Advances of chemistry for the design of antimicrobial biomaterials

Since the dawn of civilization, humankind has been engaged in a continuous struggle against infectious diseases. Even before the beginning of our era, the therapeutic properties of copper, silver, gold, and a number of natural substances were empirically discovered and successfully used for treatment and mummification. The use of copper dates back to between fifth and six thousand years before Christ. Copper could be found as the free metal inside nuggets, which did not require smelting, and later it began to be used to produce brass and bronze. The first medicinal application of copper is mentioned in the Smith Papyrus (year 1500 BC), an ancient medical text,[1] which partly reproduces the text written between the year 2600 and the year 2200 BC and describes the use of copper for sterilization of chest wounds and drinking water. Copper or its compounds were used to treat diseases such as headaches, itching, burns, intestinal worms, and ear infections as well as for general hygiene.[2]

Silver metal was already known to the Chaldeans around 1000 years BC. For thousands of years, silver was used to treat many diseases, even before it was found that infections are caused by microbes.[3] Herodotus (484 – 425 BC) deemed that Kings of Persia always drank water that was stored and transported only in silver vessels. The ancient Phoenicians, Greeks, Romans, and Egyptians stored water in silver vessels during long military campaigns.[4] An early attempt to treat wound infections with silver plates is attributed to ancient Macedonians. Hippocrates (460 – 370 BC) used silver dressings to treat ulcers and to accelerate wound healing. There is a hypothesis about the use of silver nitrate in medicine, as it was mentioned in a pharmacopoeia published in Rome in the year 69 BC, and the first mention of the use of silver nitrate for medical purposes dates back to 702 – 705 AD. Avicenna (980 – 1037 AD) used silver additives for a blood cleanser and to prevent palpitation and eliminate bad breath.[3] A number of discoveries in 1847 – 1880 indicating that infections are caused by microorganisms (works of Ignaz Semmelweis, John Snow, Casimir-Joseph Davaine, Louis Pasteur, and Robert Koch)[5] provided a rational base for medicinal use of silver. For example, in the 1880s, the German obstetrician Karl Crede employed dilute solutions of silver nitrate to treat eye infections in newborns. In 1889, William Halstead, the head of the department of surgery in the Johns Hopkins hospital, used silver foil to treat postoperative wound infections.[6] Colloidal silver solutions formed the basis of antimicrobial therapy in the first half of the 20th century, before the advent of antibiotics in the early 1940s. According to a study of the National Aeronautics and Space Administration (NASA) carried out in the mid-1960s, a silver ion concentration of 400 ppb is sufficient for water disinfection.[7] The Soviet spacecrafts Salyut and Mir used ionic silver at a concentration of 0.2 mg L^–1 (200 ppb) for water treatment.[8]

The medicinal use of gold as a therapeutic agent began in China around 2500 BC. Pure gold was used to treat furuncles, measles, pox, and skin ulcers. Some ancient sources mention the application of gold-containing products for the treatment of joint and lung diseases.[9] Archaeologists are rarely able to discover ancient medicines and, the more so, to determine their chemical composition; therefore, most of information has been derived from ancient manuscripts written by Theophrastus (371 – 286 BC), Pliny the Elder (1st century BC), and Pedanius Dioscorides (1st century BC). In this regard, of considerable interest are the results of analysis of a tin pyxis (vessel) containing residual medicine that was found on board the Pozzino shipwreck, presumably dating back to 140 – 130 BC, which showed the presence of zinc compounds.[10]

Many natural and plant-based products have been used in Ayurvedic medicine since around 5000 BC. Ancient healers prepared plant- and animal-based medicines, which were sometimes mixed with soil, minerals, and metals.[11] Over time, herbal, mineral, and metal agents started to occupy an important place in the Ayurvedic pharmacopoeia and have been regularly used in practice in various parts of India for many centuries. The ancient Egyptians also produced various therapeutic means from plants, animal products, and minerals. Detailed information about ancient Egyptian medicine can be found in a review.[12] Modern scientists admire the art of mummification, which used pine resin, myrrh, bee wax, bitumen, onions, lichen, henna, and other natural ingredients.

Despite the fact that the first antibiotic, salvarsan, was used in 1910, and Alexander Fleming discovered penicillin only in 1928, there are numerous examples of antimicrobial drugs entering the bodies of ancient people.[13] For example, traces of tetracycline were found in human skeletal remains in ancient Sudanese Nubia, dating to the years AD 350 – 550 (Ref. [14][15]) and in thigh bones of the skeletons of the late Roman period from the Dakhla oasis, Egypt.[16] It is assumed that tetracycline entered the body with food and had a protective effect, since the level of infectious diseases in Sudanese Nubia was low, and no traces of bone infections were found in Dakhla. Antibiotics produced by microbes have been used to prevent diseases for thousands of years. For example, poultices made from mouldy bread were used to treat open wounds in Serbia, China, Greece, and Egypt more than 2000 years ago.[17] Mouldy bread and therapeutic soil are mentioned in the Ebers Papyrus dated 1550 BC, which is the oldest surviving medical document.[18] The remedy reconstructed from a recently discovered 1000-year-old Anglo – Saxon prescription has proved to be effective against methicillin-resistant Staphylococcus aureus (S. aureus).[19] Despite high expectations and the initial high efficacy of antibiotic therapy, especially during World War II, the golden age of antibiotics (1940 – 1962) turned out to be quite short. Many modern antibiotics are modified versions of previous ones, and today scientists openly speak about a crisis in this field. First of all, this is due to slow introduction of new types of antibiotics into the clinical practice and to the emergence of multidrug resistance (MDR) in pathogens. These and many other problems are described in detail in the book How to Overcome the Antibiotic Crisis.[20] Global estimates of the MDR problem based on systematic data analysis were first published in the Lancet journal in 2022.[21] The six major pathogens, Escherichia coli (E. coli), S. aureus, Klebsiella pneumoniae (K. pneumoniae), Streptococcus pneumoniae (S. pneumoniae), Acinetobacter baumannii (A. baumannii), and Pseudomonas aeruginosa (P. aeruginosa) were the cause for 3.57 million MDR-related deaths in 2019. Clinical monitoring data show that the level of antimicrobial resistance of bacteria is alarmingly high, being 51% to penicillin, while the median level of E. coli resistance to third-generation cephalosporins is 36%. Moreover, E. coli resistance to ciprofloxacin varies from 8.4% to 92.9%, while that for K. pneumoniae varies from 4.1 to 79.4%.[22]

In addition to the MDR problem, a new challenge comes from nosocomial fungal infections, which pose an increased threat to patients with weakened immune systems or those who have had COVID-19. More than 150 million people suffer from severe fungal infections, which cause more than one million deaths annually. Long-term therapy with antifungal drugs, for example, in cancer patients, has led to the emergence of drug-resistant fungi such a Candida auris (C. auris).[23]

Currently, the design of bactericidal surfaces is a hot topic, with over 23 000 papers published on the subject in 2025 (search for the keywords ‘antibacterial activity’ in the Scopus database). The main feature of this review is integrated approach to the problem of fight against pathogenic microflora ranging from the historical context and setting of tasks for modern medical materials and products to fundamental mechanisms and advanced developments.

