PDMS-based microfuidic devices have also been applied to explore how bacterial colonization, competition, and dispersal occur at fow conditions. Indeed, fow can confer growth advantages to pathogens by allowing the bacteria upstream movement [42]. Similarly, the study of bioflm development is also possible using these microfuidic platforms [35, 43–48]. Several authors revealed that fow alone was able to induce the formation of polysaccharide intracellular adhesins [46] and was the major modulator of the bioflm structures [45]. Additionally, Lee et al. [35] demonstrated that the morphology of *Staphylococcus epidermidis* bioflm formation was infuenced by local hydrodynamic conditions. While higher wall shear stress limited vertical bioflm growth, resulting in a monolayer structure, cells growing in stagnant areas were able to proliferate rapidly, resulting in the formation of a large multilayer structure [35]. Likewise, bioflm thickness was also affected by fow after 48 h, increasing signifcantly at 0.010 Pa (36 ± 9 μm) and slightly at 0.0035 Pa (20 ± 4 μm). Contrarily, no increase was detected for higher shear stresses [44]. Accordingly, Kim et al. [48] revealed that quorum sensing-mediated communication during bioflm formation was generally repressed by fow, impairing bioflm growth. The comprehensive analysis of gene expression during *S. aureus* bioflm formation was successfully conducted by Moormeier et al. [49, 50] using a different microfuidic device, the BioFlux system (Fluxion Systems, South San Francisco, CA), and compared with static conditions. The BioFlux system was presented as the most prominent commercial microfuidic platform that overcomes the limitations of static well plates and conventional laminar fow chambers. In this system, bioflm formation can be followed by light microscopy in microfuidic wells, allowing rapid screening of the effects of several compounds on the viability of bioflms under hydrodynamic conditions [51]. One of the early studies performed on this platform evaluated the effect of several antimicrobials on 8 h-developed *P. aeruginosa* bioflms under controlled hydrodynamic conditions at 37 °C. Results suggested that bioflm viability measured with the plate reader agreed with those determined using plate counts and with the results of fuorescence microscope image analysis. Since then, the BioFlux system has been considered a high-throughput methodology for the study of bioflm development under defned hydrodynamic conditions [36, 49, 50, 52–54].

Although only 1 of 21 analysed studies had the specifc objective of evaluating bacterial adhesion in urinary stents, all provided a comprehensive analysis of adhesion and bioflm formation at fow conditions representative of relevant hydrodynamic regions of ureteral stents [55] and should be considered when testing a new surface or coating for these medical settings.

#### **5 Operating Conditions**

As previously shown, MRDs, fow chambers and microfuidic devices have been used to study bacterial adhesion and bioflm formation under hydrodynamic conditions that simulate the UTDs. Because the fow rate by itself provides little information about shear without taking into account the geometry of the *in vitro* fow system, it is crucial to mimic the fow conditions in a catheter or stent by using either the wall shear stress or the shear rate [1]. The wall shear rate (σ, with unit/s) is a measure of change of the fuid velocity near the wall of the tube in the radial direction toward the center of the tube. In laminar conditions, the shear rate is related to the force which the fuid fow exerts on the wall, expressed as shear stress (τ, with unit Pa), through *τ = μ* × σ, where *μ* is the dynamic viscosity of the fuid (10−3 Pa s for water). In the fow systems under study, the fow rate should be adjusted to approach an average shear rate of around 15/s as an estimate of the intraluminal urine fow, based on predictable daily urine production and internal catheter diameter [15]. Nevertheless, urinary output values are highly variable and may reach more than 10 times the mean value [56], yielding a proportional increase in wall shear rate. Some authors performed FC tests at a shear rate of 33/s, which is higher than mean values but still within the range of shear rates found in urinary catheters [25].

Regarding the fow chamber system described in this work (Fig. 2), the numerical simulations indicated that the shear rate of 15/s reported for urinary fow in catheters can be attained at a fow rate of 2 mL/s [2, 9]. On the other hand, the average shear stress in problematic zones of ureteral stents that are prone to encrustation (0.024 Pa) [55] can be obtained by operating the PPFC system at a fow rate of 4 mL/s [21]. In the case of MRDs used by our research group, the recirculation fow rates can range from 5 [12] to 53 mL/s [13] to mimic the shear forces on urinary catheters, depending on the geometry of the fow cell.

PDMS-based microfuidic devices are usually designed for a particular application, having their own architecture and geometry with specifc operating conditions. In the case of the commercially available BioFlux system, numerical simulations revealed that the average shear stress value of 0.02 Pa reported for ureteral stents [55] can be reached at a fow rate of 66 μL/h [52].

#### **6 Strengths and Limitations of Flow Platforms**

Among the advantages of fow systems are the ability to compare, for instance, the effect that different substrates, media and hydrodynamic conditions exert on a bioflm at different developmental stages. These dynamic models may also provide an evaluation of the effect that transiently occurring molecules, such as antibiotics or adherence inhibitors, have on bioflms. However, the technical disadvantages of fow reactors include increased experimental complexity as well as possible formation/trapping of air bubbles in the setup tubing (particularly severe in microfuidic systems), as this can affect fow and bioflm architecture [57].

Choosing the experimental platform for fow experiments determines what kind of data can be extracted, and care must be taken to ensure that the selected reactor fulflls the objectives of the experiments. The three platforms covered in this chapter (modifed Robbins device, fow chamber and microfuidics-based device) have benefts and limitations, which are summarized in Table 4.

**Table 4** Advantages and disadvantages of dynamic bioflm cultivation devices

#### **7 Conclusions**

To evaluate the anti-adhesive and antimicrobial performance of novel biomedical materials, a number of fow devices have been designed to recreate *in vivo* fow conditions. Shear stress and fow rate can be accurately controlled and varied in these *in vitro* fow systems, which requires prior knowledge of the fow dynamics inside the platform. After limiting their operational range, modifed Robbins devices, fow chambers and microfuidic devices are suggested as experimental setups to mimic the fow behavior in urinary catheters and stents.

**Acknowledgements** This work was fnancially supported by: LA/P/0045/2020 (ALiCE), UIDB/00511/2020, and UIDP/00511/2020 (LEPABE), funded by national funds through the FCT/ MCTES (PIDDAC); Project PTDC/CTMCOM/4844/2020 funded by the Portuguese Foundation for Science and Technology (FCT); and by Project 2SMART (NORTE-01-0145-FEDER-000054), supported by Norte Portugal Regional Operational Programme (NORTE 2020), under the PORTUGAL 2020 Partnership Agreement, through the European Regional Development Fund (ERDF). L. C. Gomes and M. J. Romeu acknowledge FCT for the fnancial support of her work contract through the Scientifc Employment Stimulus—Individual Call—[CEECIND/01700/2017] and for a PhD grant (SFRH/BD/140080/2018), respectively. R. Teixeira-Santos acknowledges the receipt of a junior researcher fellowship from the Project PTDC/BII-BIO/29589/2017— POCI-01-0145-FEDER-029589. Support from the EU COST Action ENIUS (CA16217) is also acknowledged.

#### **References**

**Open Access** This chapter is licensed under the terms of the Creative Commons Attribution 4.0 International License (http://creativecommons.org/licenses/by/4.0/), which permits use, sharing, adaptation, distribution and reproduction in any medium or format, as long as you give appropriate credit to the original author(s) and the source, provide a link to the Creative Commons license and indicate if changes were made.

The images or other third party material in this chapter are included in the chapter's Creative Commons license, unless indicated otherwise in a credit line to the material. If material is not included in the chapter's Creative Commons license and your intended use is not permitted by statutory regulation or exceeds the permitted use, you will need to obtain permission directly from the copyright holder.

# **Biomaterial-Associated Infection: Pathogenesis and Prevention**

**Martijn Riool and Sebastian A. J. Zaat**

#### **1 The Clinical Problem**

The use of medical devices, such as urinary stents, catheters, artifcial heart valves, prosthetic joints and other implants, collectively often referred to as "biomaterials" has increased dramatically over the past century, and has become a major part of modern medicine and our daily life. With the aging society, the higher demand on these devices to restore function and quality of life, combined with the ever improving technology within the medical feld, the problem of biomaterial-associated infection (BAI) is expected to increase.

Catheters, and orthopedic devices are among the most frequently used devices in human medicine [1, 2]. Catheters suspected for infection are replaced by a new catheter at a different location, since using the original location for re-implantation over a guide-wire is strongly discouraged because of the high reinfection risk [3]. Primary implantation of prosthetic joints like prosthetic hips, knees, elbows and ankles, is considered a so-called clean procedure [4], however, in 0.5–1% (hip or knee) to over 5% (elbow or ankle) of cases, infections occur [5, 6]. Revision surgery is associated with higher frequencies of infection, due to the compromised condition of the tissue, longer procedures and more extensive tissue damage during surgery.

The most common causative microorganisms in BAI are *Staphylococcus aureus*, a major pathogen in wound infections, and *Staphylococcus epidermidis*, the harmless skin commensal [6–8]. Depending on the type of device and location of application, other pathogens such as coagulase-negative staphylococci, enterococci, streptococci, *Propionibacterium acnes* and yeast can also cause BAI [9, 10].

M. Riool (\*) · S. A. J. Zaat

Department of Medical Microbiology and Infection Prevention, Amsterdam UMC, Amsterdam institute for Infection and Immunity, University of Amsterdam,

Amsterdam, The Netherlands

e-mail<EMAIL_ADDRESS><EMAIL_ADDRESS>
As early as in 1957, Elek and Conen studied the minimum infective dose of staphylococci for man in relation to suture infection [11]. In healthy volunteers, they estimated the minimum pus-forming dose of *S. aureus*—called *Staphylococcus pyogenes* in those days—on intradermal injections in absence of sutures to be 2–8 million bacteria, numbers which are improbable in case of a natural infection. However, the presence of a foreign body, a suture in this case, resulted in a dramatic reduction in the minimal inoculum required for pus production: a dose of 300 bacteria led to abscess formation. Higher inoculum doses even resulted in lesions with 'the size of an orange', caused fever and took over a week to resolve, in spite of penicillin therapy. Although this experiment clearly demonstrated the enhancing effect of the presence of a foreign body, but the authors stated that the outcome of the experiment "led to great diffculty in fnding further volunteers". Nowadays, such an experimental set-up would not be easily approved by medical ethical committees, but it did provide crucial information on the pathogenesis of BAI. Thus, it has been recognized for at least 60 years that the presence of a foreign body predisposes for infection, and this has repeatedly been confrmed in animal studies [12–15]. In rabbits, for example, only 50 colony forming units (CFU) of *S. aureus* were suffcient for infection in the presence of a cemented hip implant, whereas 10,000 CFU were required in absence of the foreign body [16].

#### *1.1 Bioflms*

Bacterial bioflm formation is considered the major element in the pathogenesis of BAI [1, 10, 17]. Bioflm formation is initiated when planktonic bacterial cells attach to the surfaces of implants (Fig. 1). BAI are often caused by bioflm-forming bacterial strains able to cover the surface of the biomaterial, resulting in complex structures consisting of bacteria, extracellular polymeric substances (bacterial products like polysaccharides, proteins and DNA) and host proteins and cells [17]. Bacteria in bioflms behave differently from planktonic bacteria, particularly in response to antibiotic treatment [18]. The complex bacterial community of a bioflm is highly tolerant to antibiotics [19]. This is partly due to the complicated structure of the extracellular polymeric matrix of the bioflm, making the bacteria less accessible to many antibiotic agents [20]. As most antibiotics target active cell processes, the slow growth or starved state of the bacteria in a bioflm may also make them more tolerant. A subpopulation of these bacteria, the so-called persisters, reaches a dormant and drug-tolerant state. Such persisters are suggested to be largely responsible for the recalcitrance and recurrence of bioflm-associated infections [21]. Moreover, bioflm-entrapped bacteria are unreachable for the human immune system.

**Fig. 1** Bioflm stages in biomaterials

#### *1.2 Tissue Colonization*

Next to bioflm formation, another important element in the pathogenesis of BAI is bacterial colonization of the tissue around implants (Fig. 2), due to dysregulation of the local immune response by the combined presence of bacteria and a foreign body [22–25]. Bacteria are inevitably introduced in the tissue wound during surgery, either originating from the patient's skin microfora or from the operation room [26]. Due to the implanted biomaterial, the effcacy of the host immune response is reduced. Already in the 1980s, Zimmerli *et al.* showed reduced neutrophil phagocytic activity in guinea pig tissue cage models infected with *S. aureus* [27]. When different challenge doses of *S. epidermidis* were injected along subcutaneously implanted catheter segments at the back of mice, the bacteria were more often found in the peri-implant tissue than on the biomaterial itself, and persisted for longer periods in the tissue than on the implant [28]. Moreover, *S. epidermidis* survives inside macrophages in tissue surrounding implants in mice (Fig. 2) [25, 28].

In a mouse subcutaneous BAI model, the possible routes of infection at the interface between implants and the surrounding tissue were studied [29]. In this study, *S. epidermidis* bacteria applied on the surface of titanium implants, both adhering and as a bioflm, relocate from the material to the surrounding tissue (Fig. 2), which is accordance with earlier studies with other types of materials [25, 28]. This suggests that it is a more general phenomenon occurring around implants manufactured from biomaterials as diverse as polymer and titanium, and with different bacterial species. In a study by Broekhuizen *et al.*, mice were treated with dexamethasone and BrdU, a nucleotide analogue that is incorporated into DNA of dividing cells and can be detected immunohistologically. Analysis of tissue samples collected at 14

**Fig. 2** Pathogenesis of the biomaterials associated infection

and 21 days after challenge with *S. epidermidis* showed regrowth of the bacteria with BrdU incorporated, which had apparently replicated between day 14 and 21, suggesting that tissue rather than the implant provides a hiding place for the bacteria [30]. Moreover, after incubation of peri-catheter tissue biopsies of deceased intensive care unit patients with BrdU, bacteria had incorporated BrdU *in situ*, proving that bacteria also reside and synthesize nucleic acids within tissue surrounding biomaterials in humans [30].

