Biocompatibility

Biocompatibility is the ability of a device material to perform with an appropriate host response in a specific situation.^[1] This concept is fundamental in biomedical engineering and medicine, ensuring that implants, prosthetics, tissue-engineered constructs, and other devices integrate safely with living tissues, such as bone, blood, or soft tissues, to support healing or functionality without toxicity, inflammation, or rejection.^[2] The evaluation of biocompatibility is guided by international standards, particularly the 2025 edition of ISO 10993-1, which outlines a risk-based framework for biological testing within a device's overall risk management process (per ISO 14971). This standard categorizes tests based on the nature and duration of body contact—such as limited (≤24 hours), prolonged (24 hours to 30 days), or long-term/permanent (>30 days)—and includes updates like new endpoints for neurotoxicity.^[3][4] Regulatory bodies like the FDA require biocompatibility data for premarket approvals, emphasizing chemical characterization of materials alongside biological testing to reduce animal use and streamline submissions.^[5] For detailed historical development, testing methods, influencing factors, applications, and future directions, see the relevant sections below.

Fundamentals

Definition and Scope

Biocompatibility refers to the ability of a material or device to interact with biological systems in a manner that supports its intended function while eliciting an appropriate host response. According to the International Union of Pure and Applied Chemistry (IUPAC), one primary definition is the "ability of a material to perform with an appropriate host response in a specific application," emphasizing the therapeutic context in biomedical uses.^[6] A more general IUPAC definition describes biocompatibility as the "ability to be in contact with a living system without producing an adverse effect," highlighting the absence of detrimental interactions.^[6] Unlike inherent material characteristics such as density or tensile strength, biocompatibility is not an intrinsic property of a substance but is highly contextual, depending on factors like the specific application, implantation site, duration of exposure, and the biological environment involved.^[7] This dependency arises because the same material may provoke varying responses—ranging from benign tolerance to inflammation or rejection—based on these variables, necessitating tailored evaluations for each use case.^[8] Assessments of biocompatibility occur at both the material level, which examines individual components in isolation, and the device level, which considers the assembled product's overall performance and potential interactions among multiple materials.^[9] In multi-material devices, such as those combining polymers, metals, and ceramics, this distinction complicates evaluations, as synergies or conflicts (e.g., catalytic degradation induced by metal ions on polymers) may only emerge at the device level, requiring integrated testing beyond isolated material analyses.^[10] The term "biocompatibility" was first introduced in scientific literature in 1970 by researchers R.J. Hegyeli and C.A. Homsy, marking the beginning of formalized discourse on material-host interactions in implants.^[11]

Importance in Biomedical Applications

Biocompatibility plays a critical role in preventing adverse host responses to implantable medical devices and biomaterials, such as inflammation triggered by foreign body reactions, thrombosis in blood-contacting devices, and systemic toxicity from leached ions or degradation products.^[12] These reactions can compromise device functionality and patient health, leading to complications like fibrous encapsulation or chronic immune activation that necessitate device removal.^[13] By ensuring materials elicit minimal detrimental biological responses, biocompatibility assessments enable the design of safer implants that integrate effectively with tissues, reducing the incidence of such events.^[14] The clinical success of biomedical applications heavily depends on biocompatibility, as non-compatible materials contribute significantly to implant failures and subsequent revisions. For instance, approximately 10% of metal-on-metal total hip replacements implanted since 1988 required revision, often due to biocompatibility-related issues like ion release and inflammatory responses.^[15] In orthopedic implants, failure rates linked to poor biocompatibility, including corrosion and allergic reactions, contribute to complications across various studies, underscoring the need for optimized material selection to enhance long-term durability and patient satisfaction.^[14] Beyond immediate clinical risks, biocompatibility has profound implications for patient outcomes, regulatory pathways, and healthcare economics. Effective biocompatibility supports better recovery times, lower complication rates, and improved quality of life, while non-compliance can delay regulatory approval; for example, adherence to ISO 10993 standards is essential for FDA clearance of medical devices.^[4] Economically, a single material failure in a medical device can cost up to $3.5 million, encompassing recalls, litigation, and revision surgeries, with broader device-related errors contributing over $1 billion annually to U.S. healthcare burdens as of 2008.^[16][17] Furthermore, biocompatibility underpins innovations in personalized medicine by allowing tailored biomaterial formulations that match individual patient physiologies, such as immune profiles or genetic predispositions, to optimize therapeutic efficacy and minimize rejection risks.^[18] This enables advancements like patient-specific implants and regenerative therapies, where precise material-host interactions drive improved outcomes in fields like tissue engineering and drug delivery.^[18]