The global technological and demographic changes in the world such as climate change, environmental pollution, and population growth, increase in the polulation density, and active migration increase the risk of infectious diseases.[24] Materials possessing bactericidal and fungicidal properties are today in-demand for many medical and civilian products such as orthopaedic and dental implants, urological catheters, wound dressings, breathing masks, filters for artificial lung ventilation (ALV), bactericidal fabrics for sportswear and medical personnel protection, and food packaging.

Infection on the implant surface may be caused by previous injury, surgical intervention, or the bacterial entry and spread throughout the body during the installation of catheters or application of dressing materials, the use of ALV devices, or the patient stay in the hospital. In a healthy body, these bacteria carried by the blood are cleared by the host immune system before they reach vulnerable areas. To minimize infection, antibiotics are prescribed, most often, for 2 – 14 days after the surgery, together with an additional oral preventive treatment. If infection is present, it is often necessary to supplement antibiotic therapy with implant removal and elimination of the infected bone.[25][26] All this gives rise to side effects, additional medical costs, and long-term rehabilitation for the patient.

Infections are usually accompanied by biofilm formation on the implant surface (Fig. 1). Bacteria can exist in the planktonic form or as a biofilm, which is defined as a structured community of bacterial cells enclosed in a self-produced polymer matrix attached to the surface.[27] However, it is noteworthy than not all biofilms develop in this way, especially under clinical conditions in vivo, where biofilms often exist as aggregates not attached to the surface.[28]

Fig. 1

Schematic view of biofilm formation (a) and SEM images of single cells and colonies of E. coli (b) and S. aureus (c). SEM images were taken from the personal archive of the author.

The biofilm formation includes several stages: first, planktonic bacteria are attached to the surface, which may be facilitated by proteins absorbed from the environment, and then they form microcolonies, which gradually increase in the mass to form extracellular matrix (see Fig. 1).[29] The planktonic cells are attached to the substrate via adhesins such as fibronectin-binding proteins. The factors that play the key role in the biofilm formation include the biofilm-associated protein, intercellular polysaccharide adhesin, and some other extracellular polymeric compounds such as extracellular DNA. The implant-associated biofilms formed by Gram-positive S. aureus and Staphylococcus epidermidis (S. epidermidis) strains and Gram-negative Enterobacter, Acinetobacter, Klebsiella, and Pseudomonas strains pose a considerable problem throughout the world and account for the increasing rate of failures of implanted devices.[30][31] Bacterial biofilms are much more resistant to antibiotics than planktonic bacteria, as they are protected by the extracellular matrix consisting of polysaccharides, proteins, and nucleic acids.[32] This is due to the fact that attachment of bacteria to the surface leads to rapid changes in the expression of certain genes responsible for the production and maturation of exopolysaccharides and mucus and activation of the protective barrier against the endogenous defence system of the body and external agents such as antibiotics.[33] The antibiotic resistance of biofilms may be due to a delay or absence of antibiotic penetration or to operation of efflux pumps within bacteria that remove antibiotics from the maturing biofilms into the extracellular matrix. Bacteria use signalling molecules (autoinducers) of the quorum sensing (QS) system to exchange information and coordinate various activities within the biofilm; this provides for their protection from the host defence factors and antibacterial agents.[34] However, the current understanding of the processes occurring in biofilms is rather limited. To fill this gap, it was recently proposed to use highly efficient imaging and mapping of these structures based on automated scanning electron microscopy (SEM) and an integrated software system that uses neural networks to perform in-depth analysis of biofilms.[35] The main bacterial pathogens involved in biofilm formation on the surface of medical implants are listed in Table 1[36].

Table 1

\[ \]

Main bacterial pathogens involved in the biofilm formation on the surface of medical implants.³⁶

(1)

According to predictions, the market of metal orthopaedic implants and medical alloys would nearly double between 2022 and 2030 to reach $30 billion.[37] This is due to increased life expectancy and population ageing [according to the World Health Organization (WHO) data, one in six people on the Earth will be over 60 years old by 2030], development of sports and popularization of extreme sports, and numerous armed conflicts. The success of implantation largely depends on how completely the emergence and growth of pathogenic microflora is prevented in the early stages.

Two-thirds of infections associated with implanted devices are caused by S. aureus strains. Titanium and its alloys are most commonly used for implants that contact with bone tissue.[38] Numerous methods for surface modification of titanium and endowing titanium surface with bactericidal properties are discussed in detail in a book by Sánchez-Bodón et al.[39] Currently, biodegradable magnesium alloys are considered to be a promising alternative to titanium alloys, despite problems that have not yet been fully solved, such as fast corrosion, hydrogen formation, and the need to excrete harmful magnesium ions; therefore, extensive studies aimed at increasing the antibacterial properties of magnesium are underway.[40]

The WHO considers wound infections, which affect 1 – 2% of population in developed countries, to be a serious and relevant problem.[41] In the US, infections associated with implants account for approximately 26% of all infections related to healthcare.[42] The cost of treatment of wound infections amounts to approximately 3% of total healthcare costs and, according to predictions, it would reach $18.7 billion by 2027.[43] The biological tissue regeneration process consists of several overlapping stages: hemostasis, inflammation, proliferation, and maturation, which include various cells and endogenous factors.[44] Medical dressings maintain a moist medium for healing, protect the wound surface from bacterial infection, and promote debridement, epithelialization, and restoration of damaged skin areas.[45] Polymer-based hydrogels saturated with medicinal agents or containing bactericidal nanoparticles are considered to be promising for would healing, as they absorb the wound exudate and activate immune cells. In order to extend the scope of applicability of hydrogels, it is necessary to find new methods for chemical and physical cross-linking (click chemistry, enzymatic reactions, crystallization, amphiphilic block copolymers, etc.), endowing materials with bactericidal characteristics and enhancing their mechanical properties and adhesion.[44]

Catheter-associated urinary tract infection (UTI) is a highly widespread infection transmitted in healthcare facilities,[46] which gives rise to secondary bloodstream infections. The duration of catheterization plays a key role in the development of bacteriuria, since the catheter surface is prone to biofilm formation. Antibiotic treatment becomes less effective due to MDR bacteria. The hospital-acquired infections related to catheterization are often caused by antibiotic-resistant Staphylococcus and Enterococcus strains and Candida fungus. After insertion of urethral stents/catheters, infection arises in 10 – 50% of patients within 7 days.[47] Urethral catheters are usually made of polytetrafluoroethylene (Teflon®), polyurethane, polyvinyl chloride, or polyethylene. To modify the surface of inert polymeric items, surface activation is needed; surface properties such as chemical composition and morphology are crucial for cell adhesion and biofilm formation. The best way to prevent the growth and development of pathogenic microflora is to inhibit the initial stage of bacterial attachment and hinder the biofilm formation, which requires the fabrication of bactericidal surfaces. There are three main approaches to surface modification of orthopaedic and dental implants for endowing them with antibacterial properties: (i) change in morphology, roughness, and wettability, (ii) attachment of antibacterial or antifouling agents (antibiotics, antifungal agents, antimicrobial peptides, anti-adhesive molecules, nitroxides, etc.), and (iii) deposition of bactericidal coatings.

