Biofunctional Surface Modification of Polymeric Implants for Biomedical Applications

ABSTRACT

The engineering of the polymer biological interface remains a challenge as it governs responses, such as protein adsorption, cell adhesion, and bacterial colonization, thereby maintaining long‐term biocompatibility. Although polymers are widely used in biomedical implants, the challenges of pristine surfaces in directly interacting with complex physiological environments remain a major challenge. Surface modification is therefore essential for tailoring polymer interfaces to achieve optimal biological and functional performance. This review provides a comprehensive overview of polymeric materials employed in biomedical implants and the surface modification strategies used to enhance their performance. Both biodegradable and non‐biodegradable polymers are discussed, with emphasis on approaches that enhance surface hydrophilicity, mechanical integrity, and interfacial bioactivity while imparting antibacterial, antifouling, and antithrombogenic properties.

Engineering the polymer–biological interface remains critical for controlling biofilm formation, protein adsorption, and cell adhesion on biomedical implants. This review highlights surface modification strategies for polymeric implant surfaces that improve hydrophilicity, mechanical stability, and interfacial bioactivity while imparting antibacterial, antifouling, and antithrombogenic functionalities for advanced implant applications.

1. Introduction

The emergence of polymers has revolutionized the biomedical industry by enabling the design of polymeric biomaterials with precisely controlled physicochemical properties. The inherent biocompatibility and adaptability facilitate their widespread integration in various biomedical applications, such as implants [1, 2], diagnostics [3], tissue engineering [4, 5, 6], controlled drug release systems, and regenerative medicine [4]. Over 50% of all biomedical device materials are polymer‐based, owing to polymers' adaptability, bioactivity, degradability, and economic feasibility. The projected reach of the global polymer market is about $38.3 billion by the year 2030. The compound annual growth (CAGR) of polymers in the biomedical industry is reported to be about 8% by 2030 [7]. The convergence of market growth and healthcare demand has driven the engineering of polymers, leading to significant innovations, such as stimuli‐responsive drug delivery systems [8], bioresorbable grafts [9], 3D‐printed scaffolds for tissue regeneration [10], and their widespread use in biomedical implants. In addition to these advantages, polymers and polymeric implants offer a wide range of chemical, mechanical, thermal, and electrical properties that can be tailored to specific applications. The chemical structure and their ability to degrade by natural biological processes differentiate them into two categories: biodegradable and non‐biodegradable polymers. Non‐biodegradable polymers are long‐chain, heterogeneous synthetic molecules built with strong C─C and C─H bonds. Common examples of such polymers include polyethylene, polypropylene, and polystyrene. These polymers are derived from petroleum through addition polymerization. On the contrary, biodegradable polymers comprise functional groups, such as esters, amides, or ethers. These molecules are readily hydrolyzed by microbial enzymes, yielding non‐hazardous products, such as CO₂, water, and biomass [11]. Polylactic glycolic acid (PLGA), polycaprolactone (PCL), and polyglycolic acid (PLA) are biodegradable polymers. Figure 1 shows examples of biodegradable and non‐biodegradable polymers used in biomedical applications. One of the key reasons for the rapid expansion of polymers is their ability to mimic biological tissues and control over shape, porosity, and topography. These processing advantages enable the creation of complex patient‐specific devices that gradually degrade in vivo, eliminating the need for secondary surgeries [12, 13, 14].

FIGURE 1.

Classification of different polymeric implant biomaterials based on their properties.

Despite polymers' outstanding properties, their interactions with the biological environment pose significant challenges, particularly regarding biofilm formation, thereby limiting their application. Biofilm on polymeric implants results in reduced device performance, infections, device failures, and high patient morbidity [15]. Hence, surface modification of polymeric implant interfaces is essential. Various strategies such as polymer grafting, surface assembled monolayers (SAM), Layer by Layer deposition (LbL), silanization, as well as simple techniques, such as dip‐coating and spin coating, depending on the application and structure can be used to modify the surface interface. Thus, for implants, the interaction between the material and the host tissue is the primary factor in the medical device's success. These modified surfaces can trigger specific biological responses, promote cell adhesion and proliferation, possess antimicrobial properties, modulate immune system components, control thrombogenicity, and exhibit antifouling activity.

In this review, we have discussed advances in surface modification techniques that have emerged as unique tools for controlling biological responses on polymeric implants. The review examines polymer‐host interactions, challenges associated with microorganism attachment, and biofilm formation. The emphasis remains on surface‐based antimicrobial strategies used to inhibit adhesion, disrupt biofilm development, and eliminate microbial cells. By integrating chemical, physical, and nanotechnology‐based approaches within a single framework, this review highlights the common principles that underlie successful surface engineering of polymeric implants. The work additionally highlights the link between surface physicochemical properties and biological performance, including cell behavior, bacterial interactions, and host response. In doing so, the review provides a coherent perspective that bridges materials science and biology, offering practical insight into the rational design of next‐generation polymeric implants and biomedical devices to enhance their functionality and safety.

2. Polymers as Biomedical Implants

The importance of polymeric implants in biomedical applications is established owing to their properties. In critical care settings, polymers such as polydimethylsiloxane, commonly known as silicone (PDMS) [16, 17], polyurethanes (PU) [18, 19], and polyvinyl chloride (PVC) [17, 20] are universally employed in devices ranging from vascular and endotracheal catheters and urinary catheters to drainage pumps and pipes. These polymeric implants are the cornerstone of fluid administration, monitoring, and life support in critical care patients, underscoring their importance. Beyond critical care, polymeric implants are also extensively used in cosmetic and reconstructive applications, such as facial implants, breast implants, and rhinoplasty frameworks, to name a few. The overall impact of polymer‐based implants across medical disciplines is therefore substantial and commendable. However, the biological polymer interface of these implants makes them vulnerable to biofilm formation, often necessitating antibiotic therapy.

Each year, healthcare‐associated infections (HAIs) are responsible for approximately 40, 000 deaths. For instance, in catheter‐associated urinary tract infections (CAUTIs), the risk remains relatively low (3%–5%) for short‐term catheterization, but escalates significantly to 3%–10% in many cases [21]. These infections account for more than a quarter of the health issues acquired during medical care in developing nations and as much as 15% in affluent countries [22]. The Lancet Global Health [23] indicated that the prevalence of healthcare‐associated infections (HAIs) is 20 times higher in low‐ and middle‐income countries (LMICs), including India. Furthermore, statistics from the International Nosocomial Infection Control Consortium (INICC) indicate that the incidence of catheter‐associated urinary tract infections (CAUTI) and central line‐associated bloodstream infections (CLABSI) in intensive care units (ICUs) across 45 low‐ and middle‐income countries (LMICs) is significantly elevated, approximately three to six times per 1000 device‐days, compared to the rates documented in ICUs in the United States. Pathogens such as Pseudomonas aeruginosa (P. aeruginosa), Klebsiella pneumoniae (K. pneumoniae)‐, Acinetobacter baumannii (A. baumannii), Escherichia coli (E. coli), Staphylococcus aureus (S. aureus), Staphylococcus epidermidis (S. epidermidis), Enterococcus faecalis (E. faecalis), and Proteus mirabilis (P. mirabilis) [22, 24] are commonly associated with implant‐associated infections.

This remains the leading cause of mortality, prolonged hospital and ICU stays, and rising healthcare expenses. The primary approach for management of infections is the administration of antibiotics. However, many bacteria are resistant to commonly used antibiotics, including advanced drugs such as third‐generation cephalosporins [25] and carbapenems [26], making infections harder to treat and leading to adverse outcomes. Hence, the research shifted to alternatives, such as surface‐engineered biomedical devices. To alter the material‐biological interface and thereby influence biofilm formation, effective surface engineering strategies require a comprehensive understanding of the interactions at the material–microbe interface that govern adhesion, growth, and maturation.

3. Formation and Combating Biofilms Affecting Polymeric Devices

The biofilm is an organized microbial consortium in which cells are surrounded by an extracellular matrix (ECM) and show resistance to the immune response and to antimicrobial agents. Formation of biofilms is a dynamic, multi‐step process in which microbial cells attach to a surface, ultimately evolving into a surface‐associated, multicellular lifestyle characterized by the deposition of extracellular materials. A critical factor in biofilm development is the Vroman Effect [27], which describes the sequential adsorption of biomolecules upon implantation. Following implantation, low‐molecular‐weight proteins, such as albumins, antibodies, and red blood cells, rapidly adsorb to the implant surface, forming a conditioning layer that facilitates bacterial adhesion. These components serve as scaffolds, facilitating cell‐to‐cell adhesion and the formation of microcolonies. Structures associated with the bacterium, such as flagella, fimbriae, and pili, facilitate approach and initial binding to surfaces. Specifically, environmental signals, such as surface contact or nutrient availability, trigger the synthesis of the second messenger cyclic di‐GMP (c‐di‐GMP), which governs the transition to a sessile lifestyle by promoting adhesin and EPS (extracellular polymeric substances) synthesis [28]. Upon sustained environmental cues and increased c‐di‐GMP levels, as Figure 2 demonstrates, the bacteria irreversibly attach and begin producing EPS components, mainly polysaccharides, proteins, and extracellular DNA (eDNA). EPS serves as a protective barrier while also playing a crucial role in the formation of new biofilms and the dispersion of existing ones [29]. This stage is primarily associated with the enhanced expression of adhesins (e.g., CdrA) and the production of species‐specific polysaccharides, such as Psl, Pel, and alginate in Pseudomonas aeruginosa, as well as the accumulation of eDNA through programmed cell lysis [30, 31].

FIGURE 2.

Biofilm development stages and alternatives to achieve infection‐resistant implant surfaces. Reproduced with permission [32]. Copyright 2024, Antibiotics.