Historical Development

Origins and Early Concepts

The concept of biocompatibility originated in the early 20th century amid initial attempts to integrate foreign materials into the human body for medical purposes, with a primary emphasis on selecting substances that exhibited minimal adverse reactions through trial-and-error approaches. Materials like gold, used in ancient dental prosthetics as early as 630 BCE by the Etruscans for stabilizing loose teeth, and painted clay, employed in rudimentary external ocular prosthetics by Egyptian and Roman civilizations around the 5th century BCE, were chosen for their perceived inertness and durability in biological settings.^[19][20][21] By the 1960s, advancements in materials science led to a more systematic exploration of inert biomaterials for implants, aiming to encapsulate them within fibrous tissue without eliciting inflammation or rejection. Vitreous carbon emerged as a prominent example during this decade, valued for its high chemical stability, biocompatibility, and resistance to degradation; it was tested in dental and orthopedic applications, such as tooth replicas and joint components, demonstrating low toxicity and stable integration in animal models.^[22][19] Pioneering researchers like C. William Hall played a crucial role in elucidating implant-host interactions, conducting foundational studies on artificial organs and tissue responses to synthetic materials, which highlighted the need for materials that could coexist harmoniously with living systems without causing systemic effects.^[23][24] The term "biocompatibility" was coined in 1970, marking a pivotal moment in formalizing these ideas; it first appeared in a conference abstract by R.J. Hegyeli at the American Chemical Society Annual Meeting and in a publication by C.A. Homsy and colleagues, who defined it in the context of selecting implant materials that promote healing and avoid incompatibility factors like corrosion or excessive fibrosis.^[11] In the early 1970s, the term rapidly entered scientific literature, signaling a conceptual evolution from strict inertness—where materials were expected to provoke no response—to "tissue compatibility," emphasizing controlled biological interactions that support function and repair.^[25][26] This shift was driven by observations that truly inert materials often led to suboptimal encapsulation, prompting a reevaluation toward materials capable of eliciting beneficial host responses.^[27]

Evolution of Key Definitions and Standards

In the 1970s and 1980s, definitions of biocompatibility primarily emphasized the absence of toxicity, as exemplified by Dorland's Illustrated Medical Dictionary, which described it as "the quality of not having toxic or injurious effects on biological systems."^[28] This view shifted toward performance-based assessments by the late 1980s and 1990s, incorporating comparisons of tissue responses to materials versus controls, as outlined in ASTM standards like F748 (1982), which provided guidelines for selecting generic biological test methods such as cytotoxicity and sensitization to evaluate material-host interactions.^[29] This evolution reflected growing recognition that biocompatibility involved not just non-toxicity but functional integration in specific biological contexts, moving from phenomenological observations to standardized toxicological evaluations.^[30] The development of early standards further formalized these concepts, with ASTM guidelines in the 1980s—such as F619 (1979) for extraction testing and F981 (1986) for long-term implantation—laying groundwork for international harmonization.^[29] This culminated in the first edition of ISO 10993-1 in 1992, which established a framework for biological evaluation of medical devices, emphasizing risk-based testing over mere toxicity avoidance and integrating ASTM principles into a global standard.^[29] A pivotal refinement occurred at the 2005 consensus conference in Sorrento, Italy, organized by the European Society for Biomaterials, where participants proposed sub-definitions tailored to applications, distinguishing biocompatibility for long-term implants (focusing on stable integration) from short-term implants (emphasizing minimal acute reactions) and tissue engineering scaffolds (requiring active promotion of regeneration).^[31] David F. Williams' 2008 reevaluation built on these developments, critiquing five common definitions—including Dorland's toxicity-focused view, his own 1999 formulation ("the ability of a material to perform with an appropriate host response in a specific situation"), the ASTM's tissue-response comparison, an expanded version of his earlier work, and the International Union of Pure and Applied Chemistry's dictionary entry—for their limitations in addressing mechanistic interactions and application-specific needs.^[32] He advocated shifting from ideals of inertness to eliciting an "appropriate host response," where materials actively support desired biological outcomes like resolution of inflammation or tissue remodeling, rather than passive non-reactivity.^[33] This perspective underscored biocompatibility as a dynamic process, influencing subsequent regulatory and research frameworks.^[32]

