Advanced strategies to thwart foreign body response to implantable devices

Abstract Mitigating the foreign body response (FBR) to implantable medical devices (IMDs) is critical for successful long‐term clinical deployment. The FBR is an inevitable immunological reaction to IMDs, resulting in inflammation and subsequent fibrotic encapsulation. Excessive fibrosis may impair IMDs function, eventually necessitating retrieval or replacement for continued therapy. Therefore, understanding the implant design parameters and their degree of influence on FBR is pivotal to effective and long lasting IMDs. This review gives an overview of FBR as well as anti‐FBR strategies. Furthermore, we highlight recent advances in biomimetic approaches to resist FBR, focusing on their characteristics and potential biomedical applications.


| INTRODUCTION
In the past years, the rising demand for implantable medical devices (IMDs) has been fostered by advances in manufacturing technologies and biomaterial science. This trend is mainly driven by the increasing geriatric population, more prone to chronic conditions, and the increased demand for organ transplantation. 1,2 Orthopedic prosthesis, 3,4 breast implants, 5,6 neural stimulators, 7,8 cardiovascular devices [9][10][11] and stents, 12 ocular and cochlear implants, [13][14][15] tissue engineering scaffolds, 16,17 and biosensors [18][19][20] are only some widely used examples of clinically approved IMDs ( Figure 1). Furthermore, IMDs can be utilized as self-regulated drug delivery and cell encapsulating systems that allow controlled sustained therapeutic delivery and cell engraftment. [21][22][23][24][25][26] The global IMDs market is expected to grow at a compound annual growth rate of approximately 6.9% from 2020 to 2027 and reach a market value of nearly US$ 155 billion by 2026. 1 However, despite many advantages that these devices potentially offer to medicine, and increasing demand in the market, most implants fail to meet the implantable devices biocompatibility criteria due to the foreign body response (FBR).
FBR induces the formation of a capsule-like dense fibrous tissue that isolates the device. FBR consists of a complex series of immune Simone Capuani and Gulsah Malgir contributed equally to this study. defense mechanisms against foreign material. Upon implantation, the first event is the adsorption of blood plasma proteins, predominantly albumin, and fibrinogen, on the material. 27 Depending on the material surface properties, proteins undergo conformational changes, resulting in the opening of protein recognition patterns that attract the innate immune system cells; neutrophils, monocytes, and macrophages. Neutrophils are the first-line responders, as their recruitment to the implantation site happens 2 days postimplantation. 28 This first phase of acute inflammation usually lasts for a week and resolves shortly thereafter. In presence of a foreign body, namely, an IMD, the acute inflammation persists and leads to chronic inflammation. 29 This phase is marked by monocyte infiltration and macrophage activation, and lasts for approximately 3 weeks. 30 Macrophages are critical components in capsule formation. 31 When activated during the inflammation period, these cells are classified as M1 and M2 phenotype macrophages. M1 macrophages secrete proinflammatory cytokines (including interleukin-1 (IL-1)) and chemokines, 32,33 while M2 macrophages upregulate the anti-inflammatory pathway and tissue remodeling. In the initial stage of immune reaction to tissue injury, the M1 phenotype population is predominant. As the chronic inflammation resolves, macrophage polarization shifts into M2 phenotype and the natural wound healing process. However, in the presence of foreign bodies, such as IMDs, this process is delayed, and proinflammatory macrophage proliferation continues. Macrophages attempt to eliminate the implant via phagocytosis by secreting reactive oxygen species (ROS) and matrix metalloproteinases. 34 However, in the case of slowly degradable or nondegradable implants, the continuous presence of the device and the inability of macrophages to eliminate the implant promotes the fusion of macrophages into foreign body giant cells (FBGCs). 35 Antigen presenting proinflammatory macrophages also induce adaptive immune system cells, B lymphocytes, and T lymphocytes to secrete proinflammatory cytokines such as interleukins to induce fibroblast activation. Following the activation of fibroblasts via secreted chemokines and cytokines, weak focal adhesion of the cells on the material surface triggers differentiation of fibroblasts into myofibroblasts via tensile forces. This process is characterized by α-smooth muscle actin (α-SMA) expression in intracytoplasmic stress fibers, which implies high contractile activity, 36 and by secretion of collagen by myofibroblasts. Ultimately, they create a dense, avascular collagen fiber network called fibrotic tissue that encapsulates the device 37,38 (Figure 2). This tissue blocks the implant-host tissue interaction, which may impair the implant function and subsequently reduce the implant lifetime. 39 Lack of vascularization, for example, reduces blood supply, limiting oxygen and analyte diffusion, and obstructs drug delivery. 40 Thus, understanding the tissue response to implantable materials in depth could be leveraged to achieve specific functions (e.g., engraftment) or avoid undesired effects (overt fibrosis) to meet the clinical needs.
