on the M1 phenotype.249 A significant amount of M1-associated cytokines are released on interaction with particles, including IL-1α, IL-1β, IL-6, IL-10, TNFα, PDGF, and EGF.250 Many papers have been published on putative mechanisms and pathways in polyethylene particle induced osteolysis. The recent review of Kandahari et al. provides a good example.251 The sequence of events is consistent with the principles of BISI covered earlier, with the added factor of how macrophagedominated inflammation leads to osteoclast-induced bone resorption. The NALP3 inflammasome appears to be involved, along with NF-κB, TNF-α, and IL-1β activation. Pathogenesis here is multifactorial but driven by increased osteoclastogenesis, proinflammatory enzymatic bone resorption and decreased osteoblastic activity. The osteoclastogenesis is mediated by the growth factors RANKL and M-CSF, which results in stimulation of monocytes and osteoclast progenitors and differentiation into mature, multinucleated osteoclasts. Several other parallel pathways are operative, including the generation of ROS in giant cells that further contributes to NALP3 activation and caspase-1 activity. Through these related pathways, the resulting clinical outcome is a dose−response relationship between wear rate and osteolysis.252 It is important to note that T lymphocytes are observed only rarely in these tissues and play no significant role in this osteolysis. The situation is different when metal-on-metal (MoM) combinations are used in total hip replacements compared to metal on polymer (MoP). Some osteolysis takes place but a small number of patients experience a significant soft tissue response. This has been a very controversial topic, with

adsorbed proteins have any significant effect. The main device characteristic that does influence cellular responses is the mesh design, including filament structure and porosity; the effects are largely mediated through the amount of material being present and filament surface area since these parameters increase the challenge to the tissue and the level of oxidative stress. It is a fine balance, influenced by patient variables and clinical technique, the most important contributors to this being the M1/M2 profile, fibroblast to myofibroblast transition245 and the activity of T-lymphocytes. Excessive fibrosis, especially involving myofibroblasts cause scarring and tissue contraction, whereas prolonged pro-inflammatory behavior (for example caused by M1 polarization in the event of ROS production in dense meshes) can lead to lack of healing, dehiscence, and erosion of tissues. Wear Debris and Joint Replacement Prostheses. One of the most successful of all implantable devices is the total joint replacement. Even here, however, there remain significant biocompatibility issues, the most important of which is the host response to the wear debris that is generated at articulating surfaces; in this respect the joint receiving most attention is the hip, although the phenomena are relevant to other systems. For several decades, these problems arose from the generation of polyethylene particles caused by wear of acetabular components, and the consequences of this debris on resorption of bone and resulting in loosening of the prostheses. It has been known for a long time that the greater the wear rate, the higher the risk of osteolysis. Although, as discussed in section 3, mechanotransduction plays a significant role in musculoskeletal biocompatibility phenomena, and mechanical force transduction in osteocytes is reasonably well-understood,246 mechanotransduction is unlikely to be a factor here. This is clearly a prime example of BISI, consistent with the names usually given to the clinical effect, “aseptic loosening” or “aseptic osteolysis”.247 Initial events are dominated by damage 20


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istration but declining during the degradation process.261 Concerns have been expressed about the potential adverse effects of gadolinium-based compounds used as contrast agents. Schmidt-Lauber et al. have found that free gadolinium derived from the particles triggers a NLRP3 inflammasome-dependent inflammatory response that leads, after IL-1β release, to fibrosis in patients with chronic kidney disease.262 Finally, of considerable importance is the host response to different forms of silica, which is relevant both to the etiology of silicosis and the fate of injected silica nanoparticles used in some therapies. Once again, activation of the NLRP3 inflammasome and caspase-1 activity are primarily involved in the inflammation-fibrosis continuum, where both the particle size and morphological form (crystalline vs, amorphous, colloidal vs pyrolytic, etc.) are important factors.263−265 This section was predicated on the often-voiced view that the biomaterial chemistry was responsible for the development of the host response. The evidence would suggest that in the vast majority of situations, this is not the case. Sterile inflammation is usually driven by nonchemical factors. A combination of mechanotransduction mechanisms and the effects of local tissue trauma are sufficient to constitute DAMPS, with activation of PRRs and their associated inflammasomes. Neutrophils are obviously involved here, contributing additional danger signals in the form of ROS and a variety of intracellular cytokines and chemokines activate common pathways downstream of PRRs. Several balancing mechanisms control the direction of the response, including macrophage polarization, the competing activity of MMPs and TIMPs with respect to ECM deposition and the variable effects of lymphocyte-originating proinflammatory and pro-fibrotic cytokines. The role of the biomaterial and device is largely determined by morphological and architectural features. It is also relevant to consider that the inflammation-fibrosis balance can change over time, for example, during the late stage of degradation of biodegradable polymers and the release of pro-inflammatory agents, especially microparticulate, acidic moieties, in either implantable devices266 or tissue engineering constructs.267 This BISI process is replicated in those situations where nondegradable microparticles are released from devices, with some variations in the end point of the inflammation−fibrosis balance depending on the circumstances, for example the osteoclastogenesis activity with polyethylene wear debris. There is a departure from this scenario where chemistry does matter and that is when biomaterial derived nanoparticles can be internalized in some cells where specific genotoxic of hypersensitivity responses may be seen. Inflammasomes are still involved with responses to nanoparticles, as seen with responses to injected metal-based therapeutic or diagnostic nanoparticles where metal intracellular ions may constitute the DAMP. It is also possible that chemistry can influence inflammasomes under some circumstances when functionalized microparticles are used.

