Bioactive biomaterials for controlling biofilms
8.6 Development of conventional antimicrobial biomaterials
As conventional antimicrobial therapies are significantly limited in the effective management of device- associated infections, actions to deal with the high incidence rate of such infections have focused on developing device- based approaches. Such approaches aim to prevent microbial biofilms developing in the first place, rather than attempting to eradicate an infection once it is already established as a sessile popula- tion. The development of bioactive, anti- infective or antimicrobial devices are the key focus of such a research. Such devices can act to inhibit microbial adherence through modification of the device surface or inciting irreversible tethering of antimicrobials to these surfaces. Additionally they can restrict microbial growth via the elution of an active agent ( Figure 8.3 ).
A range of methods have been used to modify the surfaces of polymers and/or load medical device polymers with antimicrobial agents in an effort to protect indwelling device surfaces from microbial adherence and biofilm formation. These steps aim to create bacteria- inhibitory and bactericidal surfaces. Whilst bacteria- inhibitory sur- faces inhibit both bacterial colonization and proliferation, bactericidal surfaces emit bactericides designed to kill planktonic and early colonizing microorganisms. This prevents contamination by microbes from occurring on the medical device surface and, furthermore, inhibits both bacterial colonization and the resultant biofilm forma- tion. Such materials feature a number of positive characteristics, including relatively low manufacturing costs, long shelf lives, ease of production and processing and the
Figure 8.3 Antibacterial coating of medical devices: (a) impregnation/loading of a device coated with a polymeric layer (e.g. hydrogel) containing biocidal agents; (b) permanent surface modification of a device with conjugates, with either cidal or anti- adherent activity (adapted from Vasilev et al. , 2009).
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ability of a device to function uncompromised even if an active agent is present (Lin et al. , 2001). Immersion, coating, matrix loading and drug polymer conjugates can all be used to load antimicrobial agents into polymeric matrices (Gorman and Jones, 2002).
Although the process of loading antimicrobial agents directly into a polymeric biomaterial matrix by coating or immersion is a relatively straightforward proc- ess, there is a major limitation in the difficulty of optimizing drug release from the medical device’s surface; there is no coordination of this release of antibacterial agents with the presence of infecting organisms. In accordance with Fick’s Law, the release of antimicrobial agents from the drug- loaded polymer matrix usually follows a ‘burst’ release profile; the majority of the drug is released soon after implantation, followed by a subsequent slow discharge of the antimicrobial at potentially sub- inhibitory levels. Although this level may not be sufficient to prevent infection, it may help in the selection of antimicrobial- resistant strains. The ‘burst’ release depletes the polymer reservoir concentration, leaving the device exhausted and sub- sequently susceptible to bacterial or fungal surface colonization at any point after the antimicrobial release ‘burst’. Further long- term studies are also required to address concerns that the use of medical devices coated or loaded with antibiotic agents, as a prophylaxis of device- associated infections, may increase the spread of antimicrobial resistance.
The use of antimicrobial devices in clinical practice has highlighted the impor- tance of developing efficient drug- release mechanisms as, for the majority of clinical scenarios, any uncontrolled rapid release of a drug (whether this release occurs over a few hours or a few days) would not be appropriate for the prevention of device- related infections (Lin et al. , 2001). Furthermore, in order to maintain sufficient con- centrations of antimicrobial agents for prolonged bactericidal or bacteriostatic action, a certain level of drug mass is required, yet the mass which can be incorporated into the devices is often insufficient to maintain such concentrations.
By covalently linking an agent to a monomer before polymerization, a drug- polymer conjugate can be formed, leading to the production of a drug- polymer material that is extremely resilient. The therapeutic potential of drug- polymer con- jugates for urinary catheter use, for example, has been supported by the fact that they have displayed the ability to significantly reduce both bacterial adherence and encrustation in such applications. However, this approach still presents a number of limitations, including increased manufacturing costs and reduced antimicrobial choices, depending on the compatibility of therapeutic agents with the synthetic reac- tion scheme (Gorman and Jones, 2002). In addition, if conditioning film or cellular material is deposited on the device surface, it may facilitate the establishment of a sessile microbial population by masking antimicrobial activity at the surface of the biomaterial.
