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Introduction: interaction of medical devices with the human body

In recent years, a lot of progress has been made in the areas of tissue engineering, biomaterials and polymer science. Accordingly, progress has also been made in the areas of implantable devices and drug/device combination products. Nevertheless, the biocompatibility of implantable devices remains a critical issue, limiting device functionality and presenting a significant risk for patients.

A biomaterial implanted into the body induces a foreign body response composed of nonspecific protein adsorption, cell adhesion, upregulation of cytokines and subsequent proinflammatory processes and finally encapsulation of the device by fibrous tissue.

The ISO 10993 serves the assessment of the overall biological safety of a medical device. It can be used for the valuation of potential new materials for suitability in a medical device for a proposed clinical application. Being focused on the biological evaluation of medical devices it therefore puts not only emphasize on the chemical characterization of materials (part 18) but also on the physicochemical, morphological and topographical characterization of materials (part 19). The combined surface characteristics of medical devices are decisive for the adhesion of proteins, bacteria and eukaryotic cells and several strategies on biofunctionalization of surfaces attempt to achieve an appropriate biological response.

Attachment of proteins, bacteria and mammalian cells

Biofouling

Immediately after implantation, a nonspecific attachment of native and denatured proteins, cells and other biological material takes place on the implant surface, a process called “biofouling” [1]. This nonspecific attachment is completely untypical for normal physiological processes and is therefore probably the trigger for the foreign body reaction [2]. Biofouling and the subsequent inflammatory reaction of the body are the main causes for the failure of in vivo implants [1], [2].

Preventing nonspecific protein adsorption reduces undesired biological responses, such as thrombus formation in blood-contacting devices and limits the initial stages of bacteria adhesion and hence biofilm formation [3], [4]. Therefore, protein-resistant properties of medical devices might extend their service life and biocompatibility.

A broad range of anti-biofouling strategies already exist, using biological, physical and chemical techniques to inhibit protein deposition as the initial step of biofouling. For example, surfaces bearing poly(ethylene glycol) (PEG) moieties strongly resist nonspecific protein adsorption. This has led to the wide use of PEG in biomedical applications [5], [6], [7]. In addition, micropatterns can be formed on surfaces to make them superhydrophilic or superhydrophobic like the self-cleaning surface of the lotus leave (see review [8]).

Bacterial adhesion

Bacteria adhere to virtually all natural and synthetic surfaces [9], [10] and stimulate the organism to produce extracellular polymeric substances, such as polysaccharides, proteins, nucleic acids and lipids [11], through which they embed themselves in a protective matrix. This matrix provides mechanical stability and constitutes the main difference between planktonic bacteria and bacteria adhering to a surface in a so-called biofilm mode of growth. Bacteria organized in biofilms are at least 10–1000 times more resistant to antibiotics [12] than bacteria in a planktonic state and can cope much better with unfavorable external conditions in the host immune system than planktonic ones.

Initial bacterial adhesion arises from fundamental physicochemical interaction, such as the attractive Lifshitz-Van der Waals forces, electrostatic forces, hydrogen bonding and Brownian motion forces [13]. Additionally, specific molecular interactions operate over small distances. They arise from highly specific, chemical domains on the interacting surface, such as mannose, which is specifically bound by type 1 fimbriae possessing tip-positioned adhesive protein FimH, expressed by Escherichia coli [14].

To maintain antimicrobial activity, conventional formulations of antibiotics are administered. However, concentrations below the minimum inhibitory concentration (MIC) occur frequently during anti-infective treatment and induce antibiotic resistance. Besides, many antibiotics have short half-life values and need to be administered frequently, which contributes to patient incompliance. This poor compliance often leads to treatment failure and increases the costs of health-care resources such as the requirement for additional agents and hospital admission. By maintaining a constant plasma drug concentration over the MICpem for a prolonged period, sustained-release approaches maximize the therapeutic effect of antibiotics while minimizing antibiotic resistance [15].

Mammalian cell interaction

In vitro studies indicate that initial interactions typically occur between extracellular matrix (ECM) molecules adsorbed at the surface of the biomaterial, and transmembrane integrin receptors present on the cell surface. Rapid adsorption of endogenous proteins to a material surface provide a structural framework which initiates the cellular adhesion. Specific amino acid sequences have been identified to ligate many different cellular integrins, which facilitate adhesion of a variety of cell types [16], [17]. The ECM proteins also control downstream biological processes, such as cell proliferation, differentiation, migration and even cell death.