Evaluation and Testing

In Vitro Testing Methods

In vitro testing methods evaluate the biocompatibility of materials and devices by simulating biological interactions in controlled laboratory environments, typically using cell cultures or isolated biological components, without involving whole living organisms. These tests focus on direct cellular and molecular responses to assess potential toxicity, ensuring materials do not elicit adverse effects at the initial stages of development.^[34] Cytotoxicity assays measure the potential of biomaterials to impair cell viability or proliferation, serving as a primary screen for material safety. The MTT (3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide) assay is a widely used colorimetric method that quantifies metabolically active cells by assessing the reduction of MTT to formazan crystals, which are solubilized and measured for absorbance at 570 nm using a spectrophotometer. This absorbance value correlates directly with the number of viable cells, allowing researchers to determine cytotoxicity levels through comparison to untreated controls.^[35][36] Other cytotoxicity tests, such as those outlined in ISO 10993-5, employ similar principles to evaluate extracts or direct contact of materials with cell lines like fibroblasts or epithelial cells.^[37] Hemocompatibility tests assess interactions between materials and blood components, particularly focusing on damage to red blood cells. The hemolysis assay, standardized under ASTM F756, involves incubating material extracts or surfaces with diluted anticoagulated blood, followed by centrifugation to separate intact cells from free hemoglobin. The degree of hemolysis is quantified by spectrophotometric measurement of released hemoglobin at approximately 540 nm, with results expressed as a percentage relative to positive (distilled water) and negative (saline) controls; values below 5% indicate non-hemolytic materials.^[38][34] This method helps identify potential thrombogenic or hemolytic risks for blood-contacting devices. Genotoxicity tests detect whether biomaterials can induce DNA damage or mutations, which could lead to long-term adverse effects. The Ames test, a bacterial reverse mutation assay, uses histidine-requiring strains of Salmonella typhimurium or Escherichia coli exposed to material extracts, with and without metabolic activation via S9 liver fraction; mutagenic potential is determined by increased revertant colonies compared to controls, indicating the material's ability to cause point mutations.^[2] This test is particularly valuable for screening leachables from polymers or metals in implants. Irritation tests evaluate the potential for materials to cause inflammatory responses on epithelial surfaces. Models using reconstructed human epidermis (RhE), such as those based on keratinocyte cultures forming a stratified skin-like barrier, expose the tissue to material extracts for 42 hours, followed by assessment of cell viability via MTT assay; a viability below 50% of the negative control classifies the material as an irritant.^[39][40] These 3D models mimic human skin architecture, providing a reliable alternative for predicting dermal irritation without animal use. In vitro methods offer several advantages, including cost-effectiveness due to reduced resource needs compared to animal studies, ethical benefits by avoiding live subjects, and high-throughput capabilities for screening multiple materials rapidly.^[41] However, they have limitations, such as the inability to capture complex systemic or immune responses that occur in vivo, potentially leading to incomplete predictions of long-term biocompatibility.^[13] These tests thus serve as essential preliminary steps, often requiring validation through subsequent in vivo assessments.