FBR and subsequent fibrotic encapsulation contribute to the failure 41 of many devices, including biosensor, 42 coronary stents, 43 breast implants, 44 encapsulated tissues/cells drug delivery systems, 34 and ocular implants, 45 endangering the health of the patients (Table 1). For example, the failure rate of breast implants alone due to the FBR is 30%, 58 and the failure rate of all other implantable devices is conservatively estimated to be 10%. 66 Notably, solving this critical clinical challenge could eliminate nearly $10 billion in cost to the healthcare system annually. Therefore, there is a clear need for IMDs design principles that focus on device parameters, such as size, shape, surface topography, mechanical stiffness, and wettability 67 (Figure 3), critical for the FBR.
This review first focuses on the characteristics of implantable devices and how these affect local tissue remodeling in response to immune modulation. Second, repercussions of FBR on implant performances is examined, with emphasis on drug delivery and cell encapsulation devices. Finally, this work discusses recent advances in biomimetic strategies, adopted to mitigate the FBR. More specifically, this review covers solid, nondegradable, implantable macrodevices for long-term deployment. Cardiovascular devices bear additional complexities related to their blood-contacting nature, are thoroughly discussed by other groups, and will not be covered herein. [68][69][70][71][72][73][74][75][76] Similarly, F I G U R E 1 Examples of implantable devices F I G U R E 2 Stages of foreign body reaction and fibrotic tissue formation micro-or nanoparticles and bone and joint replacements are out of the scope of this review as they have been extensively reviewed elsewhere. [77][78][79][80][81][82] 2 | IMPLANT PROPERTIES AFFECT THE DEGREE OF HOST RESPONSE Biomaterial surface properties determine the protein interaction level and biological response of immune cells, particularly the fate of macrophage polarization. Understanding biotic-abiotic interaction is of profound importance in designing implantable biomaterials with immunomodulatory properties. Tailoring the surface characteristics such as roughness, hydrophilicity, charge, size, shape, and mechanical stiffness has a potential impact in changing the direction of FBR towards the tissue repair process. In this section, we will cover the implant material parameters that induce different immune-mediated FBR.

| Surface topography
Surface topography is an essential aspect of medical implants that plays a pivotal role in material-host tissue integration. 83 It regulates the density of adsorbed protein on the surface and its interaction with the surface, which induces inflammatory cytokine secretion and macrophage fusion. Altering the surface topography at micro/nano levels can tune the degree of biofouling, focal cell adhesion, proliferation, and ultimately regulate fibrotic capsule formation. 84,85 For instance, it has been confirmed that osteoblastic cell adhesion, growth, and proliferation are correlated to the surface roughness of Ti. 86 89 have an influence on macrophage behavior in vitro. 90 This material has been studied in flat, expanded, and electrospun arrangements. Electrospun PTFE with a surface roughness of 1.08 μm reduced the macrophage cell attachment and FBGCs formation compared to flat (roughness 0.17 μm) and expanded (roughness 0.37 μm) PTFE.