The second step leads to biofilm maturation, during which microcolonies grow and undergo architectural development to form a mature, 3D structure. The biofilm matrix, primarily composed of EPS, plays a critical role in this development by providing mechanical stability and surface cohesion, maintaining nutrient sorption and retention, promoting hydration, and protecting cells from desiccation and stress factors. Notably, EPS also contributes to the establishment of physicochemical gradients and micro niches within the biofilm. Cell density‐dependent signaling (quorum sensing) regulates the expression of matrix components and virulence factors. Quorum sensing regulates the production of rhamnolipids, lectins (LecA and LecB), siderophores, and DNA [33].

EPS not only ensures the structural integrity and adhesion of biofilms but also provides protection against antibiotics, oxidative stress, and the immune response. Additionally, retention of extracellular enzymes provides an outer digestive system and serves as a storage of nutrients and a vector for horizontal gene transfer (e.g., eDNA). The structural diversity of EPS leads to spatial heterogeneity and microdomain formation. The last step is the release of cells from the biofilm into the surrounding medium, which creates new surfaces for colonization. Dispersal may occur in response to nutrient depletion or environmental stress, enzymatic degradation of the matrix (e.g., glycosidases, proteases, DNases), genetic regulation of transport and catabolism systems such as dspB‐encoded β‐hexosaminidase, which breaks down matrix constituents like poly‐β‐1,6‐N acetylglucosamine [34].

As bacterial populations proliferate, they secrete ECM components, reinforcing biofilm integrity and enhancing resistance to antibiotics and host immune defenses. This biofilm formation process not only contributes to implant failure but also enables the persistent survival of pathogenic bacteria in a dormant state, reactivating under favorable conditions. Biofilm‐associated infections are particularly concerning in long‐term medical device use. Biofilm dispersion further exacerbates this issue, as detached bacterial clusters can colonize new sites, increasing the risk of systemic infections. The long‐standing combat approach involved the administration of antibiotics and lock solutions [35, 36]. However, this preventive approach has significant downsides, including antibiotic resistance due to the emergence of multidrug‐resistant organisms, biofilm tolerance due to the protective matrix, and the risk of systemic toxicity and disruption of the host microbiome from high or prolonged systemic dosing. Hence, surface modification of the polymeric implant surfaces emerges as an effective strategy.

Figure 1, 2 summarizes effective prevention strategies that primarily focus on interrupting the early stages of biofilm development, namely microbial adhesion and initial sessile growth. Inhibiting microbial adhesion is a critical first step, which can be achieved through the application of nanoparticles (NPs), antimicrobial peptides (AMPs), photodynamic therapy (PDT) [37], and a variety of natural or synthetic anti‐adhesive compounds. Surface modification of medical devices with antimicrobial agents, enzymes, NPs, AMPs, or antiadhesive coatings further reduces the likelihood of microbial colonization by creating a hostile environment that deters bacterial attachment and biofilm initiation. Tweaking the surface energy [38, 39], morphology and surface charges can effectively reduce bacterial adhesion [40]. In addition, biochemical processes that support biofilm development offer an additional layer of prevention. Early intervention at the stages of adhesion, EPS production, and quorum sensing is crucial to effectively prevent biofilm formation, thereby enhancing the safety, functionality, and longevity of medical devices. This can be achieved by modifying the surface to prevent bacterial attachment or proliferation. The subsequent sections outline various strategies for surface‐engineering biomedical devices to prevent infections associated with implant surfaces.

4. Polymers in Biomedical Implants and their Surface Modification Strategies

Surface modifications of biomedical implants are a crucial, widely accepted technique for enhancing their performance, biocompatibility, and therapeutic efficacy. Various strategies have been developed to modify the surface properties of these devices, aiming to improve their interactions with biological systems, reduce adverse reactions, and introduce specific functionalities for targeted applications. The next section briefly discusses various modification strategies for polymeric implant surfaces, and Table 1 summarizes the biological performance and antibacterial properties of these approaches.

TABLE 1.

Biological performance and antibacterial properties of coatings on commonly used biomedical implant surfaces.

Substrate	Coating	Method of deposition	Antibacterial activity	Causative organism	Protein adsorption	Cytotoxicity	Stability /other studies	References
Polydimethylsiloxane (PDMS)	Hyaluronic Acid‐Gelatin‐Crosslinked Hydrogel (3:7)	Silanization	No live cells ∼1.5%, ∼5% (higher adhesion) (LIVE/DEAD fluorescence)	P. aeruginosa, R. pickettii S. epidermidis	—	Human‐derived adipose stem cells (hASCs) no toxicity	Stable hydrophilicity‐ 7 days under PBS environment. Swelling ratio stable (∼15%)‐7 days under PBS environment	[41]
	TBA‐LT/PDMS	Spray coating	95.2%, 74.1%	E. coli, S. aureus	—	—	—	[42]
	p(SBMA‐co‐NHSMA) p(SBMA‐co‐NHSMA) PHMB copolymer	Grafting	∼98% (70: 30), ∼98% (80: 30) ∼99% (70: 30), ∼99% (80: 30)	E. coli, S. aureus E. coli, S. aureus	—	No toxicity against C2C12 cells (in vivo)	—	[43]
	HS‐PEG‐PCN(18:26) HS‐PEG‐PCN, (18:10) HS‐PEG‐PCN, (18:0)	Grafting (PDA)	100% (1.8 mM) 100% (1.8 mM) No antibacterial activity	MRSA	BSA absorption (Minimal Δf and ΔD changes)	—	HS‐PEG‐PCN, (18:10) antibiofilm activity	[44]
	Benzalkonium chloride	Dip coating	5.3 logs (Artificial Urine Medium) No antibacterial effect	E. coli, P. aeruginosa	—	—	—	[45]
	Ag/Au‐TANPs with Gel‐DOPA matrix	Dip coating (PDA)	99%,100% (Planktonic) 60%, 70% (Biofilms)	P. aeruginosa and S. aureus planktonic P. aeruginosa and S. aureus biofilms	—	No toxicity against Human foreskin fibroblasts (BJ‐5ta cell line) and keratinocytes		[46]
Polyurethane (PU)	P‐IMCTS‐PMPC	Grafting (PDA)	99.94 ± 0.06%, 99.82 ± 0.09%	E. coli S. aureus	89.30 ± 2.12(BSA) 85.78 ± 2.14 (Lys), 90.89 ± 2.08% (BSA+Lys)	80% (L929 cells)	4 Weeks stability	[47]
	Poly‐L‐lysine	UV‐induced surface‐initiated radical graft polymerization	∼1.3 × 104 cells/mm2 (PU) ∼3 × 104 cells/mm2 (PU)	P. aeruginosa	—	No cytotoxicity against L929 cells	2 weeks (isotonic pH)	[48]
	PDA/Ag 5 PDA/Ag 10	Dip coating (PDA)	∼80%, ∼95% ∼90%, ∼98%	E. coli S. aureus	—	—	Stable release of AgNP for 7 days (ICPMS)	[49]
Polyvinylidene fluoride (PVDF)	PVDF‐g‐ATMAC (7.5 M) PVDF‐g‐DMC (7.5 M)	Grafting	99% 99%	E. coli S. aureus	98% (FRR) 96% (FRR)	—	—	[50]
Polytetrafluoroethylene (PTFE)	Silver nanoparticles	In situ nanoparticle coating (DOPA)	92% (4 h incubation) and 38% (24 h incubation) 93% (4 h incubation) 49% (24 h incubation	E. coli S. aureus		—	Reduction of NIH3T3 fibroblasts (Formation of less fibrous tissue on implantation)	[51]
	Argon (121°C for 480s)	Plasma treatment	—	—		84% viability after 144 h (Vascular smooth muscle cells)		[52]

4.1. Chemical Modifications

Chemical modification is a pivotal strategy for modifying the surface properties of biomedical devices to enhance hydrophilicity, charge distribution, and bioactivity. This approach involves incorporating specific functional groups, such as amine (─NH₂) [53] and hydroxyl (─OH) [54], which improve protein adsorption, facilitate cellular attachment, and exhibit antimicrobial activity [55, 56, 57]. Beyond simple surface functional group modifications, advanced surface engineering approaches, such as polymer grafting, self‐assembled monolayers (SAM), and silanization are notable examples.

Grafting enables the formation of dense polymer chains on the substrate, thereby allowing precise tuning of the surface interface. Depending on their nature, these polymer chains interact with the biological interface, influencing cell adhesion, protein adsorption, and the overall compatibility of the bioactive surface. The methods differ in how polymer chains or functional groups are attached to a substrate surface. According to the deposition route, this is further divided into three types. “Grafting‐to” approach involves attachment of complementary reactive groups with functionalized polymers, such as a combination of amine (─NH₂) and carboxylic acid (─COOH) [58], Azide (─N₃), and alkyne (─C≡CH) [59]. In simple terms, as depicted in Figure 3a, this approach involves presynthesized functionalized polymers (black lines with red circles) covalently attached to reactive surface groups (inverted umbrella structure), providing precise control over molecular weight and dispersity before immobilization. It is particularly advantageous when well‐defined polymer architectures must be validated before surface conjugation. Building on this approach, a copolymer of an imidazolium salt chitosan, poly(2‐methacryloyloxyethyl phosphorylcholine), and acrylic acid was grafted onto a PDA/polyethyleneimine‐coated polyurethane surface [47]. The deposition scheme for facilitating effective grafting of PU surfaces is shown in Figure 3b, where the copolymer's carboxyl groups were activated to form NHS‐activated groups, enabling subsequent covalent attachment to the modified surface. The sequential modification of PU substrates was facilitated by functionalization using polyethylenimine/polydopamine. This functionalization led to an increase in surface roughness (83.13 ± 3.34 nm). Moreover, the amino and phenolic hydroxyl groups of the functionalized surface led to a decrease in the water contact angle to 55° ± 2.64° compared to the pristine surface (90.75° ± 3.24°). Further, owing to the uniform and adherent coating of the copolymer system (IMCTS−PMPC−PAA), a reduction in surface roughness (24.53 ± 1.78 nm), thereby leading to reduced contact angle (18.09° ± 2.45°), was observed. The hydration layer of the zwitterionic polymer (PMPC), in addition to surface roughness, was attributed to the hydrophilicity of the coated surface. A similar multifunctional copolymer coating was developed by Ding et al. [44] to enhance surface properties. This study reported the grafting of amine‐reactive, thiol‐terminated PEG‐b‐PCN diblock copolymers onto polydopamine‐functionalized PDMS. The stepwise deposition is shown in Figure 3c, where reactive functional groups (─NH₂, ─OH) are first introduced onto the silicone surface via polydopamine functionalization. Subsequently, thiol groups on thiol‐terminated PEG‐cationic polycarbonate diblock copolymers (HS‐PEG‐PCN) enabled grafting to surface‐bound catechol functionalities. The polymers with varying compositions, defined by a constant MTC‐OPrBr content (18 units) and different MTC‐OEt contents (26, 10, and 0 units), were prepared and grafted onto a PDA‐modified PDMS surface. An increase in MTC‐OEt content led to enhanced hydrophilicity with HS‐PEG‐PCN (18:10) (∼68.3°), and HS‐PEG‐PCN(18:26) (∼68.9°) as compared to HS‐PEG‐PCN (18:0) (∼71°). The coating thickness, as determined by the viscoelastic model, increased significantly from 5.2 nm (0.075 mM) to 7.0 nm (0.75 mM) at different polymer concentrations. However, a further increase in concentration (1.8 mM) resulted in a smaller increase in thickness. The antibacterial performance of these PEG‐coated surfaces at two polymer concentrations (0.75 and 1.8) against Methicillin‐resistant Staphylococcus aureus (MRSA) was compared in Table 1.