The subsequent performance of the implant or tissue-engineered product depends on the extent to which the cells interact with the surfaces of these products. Medical devices being in contact with blood, e.g. stents are inherent foreign bodies in the vessel wall, inducing platelet adhesion and activating coagulation, which may lead to thrombosis. The end stage of a potential foreign body response is fibrosis, which involves walling off the device by a vascular, collagenous fibrous capsule that is typically 50–200 μm in thickness [2]. This fibrous wall confines the implanted device and prevents it from interacting with surrounding tissues.

Interestingly, the degree of protein deposition, blood coagulation, or acute inflammation induced by a biomaterial does not always correlate with the amount of fibrosis. The chemical composition of the surface of the biomaterial indeed plays a role, but its physical structure and surface topography are more important to predict the degree of fibrosis [18]. For example, implants with a smooth and non-degradable surface create a thicker fibrous capsule than rough or textured implants [19], [20]. Modern implants use different chemical and topographical modifications to regulate cellular adhesion, differentiation and de novo tissue deposition (see review [21]).

In particular, nanoscale geometrical, topographical and mechanical properties of biomaterials are essential to achieve control of the cell-biomaterials interface [22]. For example, the antifouling properties of PEG can be used when grafted to polyelectrolyte multilayer (PEM) films to reduce cell adhesion. For this purpose, PEG grafted to poly(ethylene imine) (PEI) can be deposited onto the top of PEI/PAA [poly(acrylic acid)] multilayer films. Cell adhesion is reduced on the swollen PEM films and further decreases by increasing deposition of PEI-g-PEG as the topmost layer [23].

Biofunctionalization of medical devices with polyelectrolyte multilayer coatings

PEMs are nanoscale polymer thin films, which are formed by alternating deposition of polyelectrolytes. The used polyelectrolytes are polymers whose monomer units bear an electrolyte group which dissociate in suitable, e.g. aqueous solution, leaving charges on the polymer chains and releasing counter ions in solution. Examples of polyelectrolytes include polystyrene sulfonate, polyacrylic and polymethacrylic acids and their salts, DNA and other polyacids and polybases. The build-up process is driven by electrostatic interactions, but the stability of the films is given by entropic forces.

PEMs, fabricated by layer-by-layer (LbL) self-assembly of polycations and polyanions are a versatile tool for biofunctionalization of surfaces [24], [25]. A flexible deposition process is feasible by dipping or spraying on materials with any size and geometry. The resulting films show defined topography and controlled stability with the potential for tunable drug release under physiological conditions.

During preparation of the multilayer coatings, certain deposition parameters can be controlled and modified, such as the number of deposited layers, the type of specific polyelectrolytes, salts, pH, temperature, etc. These parameters influence the physicochemical characteristics of the PEM films, such as their chemical composition and charge, flexibility and stiffness, roughness and stability, as well as its growth regime (linear vs. exponential). For example, thinner PEMs are stiffer. On the other hand, exponentially growing PEMs are in general less dense and thus more flexible.

Accordingly, certain characteristics can be selected, to modulate the biofunctionality and interaction with biomolecules and cells [26]. For example, protein resistant properties can be generated on polycation-ending PEMs that efficiently prevent the adsorption of the negatively charged model protein glucose oxidase. The protein-resistant properties of these films, based on two typical weak poyelectrolytes, poly(allylamine hydrochloride) (PAH) and PAA, can be attributed to a high film hydration becoming the predominant factor over electrostatic interactions [27].

The flexibility of polyelectrolytes in solution is another important parameter which can influence cell adhesion. It was shown that the semi-flexible polyelectrolytes sulfated chitosan and cationic guar gum result in very little protein adsorption when used as components of PEMs, compared to the flexible polyelectrolytes poly(sodium 4-styrenesulfonate) (PSS) and poly(diallyldimethylammonium chloride) (PDADMAC). The adsorption of bovine serum albumin as test protein showed to be irrespective of which polyelectrolyte is used to terminate the film [28].

Besides several patents, there exist also two products, which are manufactured and sold commercially. One is the Yassa Sheet (Shiratori NanoTechnology, Kawasaki, Japan). It is equipped with a multilayer film and contains an enzyme that controls the ethylene concentration and thereby extends the shelf life of fruits and vegetables. In addition, contact lenses, equipped with a multilayer coating are sold commercially (CIBA Vision/Novartis, Duluth, U.S.A.).