As a strategy to limit bacterial colonization and subsequent biofilm formation, the use of biomaterials combined with standard antimicrobial agents is both a simple and straightforward approach. Antimicrobial devices have become a common feature of clinical practice, despite significant limitations and conflicting opinions as to their ability to effectively prevent medical device- associated infections in the long- term.
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8.6.1 Development of bioactive and bioresponsive biomaterials for infection and biofilm control
The key challenge associated with the use of antibiotic or microbicide- loaded poly- meric systems is coordination of the release of inhibitory or cidal concentrations of active agent with the onset of infection or colonization. Typically, antimicrobial- impregnated devices (as discussed earlier in this chapter) exhibit a ‘burst release’
profile, with high concentrations of the active agent released on contact with the biological milieu, followed by a slow, steady release of active agent for hours, days or weeks. However, the previously highlighted limitations of this burst release pro- file and the possible resultant biofilm formation must be considered. Alternatives to simple polymer drug loading and passive elution approaches are therefore necessary to coordinate release of active agents at therapeutic/microbicidal or inhibitory con- centrations, with the arrival of contaminating microorganisms at the device surface and subsequent initiation of colonization and biofilm- mediated infection. Polymeric systems capable of responding to external stimuli therefore offer great promise in the development of novel biomaterials for infection control whereby release of antimi- crobial may be accurately controlled in response to infection- relevant stimuli.
Hydrogel materials are a particularly important class of biomaterials amenable to the development of ‘smart’ triggers for bioresponsive drug release. According to Ulijn et al. , (2007), hydrogels provide a suitably hydrated surface for biological interactions; provide inert surfaces capable of resisting deposition of a condition film through non- specific interactions of biological macromolecules at the surface; the potential for incorporation of biological molecules and cleavable linkers; exhibit tun- able mechanical properties; and can be designed to respond to external stimuli, lead- ing to changes in physical properties (such as swelling/collapse or sol- gel transitions).
Hydrogels have also found application as coatings for indwelling medical devices, since they exhibit good drug loading and, importantly, reduce coefficient of friction between device and biological interface thus easing patient discomfort (especially in indwelling urological catheters) and cell adhesion at the biomaterial interface (Beiko et al. , 2004).
The overarching principle of triggered (or ‘smart’) drug delivery from biorespon- sive polymeric biomaterials relies on the ability of the polymeric system to respond to a stimulus, either externally applied or originating within the system itself, and induce either a physical change ( Figure 8.4 ) resulting in modulation of drug release rate, or acts directly on the biomaterial itself to induce cleavage of the active agent from a stimulus- labile drug polymer conjugate from which the free drug is liberated ( Figure 8.5 and see also, Figure 8.6 ). Although stimuli- responsive, triggered drug release systems are typically in developmental stages on the laboratory scale and not yet available for use in patients to control biofilm and infection, these technologies which exploit a number of triggering stimuli (presence of chemical species and metabolites, enzymes, light, temperature, pH, electrical current), offer exceptional promise for the develop- ment of next- generation medical devices and device coatings which can respond to the presence of infecting microorganisms and infection- relevant stimuli and offer precise drug dosing to control biofilm formation and mitigate the consequences of
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exposure of bacteria to sub- optimal concentrations of microbicidal agents potentially leading to emergence of antimicrobial resistance. A number of these triggers are discussed below in relation to antimicrobial drug release from polymeric matrices.
8.6.2 pH-Triggered antimicrobial release
Throughout the body, variation in local pH (primarily in the GI tract) has been use- fully exploited in controlled and sustained- release formulations, which are currently available. However, pH changes associated with infection may also be usefully exploited as a trigger for the coordinated release of antimicrobials from medi- cal device coatings. pH responsive polymers contain functional groups capable of
Figure 8.4 Stimuli- responsive drug release from a polymer undergoing morphological change as a result of applied stimulus (adapted from McCoy et al. , 2010).
Figure 8.5 Stimuli- responsive drug release from a polymer to which drug is conjugated via a stimuli- labile linker (adapted from McCoy et al. , 2010).