In Vivo and Clinical Assessment

In vivo assessment of biocompatibility involves evaluating the interactions between medical devices and living organisms, particularly to detect local, systemic, and long-term biological responses that cannot be fully captured through in vitro methods. These tests typically use animal models to simulate clinical conditions, assessing tissue integration, inflammation, and potential toxicity over durations matching the device's intended use. Standards such as ISO 10993 guide these evaluations, emphasizing the selection of appropriate animal species, implantation sites, and endpoints like histopathological analysis to ensure the device's safety profile.^[42][9] Implantation studies, detailed in ISO 10993-6, focus on local effects by implanting device materials or extracts into animal tissues, such as subcutaneous or muscular sites in rabbits, to monitor acute and chronic responses. Intracutaneous reactivity tests, for example, involve injecting device extracts into rabbit skin to evaluate irritation and inflammation through macroscopic and microscopic observations, with responses compared against controls to determine biocompatibility. These studies assess tissue response evolution, including encapsulation, degradation, or integration, and are essential for devices with prolonged tissue contact, often extending to 90 days or longer based on material properties. Histopathological examination at multiple time points helps quantify inflammation scores and cellular infiltration, providing evidence of non-adverse local effects.^[42][9][43] Systemic toxicity evaluations, governed by ISO 10993-11, examine the potential for device leachables to cause widespread adverse effects through routes like implantation, oral, or intravenous administration in rodents. Acute tests detect immediate responses within 72 hours, while subchronic studies, such as 90-day implantation in rats, monitor organ function, body weight changes, and hematological parameters to identify chronic toxicity targets like the liver or kidneys. These rodent-based assays establish no-observed-adverse-effect levels (NOAELs) and are conducted under good laboratory practice (GLP) to support risk assessments, ensuring that systemic exposure does not lead to functional impairments.^[44][9][45] Pyrogenicity testing addresses the risk of fever-inducing contaminants using the rabbit pyrogen test outlined in USP <151>, where device extracts are injected intravenously into rabbits, and rectal temperature is monitored for up to three hours. A temperature rise exceeding 0.5–1.4 °C (depending on dose) indicates pyrogenic potential, helping to rule out non-endotoxin pyrogens in devices contacting blood or cerebrospinal fluid. This in vivo method remains relevant for certain medical devices when bacterial endotoxin tests are insufficient, providing a direct measure of physiological response akin to human fever mechanisms.^[46][47] Sensitization tests, per ISO 10993-10, evaluate the potential for delayed hypersensitivity reactions using animal models like guinea pigs in assays such as the Guinea Pig Maximization Test or Buehler test. These involve repeated topical or intradermal applications of device extracts followed by a challenge phase, with skin responses scored for erythema and edema to detect allergic contact dermatitis. Positive responses indicate sensitizing potential, guiding material selection to prevent clinical hypersensitivity in patients with prolonged device exposure.^[48][49] Clinical assessment extends biocompatibility evaluation into human trials, where Phase I focuses on safety in small cohorts, Phase II assesses efficacy and adverse events in larger groups, and Phase III confirms long-term performance in diverse populations. Monitoring includes tracking biocompatibility-related adverse events, such as localized inflammation or implant rejection, through metrics like visual analog scales for pain, serological markers, or imaging techniques (e.g., MRI for tissue response). Clinical data from these phases can validate preclinical findings, demonstrating no material-induced toxicities and supporting regulatory approval when adverse event rates remain below predefined thresholds.^[9][50]

Regulatory Standards and Guidelines

The International Organization for Standardization (ISO) 10993 series provides the primary global framework for the biological evaluation of medical devices, emphasizing a risk-based approach to assess biocompatibility within overall risk management processes. ISO 10993-1, titled "Biological evaluation of medical devices — Part 1: Evaluation and testing within a risk management process," outlines the general principles for identifying potential biological risks, categorizing devices based on the nature and duration of body contact (such as limited exposure under 24 hours, prolonged from 24 hours to 30 days, or permanent exceeding 30 days), and selecting appropriate tests to fill data gaps.^[51] This standard, originally published in 1992 and revised multiple times, was updated in its fifth edition in 2018 and further revised in 2025 to incorporate enhanced guidance on foreseeable misuse, bioaccumulation risks, and lifecycle considerations from manufacturing to end-use.^[51] Specific parts address targeted evaluations, such as ISO 10993-4:2017, which guides the selection of tests for interactions with blood (hemocompatibility), classifying devices into categories like external communicating and implant devices to evaluate risks such as thrombosis or hemolysis without prescribing exact protocols due to evolving scientific knowledge.^[52] Recent updates in the 2020s, including ISO 10993-18:2020 on chemical characterization and an annex in the 2025 revision of ISO 10993-1 on genotoxicity evaluation, address emerging concerns with nanomaterials by recommending physicochemical characterization to identify potential hazards like particle size and surface reactivity.^[53][54] In the United States, the Food and Drug Administration (FDA) endorses the ISO 10993-1 framework for biocompatibility assessments in premarket submissions, including Premarket Approvals (PMAs), Humanitarian Device Exemptions (HDEs), Investigational Device Exemptions (IDEs), 510(k) notifications, and De Novo classifications. The FDA's guidance document, "Use of International Standard ISO 10993-1," originally issued in 2016 and updated on September 8, 2023, clarifies how to apply this standard within a risk management process, emphasizing chemical characterization as a foundational step to reduce unnecessary animal testing and recommending a risk-based categorization of devices (e.g., surface devices for intact skin versus implant devices).^[4] This update aligns with the 2018 ISO version but anticipates further harmonization with the 2025 revision, particularly for devices involving submicron or nanotechnology materials, where extractables and leachables must be evaluated for potential systemic effects.^[4] Manufacturers are required to justify testing selections based on device history, prior data, and intended use, with the FDA reviewing submissions to ensure patient safety without mandating every ISO test unless risks warrant it. Under the European Union's Medical Device Regulation (MDR) 2017/745, effective since May 26, 2021, biocompatibility is integral to demonstrating general safety and performance requirements, particularly under Annex I, Section 10, which mandates that devices must be biocompatible to avoid adverse tissue reactions or toxicity.^[55] Manufacturers must compile comprehensive biocompatibility data—including risk assessments, chemical analyses, and test results aligned with ISO 10993—within the technical documentation outlined in Annexes II and III, retained for at least 10 years (or 15 years for implants).^[55] For Class IIa, IIb, and III devices, Notified Bodies conduct conformity assessments under Annex IX, scrutinizing this documentation to verify compliance, with expert panels reviewing clinical evaluations for higher-risk categories to ensure no unacceptable biological risks remain.^[55] The MDR harmonizes with ISO 10993 standards, requiring updates for legacy devices and emphasizing post-market surveillance to monitor long-term biocompatibility. Global harmonization efforts are advanced by the International Medical Device Regulators Forum (IMDRF), which facilitates convergence of regulatory practices across member countries, including through working groups established post-2021. The IMDRF's Software as a Medical Device (SaMD) Working Group, active since 2014 and ongoing, integrates digital health technologies into risk-based frameworks, indirectly supporting biocompatibility evaluations for hybrid devices combining software with material components by promoting consistent definitions and assessment principles.^[56] These initiatives build on earlier IMDRF efforts to align ISO 10993 application, reducing duplicative testing and enhancing international reciprocity for biocompatibility data in submissions.^[57]

