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The foreign body response demystified Yashoda Chandorkar, Ravikumar Krishnamurthy, and Bikramjit Basu ACS Biomater. Sci. Eng., Just Accepted Manuscript • DOI: 10.1021/acsbiomaterials.8b00252 • Publication Date (Web): 04 Jun 2018 Downloaded from http://pubs.acs.org on June 27, 2018

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ACS Biomaterials Science & Engineering

The Foreign body response demystified Yashoda Chandorkar1, Ravikumar K1 and Bikramjit Basu1,2,* 1

Laboratory for Biomaterials, Materials Research Centre, Indian Institute of Science, C V Raman Road, Bangalore, 560012, India

2

Centre for Biosystems Science and Engineering, Indian Institute of Science, C V Raman Road, Bangalore, 560012, India

Abstract The human body is endowed with an uncanny ability to distinguish self from foreign. The implantation of a foreign object inside a mammalian host activates complex signalling cascades leading to the biological encapsulation of the implant. This reaction by the host system to a foreign object is known as foreign body response (FBR). A deeper insight into the mechanisms of FBR is needed to develop biomaterials for better integration with the living system. In the light of recent advances in tissue engineering and regenerative medicine, particularly in the field of biosensors and biodegradable tissue engineering scaffolds, the classical concepts related to the FBR acquire new dimensions. The aim of this review is to provide a holistic view of the FBR, while critically analysing the challenges, which need to be addressed in the future to overcome this innate response. In particular, this review discusses the relevant experimental methodology to assess the host response. The role of erosion and degradation behaviour on FBR with biodegradable polymers is largely explored. Apart from the discussion on temporal progression of FBR, an emphasis has been given to the design of next generation biomaterials with favourable host response.

Keywords: foreign body response, non-specific protein adsorption, anti-fouling polymers, polymer degradation and erosion *Corresponding author; Email: [email protected] 1

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Table of Contents Abstract………………………………………………………………………………………………..……………………………………….1 Table of contents………………………………………………………………………………………………..…………………………2 Introduction………………………………………………………………………………………………..………………………………..3 Defining biomaterials and biocompatibility……………………………………………………………………..……….….6 Foreign body response: Definition and mechanism…………………………………………………………….……….7 Advantages of the foreign body response to biomaterials…………………………………………….13 Disadvantages of the foreign body response…………………………………..…………………..……….15 Evaluation of the foreign body response……………………………………………………………………..…..………..18 Host model……………………………………………………………………………………………………………………18 Collagen distribution and its implications……………………………………………………..……………..19 Histomorphometric analysis…………………………………………………………………………………..……22 Cytokines………………………………………………………………………………………………………………………23 Implant related parameters………………………………………………..……………………………………….27 An insight into tissue-biomaterial interaction……………………………………………………………………..…….29 Protein Adsorption on biomaterials…………………………………………..…………………………..……29 Cell adhesion on biomaterial surfaces…………………………………………………………………………31 The critical role of water molecules at the interface………………………………………………….33 Material properties affecting the foreign body response…………………………………..………36 Material strategies to overcome the foreign body response…………………….……………………………..40 Reduction of non-specific protein adsorption……………………………………………..……..……..40 Poly(ethylene glycol) …………………………………………………………………………………….42 Zwitterionic materials…………………..………………………………….……………………………45 Implant modifications…………………………………..………………………………………….………………..47 Angiogenic drugs…..……………………………………………………………….. …………………..47 Use of steroidal and non-steroidal anti-inflammatory drugs……………..………..48 2


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Host response to some common polymers……………………………………………………………………..…….55 Late degradation tissue response……………………………………………………………………….……55 Bulk erosion and burst release of degradation products………………………………………….58 Lack of tuneable mechanical properties…………………………………………………………….…….62 Outlook and concluding remarks……………………………………………………………………..…………………….63 List of Abbreviations……………………………………………………………………..…………………… …………………66 References……………………………………………………………………..……………………………………………………….67

Introduction Recent advances in materials science pride themselves on the ability to determine the properties of a material and draw conclusions about its performance. While this is sufficiently precise and accurate in many scenarios, it fails in the context of materials which act at the interface of biological systems. This is particularly true in the case of implanted biomaterials, where the expected behaviour of the material is grossly different from the observed behaviour. Most of the time, the implant performs worse than in lab scale experiments in in vitro models.1-3 For instance, glucose sensors fail to live up to their expected performance upon implantation and show erratic responses4. Continuous monitoring of blood parameters such as glucose levels or other hormones is still not achieved due to various difficulties arising from interactions with the host. The host recognizes most sensors and implants as ‘foreign’, and the consequences which follow may be likened to ‘reflex actions’ of the immune system of the host. Many systems designed today do not live upto their expected performance as they face stiff resistance from the host, which identifies them as ‘foreign’ and hence, a potential threat. Most of the implants today are poised to perform better if this hostile response from the host can be eliminated or at the least, alleviated. In order to create a stealth material that can bypass the defences of the host body, a few serendipitous discoveries 3



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from the past prove to be a good source of inspiration. Poly(methyl methacrylate) shards that were accidentally lodged in the eyes of aviators did not create inflammation or an adverse host tissue reaction. This was instrumental in creating early intra-ocular lenses made of poly(methylmethacrylate) (PMMA) by the British ophthalmologist Sir Harold Ridley.5. Today, they have been replaced by poly (dimethylsiloxane) (PDMS) and foldable acrylate lenses. With such advances, cataract associated blindness can be avoided. Projections state that around 14 million cataract surgeries will be needed every year in 2016-2020 in India alone 6. The use of materials in medicine is not a new concept and has been recorded in ancient times as well. The Sushruta Samhita (The Compendium of Sushruta), written by Sushruta (c.a. 600 B.C.), the famous surgeon of the ancient Indus valley civilization is arguably one of the oldest evidence for the use of artificial materials to treat a medical condition. Apart from the use of bark, tendons, hair and silk as suture materials; it also mentions the use of ant-heads as a stitching material for intestines. Such examples highlight the use of an implant to compensate or retrieve the loss of an anatomical function. However, it is important to note that these materials were used without an assessment of the host response. Their long term success itself was assumed as a proof of their efficacy. Careful material selection can increase the success rates of implants dramatically. Early implants made of brittle materials like glass or metallic materials such as stainless steel failed due to material as well as design issues.7-8 The innovative use of Teflon (PTFE) acetabular cup by Dr. Charnley marked the onset of the modern age orthopaedics 9-10. Other examples of successful translations include dental implants, breast implants, the artificial kidney, pacemakers, vascular grafts and stents, heart valves etc. The sheer range of applications of these biomedical devices and their widespread use are indicated by their everincreasing numbers. These examples highlight how biomaterials have evolved and have become an integral part of our lives today. In most of these cases, the materials that were 4


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used were “off-the-shelf”, without substantial redesign to adapt to the purpose. While this is not the best approach as far as performance is concerned, it was pragmatic and was verified by experimentation in human subjects. At present, researchers take refuge in a more conscious, logical approach that assesses material properties and the host response in a hierarchical manner for the intended application. This is particularly important as an increasing number of different materials (metals, ceramics and polymers) are being used for biomedical applications. The market has been dominated by polymeric biomaterials for the last two decades. A comprehensive list of such materials has been summarized in many reviews and books in the field