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accepting or donating protons (carboxylate, hydroxyl, amine, sulphydryl). Such func- tional groups exist in dynamic equilibrium and dissociate depending on the pH of the immediate milieu and the pKa/pKb of the functional group. In the case of polyacids, such as poly(acrylic acids) or poly(methacrylic acids), dissociation of carboxylic acids (COOH Ͳ COO- + H+) results in repulsion between adjacent COO- groups throughout the matrix, leading to polymer chain repulsion and elongation, ultimately giving rise to matrices of increased pore size, leading to increased solvent efflux and drug release. This pH-dependent swelling/de- swelling behaviour of polyacidic hydrogels was first described by Katchalsky in 1949. Recently, pH-triggered release of antifungal (miconazole) and microbicidal agents (chlorhexidine digluconate) from polymethacrylic acid (PMAA)/diurethane dimethacrylate- based dental biomaterials for the control of Candida -associated denture stomatitis has been described (Cao et al. , 2010). Antimicrobial release was faster at pH5 compared with pH7 and the matrices could be ‘washed out’ on treatment with EDTA and subsequently recharged with the same or a different class of agent, giving rise to rechargeable, infection- responsive antifungal materials for dental applications.
Perhaps the most obvious infection- dependent pH trigger which may be exploited in the body accompanies infection of urinary tract biomaterials by the Gram- negative pathogen, Proteus mirabilis . During infection P. mirabilis secretes a powerful urease which catalyses the hydrolysis of urea to ammonia and carbon dioxide, raising the pH of urine from physiological (or slightly acidic) pH to alkaline pHs of >9 (Morris and Stickler, 1998), thereby initiating crystallization of poorly soluble calcium and magnesium salts and encrustation of the device. This specific event may prove a use- ful stimulus for the infection- responsive release of antimicrobial agents, with at least one major medical device company involved in development of pH sensitive coatings for urinary catheters which shed or release drugs in response to this pH trigger.
8.6.3 Light triggered antimicrobial release
The use of light as an external stimulus for release of antimicrobial agents from biomaterials is currently an area of intense academic and industrial research activity, since light provides a potentially high degree of control for on/off pulsatile delivery strategies. Although much work has concentrated on polymeric systems for medical devices which may be easily activated by light of a given wavelength (endotracheal tubes, urinary catheters), developments in polymeric design and fibre optic technolo- gies is opening the up exciting possibilities for the extended use of light stimuli in a range of infectious complications of biomaterials. The use of photolabile linkers bear- ing pendant antimicrobial groups is one of a number of potential avenues which have been described recently, albeit as an extension of work performed on photocontrolled non-steroid anti-inflammatory drugs (NSAIDs) release from hydrogels based on copolymers of 2-(hydroxy ethyl) methacrylate and methyl methacrylate, cross- linked with ethylene glycol dimethacrylate (McCoy et al. , 2007). The same group have also demonstrated the effectiveness of light triggered photodynamic antimicrobial activity (photodynamic antimicrobial chemotherapy, PACT) against both S. epidermidis and P. mirabilis from anionic hydrogel copolymers which permanently bind a cationic
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porphyrin through electrostatic interactions as a thin surface layer (Parsons et al. , 2009). While these materials are proposed as intraocular lens materials, their potential for wider application in other indwelling polymeric devices is clear.
8.6.4 Enzymatically triggered antimicrobial release
This strategy employs the inherent enzymatic (proteolytic/lipolytic) activity of virulence factors produced by bacteria during adhesion, colonization and biofilm formation to activate the release of an antibiotic agent from the device coating via cleavage of an enzymatically labile pro- drug. Coordination of release of antimicrobi- als during bacterial attachment and formation of the biofilm should prove effective since this approach targets bacterial cells (i) before biofilm develops which reduces the accessibility of the drug to the bacteria, and (ii) before the nutrient depleted state of a biofilm community develops and the pathogens enter a metabolically quiescent state, rendering them resistant to standard antibiotic therapy.
