Inorganic ABAs mainly include metal ions (e.g., Cu²⁺, Ag⁺, Zn²⁺) and their NPs form (Ghimire and Song, 2021; Chen et al., 2023; Liu J. et al., 2024). Their antibacterial mechanisms primarily arise from multiple, complementary, and often sequential mechanisms (Figure 2). Metal ions initially interact with the bacterial envelope, inducing membrane depolarization, lipid peroxidation, protein denaturation, and osmotic imbalance. Subsequent intracellular penetration causes oxidative stress, metabolic disruption, enzyme inactivation, and, in some cases, direct genotoxicity through interactions with nucleic acids (Mahmoudi et al., 2022). These interconnected pathways collectively contribute to bactericidal efficacy.

Although the ultimate mechanisms of toxicity are similar across bacterial taxa, the architecture of the cell envelope significantly influences the kinetics and efficiency of metal ion entry. For instance, in Gram-negative bacteria, the presence of an outer membrane slows initial Cu²⁺ penetration, while periplasmic chaperones and homeostasis proteins (e.g., CopA, CueO, CusCFBA) buffer Cu ions before they reach cytoplasmic targets (Bondarczuk and Piotrowska-Seget, 2013; Solioz, 2018). In contrast, Gram-positive bacteria lack an outer membrane and periplasmic compartment (Zhang et al., 2023), leading to more rapid access to the cytoplasmic membrane and faster accumulation of intracellular Cu²⁺ (Solioz et al., 2010; Dupont et al., 2011). These structural differences modulate transport dynamics but do not fundamentally alter the primary antibacterial mechanisms.

To date, extensive research has been focused on incorporating Ag into implants. However, Ag-containing alloys may form soluble Ag⁺ that can enter systemic circulation and accumulate in tissues, raising concerns over long-term toxicity (Mo et al., 2007; Tamay et al., 2022). Moreover, the antibacterial performance of Ag⁺ is sensitive to environmental conditions such as pH, humidity, and temperature, which can undermine reproducibility in vivo. Cu²⁺ has emerged as a promising alternative due to its strong antimicrobial activity, lower cost, and essential physiological roles in enzymatic and hematopoietic processes (Yang et al., 2023). Nevertheless, Cu²⁺ release must be carefully controlled to maintain antibacterial efficacy without inducing cytotoxicity (Zhang et al., 2019).

Accumulating evidence indicates that bacteria can develop adaptive responses to metal ions, challenging the earlier assumption that metal-based agents are inherently resistance-proof. Adaptive resistance mechanisms include the upregulation of efflux systems (e.g., CusCFBA, CopA), enhanced metal sequestration by periplasmic chaperones and metallothioneins, alterations in membrane composition, and increased production of EPS that bind or immobilize metal ions (Hobman and Crossman, 2015; Chandrangsu et al., 2017). Biofilm-associated physiological changes and the induction of oxidative stress–response pathways further reduce susceptibility. Although these responses are often reversible rather than genetically fixed, they can significantly dampen antibacterial activity, particularly under chronic low-dose exposure. These findings underscore the need for metal-based coatings that minimize prolonged sublethal ion release and incorporate multi-target strategies to mitigate adaptive responses.

NPs-based ABAs introduce additional bactericidal mechanisms beyond ion release. Their physical interaction with cell walls, direct membrane penetration, and catalytic participation in intracellular redox reactions contribute to enhanced antibacterial activity (Slavin et al., 2017; Linklater et al., 2020). It should be noted that the antibacterial activity of NPs is strongly dependent on physicochemical parameters. Firstly, smaller NPs often display greater antibacterial activity due to higher surface area–to–volume ratios (Xie et al., 2020a). However, few studies have shown that larger NPs can be more effective in certain contexts (El Badawy et al., 2011; Sohm et al., 2015), indicating that size alone may not be the most critical factor determining their toxicity. Secondly, shape also plays a critical role: nanocubes and nanorods frequently outperform other morphologies due to their exposed reactive facets and favorable oxidation states (Xie et al., 2023). This observation is bolstered by analyses indicating that facets with lower stability require less energy to generate oxygen vacancies, contributing to enhanced catalytic and bactericidal behavior (Thakur et al., 2024). Thirdly, surface charge further influences activity. This is because positively charged NPs interact electrostatically with negatively charged bacterial envelopes, increasing adsorption and membrane disruption (Gao et al., 2020; Godoy-Gallardo et al., 2021).