In a recent study, we increased the surface roughness by introducing kosmotropic anions. The addition of Hofmeister anions improved the biocompatibility of PEM films built from the natural polysaccharides hyaluronic acid and chitosan: We could observe a decrease in albumin adsorption, and a subsequent anti-thrombogenic effect [29].

Last but not least, the top-down erosion principle of biodegradable PEMs naturally prevents bacterial attachment to coated surfaces [30].

Polyelectrolyte multilayer coatings with antibacterial activity

Nosocomial infections are developed in patients during their stay in a hospital. Such infections occur in 5%–10% of all hospitalizations in Europe and North America and in more than 40% in parts of Asia, Latin America and sub-Saharan Africa [31].

Staphylococcus aureus and Pseudomonas aeruginosa have become the most important causes of nosocomial infections in recent years. Their pathogenicity is mainly due to the ability to form biofilms on biomedical devices and prosthesis [32], [33]. Biofilms can attach to both, abiotic and biotic surfaces, and have therefore been implicated in the development of wound infections, non-healing wounds, but also medical device-related infections [34], [35], [36].

The initial contamination of the medical device most likely occurs from a small number of microorganisms, which are often transferred to the device via the patient’s or healthcare workers’ skin, contaminated water or other external environmental sources [37], [38]. Medical device-related infections pose a huge financial burden on healthcare services, being associated with increased patient morbidity and mortality [39].

With the increasing evidence of the presence of biofilms in healthcare-associated infections, current research has focused upon more sophisticated methods of sterilization and the modification of medical devices in order to prevent microbial growth and biofilm formation. There are two general ways to prevent bacterial attachment and subsequent bacterial infections: (1) physical and (2) pharmacological means. PEM can be used for both attempts, by building up anti-adhesive layers and by incorporating antibacterial substances, such as antibiotics, which are then slowly released (see Figure 1).

Figure 1:

Left: Scheme of a layer-by-layer (LbL) assembly of polyelectrolyte multilayer. Middle: Scheme of the main antibacterial strategies (anti-adhesive layers and releasing of antibiotic substances). Right: Fluorescence microscopic pictures showing high number of live bacteria (green) only on uncoated substrate. LbL films with antibiotics show dead bacteria (red) and anti-adhesion layer clearly reduce the number of live bacteria. Reproduced from review [40] with permission of The Royal Society of Chemistry.

Correlation between surface physicochemical properties of PEMs and antibacterial effect

Surface hydrophobicity

Superhydrophilic surfaces have good nonfouling properties due to their high affinity to water. The strongly bound hydration layer on the hydrophilic surface opposes the adsorption of molecules to the surface. This weakens the interaction between the cell surface and the substratum material and thus reduces cell adhesion. Therefore, many coatings involve highly hydrated polymers, e.g. PEG to exhibit antifouling properties [41]. Due to the strong entropy based interactions and affinity of PEG for water, such coatings significantly reduce both, protein adsorption and bacterial adhesion. Studies showed that PEMs based on the combination of PDADMAC/PAA and PAH/PAA already decrease the adhesion strength of the yeast Saccharomyces cerevisiae. However, a further reduction in initial cell attachment and adhesion strength can be observed after adsorption of PAA-g-PEG copolymer onto PEMs [42].

In addition, zwitterionic polymers with a positive and a negative electrical charge in close proximity can be used to coat material surfaces to obtain superhydrophilicity. Unlike PEG, zwitterionic polymers have a broader chemical diversity and greater freedom for molecular design [43]. Incorporation of various zwitterionic components into coatings was shown to decrease adsorption of proteins [44], [45], [46], attenuate immune responses [47], [48], and resist attachment of eukaryotic cells [49] and occasionally bacteria [50].

Recent studies on PEMs, built with PAH and copolymers of PAA and the zwitterionic group 3-[2-(acrylamido)-ethyldimethyl ammonio] propane sulfonate (PAA-co-AEDAPS) showed contradictory results: these PEMs resist the adhesion of mammalian cells, but are bacterial adhesive, due to differences in adhesion mechanisms as well as different responses to the topology and morphology of the multilayer surface [51].

On the other side, superhydrophobic surfaces also display antifouling properties and can inhibit microbial adhesion, as well [52]. Superhydrophobic PEM can be created for example via deposition of gold on six bilayers of PDADMAC and PSS, as shown in Figure 2 [53].

