Ziqian Zhou
Chunning Gu
Anyuan Shi
Junyan Jing
Wei Cheng- Nanjing Stomatological Hospital, Affiliated Hospital of Medical School, Institute of Stomatology, Nanjing University, Nanjing, China
Next-generation antibacterial implant surfaces are rapidly evolving toward intelligent, adaptive, and patient-specific designs powered by emerging technologies such as smart biointerfaces, artificial intelligence–guided material optimization, and additive manufacturing. These advances promise to fundamentally reshape strategies for preventing peri-implant infections. However, their clinical translation remains constrained by critical challenges including activation thresholds of stimuli-responsive coatings, durability of anti-adhesion layers, long-term stability after release depletion, and the persistent “race for the surface” between bacteria and host tissue. Peri-implantitis, driven predominantly by bacterial adhesion and biofilm maturation on implant surfaces, continues to compromise the longevity of dental and orthopedic implants, and conventional mechanical or antibiotic-based therapies often fail to fully eradicate resilient biofilms. Recent progress in antibacterial implant surface engineering is summarized in this review, covering two complementary strategies: (i) preventing initial bacterial adhesion through passive micro/nanostructuring, superhydrophobic or superhydrophilic surfaces, and active release-based coatings; and (ii) inhibiting the proliferation and persistence of attached bacteria via contact-killing mechanisms and controlled dismantling of the extracellular polymeric substances matrix. Mechanisms and immobilization strategies of organic (e.g., antimicrobial peptides, antibiotics) and inorganic antibacterial agents (e.g., metal ions, nanoparticles) are further compared, highlighting their advantages and limitations. Finally, the translational pathway for future antibacterial implants is outlined. By bridging mechanistic insights with emerging technologies, next-generation implant surfaces may achieve durable antibacterial function, enhanced osseointegration, and improved long-term outcomes for patients at risk of peri-implantitis.
1 Introduction
The global demand for orthopedic and dental implants continues to rise as age and bone-related diseases increase (Guglielmotti et al., 2019). Titanium (Ti) and its alloys remain the clinical standard due to their excellent mechanical strength, corrosion resistance, and biocompatibility (Nikoomanzari et al., 2022; Jiang et al., 2023; Wen et al., 2023; Liu et al., 2024b). However, this biocompatibility is not selective: the same surface properties that support osteoblast adhesion also permit the rapid colonization of planktonic bacteria, also known as “race for the surface,” predisposing patients to peri-implant infections such as peri-implantitis, whose incidence ranges from 20% to 47% (Dixon and London, 2019). These infections are a major cause of implant failure (Choi et al., 2021; Wu et al., 2021).
Beyond prevalence, peri-implantitis imposes a significant clinical burden. Patients may experience pain, suppuration, bleeding, impaired mastication, and considerable anxiety of implant (Dreyer et al., 2018). Advanced cases often require complex treatment procedures, including flap surgery, surface decontamination, regenerative interventions, or even implant removal, whose outcomes are variable and whose costs are substantial (Saleh et al., 2024). These challenges are further amplified in patients with systemic risk factors such as diabetes, smoking, or previous periodontitis (Darby, 2022). Current clinical treatments, including mechanical or laser debridement combined with antibiotics, face limitations: mechanical debridement often lacks efficacy, and repeated antibiotic use increases the risk of bacterial resistance (Hizal et al., 2015; Yin et al., 2021). Therefore, preventing peri-implantitis remains challenging, highlighting an urgent need for novel and effective strategies. Collectively, the morbidity, treatment complexity, and economic impact underscore the urgent need for improved strategies for the prevention and management of peri-implantitis.