Hospital patients on ventilators are most vulnerable to bacterial/viral infections, because the contours of the ALV apparatus can be contaminated by pathogenic microflora from the respiratory passages and the environment. The most frequent complication is ventilator-associated pneumonia, a serious hospital-acquired infection that occurs 48 h after endotracheal intubation and is difficult to diagnose and treat.[48] The most common pathogens are Gram-positive S. aureus (28.4%) and Gram-negative P. aeruginosa (25.2%) and other bacteria (26.6%).[49] Modern filter membranes for ventilators mainly operate on the principles of mechanical and electrostatic filtration, which does not provide the required antibacterial protection.

Cotton fabrics ate widely used in healthcare, the sports industry, and everyday life, but they are prone to microbial growth and, hence, they can carry bacteria and viruses. Cotton fabrics can be endowed with antibacterial properties by immobilization of antimicrobial agents on the fabric surface.[50] In recent years, this area has actively developed owing to the success of nanoindustry in the production of nanoparticles and nanofibres of bactericidal metals and introduction of these species into fabrics. However, immobilization of nanomaterials is a challenge due to the absence of a strong chemical bond with textile surfaces.[51] It is necessary to introduce a cross-linking agent, in order to improve the ability of a cotton fabric to hold nanoparticles. In addition, the bactericidal properties of textiles are deteriorated after prolonged use as the additives are gradually lost during washing. In a new, recently developed method for copper ion impregnation into a cellulose matrix, copper ions are strongly coordinated to oxygen-containing polar functional groups (e.g., hydroxyl groups) of the cellulose chains.[52]

There are several main mechanisms for suppressing the bacterial or fungal infection: (1) treatment with bactericidal ions, therapeutic agents, viruses and/or amino acid residues (antibiotics, antimycotics, bacteriophages, antimicrobial peptides), (2) formation of reactive oxygen species (ROS), which damage the bacterial wall and cause the oxidative stress in bacteria, (3) bacterial wall damage on contact with the implant surface via mechanical damage or microgalvanic interactions, (4) prevention of bacterial adhesion via formation of an anti-adhesive surface, (5) suppression of the QS system in bacterial films by using bioactive molecules, and (6) surface response to changes in specific environmental factors (Scheme 1).[53-56]

Scheme 1

Correspondingly, there are a few strategies for the fabrication of bactericidal surfaces: (i) doping with bactericidal elements (Ag, Cu, Zn, etc.),[57][58] (ii) creation of heterogeneous surfaces decorated with bactericidal metal or oxide nanoparticles (NPs),[59-61] (iii) surface modification with various antibiotics, antimycotics, antibacterial peptides, N⁴⁺ and P⁴⁺ compounds, anticoagulants, or nitroxides via covalent or electrostatic bonds,[61-63] and (iv) generation of a specific surface texture or roughness.[53][64-66] The problems of implant-associated infections and various strategies directed towards growth inhibition of bacterial colonies have been described in detail in the literature.[67][68]

The most popular method for endowing a surface with antibacterial properties is the introduction of bactericidal metals, among which Ag, Cu, Au, and Zn are most important and studied in most detail. A benefit of the approach based on the release of bactericidal ions such as Ag,[69-71] Cu,[72][73] Au,[74][75] and Zn[76][77] is that the action is local, being concentrated directly near the substrate surface. However, the mechanisms of their action on pathogens are considerably different.[27] The Ag⁺ ions have the widest range of action: they damage the bacterial cell membrane and cytoplasmic component, block the transfer of oxygen into bacterial cells, inactivate functions of enzymes, and disrupt the DNA replication.[53] The damage of the E. coli outer membrane and cytoplasmic membrane by Ag⁺ ions was demonstrated using energy-filtering transmission electron microscopy.[78] An important component of the bacterial cell wall is the heteropolymer peptidoglycan (PG), which acts as a target for Ag⁺ ions. It was shown that the Ag⁺ ions inhibit the PG elongation in S. aureus cells by enhancing the activity of autolysins (amidases and hydrolases). Bacteriolysis and destruction of E. coli cell wall by Ag⁺ ions is caused by disruption of lipopolysaccharide synthesis and breaking of the outer membrane due to degradation of lipoprotein at the C- and N-termini.[79] The DNA damage is due to the formation of two-coordinate Ag⁺ complexes within triple (guanine)≡(cytosine) and double (adenine)=(thymine) hydrogen bonds in DNA base pairs.

The bactericidal action of Cu²⁺ ions is based on their ability to penetrate through the cell membrane and generate ROS, thus inducing oxidative stress. This results in the bacterial cell damage, influx of additional Cu²⁺ ions, and destruction of DNA molecules and bacterial enzymes.[80] The antibacterial activity of Zn²⁺ ions is usually attributed to the generation of ROS, damage of cytoplasmic membrane, which disrupts the enzymatic system, and formation of surface hydroxyl groups, which prevents bacterial adhesion.[81][82] The Zn²⁺ ions can also be absorbed by E. coli cells, being thus converted to insoluble ZnO-based nanocomposites, which damage the cell membrane.[83] The excellent antibacterial activity against Gram-positive and Gram-negative bacterial strains and prevention of biofilm formation are attributable to the synergistic effect of Zn²⁺ ions and ROS.[84]

Gold NPs and nanoclusters are most often used in the antibacterial therapy as conjugates with various surface ligands (amino acids, peptides, antibiotics, antibodies, enzymes, DNA, etc.).[85] Using in situ liquid-cell transmission electron microscopy, it was shown that Au NPs surface-decorated with glutathione ligand were first attached to bacterial membrane Acetobacter aceti through physical adsorption and then internalized. The ROS generation caused destruction of the bacterial membrane and induced cell death.[86]

The antimicrobial activity of Ag⁺, Cu²⁺, and Zn²⁺ ions against various strains of bacteria and fungi was confirmed in numerous studies (see, e.g., review[87]). However, the data on the effective concentrations of bactericidal ions are highly scattered (Table 2[88-91]): 1 × 10^–4 – 5 ppm (Ag⁺), 0.15 – 450 ppm (Cu²⁺), and 0.015 – 770 ppm (Zn²⁺). It was reported that the ion concentrations of 0.25 – 1 ppm (Ag⁺) and 65 – 650 ppm (Zn²⁺ and Cu²⁺) are effective against S. aureus and E. coli strains, while exhibiting no cytotoxicity against fibroblast cells.[80] The minimum inhibitory concentrations (MICs) of metal ions released from chitosan nanoparticles saturated with ions against E. coli and S. aureus strains, respectively, were 4 and 8 ppm (Ag⁺), 256 and 448 ppm (Cu²⁺), and 768 ppm (Zn²⁺).[92] In relation to P. aeruginosa, MIC was 62.5 ppm (Zn²⁺), 12.5 – 125 ppm (Cu²⁺), and 104 ppm (Ag⁺).[93] One explanation to this large scatter of data is the difference between methods and conditions of bactericidal assays.[94] The types of used strains (pathogenic or non-pathogenic, collection or clinically isolated, MDR or non-MDR), the initial concentrations of colony-forming units (CFU), which are usually in the 10⁴ – 10⁸ CFU mL^–1 range, and the growth medium are different. Since bacteria vary in shape, size, cell wall thickness, outer membrane composition, and many other characteristics, the causes for their death can also be different. Some strains exhibit increased resistance to bactericidal ions due to their ability to regulate their cellular uptake.[95] In experiments using physiological saline solution (PSS), the initial cell concentration in the control either remains unchanged or decreases with time, while the culture medium usually promotes cell proliferation.

Table 2

\[ \]

Dependence of the antibacterial effect against E. coli and S. aureus on the type and concentration of ions.