11-13

. Correspondingly, the biomaterials market is increasing

exponentially with the advent of novel implant materials, drug loaded scaffolds for regenerative medicine, improved joint replacement materials, sutures and wires to name a few. In fact, the global biomaterials market was estimated to be $70 billion in 2016 and is expected to increase yearly with a compound annular growth rate of 13% to reach a net worth of $130 billion at the end of the year 202114. With such enormous growth in the industry and the applications of biomaterials, it is essential to understand the reasons for the success of certain materials and the failure of many others. One of the major factors determining this clinical success of a biomaterial is the host response, which is the focus of this review. This review elucidates the different material aspects and host response parameters that are currently thought to dictate final implant performance. Initially, the basic concepts of biomaterials and biocompatibility are defined. After providing a glimpse into the steps that lead to the foreign body response, the common methods used to evaluate it are discussed. Many strides have been taken in this field since Koshland had stated that the mechanism of recognition of the self from the non-self was one of the top most mysteries of modern science 15

. Different viewpoints have been proposed to solve this puzzle, and although it is still not

possible to completely overcome the FBR, strategies have been developed to mitigate it. 5



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Relevant fundamental aspects of material properties such as polymer degradation are highlighted at the end of the review. This review provides a bird’s eye view of the different approaches used to minimize the FBR, with a special focus on implant material, design and host considerations. Through the course of the arguments and discussion, the nexus between these traditional concepts and the emerging paradigm, which holds mechanotransduction and sterile inflammation responsible for the FBR is emphasized. Defining biomaterials and biocompatibility The field of biomaterials has evolved over the last few decades to encompass different material classes such as metals, ceramics and polymers. A very refined definition of biomaterials that is applicable to different classes of materials has been coined by Prof. David Williams in 2009 and is as follows: 'A biomaterial is a substance that has been engineered to take a form which, alone or as part of a complex system, is used to direct, by control of interactions with components of living systems, the course of any therapeutic or diagnostic procedure, in human or veterinary medicine' 16. It is important to mention here that these definitions too have evolved with time as the field of biomaterials progressed and as our understanding of material-tissue interaction improved. A recent definition is the following: 'Biomaterials are those materials — be it natural or synthetic, alive or lifeless, and usually made of multiple components — that interact with biological systems. Biomaterials are often used in medical applications to augment or replace a natural function. One of the pre-requisite conditions that a successful biomaterial ought to fulfil is that it should be accepted by the body, i.e. it should be compatible with the body. On the lines of the first principle of Hippocrates that a doctor should do no harm, the main requirement for 6


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biocompatibility of a material, whatever be its end application, is that the material should do no harm. This performance of implants, along with their interaction and co-existence with biological tissues is studied in the realm of the broad and elusive subject of biocompatibility. While it is relatively common to describe a successful biomaterial as biocompatible, the exact nature of biocompatibility is still uncertain

17

. The following definition of biocompatibility is well

accepted in the biomaterials community: 'Biocompatibility refers to the ability of a biomaterial to perform its desired function with respect to a medical therapy, without eliciting any undesirable local or systemic effects in the recipient or beneficiary of that therapy, but generating the most appropriate beneficial cellular or tissue response in that specific situation, and optimising the clinically relevant performance of that therapy' 17. In this regard, the property of biocompatibility of a material is specific to a particular application. Generally, most of the biomaterials induce an initial response from the host system upon their implantation. The response by the host system and its nuances are addressed in the later sections of this review.

Foreign body response: Definition and mechanism The implantation of any biomaterial elicits a local or systemic response from the host tissue. This response is termed as the host response. This is independent of the method by which the biomaterial is introduced in the body- by injection or by surgery, as all biomaterials cause a disruption of the local host tissue environment 18. The degree of the host response depends on the extent to which the homeostasis of the host was disturbed by the injury and the subsequent introduction of the foreign object, which is deciding factor in determining the biocompatibility of the material. Although a large number of biomaterials and medical 7



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devices have been implanted successfully in humans, a stealth material that can surpass the highly efficient human surveillance system has not been created

19

. This is triggered by

protein adsorption on the material (biofouling), leading to encapsulation of the implant by a dense collagenous capsule and impeding further interaction of the implant with the surrounding tissue. The steps that lead to the formation of this capsule will now be discussed. The different stages of FBR include injury, blood-material interactions, provisional matrix formation, acute inflammation, chronic inflammation, granular tissue development and fibrous capsule development

1, 20

. Each of these events is very complicated and involves a

complex interplay of inflammatory cells, mitogens, chemo-attractants, cytokines and other bioactive agents. The sequential steps involved in the host response to an implant are described below and are outlined in Figure 1. a)

Injury to the vascularized connective tissue activates the initial inflammatory

response. This is regardless of the tissue in which the biomaterial is implanted

21

. In fact,

inflammation is the reaction to local injury by a vascularized, living tissue. The main purpose of inflammation is to contain or wall off the injurious agent. At the same time, inflammation also initiates a cascade of events that may heal the injured tissue 5. The injury caused during implantation leads to changes in vascular flow and permeability, leading to a leakage of blood, fluid and proteins at the site of injury through a process known as exudation 5.

8


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Figure 1: A flowchart showing the sequence of host reactions following the implantation of a biomaterial/ medical device, along with a schematic representation of some of the stages involved in acute inflammation and chronic inflammation that ultimately lead to the formation of the fibrous capsule around the implant, depicted with blue strokes in the figure. Adapted with permission from Ref. 22.. Copyright 2002 Elsevier.

b)

The injury to the connective tissue leads to the development of a blood-based

transient provisional matrix around the biomaterial 20. Blood protein deposition, which occurs after injury at the biomaterials surface, may be regarded as a provisional matrix from the wound healing perspective

21

. Blood clot formation is a quick but complex process and

involves the activation of extrinsic and intrinsic coagulation systems (factors extrinsic to circulating blood and inside circulating blood, respectively

23

), the complement system

(which is a set of proteins that are activated in the presence of foreign bodies or pathogens 9



leading to the release of inflammatory and immune cells

24

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), the fibrinolytic system (which

regulates the delicate balance between homeostasis and fibrin deposition generating system (which regulates vascular permeability

25

24

), the kinin

) and platelets. The protein

cascades involve dynamic displacement of initially adsorbed proteins by more surface active proteins. This is known as the Vroman effect 26.

Although the inflammatory response is initiated by injury, the response is mediated by chemicals released by plasma, cells and the injured tissue. These include cytokines, growth factors, vasoactive agents, lysosomal proteases, oxygen free radicals etc. It is important to note that these mediators are controlled by feedback mechanisms, and that their action is predominantly local. The proteases and oxygen free radicals released by the lysosome play a key role in the degradation of a biomaterial 21. Acute inflammatory response is characterized by the presence of neutrophils (polymorphonuclear leukocytes). About a quarter century ago, Tang and Eaton conclusively proved that fibrinogen was primarily responsible for the acute inflammatory response; and highlighted, for the first time, the nexus between inflammation and wound healing/ clotting 27. c)

Depending on the extent of injury at the implant site, the acute inflammatory

response to biomaterials subsides quickly (within a week), as neutrophils disappear within 24-48 h

21

.. Figure 2 shows the temporal distribution of different cells and stages that are

involved in the host response. The neutrophils are the first line of defence of the host system.