Despite their potential, bacteria may also develop tolerance to NPs by increasing EPS secretion, modifying membrane rigidity, altering motility and division patterns, and activating oxidative-stress defense systems. These adaptive phenomena reinforce the need for long-term evaluation of NP-based coatings beyond short-term antimicrobial assays. Moreover, standardized NPs fabrication protocols and clinically relevant dosing frameworks remain lacking. Additionally, optimal NPs concentrations must balance antibacterial efficacy with cytotoxicity and inflammatory risks.

Beyond organic and inorganic ABAs, enzyme (e.g., lysozyme, acylase), organic cationic compounds (e.g., quaternary ammonium compounds, chlorhexidine, octenidine, cationic surfactants, chitosan), organic non-cationic compounds (e.g., furanones, triclosan), and other non-organic compounds (e.g., nitric oxide) can also be employed as ABAs, with mechanisms reviewed elsewhere (Cloutier et al., 2015).

It deserves particular attention that, bacteriophages (phages), including lytic (virulent) and lysogenic (temperate) types (Strathdee et al., 2023), represent emerging antibacterial tools as well. Specifically, lytic phages recognize bacterial hosts, inject their genomes, replicate, and release progeny through host cell lysis. In contrast, temperate phages integrate their DNA into the host’s chromosome, entering a dormant prophage state until triggered by external stimuli (e.g., UV light, heat, or chemical agents) (Bayat et al., 2021). Therefore, lytic phages are preferred for therapeutic applications due to their ability to avoid horizontal gene transfer of virulence factors (Melo et al., 2020). Although phage-functionalized implant coatings remain underexplored, successful phage monotherapy and phage-antibiotic combinations in clinical studies highlight strong translational potential (Onsea et al., 2020).

However, bacteria can deploy multiple anti-phage defense mechanisms (Stanley and Maxwell, 2018; Addo et al., 2024), including CRISPR-Cas systems, adsorption inhibition, restriction-modification systems, and EPS-mediated physical shielding, especially within biofilms. These challenges underscore the need for further investigation into phage–biofilm interactions before implant applications become feasible.

Initial bacteria-substrate interactions are governed by multiple factors, including van der Waals forces, Brownian motion, and electrostatic and hydrophobic interactions (James et al., 2017; Yongabi et al., 2020). Consequently, the physical structure and chemical characteristics of implant surfaces play a critical role in determinging irreversible bacterial adhesion. Therefore, modifying these surface characteristics has emerged as a highly promising passive strategy to repel bacteria and prevent early biofilm formation.

The inherent releasing kinetics of the loaded ABAs play a decisive role in determining the antibacterial performance of implant coatings, independent of the specific type of agent used, as discussed in Section 2. Generally, drug release follows first- or second-order kinetics, characterized by an initial burst phase in which a disproportionately large amount of the loaded agent is rapidly released after implantation, followed by a slower, sustained release stage (Chen et al., 2021). Although this mode of release can effectively eliminate pathogens within a short period, it also presents several drawbacks, including cytotoxicity to surrounding host cells such as osteoblasts, tissue irritation, and a limited effective duration. Therefore, optimal antibacterial function depends not only on the ABAs itself but also on the physicochemical characteristics of the coating. Three key parameters must therefore be considered when designing an effective sacrificial layer: the enhancement of drug-loading capacity, the regulation of release rate, and the extension of release duration (Chen et al., 2021; Ghimire and Song, 2021; Liu et al., 2024b). The primary coatings for the controlled release of ABAs are discussed herein.