(2)

A key challenge in the use of bactericidal ion release strategy is to provide a controlled release of ions. It should be borne in mind that the rate of ion release does not always depend on the concentration of the bactericidal element, and there are additional factors that affect the yield of bactericidal ions, such as the state of metal agglomeration, surface roughness, and surface oxidation kinetics.[53] It was shown that the surface roughness has a more pronounced effect on the release of Ag⁺ ions than the silver content in the surface. By appropriate adjustment of the surface roughness, it is possible to considerably increase the release of silver compared to that for smooth surface. For example, after 24 h, the amounts of Ag⁺ ions released from the multielement TiCaPCON-Ag coating deposited on the smooth and rough Ti surfaces by magnetron sputtering of the TiC_0.5 – Ca₃(PO₄)₂ target in the Ar – 15% N₂ gas mixture differed by a factor of three. The effect of surface topography on Ag⁺ release was demonstrated by Ponomarev et al.[96] Periodically spaced pillars of 10 and 50 μm diameter and 3, 1, and 0.2 μm height were formed by photolithography on a silver-coated silicon surface. The samples with the greatest pillar height showed the fastest release of Ag⁺ ions due to greater specific surface area and the presence of Ag NPs located between the pillars.

An equally important factor influencing the yield of ions is surface oxidation in a biological medium, with the oxidation rate being directly related to the surface roughness. As a rule, the higher the roughness, the greater the specific surface area and the faster the oxidation (see Section 4). An additional contribution is made by the concentration of bactericidal nanoparticles on the surface: the higher the concentration, the slower the surface oxidation. After 7 days, the amount of Ag⁺ ions released from the surface of the TiCaPCO – Ag coating deposited in an argon medium that contained 27 at.% oxygen was almost three times higher than that from the surface containing 2 at.% oxygen that was deposited in water.[97] All these factors must be taken into account for temporal control of bactericidal ion release.

As noted in some studies[53][98] and was demonstrated above, there is no direct correlation between leaching of bactericidal ions and antibacterial properties. Meanwhile, the cyto­compatibility of materials depends on the dose of the bactericide; therefore, the concentration of bactericidal ions should be thoroughly controlled and compared with the results of cytocompatibility assays.[99]

The bactericidal mechanism based on the action of ions is shown in Scheme 2 and Scheme 3. Bacterial cells are surrounded by an outer membrane (only in the case of Gram-negative bacteria) containing lipopolysaccharides, lipids, and proteins, cell wall the main component of which is PG (same as murein), and an inner cytoplasmic membrane. The thick PG layer of Gram-positive bacteria decreases the probability of penetration of bactericidal ions. The interaction of ions with bacterial cells often starts with electrostatic attraction between positively charged ions and the negatively charged bacterial surface (Scheme 2). The Ag⁺, Cu²⁺, Zn²⁺, and Mg²⁺ ions can directly interact with various structural components of the cell wall and membrane such as lipids and amino acids occurring as parts of proteins and PG and change their structure, which leads to damage and denaturation (Scheme 3). In the case of Gram-negative bacteria, ions can penetrate the outer membrane through beta-barrel proteins (porins) or membrane transport proteins. When bactericidal ions get into the cell, they deactivate respiratory enzymes and disrupt the production of adenosine triphosphate (ATP), which provides energy for various biochemical reactions. The reactions of Ag⁺ with sulfur and phosphorus in DNA and the action of ROS on DNA are main causes for the disruption of DNA replication. Silver ions can inhibit protein synthesis, which results in ribosome denaturation in the cytoplasm.[100] Copper ions react with proteins depending on the oxidation state: Cu(I) binds to sulfur donors such as cysteine and methionine, whereas Cu(II) prefers nitrogen donors such as histidine and other amino groups.[101]

Scheme 2

Scheme 3

In conclusion, it should be noted that the release and transport of bactericidal ions are influenced by the biological medium. The implant surface actively interacts with blood serum proteins such as human serum albumin (HSA), which is the main zinc(II) transporter in blood plasma.[102] The reactions of Ag⁺ with extracellular thiols and human blood components prevent the interaction of silver ions with the cells, which reduces the antimicrobial activity. Binding to silver affects the protein structure, thus inducing loss of alpha-helices and a change in the tertiary conformation.[103]

The strategy based on the release of pharmaceutical agents [antibiotics,[61][104][105] quaternary ammonium compounds,[106] halogen compounds (fluorides[107] and iodides[108]), chlorhexidine,[109] nitric oxide,[110] chitosan,[111] etc.) also implies the local action just near the substrate surface, which considerably reduces the therapeutic load. The main benefit of the surface-bound antibiotics and antimycotics is that they suppress the infection over a short period of only a few hours even in low doses[61][63] and are non-toxic for the surrounding tissues and the whole body. In order to achieve a wider range of action, mixtures of medications can be used.[63] The major drawbacks are the emergence of MDR (and related decrease in the efficacy of antibiotics), limited time of the bactericidal effect due to complete release of the medication over a short period, and undesirable side effects at high therapeutic doses.

A promising approach is to combine the strategies based on the release of antibiotics and bactericidal ions, which follow different kinetics. For example, Ag-containing TiCaPCON coating provided a persistent release of Ag⁺ ions for 7 days, while gentamicin (GM) loaded into the specially formed surface relief was released almost completely within 12 h, with accelerated release being observed during the first hour.[63] The possible interactions of therapeutic components should also be taken into account. For example, GM can interact with silver ions and form stable Ag/GM complexes, thus decreasing the concentration of free silver ions in the test solution.[112] The accelerated initial release of the medication in a low dose (in order to avoid side effects) provides fast suppression of the infection, while prolonged release of bactericidal ions provides long-term protection to prevent latent infection.

A wider range of action can also be achieved by combining various antibiotics, antimycotics, and bactericidal ions. For example, titanium hybrid bioconstructs containing Ag NPs and loaded with a mixture of GM and amphotericin B (ATB) were equally effective against the Gram-negative E. coli bacterium and GM-resistant Neurospora crassa (N. crassa) fungus.[63] This type of fungus is characterized by strong cell walls that act as a barrier against bactericidal Ag⁺ ions. The GM- (150 and 300 μg cm^–2) and ATB-loaded (100 μg cm^–2) hexagonal boron nitride h-BN coatings also effectively inhibited the growth of E. coli K-261 and N. crassa strains.[55]

The strategies that imply immobilization of therapeutic agents are based on the generation of a surface rich in functional groups to provide a strong adhesion to the substrate surface via the formation of covalent bonds (Scheme 4). The titanium surface is activated using carbodiimide chemistry and click-chemistry methods, surface functionalization using thiols, plasma amination (introduction of NH₂ groups), silanization, surface photopolymerization, etc.[113]

Scheme 4

The surface of titanium with the deposited TiCaPCON coating (Ag-loaded or Ag-free) was treated with the Ar/C₂H₄/CO₂ radio frequency plasma at a low-pressure of 2 – 4 Pa in order to form a polymer layer containing COOH groups (Scheme 5).[114] The generation of surface carboxyl groups is commonly used for immobilization of various biomolecules on plasma-treated materials.[115] For the covalent immobilization of GM, the sample was first placed into an aqueous solution of 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC), which is often used a carboxyl-activating reagent, and then into a 2% aqueous solution of GM for chemical binding to EDC through amide bonds. Anti-adhesive heparin (HP) molecules were immobilized on the GM surface via ionic bonds formed between the positively charged GM amino groups and the negatively charged COOH groups of HP.[114] After 24 h, GM- and GM/HP-loaded samples, as well as the sample with Ag NPs showed 99.999% bactericidal effect against antibiotic-resistant E. coli K261 strain.