multiple descriptors being used for the histopathologic appearance, including pseudotumor253 and ALVAL, aseptic lymphocyte dominated vasculitis associated lesion.254 There are two major, relevant, differences between MoM and MoP combinations. Polyethylene particles are not degradable within tissues or cells and, as noted above, are of micron size. Cobalt alloy particles derived from typical MoM prostheses, and of submicron size (often around 10 nm) and will corrode or dissolve within tissues, especially within cells. Thus, the chemistry of polyethylene plays no role in the host response but cobalt alloy nanoparticles, and especially cobalt ions are reactive in tissues. There has, therefore, been much discussion about the effect of metal ions in the host response, although without total agreement at this stage.217,255,256 The effects are primarily on macrophages and lymphocytes, where the internalized cobalt-containing nanoparticles release cobalt ions, which can have direct cytotoxic effects. The complex chemical characteristics of cobalt ions certainly provide danger signals, with significant inflammasome activity and resulting hypoxia like cell responses illustrated in Figure 5. Genotoxic effects may also occur. There is a variable presence of lymphocytes in the soft tissues, resulting in the equally variable observation of the ALVAL phenomenon; there have been suggestions that cobalt ions bind to intracellular proteins to form haptens, which then elicit a type IV hypersensitivity response. Although there are some doubts about the role of metal hypersensitivity in the host response to total joint replacements,257 the recent paper by Paukkeri et al. gives the best current representation of the sitaution.258 There are distinct inflammatory phenotypes that characterize the response; a macrophage-dominated response is seen in association with high metal ion blood and high cytokine levels, depicting metal-induced cytotoxicity resulting in massive macrophage infiltration and clearance of necrotic tissue. With lower metal levels, the tissue is characterized by a T-cell response reminiscent of hypersensitivity. 5.4. Case for Biomaterials-Induced Sterile Inflammation. The above examples demonstrate very clearly the significant role of BISI in the development of the host response to biomaterials. The examples chosen all relate to implanted devices, simply because there is much more evidence about them than with other biomaterial applications. There are, however, many situations where there is good, if not as yet very comprehensive, evidence about inflammasome related mechanisms concerned with other types of biomaterials; it is important to briefly reflect on these in order to appreciate the significance of these pathways. Little is known of inflammasome activation in tissue engineering applications; these would obviously not be expected under ex vivo reactor situations but could be involved in in vivo applications. Perhaps not surprisingly, chitosan, which is the deacetylated form of chitin, a β-(1,4) linked polymer of N-acetylglucosamine, features quite prominently here. Bueter et al. have shown that NLRP3 inflammasome activation is involved in the immunostimulatory properties of chitosan, with multiple mechanisms.259 The effectiveness of synthetic polymer particulate vaccine adjuvants, for example poly(lactideco-glycolide) and polystyrene, has been shown to be due to interaction with dendritic cells and activation of the NALP3 inflammasome.260 Interestingly, in vaccine or drug delivery systems involving rapidly degradable polymers, the intrinsic immunogenicity varies with the physicochemical form of the particles, maximum stimulation occurring soon after admin-

6. EFFECTS OF MICROSCALE AND NANOSCALE CHARACTERISTICS ON BIOCOMPATIBILITY It is necessary to briefly summarize the role of scale, particularly the microscale and nanoscale, on biocompatibility pathways. Although some attention has been given to the scale of internal materials structures (e.g., “nanocrystalline” or nanograin size), there is insufficient evidence of any such features having roles in biocompatibility pathways to give consideration here. This 21