Silicone is a biocompatible polymer widely implemented as an implant material for many applications, including tissue engineering  103 The study found that micropillar sizes ranging from 5-10 μm in diameter can enhance macrophage adhesion and a combination of micropillar size and density can modulate their phenotype.
Collectively, these findings suggest that surface roughness or modified surface topography obtained by adjusting height and depth of surface features can influence the FBR formation through modulation of cell adhesion patterns. Roughness smaller than 1 μm appear to have little to no effect on FBR mitigation, while surface features in the range of 1-4 μm show a potential to ameliorate implant integration. Spatially confined surfaces with a diameter smaller than the size of an immune cell can limit the spreading and activation of proinflammatory cells on the material surface. 85 Although tuning surface topography could reduce FBR, it is of paramount importance that the overall mechanical and functional properties of the device remain unaltered.

| Surface charge
The surface charge can influence the protein adsorption and the interactions between immune cells and the material at different stages of FBR. In particular, adsorption is dictated by the overall charge, present on the surface of the material rather than by atomic-scale electrostatic interaction. [104][105][106] Moreover, the balance between the surface isoelectric point and the pH of the surrounding fluid defines the charge at the material and fluid interface. Thus, a pH below or above the surface isoelectric point generates a positively or negatively charged surface, respectively. 107  and proliferation in vitro. 111 The combination of surface treatments significantly reduced cell adherence after 24 h of incubation compared to the unmodified nanostructured surface. In addition, increased macrophage polarization towards the wound-healing M2 phenotype was observed on the ion-modified surfaces.
One of the molecules most utilized to obtain biocompatible biomaterials is polyethylene glycol (PEG) which provides a shielding effect. PEG is negatively charged, and it is known to protect biomaterials against nonspecific protein adsorption. 112  Furthermore, zwitterions induce anti-inflammatory M2 macrophage expression. 124 In a recent study, neural microelectrodes were coated with zwitterionic layer consisting of poly(sulfobetaine methacrylate). 125 The treatment prevented protein adsorption, fibroblast, and microglia attachment on the electrodes and remained stable in vitro for 4 weeks. Furthermore, in a short-term implantation test, the coated microelectrodes significantly reduced microglial surface coverage compared to uncoated controls. In addition, zwitterionicmimicking materials can be developed by assembling oppositely charged macromolecules, such as the balanced charged alginate/poly ethylenimine hydrogel. 109 After 3 months subcutaneously implanted in mice, the hydrogel showed significant antifouling properties, diminished the FBR, and subsequent capsule formation.
It can be challenging to isolate and study the surface charge as a single factor avoiding other properties, such as surface wettability. 126 Nevertheless, negatively charged surfaces appear to elicit a milder response, followed by thinner capsule deposition and limited neovascularization when compared to positively charged counterparts. Furthermore, surfaces exhibiting a neutral charge prevent protein adsorption and significantly reduce the FBR.

| Surface wettability
Protein adsorption, the first stage of FBR, is generally energetically favorable towards hydrophobic surfaces. 127 On the contrary, removing the water molecules from the hydrophilic surfaces bears a higher energy barrier 114 demonstrating protein repellent features. 128

| Implant stiffness
One of the parameters of the implant that influences the FBR is stiffness. This mechanical property is a characteristic of the material and it can be measured with the elastic modulus or Young's modulus. It appears that implant materials with analogous Young's modulus to the one of the surrounding tissue can help avoid severe immune response. 152 Mismatch of Young's modulus at the biotic-abiotic interface is one of the fundamental driving forces in scar tissue formation. 153 Shear stress due to the stiffness of the material and micromotion in the brain, for example, damage the surrounding tissue resulting in enhanced proinflammatory cell activation, including reactive astrocytes. 154 Currently deployed brain implants, predominantly silicon implants, are much stiffer 155 than the brain tissue ($1-30 kPa) and can generate acute FBR that may impact their funciton. Thus, efforts have been devoted to developing softer implants to reduce FBR.