10


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Figure 2: The temporal variation in the acute inflammatory response, chronic inflammatory response, granulation tissue development, and foreign body reaction to implanted biomaterials. Reproduced with permission from Ref.2. Copyright 2017 Cambridge University Press.

d)

Acute inflammation is followed by chronic inflammation, which is marked by the

presence of mononuclear cells (monocytes and lymphocytes) at the implant site. It is less uniform histologically compared to acute inflammation and hence is a term that is used to identify a wide range of cellular responses

28

. Monocytes migrate to the wound site and

differentiate into macrophages. They constitute the second line of defence of the body 29. The macrophages release mediators of inflammation such as oxygen free radicals and reactive oxygen intermediates, degradative enzymes and acids on the surface of biomaterials 20. They also adhere to the material and attempt to phagocytose it. They are unable to digest or phagocytose a large mass such as an implant even after a final desperate attempt, which is known as 'frustrated phagocytosis'

22

. Macrophages frequently undergo fusion to improve

their efficiency and form multi-nucleated giant cells, known as foreign body giant cells 11



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(FBGCs). It is believed that these FBGCs send out signals for fibroblasts to arrive at the injury site 22. Recently, Friedl and co-workers used infra-red excited non-linear microscopy to demonstrate that immigrating macrophages are immobilized at the site of the implant, and form FBGCs. Macrophages (M1 type, i.e. pro-inflammatory) and FBGCs release vascular endothelial growth factor (VEGF) which aids in the development of an immature vascular network 30. Another type of macrophage population, which is distinct from the inflammation recruited monocyte derived macrophage population (blood derived) are tissue resident macrophages (probably thought to be derived from the embryo). These are a heterogeneous group of macrophage populations, which differ substantially in different tissues and contribute to a more tissue specific response31-32. Although the recruited monocytes often exceed the tissue-resident macrophages in their absolute number, they play a critical role in recognizing initial damage to the tissue and by triggering the initial pro-inflammatory response that results in an influx of monocytes 33. e)

The development of granulation tissue is also a part of chronic inflammation

21

.

Fibroblasts and vascular endothelial cells arrive at the implant site, signifying healing. Histologically, granulation tissue is characterized by the presence of fibroblasts and new blood vessels. Small new blood vessels are formed from branching of pre-existing vessels by a process known as neo-vascularization. Fibroblasts are active in synthesizing collagen. Granulation tissue gets its name from the characteristic pink, granular appearance on the surface of wounds usually seen after 3-5 days, depending on the extent of injury 21, 34.

12


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Figure 3: A schematic representation of the host response upon implantation of a biomaterial. Reproduced with permission from Ref. 35. Copyright 2013 Springer Nature.

The foreign body reaction consists of foreign body giant cells and granulation tissue along with inflammatory cells such as macrophages and fibroblasts, is summarized in Figure 3. Due to the collagen deposition by the fibroblasts, the implant is effectively isolated from the rest of the body. This condition, in which the fibrous capsule encapsulates the implant, along with the interfacial foreign body reaction is known as fibrosis 21. The host response to an implant is evaluated by morphological and histological observations of the implant site and relates to the biocompatibility

34

. The intensity and time scale of the response are dependent on the

extent of injury created and on implant as well as host parameters, which will be addressed in later sections.

13



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Advantages of the foreign body response to biomaterials In brief, after implantation, all foreign materials are surrounded by collagen deposition in the form of a fibrotic scar at the material-tissue interface as shown in Figure 4. Thus, the goal of biological response evaluation is to predict whether an implant presents potential harm by evaluating its performance in conditions that simulate clinical use 21.

Figure 4: A schematic depiction of the collagen deposition encapsulating an implant/ biomaterial after implantation. Also seen are the inflammatory cells at the material/ tissue interface. (Not to scale)

At this juncture, it may be worthwhile to cast a quick glance into the purpose of the foreign body response that addresses the moot question- Why has the foreign body response evolved? First and foremost, it is important to understand that immune and inflammatory reactions occur in order to protect the body from foreign objects

36

. In a world full of pathogenic

bacteria and microbes, the host innate immune response has evolved over millions of years to protect the body against such disease-causing and unknown substances. The importance of this impeccable ability of the human/mammalian body to distinguish itself from foreign bodies has allowed mankind to survive and evolve over the years. In some cases, it is also advantageous to patients. For instance, lead poisoning was rarely observed in patients who had retained lead bullet fragments. The development of an impermeable collagen capsule around the bullet fragments is thought to be the likely reason behind this

19, 37

. Since these

mechanisms have evolved to handpick bacteria and viruses, they also deal with synthetic 14


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materials which resemble such pathogens in size as well as chemistry 38. A major fraction of literature on the foreign body response is focussed on strategies to mitigate it. However, Sparks was the first to report the use of autologous in situ tissue engineered vascular grafts, obtained by subcutaneous implantation of silicone rods

39

. These were later used as bypass

grafts. This was one of the first approaches, in which the body itself was used as a bioreactor. Later on, many attempts were made to generate vascular grafts using the foreign body response of the host, and Campbell et al. successfully demonstrated this in dog model

40-41

.

The interested reader is referred to an excellent review on this topic 41. The bone implant that was developed by Masquelet also takes advantage of the fibrous tissue generated in response to a poly(methacryl methacrylate) (PMMA) cement spacer used to treat large bone defects 42. After 6-8 weeks, when the fibrous tissue around the implant is well developed, the spacer is replaced by cancellous bone autografts from the iliac crest. Here, the periosteal fibrous response to the PMMA spacer is used as a protective envelope when the autograft is implanted, and autograft resorption is prevented. Thus, Masquelet et al. effectively demonstrate a case where the innate foreign body response can be used to treat large bone defects. Disadvantages of the foreign body response Although most biomaterials elicit a foreign body response and implant encapsulation, their performance is acceptable, although not optimum in many cases. However, in some cases (which are elucidated later), the foreign body response has deleterious consequences. The immune system of the host recognizes the implants as foreign, and isolates them from the rest of the body. This is achieved through a set of events mentioned earlier, that leads to implant encapsulation by a dense collagen capsule

43-44

. In recent years, many novel implantable

devices such as joints, blood vessel substitutes, sensors, hernia meshes materials, heart 15



valves, cosmetic and reconstructive implants, and artificial organs have been developed

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19

.

For some of these applications, the formation of the collagen capsule is a liability 19. Also, in spite of the advancement in tissue-based sensors and scaffolds for tissue engineering, their optimum function is impaired by the foreign body response. The impermeable nature of the fibrous capsule results in poor mass transport and electrical communication between the implant and the tissue, thus impairing the intended implant function. Traditionally, an implant was considered to be biocompatible if it was encapsulated by an avascular layer of collagen without adversely affecting its intended performance in vivo

45

.