Scheme 5

Low-pressure plasma polymerization and carbodiimide chemistry methods were used to immobilize four different therapeutic agents (antibiotic, GM; antimicrobial peptide, indolicidin; anti-adhesive HP; and nitroxide, 2,2,5,5-tetramethyl-3-carboxyl-pyrrolidin-1-oxyl) on the surface of TiCaPCON coatings via ionic or covalent bond to endow the coatings with antibacterial properties.[116] The use of modified coatings markedly decreased the level of antibiotic-sensitive E. coli U20 strain and (except for samples with immobilized indolicidin) induced effective inhibition of the antibiotic-resistant E. coli K261 strain at an initial concentration of 10⁴ CFU mL^–1. The coatings bearing covalently attached nitroxide not only exhibited the highest antibacterial activity against E. coli K261 cells (99.99% after 8 h), but also prevented the biofilm formation.

The vancomycin (VM) antibiotic was encapsulated into a polyethylene glycol (PEG)-based hydrogel film that was bound to titanium implants by a covalent bond and then coated with a membrane made of PEG copolymer with the product of polycondensation of lactic acid and caprolactam.[117] In order to retard the drug release, the hydrogel was used as a mixture with cross-linked starch, the porous microstructure of which inhibited the hydrogel swelling. The materials demonstrated stable release of VM for almost 3 weeks in vitro and for more than 4 weeks in vivo. In the S. aureus infection model in rabbits, implants loaded with 4 mg of VM substantially reduced the imflammatory response and showed good antimicrobial activity.

The surface of titanium implants was functionalized with 3-aminopropyltriethoxysilane followed by attachment of enoxacin (EX)-linked PEG molecules via amino groups (Scheme 6).[118] The EX-bearing titanium surface effectively prevented the bacterial colonization with methicillin-resistant S. aureus (MRSA) without deteriorating in vitro cell viability, adhesion, or proliferation.

Scheme 6

GM-loaded biodegradable coatings based on polylactic and glycolic acid polymers or poly(D,L-lactide) were deposited on the surface of titanium implants, which markedly decreased the frequency of infection and improved the recovery rate after infection compared to systemic GM application.[119-121]

Implant coating by an amphiphilic polymeric material was performed by a new technique involving surface binding of a branched block copolymer of PEG with polyallylmercaptan (PAM) obtained in situ,[122] which was effectively chemisorbed on metal surfaces and self-organized to give a uniform coating. The titanium Kirschner wires were immersed into a solution of PAM, PEG thiol (6 mass %), and vancomycin (20 mg mL^–1). Then the wet titanium wires were UV-irradiated and dried at room temperature to form PEG-PAM on the sample surface.

Vancomycin and caspofungin were covalently attached to titanium surface pretreated with 3-aminopropyltriethoxysilane. The functionalized titanium samples were placed into a solution (1 mL per disc) of n-heptane/hexamethylene diisocyanate. The in vitro formation of S. aureus bacterial biofilms and Candida albicans (C. albicans) fungal pathogen was substantially reduced compared to that in the titanium control sample. In in vivo assays, the formation of S. aureus biofilms in VM-loaded samples decreased by more than 99.9% and the formation of C. albicans biofilms in caspofungin-loaded samples decreased by 89% compared to the control Ti sample.[123]

To suppress local infection, titanium implants were coated with chitosan-based hydrogels loaded with antibiotics such as ciprofloxacin, amoxicillin, and gentamicin[124] or their mixture. After 24-hour subcutaneous implantation of titanium implants coated with antibiotic-loaded chitosan particles in male Sprague–Dawley rats, analysis of SEM images showed that the samples contained less S. aureus colonies than the control sample.[125]

Antimicrobial peptides (AMPs) attached to the surface of biomaterials represent a promising alternative to traditional antibiotics for fight against infections.[126] АМPs are molecules formed by amino acid residues connected by peptide (amide) bonds; they exhibit antibacterial activity against negatively charged outer and cytoplasmic membranes of Gram-negative bacteria due to their positive charge (cationic type).[127] A small number of АМPs are anionic owing to the presence of glutamic and aspartic acids (e.g., dermcidin present in sweat); their antibacterial activity is caused by the ability of metal ions to form cationic salt bridges with negatively charged components of bacterial membranes.[128] Antimicrobial peptides interact with bacterial cell membranes, thus destroying them or causing the formation of pores, which results in cell death. АМPs can also be incorporated in the negatively charged cytoplasmic membrane, thus disrupting the cell integrity, and penetrate into the bacterial cytoplasm, binding to DNA and RNA, which also induces cell death.[129] The antimicrobial properties of АМPs, cytocompatibility, molecular structure, and mechanism of antibacterial action have been described in detail in a number of reviews.[130][131]

There are two key strategies for the attachment of АМPs to the implant surface: covalent binding of AMPs to metal surface and non-covalent immobilization. In both cases, the bactericidal effect is achieved through the controlled release of AMPs into the microenvironment surrounding the implant.[132-134] The key factors determining the efficiency of bactericidal protection are the surface density, mobility, and orientation of peptide molecules.[126] A meta-analysis of the available data on in vivo AMP modification of implants was published.[135] Synthetic HHC36 and VM were the most common АМPs used for surface modification, while S. aureus was the bacterium best studied in vivo. The main methods for AMP attachment to the titanium surface included covalent binding to a predeposited polymer surface layer, physical adsorption, and layer-by-layer self-assembly.[136-138] Functional groups present in АМPs such as NH₂ , COOH, and SH can react with functional groups on the implant surface, e.g., OH groups formed after alkaline etching or NH₂ groups resulting from silanization.[139][140]

As of September 2025, the AMP database contains 5680 peptides with various antibacterial, antiviral, and antifungal activities.[141] These peptides are constantly expressed in epithelial and phagocytic cells of the body and are activated when host cells detect an active infection or bacterial endotoxin. АМPs effectively kill antibiotic-resistant bacteria such as MRSA, P. aeruginosa, and quinoline-resistant Enterobacteriaceae.[142] АМPs are classified in terms of length, charge, structure, concentration of hydrophobic amino acids, and the source of origin.[143]

After diffusion release, АМPs can affect bacteria in two ways: (1) by damaging bacterial membranes or disrupting their vital functions or (2) by direct physical contact between the bacterial cell and cationic AMPs.[144] The practical use of AMPs is limited due to a number of problems, first of all, uncontrolled toxicity, high cost of production on an industrial scale, cytotoxic effect on eukaryotic cells, and degradation by host proteases.[145]

A promising alternative to traditional antibiotic therapy is the use of bacteriophages, that is, viruses that selectively infect bacterial cells to induce cell lysis. An advantage of bacteriophages is that they can selectively infect particular types of bacteria without affecting the bacteria of the normal microbiome flora.[146] After the bacteriophage is attached to specific receptors on the membrane of the target bacterial cell, the genetic material of the bacteriophage is inserted into the host cell. During replication of the virus inside the cell, new phages are formed; they attack new targets and destroy the infected bacterial cell. The bacteriophage can also transmit its own genetic material into the bacterial genome and replicate with each cell division of the infected bacterial cell.[147] The efficiency of bacteriophages was demonstrated against various MDR strains such as K. pneumoniae, P.aeruginosa, E. coli, and A. baumannii.