Proteases of the serine-, cysteine-, aspartyl- and metallo- type are widely distributed in many pathogenic bacteria, where they serve critical functions as virulence fac- tors mediating host immune system evasion, colonization and invasion, direct tissue damage and activation/inactivation of host cell responses (inflammation, protease production). In addition to their direct roles as virulence factors, pathogen- derived proteases function in various distinct housekeeping roles such as nutrient acquisition and facilitation of dissemination during colonization, infection and biofilm forma- tion (ZoBell, 1933). Indeed, more subtle roles for pathogen- derived proteases are constantly emerging in the areas of host proteolytic cascade interruption, disruption of normal response cytokine pathways, inactivation of protease activated receptors (PARs) or excision of other cell surface receptors and inactivation of host immu- noglobulin and endogenous protease inhibitors. While a few pathogens are known to secrete an extensive arsenal of protease virulence factors (for example virulent strains of Porphyromonas gingivalis secrete a number of metallo-, aspartyl- and serine- type proteases), the majority of pathogens studied appear to secrete at least one proteolytic enzyme. It has also been demonstrated that genes involved in phenotypic switching (including those encoding protease species) between planktonic and biofilm- forming type S. epidermidis and P. aeruginosa are activated within minutes of attachment to biomaterial surfaces (Davies and Geesey, 1995; Donlan and Costerton, 2002).
Bacterial proteases therefore offer novel and exciting therapeutic opportunities in two main areas (i) development of inhibitors to modulate protease activity in vivo and, in the strategy proposed here, (ii) utilisation of the exquisitely selective cleav- age specificities of soluble pathogen- derived proteases to effect ‘smart’, co- ordinated antimicrobial release from biomaterials, via the cleavage of tethered antimicrobial prodrugs comprising peptidyl recognition motifs for various pathogen- derived pro- teases. In this manner, antimicrobial release from biomaterials can be programmed to coincide with the presence of microorganisms and their secretion of proteolytic enzymes. Proteases are characterized by exquisite selectivity for their peptidyl substrate and, in the case of pathogen- derived proteases, exhibit an inherent lack of regulation by host- derived natural protease inhibitors (Travis and Potempa, 2000), an
Bioactive biomaterials for controlling biofilms 179
excellent example being Staph. aureus V8 protease (GluV8), a serine protease with no known proteinaceous inhibitor (Komiyama et al. , 1996). This renders pathogen- derived proteases robust chemical effectors for the proteolytic activation of antimi- crobial release from biomaterials as medical devices.
In addition, covalent immobilization of antimicrobial agents in this manner affords significant advantages over standard conventional approaches. Peptidyl ligands remain stably attached to the material surface in the absence of activating proteases and therefore circumvent the problem of ‘burst release’ of antimicrobial. Furthermore, residual protease activity persisting after release of antimicrobials and subsequent eradication of viable microorganisms should effect a controlled, prolonged release of active agent from the material ensuring eradication of persisting microorganisms. The general approach is summarized in Figure 8.6 .
A number of studies have validated the use of enzymatic triggers as stimuli for release of antimicrobial agents to control adhesion and biofilm formation at hydrogel surfaces. Tanihara et al. , (1999) described a protease- triggered antibiotic release sys- tem activated by thrombin, a serine protease of the clotting cascade. In their system, an insoluble polymer- drug conjugate was constructed by covalent attachment of gentamicin to a poly(vinyl alcohol) hydrogel via a thrombin- sensitive peptide linker (Gly-(D)-Phe-Pro-Arg-Gly-Phe-Pro-Ala-Gly-Gly). The conjugate selectively released gentamicin when incubated with thrombin- expressing Staphylococcus aureus infected wound exudate, whilst gentamicin release was not detected following incubation
Figure 8.6 Schematic representation of protease- activated prodrug. The prodrug is composed of an N terminal blocked (or polymer conjugated) protease cleavable peptide ligand and a potent therapeutic compound (drug). Upon exposure to the activating protease, the enzyme cleaves the prodrug at the peptide cleavage site, liberating the therapeutic agent from the peptidyl linker and conjugate (adapted from Gilmore 2012).
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with non- infected wound exudate. Recently, Gilmore and co- workers described the development of hydrogel- based medical device coatings bearing PEGylated peptidyl prodrugs of conventional antibiotics which are substrates for the Staph. aureus V8 serine protease. In the presence of Staph. aureus expressing V8 protease, the protease activated antibiotic- peptide conjugate is cleaved, releasing sufficient concentrations of antibiotic to effectively prevent adhesion and biofilm formation (unpublished data), thus validating this approach to controlling bacterial adhesion to material surfaces (Gilmore, 2012). In a similar approach, Komnatnyy and colleagues exploited the production of lipases by P. aeruginosa to trigger the release of ciprofloxacin from poly(ethylene glycol) materials and control biofilm formation on the polymer surface (Komnatnyy et al. , 2013).