Bacteria colonizing the surface of a biomaterial not only are a focus of a localized bioflm infection, but can also be the source of tissue colonization (Fig. 2). Conversely, bacteria residing in the tissue can be a cause of infection after reimplantation, in experimental infection [31] as well as in patients [32].

Tissue-residing bacteria can be hard to eradicate by antibiotic treatment [33, 34]. For instance, when infected prosthetic joints are removed, patients usually require a prolonged regimen of systemic and local antibiotic treatment in order to reach and kill bacteria present in the tissue before re-implantation can be performed [6, 35]. In conclusion, next to the prevention of bacterial colonization of the implant and the subsequent bioflm formation, prevention of bacterial colonization of peri-implant tissue is of vital importance.

#### *1.3 Intracellular Survival*

In the subcutaneous mouse BAI model staphylococci predominantly co-localized with macrophages in the peri-implant tissue, even when the bacteria were present exclusively on the implant surface at the start of the experiment (Fig. 2) [29]. This interesting observation suggests that the bacteria were either removed from the implant by phagocytosis, or frst detached and were subsequently phagocytosed. In this mouse model, both *S. epidermidis* [29] and *S. aureus* [36] were cultured in high numbers from the tissue and co-localized with macrophages in histology, particularly at 4 days after challenge, suggesting that these macrophages were not effectively killing the bacteria. Most likely, the local host immune response is impaired in presence of an implant, resulting in less or no clearance of bacteria. As mentioned before, neutrophils can have reduced phagocytic and bactericidal capacity in the vicinity of an implant [27, 37]. Moreover, the intracellular killing capacity of macrophages can be reduced due to altered cytokine tissue levels due to the presence of a biomaterial [25, 30, 37–39]. Staphylococci may even form small colony variants to adapt to this micro-environment, which are more resistant to antimicrobial compounds [40, 41]. Apparently, when bacteria are initially present near or on the surface of implants this results in ineffective eradication by phagocytes. This might lead to persistence of (intracellular) bacteria in the peri-implant tissue.

#### *1.4 Antimicrobial Resistance*

In addition to the diffculty of treating bioflm-encased or intracellularly residing bacteria with conventional antibiotic therapy, treating BAI is further hindered by the rising antibiotic resistance among pathogens. The World Health Organization recently endorsed a global action plan to tackle antibiotic resistance [42]. One of the key objectives of this plan is to develop novel antimicrobial drugs. The emergence of multidrug-resistant (MDR), extensively drug-resistant (XDR) and pandrugresistant (PDR) pathogens, accelerated by the selective pressure exerted by extensive use and misuse of antimicrobials, further underscores the very pressing need for the discovery of novel treatment strategies to replace or complement the conventional antibiotics. Magiorakos *et al.* defned MDR bacteria as non-susceptible to at least one agent in three or more antimicrobial categories, XDR bacteria as nonsusceptible to at least one agent in all but two or fewer antimicrobial categories, meaning bacterial isolates which remained susceptible to only one or two categories, and PDR bacteria as non-susceptible to all agents in all antimicrobial categories [43]. The occurrence of XDR and PDR strains illustrates the clinical challenges that we will be facing in the dark scenario of a possible "post-antibiotic era". Antimicrobial resistance causing limited or no treatment options in critically ill patients, stresses the importance of the development of new agents that can be used against drug-resistant bacteria. Clearly, it is vital that novel antimicrobial agents are also effective against drug-resistant Gram-negative bacteria belonging to the socalled ESKAPE panel (*Enterococcus faecium*, *S. aureus*, *Klebsiella pneumoniae*, *Acinetobacter baumannii*, *Pseudomonas aeruginosa*, and *Enterobacter* species [44]), which cause the majority of US hospital infections [45] and are associated with high morbidity and mortality [46].

#### **2 Preventive Strategies**

As explained above, in addition to bioflm formation on the implant, colonization of peri-implant tissue is an important factor in the pathogenesis of BAI. Therefore, this niche needs to be taken into consideration when designing preventive strategies against BAI. Current strategies mainly focus on the development of four types of antimicrobial surfaces: (1) antifouling/anti-adhesive surfaces, (2) tissue-integrating surfaces, (3) contact-killing surfaces, and (4) surfaces which incorporate and release antimicrobials (Fig. 3) [47]. These approaches all have their benefts and limitations, which need to be taken into account when designing an antimicrobial strategy for a particular device [48].

#### *2.1 Anti-adhesive*

Implant surfaces are ideal substrates for opportunistic bacteria to attach to, colonize, and form bioflms on. Surface properties of the implant, like surface charges, hydrophobicity/hydrophilicity and surface chemistry play a major role in initial bacterial adhesion and proliferation. Already in 1987, Gristina suggested that tissue cell integration and bacterial adhesion compete for a spot on the implant's surface, summarized as the so-called 'race for the surface' concept [49]. In case the bacteria win this race, infection instead of tissue integration would be the end result. In addition,

**Fig. 3** Antimicrobial functionality in implant surface

Gristina also suggested that colonization of the tissue around implants was a possible mechanism of infection [49]. Bacterial adhesion and subsequent bioflm formation may be prevented by modifying the physicochemical surface properties of biomaterials, for instance by using hydrophilic polymer coatings, *e.g.* immobilized poly(ethylene glycol) (PEG), as applied on contact lenses, shunts, endotracheal tubes and urinary catheters [47, 50]. Functionalization of the surface with a dense layer of polymer chains commonly known as polymer brush coatings, is another approach [34, 51]. Large exclusion volumes of tethered polymer chains result in surfaces diffcult to approach by proteins or bacteria, and these brush coating molecules may even possess antimicrobially active functional groups.

#### *2.2 Antibiotics*

In general, antibiotics are selected based on their capacity to prevent bioflm formation, but not on their ability to kill bacteria in the other niches relevant for BAI, like in peri-implant tissue and intracellularly in host cells [47]. Antibiotics often used in the treatment of BAI, such as vancomycin and gentamicin, have low or hardly any penetration into host cells, and are thereby not active against intracellular bacteria. On the other hand, rifampicin (against staphylococci) or fuoroquinolones (against Gramnegative bacilli) do target these intracellularly localized bacteria, but resistance develops rapidly against these antibiotics. The combination of vancomycin and rifampicin is often used to treat BAI, but—as vancomycin does not reach intracellular bacteria this likely results in a high risk of resistance development towards rifampicin.

Coatings releasing antibiotic are widely used for medical devices, like in sutures and central venous and urinary tract catheters. These coatings have two major disadvantages: (1) a patient can be infected with a bacterium resistant to the released antibiotic, and (2) due to the local release a gradient of the antibiotic will be created near the implant, which increases the risk to select for resistant bacteria. In view of the increasing development of resistance, the use of antibiotics for medical device is discouraged by government regulatory agencies like the American Food and Drug Administration (FDA) [48, 52].

#### *2.3 Antiseptics*

As an alternative to antibiotics, commonly used antiseptics and disinfectants may be used, as they are less known to induce resistance and in general have a broader spectrum of activity than antibiotics. These biocides, such as alcohols, aldehydes and biguanides, are extensively used in hospitals and other health care settings, and also by the general public, as an essential part of infection control practices [53]. Probably the most widely used biocide in antiseptic products (*e.g.* hand wash and oral products) is chlorhexidine, owing to its broad spectrum activity, low toxicity

and good tolerability of soft tissue. Moreover, resistance development is extremely rare and chlorhexidine has been shown to prevent infection in animal models [36] and in patients [54]. It is used topically, for surgical site preparation, and also intracorporeally [55], and as dental irrigant fuid [56]. Chlorhexidine is currently FDA approved for coatings on intravenous catheters, and these catheters have been shown to be effective in decreasing catheter-related infection in humans [57, 58].

#### *2.4 Antimicrobial Peptides*

As discussed earlier, due to the major problems arising from resistance to conventional antibiotics, there is a strong need for antimicrobials not associated with resistance development. Antimicrobial peptides (AMPs) are innate defence molecules of animals, plants and microorganisms. These amphipathic, cationic peptides commonly have antimicrobial activity against a wide variety of pathogens, including bacteria, fungi and viruses, and low risk of resistance development [59, 60]. In addition, many AMPs have immune-modulatory and wound healing activities [61]. The low risk of resistance development is due to the fact that AMPs interact with microbial membranes, mostly resulting in membrane depolarisation, permeabilization and/or disruption leading to rapid cell death, or passing of the membrane to reach intracellular targets [62]. Naturally occurring human AMPs are considered excellent templates for the development of novel synthetic antimicrobials. Indeed, native AMPs have been used as design templates for a large variety of synthetic AMPs, some of which have now entered phase 2 and 3 clinical trials [63, 64].

For biomaterials, the predominant AMP-related antimicrobial strategies are coating by tethering AMPs to the surface, or to apply the peptides in controlled release coatings. Immobilisation of AMPs on surfaces has been performed with a variety of peptides, and with many different chemistries [65–68]. Peptides should retain the structural characteristics important for their antimicrobial activity after immobilisation, to be effective on a surface. Length, fexibility, and kind of spacer connecting the peptide to the surface, the AMP surface density and the orientation of the immobilised peptides are other decisive factors for success [69]. Interestingly, even short surface-attached peptides, which are unlikely to have a free interaction with the bacterial membrane, have antimicrobial activity [70], probably due to destabilisation of the membrane by displacement of positively charged counter-ions, changing bacterial surface electrostatics and activating autolytic enzymes or disrupting the ionic balance [70].

Surface attachment of peptides may have certain disadvantages. Firstly, chemical procedures of tethering AMPs to surfaces may cause strong decrease in their antimicrobial activity, or even their inactivation [71, 72] depending on the combination of peptides and immobilization technology. Secondly, proteins, blood platelets and dead bacteria may block the antimicrobial groups on the surface. Lastly, since the antimicrobial activity is restricted to the surface of the implant, there is a lack of antimicrobial impact on bacteria in the tissue surrounding the implant.

Incorporation of AMPs in controlled release coatings has not yet been extensively developed, although AMPs such as OP-145 [73], IB-367 (Iseganan) [74] and Omiganan [75] have already reached clinical phase 2 or 3 testing for infections not associated with biomaterials [64]. Application of AMPs in antimicrobial surface coatings is however a subject of increasing interest [65–67, 76, 77].

In addition to direct antimicrobial activity, AMPs can prevent excessive activation of pro-infammatory responses by binding bacterial endotoxins such as lipopolysaccharide (LPS) of Gram-negative bacteria, and peptidoglycan (PG) and lipoteichoic acid (LTA) of Gram-positive bacteria, which leads to their neutralization. This way, AMPs combine the desired characteristics of both direct antimicrobial agents and immune-modulators. The immunomodulatory activity may be used to increase effcacy of clearance of bacterial bioflm infection [78, 79], and might help to prevent derangement of immune responses which increase susceptibility to infection [22, 80, 81].

#### **3 Conclusions and Future Perspective**

Prevention of BAI is a challenging problem, in particular due to the increased risk of resistance development associated with current antibiotic-based strategies. Here we showed the evidence of bioflms as a source for peri-implant tissue colonization, clearly showing the importance of preventive measures to be able to act both against implant and tissue colonization. Subsequently, we described different strategies to prevent BAI and other diffcult-to-treat bioflm infections. Therefore we conclude that future research should focus on the development of combination devices with both anti-fouling or contact-killing capacities—to protect the implant—and controlled release of an antimicrobial agent to protect the surrounding tissue.

#### **References**

**Open Access** This chapter is licensed under the terms of the Creative Commons Attribution 4.0 International License (http://creativecommons.org/licenses/by/4.0/), which permits use, sharing, adaptation, distribution and reproduction in any medium or format, as long as you give appropriate credit to the original author(s) and the source, provide a link to the Creative Commons license and indicate if changes were made.

The images or other third party material in this chapter are included in the chapter's Creative Commons license, unless indicated otherwise in a credit line to the material. If material is not included in the chapter's Creative Commons license and your intended use is not permitted by statutory regulation or exceeds the permitted use, you will need to obtain permission directly from the copyright holder.

# **Antibiotic-Free Solutions for the Development of Bioflm Prevention Coatings**

**Bruna Costa, Joana Barros, and Fabíola Costa**

#### **1 Introduction**

Stents and urinary catheters are commonly used medical devices, whose need is forecasted to grow considering not only the world population increase but also its aging and sedentary lifestyle [1].

Independently of the great development on biomaterials and device design, infection represents still a major cause of failure of these devices, with undeniable humane and economical costs. Different antibiotic-based solutions have appeared in the market to try to address the matter. However, there is growing evidence on the impact of antibiotic-resistant microorganisms on urinary tract medical-devices infections, and respective outcomes [2]. Within this chapter, several antibiotic-free alternatives, dedicated to the urinary tract, will be discussed.

Most device associated-infections are originated by bioflm establishment. Bacterial colonization through irreversible attachment, allows the production of extracellular matrix, forming ultra-organised three-dimensional bacterial structures, with orchestrated phenotypes that provide microorganisms resistance mechanisms to survive both the immune system and conventional antibiotics [3]. From the knowledge of these bacterial constructs, researchers have been exploring different angles of action that enforce the balance towards the infection obliteration and host recovery (Fig. 1).