However, this is not desirable as far as sensors, tissue engineering scaffolds and controlled release devices are considered

22

. The foreign body response, which results in implant

encapsulation is highly detrimental to optimum implant function. This is because of the insulating nature of the fibrous capsule, which impedes the efficient interaction between the implant and the body 43 and is the fate of all materials that are implanted in the body. Some orthopaedic pathologists argue that this does not occur at the bone-metal interface for certain metallic materials (stainless steel and titanium) and hydroxyapatite based ceramics due to their direct contact with the host bone, but other researchers have shown that such materials also cause a host response46-47. Such fibrotic reactions are undesirable for next generation biomedical applications such as electrical sensors, drug delivery devices and tissue engineering scaffolds, which unlike traditional biomaterials, are sensitive to the surrounding cellular micro-environment and respond accordingly. In fact, fibrotic reactions are devastating as far as tissue integration with the host and nutrient supply are considered 48. For instance, although a significant progress has been made in real time monitoring of analytes, a glucose sensor for treatment and management of diabetes is still not realized due to the interference of the foreign body reaction to the implanted devices. The readers are 16


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referred to detailed reviews on implanted sensors in the context of the FBR here

49-50

. The

sensor loses its functionality with time, and it is evident that this is due to the fibrous reaction that occurs at the tissue-implant interface

51

. Currently, the longest time a biosensor that is

FDA approved and is functionally active biosensor is only 7 days in vivo 36. Also, it is highly inaccurate and suffers from high fluctuations and loss of sensitivity over time, making it inappropriate for use in humans

52-53

. The foreign body reaction is also the main reason why

no tissue engineering product based on encapsulation of cells by a polymer has seen the light of the day. These cell loaded constructs are encapsulated by the fibrous capsule in the in vivo situation

54

. Apart from this, the fibrous capsule can also cause tissue distortion and pain,

most commonly observed in the case of breast implants55 and hernial meshes 17, 44, 56-57. Some of the well-known examples such as intra-ocular lens and hip prostheses are of a permanent nature and they are implanted with the intention that they remain in the body and perform their function for a long time (the longer, the better). In other words, they are based on materials that are not biodegradable. With the advent of tissue engineering, there is a demand for next-generation materials which can fulfil the demanding requirements such as biocompatibility, degradability and tuneability for such novel applications. A step further would be the class of materials with bioresorbable property, which are gaining importance for applications in stents, drug delivery vehicles and scaffolds.58-59 While biodegradeable materials diminish (degrade) over the duration of their application and are removed from the host system, bioresorbable materials are those whose degradation products are assimilated by the host through the process of metabolism to serve a particular function as a part of the medical therapy of the implant/scaffold 60.

17



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Evaluation of the foreign body response Having realized the importance of FBR in the context of biomaterials and their applications, it is necessary to develop a better understanding of the biology of the foreign body response for next-generation applications such as sensors capable of continuous monitoring of different analytes and biodegradable scaffolds for tissue engineering. The foreign body response, and in particular, the fibrous capsule represents a Herculean barrier between the implant and the analytes (literally and figuratively).

Host model The foreign body response is best assessed in a suitable host model, which is generally decided based on the application. For instance, subcutaneous implantation is commonly performed in a rat or a mouse model, while the porcine model is most commonly used today to study the host response to cardiovascular valves, as their vasculature is similar to humans 61

. From an anatomical and physiological perspective, the biological architecture in primates

closely resembles that of humans and they are the preferred models for long term in vivo studies. However, strict ethical considerations limit their usage in many countries around the world. An ideal animal model is cost-effective and can develop the desired pathology very quickly. The murine model is the most suitable for studying signalling pathways as transgenic and knock-out mice are easily available. Diabetic mice are commonly used to check the encapsulation efficiency of transplanted and encapsulated islet cells

62

. Hackam and

Redelmeier conducted a survey to find out how many animal trials reported in literature actually translated into human clinical trials. Their sample space included publications from journals, which regularly publish animal work. However, their conclusion was that only ~33 % of all positive animal studies actually translate into human clinical trials

63

. While it is 18


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convenient to use mouse models for studying FBR and related phenomena, it is not possible to translate all studies into human clinical trials due to fundamental pathological and molecular pathway differences between humans and smaller sized animals. Any comparison or conclusion drawn from such models may not be directly applicable to humans, which may also account for the low success rates of pre-clinical trials translating into clinical outcomes. The foreign body response was recently evaluated in non-human cyanomolgus macaque monkeys, as it resembles humans more closely 62. Some in vitro models wherein the surface properties such as the topography, wettability, roughness and chemistry are varied are also used to evaluate their effect on inflammatory cells such as machrophage attachment and fibroblast morphology 64. Such in vitro models are effective in screening materials, which can be tested later in vivo. Due to the complex nature of the foreign body response, it has so far not been possible to completely eliminate the need for the host and to develop an in vitro model for its evaluation, but several co-culture in vitro models have been developed to attempt this, albeit with notable shortcomings65. Collagen distribution and its implications One of the commonly used techniques to evaluate host response is by determining the collagen distribution in the vicinity of host-implant tissue interface. This involves sacrificing the host and extracting the implant at the desired time point. The fibrous encapsulation as well as the presence of inflammatory cells surrounding the implant is quantified. The distribution of collagen around the implant is generally evaluated from the Mason’s Trichrome stained (MTS) sections, where the cytoplasm, collagen and nuclei are stained red, blue and black, respectively. Contrary to H&E staining (which stains collagen, muscles and cytoplasm, all in pink), MTS exclusively stains collagen blue and hence, can be used to effectively quantify collagen. The thickness of the collagen encapsulation around the implant 19



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is an effective parameter to assess the degree of the foreign body response. The collagen distribution around the implant may be measured by calculating the blue pixel coverage using ImageJ software in the Mason's Trichrome stained section. Consequently, a thin layer of collagen is desired over a thick layer. Tissues surrounding the implant are usually considered to be a part of the fibrous capsule if they meet the following criteria: a) The collagen deposition runs parallel to the implant contour, and the cells and nuclei are embedded in the collagen deposition b) The nuclei and the cells are usually aligned parallel to the collagen distribution.66. Sometimes, voids without any tissue may be present at the implant-tissue interface. These could be either due to an artefact during histology sample preparation or may reflect tissue distribution. In several pre-clinical studies, it has been observed that capsule thicknesses and collagen densities are highly variable around any single implant, so multiple careful measurements and assessments of possible histology and harvesting artefacts that break or distort the capsule and distort thickness and density are advocated. Although determining the encapsulation around a large implant is relatively straight-forward, it can be quite challenging if the implants are small (e.g. fibres) or are porous and soft such as in the case of soft polymer foams or proteinaceous biomaterials. At this juncture, it would be of interest to know how a diffuse pattern of collagen deposition can aid in better communication and transport between the implant and the body. In an elegant series of experiments performed by Sharkawy and colleagues, the diffusion properties of the fibrous capsule have been investigated. According to their diffusion model predictions, a 200 µm thick capsule increases the time required for diffusion for a small molecule such as glucose threefold. Hence, they concluded that the fibrous capsule poses a serious barrier to 20


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diffusion even for small molecules

67

. For larger molecules, this barrier will be even more

pronounced. Researchers such as Woodward et al. have hinted that a thick, but granular and well vascularized capsule can permit diffusion to a greater extent

68

. Such a capsule is

preferred over a thin, avascular and tightly packed capsule. However, in cases where permeation of the drug through the fibrous capsule of the implant is unhindered such as in the case of Levonorgestrel impregnated subdermal Norplant implants, the issue of avascular capsule is of lesser significance69-70.

Figure 5: a, b) MTS stained images of poly(hydroxy ethyl methacrylate) and poly(carboxy betaine methacrylate) showing a clear difference in the distribution of collagen, scale bar = 100 µm,(c,d) Immunofluorescence images of macrophages, 4 weeks after implantation in a mouse model. Red: pan macrophage marker, green-pro-inflammatory TNF α marker, blue: anti-inflammatory IL-10 marker. Clearly, more anti-inflammatory cytokines are visible in the tissue surrounding the PCBMA implant, scale bar = 40 µm. (Panels a, b, c, d reproduced with permission from Ref.56. Copyright 2013 Springer Nature.)