FIGURE 3.

(a) Schematic depiction of chemical modification using the grafting‐to approach. Reproduced with permission [60]. Copyright 2020, Molecular Rapid Communications. (b) The deposition scheme of grafting to the deposition of copolymers on the biomedical implant surface. Reproduced with permission [47]. Copyright 2024, ACS Applied Polymer Materials. (c) Stepwise modification of PEG‐based modification of silicon rubber by polydopamine functionalization. Reproduced with permission [44]. Copyright 2022, Biomaterials.

Overall, the grafting‐to approach enables the incorporation of diverse functional moieties and control over polymer structure; however, it is limited by low grafting density due to steric hindrance between pre‐synthesized polymeric chains. Moreover, this technique relies on surface activation steps, such as polydopamine coating or silanization, which increase fabrication complexity. The challenges underscore the need for alternative approaches.

“Grafting‐from” approach, on the other hand, provides a high density of functional sites [61] due to outward growth of polymer chains from surface‐bound initiators, while enabling fine control over polymer brush density and thickness. This technique can be described as in situ polymerization of monomers, i.e., polymerization directly from immobilized active sites on the substrate surface, as shown in Figure 4a. The surface‐bound initiator (gray triangle) is anchored to the substrate, from which polymer chains originate. Upon activation (right), the initiator triggers the growth or extension of polymer brushes into the surrounding medium. Based on this strategy, DMC ([2‐(metacryloyloxy)ethyl]trimethylammonium chloride and ATMAC (allyltrimethylammonium chloride) quaternary ammonium salts (QAS) were grafted onto PVDF membranes under UV by a surface‐initiated polymerization process [50]. Initially, PVDF substrates were immersed in varying concentrations of monomer (initiators) (2.5–10 wt%) and then exposed to 365 nm UV radiation (100 W) (activation). This process was performed in a nitrogen environment for 30 min, followed by washing.

FIGURE 4.

(a) Schematic depiction of chemical modification using the grafting‐to approach. Reproduced with permission [60]. Copyright 2020, Molecular Rapid Communications. (b) Reproduced with permission [61]. Copyright 2011, Molecular Bioscience.

Until the mass reached equilibrium, the samples were cured by heating (40°C) under vacuum. The optimal QAS concentration was determined by grafting degree, wettability, and flux analysis of coatings at different concentrations. Based on these parameters, the QAS concentration of 7.5 wt% was found to be optimal, as it provides the maximum grafting efficiency and flux enhancement compared to 10 wt%, which leads to reduced permeability. The PVDF‐ATMC grafting of QAS at 7.5 wt% resulted in an enhanced zeta potential of ‐20.15 mV attributed to protonation of nitrogen‐containing functional groups. However, the DMC grafting showed a slight increase of ‐29.59 mV compared to ‐30 mV for uncoated PVDF surfaces. The hydrophilicity of both surfaces was reported to be ∼45°. The effectiveness of the coatings was validated through antibacterial studies against common HAI pathogens, followed by flux recovery ratio testing. FRR is a quantitative measure of the recovery of intrinsic water permeability of the membrane after fouling (e.g., protein adsorption) and subsequent cleaning. The quaternary ammonium compounds, however, are often associated with hypersensitivity [62], moreover higher concentrations may result in tissue irritation. Hence, bioavailable polymer coatings can be an efficient approach. Ultraviolet/ozone (UVO) treatment activated 3‐layered PDMS substrates grafted with POEGMA brushes were evaluated for the implant packaging application [63]. The brush polymer growth led to an increase in hydrophilicity over time, with contact angles of 71.5°, 58.6°, and 54.5° at 30, 60, and 90 min of polymer brush growth, respectively. Moreover, the grafted polymers did not influence the mechanical properties of the bulk PDMS. Although this technique enables enhanced graft density, polydispersity is a common drawback. This usually arises from differences in chain initiation (i.e., some polymers grow faster than others) or termination reactions, thereby reducing grafting efficiency. Therefore, grafting‐through emerges as an effective solution to overcome these limitations.

The grafting‐through approach depicted in Figure 5a involves surfaces functionalized with a self‐assembled monolayer bearing polymerizable groups.

FIGURE 5.

(a) Schematic depiction of chemical modification using the grafting‐through approach. Reproduced with permission [60]. Copyright 2020, Molecular Rapid Communications. (b) Outline of the Surface Functionalization and polymerization reactions. Reproduced with permission [64]. Copyright 2016, Macromolecules.

In this technique, polymer chains in solution grow and are further incorporated onto surface‐bound monomers, ensuring control over the polymer brush structure. This enables uniform chain growth (lower polydispersity). Thus, grafting‐through is an effective and dependable technique for creating stable, surface‐attached polymer layers because it often exhibits a self‐limiting growth mechanism and significant resistance to reaction conditions. Owing to these advantages, hydrophilic modification of dialysis membranes via a membrane‐assisted ATRP “grafting through” approach [64] was performed. The membranes were grafted as per the scheme shown in Figure 5b, in which they were functionalized with α‐bromoisobutyryl bromide to introduce initiator sites. In the presence of an aqueous solution of N‐hydroxyethyl acrylamide (HEAA), the modified dialysis membranes were further immersed in a solution containing CuBr/CuBr₂ and PMDETA under an inert atmosphere. This system leads to polymerization while catalysts diffuse from the external solution to activate surface‐bound initiators. At the same time, the monomer (HEAA) diffuses through the membrane, enabling polymer chain growth and establishing an inverted monomer concentration gradient. This approach leads to controlled, dense, uniform polymer brushes with reduced termination and lower polydispersity.

Other chemical surface modification strategies for engineering polymeric surfaces include plasma treatment, UV/ozone oxidation, silanization, and layer‐by‐layer (LbL) assembly. Reznicková et al. [65] reported plasma modification of the PTFE membrane by argon plasma. This was done at different temperatures (37°C, 60°C, 100°C, and 121°C) for 60 and 480 s. The maximum hydrophilicity (67° ± 1.6°) was observed at 121°C for 480 s, attributed to the introduction of carboxyl groups and the cleavage of bonds in the original polymer chains. However, hydrophobic recovery limits their applications in long‐term implant surfaces [66, 67]. An alternative approach in which UV illumination (254 and 185 nm) was applied to smooth and microbead‐embedded PDMS for varying durations [68]. This creates a hydrophilic SiOx layer on PDMS, but hydrophobic recovery occurs over time due to polymer relaxation and chain diffusion. To overcome this limitation, Mao et al. [69] reported the formation of hydrogels under UV light for surface modification with enhanced adhesion. This novel method for coating the polymer with a hydrogel under UV light. Polymer formation occurs via the generation of an amine alkyl radical, produced by the complexation of hydrogen donors under UV light. This coating method is adaptable to all kinds of biomedical devices, regardless of shape. As the PU and hydrogel are chemically bound, the adhesion strength reaches 0.258 MPa and remains stable for over 14 days.