Coatings used for controlled ABA release can be broadly categorized into conservative systems, which release drugs without external triggers, and smart systems, which respond to endogenous stimuli such as pH, enzymes, or redox conditions (Figure 2). Specifically, both nondegrading and biodegradable carriers have been used to serve as conservative release killing coatings, including hydrogels, chitosan, and HA, all of which have been explored as drug reservoirs (Shariatinia, 2019). In order to tailor the release kinetics of these coatings, methods are being devised to adjust matrix properties such as porosity, surface roughness, and functional groups (Cloutier et al., 2015). Beyond bulk matrix modification, controlled release can also be achieved by incorporating specialized nanostructures that allow fine-tuning of release behavior. These include nanotubes, nanowires, dendrimers, and nanocapsules, which offer adjustable geometric and interfacial parameters (Vallet-Regí et al., 2020). For instance, the modification of titania nanotubes (TNTs) is promising to control the release rate and prolong the release time. This is because it has been demonstrated that the parameters of TNTs, including the diameter, length, and aspect ratio of the nanotubes, significantly influence drug release (Lin et al., 2016). Zhang et al. (2017) engineered dual-diameter TNTs, with upper diameters of 35 or 70 nm and a base diameter of 140 nm, to improve the sustained release of AMPs. Their system successfully maintained release for at least 60 days, attributed to the slowed diffusion through the narrower upper openings, whereas single-diameter TNTs of 140 nm exhausted their drug payload within 42 days. Such findings highlight the importance of rational nanotube design for achieving long-term antibacterial performance.

With respect to ‘smart’ coatings, a variety of endogenous stimuli, including acidic microenvironments, elevated temperature, and abnormal enzymatic activities, have been employed to achieve controlled and on-demand release of ABAs (Figure 2). Specifically, at infection sites, rapid bacterial proliferation and local hypoxia promote anaerobic glycolysis, resulting in the accumulation of acidic metabolites that lower the surrounding pH (Hu et al., 2023). This microenvironmental shift enables the selective cleavage of acid-labile chemical bonds such as ester, imine, and Schiff-base linkages, or triggers protonation of functional groups including acrylic acid, tertiary amines, or sulfonamides, thereby inducing structural transitions that facilitate pH-responsive delivery (Yuan et al., 2018). FFor instance, a hierarchical implant coating was developed in which a polydopamine layer was first deposited onto the implant surface and subsequently functionalized with an initiator to polymerize ethylenediamine-functionalized poly(glycidyl methacrylate). Gentamicin sulfate was then conjugated to the polymer network through an acid-sensitive Schiff-base bond, enabling both sustained and on-demand drug release under acidic conditions (Jin et al., 2019).

Temperature-responsive coatings have also been explored, taking advantage of the localized temperature elevation that accompanies infection. For instance, Ti implants coated with a thermosensitive poly(di(ethylene glycol) methyl ether methacrylate) (PDEGMA) brush loaded with levofloxacin demonstrated spatiotemporally controlled drug release, wherein collapse of PDEGMA chains above the lower critical solution temperature facilitated rapid antibiotic liberation. Both in vitro and in vivo studies confirmed that such thermoresponsive coatings significantly reduced bacterial colonization and subsequent infection risk (Choi et al., 2021). Enzyme-responsive systems exploit the elevated levels of bacterial or host-derived enzymes at infected sites. Specifically, pathogenic bacteria secrete enzymes such as hyaluronidase, alkaline phosphatase, and glutamyl endonuclease, while macrophages and neutrophils recruited to the infection site can release additional hydrolases, including matrix metalloprotease-9 and cholesterol esterase. These distinct enzymatic signatures have been harnessed to trigger controlled and targeted ABAs (Elkington et al., 2005). For instance, polyphosphoesters were employed as enzyme-cleavable cross-linkers in a minocycline-loaded chitosan membrane designed to degrade in response to the heightened alkaline phosphatase levels characteristic of periodontal infections (Li N. et al., 2019).