This is better illustrated with an example from literature. The biocompatibility of hydrogels of poly(carboxybetaine methacrylate) hydrogels was evaluated by Zhang et al. in a mouse 21



Page 22 of 77

model. Subcutaneous implantation was performed using poly(hydroxyl ethylmethacrylate) (PHEMA) as a control

43

. The results showed a normal distribution of collagen around the

poly(carboxybetaine methacrylate) (PCBMA) hydrogels, whereas PHEMA showed a typical insulating collagen capsule with high density near the implant (Figure 5). Zhang et al. have thus shown that the collagen deposition around the PCBMA implant resembles the natural extracellular matrix. Histomorphometric analysis As mentioned earlier, inflammatory cells are key drivers of the foreign body response. The presence of these cells is generally monitored over time to gain understanding into the progression of the foreign body response. The most common way of achieving this is by haematoxylin and eosin (H&E) staining of the tissue surrounding the implants. This helps in the visualization and identification of different inflammatory cells, such as neutrophils, macrophages, and fibroblasts, which differ in morphology. The thin sections used for histomorphometrical analysis are usually assessed using image analysis software. The tissue response is measured by quantifying different parameters such as the density of inflammatory cells (neutrophils, monocytes/macrophages, fibroblasts), blood vasculature and distribution of collagen surrounding the implant. Neovascularization is calculated by measuring the total number of blood capillaries per unit area of the fibrous capsule. The inflammatory cells are quantified and compared to that of a control. The control sample is generally a material whose response is known beforehand. For instance, PLGA may be a suitable control for studying and comparing the host response to new biodegradable polymers as it is one of the most common biodegradable polymers. The quantification of the thickness and the distribution of collagen around the implant is a very common method to investigate the host response. The presence of blood vessels in the encapsulation is also an important indicator, as 22


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it is known that a vascularized capsule is more permeable to molecules. These are commonly stained with endothelial cell markers such as MECA 32 antibody 56. This is to ensure that the blood vessels are accurately identified, and artefacts which mimic the capillary may be produced during histology sample preparation do not appear as false positives. In general, functional angiogenesis and effective blood and oxygen/ nutrient perfusion are more important metrics than mere stained circular vessel-like cross sections that may or may not be functional. Cytokines Macrophages act as a live surveillance system in the body and originate from circulating monocytes. Macrophages secrete a variety of protein molecules called cytokines that mediate immunity and inflammation. The word ‘Cytokines’ is derived from two Greek words ‘cyto’ and ‘kinos’ meaning cell and motion, respectively. Cytokines are a class of signalling molecules that include (a) Monokines (cytokines made by monocytes), (b) Chemokines; cytokines with a chemotactic effect and (c) signalling molecules between leukocytes

Interleukins: Cytokines which function as

19

. They are released mainly by immune cells that

facilitate cell-to-cell interaction during immune response and induce cell movement towards the site of injury, inflammation or implantation2. Cytokines usually work in a cascade manner with the release of one cytokine stimulating the cells to release additional cytokines and uncontrolled or excessive release of cytokines can eventually lead to “cytokine storm”, with drastic consequences to the host 71. Cytokines exist in peptide, protein and glycoprotein forms and are the soluble factors that mediate acute and chronic inflammatory responses. Their levels give a fair idea of the immune response to the implant. For example, if the levels of IL -1β remain high even after 14 days of implantation, it indicates a protracted inflammatory response

72

. Since cytokines 23



Page 24 of 77

act over short time spans in a localized area, their concentrations are many folds higher near the implant than in the blood. It is also well known that macrophages respond to combinations of stimuli and form complex functional phenotypes 73. Gordon and co-workers have extensively studied this and aspect and have proposed that classifying macrophages into pro-inflammatory (M1) type and anti-inflammatory (M2) type is a very simplistic approach, which may not reflect the real situation of the existence of mixed phenotypes

74

. However,

this is still a common approach used to evaluate the foreign body response. The macrophage behaviour and the determination of their phenotype are closely associated with the surface chemistry of the implant, as shown conclusively involving various substrate models and surface modifications by Anderson and co-workers. 75-79 Recently, Zhang et al.56 have used a triple label immunofluorescence method to reiterate this surface dependant macrophage behaviour on novel zwiterrionic hydrogels (Figure 5c, d). In particular, they showed that more

macrophages

containing

anti-inflammatory

markers

were

found

near

poly(carboxybetaine methacrylate) (PCBMA) hydrogel implants, which showed a marked decrease in the foreign body response and more vasculature as compared to the poly(hydroxyl ethylmethacrylate) (PHEMA) control. On the other hand, the macrophages near the PHEMA control expressed pro-inflammatory markers and also showed a higher foreign body response when implanted in mice for 3 months.56 Some of the well-known cytokines and their main functions are discussed here. Tumour Necrosis Factor (TNF) TNF also known as TNF-α is a class of cytokines derived from monocyte and is a proinflammatory cytokine which can cause cell apoptosis or necrosis and is associated with a many of effects such as tumour regression, cell proliferation and differentiation. Continued release of TNF in human hosts can cause immune responses like inflammation, cachexia 24


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(wasting syndrome) and fever while higher levels of TNF in the system results in septic shock, leading to multiple organ failure and death80. It is mostly expressed in macrophages, neutrophils, Natural Killer (NK)-cells, activated lymphocytes and endothelial cells and is connected to viral infections, tissue injury and DNA damage. Interleukin-1β (IL-1 β) Interleukin-1 is also known as lymphocyte activating factor, as it is involved in the initiation of immune responses such as inflammation and associated temperature rise81. Similar to TNF, IL-1β is a pro-inflammatory cytokine activated by macrophages and is involved in cell proliferation, differentiation and apoptosis. IL-1β is expressed by injured cells, macrophages, neutrophils and NK-cells. These cytokines also function in a concerted manner with TNF to increase the immune response of the host to a foreign body. Interleukin- 10 (IL-10) Interleukin-10 is an anti-inflammatory cytokine, which mitigates the activity of proinflammatory cytokines. It is also known as cytokine synthesis inhibitory factor (CSIF) due to its anti-inflammatory activity. Excess levels of IL-10 can cause immuno-suppression and compromise the ability of the host to defend itself against infections. Hence, a balance between the levels of pro-and anti-inflammatory cytokines is the ideal response to an implant or a foreign body in the host. It inhibits the activity of NK cells and macrophages which generally express pro-inflammatory cytokines and acts as a regulatory influence on the immune response of the host.82 Transforming Growth Factor- β (TGF-β) TGF β is a multifunctional polypeptide cytokine, which is involved in cellular processes such as proliferation, differentiation and apoptosis. Many cells and most leucocytes express TGF25



Page 26 of 77

β, which plays a vital role in modulation of the immune response of the host. TGF-β also has an inhibiting effect on macrophages and monocytes that usually express pro-inflammatory cytokines83. It is also known to stimulate the formation of ECM in the post-inflammatory stage of the immune response of the host body and is also connected to wound healing mechanisms at the site of injury84. Cytokine levels are generally evaluated by performing an ELISA assay on the homogenate of the tissue surrounding the implant. Serum levels may also be evaluated, although they may not represent the true picture at the implant site. Anderson and co-workers performed proteomic analysis experiments on monocytes cultured in vitro to determine the time course of macrophages and found that pro-inflammatory cytokines such as IL-1 β and IL 6 are secreted initially 85. As time progresses, the levels of these pro-inflammatory cytokines drop down and are replaced by anti-inflammatory cytokines such as IL 10. The concentration of chemo-attractant 1L 8 and MIP 1β also decrease around this time