Ecoflex, a silicone-based material with a low stiffness (20 kPa) was assessed as a mechanically matched brain implant (MMBI) in rats. 156  Although there is a mismatch in stiffness between the hard hydrogel and the surrounding tissue, these results are in contrast with the previously discussed findings. However, the authors hypothesized that the difference in acute inflammatory response can be caused by the higher concentration of pectin in the hard hydrogel coupled with its slower degradation kinetics. In another study, PEG-phosphorylcholine hydrogels with a stiffness ranging from 3 to 165 kPa were implanted subcutaneously in mice. In this case, there was a direct relationship between the stiffness of the hydrogel and macrophages adhesion and fibrotic capsule thickness. Modulation of hydrogels stiffness has also been explored as an approach to develop mechanically matching electronic nerve interfaces for tissue regeneration. Schwann cell proliferation was compared among magnetically templated glycidyl methacrylate hyaluronic acid hydrogels with different stiffnesses. 159 The hydrogels with mechanical properties similar to fresh nerve tissue promoted cell migration and infiltration within the scaffolds.
Myofibroblast activation, and subsequent collagen deposition, can be suppressed by reducing the mechanical stress generated by implants stiffer than the surrounding tissue. 160 Coating stiff implants with a layer of soft material that matches the elastic modulus of the host tissue can prompt decreased inflammation and fibrosis in comparison to uncoated implants (preprint). 161 A stiff silicone rubber core with Young's modulus $600 kPa was coated with 0.6 kPa polyacrylamide (PAA), 6 kPa PDMS, or 60 kPa PAA. Coated and uncoated implants were implanted for 3 months in subcutaneous tissue and nerve conduits in rats. In both sites, the coated implants showed reduced α-SMA and CD68 expression in the surrounding tissue than uncoated ones. In addition, a significant decrease in fibrotic capsule thickness was observed in soft-coated implants. In a similar study, the FBR to soft silicone coating ($2 kPa) applied on stiff silicone implants (2 MPa) was evaluated post-3 months subcutaneous implantation in mice. 162 The coated implants elicited a reduced myofibroblast activation and subsequent collagen deposition than the uncoated counterparts. Moreover, the soft-coated implants showed a reduced TGF-β1 activation, a profibrotic growth factor that induces myofbroblast contractile activity and leads to the formation of fibrotic tissue. 163,164 Cell behavior is driven by mechanical stimuli, among other cues.
Therefore, it appears that materials with similar stiffness to the sur-  166 In addition, a significant neutrophil population increase was observed in the peritoneal space following microcapsules implantation. 167 Conversely, PLA and PLA/PCL blend implants elicited similar FBR following subcutaneous and intraperitoneal implantation in rats for 2, 8, and 24 weeks. 168 In another study, cylindrical PEG hydrogels were implanted in the subcutaneous space, abdominal cavity, and adipose tissue. 115 The mildest inflammatory response was induced by subcutaneous implants, followed by implants in the abdominal cavity. These abdominal implants showed an increase in macrophage infiltration and few neutrophils. However, the most robust response was observed in the adipose tissue, which is known to be a more hostile microenvironment. 169,170 While, the intraperitoneal space is an attractive site for cell transplantation due to its abundance of blood vessels, and consequently oxygen, excessive fibrosis can hamper its facile diffusion. Nonetheless, fast tissue integration on implants, a specific trait of this implantation site, might be the desired effect in specific applications. 171 In a recent study, fibrosis-generating biomaterials were implanted across different species and sites in an effort to explore the variations in FBR between the subcutaneous space and the immune-privileged intrauterine environment. 