Preventing bacterial biofilm formation is essential to mitigating peri-implantitis (Daubert and Weinstein, 2019; Lee et al., 2022). Biofilm development on implant surfaces, a key driver of peri-implantitis, occurs through a multi-stage process (Lencova et al., 2021; Sauer et al., 2022): (i) initial, reversible bacterial adhesion, mediated by non-covalent interactions under dynamic fluid flow, (ii) proliferation and matrix production, during which bacteria secrete extracellular polymeric substances (EPS) to form a mechanically robust, often irreversible biofilm, and (iii) biofilm maturation, detachment, and re-colonization, fragments or planktonic cells disperse and seed secondary infection foci (Figure 1). Mature biofilms possess high mechanical stability due to their EPS-rich matrix, which hinders antibiotic penetration and reduces the efficacy of conventional debridement therapies. These effects ultimately challenge current clinical treatments. These features underline the clinical importance of surface engineering approaches that can interrupt biofilm formation at its earliest stages and weaken the structural integrity of established biofilms.
Figure 1. Bacterial biofilm formation leading to peri-implantitis and associated antibacterial strategies. EPS: extracellular polymeric substances.
This review evaluates recent advances in antibacterial implant surface engineering aimed at preventing peri-implantitis. Two primary antibiofilm strategies are highlighted, aligned with the biofilm development stages: (i) preventing initial bacterial adhesion and (ii) inhibiting the proliferation of adhered bacteria (Figure 1). In particular, antibacterial agents (ABAs) play a central role in these strategies and are categorized based on their application in surface modifications: release-killing mechanisms to prevent initial adhesion and contact-killing mechanisms to suppress bacterial proliferation. Common classes of ABAs and their bactericidal mechanisms are firstly summarized, followed by coating techniques for their integration onto implant surfaces. Additional strategies, including modifications to surface physical and chemical properties to prevent adhesion and EPS-targeting methods that weaken biofilm integrity, are also discussed. Finally, scientific and clinical perspectives, including long-term stability, biofilm detachment, adaptive resistance, and translational considerations, are provided. This review aims to support both researchers and clinicians in advancing preventive solutions for peri-implant infections.
2 Overview of antibacterial agents and their bactericidal mechanisms
ABAs are broadly categorized into organic agents (e.g., antibiotics, antimicrobial peptides (AMPs)) and inorganic agents (e.g., metal ions and their nanoparticle (NP) forms) categories (Figure 2). When integrated into implant surfaces, these agents operate via two principal antibacterial modes: contact-killing, in which immobilized ABAs eliminate adherent bacteria, and release-killing, in which ABAs are incorporated into degradable or porous matrices for localized delivery against planktonic pathogens. This section outlines the underlying bactericidal mechanisms of major ABAs. Their associated coating strategies for release- and contact-killing are detailed in Sections 3.2, 4, respectively.
Figure 2. Bacterial biofilm formation leading to peri-implantitis and associated antibacterial strategies. EPS: extracellular polymeric substances.
2.1 Organic antibacterial agents and their bactericidal mechanisms
2.1.1 Antibiotics
Antibiotics, including gentamicin, cephalothin, carbenicillin, amoxicillin, cefamandol, tobramycin, and vancomycin, primarily target essential cellular processes, particularly genetic information processing (Figure 2). For instance, rifampin and related rifamycins inhibit the DNA-dependent RNA polymerase encoded by rpoB (Surette et al., 2022). Aminoglycosides, driven by the number and basicity of their amino groups, bind to ribosomes and induce mistranslation, resulting in toxic misfolded proteins (Aguirre Rivera et al., 2021). Some antibiotics, such as the glycopeptide complestatin, directly bind to peptidoglycan stem units and inhibit cell wall synthesis by blocking remodelling enzymes. For instance, Culp et al. (2020) demonstrated that complestatin inhibits peptidoglycan hydrolases essential for cell wall turnover. Despite these mechanistic insights, the precise bactericidal sequence, whether disruption of protective barriers (cell envelopes) or damage to bacterial genetic identity (DNA), remains incompletely understood. Clarifying these processes is crucial for optimizing current drugs and designing next-generation antibiotics.