85

. Anderson et al. also

showed that the adhesion of macrophages is less on hydrophilic surfaces, but such macrophages are phenotypically different and produce more cytokines. However such macrophages also show a decrease in cytokine production with time. It is generally accepted that TGF-β also plays an important role in the induction of ECM formation.86-87

26


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Figure 6: A general scheme showing the common parameters which are considered for evaluation of the foreign body response. (a) The collagen density around a subcutaneous PLGA implant after 4 weeks of implantation is mapped from the Mason Trichrome Stain images, (b) Histopathology analysis and quantification of common anti-inflammatory cells, (c) Changes in the implant morphology observed after implantation. (Top left): PLGA implant before implantation, (Top right) The implant is opaque 2 weeks post-implantation, (Bottom left)the implant is surrounded by a dense, thick fibrous capsule after 4 weeks of implantation, (Bottom right) The morphology of the implant is distorted 16 weeks after implantation. (d) The Scanning Electron micrograph images of a surface eroding salicylic acid polyester (SAP) and bulk erosion PLGA, insets show cross-sections.(e) The differential scanning thermograms of PLGA, which shows an endothermic peak of water as a result of bulk erosion; such a peak is absent in SAP. (a,c,d,e) reproduced with permission from Ref.88. Copyright 2015 American Chemical Society.

Implant related parameters Different characteristics of the implant influence the foreign body response. The material properties such as the degradation of the implant, the implant morphology and implant surface chemistry can be evaluated after animal sacrifice. It is always advisable to perform 27



Page 28 of 77

multiple analyses on the implant as soon as possible after explantation to avoid additional degradation of the implant. These parameters are best illustrated with reference to a biodegradable implant. Our group has developed a salicylic acid based cross-linked polyester (SAP) for slow release of salicylic acid, which was found to be cytocompatible with C2C12 mouse myoblast cells. From these encouraging in vitro results, we assessed the potential of salicylic acid to reduce the foreign body response in a mouse model. Figure 6c shows the photographs of the implants after sacrifice. The geometry of SAP implants was maintained throughout the period of implantation while the PLGA implant showed distortion and opacity. This difference in morphology was also visible at a length scale of microns (Figure 6d). The SAP implants showed some roughening of the surface, but the bulk was intact. On the other hand, PLGA showed the presence of cracks, and the bulk sample also showed signs of roughening. These morphological observations indicate that degradation of SAP is restricted to surface regions only and does not extend to the bulk of the implant. It is also interesting to note that the degraded areas corroborate with the size of the FBGCs, so it is our hypothesis that a high concentration of enzymes at the polymer surface leads to degradation of the local surface, a similar behaviour of localized degradation was also observed in poly (carbonate urethane) and poly (ether urethane) in vivo and has been studied extensively by Anderson and coworkers89-91. All the evidence points to the different mechanisms of erosion that the two implants follow- SAP undergoes surface erosion, while PLGA undergoes bulk erosion. Chemical investigations may be used to reinforce such observations in addition to morphological evidence. The differential scanning calorimetry thermograms also showed a difference between the two polymers (Figure 6d). However, the first heating cycle of PLGA (16 weeks post-implantation) shows a broad endothermic peak, corresponding to water 28


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around 80-120 0C. Such a peak was not observed for SAP. The presence of such nonfreezing bound water also indicates the presence of hydrophilic groups due to degradation of PLGA. Acid and hydroxyl groups are formed during degradation due to ester hydrolysis. Such observations signify the penetration of water in to the PLGA matrix and are a strong evidence for bulk erosion. The above example proves that an in vitro investigation of the implant can reveal a lot about the behaviour of the implant in vivo and presents a powerful methodology in the evaluation of the foreign body response. Generally, the degradation behaviour of implants in vivo is grossly different from those in vitro. The erosion phenomena can be effectively assessed in this manner in the absence of extensive data from post-operative implants from humans as implant retrieval is done only in the event of a failure. An insight into tissue-biomaterial interaction Biomaterials which can reduce the foreign body response are desirable. An understanding of the interaction of the non-physiological surfaces of biomaterials with the host tissue provides a rationale for the development of biomaterials. This is the first step, which leads to a cascade of further steps that ultimately decide the fate of an implant. Hence, it is of general interest to understand these preliminary events that occur after implantation of a biomaterial.

Protein Adsorption on biomaterials After implantation, the first molecules to reach the implant surface are water molecules (within a few nanoseconds), forming a water shell on the biomaterial surface. The proteins and the biomolecules in plasma/ cell culture media also have a hydration shell of water molecules. The water shell on the biomaterial interacts with the hydration shell of the biomolecules, and this interaction decides the orientation and coverage of the adsorbed proteins and whether they denature or not. This interaction is governed by the kinetic and 29



Page 30 of 77

thermodynamic processes at the interface. They may be expressed in terms of hydrophilicityhydrophobicity at a macroscopic scale in terms of water-water (cohesive) and water-surface (adhesive) forces at the microscopic level

92

. This process of protein adsorption, which is

better known as 'biofouling' 19 has been illustrated schematically in Figure 7.

Figure 7: A schematic illustration of the events that occur after implantation of a biomaterial. The water shell that is formed (on a ns timescale) affects protein interaction which begins at a µs to ms timescale and continues for much longer times. Cells reach the surface and ‘see’ this layer of adsorbed proteins. Adapted with permission from Ref. 92. Copyright 2002 Elsevier.

30


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The consequences of protein adsorption are manifold, as this step occurs even before the arrival of cells on a biomaterial substrate. When cells reach the surface of the biomaterial they 'see' the adsorbed layer of proteins, whose properties are determined by the water shell. The surface bound protein layer is influenced and determined by the properties of the biomaterial. If the fragile quaternary structure of the proteins is not retained in the adsorbed layer, a loss in their activity is observed. However, it is this step of protein adsorption on the surface of the biomaterial that renders even synthetic biomaterials recognizable to the immune system 5. One of the strategies adopted to counter the protein adsorption is to modify the implant to surface to reduce or eliminate the initial protein adsorption by retaining the bound water later making the surface non-fouling (Figure 7b, d and f). This non-fouling nature has been reported in several systems such as crown ether, PEG and its modifications, plasma deposited PEO, PEG-like glyme, diglyme and triglyme surfaces.93-94 While such surfaces are highly resistant to protein adsorption, it has been reported that this leads to complement activation, possibly due to plasma irradiation generating functional groups that activate the complement system resulting in host response 95.

Cell adhesion on biomaterial surfaces Cellular recognition becomes possible due to the interaction of the adhesion receptors with the adhesion proteins (integrins). Focal contacts are formed due to interactions of cell receptors with ligand motifs adsorbed on the surface. These focal contacts later mature into focal adhesions. Figure 8 depicts the various steps that occur during cell adhesion on a biomaterial substrate. Cells gradually spread and attach firmly to the substrate through ligand-receptor binding sites, as shown in Figure 8. The focal adhesion points exert control on cell growth and proliferation through mechanical forces that cause changes in cell shape and cytoskeletal tension

96-97

. The mechanical properties (elastic stiffness) of the substrate 31



Page 32 of 77

also influence the focal adhesion structures resulting in mechanisms that sense matrix elasticity and modulate cell/tissue response accordingly as demonstrated in vivo effectively in poly (2-hydroxyethyl methacylate) and silicone implants of variable stiffness98-101.