172 Minimal intrauterine fibrosis was observed in NHP, whereas a strong fibrotic FBR was provoked by the same biomaterials implanted subcutaneously in mice. In this setting, subcutaneous sham surgeries led to negligible fibrosis, excluding tissue disruption as the major factor for FBR discrepancy. Different from subcutaneous implantation, intrauterine placement causes negligible tissue disruption, which could justify FBR discrepancy. Thus, the authors speculate that uterine immune privilege could play a role in minimizing fibrosis. In another study, FBR to collagen discs implanted in the left ventricular epicardium and the subcutaneous space was investigated. 173 Notably, discs in the epicardium exhibited a stronger inflammatory response with a higher influx of macrophages, PMNs, and angiogenesis. Moreover, distinct subcutaneous locations contributed to differences in fibrotic capsule thickness. 174 The fibrotic capsule was five times thicker in devices implanted in the middorsal space compared to the scapular site. This discrepancy could be attributed to the different shear forces on the implant that occur in the specific sites; hinting that the microenvironment is not the sole key determinant in FBR variation. 144 The implantation site is also highly critical for implantable sensors as the FBR can potentially impair the sensor function. For example, intravascular sensors can provide accurate measurements, but activation of the coagulation cascade is a major concern. 175 Finally, recent studies suggest that the key to understanding the FBR variation at different implantation sites may be to determine the tissue-resident macrophages population. The behavior of these macrophages is influenced by the niche in which they reside. 176  Orchestrating the host tissue response, macrophages are pivotal at controlling the device function. A comparative study implanting CGM sensors in wild-type and macrophage-depleted mouse models proved that following 4 weeks of implantation, the former model caused accumulation of macrophages that limits the sensor functionality. In contrast, the latter enhanced the device performance, indicating that sensor impairment could be macrophage-associated. 192 In another study, the same group reported that, after peritoneal injections of mouse macrophages in the proximity of CGM in mouse model, glucose levels measured by the sensor were lower than blood glucose levels. 193 CCL2 and CCR2 are leukocyte chemotactic factors that contribute to monocyte/macrophage activation and eventually the formation of FBGCs. Utilizing CCL2 and CCR2 knockout mouse models, the same group reported that the monocyte/macrophage accumulation was significantly reduced compared to the wild-type mice. 194 Furthermore, the relative low difference between sensor glucose level readout and blood glucose levels indicates that sensor accuracy was improved. This might also be attributed to the indirect inhibition of TGF-β signaling receptor that contributes to the reduced analyte diffusion in sensors. 195 Alternatively, zwitterionic polymer modification on commercial CGMs abrogated the cross-talk between inflammatory cells and sensor electrode surface, resulting in a reduced capsule formation compared to the uncoated control implant. 123 The treatment reduced the immune response towards the sensor in mice and NHP. In addition, no recalibration was needed, and the sensors accurately measured glucose levels throughout the study.
To summarize, drug and analyte diffusion within the surrounding of the device can be significantly impaired in the event of severe and uncontrolled FBR to the device or to the drug formulation, especially in the long-term. While it is intuitive that extremely collagen dense and poorly vascularized fibrotic tissue can be a physical obstacle to molecules diffusion and biodistribution, it appears that a milder FBR can still grant satisfactory device performances. However, more research is needed to study how different characteristics of the fibrotic capsule affect molecules diffusion.

| The effect of fibrosis on cell encapsulation devices
Cell transplantation is a promising approach that entails administering living cells to patients as replacement therapy to treat various disorders.
Transplanted cells can deliver therapeutic molecules in a sustained fashion or in response to stimuli. Typically, these cells can originate either from a donor or can be engineered or obtained from animal sources.