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be modified where the matrix has a fibrillar nanoarchitecture, which can be manipulated to control differentiation.279 These effects can be enhanced when the nanofibrillar structure is derived from self-assembling supramolecular structures with peptide sequences that are able to provide chemical cues as well as mechanical stimulation.280 The conceptual and practical differences between 2D topography, 3D topography, and 3D architecture are potentially important. In practical applications, it is the 3D architecture that controls the host responses, whether that relates to stem cell differentiation within a nanofibrillar structured gel or the cellular invasion and ECM formation within a mesh of otherwise porous biomaterial. Interesting work on the host responses within porous materials indicates that macrophage phenotype can be influenced solely by the pore size, where apparently a uniform interconnected porosity of 34 μm produces the optimal response.281 2D topography is far less relevant in relation to cellular behavior in tissue environments, with 3D topography somewhere in between. 6.2. Particles. The processes by which polyethylene microparticles and cobalt alloy nanoparticles in the form of wear debris in joint replacements influence the host response have already been discussed. Particles may influence biocompatibility in situations like these where they are released from devices by wear or degradation processes or when the product is itself in particulate form, as with drug delivery systems or contrast agents; the manner in which they influence the host response is therefore of some importance. There may be influences of size, shape, volume, chemistry, mode, and site of entry on the response, but generally the mechanisms will be in line with those already discussed in sections 3, 4, and 5. The main considerations are those of size and shape with respect to processes of interactions with cells.282 Microparticles in the size range 1−10 μm can be taken up only by cells that have phagocytic capability, whereas smaller particles can be taken up by cells by a variety of mechanisms, with no limit to the type of cell; this is one of the reasons for the far greater interest in nanoparticles for delivery of a pay-load to a variety of cell types. Microparticles have difficulty in crossing most biological barriers and tend to remain at the site of delivery, in contrast to nanoparticles, which can be cleared from most sites very quickly.283 Mitragotri published extensively on microparticle size a decade ago.284−286 Polymer particles of diameter 2−3 μm exhibit maximum phagocytosis and attachment, it being postulated that the typical ruffled nature of macrophage membranes allows for greater contact between particles and the membrane at this size. Pacheco et al. agree that this is the range of maximum internalization when opsonized particles are investigated, and note that maximum Fc ligand density is found at this range.287 Shape and surface texture also affect the process; Vaine et al. have shown that such effects are associated with NLRP3 inflammasome activation and the secretion of IL1β.288 The fate of nanoparticles has already been addressed in sections on mechanotransduction, protein coronas and host responses to metal wear debris. This subject has recently been reviewed by Gustafson et al.289 Nanoparticles are internalized by variations of pinocytosis mechanisms. Clathrin-mediated endocytosis is responsible for internalization of particles in the range 100−350 nm and caveolin-mediated endocytosis takes place with particles of 20−100 nm. A variety of PRRs are able to recognize patterns in tissue associated nanoparticles, involving the DAMPs processes already discussed. TLRs,

section is concerned with mechanisms related to topography and particles. 6.1. Topography and Architecture. The roughness of a biomaterial surface has often been considered as a determinant of the host response. For many years, surface finish at the microscale was thought to be the most relevant. Manufacturers of dental implants, discussed in an earlier section, have introduced several different surface technologies in attempts to improve osseointegration, such as grit-blasting, acid etching, and anodization. These give varying levels of roughness at the micron level, but generally it is difficult to show any clinically relevant difference between them.268 Indeed a recent metaanalysis of the literature indicates few statistically significant differences between commercial implants with different degrees of “microscale-roughness”.269 There are no other significant examples of the use of 2D microscale roughness used to influence the host response to medical devices and it is concluded that there is no relevance to biocompatibility pathways, and indeed no mechanisms by which such effects could take place. The situation with nanoscale topography is likely to be different, although the evidence, obtained largely through in vitro and in vivo studies rather than from clinical experience, is equivocal. With respect to the dental implant model, several in vivo studies have shown that nanostructured surfaces promote osseointegration and reduce inflammation compared to microstructured devices,270,271 with reduced presence of inflammatory cytokines such as TNF-α and higher levels of markers of osteogenic activity. Biggs et al. have reviewed potential regulators of cell behavior on nanostructured surfaces,272 and concluded that effects are related to modulation of focal adhesion formation, cytoskeletal development and cellular spreading and subsequent functional differentiation of cells through integrin-specific signaling pathways. The authors go on to say that because a diverse variety of signals affect cellular differentiation, it is likely that no single signaling pathway is responsible for regulating adhesion mediated cell function in these situations. There is no reason to assume that new biocompatibility pathways are involved, just a variation of mechanotransduction processes. It is interesting to note that nanopatterning also appears to reduce both inflammation and fibrosis in soft tissue applications, including devices in the brain.273,274 With stem cell differentiation, there is stronger support for an influence of nanotopography, although again the multiplicity of factors, both concerning the parameters of nanotopography, the plasticity of stem cell behavior, and the effects of biochemical agents, means that the situation is far from clear. There are, in fact, very few studies of stem cells on 2D nanostructured surfaces, but mechanisms of such effects reflect similar pathways to those discussed previously with respect to the effects of substrate stiffness.275 There are influences of the size and shape of the nanostructure features, and their spacing and periodicity, which control orientation and contact morphology on the surfaces.276,277 As discussed by Das and Zouani,278 transitioning from considerations of 2D structure to 3D in relation to stem cell behavior is not trivial, but is of immense importance in terms of practical tissue engineering applications. The effects of stiffness are usually different when the cells are contained within a gel substrate and cell shape can be significantly altered. Cells normally adopt an apical-basal polarization on 2D substrates but stem cells will not do this in 3D matrices. Such features will 22