Silanization emerges as an approach overcoming the limitations of plasma and UV light activation. This method allows stable hydrophilic modification. Thus, Joo et al. [41] reported modification of the PDMS surface through silanization to alleviate medical device‐induced fibrosis. Further, the functionalized surfaces were conjugated with gelatin using (3‐aminopropyl) triethoxysilane (APTES), followed by EDC/NHS chemistry. A hydrogel composed of oxidized hyaluronic acid (CHO─HA) and gelatin amine (Gel─NH₂) crosslinked via a Schiff‐base reaction was coated onto the gelatin‐conjugated PDMS. CHO─HA/Gel─NH₂ with three different ratios 5:5, 3:7, and 7:3 led to hydrophilic modification of the coated surface with water contact angle of 33.14° ± 11.76°, 45.25° ± 4.53°, and 51.52° ± 2.45°. Based on a comparative assessment of these parameters, the optimal formulation was identified and is summarized in Table 1. Among these approaches, Layer‐by‐layer (LbL) offers exceptional versatility. These multifunctional coatings are formed by alternating deposition of oppositely charged polyelectrolytes onto charged surfaces. An example of such a coating on intraocular lens (IOL) is reported. In this study, Liu et al. [70] reported a superhydrophobic and transparent antibacterial coating on intraocular lenses (IOLs) fabricated via LbL electrostatic self‐assembly. The primary layer was coated with (sodium 4‐styrenesulfonate) (PSS)/poly (allylamine hydrochloride) (PAH)_, and dipped for 4 cycles. Further (PSS/PAH)₄ was coated with a secondary layer of alternating SiO₂ and PAH layers. The entire coating was further fluorinated using 1‐H,1‐H,2‐H,2‐H‐perfluorooctyltriethoxysilane (PFOTS). The coating was denoted as (SiO₂/PAH) _n ‐F (n‐dipping cycle). An optimal contact angle of 162.8° was reported for (SiO₂/PAH) 6.5‐F. The coated IOLs maintained image resolution at a 3 mm‐diameter region, making them physiologically equivalent to the human pupil. Further using a similar LbL approach, Peng et al. [43] fabricated two types of hydrophilic coatings to achieve antifouling as well as antibacterial silicone rubber surfaces, p(SBMA‐co‐NHSMA) copolymers, and other PHMB polymers. The grafting of SBMA ([2‐(methacryloyloxy)ethyl]dimethyl(3‐sulfopropyl)ammonium hydroxide (sulfobetaine methacrylate)) and NHSMA (N‐hydroxysuccinimide methacrylate) was done by APTMS/APTES‐mediated silanization of silicone, thereby leading to the fabrication of an amine‐terminated surface, providing reactive sites for the grafting of p(SBMA‐co‐NHSMA) copolymers. The unreacted NHSMA pendants became sites for further immobilization of amino‐terminated PHMB. The hydrophilic nature of the coated polymers was attributed to SBMA, as the water contact angle was ∼70° with or without PHMB, compared to uncoated silicone surfaces (114°). Despite their ability to modify surface properties, conventional chemical approaches often face limitations in biological availability

4.2. Biomolecule Immobilization

Biomolecules, such as proteins, peptides, growth factors, and naturally occurring molecules, can be immobilized on polymeric surfaces to attain bioactive surfaces. This approach leads to the bio‐integration of biomedical devices with the host system through functionalization, thereby improving cell adhesion, promoting cellular differentiation, and facilitating tissue regeneration. Antimicrobial drugs as coating materials for combating biofilms were among the first strategies adopted and commercially accepted for coating biomedical devices, such as catheters. Commonly used drugs, ciprofloxacin and chlorhexidine, are used individually and in combination in the fabrication of antibacterial surfaces. In this study, sterilized PVC substrates were irradiated with UV light for an hour, followed by immersion in bicine buffer (10 mM, pH 8.5) containing dopamine, as shown in Figure 6a. Two different concentrations (0.5 and 2 mg/mL) of both the individual and the combination were prepared. The roughness increased with increasing drug coating concentration, imparting a hydrophobic characteristic to the surface. These coatings were evaluated against common ventilator‐associated pneumonia‐related pathogens [71]. The combination coating (CIP/CHX) performed better than the individual coatings, resulting in a subsequent reduction in dual‐ and triple‐species biofilms. However, concentrations above 0.5 mg/mL exhibited cytotoxic effects; therefore, subsequent studies were conducted at this threshold. Beyond cytotoxicity, antimicrobial resistance due to the overuse of antimicrobial drugs is reported. Therefore, the emergence of antimicrobial resistance to commonly used antimicrobial agents has become prominent [72]. Building on this approach, silk fibroin nanoparticles loaded with rapamycin and paclitaxel [73] were coated on PDMS. Figure 6b shows the sequential modification of PDMS via texturing and coating. In short, PDMS was prepared on the lotus leaf template, which served as a mold; the resultant structured surface was the lumen. The outer surface was reinforced with a braided polydioxanone monofilament, which was later incorporated with the prepared drug‐loaded silk fibroin microspheres. The morphological evaluation of the lotus leaf‐based lumen of the stent reported papillae (5–8 µm) with a spacing of 10 µm, microstructured surface led to superhydrophobic (∼122°) modification. These modified surfaces were molded into cylinders and implanted in mice as ureteral stents to evaluate their efficacy. The compression modulus (1.81–2.73 MPa) of the fabricated stents exceeded that of commercial stents, and the elastic recovery rate after 100 compression cycles remained consistent, ranging from 89.32% to 95.23%, exhibiting exceptional mechanical properties.

FIGURE 6.

(a) Combination of antimicrobial drugs coated on PVC substrates via dopamine functionalization. Reproduced with permission [71]. Copyright 2024, ACS Applied Materials & Interfaces. (b) PDMS tubes modification by microfabrication and immobilization of biomolecules (silk fibroin), respectively, for ureteral stent application. Reproduced with permission [73]. Copyright 2025, International Journal of Biological Macromolecules. (c) Polysiloxane Nanofilaments Infused with silicone oil for intranasal splints. Reproduced with permission [74]. Copyright 2021, ACS Biomaterials Science & Engineering.

Moreover, the fabricated surfaces showed sustained release of rapamycin and paclitaxel drugs, which was validated by drug release studies for over ∼7 days. The release rates from the microspheres were 40.36% for rapamycin and 38.21% for paclitaxel. Although the system demonstrates sustained drug release, the immunosuppressive nature of rapamycin and its lack of antimicrobial activity may increase the risk of bacterial adhesion and biofilm formation.

In addition, simple structures such as small organic molecules and salts exhibit antibacterial properties [75]. Kasapgil et al. [74] coated polyether‐based intranasal splints using a process as shown in Figure 6c. Briefly, they reported coating polyether‐based splints and catheters with n‐propyltrichlorosilane, later lubricating them with high‐viscosity silicone oil to prevent bacterial adhesion. The infusion of 1000 cst oil onto the coated surface makes the surface hydrophobic and reduces the blood sliding angle to 16.5° ± 2.2°, thereby rendering the surface antibacterial and antithrombotic.

Metabolites and natural compounds with inherent bioactivity are also used to tackle the issue. One such study by Bartosova et al., [76] reported the application of phenolic monoterpenes on biomedical fibers to improve antibacterial efficacy. The study analyzed four major compounds (cinnamaldehyde, eugenol, carvacrol, and thymol) at various concentrations and found differing antibacterial effects. Among them, cinnamaldehyde (Cin) exhibited the most potent antibacterial activity, while eugenol, carvacrol, and thymol were comparatively less effective against S. aureus and E. coli. The antibacterial effect of PLA/6Cin shows larger inhibition zones than that of PLA/4Cin. This modification also resulted in improved biocompatibility without significantly compromising mechanical integrity. PLA/6Eug, for instance, maintained tensile strength close to that of pristine PLA (33.8 vs. 39.8 MPa), ensuring material robustness. Another enzyme, horseradish peroxidase, was coated onto polyurethane by Kreider et al. [77]. Polyurethane (2C‐PUR) was modified with PDMS as a surface‐active modifier, and later this PUR (a copolymer of polyurethane and PDMS) was used. This surface was further activated via air plasma, converting hydrophobic (─CH₃) groups to hydrophilic (─OH), followed by silanization, enabling stable covalent linkage sites. The salinized surfaces were immersed in glutaraldehyde, leading to the conversion of imine to stable secondary amine. These modified surfaces were immersed with horseradish peroxidase (2 mg/mL) for 14 h. These coatings have been proposed for biomimetic applications.

Further, Biopolymers, such as poly‐L‐lysine, can be grafted onto the polymer surface to form a hydrogel‐based coating on biomedical tubings [78]. Herein, polyurethane sheets were used as a model substrate. The sheets were immersed in a mixture of a UV‐sensitive agent (benzophenone) and acetone for 5 min via hydrogen abstraction. Sulfobetaine methacrylate and hydroxyethyl methacrylate (3:7 molar ratio, 10 wt%) monomers were used for the fabrication of the coating. Post‐immersion, the sheets were irradiated by UV (365 nm, 5 min). Residual monomers were removed by washing and drying, followed by aldehyde functionalization with the biopolymer poly‐L‐lysine in PBS buffer for 24 h. Thus, an antifouling hydrogel layer was grafted onto the polyurethane surface, followed by oxidation of hydroxyl groups to aldehydes. Subsequently, a bactericidal poly‐L‐lysine layer was conjugated via Schiff base reactions, forming a cross‐linked dual‐layer hydrogel coating. Biopolymers such as heparin in combination with castor oil In this study, urinary catheters were fabricated via dip‐coating. A robust 3D network of CMC and PVP through UV irradiation lead to entrapment of AgNPs synthesized by the chemical reduction of silver nitrate with sodium citrate and glucose as precursor.and chitosan for modifying polyurethane‐based biomaterials was reported to enhance their bioactivity and mechanical properties [79]. These modifications have resulted in improved mechanical strength, reaching up to 1.00 MPa, comparable to that of human venous structures, primarily due to hydrogen bonding between hydroxyl (─OH) and carboxyl groups (─COOH) functional groups present in the organic oil and biopolymer. Such enhancements ensure that the modified materials can withstand physiological stresses while maintaining their functional properties. Other biopolymers, such as chitosan [80], cellulose, hyaluronic acid [81] show great potential. Additionally, heparin coatings [82, 83] and other anticoagulants as coating for thrombogenic coatings are reported. Though biomolecules are effective, they pose certain limitations, particularly in antimicrobial drugs; the increasing prevalence of antimicrobial‐resistant strains limits their use in long‐term applications. In addition, enzymes under harsh pH tend to degrade. Hence, robust systems provide both stability and effective antimicrobial activity.

4.3. Nanotechnology‐Based Approaches

Nanomaterials, including nanoparticles and nanostructured coatings, enable precise molecular‐scale functionalization, enabling the development of advanced biomedical surfaces with tailored properties. These nanoparticles can be synthesized using various processes, such as chemical [84], biological [85], and physical methods. The integration of nanomaterial‐based coatings onto medical devices enhances their mechanical durability, wear resistance, and antimicrobial efficacy. Additionally, incorporating nanoparticles into device surfaces enables controlled drug release and bioimaging [86, 87].