Despite their conceptual appeal, stimulus-responsive coatings face an important limitation: their reliance on sufficiently strong microenvironmental cues to activate drug release (He et al., 2022; Hu et al., 2023; Zhang et al., 2024). In early, mild, or spatially heterogeneous infections, the magnitude of pH changes, temperature elevation, or enzymatic activity may remain below the activation threshold required to initiate release. Such under-activation poses the risk of a “false negative” scenario in which pathogenic bacteria are present but the smart coating remains inert, thereby compromising early-stage antibacterial defense. This vulnerability highlights the need for stimulus-responsive platforms to incorporate design strategies that reduce dependence on a single trigger. These may include integrating a low-level baseline release to maintain protective drug concentrations, or developing multistimuli-responsive systems that combine pH, enzymatic, or redox sensitivities. Furthermore, smart coatings may function most effectively when integrated with complementary mechanisms, such as contact killing or passive anti-adhesion, to prevent bacterial colonization even when stimuli are insufficient to activate release. Although these adaptive platforms offer promising opportunities for infection-targeted therapy, questions regarding biocompatibility, manufacturability, response efficiency, and long-term stability must be addressed before their clinical translation.

To provide a consolidated comparison of the diverse antibacterial surface modification strategies reviewed in Sections 3 and 4, Table 1 summarizes their underlying mechanisms, advantages, limitations, and associated risks of antibacterial resistance.

Four primary technologies are used to directly coat metal ions and their NPs derivatives onto implant surface, including electrochemical techniques, hydrothermal treatment, physical methods consisting of plasma immersion ion implantation (PIII) and electron beam evaporation (EBE), and in situ redox reactions (Figure 2). For electrochemical techniques, electrochemical deposition and micro-arc oxidation (MAO) are commonly used to create bioactive coatings (Shen et al., 2022). In particular, electrochemical deposition can proceed via anodization and cathodic reduction. While anodization efficiently generates ordered nano- or microstructures on Ti, incorporating positively charged metal ions into anodic layers remains challenging. In contrast, cathodic reduction facilitates ion incorporation, with deposition efficiency strongly dependent on current density. Specifically, high pulse potentials favour the reduction of target ions into NPs, while insufficient current density leads to their co-deposition with bioceramics. For instance, chitosan-mediated pulse electrochemical deposition successfully produced a HA/Ag composite coating on Ti, in which Ag nanoparticles were uniformly embedded within the HA matrix (Yan et al., 2017). MAO remains a robust strategy to fabricate metal ion-doped coatings with nano/micropores architectures. The morphology and chemical composition of MAO coatings can be modulated by adjusting electrolyte composition, applied voltage, and by introducing chelating agents, such as d-gluconate, acetic acid, or oxalate, to prevent premature metal salt precipitation and enhance ionic migration toward the anode (Sobolev et al., 2019).

Hydrothermal treatment offers an alternative means of introducing metal ions by enabling ion exchange or chemical reactions within a sealed high-temperature autoclave. The extent of metal ion incorporation is largely determined by dopant concentration, reaction temperature, and treatment duration (Xu et al., 2020; Gao et al., 2021). For physical deposition approaches, PIII immobilizes metal ions onto implant surfaces by ionizing metal targets under pulsed high voltage without altering surface morphology, while EBE provides conformal, fine-grained coatings and is often applied after pre-patterning the substrate with desired nano/microstructures (Stolzoff et al., 2017; Yu et al., 2017; Li Q. et al., 2019). With respect to in situ redox reactions, they are commonly employed to load AgNPs onto implant surfaces using methods such as ultraviolet irradiation or dopamine self-polymerization (Xu et al., 2017). Through these approaches, Ag in the solution is reduced in situ to form AgNPs, which are subsequently integrated into the coating on the implant surface.