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Figure 6. Indirect mechanisms that can lead to genotoxicity. Nanomaterials may result in oxidative stress or inflammatory responses that in turn have the potential to damage DNA and alter transcriptional patterns. Reproduced with permission from ref 290. Copyright 2009 Elsevier.

essentially immobile microparticles and mobile nanoparticles that may be able to translocate through the body and become internalized within certain cells. The role of inflammasomes during these processes and the possibility of cytotoxicity and genotoxicity in association with internalized particles have to be taken into account.

mannose receptors, scavenger receptors and Fc receptors appear to be primarily involved. Once internalized, nanoparticles may suffer a variety of fates, most of which concern functionality rather than biocompatibility. However, the potential for both hypersensitivity and genotoxicity,290 the latter possibility being shown in Figure 6, have to be considered. 6.3. The contribution of scale to biocompatibility pathways. It might be expected that the scale of a biomaterialbased product should have some influence on the host response, especially when considering that such products range from large tissue-replacing medical devices to injected nanoparticles. This may be borne out in practice to some extent, but that does not imply that there will necessarily be scale-related differences in biocompatibility pathways. With respect to the morphology of the interface between biomaterial and host, the 2D interfacial shape plays little part in these pathways until the topographical features are measured on the nanoscale, where interactions between pattern features are of a similar size to crucial biological elements, giving effects related to modulation of focal adhesion formation, cytoskeletal development and cellular spreading and subsequent functional differentiation of cells mediated through integrin-specific signaling pathways. The influence of changing from a prescribed 2D or 3D topography to a 3D architecture is very important. At the microscale this involves porous structures, while at the nanoscale, this primarily involves gel nanoarchitecture, which certainly can influence stem cell differentiation. Many of the features of scale-influenced biocompatibility phenomena relate to mechanostransduction, involving mechanisms already discussed. With particles, the same biocompatibility pathways that have been described, especially a combination of mechanotransduction and BISI will apply, the main observations relating to the differences between

7. EFFECTS OF PHARMACOLOGICAL ACTIVITY ON THE BIOCOMPATIBILITY OF BIOMATERIALS The scene has now been set to bring together the mechanisms and influences elaborated in the last four sections into a generic framework of biocompatibility pathways. Before doing so, however, it is necessary to consider, briefly, how some other factors, not necessarily biomaterials related, may significantly affect the host response and clinical outcomes. In this section, the influence of pharmaceutical agents is discussed. 7.1. Clinical Evidence of the Amelioration of Undesirable Host Responses and Enhancement of Appropriate Responses. There are already a number of examples where pharmacological agents are used in conjunction with biomaterials to modify host responses. It is emphasized here that this is not concerned with drug delivery systems per se, nor does this address the control of infection. Also, this does not include the use of biologically active agents alongside scaffolds in tissue engineering because the modification of biocompatibility is not the objective here. Two areas are briefly discussed here. Drug Eluting Stents. The most obvious example is the drug eluting stent, where various stent platforms (stent, coating, drug) have been in clinical use for more than a decade.291 The objective is to interfere with the mechanisms of inflammation and neointimal proliferation. From the early uses of angioplasty and stents, there have been concomitant pharmacological therapies to control some of the adverse effects associated with 23