Despite their efficacy, bacteria associated with peri-implantitis commonly exhibit not only classical genetic antibiotic resistance but also biofilm-related tolerance and adaptive resistance. These include reduced metabolic activity, efflux pump upregulation, and EPS-mediated diffusion barriers (Belay et al., 2024). These adaptive responses make local delivery coatings less effective over time and underscore the need for alternative or combination strategies. Therefore, mitigating antibiotic resistance requires multifaceted strategies. Combination therapy is one effective approach. For instance, a lipid-based coating co-loaded with amikacin and vancomycin achieved 5-log and 3-log reductions in Staphylococcus aureus (S. aureus) and Pseudomonas aeruginosa (P. aeruginosa), respectively, after 24 h of exposure (Jennings et al., 2015). Moreover, newly developed antibiotics with dual or unique targets also show promise. For example, SCH-79797 disrupts both folate metabolism and membrane integrity and exhibits low resistance frequencies against diverse pathogens (Martin et al., 2020).
2.1.2 Antimicrobial peptides
AMPs, short peptide (<50 amino acids) (Lai et al., 2022), such as GL13K, hLF-11, LLL-37, and Tet213 (Baquero and Levin, 2021), primarily kill bacteria through electrostatic interactions with negatively charged bacterial membranes (Figure 2). They cross the cell envelope along an electrostatic gradient, after which their hydrophobic domains insert into the lipid bilayer, enabling pore formation, membrane disruption, cytoplasmic leakage, and eventual cell death (Brogden, 2005). These mechanisms indicate that AMPs are less likely than traditional antibiotics to induce stable genetic resistance, making them promising candidates for combating healthcare-associated (nosocomial) infections, including those related to implants.
Natural AMPs, derived from amphibians, insects, mammals, and fish (75.65%), plants (13.5%), and bacteria (8.53%) (Hazam et al., 2019), face major clinical challenges, particularly susceptibility to proteolytic degradation (e.g., trypsin, chymotrypsin, serum proteases) (Torres et al., 2019; Mookherjee et al., 2020). Synthetic AMPs alleviate these issues through strategies like sequence modification, cyclization, peptidomimetics, dimerization, and nanoscale processing (e.g., truncation, amino acid substitution, hybridization) (Erdem Büyükkiraz and Kesmen, 2022; Lai et al., 2022). For instance, substituting natural amino acids with unnatural derivatives, such as D-amino acids, halogenated amino acids, or β-amino acids, enhances resistance to proteolysis (Lai et al., 2022).
However, recent studies show that bacteria can also mount adaptive resistance to AMPs through transient, non-genetic mechanisms, such as membrane charge modification, lipid A remodeling, increased capsule formation, and activation of efflux pumps (Kraus and Peschel, 2006; Andersson et al., 2016). Although these mechanisms do not usually lead to permanent resistance, they may reduce AMPs efficacy in chronic or established biofilms. Thus, AMPs-based coatings must consider both bactericidal potency and potential adaptive responses within the peri-implant microenvironment. Despite these challenges, AMPs remain highly promising due to their broad-spectrum activity and minimal cross-resistance with conventional antibiotics.
2.2 Inorganic antibacterial agents and their bactericidal mechanisms
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.
2.3 Other antibacterial agents
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.
3 Preventing initial bacterial attachment onto implant surfaces
3.1 Passive anti-adhesion strategies through implant surface modification
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.
3.1.1 Surface physical property modification with micro–nanostructuring techniques
Surface physical properties critically shape bacterial adhesion by modulating surface roughness and wettability, which in turn influence van der Waals forces, electrostatic and hydrophobic interactions (Linklater et al., 2021). Accordingly, micro-nanostructuring techniques have been widely applied to engineer surface features ranging from micrometer to nanometer scale, thereby altering implant topography to deter bacterial colonization (Jia et al., 2017; Linklater et al., 2017). A series of surface patterns have been designed and can generally be categorized into protrusive structures such as nanorods, nanowires, nanoripples, microgrooves, and nanopillars, and into recessed architectures including nanotubes and microwells, many of which are inspired by natural bactericidal surfaces (Figure 3). For instance, Ivanova et al. (2012) reported that cicada wings possess nanopillars with diameters of 60–100 nm, heights of approximately 200 nm, and inter-pillar spacing of 170 nm, which mechanically disrupt Gram-negative bacterial cells. Similarly, dragonfly wing nanospikes measuring 50–70 nm in diameter and about 240 nm in height were later shown to inactivate both Gram-negative and Gram-positive bacteria (Ivanova et al., 2013).