Strategies for immobilizing AMPs and antibiotics generally fall into three categories: physical adsorption, affinity-based binding, and covalent conjugation (Sun Z. et al., 2023). In this section, AMPs immobilization is primarily focused on. Physical adsorption of AMPs is the simplest approach, yet it often results in inconsistent loading capacity and poorly controlled release kinetics. To address these limitations, metal deposition combined with engineered metal-binding peptide has emerged as an effective strategy to achieve more predictable loading and release behaviour. For instance, an engineered AMP, tet127, fused to a HA-binding peptide, was successfully immobilized on calcium phosphate/HA nanotubular coatings and demonstrated potent antibacterial activity against both Gram-positive (S. mutans) and Gram-negative (E. coli) bacteria (Yazici et al., 2019). Although affinity-based chemical binding offers another relatively straightforward route, it frequently requires complex reaction steps or cytotoxic reagents. Thus, the development of chimeric peptides that integrate both antimicrobial and metal-binding domains has become an attractive alternative. For instance, GL13K modified with a Ti-binding motif through a flexible GSGGG linker effectively inhibited bacterial colonization and biofilm formation, illustrating the versatility of such engineered multifunctional peptides (Wisdom et al., 2020). Nevertheless, adsorption- or affinity-based methods may suffer from low long-term stability in biological environments, potentially limiting their clinical reliability.

In contrast, covalently immobilization of AMPs offers a more stable and controlled means of surface functionalization. For instance, thiol-halogen or thiol-maleimide conjugation of C-terminal Cys-modified hLf1-11 onto silanised implant surface with either (3-chloropropyl)triethoxysilane or (3-aminopropyl)triethoxysilane proved 3 to 6-fold higher antibacterial effect against S. sanguinis and Lactobacillus salivarius than the implant without silanisation (Godoy-Gallardo et al., 2014). It should be noted that, AMPs can also be incorporated into multifunctional, polymer-based coatings. For instance, a tri-functional coating through electrodeposition of PEG film and subsequent functionalization with N-succinimidyl-3-maleimidopropionate and the RGD-Cys-LF1-11 peptide demonstrated excellent to prevent the colony formation of S. sanguinis (Hoyos-Nogués et al., 2018). However, it should be noted that the antibacterial effect of such hybrid coatings is often predominantly attributed to their anti-adhesive properties rather than direct bactericidal action. Additionally, combined usage of AMPs appears to exhibit better antibacterial effect. For instance, the implant with the co-immobilization of GL13K and LamLG3 displayed greater antibiofilm efficiency against S. gordonii than either peptide alone (Fischer et al., 2020), highlighting the potential of synergistic AMPs formulations.

The EPS matrix plays an essential role in biofilm cohesion, structural maturation, and mechanical resilience (Flemming et al., 2023). Consequently, targeting EPS components or disrupting their synthesis offers a powerful strategy not only to inhibit biofilm proliferation but also to achieve controlled weakening or detachment of biofilm biomass. Specifically, EPS consists largely of extracellular DNA, exopolysaccharides, and exoproteins, and enzymatic degradation of these components has demonstrated substantial potential for biofilm control. For instance, selective hydrolysis of branched exopolysaccharides can destabilize granular sludge, a specialized biofilm structure, resulting in its disintegration (Wang et al., 2020). Similarly, DNase-functionalized nanomaterials, such as gold nanoclusters, have been shown to degrade extracellular DNA effectively, disperse up to 80% of biofilm biomass, and eliminate approximately 90% of embedded bacteria (Xie et al., 2020b). With respect to EPS synthesis, the common consensus is that quorum sensing (QS), mediated by QS molecules (e.g., acyl-homoserine lactone, autoinducer-2, bis(3′–5′)-cyclic dimeric guanosine monophosphate), is the crucially involved in EPS production (Liu et al., 2024c). Therefore, blocking QS through quorum quenching, such as introducing quorum quenching microorganisms that are able to produce functional enzymes/inhibitors to degrade/inhibit QS signals (Xu et al., 2024).