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ACS Biomaterials Science & Engineering these methods of percutaneous coronary interventions.292 Initially with bare metal stents, patients were often given large doses of heparin and overlapping oral anticoagulation. This in turn resulted in significant morbidity and mortality associated with hemorrhagic complications; following this, dual antiplatelet therapy (DAPT) with aspirin and an ADP-receptor inhibitor was introduced, reducing thrombosis and bleeding complications. Nevertheless, the acute vessel injury associated with the intervention and the resulting in-stent restenosis, lead to the introduction of drug-eluting stents, where the localized delivery of a pharmacological agent could potentially improve the biocompatibility of the device and reduce complications. The platforms have involved a group of alloys for the stents, including stainless steel, cobalt−chromium, and platinum− chromium; the polymers include poly(ethylene vinyl acetate), poly(n-butyl methacrylate), poly(vinylidenefluoride-co-hexafluoropropylene), and poly(styrene-b-isobutylene-b-styrene),293 and active agents everolimus, sirolimus, zotarolimus, and paclitaxel; there are more than 60 types of platform with regulatory approval in Europe so that the number of platforms is constantly increasing. It is noted that recent developments include the use of totally bioresorbable polymeric stents with or without drug elution;294 long-term biocompatibility is not yet clear. The mode of action of these agents used in drug eluting stents does vary but they are primarily of immunosuppression, anti-inflammatory and antiproliferation nature. For example, sirolimus binds to the immunophilin FK506-binding protein 12 (FKBP12), which is upregulated in neointimal SMCs, and the sirolimus-FKBP12 complex then inhibits the growth-regulating mTOR. Rapamycin also forms a complex with FKBP12 but affects mTOR in a different way, resulting in variations in pharmacokinetics and toxicity. The C-terminus of mTOR contains the FKBP12−rapamycin binding domain. Because mTOR is involved with a crucial event in the cell cycle, the transition between the G1 and S phase in which DNA replication occurs, rapamycin is seen to have a cytostatic effect, inducing cell cycle arrest in late G1 phase.295 Pharmacological Effects of Bone Formation. The use of biomaterials in many orthopedic, maxillofacial and spinal applications is predicated on an ability to encourage bone healing or regeneration or discouraging resorption. In a few situations, these properties are enhanced by the use of a pharmacologically active agent, either systemically or released from the device itself. Two agents are of primary interest, a bone morphogenetic protein (BMP) and bisphosphonate. Several growth factors are known to promote bone healing, including BMP-2, BMP-7, TGF-β, PDGFs, and FGFs. Regulatory approval has focused on the BMP family, and especially on the recombinant version of BMP-2, rhBMP-2. This has two roles in bone formation in these situations, the recruitment of endogenous stem cells from adjacent issues into the environment of the device and then the direction of differentiation of these cells.296 Ever since a product involving a rhBMP-2 soaked collagen sponge contained within a titanium cage was given approval for clinical use for fusion in the lumbar spine in the US in 2002, there have been very many publications dealing with clinical outcomes and complications,297−299 but very few dealing with the mechanisms by which rhBMP-2 exerts effects on bone formation. The consensus of opinion300−302 is that BMP-2 operates through receptors composed of two types of serine-kinase receptor chains BMPR 1A and BMPR 1B. rhBMP-2 has a high affinity

for these receptors which then transmit signals through both Smad-dependent and Smad-independent (including MAPK) pathways, a major effect being activation of Smad 1/5/8, which leads to increased osteogenic differentiation. It has been known for some time that bisphosphonates, which are nonhydrolyzable analogs of pyrophosphate, can inhibit bone resorption.303 With respect to biomaterials and biocompatibility, the main relevance of bisphosphonates is their use as coatings on devices to enhance bone healing.304 There are several forms of bisphosphonate, largely depending on the nature of the side groups attached to the molecule’s central carbon atom (usually −OH, −Cl, or −H) and on whether the molecule contains nitrogen. The most widely used products are alendronate and zoledronate, both of which contain −OH groups and nitrogen. These molecules target the mevalonate pathway in bone remodeling; this pathway is required for prenylation and targets proteins such as Rho and Ras, which are important for cytoskeletal organization and cell morphology. Without prenylation, the activity of osteoclasts is decreased and apoptosis occurs. Thus, bisphosphonates can be used to upset the osteoblast - osteoclast balance and improve bone formation and implant fixation, with a number of coating technologies being available.305 It should also be noted that bisphosphonates have been shown to enhance the proliferation of bone marrow stromal cells and initiate osteoblastic differentiation.306 7.2. Prospects for Pharmacologic Control of Biocompatibility. The two examples given in section 7.1 could be considered rather crude, even if clinically quite successful, attempts to enhance the biocompatibility of some widely used devices. There are still problems with dosage and side effects associated with the very high doses of powerful agents. Although the observed beneficial effects are consistent with the mechanisms of biocompatibility pathways discussed in this paper, the use of these agents has not been designed with these mechanisms in mind. There are some recent examples of pharmacologically active agents being used with a closer connectivity to these mechanisms. For example, two recent studies have targeted the EMT in attempts to minimize PCO in cataract surgery. Tan et al. have shown that the molecule Spry2 blockades Smad 2 and ERK1/2 pathways and inhibits TGFβ induced EMT, potentially improving the biocompatibility of intraocular lenses and reducing the incidence of PCO.307 Tiwari et al. targeted the fibronectin type III repeats in tenascin-C, an ECM protein that is critical for EMT using the single-chain variable fragment TN64, a process that is mediated by Smadindependent integrin β-catenin-FAK signaling pathway.308 These types of advance reflect the considerable efforts to target the inflammasome in some chronic inflammatory diseases, as discussed by Ozaki et al., who particularly address the possibility of targeting NLRP3,309 by Ravindran et al., who target inflammasome activation in gut inflammation310 and by Paul-Clark et al., who discuss targeting the various signaling pathways and ligands associated with PRRs in the context of DAMPS and sterile inflammation.311 There are some indications that the field of biomaterials and medical technology are moving in this direction, with notions of “biomaterials reprogramming the microenvironment”312 and “implants in personalized medicine”.313