Factors Influencing Biocompatibility

Material and Surface Properties

The biocompatibility of materials is profoundly influenced by their chemical composition, which dictates interactions with physiological environments. In metallic biomaterials, such as titanium and its alloys, the spontaneous formation of a thin TiO₂ oxide layer (typically 10–30 nm thick) upon exposure to air provides essential corrosion resistance by acting as a passive barrier that minimizes ion release in bodily fluids.^[58] This layer enhances overall biocompatibility by promoting selective protein adsorption, such as albumin and fibronectin, which supports cell attachment and osseointegration without eliciting significant toxicity.^[58] For polymers, compositional factors like hydrophilicity play a key role; surfaces with lower water contact angles (indicating higher hydrophilicity) exhibit improved biocompatibility by reducing monocyte adhesion and inflammatory responses compared to hydrophobic counterparts.^[59] Contact angle measurements, often below 90° for hydrophilic polymers, quantify this property and guide material selection for minimizing restenosis in applications like drug-eluting stents.^[59] Surface modifications further tailor biocompatibility by altering physical characteristics at the material-tissue interface. Topographical features, including nanoscale roughness, are critical; increased surface roughness, as measured by scanning electron microscopy (SEM) or atomic force microscopy (AFM), can enhance cell adhesion and proliferation by providing cues that influence filopodia extension and mechanosensing.^[60] For instance, anodized titanium surfaces with controlled nanopores (e.g., 57 nm radius) demonstrate improved bioactivity through heightened early-stage osteoblast attachment.^[60] Coatings represent another vital modification strategy; poly(ethylene glycol) (PEG) layers confer anti-fouling properties by forming a hydration barrier that significantly reduces nonspecific protein adsorption, thereby preserving biocompatibility and stealth-like behavior in biological media.^[61] These coatings support applications where minimizing immune recognition is essential.^[61] Degradation profiles of materials are equally determinant for biocompatibility, particularly in resorbable systems where controlled breakdown prevents adverse local responses. Bioresorbable polymers like poly(lactic-co-glycolic acid) (PLGA) undergo hydrolytic degradation through cleavage of ester bonds, yielding lactic and glycolic acid byproducts that are metabolized via the tricarboxylic acid cycle and safely excreted.^[62] The degradation rate, tunable by the lactide:glycolide ratio (e.g., fastest for 50:50 PLGA), ensures biocompatibility by avoiding accumulation of acidic fragments that could cause inflammation, with FDA approval affirming its safety for controlled drug delivery.^[62] An illustrative example is Mg-Zn-Ca-based metallic glasses, which achieve biocompatibility in degradable stents via their amorphous structure that promotes uniform corrosion resistance, forming protective ZnO/Zn(OH)₂ layers and limiting hydrogen evolution to rates as low as 0.010 mm/year in simulated body fluids.^[63] This controlled corrosion supports tissue remodeling without excessive degradation products.^[63]