However, the clinical relevance of EPS degradation extends beyond bactericidal effects. A crucial and often overlooked consideration is the distinction between uncontrolled, naturally occurring biofilm dispersal and therapeutically induced controlled detachment. Uncontrolled dispersal can release highly virulent aggregates or planktonic cells capable of rapidly colonizing adjacent tissues, thereby exacerbating peri-implant infections. In contrast, controlled EPS modulation aims to locally weaken the biofilm matrix in a predictable, time-specific manner, facilitating the subsequent removal or killing of released cells. Enzymes such as DNase I, dispersin B, alginate lyases, glycoside hydrolases, and proteases offer targeted degradation of key matrix components (Daboor et al., 2021; Kaplan et al., 2024; Li et al., 2025), reducing mechanical integrity while enhancing susceptibility to antimicrobials and immune clearance.

Given the risks associated with uncontrolled shedding, EPS-degrading strategies are most effective when combined with complementary antibacterial modalities. Examples include enzyme-mediated matrix weakening followed by local antibiotic delivery (“detach-then-kill”), EPS degradation on surfaces engineered to be anti-adhesive to prevent re-attachment, or enzymatic pre-treatment followed by mechanical debridement in peri-implantitis therapy (Matthes et al., 2021; Souza et al., 2025). Moreover, although QS inhibition represents another promising route for reducing EPS synthesis and preventing biofilm maturation, its application in peri-implant settings remains limited due to potential off-target effects on host tissues and immune signalling.

One of the most critical unmet needs is long-term antibacterial stability. Many release-based or anti-adhesive coatings exhibit excellent short-term performance but gradually lose their activity due to burst release, depletion of loaded ABAs, lubricant loss, or chemical and enzymatic degradation under physiological conditions. Once the concentration of antibiotics, metal ions, or AMPs falls below inhibitory thresholds, the implant becomes vulnerable to recolonization by peri-implant pathogens. Moreover, residual sub-inhibitory levels of ABAs may promote tolerant or resistant phenotypes, while coatings that leave behind an inert or unprotected Ti surface may ultimately confer little advantage over unmodified implants. Time-dependent degradation and loss of superhydrophobicity, hydration layers (e.g., PEG), or liquid-infused lubricants have been documented for multiple anti-fouling systems, underscoring that initial anti-adhesion does not guarantee long-term clinical reliability in the mechanically and microbially dynamic oral environment. In this context, strategies such as more durable metallic or alloy-based antimicrobial surfaces, hierarchical architectures that couple an initial release-killing phase with underlying contact-killing micro/nanostructures, self-renewable polymeric layers, and stimuli-responsive systems that can be re-activated by infection-related cues represent particularly attractive directions.

These durability issues are tightly coupled to the race for the surface. Even when a coating successfully suppresses bacterial colonization during the early postoperative period, long-term protection will only be achieved if the surface simultaneously supports rapid host cell attachment, robust osseointegration, and stable peri-implant soft-tissue sealing. Antibacterial strategies should therefore be framed within a host–microbe competition paradigm in which early “sterilization” alone is insufficient, and the surface must also be optimized to favour long-lived host coverage once the initial antibacterial effect wanes. This dual requirement calls for coatings that balance antimicrobial potency with excellent cytocompatibility and pro-integration properties.

Standardized and clinically relevant evaluation models are also essential. Current in vitro biofilm assays often rely on mono-species, static systems that poorly recapitulate the polymicrobial, flow-exposed, and immune-modulated environment of peri-implant infections. In addition to assessing early reductions in bacterial load, next-generation models should explicitly consider biofilm maturation, detachment, and re-colonization. This is particularly important for EPS-targeting or dispersal-inducing strategies, where controlled weakening of the matrix is desired but uncontrolled shedding of viable aggregates could seed new infection sites. Hence, in vitro and in vivo models should evaluate not only short-term biomass reduction but also the fate of detached cells whether they are effectively cleared locally, killed by concurrent antibacterial measures, or disseminate to distant surfaces. Multispecies, flow-based systems and longitudinal in vivo studies that integrate mechanical challenges, salivary components, and host immune responses will be crucial for understanding the long-term behaviour of antibacterial coatings over the entire implant lifespan.