The main benefits of the bacteriophage therapy include high specificity (bacteriophages can affect a specific bacterial strain), the absence of antagonistic effect when bacteriophages are mixed, low toxicity, and the ability to penetrate into biofilms.[148] However, evaluation of the pharmacokinetic behaviour of bacteriophages is a complicated task, since inside bacterial cells, they interact with numerous microbial pathogens that are recognized and suppressed by the human immune system. Determination of the infectious titers of bacteriophages is still labour-consuming, subject to errors, and ineffective.[149] The bioavailability of bacteriophages that are not intravenously administered is considered to be low, while data on intravenous administration are currently limited. Personalized combinations of bacteriophages and antibiotics may be of interest to expand the scope of antibacterial therapy; however, it should be borne in mind that these combinations may give rise to both synergistic[150] and antagonistic effects,[151] depending on the mechanism of inhibition of bacteria by the antibiotic used in combination with the bacteriophage. Furthermore, the use of bacteriophage cocktails does not always provide an advantage over antibiotic therapy.[152]

Using various chemical methods, it is possible to modify surfaces and endow them with particular functions. Antiadhesive or antifouling coatings prevent bacterial adhesion and may consist of hydrophilic polymers, superhydrophobic materials with low surface energy, or have a structured surface. The antiadhesive properties of a surface depend on several key parameters: topography, roughness, wettability, and charge. Many materials occurring in nature have inspired researchers to design artificial surfaces resistant to bacteria, which may mimic the gecko skin,[153] lotus leaves,[154] and insect wings.[65][155] The antiadhesive properties of a surface are usually due to low contact area; therefore, bacterial adhesion is influenced most appreciably by nanostructures with a high aspect ratio (nanospikes, nanotubes, nanoripples) and structured surfaces with approximately 1 μm structural elements (micro-wells, micro-grooves, submicro-pillars, and micro-protrusions).[156] The effect of surface wettability on the bacterial adhesion is highly ambiguous, and the available data are quite contradictory: readers can find relevant information in recently published reviews.[156-158] The superhydrophilic and superhydrophobic surface characteristics are strongly influenced by bound water and trapped air, which prevent bacteria from direct contact with solid surfaces, which reduces bacterial adhesion.[159]

The repulsion of bacteria from a surface may arise if there is a physical barrier to the proteins in the bacterial cell wall caused by electrostatic interaction or low surface energy. Various types of adhesive polymer materials are described in the literature: hydrophilic polymers, among which PEG is used most commonly; chitosan-based hydrogels; mixtures of peptides and polymers; nanoparticle-doped hydrogels; and poly-zwitter-ion polymers,[137] which belong to the type of materials with equal numbers of cations and anions along polymer chains[160] and contain positively and negatively charged groups, making them highly hydrophilic.[161] The biological fouling of polymers results in the formation of conditioning films (macromolecules), which can grow to form biofilms (microorganisms) and lead to macroscopic biofouling (macroorganisms).[162] Antiadhesive coatings often contain enzymes immobilized on the surface that either prevent bacteria from attaching or can break the biofilm matrix.[163]

The van der Waals forces and electrostatic interactions largely determine the bacterial adhesion to the surface. In view of the fact that bacteria are usually negatively charged due to carboxyl, amino, and phosphate groups on their cell walls, they often have more pronounced adhesion to positively charged surfaces. Meanwhile, numerous facts indicate that bacteria can overcome electrostatic repulsion and strongly bind to negatively charged surfaces.[164]

The free surface energy is a critical initial factor for biofilm formation, but its role is ambiguous. The former paradigm according to which bacteria prefer to be attached to hydrophobic surfaces is largely outdated. In particular, in a biological medium, the surface is instantaneously covered by a layer of organic molecules (proteins, polysaccharides), which changes the initial surface free energy. In addition, bacteria can recognize the surface properties and regulate gene expression where they switch from reversible to irreversible attachment and initiate matrix formation.

A specific biofilm control mechanism is the use of bioactive molecules capable of suppressing the QS system. Over the past two decades, quorum sensing inhibitor (QSI) and quorum sensing quencher (QSC) enzymes have been developed. These enzymes are capable of inhibiting various parts of QS signalling pathway and affecting signalling molecules and their receptors. Synthetic QS inhibitors include derivatives of furanone, benzimidazole, benzothiazole, pyranone, quinolines, pyrrole, indole, pyridine, pyrimidine, thiazole, thiadiazole, thiazolidine-2,4-dione, cinnamaldehyde, and esters.[165] In view of the fact that during evolution, plants have developed chemical mechanisms to combat external pathogens, secondary metabolites of various classes of plants such as terpenes, quinones, coumarins, stilbenes, alkaloids, curcuminoids, flavonoids, phenolic compounds, and their derivatives and analogues are suitable candidates for QS system inhibition.[166]

Bioactive molecules have been used for surface functionalization of biomaterials. For example, the antibacterial and cytocompatible peptoid GN2-Npm9 was electrostatically attached to the chemically pretreated Ti6Al4V surface. In the presence of the peptioid, the adhesion of P. aeruginosa MMA83 bacteria markedly decreased, which was attributed to QS disruption.[167] Quorum sensing inhibitors were also used to prevent the biofilm formation on the surface of catheters.[168][169]

Generation of ROS such as hydroxyl radicals (^•OH), superoxide anions O₂^–•, and hydrogen peroxide (H₂O₂) may introduce imbalance (called oxidative stress) between the cellular production and consumption of ROS, which impairs the cell integrity.[170] The oxidation process can lead to the death of microorganisms if it causes oxidative damage to proteins, lipids, and nucleic acids. The O₂^–• anion is formed via cellular redox metabolism mediated by electron transfer proteins. The dismutation of O₂^–• affords H₂O₂ , which can be reduced to hydroxyl radicals by the Fenton reaction. In addition, O₂^–• can react with nitric oxide radical (^•NO) to give reactive nitrogen species.[171]

A combination of ROS and bactericidal ions produces a synergistic antibacterial effect.[55][84] The generation of ROS can be used to prevent the formation of surface biofilm and to kill planktonic bacteria.[172][173] The lifetimes of free radicals such as superoxide (O₂^–•), hydroxyl (OH^•), peroxyl (ROO^•), and alkoxyl (RO^•) amount to fractions of a second, while hydrogen peroxide (H₂O₂) and organic peroxides (ROOH) can be preserved for a few hours.[174] Current research focuses on the development of heterogeneous materials capable of generating ROS both by themselves and after exposure to various sources of radiation.[175-178] These materials are typically photocatalytic semiconductors that generate ROS as a result of electron transition from the valence band to the conduction band to give holes.[179][180] In order to decrease the band gap and enhance the visible and UV absorption, new composites are being developed or known photoactive materials (TiO₂, CuO, ZnO) are being doped with additional elements (e.g., Zn, Ca, Au, graphene, etc.).[181-183] The surface of the material is often decorated with NPs of transition and noble metals (Cu, Ag, Au, Zn, etc.) to prevent fast recombination of electrons and holes.[184-186] The electron work function is higher for metal NPs than for semiconductors, which increases the lifetime of the electron–hole pair. The efficiency of this approach was demonstrated in relation to the Ag/TiO₂ and Pt/TiO₂ systems, which efficiently generated ROS on exposure to UV radiation or visible light.[182] Bimetallic Pt – Fe NPs deposited on the surface of Ti(C,N)-based coatings also enhanced the generation of ROS after UV irradiation.[186]