Zinc Oxide (ZnO) coated urinary catheters were fabricated by Perelshtein et al. [88] using sonochemical coating. In addition, Zinc Oxide‐Carbon and Zinc Oxide‐Silicon were fabricated by spray coating and electron‐beam physical vapor deposition (EB‐PVD), respectively. The stability and antibacterial properties of the coated catheters were evaluated for urinary catheter application. After 14 days of immersion in an artificial urine medium, ZnO‐C coatings were reported to be the most stable; however, ZnO coatings showed approximately a 90% reduction in coating thickness. The same effect was observed in the coatings' antibacterial activity as well. The ZnO‐C coatings were found to be the most stable and to exhibit better antibacterial activity than other coatings. Hence, nanoparticle adhesion is an important factor in coating. These can be achieved through some of the methods such as silanization [89], plasma modification mentioned in previous section. Polydopamine (PDA)‐assisted deposition or coating methods is one such widely used approach. In this method, polydopamine PDA acts as an adhesive agent, leading to ─NH₂ and ─OH functionalization of the surface. An example of such coatings is demonstrated by Cai et al.  , where PDA is used as an adhesive layer for carboxymethyl cellulose (CMC)‐Ag coatings [90]. A robust 3D network of CMC and PVP through UV irradiation lead to entrapment of AgNPs synthesized by the chemical reduction of silver nitrate with sodium citrate and glucose as precursor.The deposited coatings showed no changes post‐abrasion. Moreover, the surface roughness decreased with the coatings, and the coatings exhibited superhydrophobic properties with a lower surface free energy (6.9 ± 14) compared to 30.9 ± 1.7 for the uncoated surface.

Beyond surface functionalization, binding agents/matrix materials such as HPMC (Hydroxypropyl Methylcellulose) can be employed, as reported by Malhotra et al. [91]. This study coated PDMS and (TIVAPs) with phytoassisted ZnO nanoparticles. In this study, zinc oxide nanoparticles were synthesized from ethanolic extracts of Eupatorium odoratum, acting as a reducing and stabilizing agent. Subsequently, the synthesized ZnO nanoparticles were coated on the catheter's surface using HPMC, enabling adhesion via non‐covalent interactions. Both the outer surface and the catheter lumen were coated by immersion and infusion, respectively, resulting in hydrophilic modification. The long‐term stability and durability of this coating remain uncertain, necessitating further evaluation to ensure reliable performance in biomedical applications. A rather advanced approach, nanofabrication for coating, can also be used. Based on this strategy, zinc oxide was coated on a polyurethane acrylate coating. Zinc nitrate dihydrate and hexamethylenetetramine were used as a precursor for hydrothermal growth [92]. Initially, a polyurethane acrylate precursor was spin‐coated onto an anodic aluminum oxide template with a 150 nm diameter. The coated template was then sandwiched between Teflon films and imprinted under vacuum to facilitate polymer infiltration into the nanopores. The polymer was subsequently cured under ultraviolet irradiation, forming a nanopillar‐structured coating after demolding. Later, a seeding solution containing zinc acetate dihydrate and triethylamine was heat‐treated at 85°C, followed by 120°C for 1 h. The coated seed solution was subjected to hydrothermal growth of zinc oxide nanostructures using zinc nitrate and hexamethylenetetramine as precursors. These nanostructured coatings showed efficient antibacterial activity owing to the hydrophilicity and mechanical action of the fabricated nanopillars. However, the scalability of these coatings onto biomedical implants may limit the application.

Silver nanoparticles (AgNPs) were among the earliest antimicrobial nanomaterials to gain significant recognition and have been extensively studied for their broad‐spectrum antibacterial efficacy. The potency and broad‐spectrum antimicrobial activity of the silver nanoparticles effectively mitigate bacterial colonization and biofilm formation [93]. Owing to these properties, AgNPs have been widely incorporated into coatings for medical devices, such as catheters, wound dressings, and implants. In a related approach, Nardo et al. [51] coated PTFE membranes with Ag⁺ ions via 3,4‐dihydroxy‐DL‐phenylalanine (DOPA) functionalization. In this technique, functionalized PTFE membranes were incubated in 50 mM AgNO₃ solution, leading to the formation of AgNP nanoparticles via catechol‐mediated chelation of Ag⁺ ions, followed by in situ reduction to metallic Ag⁰ on the membrane. Similarly, polydopamine/chitosan was enhanced by the deposition of silver nanoparticles (AgNPs) [49] using silver nitrate as a precursor as in Figure 7a. In this approach, sterilized urinary catheters were primarily coated with polydopamine (PDA), which served as an interfacial layer and a reducing agent (AgNO₃ at 5 and 10 mg/mL) for the formation of AgNPs. An average particle size of 30–50 nm was obtained. The release studies of the coated surface showed a consistent release for ∼7 days.

FIGURE 7.

(a) Schematic of PDA‐assisted dip‐coating of AgNP on silicone catheter surface for antibacterial modification. Reproduced with permission [49]. Copyright 2019,Heliyon. (b) Sonochemical coating of silver nanoparticles on PDMS via ultrasonic cavitation‐induced in situ reduction and uniform surface deposition. Reproduced with permission [46]. Copyright 2024, ACS Applied Materials & Interfaces.

However, due to the potential cytotoxicity of silver nanoparticles, combining them with natural metabolites as reducing agents to enhance their biocompatibility is a promising strategy. In this context, silicone‐based biomaterials were coated with silver/gold‐tannin nanoparticles (Ag/Au‐TANPs), capped with tannic acid, as shown in Figure 7b. This strategy employs dopamine‐functionalized gelatin as a matrix to ensure uniform nanoparticle immobilization via catechol‐mediated interactions, using the sonochemical method [46]. Nanoparticles such as zinc oxide, titanium dioxide, magnesium oxide, and silica have become more prominent for biomedical applications, with a focus on green synthesis approaches. Of these, typically recognized as safe (GRAS) nanomaterials such as ZnO and TiO₂ have been extensively studied owing to their excellent safety profiles. Furthermore, the introduction of these nanoparticles into medical devices has led to improved performance. A combination of silver and titanium oxide, coated on silicone catheters, exhibits strong self‐sterilizing and self‐cleaning properties [94]. Even after acid treatment with 1 M HNO₃, the antibacterial activity remained strong, suggesting that Ag nanoparticles were not only surface‐deposited but also embedded within the TiO₂ layer. The study concluded that 90 s of UV exposure to 0.2 mM AgNO₃ solution offered an optimal balance between performance and efficiency. Overall, Ag/TiO₂‐coated catheters present a promising solution for reducing catheter‐related infections in clinical settings. Despite these promising outcomes, concerns regarding long‐term stability persist, as the gradual degradation of coatings and the potential leaching of active agents in physiological environments (e.g., blood and urine) could limit their sustained efficacy. Addressing these challenges requires further optimization of coating composition and bonding strategies to ensure prolonged antimicrobial performance while maintaining biocompatibility and structural integrity.

Copper‐based coatings on PVC substrates as an alternative for thrombogenic coatings were assessed. Different coatings such as NE/HD (norepinephrine/hexanediamine)‐coated, antimicrobial peptide@NE/HD‐coated tubes, ABP@Copper(Cu)ions‐NE/HD, and Cu‐NE/HD along with pristine PVC were evaluated. Thrombus weights and occlusion rates significantly decreased in the Cu‐NE/HD and ABP@Cu‐NE/HD coated PVC [95]. Also, ABP@Cu‐NE/HD groups showed lower occlusion than the near‐complete blockage observed with cupric ions‐coated catheters. Both in vitro studies and ex vivo perfusion experiments demonstrated that the NO catalyzed by Cu²⁺ (i.e., ABP@Cu‐NE/HD) in the coatings exhibits superior anti‐thrombotic efficacy across various experimental conditions, irrespective of grafted ABP molecules, antibacterial peptide, or copper ions. The hydrogel was formed by chemical crosslinking of the polymers polyvinyl alcohol and polyethylene glycol acrylate (PEGDA), and by physical networking of the polymers with recombinant mussel protein (MFP‐5) in a slightly acidic medium. The sliding property of the coated catheter was evaluated using frictional forces, and the friction coefficients were 0.15 N and 0.04, respectively.

Furthermore, nanomaterials such as graphene‐based [96, 97], titanium oxide [98], chitosan nanoparticles, which impart antibacterial and self‐cleaning functionalities, have shown promise for further enhancing the efficacy of antiadhesive coatings. Furthermore, green methods such as using plant extracts [99, 100, 101] and biopolymers [102] for nanoparticle synthesis have been developed to impart biocompatible, inherently bioactive properties, thereby enabling the fabrication of antibacterial, antifouling, or antithrombogenic surfaces.

5. Biomedical Applications of Bioactive Surfaces with Antimicrobial Coatings

5.1. Antibacterial Surfaces

Antibacterial surfaces are engineered to incorporate bactericidal agents that eliminate microbial threats before biofilm formation (planktonic bacteria), thereby mitigating infection risks in biomedical applications. These can be achieved by modulating parameters such as surface roughness, wettability, and surface chemistry, thereby inhibiting or completely blocking bacterial activity on the coated surface [103]. These surfaces can be broadly categorized by mechanism of action, with active, passive, and stimulus‐responsive antibacterial coatings representing key approaches. Each strategy offers unique advantages in minimizing bacterial colonization, disrupting adhesion, or actively eliminating pathogens.