Another important challenge is the mitigation of bacterial resistance. Persistent exposure to sublethal concentrations of conventional antibiotics or metal ions risks accelerating resistance development and undermining long-term implant success. Multi-modal strategies that combine physical bactericidal topographies (e.g., high-aspect-ratio nanostructures) with AMPs, metal ions, or QS/biofilm-disrupting agents may reduce selection pressure on any single target and limit bacterial adaptation. In addition, coatings designed to destabilize the EPS matrix or interfere with QS can sensitize biofilms to host immunity and systemic therapeutics, further lowering the likelihood of resistant niche formation. Ultimately, long-term antibacterial performance must be evaluated in conjunction with resistance evolution, mechanical durability, and host response in realistic simulated and in vivo environments.

Bridging the gap between bench-top innovation and clinical implementation requires not only scientific advances but also attention to manufacturability, cost, and regulatory feasibility. Many of the most sophisticated surface modification techniques, such as nanolithography or atomic layer deposition, yield exquisitely controlled nanoarchitectures but remain expensive, time-consuming, and difficult to scale to high-throughput, good manufacturing practice (GMP)-compliant production. By contrast, clinically adopted implant surface treatments, including sandblasting, acid etching, and anodization, owe much of their success to their simplicity, reproducibility, and low cost. Future antibacterial coatings must therefore be designed with scalable, cost-effective processes, using methods such as plasma-enhanced chemical vapor deposition, electrochemical anodization, or self-assembled polymer layers that can be integrated into existing manufacturing pipelines. A clear-eyed comparison between complex multifunctional coatings and simpler, industrially proven processes is necessary to ensure that proposed technologies are not only scientifically compelling but also commercially viable.

Personalized and adaptive coatings represent another promising direction yet pose additional translational challenges. Patient-specific factors such as oral microbiome composition, systemic health, immune competence, and local tissue quality strongly influence peri-implantitis risk. Coatings that can dynamically respond to individual microbial profiles or inflammatory signals, through embedded biosensors and on-demand ABAs release, could, in principle, provide tailored protection with minimal side effects. However, such systems intrinsically increase design complexity, quality control requirements, and production costs, and they must satisfy stringent safety and reliability standards. To make personalization realistic, integration of patient-derived data into coating design will require harmonized clinical datasets, robust risk stratification tools, and pragmatic compromises between customization and standardization.

Successfully navigating regulatory pathways and clinical trials is equally crucial. Very few antibacterial surface technologies have progressed from preclinical promise to regulatory approval, in part because they often function as combination products that must satisfy requirements for both devices and drug/biologic components. Multidisciplinary collaboration among materials scientists, microbiologists, clinicians, and regulatory experts is essential to define acceptable performance endpoints, design appropriate preclinical tests, and plan stepwise clinical evaluation. Early-phase trials may be best targeted to high-risk patient subsets, such as those with a history of recurrent infections or severe peri-implantitis, to demonstrate proof-of-concept and risk–benefit advantages. Real-world evidence from post-marketing surveillance and registries will also be important to support broader adoption once safety and efficacy have been established.

The future of antibacterial implant surfaces will likely be shaped by the convergence of responsive materials, mechanobiology, artificial intelligence (AI), and additive manufacturing. However, for these technologies to move beyond conceptual appeal, their translational pathways and current roadblocks must be clearly articulated.

Smart biointerfaces based on stimuli-responsive materials exemplify both the promise and the complexity of next-generation coatings. Systems that respond to pH shifts, enzymatic activity, redox conditions, or oxidative stimuli can, in principle, deliver ABAs on demand, regenerate antibacterial activity, or remain quiescent in healthy tissue. Yet, as highlighted by the “activation threshold” problem, these coatings inherently depend on sufficiently strong local cues. Early-stage, mild, or spatially heterogeneous infections may not generate the magnitude of pH change, enzyme activity, or oxidative stress needed to trigger release, creating a risk of “false negative” scenarios in which pathogenic bacteria are present but the coating fails to turn on (Mitra et al., 2020; Hu et al., 2023). Future studies must therefore evaluate not only trigger specificity and on–off responsiveness but also under-activation risks under subtle early-infection or chronic conditions. From a translational perspective, smart coatings will need to demonstrate long-term in vivo biosafety and stability, including for both the responsive matrices and their degradation products. They must also achieve reliable activation in clinically relevant microenvironments, scalable GMP-compatible fabrication with reproducible loading and release profiles, and compliance with combination-product regulatory requirements, including sterilization compatibility and lot-to-lot consistency. In most realistic scenarios, smart release should be combined with baseline protection, contact killing, and anti-adhesion mechanisms rather than serving as the sole line of defense.