Consequently, in various cases, cell transplantation requires some  22,[196][197][198][199] The first crucial step in designing a cell encapsulation system is selecting an optimal material. 200 Alginate is the prevalent choice for cell microencapsulation platforms, however, adjustments are required to achieve solid engraftment and prevent fibrotic isolation. 63 can form within the device and impair cell viability and function. Therefore, tuning FBR is fundamental to balance fibrosis reduction around the implant while still obtaining angiogenesis in its proximity. 214  These effects presumably contributed to the FBR-induced hypoxia that led the insulin-producing cells macroencapsulation device developed by ViaCyte to failure in a phase I clinical trial. 222 A different scenario needs to be considered for cell encapsulation technologies in the context of cancer immunotherapy. In this application, the devices are usually implanted for less than a month. Therefore, the long-term viability of the cells is not a concern. Contrarily to other cell encapsulation systems, these platforms aim to boost a strong inflammatory reaction resulting in a milieu conducive for antitumor immune response. 223 Based on the specific application or approach, the FBR can be exploited in different manners. Generally, macroencapsulation devices require a certain degree of inflammatory response to generate neovasculature that can deliver oxygen and nutrient to the graft. Vice versa, microencapsulation platforms aim at evading the immune response entirely to prevent the formation of a fibrotic capsule that will limit oxygen and nutrient diffusion to the cells.

| NOVEL BIOMIMETIC STRATEGIES TO MODULATE FBR
Conventional strategies for mitigating FBR and scar tissue formation employ immunosuppressive agents such as DEX. 224 However, occurrences of detrimental side effects of those agents are inevitable.
Instead, nature-inspired biomimetic surface modifications are attractive, widely investigated options to stealth the implants from the immune system and promote tissue-device integration. In this section, we will cover the recent biomimetic approaches for implantable biomaterials (Table 2).

| Zwitterionic molecule coating
Zwitterionic polymers that alleviate FBR have become attractive candidates as a coating strategy for implantable devices. 241 For instance, zwitterionic poly(carboxybetaine methacrylate) hydrogels can be effective for more than 3 months to avoid macrophage recognition and fibrotic capsule formation in mice. 124 Similarly, inspired by the naturally occurring immunological tolerance mediated by phosphoserine (PS), zwitterionic PS hydrogel discs demonstrate antiadhesive properties when cultured in fibrinogen-rich culture in vitro. 225 Triazole-zwitterionic (TR-ZW) hydrogels used for islet encapsulation and transplantation can also reduce FBR. 242 In addition, they can  Blood-contacting devices such as cardiovascular stents have limited long-term clinical success due to the in-stent restenosis, a series of events including thrombosis, platelet aggregation on the metal stent, and reduced re-endothelialization. 248 Heparin, the most clinically used anticoagulant molecule, has been known to prevent the early stage of thrombosis. 249  showed a significant decrease in the fibroblast thickness and elimination of the immune cell filtration after 28 days of subcutaneous implantation in mice. 233 The same group found that a low degree of methylated pectin incorporated in alginates can effectively prevented TLR-2 activation through electrostatic interaction and consequently suppressed immune activation. The material, used to encapsulate rat islets, prolonged their survival in a xenogeneic graft in mice compared to alginate and high degree pectin capsules. 256 However, it is still unclear how the degradation of the hydrogel compound will affect the long-term cell survival in the cellular envelope and suppression of overt fibrosis.
Due to its high biocompatibility, 257  drug delivery matrixes and tissue engineering scaffolds. 258 For example, in a recent study, a group designed a pancreatic islet encapsulating microcapsule with a shell made of alginate or agarose gel, and a core that was incorporated with a silk scaffold encapsulating the islet cell. This creates a more realistic ECM-like structure that is crucial for long-term islet survival. Furthermore, creating an additional interior layer with silk can prevent immune cell filtration and induced an antiinflammatory pathway. 259 The bilayered RGD peptide are silk tissueengineered vascular grafts that mimic the blood vessel structure. 258 An Poly-L-lactic acid (PLLA), with a low biodegradation rate and mild immune response, has been clinically used as an implantable material.