The separation of electrons and holes is usually activated by some external source (ultrasound, radiation).[63][185-192] The most studied type of external action is UV irradiation, which activates transition metal oxides (Ti, Zn, Cu), bimetallic systems (Cu – Bi), nonmetals (graphene nitrides and oxides, polymers), and various composites.[183][186][190-194] However, this method can be used only before the implant placement, for example, during sterilization, as UV radiation does not penetrate into a tissue to a depth more than 1 mm. Inflammation and implant-associated infection can also occur after implantation during the patient rehabilitation period; therefore, the possibility of implant surface activation directly in the implantation area opens up new opportunities for fighting infections.[195]

In recent years, X-ray therapy methods have been increasingly used, since X-rays penetrate deep into tissues. X-ray luminescence is used more and more often for imaging, biosensorics, and theranostics.[196] Soft X-ray irradiation enhances the generation of ROS in mesoporous titanium peroxide used in tumour therapy.[197] The ROS activation with X-rays may be an effective alternative to UV radiation and will allow X-rays to activate the material directly at the site of infection. The available data indicate that X-ray irradiation may give rise to electron – hole pairs in TiO₂.[198]

The current knowledge of the effect of X-rays on the ability of composite coatings to generate ROS is limited; this approach has not been used until recently to combat bacterial infections. Recently, light-sensitive Si-doped TiCaCON films decorated with Fe and Pt NPs possessing ROS-mediated photoinduced antibacterial activity have been developed. After short-term (5 – 10 s) low-dose X-ray irradiation, the films generated much more ROS than they generated after UV illumination for 1 h. The Fe/TiCaCON – Si films showed enhanced capacity for biomineralization, good cytocompatibility, and excellent antibacterial activity against multidrug-resistant hospital E. coli U20 and K261 and MRSA MW2 strains.[199]

The formation of ROS upon X-ray irradiation of a zinc-doped coating produced by plasma electrolytic oxidation (PEO) was reported for the first time by Popova et al.[84] The zinc-containing TiO₂-based coating with added Na, Ca, Si, and K provided effective and versatile bactericidal protection without compromising the biocompatibility. The excellent cytocompatibility of PEO – Zn samples was confirmed by three methods: monitoring of MC3T3-E1 cell proliferation, evaluation of the viability of sheep osteoblast cells using calcein-AM staining and fluorescence microscopy, and incubation with spheroids based on primary osteoblasts and NIH3T3 mouse embryonic fibroblasts. After 24 h, the PEO – Zn coatings completely inactivated four types of strains, Gram-positive S. aureus CSA154 and ATCC29213 and Gram-negative E. coli K261 and U20, and prevented the formation of E. coli U20 and K261 biofilms. The excellent antibacterial activity is attributable to the synergistic effect of Zn²⁺ ions in a safe concentration and ROS generated upon UV irradiation or short-term soft X-ray irradiation. The results markedly expand the options for using PEO coatings on the surfaces of titanium implants.

The search for alternative strategies for preventing bacterial colonization, instead of methods based on the release of bactericidal ions or antibiotics, resulted in the development of contact-killing surfaces that do not contain toxic substances. These approaches imply generation of a specific surface nanorelief, incorporation of cationic compounds into the coating, or the use of microgalvanic effects (Fig. 2).

Fig. 2

Schematic view of destruction of bacteria upon contact with the surface due to mechanical damage or microgalvanic interaction. The atomic force microscopy image (above) and Kelvin probe force microscopy image (below) of Pt nanoparticles on the surface of TiCaPCON coating were taken from the author’s personal archive.

The bactericidal activity of nanostructured surfaces and their ability to kill bacteria by mechanical interaction with cell walls was demonstrated for various nanomaterials such as black silicon,[200][201] carbon nanostructures,[66][202] zinc oxide nanorods,[203] copper oxide nanosheets,[204] CuO_x NPs,[205] and hexagonal BN (h-BN) nanosheets/nanoneedles.[55] It was assumed that cell death is caused by the ability of vertically oriented nanotubes to accumulate and release elastic energy upon deformation caused by contact with the cells. A similar mechanism may also be involved in the case of nanosheets. The results of TEM analysis clearly demonstrated that membranes of E. coli and Bacillus subtilis cells were damaged by the nanostructured surfaces of h-BN[55] and CuO_x NPs,[205] respectively. A comparison of the bactericidal properties of a nanostructured surface and a surface loaded with an antibiotic was first reported by Gudz et al.[55] The h-BN coatings formed by nanosheets and nanoneedles were shown to have a high antibacterial activity (> 99% after 24 h) against the E. coli K-261 strain with an initial concentration of 10⁴ CFU mL^–1, which is comparable with a GM-loaded h-BN sample (150 μg cm^–2). The mechanical damage to the bacterial wall is the first stage of destruction and may be accompanied by the penetration of bactericidal ions or ROS into the cell, thus accelerating the cell death.

Microgalvanic effects on the implant surface may be caused by local electric currents arising due to potential differences between system components (e.g., metal NPs on the surface and the surrounding conductive matrix) in a conductive biological medium. As a result of dissolution of NPs, which act as cathodes in the system, the electrons that are formed in reaction (7) pass through the surface layer of the coating, which acts as an anode, towards metal NPs and are involved in depolarization reactions (8) and in reactions with molecules of the medium (9). The electrical circuit is closed due to transfer of negatively charged ions through the biological medium to the anode. In the case of Ti-based coating, the above-mentioned chemical reactions can be described in the following way:[54]

The microgalvanic interaction between the Ag NPs and the Ti matrix was studied by measuring the electrochemical polarization and zeta-potential.[206][207] The Ag-modified samples considerably decreased proliferation of the S. aureus and E. coli bacteria despite the low Ag content (< 10 ppb). It was suggested that proton-depleted regions formed during cathodic reactions disrupt the proton electrochemical gradient in the intermembrane space of bacteria.[207] To intensify the microgalvanic effect, the Ag-modified Ti coatings were doped with zinc.[208][209] Although Ag- and Zn-modified coatings demonstrated an increased bactericidal effect, it is difficult to draw any reasonable conclusions about the contribution of microgalvanic processes, since both elements are bactericides. The antibacterial effect of thin magnesium coatings against S. epidermidis was explained by various factors, including microgalvanic interactions.[210][211] Direct experimental evidence for the formation of galvanic couples in the Pt/TiCaPCON system was obtained using Kelvin probe microscopy.[54] Platinum NPs had a more positive potential than the surrounding matrix, with the potential difference being approximately 5 – 10 mV. Electrochemical study also revealed a considerable difference between the free corrosion potentials of Pt NPs (990 mV) and TiCaPCON coating (150 mV). The resistance of the TiCaPCON coating equal to 29 Ω cm^–² implies that it can conduct current. It was show that Pt/Fe NPs, which are positively charged relative to the TiCaPCON matrix (potential difference of ~ 30 mV), become negatively charged (potential difference of 40 – 50 mV), after being kept in PSS.[186]