8. EFFECTS OF CLINICAL AND PATIENT VARIABLES ON BIOCOMPATIBILITY As a final comment on biocompatibility pathways and mechanisms, questions arising with the influence of clinical 24


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Figure 7. Overarching biocompatibility paradigm. This summarizes the collection of phenomena that constitute the framework of biocompatibility pathways, starting with the three initial consequences of contact between a biomaterial and a physiological environment, then examples of the general effects within each of these three scenarios and examples of some specific mechanisms and pathways. Note that this is a nonexhaustive list of examples and that there are many situations, depicted by vertical dotted lines, where there are interactions between different mechanisms that contribute to the overall control.

sterile inflammation being up-regulated by raised levels of stress, both mechanically and physiologically. With respect to patient variables, it would appear intuitive that variables such as gender, age, body mass index (BMI), diabetes, other comorbidities, lifestyle issues (exercise, recreational drug use, smoking etc.) also influence outcomes of biomaterial-based therapies. The evidence, however, is not clear. Obesity has attracted most attention. Although a high BMI does influence the incidence of local complications in joint replacement,319,320 these are biomechanical, not biocompatibility related. Surprisingly it has been difficult to provide evidence that postoperative physiotherapy makes much difference to outcomes. It is also surprising that obesity does not significantly influence outcomes of either hernia repair321 or SUI mesh use.322 Interestingly, obesity does give higher rates of restenosis with the use of bare metal intravascular stents, but this is eliminated with drug eluting stents.323 There is little in the literature to suggest that patient variables have any significant effects on biocompatibility mechanisms. There is a growing body of evidence, however, that such variables do have some influence on sterile inflammation and especially on the activation of inflammasomes. Liu et al. report that the activation of the NLRP3 inflammasome induces vascular dysfunction in

and patient variables have to be addressed. There is no doubt that clinical skills and judgments and individual patient details have an influence, and sometimes a very significant influence on clinical outcomes. A great deal of attention has been paid to the effects of the so-called learning curve, where skills and outcomes improve over a period when a clinician is gaining experience with a technique. This has been discussed in detail with respect to many implanted devices, for example the insertion of deep brain stimulators, 314 hip resurfacing arthroplasty,315 TAVI heart valves,316 laparoscopic hernia repair,317 and cataract surgery.318 In most situations it takes 50−100 procedures before a clinician is optimally proficient. Frequently, analyses of risk factors for invasive procedures (as already discussed for intravascular stents) show that lack of sufficient clinical experience is the most significant factor. Crucially, the manifestation of deficient skills may be an undesirable host response (erosion of SUI mesh, POC with intraocular lens, peri-anastomosis hyperplasia in vascular grafts, etc.), for which the biomaterial and/or device is erroneously considered to be at fault. However, there are no new mechanisms of biocompatibility at play, the two main mediators of the host response, mechanotransduction, and 25


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ACS Biomaterials Science & Engineering obese rats,324 Shahzad et al. show that NLRP3 inflammasome activation aggravates diabetic nephropathy,325 Fairweather has discussed the effect of gender on inflammation during atherosclerosis,326 and Feldman et al. have reviewed the role of DAMPs as mediators of sterile inflammation in aging-related pathologies.327

associated with host responses to biomaterials. Mechanotransduction controls flow-dependent vascular remodeling and is primarily responsible for host responses to intravascular stents and vascular grafts. Mechanotransduction plays a major role in determining stem cell differentiation pathways within hydrogels, especially seen through the effect of hydrogel stiffness. Mechanotransduction also influences NP internalization on the basis of hardness differences between NPs and cell membranes. • It is further proposed that sterile inflammation, referred to here as biomaterials-induced sterile inflammation (BISI), is superimposed on mechanotransduction to guide and determine the balance between inflammation and fibrosis. • Central to the mechanisms of BISI are the ubiquitous damage-associated molecular patterns (DAMPs) that are initiated at the moment of biomaterial-host component contact, often considerably influenced by the traumatic events of that contact, and the activation of one or more inflammasomes, in association with pattern recognition receptors, and subsequent cytokine and chemokine activity. Pro-inflammatory, anti-inflammatory, and profibrotic pathways are all available, the orchestration of which is mediated by the nature of the DAMPs (including the generation of ROS), and the ensuing balances between MMPs and TIMPs, between ECM deposition and breakdown and crucially between M1 and M2 macrophages. Also of potential significance are epithelial to mesenchymal transformations, which can significantly alter fibroblast and, especially, myofibroblast activity. • With currently used biomaterials, the generic chemical nature (e.g., polypropylene, titanium, carbon) is not the critical determinant of the host response. Agents released form the materials (e.g., additives, degradation products, corrosion products) can constitute the irritant that initiates the DAMP. • Microtopography is of minor importance in biocompatibility pathways. Nanotopography may play some role through the modulation of focal adhesion formation, cytoskeletal development and integrin-specific signaling in the functional differentiation of cells, although these processes may be considered variations of mechanotransduction phenomena. More important than topography is material or construct architecture, possibly with 3D microscale meshes and almost certainly with 3D nanoarchitecture of hydrogels in the differentiation and function of stem cells in tissue engineering. • It is possible to modulate biocompatibility through the sustained release of biologically active agents. So far the clinically applied processes have been rather crude, with little control over pharmacokinetics and the avoidance of side effects. However, considerable opportunities arise if highly specific targets associated with inflammasomes and the inflammation−fibrosis balance are used. • Biocompatibility outcomes are often influenced by clinical skills and patient variables. However, these effects are based on changes to mechanical and physiological environments and do not involve new pathways. There are certain aspects of biocompatibility that have not been addressed in this analysis. These include the influence of infection and some potential systemic effects involving, for