Figure 8: A schematic depicting the steps in the progression of anchorage-dependent mammalian cell adhesion. (a) shows the initial contact of cells with the adsorbed protein layer, (b) shows the interactions between cell surface receptors and cell adhesion ligands that are adsorbed on the biomaterial, (c) shows extensive cell spreading on the biomaterial substrate with enhanced ligandreceptor interactions. Adapted with permission from Ref.97. Copyright 2004 Academic Press.

The modification of protein adsorption is also a tool to modulate the immune response. The cells respond specifically to the adsorbed proteins and are virtually blind to the biomaterial. Hence, adsorption of proteins on the biomaterial surface plays a key role in determining the biocompatibility of an implant. Protein adsorption also plays an important role when nanoparticles come into contact with tissues. For instance, bare nanoparticles are surrounded 32


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by active molecules (mainly proteins) called protein coronas. This increases the risk the immunogenicity of nanoparticles, as discussed in detail in a review elsewhere 102. The critical role of water molecules at the interface While the role of water molecules at the interface is undisputed as far as the ‘biocompatibility’ of a material is concerned, the mechanisms behind the cause of such phenomenon are still not clearly understood. Different views on this topic exist in literature. For a historical perspective on the water molecules at the interface, the interested reader is directed to more comprehensive literature on this topic

103

. Here, we attempt to elaborate a

few such schools of thought that were developed in this aspect. Vogler argues that the structure of water at surfaces is a manifestation of surface hydrophobicity and vice versa

104

.

Typically, the distinction of the functional surfaces in contact with the biological fluid was based on the characteristic different protein adsorption that was observed for hydrophobic surfaces (surface tension < 30 dyne/cm) when compared to hydrophilic surfaces (surface tension > 30 dyne/cm)

104

. Significant attention has been focussed on understanding the

interactions of the interfacial water. Morita and co-workers have introduced the concept of ‘intermediate water’, and hypothesize that his layer plays a deterministic role in making a polymer bioinert, a term which is generally used in reference to non-degradable implants, of a permanent nature, that do show the presence of the foreign body capsule, but are still tolerated by the host. Different types of surfaces have been developed with the aim of reducing protein adsorption and improving acceptance by the body. These are described in detail in section 7. However, the intriguing role of water warrants more research in to the contribution of water molecules. According to Tanaka et al.., water in a hydrated polymer may be classified into three different types: (a) Freezing water, which may be (i) free freezing water or (ii) bound freezing water (also known as intermediate water) and (b) non-freezing 33



(also

known

as

non-freezing

bound)

water

Page 34 of 77

(figure

9

a)

105

.

Figure 9 A schematic showing the different kinds of hydrated water in a polymer as observed from DSC thermograms. (b) The molecular structure for poly(2-methoxyethyl acrylate) (c) The DSC thermogram showing the crystallization peak for freezing bound water and (d) The different vibration bands corresponding to free water, freezing bound water and non-freezing bound water. Reproduced with permission from Ref.105. Copyright 2013 Springer Nature.

These types of water layers were experimentally proven for poly(2-methoxyethyl acrylate) (PMEA) using DSC

106

, attenuated total internal reflection Fourier transform infra-red

spectroscopy (ATR-FTIR) and Nuclear Magnetic Resonance (NMR) spectroscopy. They hypothesize that the presence of the intermediate water layer determines if the polymer would be biocompatible or not; and this layer is absent in polymers which are not bioinert. In other 34


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words, the intermediate water layer may be used as an index of biocompatibility. The intermediate water layer may be identified from the crystallization peak, which appears at a temperature < 0 oC. The non-freezing bound water does not freeze even below -100 oC. It was shown that this occurs due to strong interactions between the carboxylic groups on the polymer and the water molecules; typically two carbonyl groups interact with one water molecule in a C=O-----H-O-H-----O=C manner

105

. IR spectroscopy also provides evidence

for the presence of non-freezing bound water, which shows a vibration band at 3600 cm -1, which is characteristic for the O-H stretch of water molecules that form hydrogen bonds with –C=O groups. Thus, the intermediate water layer is emerging as the main player in directing protein adsorption and consequently in determining the foreign body response to an implant. Here we would like to warn the readers that although advances in instrumentation have greatly improved our understanding of water-surface interactions, some researchers still regard the physical chemistry behind how this occurs as debatable. This is partly fuelled by the fact that the literature on ‘protein adsorption’ is highly controversial, and hydrophilic or hydrophilic surfaces do not represent the complex physical phenomenon of water at surfaces. Nevertheless, as will be described later, achieving bio-inertness by tuning protein adsorption of the surface is a major strategy; the roots of which originate in the behaviour of water molecules at the biomaterial interface. While the actual thickness of the adsorbed layers of water and proteins on the biomaterial surface may be of the orders of nm, and may seem insignificant to cause the foreign body response, their influence should be considered in a different perspective. The gravity of the situation is better understood when we take into account that altered protein conformation can cause exposure of protein domains which can trigger cell adhesion and ECM secretion. Material properties affecting the foreign body response 35



Page 36 of 77

The foreign body response is influenced by many factors, which may be broadly classified based on (a) the design of the implant and (b) the host. These properties are elucidated in Figure 10. Sanders et al. demonstrated that small fibres with a thickness of < 6 µm generated a reduced foreign body response, with less macrophages in a rat model 66. Here the authors had proposed that if the area of contact between the cell and the implant is less than a certain threshold value, an inflammation response is not produced. However, such a threshold concept has not been experimentally demonstrated

66

. In fact, recently, Anderson and co-

workers proved that alginate spheres of 1.5 mm diameter produced a reduced foreign body response and fibrosis (Figure 10 b, c). The performance of similar spheres of a smaller size (0.5 mm diameter) was much worse with an enhanced FBR. They also tested this concept for a wide spectrum of materials such as metals, plastics and ceramics, and found this size of spheres (≈ 1.5 mm) to be optimal for curbing the FBR 62. This is rather counter-intuitive, as larger spheres of 1.5 mm diameter showed a lower FBR than smaller spheres. They also emphasized the importance of a spherical geometry of the implant, as both size and geometry of the implant influence the host response. Prior to this, Matlaga et al. demonstrated that a spherical implant generated the least FBR amongst a triangular, spherical and a pentagon shaped implant

107

. Incidentally, Li et al. also showed that the fibrous deposition around a

rectangular implant is non-uniform, and the corner rings often differ greatly from the rest of the implant, and show substantially low encapsulation

108

. Here it is important to interpret

corner rings as smooth interfaces where one implant plane meets the other, and not as extremities (also see figure 4). Sharp tips and edges in the design of an implant lead to elevated local strains (as expected intuitively) and should be avoided in implant design 109 as sharp discontinuities can cause severe tissue damage. The orientation of the implant is also a critical factor, Sanders et al. suggest that the FBR is a local response, and is not uniform throughout the implant, as shown in figure 10c

110

. In 36


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particular, the height of the implant is an important factor, as adjacent collagen layers are displaced, and the low pressure areas created due to this are replaced with fibrous tissue 4. Hence, thinner implants are generally advocated. Another important factor that strongly influences the FBR is the elastic modulus of the implant. A mismatch in mechanical properties has undesirable consequences. In the case of soft tissues, a material with elastic modulus higher than the tissue will not only lead to a higher foreign body response

111

, but

will also cause pain 44. Finite element simulations on a microelectrode implanted in the brain by Kipke and co-workers showed that the strain fields caused due to the ‘micro-motion’ of an implanted microelectrode in the brain are much larger when the modulus of the electrode was ~ 200 GPa than that caused by a softer (6 MPa) electrode 109. An important point to note here is that the displacements were small (1-2 µm). It is not hard to imagine that large displacements of such implants (which could be caused due to various reasons such as the host walking etc.) can be devastating. For instance, silicon is ~106 times stiffer than the brain and this leads to pressure on the surrounding tissue, and can even cause tissue tearing

109

.