Active antibacterial surfaces include contact‐killing surfaces (i.e., post‐bacterial interaction) functionalized with antimicrobial peptides (AMPs), cationic polymers, or enzymatic agents that induce bacterial membrane disruption upon direct contact. A notable example of such a coating was developed by Lu et al. [104]. The team engineered a peptide‐polymer‐modified thermoplastic polyurethane surface exhibiting broad‐spectrum antibacterial activity. The peptide‐polymer (D, L‐lysine):(γ‐benzyl‐L‐glutamate) = 90:10 combination facilitated antibacterial efficacy by disrupting the permeability of gram‐negative bacterial membranes. The modified surface demonstrated 99.9% reduction in E. coli, 97.5% in MRSA, 99.9% in Staphylococcus haemolyticus, and 91.2% in Pseudomonas aeruginosa after 2.5 h of exposure. The shrinkage and severe damage to the bacterial membrane (Figure 8a) confirm the surface's active antibacterial activity. The in vivo studies further indicated a 0.85 log reduction of MRSA after 7 days of implant incubation. A similar example of contact killing was reported on dopamine‐functionalized PDMS slides, coated by synthesizing a linear block polymer via γ‐methylthiopropyl‐L‐glutamic acid‐based NCA, N‐tert‐butoxycarbonyl‐L‐lysine NCA, and L‐phenylalanine NCA in DMF. This polymer was later crosslinked into a single‐chain nanoparticle, in powdered form, synthesized by intramolecular cross‐linking of a linear polymer (P1) using 1,4‐diiodobutane in DCM, stirred for 72 h. The study showed that slides coated with LP‐KF (i.e., linear block polymer) showed 91% and 92.6% killing of S. aureus and E. coli, respectively, compared to uncoated surfaces. However, NP‐KF demonstrated superior activity, completely inhibiting the growth of both bacteria. The contact‐killing mechanism of the coatings was confirmed by SEM; it further revealed that NP‐KF showed higher antibacterial adhesion than LP‐KF. The coatings were confirmed to exhibit negligible toxicity towards RBC, indicating hemocompatibility. The cell viabilities of NIH 3T3 and HEK 293T cells were 89.7% and 90.6% for LP‐KF coatings and 88.2% and 91.5% for NP‐KF coatings, respectively, indicating low cytotoxicity [105]. The in vivo studies further confirmed the superiority of NP‐KF coatings, as despite a 4‐log reduction in the host system for up to 3 days on both coated surfaces, inflammation was observed around LP‐KF. Both the polymer‐coated showed significant reduction and efficient antibacterial activity, although fouling by these coatings can further lead to biofilm formation. Hence, coatings that prevent adhesion are preferred.

FIGURE 8.

(a) Active antibacterial activity of TPU coatings with shrinkage in the membrane of bacteria attached to the surface. Reproduced with permission [104]. Copyright 2023, Journal of Materials Chemistry B. (b) Passive antibacterial activity of ion‐coated polymer brush showing no adhesion of bacteria on the surface. Reproduced with permission [106]. Copyright 2015, Applied Material & Interfaces.

On the other hand, passive antibacterial surfaces are designed to prevent bacterial adhesion by modifying physicochemical properties. This technique, unlike active antibacterial surfaces, prevents biofouling. These surfaces involve tuning surface energy [103], hydrophobicity [107], or nanostructured topographies [108], to minimize microbial attachment while maintaining biocompatibility. Hydrophilic polyurethane catheter (polyvinyl propylene‐PDMS) coatings operate via the traditional passive antifouling mechanism [109], whereby increased surface hydration reduces protein adsorption and, in turn, bacterial adhesion. The high repellence against E. coli and S. aureus (96.2% and 95%, respectively) indicates that these coatings are effective during the initial phase of bacterial colonization (i.e., prevent adhesion). Similarly, the counterion‐mediated pTMAEMA brush system is a more advanced approach that allows dynamic tuning of surface properties [106]. These brushes were further enhanced with different counterions (e.g., PP⁶⁻, citrate³⁻, SO₄ ²⁻, Cl⁻). The quantification of dead and live bacteria for all coatings is shown in Figure 8b. The pTMAEMA‐PP6⁻ showed the highest antifouling activity, with no adhesion for all incident bacterium cultures. This activity was attributed to enhanced hydrophilicity of the coatings (8°) as compared to pristine brushes (60°). In addition, the ζ‐potential of pTMAEMA from +166 mV (in DI water) to −142 mV (in PP solution) indicates a charge reversal due to counterion binding. This reversal of the negative charge leads to electrostatic repulsion, thereby enabling control over microbial adhesion. The coatings showed high antifouling regeneration efficiency (>99% over repeated cycles). This dependence on regulated counterion environments raises questions about its feasibility in vivo, where the ionic composition is tightly regulated and not easily modulated. However, such results are typically obtained in controlled situations and may not fully represent the complexities of physiological conditions. In practical applications, the rapid evolution of conditioning films formed by proteins and biomolecules can mask surface characteristics, thereby reducing antifouling efficacy over time. Moreover, passive antibacterial coatings restrict bacterial adhesion on the surface, thereby contributing to antifouling behavior; however, they do not always modulate protein adsorption, an important aspect of antifouling coatings. The deposition of the protein may act as a secondary layer, masking the active hydration layer and thereby compromising the antibacterial efficacy of the coatings. Moreover, coating stability under physiological conditions (varying pH and temperature) is an issue. Thus, there is a pressing need for advanced coatings that combine stability with stimulus‐triggered activation to achieve efficient antibacterial performance.

Stimulus‐responsive surfaces exhibit antimicrobial functionality activated by external stimuli, such as pH, temperature, light, and an electric field. These dynamic materials undergo reversible structural modifications or controlled drug release in response to environmental triggers, enabling precise, localized antibacterial effects. Building on this strategy, ultrasonically activated coatings fabricated by Duan et al. [110] was reported. These ZnO NPs‐incorporated polyvinylidene difluoride‐hexafluoropropylene (ZnO‐PVDF‐HFP) coatings were fabricated by dip‐coating on silicone catheters. The detailed deposition scheme is shown in Figure 9a. The coatings were exposed to ultrasonic stimulation (100 W for 1 h). The effects of all coating components (i.e., ZnO and the polymer matrix) are individually studied to assess the efficiency of the coated surface. Moreover, the coatings’ antibacterial effect with and without ultrasonication were evaluated. The polymer matrix (PVD‐HFP) showed no antibacterial activity on its own, irrespective of ultrasonication. However, ZnO nanoparticles showed an enhancement to ∼60%–63% for both S. aureus and E. coli. However, the most efficient antibacterial activity was reported for ZnO‐PVDF‐HFP coatings, achieving ∼96% inhibition of both E. coli and S. aureus with ultrasonication. The difference in antibacterial activity between the coatings with and without ultrasonication are evident in Figure 9b.

FIGURE 9.

(a) Foley catheter surfaces were coated with ZnO‐PVDF‐HFP to achieve piezoelectrically activated antifouling. Reproduced with permission [110]. Copyright 2024, Materials Today Bio. (b) Difference in antibacterial activity of ZnO‐PVDF‐HFP coatings with and without ultrasonication. Reproduced with permission [110]. Copyright 2024, Materials Today Bio. (c) Comparative analysis of antibacterial activity of PT8‐I, PT8‐Cu, and PT8‐NFZ against P. mirabilis for different pH ranges. Reproduced with permission [111]. Copyright 2023, Journal of Materials Chemistry B. (d) Comparative SEM images over 15 days comparing antibacterial activity of PEG‐nanoparticle composite coated surfaces. Reproduced with permission [112]. Copyright 2025, International Journal of Biological Macromolecules.

Further, pH‐responsive antibacterial coating was fabricated by Miao et al. [111] on urinary catheters. In this study, polyzwitterionic hydrogel coatings were coated onto a PDMS surface. The hydrogel was fabricated by polySBMA and Tannic acid (TA) hybrid network via hydrogen bonding and electrostatic interactions. Three different coatings based on antimicrobial agents, including PVP‐I, Cu, and nitrofurazone (NFZ), were fabricated via a one‐step impregnation method. The hydrogel coatings exhibited pH‐responsive release behavior, with significantly greater antimicrobial release at pH 10 than at pH 7. All coatings showed enhanced bactericidal activity under alkaline conditions, which correlated with increased release. As shown in Figure 9c, the PT8‐I (hydrogels with PVP‐I) achieved complete bacterial eradication at both pH levels. Bacterial adhesion was markedly reduced on all antimicrobial‐loaded coatings, with PT8‐I showing no detectable bacterial attachment. However, the relevance of these coatings in alkaline conditions (physiological environments), including infection sites, which are more commonly neutral to slightly acidic are not explored. Therefore, the practical applicability of the pH‐triggered mechanism of the PT8‐I may be limited. In contrast, Li et al. [113] demonstrated a more clinically relevant system showing antibacterial activity triggered by acidic pH, elevated glucose levels, and dithiothreitol (DTT) [113]. In this study, cationic antibacterial agents, montmorillonite (MMT), chlorhexidine acetate (CHA), and PPPB were prepared, demonstrating possible applications across a wide range of substrates. Optimal release of CHA (70%) was observed at a slightly acidic pH of 6.0 and glucose and DTT concentrations of about 1 mg/mL. These coatings showed 99% antibacterial activity against E. coli and S. aureus. Furthermore, the coatings demonstrated effective results in infection‐induced mouse models for over 14 days. The reusability of LbL surfaces was reported for six cycles [113].