J. Barros · F. Costa (\*)

B. Costa

i3S – Instituto de Investigação e Inovação em Saúde, Universidade do Porto, Porto, Portugal

INEB – Instituto de Engenharia Biomédica, Universidade do Porto, Porto, Portugal

FEUP–Faculdade de Engenharia, Universidade do Porto, Porto, Portugal e-mail<EMAIL_ADDRESS>
i3S – Instituto de Investigação e Inovação em Saúde, Universidade do Porto, Porto, Portugal

INEB – Instituto de Engenharia Biomédica, Universidade do Porto, Porto, Portugal e-mail<EMAIL_ADDRESS>
#### **2 Natural Polymer Coatings**

Apart from the different synthetic polymers with non-fouling properties described in previous chapters, naturally-occurring polysaccharides are presently being explored to design environmentally friendly and non-toxic materials. Several examples can be found in the literature, ranging from hyaluronic acid, heparin to ulvan or dextran [4–10], with potential for urinary applications. Gadenne et al., explored ulvans, with different molecular weight and sulfate ratio, for bacterial adhesion inhibition. Ulvans inhibited 36–88% of *Staphylococcus aureus* adhesion comparing to control [7]. Ruggieri et al., showed that latex catheters coated with a complex of heparin with tridodecylmethylammonium chloride were capable of reducing 53–84% *Escherichia coli* adherence comparing to controls: untreated latex, tefon coated latex (Bardex) and vinyl catheter [10]. Tenke et al., performed a 20 patients pilot assay and claimed that heparin-coated ureteral stents remained unaffected by encrustations and bioflm after 6 weeks [11]. Later, Lange et al., showed that heparin-coated Radiance © ureteral stents (Cook® Medical) failed in the prevention of *E. coli*, *Klebsiella pneumoniae*, *Enterococcus faecalis*, *S. aureus* and *Pseudomonas aeruginosa* adhesion and bioflm formation, while triclosaneluting stents had an evident inhibitory effect on bacterial adherence for 7 days [8]. So further studies are needed to conclude about heparin coatings potential. A copolymer of polyurethane with dermatan sulfate (DS) was developed as new non-adhesive material, showing a signifcant *E. coli* adhesion decrease (29–57%) with increasing DS content [12].

Carboxymethyl chitosan was explored as an antimicrobial coating onto medical-grade silicone. Higher anti-bioflm effciency was found against *E. coli* than *P. mirabilis* under fow-conditions. This effect can be explained by *P. mirabilis* high motility, which favors bioflms establishment downstream of an infected site [13]. Bracic et al., evaluated the anti-bioflm properties of colloidal polysaccharide complexes [chitosan, carboxymethyl chitosan, and hyaluronic acid in combination with a lysine-based surfactant (HA-MKM)] grafted on silicone sheets and tubes. All coatings showed antibacterial and antifungal properties, being HA-MKM the only solution capable of suppressing bioflm growth by ~ 50–75% during 18 h [14].

Recently, the anti-adhesive potential of cyanobacteria-based polymer coating (CyanoCoating) was reported against *Proteus mirabilis*, *E. coli*, *P. aeruginosa*, *K. pneumoniae*, Methicillin-resistant *Staphylococcus aureus* (MRSA) and *Candida albicans*. CyanoCoating hydrophilicity, negative charge and smooth surface may explain its broad anti-adhesive effciency against all the uropathogens tested (68–95%), even in the presence of artifcial urine (58–100%) [4]. Also, this antiadhesive coating prevented big crystals deposition, reducing encrustation. CyanoCoating could also withstand ethylene oxide sterilization [5].

Biosurfactants represent an alternative strategy to promote anti-adhesiveness to the surfaces, which is thoroughly explained at Chap. 20.

#### **3 Metal Alternatives**

Metal alternatives such as silver (Ag), gold nanoparticles, copper oxide (CuO) or zinc (Zn), have been explored for urinary medical devices. The use of Ag or Ag alloys has been broadly exploited, having a wide expression in the market.

Gold nanoparticles antimicrobial effect is associated to bacterial membrane potential disruption, ATP levels reduction and tRNA inhibition [15]. Gold nanoparticles have been tested against important uropathogens, including *S. aureus*, *K. pneumonia*, *P. aeruginosa*, and *E. faecalis*, suppressing their bacterial growth at 24 h. However, the antimicrobial effectiveness diminished when used at longer-term, raising concerns about possible gold-resistance emergence [16].

Copper (Cu) promotes bacterial DNA degradation, enzymes inactivation and cell wall disruption [17–19]. Agarwala et al., tested CuO against *E. coli*, *P. mirabilis*, *E. faecalis*, *Pseudomonas* sp., MRSA and *S. epidermidis*, showing promising antibioflm activity even at sub-Minimum Inhibitory Concentrations (MIC) [20]. Rtimi et al., incorporated Cu alone or in combination with Ag onto polyurethane catheter surfaces using a new magnetron sputtering coating technique [21]. Cu–Ag hybrid coating catheters accelerated *E. coli* K12 inactivation (≤ 5 min) compared to Cu or Ag coating catheters (30 min) [21].

Zinc antimicrobial activity has been associated with hydrogen peroxide production [22]. Zn has been combined with CuO to mitigate bacterial colonization [17]. Shalom et al., showed that Zn-doped CuO nanoparticles coated catheters reduced *E. coli*, *S. aureus*, and *P. mirabilis* bioflm formation (> 90%) under fow conditions for 24 h [23]. Moreover, these coated catheters prevented bioflm formation over 7 days in a catheter-associated urinary tract infection (CAUTI) rabbit model [23].

Despite, the promising antimicrobial effects, well-designed toxicity and irritation studies are still needed.

#### **4 Chlorhexidine**

Chlorhexidine (CHX) is a well-known antiseptic agent used for skin, dentistry, and in medical devices [24]. CHX is broad-spectrum bacteriostatic at low and bactericidal at high concentrations [24]. Recently, CHX has also been tested as a coating on urinary catheters [25–27]. Shapur et al., explored a CHX-releasing ethylcellulose varnish as antimicrobial coating, showing a 94% reduction of *P. aeruginosa* bioflm formation on catheters coated with 1% CHX [26]. Later, Segev et al., proved the anti-bioflm effectiveness of 1% CHX ethylcellulose-varnish coated urinary catheters using a dog model [25]. Zelichenko et al., evaluated growth inhibition on ureteral stent segments coated with 1% and 2% CHX, showing that 2% CHXvarnish prevented ≥ 99.9% bioflm formation of *E. faecalis*, *P. aeruginosa* and *E. coli* up to 2 weeks [28]. Gefter et al., compared anti-bioflm properties and dissolution kinetic of two sustained-release CHX-varnishes (ethylcellulose or Eudragit® RL) under the static or fow-conditions [27]. In both situations, ethylcellulose coatings had longer sustained release of CHX (for at least 2 weeks), which resulted in an inhibition of ≥ 90% *P. aeruginosa* bioflm formation at 24 h [27]. Wood et al., developed a CHX hematophosphate nanoparticles (NP) ethylene-vinyl acetatebased coating [29]. The NP-coated surfaces inhibited MRSA and *P. aeruginosa* growth (measured at 24 h), and allowed for CHX sustained release over 56 days [29]*.* Phuengkham et al., spray-coated CHX-loaded polycaprolactone nanospheres onto silicone surface, reducing *S. aureus* (3 logs), *S. epidermidis* (2 logs) and *E. coli* (3 logs) bioflm formation over 7 days [30]. Then, Srisang et al., using the same coating reported 4 logs of reduction after 4 days, and 2 logs after 12 days of *E. coli*, *S. aureus*, and *C. albicans* tested in artifcial urine [31]. Gaonkar et al., compared *in vitro* three different impregnated silicone catheters on urinary tract model: CHX– triclosan, CHX–Ag–sulfadiazine–triclosan, and nitrofurazone-coated catheters. CHX–triclosan catheter suppressed *P. mirabilis* growth for 20–30 days, compared to 4–10 days observed on the CHX–Ag–sulfadiazine–triclosan or nitrofurazonecoated catheters [32].

Despite, the extended protection period and promising antimicrobial effects, well-designed toxicity and safety studies are desirable to validate these coatings safety to patients.

#### **5 Triclosan**

Triclosan was the frst compound to be approved for clinical use in ureteral stents, having potent broad-spectrum antimicrobial and anti-infammatory activity. Cadieux et al., frst tested triclosan impregnated ureteral stents on a *P. mirabilis* rabbit urinary tract infection model: 69% triclosan-stents showed no CFU counts and the remaining 31% had fewer CFU than controls. Also, triclosan group presented bladders with signifcantly less infammation, although no signifcant difference in encrustation was observed among the groups [33]. However, in a long-term application (3 months) no clinical beneft was observed in terms of urine and stent cultures or overall subject symptoms in triclosan-eluting stents patients. Nevertheless, their use did result in decreased antibiotic usage and fewer symptomatic infections [34]. Later, a prospective randomized trial, reported that triclosan-eluting stent cannot reduce infection rates alone compared with antibiotic use [35]. This stent can, however, reduce several stent-related symptoms, and may have a role in combination with standard antibiotherapy.

#### **6 Antimicrobial Peptides**

Antimicrobial peptides (AMPs) may constitute an alternative to fght antibiotic resistance [36–39]. AMPs are part of the innate immune system of many organisms, having broad-spectrum, high anti-bioflm activity, and even immunomodulatory potential [40–43]. Due to an unspecifc mode of action, targeting the bacterial membrane or affecting multiple targets within bacteria, AMPs are much less likely to induce resistance [44]. Monteiro et al., immobilized the AMP Chain201D on model self-assembled monolayers [45]. Chain201D has broad antimicrobial activity against relevant uropathogens (bacteria and yeast), being highly stable in a wide range of temperatures, pH and salt concentrations [45]. Increased amounts of grafted AMP led to higher numbers of adhered/dead bacteria, revealing a concentration-dependent behavior. Chain201D surfaces could bind and contact kill 89% of *E. coli* and 99% of *S. aureus* adherent bacteria, suggesting a good candidacy for urinary applications [45].

Minardi et al., compared the effcacy of the AMP Tachyplesin III-coated ureteral stent alone or combined with piperacillin–tazobactam (TZP) intraperitoneal injection in the prevention of *P. aeruginosa* bioflm in a rat model. Tachyplesin III combined with TZP showed effcacies higher (3 logs of reduction) than each single therapy [46].

Lim et al., conjugated an engineered arginine–tryptophan rich AMP (CWR11) onto polymethylsiloxane (PDMS) surfaces and catheters through different chemistries [47, 48]. The CWR11-PDMS slides displayed excellent bactericidal effect against *E. coli*, *S. aureus* and *P. aeruginosa*, preventing ~ 92% *P. aeruginosa* bioflm formation after 24 h. The CWR11-silicone Foley catheter antimicrobial properties were retained for at least 21 days, with negligible cytotoxicity [47].

Li et al., grafted two broad-spectrum and salt-tolerant arginine/lysine/tryptophanrich AMP (RK1 and RK2), onto PDMS surfaces via an allyl glycidyl ether (AGE) polymer brush interlayer. AMP-PDMS killed over 80% of *E. coli*, *S. aureus*, and *C. albicans* in either media, PBS or urine, and impaired bioflm formation for up to 3 days [49].

Mishra et al., developed a Lasioglossin III AMP chemically modifed with a cysteine residue (CysLasio-III) to selectively immobilize covalently onto commercial silicone catheter [50]. CysLasio-III-catheter showed signifcant anti-bioflm properties, reducing 40% and 60% of *E. coli* and *E. faecalis* bioflm, respectively [50].

Lo et al., used polymer brushes to graft different AMPs (E6, Tet20, Kai13, and Tet26) to surfaces. *In vitro* tests revealed E6 was the most effective against *P. aeruginosa,* decreasing ~ 94.1% of bacterial adhesion [51]. Later, Yu et al., similarly grafted E6 on polyurethane tubing, reporting a > 4 logs reduction in *P. aeruginosa* adhesion to the tube and 3 logs in the bladder in a CAUTI mouse model [52].

Pinese et al., developed a silylated analogue of the AMP Palm–Arg–Arg–NH2 [1], to directly graft onto a plasma-activated PDMS catheter [53]. The authors suggested a dual anti-adhesive/bactericidal effect of the coating, since a decrease of ~ 75% *E. coli*, *P. aeruginosa* and *S. aureus* adhesion was observed and 92% of bacteria were killed on peptide-grafted catheters within 1 h. This AMP-catheter was superior to a commercial Ag-based silicone catheter (Covidien) against *S. aureus,* with earlier and persisting activity over 2 weeks [53].

Chua et al., compared an AMP CP11-6A-coated silicone catheter to an Ag-hydrogel-coated and an uncoated catheter using *E. coli* inoculated human urine. Within 3 days, both uncoated and Ag-coated catheters were heavily colonized, while CP11-6A-coated catheter showed negligible bioflm colonization and no detectable "bacteriuria" [54].

Although progress has been made, and many AMP-based coatings have impressive antimicrobial activity, further studies are needed to establish clinical signifcance.

#### **7 Nitric Oxide**

Nitric oxide (NO) has been used as an antimicrobial, showing great potential for biomedical applications [55], although poorly explored for urinary devices [56– 58]. NO covalently binds DNA, proteins and lipids, thus inhibiting or killing pathogens [59]. NO-donating polymers may provide localized treatment with minor toxicity. Nevertheless, further studies are needed to understand the NO effects in the bladder, since it is known that NO plays an important role in other biological conditions (e.g. vasodilatation, neurotransmission) [56]. Regev-Shoshani et al., showed that gaseous NO-impregnated catheters have a sustained NO release over 14 days with stable storage. The NO release was faster in urine than in water, suggesting pH infuence in the release, which might have implications at patient level [56]. Colletta et al., developed *S*-nitroso-*N*-acetyl-dpenicillamine impregnated Foley catheters with outstanding anti-bioflm effect, reducing *S. epidermidis* (3.7 logs) and *P. mirabilis* (6 logs) bioflm formation after 14 days [60]. Later, Ketchum et al., applied tertiary *S*-nitrosothiol and *S*-nitroso*tert*-dodecyl mercaptan (SNTDM) as NO donors onto catheter tubings. NO-tubings reduced *S. aureus* colonization (4 logs) on SNTDM-impregnated catheters at 1 week, maintaining high anti-bioflm effcacy (3 logs of reduction) even after 3 weeks [61].