9. OVERARCHING BIOCOMPATIBILITY PATHWAYS Arising from this extensive analysis of biocompatibility phenomena, which especially takes into account clinical outcomes following the use of biomaterials and related products, it is possible to identify a new paradigm that is able to generically define the mechanisms that drive the events of the host response and the pathways that determine the eventual outcomes. These have been discussed within the various sections of this paper and are brought together here. The overall scheme is summarized in Figure 7. The critical points are as follows: • As soon as a biomaterial comes into contact with components of a living system, which could involve an implanted device, a tissue engineering construct, or a drug delivery system, three events are simultaneously triggered. These are a perturbation of the mechanical environment, the perturbation of the physiological environment, and the adsorption of macromolecules of the environment (especially proteins) onto the biomaterial surface. • Under most circumstances, the protein adsorption process, and any rearrangements of the interfacial region, have only minor effects on subsequent events. • The exceptions are those situations where reorganization of the interfacial region results in exposure to conformational altered glycoproteins, especially FN, which assist in 3D ECM formation that can be beneficial to bone formation, and the formation of the PC on certain NPs, which will influence NP translocation and internalization, with potential consequential effects on nanotoxicity, nanogenotoxicity, and immune responses. • In theory, protein adsorption could have a role in biomaterial-induced thrombogenicity and complement activation through cascade processes. Serious questions have been raised over the significance of contact activation in thrombus formation, and the relevance of in vitro data to material surface−blood interactions. It is suggested that with currently used devices, hemodynamics and pharmacological control of coagulation tendencies are more important than material surface physicochemical characteristics. Complement activation may take place on biomaterials within the circulatory system but this is only clinically relevant with large surface area extracorporeal devices. • The perturbation of mechanical and physiological environments initiates the two essential (and related) biocompatibility pathways, mechanotransduction and sterile inflammation. • It is proposed here that mechanotransduction, which is concerned with the molecular and cellular processes involved with the conversion of mechanical stimuli into biochemical signals, is the primary, baseline phenomenon in biocompatibility. • Several well-understood mechanotransduction pathways, for example, the Wnt/β-catenin pathway, are also 26