Light and temperature‐responsive coatings utilize photothermal or thermally activated mechanisms to enable localized antimicrobial treatment. NIR‐activated antibacterial coating was developed by Huang et al. via hot‐electron‐induced polymerization triggered by xenon light, in which poly((2‐dimethylamino)ethyl methacrylate) (PDM) brushes were grafted onto gold nanoparticles (AuNPs) [114]. The formed cationic polymer‐grafted AuNPs were subsequently combined with phytic acid (PA) to prepare PA/AuNPs‐g‐PDM (PAP) aggregates. The aggregates were deposited on titanium substrates by a simple surface adherence and gravitational settling method. The same methodology was applied to several substrates, including PDMS, but antibacterial tests were not conducted on them, opening a new avenue toward antibacterial coatings. Ti‐PAP coatings could efficiently synergize the contact‐killing and photothermal antibacterial activities, with 97.34% and 99.97% killing efficiency against S. aureus and E. coli adhesion under near‐infrared (NIR) irradiation (808 nm, 1 W/cm², 10 min). Without NIR activation, the inactivation efficiencies were 91.05% and 96.06% for S. aureus and E. coli, respectively, further demonstrating the coatings' dual functionality. The in vivo studies showed a reduction of 99.42% reduction in S. aureus colonization with minimal host tissue damage. Another study with similar stimuli activation, was developed by Wang et al. to obtain an antibacterial coating that consisted of a TA/Cu‐PEG hybrid film [112] on polyurethane surfaces. In this study, sequential adsorption of a tannic acid (TA)/Cu ion complex and PEG‐NH₂ was performed. The TA/Cu complex was designed as a dual‐role agent by NIR light, a photothermal agent, and a passive‐PEG immobilization linker. The resultant surface displayed ∼97% protein resistance. After 5 min of NIR irradiation (808 nm, 2.0 W/cm²), the coatings showed >95% bacterial killing and >90% bacterial antibacterial efficiency against S. aureus, E. coli, and MRSA. Further, in vivo studies revealed TA/Cu‐PEG‐modified PU implants showed up to 3.4 log reduction against S. aureus colonization and significantly reduced inflammatory response.

The in vitro studies showed effective inhibition of biofilm formation for 15 days with NIR radiation. Biofilm growth over 15 days, as shown in Figure 9d, demonstrated efficient inhibition of bacterial adhesion on the coated surfaces. However, without NIR activation, the coatings showed increased biofouling, which can be problematic. Thus, without such stimuli, their antibacterial effectiveness decreases. This dependence on external activation limits their therapeutic applications, as it limits penetration of light and controlled irradiation into deep tissues or into implantable medical devices.

Although antibacterial coatings are primarily aimed at removing or inhibiting microbial growth (either actively or responsively), biofouling remains a challenge. This challenge can be addressed through passive antibacterial coatings; however, protein adsorption remains an issue. Thus, antifouling coatings aim to prevent the initial attachment of biological pollutants (proteins, cells, etc.) and thus prevent biofouling on implant surfaces. In a variety of biomedical applications, particularly for long‐term implantation, surface fouling is a precursor to infection, inflammation, or device failure. Hence, incorporation of antifouling properties has emerged as an alternative strategy to improve the long‐term functionality and security of biomedical devices.

5.2. Antifouling Surfaces

Antifouling surfaces are designed to prevent the adhesion of proteins, bacteria, fungi, and other microorganisms, thereby inhibiting biofilm formation. In biomedical implants, biofouling is a critical concern, particularly in blood‐contacting devices, where the initial adsorption of proteins forms a conditioning layer that facilitates bacterial adhesion. This process not only promotes infections but also accelerates mechanical degradation, ultimately compromising device performance and longevity. Antifouling coatings function through two primary mechanisms: first, preventing the adhesion of fouling agents, and second, neutralizing contaminants upon contact [115]. The challenge posed by conditioning film formation can be mitigated by engineering surfaces that inherently resist biofilm accumulation and microbial adhesion, as well as protein adsorption. This can be achieved through surface topographical modifications [116, 117], or chemical functionalization [118], to hinder the initial attachment of biological entities.

Hydrophilic modification of Foley urinary catheters by a fatty acid (oleamide) and PDMS composite to achieve an antifouling surface [119] was reported to achieve antifouling. This was done by incorporating different concentrations of oleamide (weight percentages) into PDMS (OPC). These coated surfaces showed a significant decrease in biofilm adhesion with increasing OPC concentration. The activity of these coatings against E. coli was better than that against S. aureus, as E. coli adhered even at a higher OPC concentration of 2.5%. However, E. coli showed reduced adhesion to OPC at 1.25%. The surface resisted adhesion of the microbial species but showed no killing effect on S. aureus. At 5%, no adhesion of either species was observed, and biofilms were inhibited by 60.95% for E. coli and 90.92% at 5%. This surface, therefore, showed a synergistic effect on antimicrobial and antifouling activity against E. coli. In contrast, ferrocene‐containing multilayer system based on PEG: DMAEMA forms an antifouling coating that inhibits biofilm formation [120]. The combined antibacterial and antifouling efficacy was attributed to a low‐energy PEG surface resisting bacterial adhesion, and further DMAEMA (((Dimethylamino)ethyl methacrylate)) incorporation providing contact‐killing activity. As the number of bilayers increases (>6), bacterial attachment is almost completely suppressed, indicating a highly effective antibiofilm surface. Additionally, the modular polymer design enables tunable functionality (e.g., charge and composition), offering greater control over performance. However, the coating's stability and potential leaching remain concerns. In contrast, zwitterionic surfaces exhibit improved antifouling performance due to the formation of a strong hydration layer.

Zwitterionic polymer brushes coated surfaces are another prevalent approach to obtain antifouling surfaces. The antifouling efficacy of these coatings depends on the thickness and grafting density of the polymeric brushes. Ke et al. [121] reported antifouling modification of central venous catheters (CVCs) using a zwitterionic polymer brush coating [121]. Figure 10a illustrates a stepwise grafting of (poly[2‐(methacryloyloxy)ethyl phosphorylcholine]) pMPC on CVC. This study showed optimum growth of pMPC brushes with a length of 65 nm and a grafting density of 254. 62 µg/cm². All intermediate stages, along with the final coated surfaces (pMPC), were systematically evaluated for antifouling performance. The super‐hydrophilic CVC (∼4°) grafted with pMPC polymer brushes maintained a stable hydration layer, owing to the improved water‐binding capacity of the coated surface. Figure 10b shows reduced fibrinogen adsorption on pMPC‐coated CVC, reduced via its hydration layer; the surface also preserves protein conformation, limiting exposure of platelet‐binding sites and thereby mitigating thrombosis compared to other surfaces. In addition, the coated surfaces showed antibacterial activity against E. coli (99.77%) and S. epidermidis (99.58%). This coating, however effective, may increase processing complexity for clinical translation. In contrast, a simple approach of grafting polymer brushes mimicking the endothelium was fabricated [122] on CVC. The surface was fabricated by dip‐coating ene‐functionalized dopamine in the presence of lysine, resulting in covalent coupling and the formation of Cu²⁺ ions. Subsequently, the functionalized surface was immersed in sulfobetaine methacrylate and a photo‐initiator, leading to the formation of poly(sulfobetaine methacrylate) brushes embedded with cupric ions. The grafted brushes led to hydrophilic modification, with a contact angle of ∼19°, compared to ∼63° for PDA/Lys/Cu (the intermediate step before grafting). A similar trend in surface roughness was observed: it increased with functionalized CVC (PDA/Lys/Cu), then decreased after the growth of poly(sulfobetaine methacrylate) (pSB) polymer brushes. The antibacterial activity against both S. aureus and E. coli was 99.9% and stable for 3 days. Figure 10c shows near complete elimination (99%) through FTIC fibrinogen fluorescence compared to pristine surfaces. Thereby showing effective antifouling activity. The biofilm growth on the coated surface remained negligible after 14 days of incubation in bacterial culture, attributed to the synergistic antibacterial and antifouling activities of the PDA/Lys/Cu‐SB coatings. Both strategies involve polymer brushes, but pMPC focuses on passive antifouling via a stable hydration barrier and improved hemocompatibility, while pSB uses both passive and active approaches, leading to trade‐offs in stability and safety.

FIGURE 10.

(a) Stepwise surface‐functionalization approach for fabricating pMPC‐coated central venous catheter. Reproduced with permission [121]. Copyright 2024, Bioactive Materials. (b) Reduction of fibrinogen adsorption with subsequent coating and near complete inhibition on pMPC‐coated surface. Reproduced with permission [121]. Copyright 2024, Bioactive Materials. (c) FTIC fibrinogen fluorescence of PDA/Lys/Cu coating with Cu concentration of 400 and 800. Reproduced with permission [122]. Copyright 2023, Advance Healthcare Materials.

In addition to the hydrophilic approach, superhydrophobic surfaces are widely reported to show antifouling activity. Building on this approach, ZnO─Cu and ZnO nanoparticles modified with FAS‐17, a low‐surface‐energy modifier [123]. These nanoparticles were further coated onto the silicone rubber tube using PDMS as a binder and tetrahydrofuran, resulting in a uniform coating. These coatings, owing to the superhydrophobic (SH) nanoparticle coating (∼155°), showed enhanced reduction in E. coli adhesion with ∼3 log reduction for (SH‐ZnO) and >5 log₁₀ (SH‐ZnO‐Cu) in a 7‐day drip‐flow bioreactor. However, the antifouling properties evaluated by platelet adhesion reduction were not optimal (49.7% ± 4.8%). Increased antifouling properties as compared to superhydrophobic surfaces were observed in lubricant‐infused slippery surfaces (LIS). These surfaces were fabricated by infusing silicone oil into the superhydrophobic‐coated tube via capillary action throughout the entire tube (i.e., the lumen and the outer surface). All LIS‐infused coatings (∼100°) showed superior antifouling properties, with LIS‐ZnO showing the highest efficiency, achieving 51.4% ± 2.4% reduction in platelet adhesion and ∼3 log reduction in antibacterial activity. In addition, the coatings maintained good fibroblast compatibility, indicating their potential to improve the biocompatibility and safety of indwelling medical devices.

Thus, beyond hydrophilic and superhydrophobic strategies, bioinspired antiadhesive surfaces [124, 125] and slippery liquid‐infused porous surfaces (SLIPS) [126, 127] have gained increasing attention as practical passive antimicrobial approaches. Almousa et al. coated PVC‐based substrates with negatively charged polymer brushes [128]. The negative charge on the substrates made them more effective against S. aureus, and the brushes' hydrophilicity reduced bacterial and protein adhesion by 92% compared with uncoated PVC substrates.