Host Biological Responses

When a biocompatible material is introduced into the body, it elicits an acute inflammatory response as the initial host biological reaction, primarily involving the recruitment of neutrophils and activation of the innate immune system. This phase, occurring within hours to days, is characterized by the foreign body reaction, where macrophages adhere to the material surface, fuse into multinucleated giant cells, and release pro-inflammatory cytokines such as interleukin-1 (IL-1), tumor necrosis factor-alpha (TNF-α), and interleukin-6 (IL-6) to orchestrate the inflammatory cascade.^[64][65] These cytokines amplify vascular permeability and chemotaxis, promoting clearance of potential pathogens but also contributing to localized tissue damage if unresolved.^[66] If the material persists, the response transitions to a chronic phase, typically after 2-4 weeks, dominated by persistent macrophage activation and fibroblast proliferation leading to fibrosis and encapsulation. Fibrotic capsules form as collagen deposition isolates the implant, potentially impairing device function by creating a barrier that limits nutrient diffusion and integration.^[67] Additionally, hypersensitivity reactions, particularly Type IV delayed-type hypersensitivity mediated by T-lymphocytes, can occur in sensitized individuals, involving antigen presentation by dendritic cells and subsequent CD4+ T-cell activation that sustains inflammation without humoral antibodies.^[68][69] In favorable cases, host responses can promote healing and integration, as seen in osseointegration of bone implants where direct bone-to-implant contact is achieved through osteoblast differentiation and mineralization. This process is regulated by bone morphogenetic proteins (BMPs), particularly BMP-2 and BMP-7, which activate Smad signaling pathways to enhance osteogenesis and vascularization at the implant interface.^[70][71] To mitigate adverse responses, strategies for immune modulation include the application of stealth coatings, such as polyethylene glycol (PEG)-based layers on biomaterial surfaces, which reduce protein adsorption and minimize complement system activation by inhibiting C3 opsonization and subsequent anaphylatoxin release.^[72][73] These coatings promote a more tolerant immune environment, potentially shifting macrophage polarization toward anti-inflammatory phenotypes and improving long-term biocompatibility.^[74]

Applications

Implantable Medical Devices

Implantable medical devices encompass a wide range of prosthetics and therapeutic tools designed for long-term or temporary placement within the body, where biocompatibility is paramount to minimize adverse tissue reactions, ensure structural integrity, and promote functional integration. These devices, such as orthopedic implants and cardiovascular stents, must resist corrosion, support cellular adhesion without toxicity, and avoid chronic inflammation to achieve clinical success. Biocompatibility in this context involves not only material inertness but also the orchestration of host-device interactions that facilitate healing and prevent complications like thrombosis or rejection.^[75][2] Orthopedic implants, particularly hip replacements, frequently utilize cobalt-chromium (CoCr) alloys due to their high strength, wear resistance, and favorable biocompatibility profile, which supports long-term osseointegration—the direct structural and functional connection between bone and implant surface. CoCr alloys, such as Co-28Cr-6Mo, exhibit low cytotoxicity and stimulate bone neoformation, as demonstrated in studies on additively manufactured variants that enhance surface porosity for improved cellular attachment and vascularization. However, challenges arise from potential ion release, such as cobalt and chromium, which can induce hypersensitivity reactions in susceptible patients, necessitating rigorous preoperative screening and alloy modifications like titanium coatings to bolster osteointegrative properties without compromising mechanical performance.^[76][77][2] In cardiovascular applications, drug-eluting stents represent a cornerstone of biocompatibility engineering, featuring metallic scaffolds coated with biocompatible polymers that release antiproliferative agents like sirolimus or paclitaxel to inhibit neointimal hyperplasia and prevent restenosis—the re-narrowing of the artery post-implantation. These polymers, often durable types like poly(ethylene-co-vinyl acetate) blended with poly(n-butyl methacrylate), ensure controlled drug release while maintaining hemocompatibility to reduce thrombosis risk, with clinical trials showing restenosis rates dropping below 10% in the first year compared to bare-metal stents. Despite these advances, late-stage hypersensitivity to polymer degradation products can lead to incomplete endothelialization, prompting the development of bioabsorbable polymers that degrade into non-toxic byproducts, thereby enhancing long-term vessel patency and biocompatibility.^[78][79][80] A notable case study highlighting biocompatibility risks involves silicone breast implants, where textured surfaces were linked to breast implant-associated anaplastic large cell lymphoma (BIA-ALCL), a rare T-cell lymphoma, prompting global recalls starting in 2011 and culminating in the FDA's 2019 voluntary recall of Allergan's BIOCELL textured implants after identifying over 1,500 cases worldwide as of 2024. The textured silicone elastomer, intended to reduce capsular contracture by promoting tissue ingrowth, inadvertently fostered chronic inflammation and bacterial biofilm formation in some patients, elevating lymphoma risk to approximately 1 in 3,000 for textured devices (with estimates ranging from 1 in 3,000 to 1 in 30,000) versus negligible for smooth ones. This incident underscored the need for surface texturing strategies that balance adhesion benefits with immune surveillance, leading to enhanced ISO 10993 biocompatibility testing protocols for polymeric implants.^{[81][82][83][84][85]} Multi-material challenges are exemplified in pacemaker leads, which combine conductive metals like platinum-iridium coils with insulating polymers such as silicone or polyurethane to transmit electrical impulses while isolating tissues from current leakage. These hybrid constructions demand balanced biocompatibility, as metal-polymer interfaces can degrade over time, releasing particulates that trigger inflammation or thrombosis, with historical issues like polyether polyurethane biodegradation causing lead fractures in up to 5% of cases within five years. Surface coatings, including antimicrobial silver ions or heparin, have been employed to mitigate infection risks and improve endothelialization at the lead tip, ensuring device longevity exceeding 10 years in most patients through optimized material synergies that prevent galvanic corrosion and hypersensitivity.^[86][87][88]