However, increasing the immunomodulatory effect is crucial to increase the long-term fate of the PLLA-based devices. Immobilizing the macrophage CSF-1, a hematopoietic growth factor responsible for tissue repair on the PLLA scaffolds, can enhance biocompatibility with reduced IL-1β, TNF proinflammatory cytokine levels, and increased CD68 + and CD206 + levels up to 28-day postimplantation in IL-1β reporter C57BL/6 mice. 235 Taken together, these findings confirm the short-term effectiveness of bioactive molecule coating in directing the immune cascade pathway towards the tissue repair phase. However, long-term fate of the coating remains challenging and requires extensive design testing both in vitro and in vivo.

| Other biomimetic approaches
Patterning the implant surface can influence protein adsorption and cell adhesion. Surfaces that mimic the natural tissue texture can significantly reduce the curvature at the biomaterial-tissue interface, leading to reduced fibrosis and cellular morphology. 240  Notably, these findings suggest that mimicking ECM in designing biocompatible implantable devices is a valid strategy to overcome intense FBR. However, the long-term stability of coating materials is still unclear, and requires extensive research.

| CONCLUSION AND FUTURE CONSIDERATIONS
Implantable devices have been clinically employed for decades. However, there is no gold standard to prevent or modulate the FBR.
Therefore, understanding the mechanisms of fibrotic tissue response to implantable devices is fundamental. Implant parameters, including surface wettability, topography, shape, and size, determine the degree of protein adsorption, and the proinflammatory response which may ultimately result in scar tissue formation. Considering the potential outcomes from a clinical perspective, designing innovative implantable materials to control the protein adsorption process and avoid the immune response is crucial to elucidate the implant's performance for a prolonged period. However, the implant properties that affect FBR are tightly interconnected. Multiple studies reveal that an increase in surface roughness can generate air pockets within the grooves of the surface and lead to higher hydrophobicity. Conversely, the liquid can penetrate the grooves at lower roughness, producing more hydrophilic surfaces. [263][264][265] Furthermore, the wettability of a material is highly dependent on its surface chemistry. The charged or polar functional groups exposed on the surface, either naturally, or due to a superficial treatment, determine the overall charge that interacts with water molecules. 266,267 In addition, functionalization or optimization in formulation aimed at improving the biocompatibility of a material can significantly alter its mechanical properties, thus affecting the overall stiffness and durability of the implant. 140,268 Therefore, different biomaterial features should be rigorously characterized to achieve the desired FBR mitigation and preserve device functionality.
The International Standard ISO 10993-1 principles for the biological evaluation of medical devices provide essential guidelines for in vitro/in vivo testing. 269 In vitro biocompatibility studies are mainly performed in 2D cell culture, failing to mimic the complex 3D physiological environment. Macrophages in 2D cell culture models show different phenotypes and responses to stimuli compared to in vivo settings. 270 In addition, short-term biocompatibility assessment (less than a month of evaluation) may lead to biased outcomes. 271 Therefore, biomaterial compatibility test duration should be carefully selected and in vitro findings should be validated in vivo. 272 To this extent, selecting the most suitable animal model to mimic the FBR in humans is paramount, as different species or strains can produce substantially distinct FBR. Eventually, scaling the platform to clinical translation requires further consideration. Variation in immune response among individuals, which can be related to underlying conditions or aging, needs to be accounted for. Creating a dynamic immune cell model in a lab on a chip platform can be a solution for a personalized evaluation of FBR which will be a promising strategy to reduce laboratory animal use. 273 Finally, recent progress in implementing biomimetic strategies to control FBR holds promises towards curtailing the immune response to implantable devices. For instance, CorNeat Vision's biomimetic nonbiodegradable implant that mimics ECM topography is currently under clinical trial (NCT04485858). However, much progress is required, particularly in implementing high throughput screening platforms in the early stage of device development to pave the way for clinical translation. Gulsah Malgir declares no conflict of interest.

DATA AVAILABILITY STATEMENT
Data sharing not applicable to this article as no datasets were generated or analyzed during the current study.