#### **8 Quorum-Sensing Disrupters**

Quorum-sensing (QS) corresponds to a cell–cell communication process, based on signaling molecules secreted by adhered bacteria to determine if a suffcient number of microorganisms is present that can initiate the expression of a particular bioflmassociated phenotype [62]. Quorum quenching, can severely hinder bioflm formation, diminishing bacteria antimicrobial resistance [63].

Ureteral stents coated with QS-inhibitor RNAIII-inhibiting peptide (RIP) reduced *S. aureus* adhesion (2 logs) to stents implanted in rat bladders. No bacteria were detected either on the stent or urine when the peptide treatment was combined with teicoplanin (which achieved only 3 logs reduction in single teicoplanin therapy) [64].

A combination of alpha-amylase and acylase was tested as a layer-by-layer coating on silicone urinary catheters [65, 66]. Alpha-amylase interferes with assembly of the extracellular matrix and acylase degrades small quorum signaling molecules. When tested *in vitro*, this coating reduced *P. aeruginosa* (> 40%), *S. aureus* (> 30%) and mixed-species bioflms (> 50%), although planktonic growth was not inhibited. *In vivo* (rabbit model) bioflm formation on the catheter's balloon section was decreased by 90%, although this was not seen in the lumen of the catheter. Furthermore, the quorum quenching action of single acylase, reduced signifcantly *P. aeruginosa* bioflm formation on a catheter under static and dynamic conditions [66].

Recently, furanone, a QS-inhibitor, was also tested as a coating for urinary catheters (latex, silicone and polyurethane), preventing *Candida* sp. bioflm formation in more than 80% [63].

#### **9 Extracellular Matrix Degrading Enzymes**

Exopolysaccharides are a crucial component of bioflm architecture, providing protection against drugs and host immunity [67]. Therefore, enzymes capable of degrading one of bioflm matrix components (proteins, extracellular DNA, polysaccharides), would impact on bioflm establishment [68].

Alpha-chymotrypsin (α-CT), a serine endopeptidase that cleaves peptide bonds [69, 70], was grafted on polyethylene surfaces, signifcantly reducing adherent cells, affecting the bioflm thickness, roughness and coverage. Additionally, the biovolume of the polysaccharide matrix decreased [70].

Also, recombinant human DNase (rhDNase), has effciently inhibited *S. aureus* bioflm formation at 1–4 μg/L [71].

Cellobiose dehydrogenase (CDH), an oxidative enzyme that produces hydrogen peroxide, was grafted CDH onto PDMS catheters by ultrasonic waves, reducing *S. aureus* growth (60%), biomass deposition (30%), and bioflm production (70%) when compared to control catheters [72]. Thallinger *et al.*, tested CDH-grafted urinary catheter in artifcial urine, over 16 days, observing a 60% reduction of the viable *S. aureus*, and a 70% bioflm formation decrease, comparing to control [73].

Glycoside hydrolases (Ghs) selectively target and hydrolyze the glycosidic bonds of exopolysaccharide components of the bioflm matrix [74, 75]. Asker et al., used Ghs that specifcally targets Psl, a neutral exopolysaccharide, grafted to silica glass, PDMS, or polystyrene surfaces. PslGh-grafted surfaces reduced adhered *P. aeruginosa* in 3 logs up to 8 days, suggesting this strategy keeps bacteria in a planktonic state more susceptible to antimicrobials [76].

More studies might further validate this strategy, alone or in combination with other antimicrobials [75].

#### **10 Bacteriophages**

Bacteriophage (phage) therapy is proposed as a safe and effective strategy to address bioflms and multidrug-resistant pathogens, without impairing the resident microbiota [77, 78].  Phages are self-replicating at infection sites, producing new phage progeny that can migrate to a new focus of infection. Additionally, phages encode enzymes that degrade bioflm matrix [77, 78, 80]. Phages-impregnated catheters have been used against common uropathogens [78, 80–83].

Curtin et al., reported a signifcant reduction of coagulase-negative *S. epidermidis* bioflm formation (2.34 logs) over 24 h on hydrogel Foley catheter impregnated with phage 45,682. Later, Carson et al., showed that T4 or coli–proteus phages impregnated hydrogel-coated Foley catheters were able to reduce 90% of *P. mirabilis* or *E. coli* bioflm formation over 24 h, respectively [78].

Lehman et al., showed hydrogel catheters pretreated with anti-Pseudomonas and anti-Proteus phage cocktails have anti-bioflm activity against single- (1.5 logs *P. aeruginosa* and 2.5 logs *P. mirabilis*) and dual-species (4 logs and 2 logs, respectively) bioflms, over 48 h [77]. Previously, the same group reported that anti-Pseudomonas phage cocktail catheters is more effective than single phage-loaded catheters [82].

Milo et al., showed that the dual-layered polymeric coating, based on PVA hydrogel impregnated with phage, capped by a pH-sensitive polymer (EUDRAGIT S 100), was able to prevent *P. mirabilis*-associated encrustation of the catheter lumen through the pH-triggered release of phage [80].

Liao et al., compared the effcacy of 4 silicone catheter segments pretreated with sterile media, *E. coli* HU2117, anti-pseudomonal phage (ΦE2005-A) and *E. coli* HU2117 plus phage ΦE2005-A, on prevention *P. aeruginosa* bioflm formation for 72 h. A synergistic effect between pre-established bioflm of *E. coli* HU2117 and phage ΦE2005-A was observed, reducing effciently (4 logs after 24 h and 3 logs after 72 h) *P. aeruginosa* adherence to catheters [83].

#### **11 Conclusions**

The urinary tract device-associated infections prevalence and the rise of multidrugresistant microorganisms have prompt researchers towards the development of antibiotic-free solutions. A broad number of alternatives have been proposed, however, given the wide variability of results for different strategies, there remains a tremendous need to validate their clinical signifcance, particularly assuring patient safety. Additionally, most of these strategies might be advantageous while in combination with current therapies, so further studies are needed.

#### **References**

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# **Plasma Based Approaches for Deposition and Grafting of Antimicrobial Agents to Polymer Surfaces**

**Todorka Gancheva Vladkova and Dilyana Nikolaeva Gospodinova**

## **1 Introduction**

Improved protection of urinary stents against infections is a signifcant current challenge because of the increasing microbial resistance to the conventional antibiotics and negative issues for the patients. Formation of crystalline bioflms of pathogenic microbial cells is the leading cause of urinary stent associated infections.

On many parameters polymeric materials satisfy the basic requirements and are widely used for the fabrication of urinary stents, silicones and polyurethanes being preferable ones currently [1]; and biodegradables attracting interest lately. However, the nonsuffcient microbial resistance, bioflm formation and encrustation are their common gap.

A lot of approaches, antimicrobial agents and techniques are under a study to mitigate the problem by creation of contact killing, releasing or low adhesive surfaces do not allowing attachment of microbial cells [2–4].

The plasma treatment has a number of advantages that make it preferable in many strategies for the development of antimicrobial biomaterials. The control over the plasma processing parameters allows control over the surface chemistry, charge, structure, morphology, hydrophilic/hydrophobic balance, etc. Due to a variety of biomaterials and bacteria, causing urinary tract infections, plasma assisted antibacterial strategies need in tailoring to each specifc surface [5–8].

D. N. Gospodinova

T. G. Vladkova (\*)

Department of Polymer Engineering, University of Chemical Technology and Metallurgy, Sofa, Bulgaria e-mail<EMAIL_ADDRESS>
Department of Electrical Apparatus, Faculty of Electrical Engineering, Technical University of Sofa, Sofa, Bulgaria

#### **2 Physical Plasma and Plasma Processes**

Plasma is a multicomponent system obtained by a partial ionization of gas. The plasma consists of positively and negatively charged ions, negatively charged electrons, radicals, neutral and excited atoms, highly energetic molecules and molecular fragments [8–12]. According to the temperature of the ions, the plasma refers to *low temperature* (*cold*) or *high temperature plasma* [13].

Surface modifcation of polymeric biomaterials is performed in low temperature plasma created by ionization of inert (Ar, Ne, He) or reactive (O2, N2, NH3, CO2) gases or volatile monomers at low (RF) or atmospheric pressure by applying energy in the form of heat, direct or alternating electric current, radiation or laser light [8, 12].

Depending on the mode of plasma treatment, four types modifcation processes happen on the polymer surface: sputtering, etching, ion implantation and plasma polymer deposition [12].

*Plasma sputtering* is the plasma physical degradation, limited to the outermost layer of the polymeric biomaterial, as it simplifed presented in Fig. 1.

Plasma sputtering is used for sterilization of sensitive to temperature or radiation biomaterials, removal of surface contaminations and deposition of sputtered thin coatings [12, 14].

*Plasma etching* (Fig. 2) is a process at which the loss of the exposed polymeric material is deeper and the adsorption of energetic species is followed by a product formation, prior to a product desorption.

**Fig. 1** Simplifed presentation of plasma sputtering: the plasma degradation is limited to the outermost layer of the polymeric material

**Fig. 2** Simplifed presentation of plasma etching: the loss of the exposed polymeric biomaterial is deeper and the adsorption of energetic species is followed by a product formation, prior to a product desorption

**Fig. 3** Simplifed presentation of plasma ion implantation: the presenting in the plasma excited species react directly with the polymer surface and induce grafting of new functional groups

Plasma etching is aimed at: removal of impurities, surface nanopatterning, crosslinking of surface polymer chains and generation of surface functional groups [11, 15].

*Plasma ion implantation* is a process at which the presenting in the plasma excited species react directly with the polymer surface and induce grafting of new chemical groups (amine, hydroxyl and others), Fig. 3 [11, 15].

*Plasma polymer deposition* (PPD) is a process in which a thin polymer-like flm is formed over the surface of the substrate polymer (Fig. 4).

PPD happens inside the plasma reactor but outside the plasma zone where activated gas species polymerize onto the cold substrate. The generated flms are

**Fig. 4** Simplifed presentation of plasma polymer deposition: a thin polymer-like flm is formed over the surface of the original biomaterial

commonly referred as plasma polymers although they do not be formally classifed as polymers because they do not consist of repeating monomer units [10, 15]. Using plasma of different gaseous substances: allylamine, octadiene, aldehydes, ethanol, acrylic acid, perfuorooctane, etc.) and optimizing the operation conditions it is possible to create thin coatings with different functional groups and varied properties (hydrophilic/hydrophobic, positively/negatively charged or non-charged, soft or hard, etc.), as it is evident from Fig. 5.

*Plasma-enhanced chemical vapor deposition* (PECVD) is the most common plasma polymerization technique. Magnetron sputtering, liquid-assisted deposition, plasma-assisted thermal evaporation, etc. are other technics to polymer surface modifcation [11, 15–17]. *Ion beam processing* includes: ion implantation; ion beam texturing and sputtering, etc. The ion-beam modifcation is durable and no surface "reconstruction" happens in contrary to the plasma treated surface. Plasma immersion ion implantation and ion treatment by plasma exposure are two relative new possibilities. The above methods could be utilized in antimicrobial approaches to the creation of functionalized, or low adhesive surfaces with controlled topography and surface energy [11, 18]. Some examples of plasma treatments are: ion-plasma modifcation of polyvinylchloride microfltration membranes [19]; treatment of polyacrylonitrile membranes in DBD discharge in air (including magnet stimulated), [20–22]; electron beam cross-linking of silicone rubber [23], plasma pre-treatment of collagen and keratin base materials [24]; generation of self-organized patterns in cold atmospheric plasma-activated liquids [25] etc.

## **3 Plasma Approaches to Antimicrobial Surfaces Development**

Polymer surfaces with antimicrobial properties can be developed either by coating deposition, antimicrobial agent grafting or affecting the surface topography and free energy. These approaches act through distinct mechanisms: contact killing, killing in solution or stimuli responsive killing and bacterial repellence [3, 4, 26]. Plasma strategies for developing of such surfaces are in the focus of many current investigations [27]. They could be referred to the following main groups: deposition of antimicrobial plasma coatings, plasma based surface functionalization and covalent immobilization of antimicrobial molecules.

#### *3.1 Deposition of Antimicrobial Plasma Coatings*

Deposition of plasma coatings decreasing the microbial adhesion, contact killing or releasing antimicrobial agent (passively or in response of external stimuli) is accepted now as a promising strategy to creation of antimicrobial surfaces [4, 8].

Plasma *deposition of nanocomposite coatings*, containing metal or metal oxides nanoparticles (Ag, Cu, Ti, TiO2, etc.), is becoming an important step in the manufacturing of antimicrobial polymeric biomaterials [8]. Silver is one of the most utilized antibacterial components of plasma coatings. It is included in different forms: as phospholipid encapsulated nanoparticles [28]; hybrid silver-poly(*l*-lactide) nanoparticles [29]; polyvinyl-sulphonate-stabilized nanoparticles [30]; as Ag/ SiO(x)C(y) plasma polymer [31]; as (AgNPs)-loaded coatings with a second *n*heptylamine layer [32]; plasma-sprayed silver-doped hydroxyapatite coating [33, 34]; silver-doped diamond-like carbon coatings deposited via a hybrid plasma process [35]; Ag clusters incorporated in a:C (Ag/a:C) matrix produced by plasma gas condensation process [36]; silver/montmorillonite biocomposite multilayers [37], hexamethyldisiloxane nanocomposites [38]; as plasma coated AgNPs [39]; etc. Piszek and Radtke [40] discuss chemical vapor deposition (CVD) and atomic layer deposition (ALD) as a tools for fabrication of silver layers, nanoparticles, and nanocomposites together with the release of silver ions from nanoparticles or nanolayers as well as the antimicrobial activity of these materials.