Tissue Engineering and Regenerative Medicine

In tissue engineering and regenerative medicine, biocompatibility is defined as the ability of a scaffold or matrix to serve as a substrate that supports appropriate cellular activity, including molecular and mechanical signaling, to optimize tissue regeneration while avoiding undesirable local or systemic host responses. This definition, proposed by David F. Williams, emphasizes the material's role in facilitating cell adhesion, proliferation, differentiation, and extracellular matrix production without eliciting toxicity, inflammation, or genotoxicity. Such scaffolds must degrade controllably, releasing non-harmful byproducts to support long-term integration and function in the host environment. Hydrogels, such as alginate-based ones, are widely used for cell encapsulation in tissue engineering due to their high water content and biocompatibility, mimicking the native extracellular matrix to promote cell viability.^[89] These scaffolds require high porosity, typically exceeding 90%, to enable efficient nutrient and oxygen diffusion while allowing waste removal, which is critical for maintaining cell survival in avascular constructs.^[90] For instance, ionically crosslinked alginate hydrogels achieve uniform pore sizes that support uniform cell distribution and metabolic exchange, enhancing biocompatibility by preventing hypoxic stress or necrosis in encapsulated cells. Decellularized extracellular matrices (dECM) represent another key approach, providing natural scaffolds for organ regeneration by retaining bioactive cues like growth factors and structural proteins while removing cellular components to minimize immunogenicity.^[91] Effective decellularization protocols ensure low residual DNA and alpha-gal epitopes, resulting in scaffolds that elicit minimal immune responses and promote host cell repopulation for functional tissue reconstruction, as demonstrated in cardiac and vascular applications.^[92] Integration of induced pluripotent stem cells (iPSCs) into these scaffolds further advances regenerative therapies, but biocompatibility demands strategies to prevent teratoma formation from residual undifferentiated cells, which could lead to tumorigenic risks.^[93] Purification methods, such as small-molecule treatments like PluriSIn#1, selectively eliminate undifferentiated iPSCs by over 10^6-fold, ensuring safe transplantation without compromising differentiated cell potency or scaffold integration.^[93] This approach maintains the scaffold's supportive role for iPSC-derived lineages, fostering directed differentiation and vascularized tissue formation in vivo.^[94]