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(9) Williams, D. F. Essential Biomaterials Science; Cambridge University Press: Cambridge, U.K., 2014; Chapter 8, Contemporary and Future Biomaterials. (10) Williams, D F Essential Biomaterials Science; Cambridge University Press: Cambridge, U.K., 2014; Chapter 3, Biocompatibility Pathways. (11) Iskratsch, T.; Wolfenson, H.; Sheetz, M. P. Appreciating force and shape − the rise of mechanotransduction in cell biology. Nat. Rev. Mol. Cell Biol. 2014, 15, 825−833. (12) Miller, C. J.; Davidson, L. A. The interplay between cell signaling and mechanics in developmental processes. Nat. Rev. Genet. 2013, 14, 733−744. (13) Murphy, W. L.; McDevitt, T. C.; Engler, A. J. Materials as stem cell regulators. Nat. Mater. 2014, 13, 547−557. (14) Mitchell, M. J.; King, M. R. Computational and experimental models of cancer cell response to fluid shear stress. Front. Oncol. 2013, 3, No. 44, DOI: 10.3389/fonc.2013.00044. (15) Davies, P. F.; Civelek, M.; Fang, Y.; Fleming, I. The atherosusceptible endothelium; endothelial phenotypes in complex hemodynamic shear stress regions in vivo. Cardiovasc. Res. 2013, 99, 315−327. (16) Shao, Y.; Sang, J.; Fu, J. On human pluripotent stem cell control; the rise of 3D bioengineering and mechanobiology. Biomaterials 2015, 52, 26−43. (17) Fedorchak, G. R.; Kaminski, A.; Lammerding, J. Cellular mechanosensing: Getting to the nucleus of it all. Prog. Biophys. Mol. Biol. 2014, 115, 76−92. (18) Humphrey, J. D. Stress, strain and mechanotransduction in cells. J. Biomech. Eng. 2001, 123, 638−641. (19) Huiskes, R.; Weinans, H.; van Rietbergen, B. The relationship between stress shielding and bone resorption around hip stems and the effects of flexible materials. Clin. Orthop. Relat. Res. 1992, 274, 124−134. (20) Qin, Y.-X.; Hu, M. Mechanotransduction in musculoskeletal tissue regeneration: Effects of fluid flow, loading and cellular-molecular pathways. BioMed Res. Int. 2014, 2014, No. 863421. (21) Lara-Castillo, N.; Kim-Weroha, N. A.; Kamel, M. A.; Javaheri, B.; Ellies, D. L.; Krumlauf, R. E.; Thiagarajan, G.; Johnson, M. I. In vivo mechanical loading rapidly activates β-catenin signaling in osteocytes through a prostaglandin mediated mechanism. Bone 2015, 76, 58−66. (22) Thorfve, A.; Lindahl, C.; Xia, W.; Igawa, K.; Lindahl, A.; Thomsen, P.; Palmquist, A.; Tengvall, P. Hydroxyapatite coating affects the Wnt signaling pathway during peri-implant healing in vivo. Acta Biomater. 2014, 10, 1451−1462. (23) Guo, R.; Merkel, A. R.; Sterling, J. A.; Davidson, J. M.; Guelcher, S. A. Substrate modulus of 3D-printed scaffolds regulates the regenerative response in subcutaneous implants through the macrophage phenotype and Wnt signaling. Biomaterials 2015, 73, 85−95. (24) Spatz, J. M.; Wein, M. N.; Gooi, J. H.; Qu, Y.; Garr, J. L.; Liu, S.; Barry, K. J.; Uda, Y.; Lai, F.; Dedic, C.; Balcells-Camps, M.; Kronenberg, H. M.; Babij, P.; Pajevic, P. D. The Wnt inhibitor sclerostin is up-regulated by mechanical unloading in osteocytes in vitro. J. Biol. Chem. 2015, 290, 16744−16758. (25) Delgado-Calle, J.; Riancho, J. A.; Klein-Nulend, J. Nitric oxide is involved in the down-regulated of SOST expression induced by mechanical loading. Calcif. Tissue Int. 2014, 94, 414−422. (26) Picke, A.-K.; Salbach-Hirsch, J.; Hintze, V.; Rother, S.; Rauner, M.; Kascholke, C.; Möller, S.; Bernhardt, R.; Rammelt, S.; Pisabarro, M. T.; Ruiz-Gómez, G.; Schnabelrauch, M.; Schulz-Siegmund, M.; Hacker, M. C.; Scharnweber, D.; Hofbauer, C.; Hofbauer, L. C. Sulfated hyaluronan improves bone regeneration of diabetic rats by binding sclerostin and enhancing osteoblast function. Biomaterials 2016, 96, 11−23. (27) Virdi, A. S.; Irish, J.; Sena, K.; Liu, M.; Ke, H. Z.; McNulty, M. A.; Sumner, D. R. Sclerostin antibody treatment improves implant fixation in a model of severe osteoporosis. J. Bone Joint Surg. 2015, 97, 133−140.

example, reproductive toxicology and mutagenesis/carcinogenesis. In conclusion, a system of biocompatibility pathways has been generated; these pathways are able to explain a wide range of clinical biocompatibility challenges, including nanoparticle translocation and internalization, intraocular lens opacification, leukocyte dominated responses to metallic wear debris in joint replacement, stem cell differentiation of nanostructured hydrogels, tissue responses to incontinence meshes, and restenosis of intravascular stents. Perhaps even more importantly, the identification of these molecular pathways of biocompatibility offers prospects of the control of the host response by targeting specific points in these pathways, for example, the inhibition of epithelial to mesenchymal transformation that can result in excessive fibrosis, and the inhibition of activation of the NLRP3 inflammasome following exposure to biomaterial-induced stresses; this should lead to a more effective translation of biocompatibility understanding into better clinical outcomes. As a final point, if the mechanisms and pathways described in this paper are valid, then the whole process of evaluating the “biological safety” of materials-based systems for regulatory approval has to be challenged and, indeed, changed. A series of in vitro tests concerned with leachables derived from material samples cannot begin to address the complexity of the phenomena that take place within the host; this is the main reason why devices approved for clinical use often fail to give the ultimate desired performance.

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AUTHOR INFORMATION

Corresponding Author

*E-mail: [email protected]. Fax: +1 336 713 7290. Tel: +1 336 671 8895. ORCID

David F. Williams: 0000-0001-5173-3157 Notes

The author declares no competing financial interest.

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Biocompatibility Pathways: Biomaterials-Induced Sterile Inflammation

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