Figure 3. Implant surface modifications to prevent bacterial adhesion.
Micro–nanostructuring involves three steps (Yang et al., 2022; Rahvar et al., 2025): (1) Surface cleaning (e.g., ultrasonic treatment, organic solvents like acetone and ethanol, and distilled water), (2) preparation of a micro/nano-polished surface through machining or acid etching (e.g., electronic polishing with acetic acid, sulfuric acid, and hydrofluoric acid, or chemical polishing with HF and nitric acid), and (3) fabrication of the designed micro/nanopatterns. It should be noted that diverse fabrication methods have been developed to achieve specific micro/nano patterns. Hydrothermal treatment, anodization, and reactive ion etching (RIE) are typically used to generate nanopillars, nanospikes, and nanowires (Jin et al., 2016; Wandiyanto et al., 2020; Xue et al., 2021). Microwells can be produced through mask lithography (Yang et al., 2015), RIE (Munther et al., 2018), electron beam lithography (Alalwan et al., 2018), and laser ablation (Parmar et al., 2018). Nanoripple structures are often formed using femtosecond laser–based interference and ablation (Dar et al., 2016; Liang et al., 2016), and microgroove structures can be produced by machining, grinding, laser ablation or RIE.
The anti-adhesion effectiveness of micro/nano-patterned surfaces depends strongly on bacterial species and the geometric parameters of the structures, including size, width, shape, porosity, and height. For instance, Streptococcus sanguinis (S. sanguinis) can adhere to metallic surfaces with roughness value as low as 0.5 nm, whereas P. aeruginosa shows limited attachment on nanosmooth surfaces with roughness below 1 nm. (Truong et al., 2015). This species-dependent adherence difference may reflect differences in size, as bacteria preferentially colonize recesses features larger than their own dimensions. Given that S. sanguinis measures 0.5–1.0 μm (Whitman et al., 2015), and P. aeruginosa approximately 0.3–0.8 μm (Holt and Ebersole, 2005), nanoscale roughness, being orders of magnitude smaller, may provide insufficient attachment sites for P. aeruginosa while still permitting S. sanguinis adhesion.
Recently, it is also intriguing to find that surfaces with high-aspect-ratio patterns (e.g., nanopillars, nanospikes, nanowires) can exert physico-mechanical forces that directly comprise bacterial cell integrity, thereby preventing colonization. Three principal mechanisms have been proposed. The first involves excessive stretching of the bacterial cell wall when sharp, rigid nanostructures cause deformation beyond its elastic threshold, leading to membrane rupture (Pogodin et al., 2013). The second mechanism relates to the storage and release of elastic energy within flexible nanostructures, which bend upon contact and subsequently recover, generating forces capable of damaging bacterial membranes (Linklater et al., 2018). Notably, the magnitude of this stored energy can be estimated using Euler beam theory. The third mechanism is the “nanoknife” effect, where sharp nano-edges puncture bacterial envelopes, disturb osmotic balance, and induce cell lysis (Pham et al., 2015). Notably, susceptibility to these physico-mechanical actions varies among bacterial species and is influenced largely by differences in peptidoglycan thickness between Gram-positive and Gram-negative bacteria (Arnoldi et al., 2000).