Synthetic, elastomers, which have unique mechanical properties and can undergo large, reversible deformations are considered to be the panacea to these drawbacks of currently existing polymers. The implant-tissue mechanics involves both, the relative motion of the implant with the tissue and modulus mismatch. Both these aspects ought to be minimized in order to reduce wear debris due to friction at the interface112-113. Guck and co-workers have also shown that adapting the stiffness of implants to that of the physiological stiffness of the organ can also alleviate the foreign body response

111

. In fact, they suggest that a surface

coating of the required mechanical stiffness to might suffice, but this is yet to be confirmed by focused studies 111.

37



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Figure 10: (a) An overview of the different factors which influence the foreign body response. (b) Laparoscopic images of alginate spheres after 28 days of implantation in the intraperitoneal cavity of non-human monkeys. While spheres of 1.5 mm diameter were free, those with a diameter of 0.5 mm were completely embedded in the fat tissue, as is also visible in the tissue biopsy (image3, top row). The retrieved hydrogels were transparent, without obvious encapsulation. Reproduced with permission from Ref.62. Copyright 2015 Springer Nature. (c) 38


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The distribution of collagen after 18 months of implantation of a disc shaped hydroxyapatite implant in rats. Reproduced with permission from Ref 108. Copyright 1999 John Wiley and Sons.

Material strategies to overcome the foreign body response To develop effective strategies to reduce the foreign body response, the study of the foreign body response to gain an in-depth understanding is advocated 19. On similar lines, Ratner has compared normal wound healing with wound healing in the presence of a biomaterial 22. He has attempted to find an explanation to the question: Why do the so-called ‘biocompatible’ materials shut off normal healing? He observes that while the human body has an excellent capacity for healing, wounds with implants do not heal like normal wounds. In this seminal shift in the paradigm of biocompatibility, he has critically assessed the differences in protein adsorption on the biomaterial, which is one of the initial stages in the foreign body response. Reduction of non-specific protein adsorption Wound healing was assessed by Ratner and co-workers in an experiment where mammals were implanted with implants made of materials commonly used in medicine such as teflon, polyurethane, silicone rubber, polyethylene, poly(methyl methacrylate) (PMMA), PHEMA, Dacron, gold, titanium and alumina 22. All these implants healed in a similar manner after 1 month.

39



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Figure 11: A schematic showing the hypothesis that non-specific protein adsorption causes a foreign body response. The conformation and orientation of proteins is also distorted during non-specific protein adsorption. The foreign body response may be overcome by preventing protein adsorption or by modifying the biomaterial to release bioactive agents such as anti-inflammatory/ angiogenic drugs/ Vascular Endothelial Growth Factor (VEGF). 40


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However, these materials showed differences in protein adsorption and hence, different cell growth behaviour in vitro. This observation was rather confounding. The only common aspect in all materials in vivo was the adsorption of a complex protein mixture, (> 200 proteins) in different orientations and conformations. Ratner has hypothesized that this protein layer is the key cause of the foreign body response. In nature, such non-specific protein adsorption is never observed. Hence, he proposed the reduction/ elimination of this non-specific protein adsorption on the biomaterial as one of the main strategy to reduce the foreign body response. This approach has been schematically depicted in figure 11. Different methods have been adopted to reduce non-specific protein adsorption on biomaterial surfaces. The most common one is the use of surface coatings. Biocompatible coatings or surface modifications are used to mask the underlying implant surfaces and to improve surface properties of implants. Generally, a coating that creates a hydrophilic surface is used 36. Hydrophilic poly (ethylene glycol) and zwitterions are two major classes of nonfouling materials that are based on total reduction of protein adsorption

114

. Each of these

classes is described below. Poly(ethylene glycol) Surface grafting using poly(ethylene glycol) (PEG) based materials (PEGylation of surfaces) has been used since 1980s to inhibit protein adsorption in biological media

115

. PHEMA is

also used to create hydrophilic surfaces. The structures of PEG and PHEMA are shown in Figure 12 (a, b). Such hydrophilic surfaces can reduce non-specific protein adsorption, but the reduction is not sufficient enough to prevent adhesion of cells/bacteria

116

. Many

researchers observed that when scaffolds coated with PEG or PHEMA were implanted, collagen encapsulation was observed around them22, 117-118. This is due to enzymatic oxidation

41



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of the hydroxyl groups in these implants to aldehydes and acids. These surfaces allow protein adsorption and cell attachment

118

. It is also likely that the manifestation of the physical

chemistry of the polymer at the interface plays an influential role in determining the fate, apart from the water binding differences. . Although PEG has been considered as one of the best non-fouling materials, it is decomposes in the presence of oxygen and transition metal ions and in most biological media

118-119

. Hence, the main drawback of PEG is its

susceptibility to oxidation damage, which has severely restricted its use in long term applications in biological system

120

. In spite of this, there is currently no reported evidence

of oxidation induced damage of PEG-based materials in vivo.

42


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Figure 12: Chemical structures of some commonly used hydrophilic coating materials (a) Poly(ethylene glycol), (b) poly(hydroxy ethyl methacrylate), (c) zwitterionic poly(sulfobetaine), (d) zwitterionic poly(carboxy betaine), (e) cationic poly(carboxybetaine) ester, (f) mixed charge polymers.

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Zwitterionic materials The use of zwitterionic materials is one of the earliest strategies to reduce protein adsorption on implant surfaces. Some polymer systems like phosphorylcholine (belonging to the class phospholipid polymers) were extensively studied as non-fouling surfaces by Chapman’s group and Ishihara’s group over the last two decades.121-127 As a result, many clinical devices currently make use of zwitterion based biomaterials as stents, lenses and non-adherent membranes to reduce foreign body response. Recently, Jiang and co-workers have shown that zwitterionic materials such as poly(sulfobetaine) and poly(carboxybetaine) and poly(carboxybetaine ester) (Figure 12 c, d, e) are more efficient at reducing non-specific protein adsorption

116

. These materials have cationic and anionic groups on the same

monomer residue. Mixed charge materials (Figure 12 f), though not zwitterions, are equivalent to zwitterions if the charges are uniformly distributed at a molecular level. In these materials, two different moieties (unlike in zwitterions) having opposite charges selfassemble (depending on polymerization conditions) on the surface to homogenously spread positive and negative charges on the surface at a molecular scale, while remaining neutral as a whole resulting in a zwitterionic behaviour and resistance to protein adsorption128-129. They bind water molecules very strongly due to electrostatically induced hydration as opposed to hydrogen bonded water molecules on hydrophilic surfaces, which is a precursor for nonfouling properties

116, 126, 130

. Zwitterions and mixed charge materials form a class of very

robust non-fouling surfaces, and are known as ultra-low fouling materials, as they allow

The foreign body response demystified - ACS Publications - American

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