Vladkova et al. [41] developed new functional coatings for medical devices, employing magnetron co-sputtering to deposit triple TiO2/SiO2/Ag nanocomposite thin flms. Combining the antimicrobial activity of the TiO2 and Ag with the dispersing effect of the SiO2 these coatings demonstrate strong inhibitory effect toward *E. coli* and *P. aeruginosa* growth. Direct contact and eluted silver mediated killing were experimentally demonstrated as mechanism of antibacterial action of these coatings [41]. Kredl et al. [42] use DC plasma air jet to deposit Cu coating on PDMS and acrylonitrile butadiene styrene ABS triple co-polymer surfaces. Good antimicrobial activity against Gram-negative and Gram-positive test bacteria with clinical signifcance was found by Stoyanova et al. [43] for RF magnetron co-sputtered Ag and Cu doped TiO2 coatings. Woskowicz et al. [44] just reported MS-PVD plasma treatment of polypropylene surface utilizing sputtering of Ag, Cu and their oxides in order to impart antimicrobial activity [44].

Deposition of *plasma coatings releasing antimicrobial agents either passively or in response to external stimuli* is another option to limit bacterial colonization. Releasing metal nanoparticles (silver, coper or tin), amino-hydrocarbon coatings, prepared by plasma immersion ion implantation are an example discussed as alternative of the antibiotics releasing ones [12].

Plasma coatings that *release antimicrobials in response to external stimuli* are produced by 'sandwiching' of antimicrobials between two plasma polymer layers, plasma polymer over coating or nano-templating for creation of antimicrobial reservoirs [45–47]. A novel approach to generate hydrogel coatings through atmospheric-pressure plasma polymerization includes: plasma pre-treatment of the substrate leaving reactive surface radicals; plasma-induced polymerization of the monomer units and cross-linking the polymer chains into a polymer network [48].

*Plasma coatings inhibiting bacterial adhesion are* based on the idea for creation of surfaces decreasing microbial adhesion down to levels do not allowing attachment or allowing easy detachment of microbial cells. Non-toxicity is the main advantage of this strategy together with some others. Ykada et al. [49] proved that the work of adhesion in aqueous media, W1,2w approaches to zero when the water contact angle (WCA) or surface tension, γc approaches to zero, i.e. low adhesive are strong hydrophilic or strong hydrophobic surfaces. Surface enrichment of relevant functional groups and topographical modifcations are main ways to creation of such by plasma treatment [4, 11]. When plasma processes are combined with nanotexturing, remarkable wetting states such as superhydrophobicity and superhydrophilicity can be achieved [50–52]. Nwankire et al. [53] deposited superhydrophobic (WCA above 150°) siloxane coatings using atmospheric pressure plasma jet system and hexamethyldisiloxane (HMDSO), tetramethyl cyclotetrasiloxane (Tomcats) or a mixture of Tomcats and fuorosiloxane as liquid precursors [53].

*Diamond like carbon (DLC) plasma coatings* are also used for creation of low adhesive polymer surfaces. A fuid precursor, generally used for production of DLC is hydrocarbon (methane) or silicone. The hydrogen to carbon ratio has a dramatic infuence on the characteristics of the DLC coatings. But overall they characterize by excellent biocompatibility and low friction coeffcient [11, 54]. In 2007, Laube et al. [55] discuss DLC coatings as a new strategy for decreasing the formation of crystalline bacterial bioflms on ureteral stents. A preliminary study with ten patients having indwelled DLC coated stent demonstrates quite promising results: signifcantly decreased friction, encrustation and bioflm formation. It was concluded that further investigation in larger patient groups is necessary for their confrmation. Unfortunately, no reports were found about that [55]. The doping with antimicrobial metals (copper, silver, and other) or elements increasing the hydrophobicity are conventional tools to improve the DLC coatings resistance to bacterial bioflms formation. Ren [56] reports increased resistance to bacterial colonization of anti-adhesive Si-and F-Doped DLC coatings and micro-nanostructured surfaces than non-doped DLC coatings.

Another promising candidate for antimicrobial protection of biomaterial are *oxazoline-derived plasma polymer (PPOx) coatings.* Bacteria may attach in small numbers to the deposited under appropriate conditions PPOx coatings but would not proliferate to form bioflms, that is very interesting for development of low fouling coatings to indwelling medical devices [57, 58]. A simple and effcient strategy for preparation of poly(2-oxazoline)-based coatings on polytetrafuoroethylene (PTFE) substrate, using diffuse coplanar surface barrier discharge (DCSBD) as a cold plasma source was just reported [59].

## *3.2 Plasma-Based Surface Functionalization and Antimicrobial Agents' Immobilization*

Plasma deposited antimicrobial coatings improve the antibacterial activity but the effect is not enough durable for some applications and toxicity concerns remain. Covalent grafting of antimicrobial agents is a subject of many investigations aimed at a long-lasting effcacy and a reduced toxicity [5]. Plasma treatment is an easy way to creation of functionalized surfaces with antimicrobial activity or such that can be utilized for covalent immobilization of antimicrobial agents including on nanoscale topographies [4, 60].

Conventionally, an antimicrobial agent immobilization is carried out in two steps: enriching the topmost polymer surface with reactive functional groups by cold plasma treatment (RF or atmospheric pressure plasma, lately preferable) followed by covalent binding using known chemical reactions [4, 60].

The oxygen-rich surfaces, containing hydroxyl, carboxyl and carbonyl groups promote the cellular attachment because of ionic interactions with molecules, mediating the cell adhesion [61]. The amine-rich polymer surfaces created by plasma treatment, using allylamine, ethylendiamine, propylamine, butlyamine or heptylamine as starting monomer are positively charged and facilitate the electrostatic adsorption of negatively charged proteins, that maybe confer their biocompatibility [62].

Plasma deposition of ring opening monomers (oxazolines, pyrrole, furfuryl, thiophene, aniline, etc.) generates surface chemistries that are not achievable via other ways. Careful tuning the plasma deposition condition is very important to tailor the amount of functionality suit for any specifc application and to ensure that flm reactivity can be maintained for relevant time [57, 63–67]. Plasma deposited polyoxazolines (POx) contain isocyanate-, nitrile groups and intact oxazoline rings. This provides unique opportunities to carry out binding reactions with biomolecules, nanoparticles and various ligands that contain carboxyl groups in their structures [57, 68]. Plasma-assisted processing and catechol chemistry as well as the use of natural antimicrobial agents to produce synthetic antibiotic-free antibacterial surfaces are a particularly hot topic discussed now [69].

*Immobilization of biologically active molecules* (antibiotics and other antimicrobial agents) to polymer surfaces is for a long time studied [4].

Examples of plasma assisted *attachment of antibiotics* are grafting of triclosan and bronopol on oxygen plasma pre-treated polyvinylchloride; grafting of gentamycin to polyvinylidene fuoride after plasma-induced graft polymerization of acrylic acid; etc. [27]. The effcacy of cold plasma for direct deposition of antibiotics is discussed lately as a novel approach for localized delivery and retention of the effect. Ampicillin and gentamicin, deposited onto two types of surfaces: polystyrene micro-titer plates and stainless steel coupons confrmed that the plasma process bonds the antibiotics to the surfaces and ensures localized retention of the antibiotic activity against planktonic and sessile *E. coli* and *P. aeruginosa* [70].

Plasma treated polymer surfaces, enriched with reactive functional groups are utilized in other bioactive molecules immobilization such as peptides, proteins, quaternary ammonium compounds, etc. via corresponding chemical reactions [11, 60]: Plasma pre-treated expanded poly(tetra fuoro ethylene) (PTFE) was peptide immobilised after acrylic acid (AA) grafting and diNH2PEG coupling [71]; RF acrylic acid plasma treated silicone surface was immobilized with avidin protein [72].

Plasma treatments, opening a way to biofunctionalization of chemically inert polymers (such as PDMS is) are of especial interest. Trying to combine some advantages of both: ion-beam and plasma treatment, namely the durability of the modifying effect of the ion-beam with the simplicity of the plasma as compared to ion-beam equipment, we developed a special irradiation technique, plasma based Ar+ beam, to activate the PDMS surface for further hybrid functionalization [18, 73, 74]. Assuming that the existence of an ion-fow in the plasma volume could strength the surface modifying effect including its durability, a parallel plate reactor equipped with a serial capacitance (Fig. 6) was employed to obtain an ion fow in the plasma volume. The vary of the discharge power ensures varied density of the ion fow [10].

**Fig. 6** Parallel plate single-wafer reactor in variant of plasma based Ar+ beam mode of surface treatment

Plasma based Ar+ beam can initiate multistep surface modifcation procedure including antimicrobial agent (peptide/protein) immobilization via fexible spacer. The principle scheme of this experimental approach is presented in Fig. 7.

This multistep procedure opens a new way to obtain four types modifed PDMS surfaces: (1) partially mineralised (moderate hydrophilic, with O-containing groups and free radicals); (2) chemically grafted with AA (moderate hydrophilic, with – COOH functional groups); (3) diNH2PEG-coupled (strong hydrophilic, with –NH2 functional groups; PEG acting as fexible spacer); and (4) biomolecules immobilized (collagen, antimicrobial peptide, or other). The chemical composition, surface topography and roughness as well as the surface hydrophilic/hydrophobic balance, surface free energy, its components and polarity were controlled on every stage of the modifcation procedure by means of XPS, AFM, SEM and equilibrium contact angle measurements.

**Fig. 8** A simplifed sketch of the cascade process that results in a partial mineralization and O-containing functional groups formation on the PDMS surface due to a plasma Ar+ beam treatment

Figure 8 demonstrates the cascade process of partial mineralization and functionalization of cross-linked PDMS. It is evident where and how appear a number of active centres such as free radicals and oxygen-containing gropes that could be utilised in further chemical grafting of desired functionalities.

The partially mineralized surface layer was similar to that obtained after a conventional ion-beam. The PDMS surface hydrophilisation was due to surface polarity increase as a result of polar groups' accumulation, this effect depending on the discharge power. The above presented multi-step procedure has a potential to be used whenever need arises to control chemical activity, hydrophilic/hydrophobic balance and biocontact properties of chemically inert polymers for application as antimicrobial biomaterials for cells culture, processing of biosensors, indwelling medical devices, etc.

In 2019, Tran et al. [75] develop a single step plasma process for covalent binding of antimicrobial peptides on catheters to suppress bacterial adhesion. Plasma immersion ion implantation (PIII) was demonstrated as a single step treatment leading to covalent coupling of antimicrobial peptides to both internal and external surfaces of PVC catheter tubing, reducing 99% of bacterial adhesion. Ye et al. [76] created self-sterilizing surfaces using a single-step solvent less grafting method. A grafting process was conducted by vapor deposition of a crosslinked poly(dimethylaminomethyl styrene-*co*-ethylene glycol diacrylate) (P(DMAMS-*co*-EGDA)) prime layer, followed by *in situ* grafting of poly(dimethylaminomethyl styrene) (PDMAMS) from the reactive sites of the prime layer. This hybrid coating demonstrates more than 99% bacterial killing against both Gram-negative *E. coli* and Gram-positive *B. subtilis* [76]. Surface-grafted polymers, known as polymer brushes, become an important tool for surface modifcation and functionalization.

Wang et al. [77] review the recent progress in the surface-grafting of polymers, including their formation and utilization in functional materials for electronics, medical devices, etc. O2-functional groups, introduced by oxygen-plasma treatment, on plasma polymerized HMDSO surface, are utilized in binding of pharmaceuticals and anti-microbial peptides inhibiting the bioflms accumulation [78, 79]. The plasma treatment effectiveness as a tool to direct PET surface modifcation or to surface functionalization prior to immobilization of chitosan was evaluated using different discharge types: DC-discharge (at the cathode or at the anode) or AC-discharge [80]. The use of cold RF- and atmospheric pressure plasma-assisted polymerization for subsequently immobilization of various biomolecules for biomedical applications is discussed lately [81].

#### **4 Concluding Remarks**

Plasma treatment of polymers under corresponding operation conditions allows deposition of contact killing, releasing (including controlled release) or lowadhesive antimicrobial coatings, as well as polymer surface functionalization and durable immobilization of antimicrobial molecules. Most of the plasma technologies are developed in laboratory conditions and the surface engineered biomaterials are tested *in vitro*.

The use of plasmas facilitates modifcations which are diffcult or unable to achieve by conventional physical or chemical methods, like for example the stable attachment of biologically active molecules onto chemically inert polymer surfaces.

For the step "from laboratory into clinical practice" it is essential to examine the *in vivo* antimicrobial action by using appropriate animal models and human groups.

#### **References**

oxazoline)-based coatings designed for biomedical applications. Sci Rep. 2020;10(1):9478. https://doi.org/10.1038/s41598-020-66423-w.

**Open Access** This chapter is licensed under the terms of the Creative Commons Attribution 4.0 International License (http://creativecommons.org/licenses/by/4.0/), which permits use, sharing, adaptation, distribution and reproduction in any medium or format, as long as you give appropriate credit to the original author(s) and the source, provide a link to the Creative Commons license and indicate if changes were made.

The images or other third party material in this chapter are included in the chapter's Creative Commons license, unless indicated otherwise in a credit line to the material. If material is not included in the chapter's Creative Commons license and your intended use is not permitted by statutory regulation or exceeds the permitted use, you will need to obtain permission directly from the copyright holder.

# **Antimicrobial Biosurfactants Towards the Inhibition of Bioflm Formation**

**Inês Anjos, Ana F. Bettencourt, and Isabel A. C. Ribeiro**

#### **1 Introduction**

Nowadays, infections associated with urinary tract medical devices, have become a common health issue. The fact that their surfaces are prone to microbial colonization and bioflm formation is certainly a problem. As a result, these medical devices usage can be a source of extreme concern, especially for critically ill patients [1].