3.1.2 Surface chemical property modification with functional coatings
3.1.2.1 Superhydrophobic/superhydrophilic coatings
A superhydrophobic surface, defined by a water contact angle exceeding 150° and a sliding angle below 10° (Rasitha et al., 2024), functions by trapping air within its hierarchical micro/nanostructures to form a stable plastron layer that inhibits bacterial attachment at the gas–liquid interface (Figure 3). In most cases, the fabrication of superhydrophobic surfaces involves first increasing the surface roughness and then modifying the surface chemistry to reduce surface free energy. The specific techniques used to generate surface roughness vary widely (Souza et al., 2020). Sandblasting, acid etching, laser radiation, and anodization are commonly applied for this purpose (Sun X. et al., 2023). During sandblasting, microscale particles such as TiO2, Al2O3, SiO2, and hydroxyapatite (HA) are propelled onto implant surfaces using compressed gas to generate irregular topographies (Accioni et al., 2022). Acid etching, often combined with sandblasting in the widely used SLA process, produces micropits through reactions with mixture of HCl, H2SO4, HNO3, and HF (Mandracci et al., 2016), and has become a predominant commercial method because of its simplicity and effectiveness. Laser-based modification further allows precise tailoring of surface microstructures, as the morphology of the resulting pits or protrusions depends on laser type, energy, and scanning direction (Al-Zubaidi et al., 2020).
To reduce surface free energy, coatings composed of nonpolar materials such as hydrocarbons or fluorocarbons are typically employed. In particular, fluorinated chains exhibit low cohesive energy and unique physicochemical features, including low dielectric constants, high vapor pressures, enhanced compressibility, and high surface activity in aqueous media (Kovalchuk et al., 2014). For instance, a composite fluorinated coating consisting of fluorinated polydimethylsiloxane, fluorinated silicon dioxide, and a perfluoroalkyl polymer achieved contact angles above 150° for both water and hexadecane through a sequential spray-coating process (Huang et al., 2022). Notably, various deposition, such as layer-by-layer assembly, chemical vapor deposition, electrochemical deposition, and sol-gel processing have been reported for applying low-surface-energy coatings onto implant surfaces (Rasitha et al., 2024).
It should be noted that one-step methods have also been developed to fabricate superhydrophobic surfaces. For instance, glow discharge plasma treatment with argon, oxygen, and hexamethyldisiloxane gases can simultaneously etch Ti and form a superhydrophobic layer. In vitro assays using saliva inocula demonstrated an eight-fold reduction in microbial adhesion after 2 hours, highlighting the potential of these surfaces for controlling early biofilm formation (Souza et al., 2020). Despite these advantages, the long-term stability of superhydrophobic coatings remains a major challenge, as prolonged exposure may diminish antiadhesive properties or even promote bacterial colonization. For instance, Hwang et al. (2018) found that the superhydrophobic surface might promote bacterial adhesion after long-term exposure. Consequently, strategies for maintaining long-term performance warrant further investigation. As an alternative, superhydrophilic coatings have emerged that rely on a contrasting mechanism (Makowiecki et al., 2019): the formation of a strongly hydrated surface layer that physically repels bacteria and creates an unfavourable microenvironment for microbial growth (Figure 3). However, relatively few studies have examined their fabrication and long-term stability.
3.1.2.2 Antibiofouling polymers
Hydrophilic polymers have been extensively explored as antibiofouling materials because their large molecular structures, composed of repeating subunits, generate highly hydrated interfaces that reduce bacterial adhesion (Figure 3). Their high polymer density and tightly bound water layers create an energetically unfavourable surface for bacterial approach, functioning simultaneously as a physical barrier and a free-energy shield. Among these materials, polymer brushes, densely packed polymer chains end-grafted to a substrate and extending into the surrounding medium, provide a tunable platform for controlling interfacial energy and surface interactions. (Kobayashi et al., 2012). Currently, polyethylene glycol (PEG) has long been regarded as the gold standard for antifouling coatings (Maan et al., 2020), and its success has inspired a broad range of alternative strategies, including lubricin, polyoxazolines, polyglycerol dendrons, polyvinylpyrrolidone, polysaccharides, polypeptoids, polyacrylamide, and zwitterionic polymers, all of which build upon the principles established by PEG-based systems (Knowles et al., 2017; Li et al., 2019c). Therefore, its application as fouling-resistant coatings is particularly discussed herein.