The surface of Gram-positive and Gram-negative cells carries a negative charge, arising upon dissociation or protonation of carboxyl (COOH), phosphate (HPO₄, H₂PO₄, HPO₄⁻), and amino (NH₃⁺) groups. On approaching the charged surface of an implant, these groups can associate or dissociate on the bacterial surface and thus change their conformation. In the initial stage of interaction of bacteria with the surface, the major role belongs to van der Waals forces and electrostatic interactions.[212] Since the overall bactericidal effect may be caused by simultaneous action of several factors, identification of the contribution of each single factor is a challenging task. Therefore, additional model tests were carried out, which made it possible to exclude the contribution of ions and ROS and confirm that E. coli bacteria are inactivated on contact with the surface as a result of microgalvanic interactions.[186] It was shown experimentally that the negative charge is transferred from the substrate (TiN, TiC, or TiCN) to Pt NPs, which gives rise to electron-enriched Pt sites.[213]

The release of a bactericidal component from the implant surface may be due to changes in specific factors of the medium such as pH, temperature, enzymatic activity, light, etc. For example, chronic wounds form an acidic microenvironment caused by bacterial activity; therefore, biodegradable polymers can release encapsulated AMPs at low pH upon contact with the infected wound bed.[214] The pH-sensitive natural polymers include hyaluronic acid, alginates, and chitosan. The two last-mentioned materials respond also to changes in the ionicity of the environment: Ca²⁺ (alginate) and Mg²⁺ (chitosan); this leads to swelling/dehydration.[215] Other examples are thermosensitive hydrogels, which maintain their integrity at normal body temperature and break down at elevated temperature during inflammation, releasing the therapeutic drug at the site of infection.[216] Thermosensitive and biodegradable hydrogels based on poly(lactic-co-glycolic acid) (PLGA) for drug delivery to various organs have been described in detail in a recent review.[217] A microneedle patch sensitive to near-infrared light and capable of releasing AMPs was developed for the treatment of wound biofilms.[218] The IR780 iodide as a photothermal conversion agent and molecularly engineered W379 peptide were loaded into soluble polyvinylpyrrolidone (PVP) microneedle patch, which was later coated with 1-tetradecanol (TD). After the patch placed onto a wound was exposed to light, the iodide converted light to heat, which induced melting of TD, dissolution of PVP microneedles, and release of the loaded W379 peptide. Also, there is a group of polymers that respond to a change in the electric field [e.g., polypyrrole, polyaniline, poly(3,4-ethylenedioxythiophene), chitosan] or magnetic field (e.g., PEG, dextran, polyvinyl alcohol, polyethyleneimine, etc., containing embedded magnetic particles).[215]

The hydrogel that comprised xylane hemicellulose and biodegradable polymer based on acrylic acid and polyethylene glycol diacrylate was functionalized with Fe₃O₄ NPs.[219] The resulting polymer material could release the therapeutic agent both under the influence of magnetic field or upon pH change.

Smart composites responding to pH change were developed for applications in dentistry. Bacteria in dental plaque metabolize sugars, which increases the acidity, while the composites release therapeutic ions that inhibit bacterial growth in response to the decrease in pH.[220] It was demonstrated that the release of bactericidal Zn²⁺ ions from a glass powder composed of SiO₂, ZnO, CaO, and fluorine is accelerated in acidic media at pH 4.5, which can also be used to treat dental caries.[221][222]

It is known that the synergistic effect implies the enhancement of two or more factors when they act together. Sukhorukova et al.[53] considered the synergism in the effects of the chemical composition and surface roughness on the release of bactericidal Ag⁺ ions. The authors compared five types of pretreated titanium surfaces (subjected to polishing, sand blasting, pulsed electrospark deposition, and laser treatment in two modes to produce surface grids and holes). The samples were then coated with TiCaPCON – Ag film. It was shown that by changing topographical characteristics of the surface, with the content of the bactericidal element being the same, the release of Ag⁺ ions can be both accelerated 2.5-fold and suppressed almost completely. Another important factor that crucially affects the release of ions is the surface oxidation rate. There is an intricate relationship between the density of Ag NPs on the implant surface, surface roughness, and oxidation kinetics. Electrochemical measurements demonstrated that an increase in the concentration of Ag NPs on the surface slows down the growth of the oxide layer, while an increase in the surface roughness accelerates oxidation due to increasing specific surface area, which retards the release of bactericidal ions. The interplay of these factors resulted in the following dependences: for Ti samples coated with TiCaPCON containing 2 at.% Ag, the number of leached Ag⁺ ions increased with increasing roughness, while the sample containing 1 at.% Ag followed the opposite trend: an increase in roughness resulted in decreasing leaching of surface Ag⁺ ions.

The release of bactericidal ions can also be affected by the presence of a second therapeutic component. For example, the presence of GM retarded leaching of Ag⁺ ions from the surface of TiCaPCON – Ag-GM samples into PSS,[63] which may be attributable to GM reaction with silver ions to give stable Ag/GM complexes, which decreases the concentration of free silver ions in the solution.[112] The GM release rate from the TiCaPCON – Ag-GM sample was also lower than that for Ag-free material.

Antibacterial activity assays showed that implants with a cellular surface structure loaded in 4 and 0.4 mg mL^–1 GM solutions completely suppressed E. coli bacteria after 3 and 24 h, respectively. The synergistic effect can be manifested when combinations of antibiotics and bactericidal ions are used. For example, in order to evaluate the bactericidal effect after the active release of the antibiotic and Ag⁺ ions in the first 24 h, the samples were initially kept in PSS for 24 h. The coatings containing both GM and Ag demonstrated a strong antibacterial effect (decrease in the CFU count by 99.9%) during incubation with the S. aureus for 24 h, whereas samples containing either GM or Ag separately exhibited no antibacterial activity.[63]

The kinetics of bactericidal ion release from the implant surface can be regulated by forming a heterogeneous surface containing various NPs. For example, an equilibrium potential is established between the neighbouring bactericidal Ag (cathode) and Zn (anode) NPs, as shown in Scheme 7. The free corrosion potential of zinc shifts to more positive values, while the potential of silver shifts to the negative side. As a result, Ag NPs are under anodic protection of Zn NPs and do not dissolve until the Zn NPs are completely dissolved.[54] A similar situation is observed in the Pt – Zn NP pair. Meanwhile, in the Ag – Zn and Pt – Zn pairs, cathode NPs accelerate the dissolution of zinc with respect to the dissolution in the absence of cathode nanoparticles. This strategy makes it possible not only to combine different bactericidal ions, but also to control the rate or time of ion release.

Scheme 7

There are numerous methods for surface modification to endow a surface with bactericidal properties.[223-225] The particular choice of the most appropriate method depends on the type of material (metal, polymer, organic or bioresorbable materials). Conventionally, two groups of methods can be distinguished: (1) wet chemistry methods for deposition of coatings from solutions such as electrochemical deposition [anodization, micro arc oxidation (MAO), PEO], sol – gel process, centrifugation, immersion into a solution and impregnation and (2) vapour deposition of coatings [physical vapour deposition (PVD) and chemical vapour deposition (CVD)]. Each method has its own benefits and drawbacks, with their comparison being beyond the scope of this review; however, note that PVD and CVD methods typically require the use of special expensive equipment.