Urinary tract related infections (UTIs) are among the most frequent HAIs comprising 27% in Europe and 36–40% in the USA [2, 3]. Unfortunately, intensive care units (ICU) also have become a stage of HAIs events where several reported infections among ICU patients can be attributed to catheter-related urinary tract infections (CRUTIs) [4, 5]. UTIs in ICUs have been reported as 1.1 per 1000 patient-days in Europe and most of these are, CRUTIs i.e.97.4% [3]. Nevertheless, deviceassociated UTI are not exclusively catheter-related. Consequently, among patients undergoing ureteral stents, 38% develop UTI while 45–100% of them have bacteriuria [3].

The incidence of UTIs calls for a well understanding of their pathogenesis, alongside with rapid interventions before and after bacterial colonization and bioflm formation to prevent such infections and to diminish its negative impacts.

Bioflm is a mono or multilayer of interconnected microorganisms surrounded with extracellular matrix (ECM) interfacing a liquid medium. Scientists have described bioflm formation as a process of multi-steps including adhesion, aggregation, maturation, and detachment. Drawbacks associated with bioflm development are substantial and compose a challenge for UTIs management and prevention. In addition to their ability to detach as planktonic to colonize on other surfaces inside the body systems, the diversity of bioflm microorganisms also contributes to

I. Anjos · A. F. Bettencourt · I. A. C. Ribeiro (\*)

Research Institute for Medicines (iMed.ULisboa), Faculty of Pharmacy, Universidade de Lisboa, Lisbon, Portugal

e-mail<EMAIL_ADDRESS><EMAIL_ADDRESS><EMAIL_ADDRESS>
https://doi.org/10.1007/978-3-031-04484-7\_23

the failure of antimicrobial treatment [6]. Additional potential protection of bacteria is provided by the bioflm composition which acts as a guard against antibiotics and immunity system.

Due to the resistance of mature bioflms and the risks associated to bioflm manipulation or eradication, prevention of bioflm generation is advantageous. Therefore, the majority of the strategies that have been proposed aim at preventing the early stages of bioflm formation on the catheter surface [7]. A possible strategy may be the use of surface active molecules with antimicrobial activity such as biosurfactants.

Biosurfactants are surface-active molecules that have granted priority in research and industrial studies. They are identifed as amphiphilic biomolecules produced by a wide range of microorganisms as secondary metabolites, owning the ability of surface tension reduction like the industrial surfactants but with the advantage of being eco-friendly molecules that can be produced from renewable resources. Moreover features associated with biosurfactants include low toxicity, biodegradability, cost-effectiveness, and biocompatibility [8]. Due to their unique properties, biosurfactants can be used in several applications regarding pharmaceutical, food, agriculture, petroleum and cosmetic industries [9, 10].

According to their structure, biosurfactants have been classifed into glycolipids (e.g. rhamnolipids and sophorolipids), lipoproteins or lipopeptides (e.g. surfactin), conjugated phospholipids and fatty acids (e.g. polymyxin) and polymeric biosurfactant (e.g. liposan) [8].

Glycolipids are the most used biosurfactants due to their surface-active properties, e.g. dispersion, emulsion, foaming, solubilization, wetting and penetration [8, 11]. In glycolipids monosaccharide residues are linked to an hydrophobic group [12] and this class comprises sophorolipids, rhamnolipids, trehalose lipids, cellobiose lipids and mannosylerythritol lipids, which are the most studied [8, 13].

Besides their surface-active properties, some biosurfactants also detain interesting biological activities and the most studied have been sophorolipids and rhamnolipids.

Among the glycolipid class, sophorolipids have been assigned for antimicrobial properties and are also considered as potential anticancer candidates considering their ability for apoptosis induction among different types of cells such as liver and leukemia cancer cells [14, 15]. Additional sophorolipids revealed sperm immobilization and death through micelles formation combined with anti-HIV via inhibition of virus duplication [16]. Also, rhamnolipids have been pointed to present antifungal, antimicrobial, antiviral and anti-adhesive properties which makes them suitable for a variety of industrial, environmental, agricultural or medical applications [17].

#### **2 Sophorolipids and Rhamnolipids Antimicrobial Activity**

To understand why sophorolipids and rhamnolipids have been proposed for bioflm inhibition or disruption it is important to study their antimicrobial properties.

#### *2.1 Sophorolipids*

Sophorolipids can be produced by several microorganisms with a considerable yield [13] and present the favorable characteristics of being antibacterial, antiviral, antimycoplasma, antifungal and antialgal agents. Furthermore, sophorolipids can occur either in an acidic (non-esterifed) or lactonic (esterifed) form as illustrated in Fig. 1. Usually, lactonic sophorolipids present higher antimicrobial activity while acidic SLs display higher solubility and foaming characteristics [13, 18].

Sophorolipids' antimicrobial activity effect is assigned to their ability to change the hydrophobic properties of bacterial surfaces and to burst the cellular membrane resulting in the release of intracellular content and death. Sophorolipids may act also as antifungal since they are able to inhibit their movement and induce their lysis [14, 15].

An example of sophorolipids' antimicrobial activity can be verifed in Lydon et al. studies [19] when investigating the antimicrobial potential of sophorolipids produced by *Starmerella bombicola*. The acidic sophorolipids proved to have antimicrobial activities against the nosocomial infective agents *Enterococcus faecalis* and *Pseudomonas aeruginosa*, with signifcant reduction in colony forming units (CFU) at concentrations of 5 mg/mL. In addition, *in vivo* experiments using a mouse skin wounding assay revealed that acidic sophorolipids could be used as a component of antimicrobial creams to reduce the risk of wound infection during healing [19]. Moreover, the antimicrobial activity of sophorolipids was also shown by Dangle-Pulate et al. [20]. The biosurfactants obtained from *Candida bombicola*, with glucose and lauryl alcohol media supplementation, were able to prevent bacterial colonization of *Escherichia coli* (30 μg/mL and 2 h) and *P. aeruginosa* (1 μg/ mL and 4 h) as well as *Staphylococcus aureus* (6 μg/mL and 4 h). The suggested mode of action of antibacterial sophorolipids was pointed out as cellular membrane disruption causing the loss of all cytoplasmic components leading to cellular death [20].

**Fig. 1** Illustration of the chemical structure of acidic (**a**) and lactonic (**b**) sophorolipids. R1=R2=H; R1=H and R2=COCH3; R1=COCH3 and R2=H; R1=R2=COCH3

#### *2.2 Rhamnolipids*

In the past three decades, rhamnolipids have also acquired some recognition for presenting some valuable characteristics, such as antifungal, antimicrobial, antiviral and antimycoplasma activity [21, 22]. Besides presenting antimicrobial activity, rhamnolipids also present antiadhesive properties that can be used as an antimicrobial strategy by coating the surface of medical devices and perform changes on surface's hydrophobicity [18].

Rhamnolipids have structures (Fig. 2) and properties similar to that of detergents and have been reported to intercalate into the membrane phospholipid bilayer, facilitating the permeability of the membrane and fow of metabolites [23].

Their antimicrobial mechanism can explain the results of the antimicrobial activity of the rhamnolipid extract, obtained by Ndlovu et al. [23], observing pronounced activity against a broad spectrum of opportunistic and pathogenic microorganisms, including antibiotic resistant *S. aureus* and *E. coli* strains and the pathogenic yeast *Candida albicans*, when using the agar disc susceptibility method [23]. Moreover, the antimicrobial properties of these compounds were also evaluated by Lotfabad et al. [24] who studied rhamnolipids produced by two indigenous *P. aeruginosa* strains. In this study, preliminary disc diffusion assay showed that all examined Gram-positive bacteria (i.e. *S. aureus* ATCC 29213, *Staphylococcus epidermidis* ATCC 12228 and *Bacillus cereus* ATCC 6051) were inhibited by biosurfactants produced by both MR01 and MASH1 strains [24]. Another study, conducted by de Freitas Ferreira et al. [25] investigated the antimicrobial activity of rhamnolipids under different pH values and assessed an antimicrobial activity against the Grampositive pathogens, *Listeria monocytogenes*, *B. cereus* and *S. aureus*. *B. cereus* was the most sensitive bacteria showing a MIC value of 19.5 μg/mL, and a bactericidal activity at 39.1 μg/mL of rhamnolipids [25].

**Fig. 2** Illustration of chemical structure for mono-rhamnolipids (**a**) and di-RLs (**b**). Fatty acid moieties may present a length that usually ranges from 8 to 14 carbon atoms (1 < n1 and n2 < 7)

#### **3 Fighting Bioflm Formation with Biosurfactants**

Biosurfactants are molecules that have amphipathic structures, which allow the interaction with cellular membranes, such as the bilayer membrane [24, 26]. The interaction of the biosurfactants with the cellular membrane causes changes and perturbations that will lead eventually to the disruption of the cellular membrane and ultimately the release of the cytoplasmic content of the cell and consequently important metabolites [24, 27]. These properties can be used to prevent bioflm formation on medical devices surfaces with the intention of reducing their related infections.

Due to presenting antimicrobial activity, sophorolipids have been explored on bioflm inhibition studies (Table 1). These studies have been developed mostly *in vitro* (e.g. using common microtiter plates) by evaluating the ability of sophorolipids' bioflm disruption or bioflm formation inhibition. Examples include the study of Mukherji and Prabhune [28] that observed the ability of sophorolipid mixtures, produced by *C. bombicola* when the media was supplemented with different plant essential oils, to inhibit *Vibrio cholerae* bioflm formation on glass slides [28]. Moreover, Sen et al. [29] investigated the effcacy of a sophorolipid mixture (SL-YS3) produced by *Rhodotorula babjevae* YS3 towards *Trichophyton mentagrophytes* bioflm. In this study a bioflm eradication around 80% was observed on microtiter plate assays when a concentration of 2 mg/mL was used. Moreover, when observing bioflms disruption on pre-sterilized glass coverslips by scanning electron microscopy (SEM) or confocal laser scanning microscopy (CLSM) a considerable reduction was also observed. The therapeutic effcacy of this sophorolipid mixture on experimentally induced dermatophytosis in mice infected with *T. mentagrophytes* was also evaluated. SL-YS3 showed therapeutic effects and also its ability to regulate collagen deposition together with proper matrix and spatial arrangement, thereby contributing to the healing of the infected skin tissue as compared to the untreated control [29].

The use of these natural biosurfactants to prevent bioflm formation on the surface of medical grade silicone, a common material used in catheters and stents fabrication has also been studied in order to evaluate its potential to reduce related infections. This was frst investigated by Pontes et al. [30] who observed that sophorolipids when adsorbed to silicone could reduce *S. aureus* and *E. coli* bioflm formation. Comparing to plain silicone a reduction of 3 log units on *S. aureus* surface colonization was observed when using a solution with a concentration of 1.5 mg/ mL to promote sophorolipids adsorption. Moreover, a 50% decrease on *E. coli* bioflm formation was also observed (Fig. 3, [30]).

More recently, Ceresa et al. [31] studied the effect of acidic congeners, C18 lactonic sophorolipids and mixture of acidic and lactonic sophorolipids on the disruption of *S. aureus* ATCC 6538, *P. aeruginosa* ATCC 10145 and *C. albicans* IHEM 2894 pre-formed bioflms on medical grade silicone. All three tested mixtures (when at a concentration > 0.1% w/v) were able to disrupt bioflms up to 70%, 75%

**Table 1** Examples of research studies presenting the potential application of the biosurfactants (BS) sophorolipids (SLs) and rhamnolipids (RLs) in preventing bioflm formation

**Table 1** (continued)

and 80% regarding *S. aureus*, *P. aeruginosa* and *C. albicans*, respectively. Moreover, acidic sophorolipid (0.8% w/v) pre-coated silicone discs reduced *S. aureus* bioflm by 75% while *C. albicans* reached 68–70% [31].

It is also of great importance to study the development and ways to prevent bioflm formation onto microfuidic systems. Nguyen et al. [32] demonstrated that sophorolipids had a stronger effect than chemical surfactants such as sodium dodecyl sulphate, Tween20 and Tween80 when disrupting established *P. aeruginosa* PAO1 bioflms grown in microfuidic channels. The authors noticed that although presenting antibioflm properties, sophorolipids did not seem to have antibacterial effects on PAO1. When testing these compounds on a mutant strain that overexpresses extracellular polymeric substances they observed that sophorolipids detached and disintegrated bioflms from glass surfaces [32].

Rhamnolipids have been recognized for their antiadhesive and bioflm dispersion effects. It is suggested that the antibioflm activity of rhamnolipids occurs through

**Fig. 3** Anti-sessile activity of silicone specimens adsorbed with sophorolipids towards *S. aureus* (**a**) and *E. coli* (**b**). Sessile cells CFU counts were performed when different concentrations of SLs were tested (0.1–3 mg/mL). (*Reproduced with permission* [30] *Copyright 2016, Elsevier*)

interference with quorum sensing of bioflm cells, which leads to detachment of microorganisms. This interference has been attributed to the inhibition of intracellular lipidic signals, however, it is also reported that rhamnolipids solubilize ECM proteins via micelles formation [8, 33]. Furthermore, other investigations have pointed out that rhamnolipids can manipulate bioflm-associated channels, altering the oxygen and nutrient supply for sessile microorganisms. Moreover, it has been observed that they may enhance interconnections changes as well, so they affect not only quorum sensing of bacteria but also, they do not allow them to develop genetic mutations or resistance [34]. A previous study of Davey et al. [35] mentioned that rhamnolipids can inhibit both intracellular contact and cell–surface contact allowing the detachment and preventing attachment of microorganisms [35]. Other studies pointed out that the mechanism of action of rhamnolipids is through modifcation of bacterial surface components. These studies proved that rhamnolipids could sharply increase the hydrophobicity of the cell surface by removing out the lipopolysaccharides parts from the outer membrane. Also, they showed that a low concentration of rhamnolipids is required to induce the disruption of cells [36, 37]. Similarly, the antifungal activity of rhamnolipids was explained by the disruption of the cytoplasmic membrane [38]. Additional investigations regarding the antibacterial effect of rhamnolipid showed that monorhamnolipids exhibit bacteriostatic effect while dirhamnolipids were able to kill the bacteria and show bactericidal effect [34].