PEG is a water-soluble, non-toxic, highly flexible, and biocompatible polymer with low immunogenicity and a very large exclusion volume, making it highly effective at preventing non-specific adsorption. PEGylation, the grafting of PEG chains onto surfaces, remains the most widely used method for creating linear PEG brushes capable of resisting bacterial and protein attachment. The superior repelling potential of PEG derivatives was first demonstrated by Prime and Whitesides (Prime and Whitesides, 1991), who showed that extending ethylene oxide (EO) chain length markedly enhanced resistance to adhesion, as longer PEG chains form more complete and sterically demanding hydration layers. Subsequent studies using a variety of PEG architectures and grafting methods have consistently confirmed strong antifouling performance across numerous substrates, including PEG-tethered gold surfaces, glass modified with PEG-containing copolymers, and PEGylated polyaniline nanofibers.
Despite its advantages, PEG presents several limitations that restrict its clinical translation. Firstly, cleavage of EO units yields aldehyde-terminated chains that readily react with amine-functionalized proteins, making PEG susceptible to oxidative degradation and enzymatic cleavage in most biochemically relevant environments, therefore limiting its long-term stability. Secondly, immobilizing PEG on different substrate chemistries often requires complex and highly specific surface reactions, which increase fabrication difficulty and cost, particularly for large-scale or industrial applications (Yi et al., 2019). Thirdly, the strong hydration of PEG chains also leads to significant swelling in aqueous environments, reducing mechanical stability and making coatings more prone to deformation or detachment (Xie et al., 2019). Additionally, temperature sensitivity presents an additional challenge, as elevated temperatures above approximately 35 °C can cause collapse of the hydration layer, thus increasing hydrophobicity and diminishing antifouling performance (Rahvar et al., 2025).
3.1.2.3 Others
Beyond the aforementioned functional coatings, several other anti-adhesion strategies have been developed. For instance, surface charge modification represents an effective approach, as anionic polyelectrolyte multilayers assembled through layer-by-layer deposition have demonstrated strong resistance to adhesion by P. aeruginosa, Escherichia coli (E. coli), and S. sanguinis through electrostatic repulsion (Zhu et al., 2015). Moreover, liquid-infused surfaces (LISs), characterized by an anchored lubricant layer that creates a smooth, slippery interface, offer another route to antibiofouling functionality. LISs maintain stable liquid-repellent behaviour even at very low contact angles and may be fabricated either by applying lubricant-compatible chemistries to pre-roughened surfaces or by roughening substrates that inherently exhibit strong affinity for the lubricant (Villegas et al., 2019). However, LISs face challenges related to cytocompatibility and long-term stability that must be addressed before clinical implementation can be achieved.
3.2 Active anti-adhesion strategies through release killing of antibacterial agents-loaded coatings
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.
Table 1. Summary of major antibacterial implant surface modification strategies.
4 Contact killing and controlled EPS matrix modulation to suppress the proliferation of attached bacteria
4.1 Immobilization of metal ions and their derivatives
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.
4.2 Immobilization of antimicrobial peptides and antibiotics
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.
4.3 Controlled enzymatic modulation of the EPS matrix to destabilize biofilms
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.
5 Future perspectives: bridging fundamental research to clinical applications
Building upon the comparative framework summarized in Table 1, the following subsections highlight the key translational challenges and emerging opportunities that span across these diverse antimicrobial strategies.
Despite substantial advances in antibacterial implant surface engineering, peri-implantitis remains a major cause of implant failure and patient morbidity in both dental and orthopedic settings. The classical “race for the surface” paradigm highlights that, immediately after implantation, host cells and bacteria compete for the same implant interface. Therefore, long-term success depends not only on the early suppression of bacterial adhesion but also on the rapid establishment of a stable, host-integrated soft- and hard-tissue seal. Short-lived antibacterial effects that merely delay bacterial colonization without supporting durable tissue integration are unlikely to provide lasting clinical benefit. To translate promising laboratory concepts into clinically viable implant surfaces, future work must address key challenges related to long-term durability, biosafety, resistance development, and manufacturability, and must consider realistic translational pathways rather than purely visionary scenarios.
5.1 Addressing current challenges in antibacterial surface modifications
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.
5.2 Strategies to accelerate clinical translation
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.
5.3 Emerging trends and future technologies
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.