Mechanobiology-inspired strategies offer a complementary avenue by harnessing biomechanical cues that are intrinsically present in the peri-implant niche, such as fluid shear, occlusal loading, and immune cell–material interactions. Surfaces that modulate bacterial adhesion and EPS organization under physiological shear stresses, or that bias macrophage polarization toward pro-resolving phenotypes, could synergize with chemical approaches to create robust, host-integrated antibacterial defenses. However, translating these concepts will require quantitative models linking surface mechanics, biofilm architecture, and immune responses, as well as standardized testing conditions that capture relevant mechanical regimes.

AI–driven design has the potential to accelerate the discovery of multifunctional coatings by predicting optimal combinations of surface chemistry, topography, and mechanical properties for defined microbial and host contexts. At present, the main bottleneck is not algorithmic sophistication but the lack of large, standardized, and well-annotated datasets integrating bacterial behavior, coating performance, host responses, and clinical outcomes (Butler et al., 2018; Najeeb and Islam, 2025). Existing biomaterials and implant datasets are often small, fragmented, generated under heterogeneous experimental conditions, and lack clinically relevant endpoints such as long-term biocompatibility, degradation behavior, and in vivo success rates. Similar limitations are seen in dental and implant AI applications, which frequently rely on single-center, retrospective data with limited external validation. For AI-guided coating design to become truly impactful, curated multi-center databases, standardized reporting of composition and test protocols, and interpretable models that can be prospectively validated in preclinical pipelines will be required. In the near term, AI is best viewed as a tool that operates in a closed design–test–refine loop, augmenting but not replacing rigorous experimental and clinical research.

Additive manufacturing for antibacterial, patient-specific implants illustrates both the power and constraints of advanced fabrication. Metal-based techniques such as laser powder bed fusion and electron beam melting operate at very high temperatures that are fundamentally incompatible with heat-sensitive ABAs, including most antibiotics, AMPs, enzymes, and polymeric bioactive coatings (Dzogbewu and du Preez, 2021; Long et al., 2023; Alam et al., 2025). As a result, antibacterial functionality cannot usually be co-printed with the metallic framework. Instead, it must be introduced through post-printing surface modifications such as sol–gel deposition, hydrothermal treatment, or electrophoretic coating. Moreover, the complex, often highly porous geometries produced by 3D printing pose additional challenges for uniform coating coverage, strong adhesion, and long-term stability under cyclic loading and clinical sterilization procedures. These technical issues are compounded by high production costs, stringent quality assurance requirements, powder handling safety, and regulatory demands for batch consistency in patient-specific devices (Dzogbewu and du Preez, 2021; Long et al., 2023; Alam et al., 2025). When 3D-printed antibacterial implants are compared with simpler, well-established treatments such as SLA or anodization, which are low-cost, robust, and easily standardized, it becomes clear that their commercial viability will depend on carefully balancing functional sophistication with manufacturing pragmatism.

Finally, a realistic future outlook must integrate scientific innovation with economic and regulatory feasibility. Coatings that rely on multi-step wet chemistry, high-purity biological reagents, controlled environments, or cold-chain logistics will inevitably be more expensive and challenging to scale than single-step, low-temperature, and purely inorganic treatments. By explicitly acknowledging these constraints and outlining the technical and regulatory milestones required (e.g., long-term durability, resistance risk management, activation reliability, manufacturability, and cost-effectiveness), future research can chart a more credible translational pathway. In this way, antibacterial implant surfaces can gradually evolve from promising experimental concepts into robust, clinically approved technologies that sustainably prevent peri-implant infections across the full lifetime of the implant.