Challenges and Future Directions

Current Limitations and Risks

One major limitation in achieving reliable biocompatibility lies in the inherent variability of patient-specific responses, which can lead to unpredictable adverse outcomes. Factors such as age influence immune system function, with immunosenescence in older individuals resulting in altered inflammatory responses and reduced adaptive immunity, potentially exacerbating foreign body reactions to implants.^[95] Allergies represent another key source of variability; for example, nickel hypersensitivity, affecting 10-15% of the general population, can trigger systemic reactions or in-stent restenosis in patients receiving coronary stents containing nickel alloys like cobalt-chromium.^[96] These individual differences, including genetic predispositions and comorbidities, complicate standardized predictions of material-host interactions.^[97] Long-term degradation of biomaterials poses significant risks, particularly in orthopedic implants where mechanical wear generates particulate debris that drives chronic inflammation. In total hip and knee replacements, ultra-high molecular weight polyethylene (UHMWPE) wear particles elicit macrophage activation and osteoclastogenesis, leading to periprosthetic osteolysis and aseptic loosening, which accounts for a substantial portion of revision surgeries.^[98] This process underscores the challenge of maintaining material integrity over decades, as even minor degradation can initiate a cascade of biological responses compromising implant stability.^[99] Device-specific failure modes highlight persistent biocompatibility challenges despite engineering efforts. Ventricular assist devices (VADs), essential for end-stage heart failure, continue to face high thrombosis rates—up to 8.4% in some cohorts—even with heparin-coated surfaces designed to enhance hemocompatibility, often requiring intensified anticoagulation that introduces bleeding risks.^[100] Such issues arise from complex blood-material interactions, including platelet activation and fibrin deposition, which current coatings inadequately mitigate in high-shear environments.^[101] Gaps in current biocompatibility knowledge are particularly evident with nanomaterials, where chronic low-dose exposure effects remain underexplored. Regulatory testing often overlooks subtle, long-term toxicities such as nanoparticle accumulation in organs, which may induce oxidative stress or immune dysregulation without acute symptoms.^[102] Furthermore, the immunological properties of nanomaterials in vivo are incompletely characterized, limiting safe translation to clinical applications like drug delivery systems or tissue scaffolds.^[103]

Emerging Materials and Technologies

Nanomaterials have emerged as a promising class of biocompatible materials, particularly graphene oxide (GO) coatings, which exhibit strong antibacterial properties while minimizing host inflammatory responses. Studies post-2020 have demonstrated that GO coatings on substrates like mullite foams and poly-ether-ether-ketone (PEEK) implants effectively inhibit bacterial growth, such as Escherichia coli and Staphylococcus aureus, through physical disruption of cell membranes and oxidative stress induction, achieving up to 99% bacterial reduction in vitro.^[104][105] Furthermore, these coatings promote reduced inflammation in vivo, as evidenced by lower levels of pro-inflammatory cytokines like TNF-α in animal models of implantation, attributed to the material's high surface area and tunable functionalization that supports tissue integration without excessive immune activation.^[106][107] Smart materials, including pH-responsive shape-memory polymers (SMPs), offer dynamic control over biocompatibility in biomedical applications, enabling targeted drug release in response to physiological cues. These polymers, often based on polyurethane or polyester networks, recover their original shape upon pH shifts, such as from acidic tumor microenvironments (pH 6.5) to neutral physiological conditions (pH 7.4), facilitating controlled elution of therapeutics like hemin with release rates exceeding 95% over extended periods.^[108] Recent formulations demonstrate excellent biocompatibility, with cell viability rates above 90% in fibroblast and osteoblast cultures, and minimal cytotoxicity due to their biodegradable linkages that degrade into non-toxic byproducts, enhancing long-term implant safety.^[109][110] Bioinspired designs, such as peptide amphiphiles (PAs) that mimic the extracellular matrix (ECM), provide nanofibrous scaffolds with enhanced cellular compatibility by recapitulating natural tissue architecture and bioactivity. These self-assembling PAs, incorporating motifs like RGD for integrin binding, form hydrogels that support cell adhesion, proliferation, and differentiation, with biocompatibility confirmed through in vitro assays showing over 85% viability in neural and mesenchymal stem cells.^[111][112] In wound healing models, PA nanofibers promote ECM remodeling and reduce scar formation by upregulating collagen deposition, offering a biocompatible alternative to synthetic scaffolds.^[113] Recent developments from 2021 to 2025 in CRISPR-edited cells have advanced biocompatibility in regenerative therapies by engineering reduced immunogenicity for allogeneic transplants. CRISPR-Cas9 editing of HLA class I and II genes in regulatory T cells or donor organs, such as kidneys, disrupts immune recognition sites, potentially enabling "universal" cells or organs that evade host rejection.^[114] Clinical trials initiated in 2025 for HLA-edited kidney transplants, such as NCT07053462, are underway to evaluate this approach for off-the-shelf cellular therapies. However, challenges include ethical considerations around gene editing in organs and stringent regulatory requirements to ensure safety and efficacy.^[115][116]