6 Conclusion
The development of antibacterial implant surfaces is entering a pivotal phase in which mechanistic understanding, materials innovation, and translational science must converge to overcome the persistent challenge of peri-implant infections. Across the strategies reviewed, ranging from passive anti-adhesion designs and controlled-release systems to contact killing and EPS-matrix modulation, one message is clear: durable antibacterial stability and optimal biocompatibility can only be achieved through integrated, multi-layered surface engineering rather than reliance on any single mechanism. The evidence synthesized in this review suggests a coherent pathway forward.
First, long-term stability must be built on hierarchical or staged architectures that couple early-phase release killing with sustained contact-killing structures or anti-adhesive layers, thereby addressing the inevitable depletion of loaded agents and reducing the risk of recolonization. Second, successful implants must balance antibacterial performance with host integration by incorporating bioinspired chemistries, osteoconductive nanostructures, or matrix-mimicking interfaces that actively promote rapid soft-tissue and bone attachment, an essential countermeasure in the “race for the surface.” Third, coating reliability and reproducibility require standardized and clinically relevant testing models, including multispecies biofilms, dynamic flow systems, and long-term mechanical and chemical stability assessments that mirror practical oral or orthopedic environments. Fourth, antibacterial strategies should mitigate resistance by combining orthogonal mechanisms, including physical, chemical, and biological approaches, to minimize selective pressures that enable adaptive bacterial phenotypes.
Finally, emerging tools such as stimuli-responsive “smart” biointerfaces, mechanobiology-driven surface designs, AI-guided optimization, and additive manufacturing offer unprecedented opportunities to create implant surfaces that are adaptive, patient-specific, and clinically durable. Realizing this vision will require GMP-compatible fabrication routes, validated activation thresholds for intelligent coatings, robust datasets for AI training, and regulatory frameworks that align with multifunctional implant technologies. By uniting mechanistic insights with these advancing technologies, the next-generation of implant surfaces can achieve sustained antibacterial function, seamless integration with host tissues, and long-term clinical reliability, ultimately improving patient outcomes and establishing a new standard of care in implantology.
Author contributions
ZZ: Writing – review and editing, Conceptualization, Writing – original draft, Visualization. CG: Writing – review and editing, Visualization. LG: Investigation, Writing – original draft. AS: Writing – original draft. JJ: Writing – original draft. WC: Writing – review and editing, Conceptualization, Funding acquisition, Project administration.
Funding
The author(s) declared that financial support was received for this work and/or its publication. The authors would like to acknowledge the support of the Nanjing Foundation for Development of Science and Technology (202305028).
Conflict of interest
The author(s) declared that this work was conducted in the absence of any commercial or financial relationships that could be construed as a potential conflict of interest.
Generative AI statement
The author(s) declared that generative AI was not used in the creation of this manuscript.
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Keywords: antibacterial agents, biofilm, contact killing, peri-implantitis, release killing, surface modification
Citation: Zhou Z, Gu C, Guo L, Shi A, Jing J and Cheng W (2026) Preventing peri-implantitis: strategies, mechanisms, and clinical perspectives for inhibiting biofilm formation on implant surface. Front. Mater. 12:1739290. doi: 10.3389/fmats.2025.1739290
Received: 04 November 2025; Accepted: 09 December 2025;
Published: 07 January 2026.
Edited by:
Muhammad Zohaib Nawaz, Jiangsu University, ChinaReviewed by:
Babbiker Mohammed Taher Gorish, Omdurman Islamic University, SudanMahammed Ilyas Khazi, Bursa Uludag Universitesi, Türkiye
Copyright © 2026 Zhou, Gu, Guo, Shi, Jing and Cheng. This is an open-access article distributed under the terms of the Creative Commons Attribution License (CC BY). The use, distribution or reproduction in other forums is permitted, provided the original author(s) and the copyright owner(s) are credited and that the original publication in this journal is cited, in accordance with accepted academic practice. No use, distribution or reproduction is permitted which does not comply with these terms.
*Correspondence: Wei Cheng, ZGVudGlzdF9uakAxNjMuY29t