Figure 2:

(A) SEM image of dendritic gold clusters formed on an ITO electrode modified with a PDADMAC/PSS polyelectrolyte multilayer by electrochemical deposition at −200 mV (vs Ag/AgCl). (B) Shape of a water droplet on the surface of dendritic gold clusters showing the super hydrophobicity of the coated surface. (Reprinted with permission from [53]. Copyright 2004 American Chemical Society.)

Recent research confirms that both, superhydrophobic and superhydrophilic surfaces might prevent biofilm formation: When the surface roughness is in an appropriate range, the system reaches a state, in which air is trapped in the grooves between surface features and thus prevents wetting [54].

Surface charge

Most bacterial cells are negatively charged; thus, in general, a positively charged surface is more susceptible to bacterial adhesion, whereas a negatively charged surface is more resistant [55]. For example, the negatively charged, antiadhesive polymer heparin can be assembled into multilayers with the positively charged, antibacterial biopolymer chitosan, taking advantage of the combined beneficial polymer characteristics [56].

Meanwhile, surfaces presenting certain cationic groups, such as quaternary ammonium, PEI, or chitosan, have particular antimicrobial activities and thus can kill attached bacterial cells [41]. Different studies in the past showed that specific adjustment of multilayer assembly and post-assembly conditions (e.g. pH value) can serve to expose mobile cationic charge and create antimicrobial PEMs without the addition of specific biocidal species. For example, this could be utilized for PEMs comprising PSS and PAH, assembled at high pH and subsequently immersed in low pH solutions. This system underwent a reversible pH-dependent swelling transition, and antimicrobial functionality at physiological pH conditions could be turned on and off with suitable pH treatment. As a result, the viability of gram-positive S. epidermidis and gram-negative E. coli was effectively reduced on PSS/PAH multilayers displaying accessible cationic charge [57].

Mechanical stiffness

An additional parameter that can regulate the adhesion and subsequent proliferation of viable bacteria is the stiffness of the coated substrate. For example, a study on biofilm formation on poly(dimethylsiloxane) with varying stiffness (2.6–0.1 MPa) showed that an increase in stiffness reduced not only the adhesion and growth of E. coli and P. aeruginosa but also the size and antibiotic susceptibility of attached bacteria, indicating that bacteria are able to sense and respond to material stiffness during biofilm formation [58].

In another systematic study, weak PEM films were employed, comprised of the synthetic polymers PAH and PAA, assembled over a range of conditions, to explore the physicochemical and mechanical characteristics of material surfaces controlling adhesion and growth of S. epidermidis (Gram-positive). AFM-enabled nano-indentation of hydrated PEMs indicated an average film stiffness ranging over two orders of magnitude from the stiffest PEM assembled at pH 6.5 (E = 80.4 MPa) to the most compliant PEM assembled at pH 2.0 (E = 0.8 MPa). The authors found that adhesion of viable S. epidermidis correlates positively with the stiffness of these polymeric substrata, independent of the roughness, interaction energy, and charge density of these materials. Quantitatively similar trends were observed for E. coli (Gram-negative), showing lower bacterial numbers on softer PEMs [59], as usually observed for eukaryotic cells.

Another study showed contrary effects with photo-cross-linkable polyelectrolyte films, whose nanomechanical properties were varied under UV light illumination, with minimal variations of film chemistry and microstructure. Here, E. coli and Lactococcus lactis grew faster on soft (Young’s modulus 30 kPa) than on stiffer PEM films (Young’s modulus 150 kPa) [60] (see Figure 3). The films were prepared from biopolymers: poly(l-lysine) and a hyaluronan derivative modified with photoreactive vinylbenzyl groups (HAVB).

Figure 3:

The adhesion and the growth of E. coli were studied on softer non-cross-linked and stiffer photo-cross-linked PLL/HAVB films to investigate how the film stiffness influences the bacterial behavior. The optical microscopy images show that the gram-negative E. coli exhibited a more rapid growth on non-cross-linked softer films compared to the stiffer ones. (Reprinted with permission from [60]. Copyright 2013 American Chemical Society.)

Even though in each of these studies, the chemical composition and microstructure of the films might play an additional role, it could be concluded that bacterial adhesion is especially favored on very soft coatings (30/100 kPa) and reduced only on surfaces with certain, intermediate stiffness (up to 2.6 MPa). Coatings with higher rigidities (80.4 MPa) again seem to be more susceptible for bacterial growth.

Antibacterial PEMs with pharmacological function

Despite significant progress in designing surfaces that inhibit bacterial adhesion and creating antifouling coatings via surface modification, the use of antibacterial agents to prevent infection still might be clinically unavoidable [15], [61].