Although the focus of this problem is on preventing microbial and protein attachment to control biofilm formation and device‐associated infection, another critical issue is thrombosis. The first adsorption step with plasma proteins, particularly for blood‐accessing devices (such as catheters, vascular grafts, and stents), can promote bacterial colonization of plasma proteins. This introduces the coagulation cascade, which leads to blood clotting and thrombus formation.

5.3. Antithrombogenic Surfaces

Antithrombogenic coatings aim to prevent platelet adhesion, activation, and fibrin clot formation, thereby minimizing thrombosis on dynamic blood‐contacting surfaces. Thrombosis, a pathological process of blood clotting, poses a significant challenge for biomedical implants, including vascular grafts, stents, catheters, and prosthetic valves. The underlying mechanism involves the adsorption of plasma proteins onto the implant surface, triggering the intrinsic coagulation pathway. Factor XII (FXII) undergoes autoactivation, initiating a cascade of enzymatic reactions culminating in platelet adhesion, activation, and thrombin generation, which leads to fibrin deposition and clot formation. This process is further exacerbated by leukocyte recruitment and the release of inflammatory cytokines, which promote clot stabilization and may result in vascular occlusion. Given the potentially life‐threatening consequences of thrombosis, including embolism, device failure, and infection, the development of antithrombogenic surfaces is a critical focus in biomedical engineering.

To mitigate thrombogenicity, implant surfaces must be engineered to resist protein adsorption, platelet adhesion, and inflammatory activation. Hydrophilic coatings fabricated by zwitterionic polymers create a hydration barrier that prevents non‐specific protein interactions. The poly(xylitol−sulfobetaine−cefuroxime sodium) (pXASBCS) hydrogel coating was coated on the PDMS catheter surface [129]. The hydrogel‐coated catheter showed hydrophilic modification, with contact angles of ∼11° and ∼8° for pXABCS₁ and pXABSB₃, respectively. The samples, however, showed increased values of 15.0° and 13.5° for pXABCS₁ and pXABSB₃, respectively, after 8 days, attributed to hydrophobic recovery. All hydrogels showed efficient mechanical properties and interfacial binding, with pXABSB₃ being the most promising. This coating (pXABSB₃) showed promising results both for in vitro and ex vivo rabbit model studies. The in vitro models showed a 74% reduction in thrombus formation, followed by ∼56.8% in the rabbit models (as shown in Figure 11a), indicating partial thrombus elimination owing to the hydration layer of the coated hydrogels. A recent report on sulfobetaine‐based hydrogel coated with erythromycin SB‐M₂E₇ [130] showed enhanced thrombogenic activity as compared to pXABSB₃ with occlusion rate of ∼3% (∼95.5% reduction) as in Figure 11b. Comparatively, another study showed that pMPC brushes [121] enhanced antithrombogenicity ex vivo (rabbit models). The brushes showed a ∼98% reduction in thrombosis, with thrombus weight of 2.31 ± 0.56 mg, compared to uncoated, which achieved near‐complete thrombus elimination (24.82 ± 0.78 mg) as in Figure 11c, compared to pristine polymeric implant surface.

FIGURE 11.

(a) Image comparison of pristine and poly(xylitol–sulfobetaine–cefuroxime sodium) (pXASBCS) coated catheters and associated thrombus weight. Reproduced with permission [129]. Copyright 2025, ACS Applied Materials & Interfaces. (b) Crossectional Image and SEM image comparison showing occlusion in the lumen of the zwitterionic hydrogel coating reinforced by microgels (SBM₂E₇) coated catheters post 48 h implantation. Reproduced with permission [130]. Copyright 2024, ACS Applied Polymer Materials. (c) Reduction in blood adhesion with the progression of coating on polyurethane catheters and associated thrombus weight. Reproduced with permission [122]. Copyright 2023, Advance Healthcare Materials.

Other compounds, such as vascular endothelial growth factors (VEGFs) [131] have been reported as antimicrobial agents and have a major effect on vascular grafts, markedly improving endothelialization and reducing the risk of thrombosis along the graft, ultimately prolonging its lifespan. However, in vivo studies of the same compound have shown unprecedented growth of FBGC (Foreign Body giant cells) in vivo rat models. Likewise, bone morphogenetic proteins (BMPs) [132] have been immobilized on orthopedic implants, inducing osteogenesis that facilitates bone healing and integration with prosthetic bodies. Furthermore, heparin‐based coatings [133, 134, 135, 136] are the most prevalent, spanning commonly used blood‐contacting biomedical devices. However, associated complications, such as heparin‐induced thrombocytopenia (HIT) [137], call for an alternative approach.

Heparin is another commonly used coating for achieving antithrombogenic surfaces. PVC tubes used for extracorporeal circulation are commercially coated with heparin. However, due to complications associated with the coating alternative, nanoparticle‐based coatings are being considered. In this reported approach, Sun et al. coated PVC with a silicon oxide‐stearic acid mixture, yielding a superhydrophobic (∼145°) surface [138]. The fabricated surface extended clotting time by 73% compared to a commercial heparin‐based coating. Additionally, fibrinogen adsorption was reduced to 67%, demonstrating efficient antifouling activity attributed to the physical barrier mechanism. Another such study of replacing e‐PTFE membranes used as vascular grafts. However, their synthesis requires higher temperatures and greater energy. Thus, Wang et al. [139] proposed fabricating tubes by electrospinning with PEO and PVA as spinning aids, which were subsequently sintered to yield pure PTFE tubes. These tubes were further activated to be coated with heparin to make them bioavailable and antithrombotic. The electrospun PTFE tubes were plasma‐treated, yielding surface ─COOH and ─OH functional groups, which allowed polydopamine to bind to the surface oxygen moieties. The ─NH groups on polydopamine further promoted the adhesion of heparin to the PTFE surface. This coated surface exhibited all the properties of e‐PTFE grafting, including a similar hemolytic rate and bioavailability, and a superior antithrombotic surface compared with the original e‐PTFE [139]. In addition, blending base polymers to achieve the desired surface interface is also an efficient alternative. Choi et al. [140] incorporated novel surface‐migrating oligomers with fluorine backbones into polyvinyl chloride films. The incorporation of these oligomers conferred antithrombogenicity to the surface, a desirable property for blood‐contacting medical devices [141]. The modified devices showed less fibroblast adhesion as compared to pristine PVC. The modified PVC surface was biocompatible and had less protein adsorption.

Despite these advancements, challenges remain in ensuring long‐term durability, scalability, and in vivo stability of these coatings. Future research is directed toward stimuli‐responsive antithrombotic surfaces [134, 142, 143] that dynamically adapt to changes in the vascular microenvironment, thereby offering personalized, adaptive strategies for thrombosis prevention in biomedical implants.

6. Conclusions and Future Perspectives

Advances in surface chemistry, biomaterials engineering, and nanotechnology have enabled significant progress in the development of biofunctional coatings for polymeric implants. Surface modification techniques have enabled inert polymers to acquire biocompatible surfaces for biomedical applications. By altering the surface chemistry and the biological‐material interface, coatings can reduce bacterial adhesion, enhance antifouling performance, improve hemocompatibility, or promote favorable interactions between the host and the tissue. In this review, we examined coating strategies utilized to enhance the biological performance of polymeric implants, emphasizing methods tailored to meet specific functional requirements. Some of them are antibacterial coatings that prevent infections around implants, antifouling surfaces that prevent proteins from adhering to the polymers and microbes from growing on them, and hemocompatible coatings that reduce the risk of blood clots. There are many different types of material platforms, such as polymer grafting, bioinspired adhesive layers, biomolecule immobilization, and nanostructured coatings. These demonstrate how rational surface engineering can be used to tailor the interface between polymeric substrates and biological environments.

A major limitation of most current coating approaches is their reliance on multistep synthesis and solvent‐intensive fabrication processes, which hinder scalability and reproducibility in industrial settings. Future progress will depend on establishing clear structure–property–performance relationships linking coating architecture and surface chemistry to long‐term biological outcomes. The regulatory and biocompatibility of coatings are central in clinical translation. In addition, long‐term toxicity, immune response, and assessment of breakdown products remain critical for nanoparticles, antimicrobial agents, and coatings based on complex hybrid materials. The stability of such coatings in physiological conditions, where they are subjected to mechanical stress, enzyme activity, and constant fluid flow, remains a challange. Thus, uture efforts with polymeric implant coatings will probably center on more robust surface‐anchoring technologies with simpler manufacturing techniques and greener materials. These advancements are excpected to ensure long‐term durability, scalability, and in vivo stability of coatings. Another promising avenue for future research is directed toward stimuli‐responsive antithrombotic surfaces that dynamically adapt to changes in the vascular microenvironment, thereby offering personalized, adaptive thrombosis‐prevention strategies for biomedical implants.

Conflicts of Interest

The authors declare no conflicts of interest.

Biographies

Ritika Tripathi received her master's degree in nanotechnology from the National Center for Nanoscience and Nanotechnology, University of Mumbai, in 2020. She is currently pursuing her Ph.D. under the supervision of Dr. Suchitra and the co‐supervision of Prof. Reddy. Her research focuses on the surface modification of biomedical implants using biopolymers and nanomaterials.

M. Sudhakara Reddy received his Ph.D. from the University of Madras in 1994. He has been associated with the Thapar Institute of Engineering and Technology since 2000. Over the duration he has supervised several Ph.D. and M.Tech. scholars. His research interests include microbiology, molecular biology, biotechnology, metatranscriptomics, and plantmicrobe interactions.

Suchitra Rajput received her Ph.D from the Indian Institute of Technology in 2008, working on MgB₂ superconducting tapes during her PhD and focusing on wires during her postdoctoral research. She has been associated with BML Munjal University since 2017. She specializes in physics and has been educating graduating engineers for the last 15 years, and uses her tireless energy to encourage research scholars to work hard and succeed. She has published several journal papers and convened conferences. Her research interests include coatings, water treatment, and biomedical implants.