Due to their antimicrobial and antiadhesive properties rhamnolipids have been the target of many studies focused on diminishing bioflm formation in different surfaces such as polystyrene, silicone and medical devices surfaces.

Different studies aiming the investigation of rhamnolipids antibioflm, antiadhesive and bioflm dispersion effects have been performed and some examples can be found in Table 1. Gomes and Nitschke [39] used different concentrations of rhamnolipids from *P. aeruginosa* to evaluate their capacity in reducing the adhesion and bioflm formation on polystyrene surfaces. When using a rhamnolipids solution at 1% an adhesion reduction of 57.8 and 67.8% was observed for *L. monocytogenes* and *S. aureus*, respectively. Moreover, rhamnolipids were also effective in preventing the adhesion of bacterial mixed cultures. When at a concentration of 0.25%, rhamnolipids removed 58.5% of *S. aureus* bioflm, 26.5% of *L. monocytogenes* bioflm, 23.0% of *Salmonella enteritidis* and 24.0% of the mixed culture bioflm [39].

Moreover, Aleksik et al. [40] compared the antibioflm ability against *P. aeruginosa* PAO1 of di-rhamnolipids produced by *Lysinibacillus* sp. BV152.1 with the commercial di-rhamnolipids. The authors observed that di-rhamnolipids produced by *Lysinibacillus* sp. BV152.1 were more effective in reducing bioflm than the commercial ones and an inhibition of 50% was observed with 50 μg/mL and 75 μg/ mL, respectively. The authors also observed that amide derivatization of both dirhamnolipids improved the inhibition of bioflm formation and dispersion, and that the morpholine derivative was the most active causing more than 80% bioflm inhibition at concentrations of 100 μg/mL [40].

Besides their activity towards bacterial bioflms, rhamnolipids, also have revealed activity on *C. albicans* bioflms. Di-rhamnolipids produced by *P. aeruginosa* when at a concentration of 0.16 or 5 mg/mL were able to reduce pre-formed bioflms on polystyrene by 50% and 90%, respectively. In this study the infuence of rhamnolipids in disrupting *C. albicans* bioflms was proven and the authors suggested their exploration as a potential alternative to the available conventional therapies [41].

Comparison of rhamnolipids with other antimicrobial compounds or surfactants has also been performed. For example, Quinn et al. [42] compared rhamnolipids with antibiotics antibioflm activity. The effect of a rhamnolipid mixture, containing mono- and di-rhamnolipids (20 μg/mL) in pre-existing bioflms was observed as a reduction of 88.4 ± 5.8, 74.5 ± 6.6% and 85.6 ± 3.9% against *B. subtilis*, *Micrococcus luteus* and *S. aureus*, respectively. A lower antibioflm effect was observed with the antibiotics ampicillin, chloramphenicol and kanamycin (5 μg/mL) [42].

Moreover, Shen et al. [43] evaluated sodium lauryl sulfate, rhamnolipids, and N-acetylcysteine ability to eradicate mature bioflms and inhibit new bioflm formation of *Helicobacter pylori*, *E. coli*, *P. aeruginosa*, *S. aureus*, and *Streptococcus mutants*. The authors observed that sodium lauryl sulfate and rhamnolipids successfully inhibited the formation of those fve bacterial bioflms in a dose-dependent manner even at concentrations below the minimal inhibitory concentrations. This suggests that their antibioflm activities are unrelated to their antibacterial activities and that had already been observed by Quinn et al. [42] when comparing rhamnolipids antibioflm activity to antibiotics [43].

These results have led to the hypothesis of using these biosurfactants on medical devices to prevent their related infections. Therefore, some papers have also investigated the potential of rhamnolipids inhibition of different strain bioflms on silicone rubber or medical grade silicone. For example, Rodrigues et al. [44] studied the ability of rhamnolipids to interfere in the adhesion of bacteria and yeasts isolated from explanted voice prostheses onto silicone rubber. The authors concluded that the number of cells adhering onto silicone rubber treated with biosurfactant was reduced and that declines of 50% on the number of cells were attained for *S. epidermidis*, *Streptococcus salivarius*, *S. aureus* and *C. tropicalis*. Nevertheless, *C. albicans* and the bacterial strain *Rothia dentocariosa* showed a lower decrease in the number of attached cells after 4 h (20–28%) [44].

Studies on antibioflm activity of rhamnolipids have also been realized on other medical devices such as catheters. Bioflm formation by *S. epidermidis* is a cause of infections related to peritoneal dialysis. Pihl et al. [45] used a peritoneal dialysis catheter fow-cell model in combination with confocal scanning laser microscopy and atomic force microscopy to study bioflm formation by *S. epidermidis* and observed a reduction in the covering of bioflm with exposure to the supernatant from two *P. aeruginosa* strains (i.e. rhamnolipids). The exposure to this supernatant originated a coverage of only 10% in bioflm when compared to untreated samples [45].

Additionally, when adsorbed to silicone elastomeric discs the rhamnolipids 89, produced by *P. aeruginosa* 89, were able to reduce *Staphylococcus* spp. bioflm formation, by 70 and 50% regarding biomass and 72 and 63% regarding cell metabolic activity (at 72 h) for *S. aureus* and *S. epidermidis*, respectively. SEM analysis also corroborated these results making R89 a promising antibioflm coating for silicon catheters [46].

Recently, Bettencourt et al. [47] developed chitosan–rhamnolipid nanoparticles intended to fght *S. aureus* infections. The obtained particles showed an antimicrobial synergic effect between chitosan and rhamnolipids produced by *P. aeruginosa* when testing their antimicrobial activity towards *S. aureus*. Regarding antibioflm activity of the produced particles a reduction of 99% on bioflm formation on medical grade silicone could be observed making these particles an interesting approach to prevent *S. aureus* related infections such as the medical devices-related (Fig. 4) [47].

#### **4 Conclusion**

Infections associated with urinary tract medical devices are a common health concern, in particularly, when associated to bioflm formation on their surfaces.

Among multiple strategies to fght those infections, biosurfactants as glycolipids can be a valuable tool for bioflm inhibition or disruption. In particularly, multiple *in vitro* studies concerning sophorolipids and rhamnolipids confrms the antimicrobial activity of those compounds.

Further, sophorolipids or rhamnolipids potential role to prevent bioflm associated infections, using different surfaces like medical grade silicone as example of

**Fig. 4** Illustration of rhamnolipids–chitosan particles (RLs–CSp) antimicrobial mechanism of action towards *S. aureus*. Particles may deliver rhamnolipids as encapsulated onto RLs–CSp. (**A1**) or/and adsorbed to its surface (**A2**). First, electrostatic attraction of RLs–CSp (negatively charged) to *S. aureus* membranes (positively charged) takes place (**B1**, **B2**). Later, RLs are released from the particles, enter into membranes leading to cell damage and death (**B3**). (*Reproduced with permission* [47] *Copyright 2021, Elsevier*)

common material used in catheters and stents fabrication, shows the capacity of those biosurfactants in reducing the adhesion and bioflm formation.

Finally, new trends in the delivery of these biosurfactants, namely by their inclusion in nanoparticulate systems paves the way for newer clinical applications.

Overall, sophorolipids and rhamnolipids due to their multiple antimicrobial/antiadhesive effects might be an interesting approach to fght urinary tract medical devices associated infections.

**Acknowledgements** The authors thank Fundação para a Ciência e Tecnologia (FCT) for the fnancial support under Project PTDC/BTM-SAL/29335/2017, UIDB/04138/2020 and UIDP/04138/2020 (iMed.ULisboa).

#### **References**

1. Muszanska AK, Nejadnik MR, Chen Y, Van Den Heuvel ER, Busscher HJ, Van Der Mei HC, et al. Bacterial adhesion forces with substratum surfaces and the susceptibility of bioflms to antibiotics. Antimicrob Agents Chemother. 2012;56(9):4961–4.

**Open Access** This chapter is licensed under the terms of the Creative Commons Attribution 4.0 International License (http://creativecommons.org/licenses/by/4.0/), which permits use, sharing, adaptation, distribution and reproduction in any medium or format, as long as you give appropriate credit to the original author(s) and the source, provide a link to the Creative Commons license and indicate if changes were made.

The images or other third party material in this chapter are included in the chapter's Creative Commons license, unless indicated otherwise in a credit line to the material. If material is not included in the chapter's Creative Commons license and your intended use is not permitted by statutory regulation or exceeds the permitted use, you will need to obtain permission directly from the copyright holder.

# **Novel Antimicrobial Strategies to Combat Biomaterial Infections**

**Zoran M. Markovi**ć **and Biljana M. Todorovi**ć **Markovi**ć

#### **1 Introduction**

Bacteria are present in nature everywhere and the combat with them has the major priority especially in various industrial settings (i.e. food industry) or medical devices [1]. It was established earlier that most of bacteria found in nature exist in the form of bioflms (attached to surface of different objects and not as free foating organisms). Therefore, bioflm formation can be defned as a multistage process. It starts with bacteria adhesion to surface and continues with the formation of extracellular matrix. This matrix is composed of one or more polymeric substances (proteins, polysaccharides, humic substances, extracellular DNA) [1, 2]. Bacteria adhesion to surfaces depends on different surface parameters: wettability, roughness, chemistry, and charge of materials as well as of the nature of bacterial surface, environmental factors and the associated fow conditions etc. [3, 4].

There are several possible strategies to reduce or prevent bacterial infections among different populations: patients and medical staff [5]. Traditional hospital sterilization strategies are based on usage of high level disinfectants: hydrogen peroxide, peracetic acid, glutaraldehyde and low level disinfectants: alcohols, hypochlorites, iodine and iodophor. Advanced sterilization technology focuses on chemical-free technology such as UV rays or gas plasma components. However, there are several disadvantages of both chemical and chemical-free approaches. Firstly, they are toxic to some extent so medical personnel and patients have to evacuate the premises. Secondly, the quality of sterilization is proportional to human labor invested by cleaning personnel [6]. One of the alternative strategies independent of human labor, is to produce antibacterial coatings to reduce or eliminate bacteria colonization on surfaces by leaching of biocides, antibacterial surfaces

Z. M. Marković (\*) · B. M. T. Marković

Vinča Institute of Nuclear Sciences—National Institute of the Republic of Serbia, University of Belgrade, Belgrade, Serbia e-mail<EMAIL_ADDRESS>
with deposited metals such as copper, silver or gold, formation of superhydrophobic surfaces and surfaces encapsulated by photoactive nanoparticles [7–12].

The major drawback of biocides and metal deposited surfaces is their leaching from the surface in the environment. In this way those surfaces lose their antibacterial properties after some time. Besides, these surfaces develop bacterial resistance which causes more than 33,000 deaths and costs 1.5 billion euros per year in Europe [13]. The increase of patients infected in hospitals (in the developing countries the infection rate is 75%) was noticed [14]. The cost and cytotoxicity of the agents mentioned above might be a problem as well. As the price of the best antimicrobial additives (silver, titanium, gold, chitosan) is too high, companies are looking for cheaper and safer additives with strong antimicrobial potential. Permanent cytotoxicity of certain antimicrobial agents in concentrations larger than needed for antimicrobial action may cause many problems. A further limited factor of these materials usage is that silver and copper nanoparticles are prone to oxidation. After a certain time they don't show antibacterial effects.

In recent years new types of antibacterial surfaces have been designed by encapsulation of different photoactive nanoparticles in polymer matrices (polyurethane or dimethylsiloxane) [5, 15, 16].

#### **2 State of the Art**

Photodynamic therapy (PDT) is a treatment which includes the usage of light sensitive drugs in the healing of various diseases (for example skin or eye cancers). Antibacterial PDT (APDT) is used to eliminate multidrug-resistance pathogenic bacteria [17]. Based on principles stated above it is possible to design antibacterial surfaces from photoactive nanoparticles (in the form of hybrids or thin flms/coatings) or by encapsulation of photoactive nanoparticles into various polymer matrices. One of the properties of these nanoparticles is their ability to produce reactive oxygen species-ROS (singlet oxygen, superoxide, hydroxyl radicals, hydrogen peroxide) or heat [18, 19]. ROS eradicates multidrug resistant bacteria, quickly disappears and does not represent a danger to the environment. Heat causes denaturation of bacteria but requires additional means for its control.

Photoactive nanoparticles called photosensitizers (PSs) produce ROS by the following mechanism: PSs have been excited to a singlet excited state by ultraviolet or visible light. From this state electrons are moving to a triple state or return to a ground state. Singlet oxygen can be generated if they transfer their electrons or energy to molecular oxygen as shown in Fig. 1. Molecular oxygen causes oxidative damage of bacteria cells. Since molecular oxygen simultaneously attacks several sites in bacteria, the bacteria are unable to mutate and develop resistance [20–23].

Different nanoparticles can be used as PSs: pristine and doped carbon quantum dots (CQDs) and graphene quantum dots (GQDs), chitosan nanodots (ChiNDs), ultra short single wall carbon nanotubes (US SWCNTs), gold nanoparticles (AuNPs)—Fig. 2. It was earlier reported that polymers (polyurethane,

**Fig. 1** Mechanism of singlet oxygen production by PSs

**Fig. 2** Photoactive nanoparticles used potentially in APDT

polydimethylsiloxane) doped with different molecules and nanoparticles [porphyrin, methylene blue (MB), crystal violet (CV)/ZnO, Au–MB, CQDs/Ag] eradicate wide range of bacteria [*Staphylococcus aureus* (*S. aureus*), *Staphylococcus epidermidis* (*S. epidermis*), *Saccharomyces cerevisiae, Escherichia coli* (*E. coli*), *Bacillus subtilis* (*B. subtilis*)] effectively under visible light [12, 24–30].