Although many antimicrobial implant coatings have already been developed, drug dosage and release rate cannot be easily tuned, and local sustained delivery is still difficult to achieve [15]. One possible approach is the use of specific PEMs that allow versatile and tunable release of drugs like antibacterial substances and make them additionally attractive as coatings for implantable devices [62].

Antibiotics

Antibiotics either kill or inhibit the growth of bacteria: Bactericidal antibiotics target the cell wall (e.g. penicillins) or interfere with essential bacterial enzymes (e.g. rifamycins) and kill the bacteria. Bacteriostatic ones target the bacterial protein synthesis (e.g. macrolides) and inhibit bacterial growth.

The sustained release of antibiotics from coated surfaces can be used to provide a high local concentration and kill bacteria before biofilm formation [63]. The amounts of biocide agents in the blood or surrounding tissue can also be effectively controlled through drug delivery surfaces and help to improve the biocompatibility.

Cationic antibiotics, e.g. tobromycin, gentamicin, and polymyxin B can be easily immobilized into PEMs resulting in bio-responsive, controlled-release antibacterial coatings. For example, films were constructed by direct assembly of tannic acid (partially ionized at pH 7.5), combined with cationic antibiotics. These coatings stayed stable in phosphate-buffered saline at pH 7.4 for as long as 1 month in the absence of bacteria and released antibiotics only upon pH lowering [64]. More recently, PAA-gentamicin/PEI multilayer films were constructed via the LbL self-assembly method. In the biofilm inhibition assay, this multilayer film effectively inhibited bacterial adhesion and biofilm formation while showing only low cytotoxicity against human lens epithelial cells [65].

Silver

Another, increasingly used biocidal substance is silver, which is toxic for bacteria, fungi and viruses at concentrations of 10−9–10−6 M [66]. Especially silver nanoparticles have gained considerable attention over the last few decades due to the increasing bacterial resistance to commonly used antibiotics [67]. There are several pathways known for the interactions between silver compounds and bacteria [68]. These interactions can be classified into direct interactions of the silver compounds with (i) the bacterial cell wall, (ii) DNA, (iii) enzymes and membrane proteins, and (iv) interactions based on the silver-induced formation of reactive oxygen species [67].

Recent publications described the incorporation of silver nanoparticles into dopamine-modified alginate/chitosan multilayer. With this method, it was possible to modify the surface of titanium alloy and improve its antibacterial property against E. coli and S. aureus (see Figure 4) [69].

Figure 4:

Antibacterial activity of titanium against E. coli and S. aureus: (a) uncoated, (b) dopamine-modified alginate/chitosan multilayer, and (c) silver nanoparticles in dopamine-modified alginate/chitosan multilayer. (Reprinted from [69] Copyright 2013, with permission from Elsevier.)

In another study, the surface of poly(ether sulfone) membranes was modified with PSS, PDADMAC, and silver nanoparticles to reduce organic and biological fouling. These membranes could be useful for organic and biological fouling prevention during water and wastewater treatment [70]. In a murine wound infection model, the antibacterial efficacy of silver-impregnated PEMs immobilized on a biological dressing was shown. In detail, positively charged 20 nm diameter poly(vinylpyrrolidone) coated silver nanoparticles were incorporated into PEMs of the oppositely charged polyelectrolytes PAH and PAA [71].

Although silver ions have low toxicity, they can cause cell death in the eukaryotic host tissue: Concentrations of 10 μg/mL silver nanoparticles were reported to induce morphological changes in mammalian stem cells, reduce mitochondrial function and therefore cell viability [72]. After permeabilization of the eukaryotic mitochondrial cell membrane, silver ions induced mitochondrial swelling, accelerated cellular respiration and cytochrome c release [73]. Additionally, mitochondria independent toxicity was described for erythrocytes where silver nitrate induced eryptosis. Cell shrinkage and hemolysis were observed with 50–100 nM silver nitrate. The resulting reduced number of erythrocytes in the blood could lead to anemia [73].

Chitosan

Chitosan has several advantages over other disinfectants, as it possesses a higher antimicrobial activity, a broader spectrum of activity, a higher kill rate and lower toxicity towards mammalian cells [74], [75]. Chitosan is a hydrophilic, cationic biopolymer which exerts either bacteriostatic or bactericidal activity by disrupting the inner organelles or disturbing the metabolism of certain bacterial strains [76], [77]. The actual mechanism of inhibition is not yet fully understood. The most feasible hypothesis is a change in cell permeability due to interactions between the positively charged polysaccharide (at pH lower than 6.5) and the negatively charged bacterial membrane. Bacterial growth is then inhibited, when the positively-charged polymer combines with anionic components such as N-acetylmuramic acid, sialic acid and neuraminic acid, on the cell surface [78].

Bound, low molecular weight chitosan could then penetrate the cell wall of bacteria, bind to DNA and inhibit DNA transcription [79]. Bound, high molecular weight chitosan could alter cell permeability or form an impermeable layer around the cell and block any transport over the cell wall [80], [81], [82]. Antibacterial effects of chitosan and its derivatives were shown both, when used in solution and when immobilized on surfaces [83]. Chitosan as polymer showed higher antibacterial activity than chitosan oligomers [84].

Several studies showed that chitosan is more effective for gram-negative bacteria than Gram-positive bacteria [76], [85], [86]. The MIC values of acid-soluble chitosan are much lower (0.03%–0.1%) than those of water-soluble chitosan (0.05%–0.8%) [87]. The acid-soluble chitosan with a deacetylation degree of 0.99 and low viscosity was shown to be most effective in inhibiting bacterial growth [88].

Chitosan can be deposited via LbL assembly: In our own studies, the alternative deposition with hyaluronic acid on polyester and cotton textile proved to be biocompatible for mammalian cells and showed antibacterial functionality against E. coli, S. aureus and Klebsiella pneumoniae [89]. This approach could be used to coat different textiles in order to kill pathogens and to minimize their transmission in the hospital. Examples for possible applications would be coatings of wound dressings, compresses, surgical clothing, bed linen and towels.

In a similar study, LbL films were assembled on cotton samples, using chitosan and alginates to suppress the growth of S. aureus and K. pneumonia [90]. Likewise, chitosan and pentasodium tripolyphosphate were assembled on cotton fabrics to improve inhibition against E. coli and S. aureus [91]. Additional studies, applying alternative deposition with anionic lentinan sulphate showed antibacterial activity against P. aeruginosa [92].

Besides the described desired antibacterial effects, chitosan also promotes plasma protein adsorption, platelet adhesion and activation, and thrombus development, all of which is appropriate for bacteria adhesion [93]. In order to overcome these disadvantages, another study combined chitosan with carrageenan multilayers, to provide additional anti-adhesion properties [94].

Summary/outlook

Different ways of surface modification are known and widely used for finishing of different medical products. Here, we summarized the application of the layer-by-layer (LbL) coating technique for preparation of polyelectrolyte multilayer (PEM). This technique is versatile and can be used for fine adjustment of different surfaces. The PEM based coatings have a thickness of only a few nanometers. They can form robust coatings for long-lasting surface modifications or can be prepared as slowly degradable coatings for controlled release of active substances. The large variety of used and suitable polyelectrolytes enables the formation of coatings with well-predefined and biocompatible properties.

This review outlined specific PEM-based strategies to control the interactions between surfaces and biological systems. The results are important for the development and modification of surfaces from different medical implants. The general aim is to assure better contact between the biological tissues and the implants, and decrease the risk of bacterial infections caused by the implantation. The physical properties of the surfaces like hydrophobicity, hydrophilicity, charge and charge density, as well as roughness play an important role to control the adhesion and proliferation of mammalian cells. These properties, but often also the combined release of active compounds such as silver or antibiotics from the surfaces, can be used to suppress the adhesion and growth of bacteria on these materials. The special properties of chitosan are discussed in more detail within this review. Chitosan can suppress the bacterial adhesion, both because of its charge and because of its interaction with the bacterial cell wall.

Acknowledgments

Some of this work was performed as part of the projects ANKLANA (BMBF KMU innovative, FKZ: 13N13010), EnzCap (BMBF Strategy process “Nächste Generation biotechnologischer Verfahren – Biotechnologie 2020+”, FKZ: 031A171B) and Systemimmunologie an biologisch-technischen Grenzflächen (subproject of “Forschungscampus BioMedTech”).

Author’s statement

  • Conflict of interest: Authors state no conflict of interest.

Materials and methods

  • Informed consent: Informed consent has been obtained from all individuals included in this study.

  • Ethical approval: The research related to human use has been complied with all the relevant national regulations, institutional policies and in accordance the tenets of the Helsinki Declaration, and has been approved by the authors’ institutional review board or equivalent committee.

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