CQDs and GQDs are zero dimensional carbon nanomaterials with lateral dimension smaller 10 nm. These materials have very interesting properties: high chemical stability, resistance to photo-bleaching, very good solubility in water or organic solvents, high photoluminescence and simple route for high yield synthesis—Fig. 3a, b. Most interesting biomedical property is their ability to generate ROS when

**Fig. 3** (**a**) Top view AFM image of CQDs; (**b**) PL spectra of CQDs; (**c**) top view AFM image of ChiNDs; (**d**) PL spectra of ChiNDs

they are triggered by visible light and lack of cytotoxicity. Their functionalization (by different functional groups) and modifcation (by doping with different heteroatoms for example) contribute to improvement of ROS generation as well as reduction of energy required for triggering of ROS production [31–36]. ChiNDs are novel class of dots with lateral dimension between 20–50 nm, tunable photoluminescence and high chemical stability—Fig. 3c, d. Due to high surface/volume ratio ChiNDs should be more effcient than commercial bulk chitosan in bacteria eradication. There are only few reports on synthesis of ChiNDs by gamma irradiation [37].

AuNPs have been widely studied in biomedicine due to their unique properties and multiple surface functionalities. Spherical AuNPs possess high surface-tovolume ratio, excellent biocompatibility, low toxicity, surface plasmon resonance and ability to quench fuorescence. Hybrids of AuNPs and CQDs produce ROS better than CQD alone [38].

US-SWCNTs are ultrashort 5–10 nm segments of single-walled carbon nanotubes (SWCNTs), with average width of 1 nm and semiconducting nature [39]. They are soluble in polar organic solvents, acids, and water. This high solubility in organic solvents coupled with their short length, should enable these US-SWNTs to be dispersed and incorporated as single tubes into other materials to form

composites. Due to their similarity with CQDs and GQDs, US-SWCNTs should be potent ROS generators triggered by infrared light.

In our earlier investigation we established that pristine and doped CQDs and GQDs can be very toxic against different types of bacteria strains but only under blue light irradiation [34]. By depositing CQDs as very thin flms (only 3 nm) on glass and SiO2 substrates CQDs show good antibacterial activity against *S. aureus* and *E. coli* and moderate antibiofouling effect toward *Bacillus cereus* (*B. cereus*) and *Pseudomonas aeruginosa* (*P. aeruginosa*) under blue light [33]. By encapsulating CQDs in polyurethane and polydimethylsiloxane antibacterial activity of these nanocomposites enhances several orders of magnitude [5, 16]. Different authors reported earlier that CQDs/TiO2, CQDs/Ag or CQDs/ZnO nanostructures as well as CQDs functionalized with (ethylenedioxy)bis(ethylamine)-EDA, N, S doped CQDs and CQDs @hematite composites show good antibacterial potentials against *S. aureus*, *E. coli*, *K. pneumoniae, B. subtilis* [30, 40–44]. CQDs/EDA nanostructures have higher fuorescence quantum yield compared to pristine CQDs and mixed with H2O2 show synergistic effect and thus can inhibit bacteria growth in smaller concentrations of each individual chemical [45].

## **3 Mechanism of Antibacterial Activity CQDs and Their Hybrids**

Antibacterial activity of CQDs and their composites with different materials is based on the production of ROS. Generated singlet oxygen attacks bacterial wall membrane and contributes to lipid peroxidation. The bactericidal effciency of CQDs/polymers depends on the lifetime of generated singlet oxygen [5, 16]. Luminescence method of singlet oxygen production indicates that luminescence of singlet oxygen come from the CQDs located in the interior of polymer matrix. Thus the contribution of the CQDs nearby polymer surface is negligible.

CQDs doping (for example with nitrogen) improves their antibacterial activity by the formation of amide and amino groups. Electrostatic interaction between protonated forms of amines and amides and the lipids of bacterial membrane induces bacterial dead [46].

In the case of CQDs/TiO2 composites TiO2 generates ROS-electrons of TiO2 transfer from valence band to conduction band and thus form holes in the valence band whereas CQDs under visible light emit shorter wavelength and excite TiO2 again [40]. Antibacterial effect of CQDs@hematite is achieved by electron–hole generation on the surface of this nanocomposite. The electrons in the conduction band react with molecular oxygen and thus produce hydroxyl radicals through an oxidative stress [44].

Agents applied in PDT should have low cytotoxicity. In our previous studies we established that CQDs had low dark cytotoxicity [47]. But it was also reported that cancer cells as well as normal cells might be less sensitive to phototoxicity of GQDs than bacteria strain due to different level of isocitrate dehydrogenase in the cells. Singlet oxygen affects the level of isocitrate dehydrogenase and the cells with lower level of isocitrate dehydrogenase are more sensitive to death by singlet oxygen [48].

Apart from ROS generation and surface functionalization of CQDs, surface wettability and roughness affect the bacterial death. But the effect of surface roughness is limited by the shape and size of bacteria. Namely, bacteria adhere to surfaces which features correspond to their own diameters [4].

#### **4 Conclusion**

In this chapter we discussed new light triggered strategies to combat bacterial infections and possible usage of photoactive polymers for these purposes. Photoactive antibacterial polymers are highly promising solution for novel medical devices. To enable their wise usage for the treatment of urinary infections some changes must be made. For example, the effectiveness of photoactive polymers inside human body can be increased by incorporation of micron sized electronic devices (light emitting diode, light detector, pH sensor, radio frequent device) into polymer matrices. The smart medical device should have multifunctional role: the detection of bioflm formation, the eradication of the formed bioflms by APDT and transferring information to medical staff in real time.

**Acknowledgements** Authors thank for support to the Ministry of Education, Science and Technological Development of the Republic of Serbia (451-03-2/2020-14/20-0302002) and for funding by a STSM grant from the COST Action CA16217 "ENIUS" and funded by the COST (European Cooperation in Science and Technology).

#### **References**

**Open Access** This chapter is licensed under the terms of the Creative Commons Attribution 4.0 International License (http://creativecommons.org/licenses/by/4.0/), which permits use, sharing, adaptation, distribution and reproduction in any medium or format, as long as you give appropriate credit to the original author(s) and the source, provide a link to the Creative Commons license and indicate if changes were made.

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# **Light-Activated Polymer Nanocomposites Doped with a New Type of Carbon Quantum Dots for Antibacterial Applications**

**Mária Ková**č**ová, Eva** Š**pitalská, and Zdenko** Š**pitálský**

#### **1 Introduction**

Carbon quantum dots (CQDs) are relatively new carbon allotrope. The frst mention dates from 2004 when Xu et al. described new fuorescent material after electrophoretic purifcation of carbon nanotubes [1]. It triggered an investigation of new CQD research of synthesis, properties CQDs, and applications. As can be seen from Fig. 1, the number of publications about CQDs is increasing exponentially during the last decade with 3419 papers in the last years according to Web of Science. Antibacterial properties of CQDs represent 3.2% of them; however, their citations are exponentially increasing exceeding the number over 2900 in last year. These data suggest a signifcant future of CQD as an antibacterial material.

CQDs are quasispherical carbon particles with a size less than 10 nm with crystalline *sp*2 cores of graphite and quantum effects. A subclass of CQDs are graphene quantum dots (GQDs), and they have a structure of one or several graphene layers with diameter < 10 nm with higher crystallinity than CQDs. In both cases, however, CQDs are functionalized by functional groups on their surface, which can improve the optical properties, solubility, and chemical stability and generally increase the surface variability and complexity of CQDs.

CQDs have many hydrophilic functional groups at the edges or on the basal plane. Specifc hydrophilic functional groups in the CQDs include epoxy and hydroxyl groups. Hydrophilic CQDs are very well soluble in water and other polar solvents due to their chemical composition; therefore, they differ from other

M. Kováčová · Z. Špitálský (\*)

Polymer Institute, Slovak Academy of Sciences, Bratislava, Slovakia e-mail<EMAIL_ADDRESS><EMAIL_ADDRESS>
E. Špitalská

Institute of Virology, Biomedical Research Center of the Slovak Academy of Sciences, Bratislava, Slovakia e-mail<EMAIL_ADDRESS>
**Fig. 1** Number of CQD publications (green bar), the number of antibacterial properties of CQDs publications (blue line) and their number of references (red line) according to Web of Science

carbon-based nanomaterials. Additional advantages of CQDs include their nontoxicity and biocompatibility. The second types of CQDs are hydrophobic CQDs containing carboxyl and carbonyl functional groups. In comparison to hydrophilic CQDs, hydrophobic CQDs are more effective in producing reactive oxygen species (ROS) responsible for the antibacterial activity of CQDs.

CQDs have extensive application usage from sensors through photoelectrochemical water splitting, chemiluminescence, LEDs, photovoltaic solar cells up to photocatalysis and readers can found several reviews about it [2–7]. CQDs also play an important role in medicine. CQDs are used in intracellular ion detection, toxin detection, pathogen, vitamin, enzyme, protein, nucleic acid, and biological pH value determination [8]. QDs also have great utility in bioimaging, biosensing (for example, QD modifcation with metal ions or biomolecules), fuorescence labelling of cellular proteins (biolabelling), genetic technologies, and cell motion tracking [9, 10]. Despite the broad range of biomedical applications, we would like to focus on antibacterial properties of pure CQDs and their polymer composites. The antibacterial effect of CQDs is based on noninvasive photodynamic therapy (PDT). PDT can cause a specifc biological response on the cellular or subcellular level, such as apoptosis, programmed death, or necrosis, a nonprogrammed pathway [11]. During

**Fig. 2** Mechanism of photodynamic therapy, or so-called Jablonski diagram

this process, CQDs absorb light (photons). One electron absorbs this energy and moves into a higher excited single state. This state is in nanoseconds, and it can emit light and lose its energy or dissipate as heat. The excited photosensitizer (PS) in the singlet state may also undergo the process known as an intersystem crossing. The spin of the excited electron inverts to form the relatively long-lived (microseconds) excited triplet-state that has electron spins parallel. The long lifetime of the CQD triplet state is explained by the fact that the loss of energy by emission of light (phosphorescence) is a spin-forbidden process as the CQDs would move directly from a triplet to a singlet-state [12]. In the presence of molecular oxygen, part of the energy can be transformed into an oxygen molecule which changes to ROS. One of the most important types of ROS is singlet oxygen (1 O2). Singlet oxygen and other ROS react with a wide range of biological targets and are known to be involved in cellular signalling and cell damage [13]. Therefore CQDs act as indirect antibacterial materials when after the illumination with light generates singlet oxygen. This whole process is presented in Fig. 2.

#### **2 Antibacterial Effect of CQDs**

#### *2.1 Pure CQDs*

Jhonsi et al. showed that antibacterial and antifungal activities of CQDs on *Escherichia coli* and *Candida albicans* increased linearly with an increase in CQDs' concentration. It reveals that CQDs can effectively inhibit the growth of the bacteria in a concentration-dependent manner. The results revealed that CQDs inhibit the growth of *E. coli* and *C. albicans* more effectively compared to other tested microbes, *Staphylococcus aureus, Klebsiella pneumoniae,* and *Pseudomonas aeruginosa* [14]. Nie et al. demonstrated that CQDs are effective photosensitizers for *in vitro* PDT, and revealed detection limit inactivation (99.9999 + %) of *E. coli* and *S. aureus* upon visible light illumination (*λ* ≥ 420 nm, 65 ± 5 mW/cm2 ; 60 min) [15]. The antibacterial effect of photoexcited GQDs using blue light (465–475 nm) signifcantly affected the viability of *E. coli* or *S. aureus* [16]. Sun et al. designed an antibacterial system combining GQDs with a low level of H2O2, which could inhibit the growth of *E. coli* and *S. aureus* (106 CFU/mL) bacteria and assessed the antibacterial effcacy of patches from GQDs *in vivo* using Kumming mice as a model. Results of that study indicated that GQDs have potential use for wound disinfection [17]. The results of the study of Kuo et al. showed that a nitrogen-doped graphene quantum dot (N-GQD), serving as a photosensitizer, was capable of generating a higher amount of ROS than a nitrogen-free GQD in PDT when photoexcited for only 3 min of 670 nm laser exposure, indicating highly improved antimicrobial effects. The N-GQD (5.1%) effciently exerted an antibacterial effect, resulting in 100% elimination after a 3-min exposure [18]. Sulfur-doped CQDs (S-CQDs) and N-CQDs were evaluated for bactericidal activity against *E. coli* and *B. subtilis* subsp*. subtilis* (6 × 106 cells/mL). Antibacterial activity was slightly higher against *B. subtilis* than against *E. coli* for both S-CQDs and N-CQDs with greater effectiveness of N-CQDs compared to S-CQDs [19].

#### *2.2 Doped CQDs*

Nanorods of CQDs–ZnO had strong antibacterial activity under visible light irradiation, and a concentration of 0.1 mg/L was able to kill more than 96% of bacteria *E. coli* and *S. aureus* [20]. Feng et al. realized bacterial inactivation on a CQD/TiNT flm. Use of CQD/TiNTs led to almost complete inactivation of *S. aureus* and *E. coli* (2 × 107 CFU/mL) within 10 min using 365 nm UV irradiation [21]. Nitrogen and zinc doped CDs displayed good bactericidal activity against *E. coli* (107 CFU/mL) and *S. aureus* (108 CFU/mL) under visible-light radiations [22].