Antibiotic-eluting medical devices for various applications

Journal of Controlled Release 130 (2008) 202–215

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Journal of Controlled Release

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Review

Antibiotic-eluting medical devices for various applications

Meital Zilberman ⁎, Jonathan J. ElsnerDepartment of Biomedical Engineering, Tel-Aviv University, Tel-Aviv 69978, Israel

⁎ Corresponding author. Department of Biomedical EE-mail address: [email protected] (M. Zilberman

0168-3659/$ – see front matter © 2008 Elsevier B.V. Aldoi:10.1016/j.jconrel.2008.05.020

A B S T R A C T
A R T I C L E I N F O
Article history:
Infection is defined as a hom Received 10 March 2008Accepted 16 May 2008Available online 6 August 2008
Keywords:Drug releaseAntibioticsWound dressingsOrthopedic implantsPeriodontal devices

eostatic imbalance between the host tissue and the presence of microorganisms.It is associated with a large variety of wound occurrences ranging from traumatic skin tears and burns tochronic ulcers and complications following surgery and device implantations. If the wound setting managesto overcome the microorganism invasion by a sufficient immune response then the wound should heal. Ifnot, the formation of an infection can seriously limit the wound healing process. Evidence of increasingbacterial resistance is on the rise, and complications associated with infections are therefore expected toincrease. The main goal in treating various types of wound infections is to decrease the bacterial load in thewound to a level that enables wound healing processes to take place. Conventional systemic delivery ofantibiotics entails poor penetration into ischemic and necrotic tissue and can cause systemic toxicity withassociated renal and liver complications, which result in a need for hospitalization for monitoring.Alternative local delivery of antibiotics by either topical administration or by a delivery device may enablethe maintenance of a high local antibiotic concentration for an extended duration of release withoutexceeding systemic toxicity. The present review describes approaches for local prevention of bacterialinfections based on antibiotic-eluting medical devices. These devices include bone cements, fillers andcoatings for orthopedic applications, wound dressings based on synthetic and natural polymers, intravasculardevices, vascular grafts and periodontal devices. Part of the review is dedicated to our novel composite drug-eluting fibers and structured drug-eluting films, which are designed to be used as basic elements of variousdevices. In this review emphasis is placed on processing techniques, microstructure, drug release profiles,biocompatibility and other relevant aspects necessary for advancing the therapeutic field of antibiotic-eluting devices.

© 2008 Elsevier B.V. All rights reserved.

Contents

1. Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2032. Musculoskeletal and orthopedics-related devices . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 203

2.1. Antibiotic-loaded bone cements and fillers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2042.2. Antibiotic-loaded implant coatings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2052.3. Structured antibiotic-loaded bioresorbable films . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 206

3. Wound dressings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2073.1. Wound dressings based on synthetic polymers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2083.2. Wound dressings based on natural polymers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2083.3. Composite fiber structures loaded with antibacterial drugs for wound healing applications . . . . . . . . . . . . . . . . . . . . . . 210

4. Periodontal devices . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2114.1. Non-degradable devices . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2124.2. Degradable devices . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 212

5. Intravascular devices and vascular grafts . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2136. Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 213References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 213

ngineering, Faculty of Engineering, Tel-Aviv University, Tel-Aviv 69978, Israel. Tel.: +972 3 6405842; fax: +972 3 6407939.).

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http://dx.doi.org/10.1016/j.jconrel.2008.05.020

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203M. Zilberman, J.J. Elsner / Journal of Controlled Release 130 (2008) 202–215

1. Introduction

Infection is defined as a homeostatic imbalance between the hosttissue and the presence of microorganisms at a concentration thatexceeds 105 organisms per gram of tissue or the presence of beta-hemolytic streptococci [1,2]. The emergence of infection is associatedwith a large variety of wound occurrences, ranging from traumatic skintears and burns to chronic ulcers and complications following surgeryand device implantations. If the wound setting is able to overcome themicroorganism invasion by a sufficient immune response then thewound should heal via the common four-phased process of coagulation,inflammation, proliferation and remodeling [3]. If not, the formation ofan infection can seriously limit thewound healing process, can interferewithwound closure andmay even lead to bacteremia, sepsis andmulti-system failure. People who suffer from immunosuppressive disordersobviously face a higher risk of infection. However, evidence of increasingbacterial resistance is on the rise, and complications associated withinfections are therefore expected to increase in the general population.Concern among health care practitioners regarding the risk of woundinfection is justifiable not only in terms of increased suffering to thepatient, but also in view of its economic burden to society.

The main goal of treating the various types of wound infectionsshould be to reduce the bacterial load in the wound to a level at whichwoundhealing processes can take place. Conventional systemic deliveryof antibiotics for both prevention (prophylaxis) and curing suffers fromthe drawbacks of systemic toxicity with associated renal and livercomplications, poor penetration into ischemic andnecrotic tissue typicalof post-traumatic and postoperative tissue, and need for hospitalizedmonitoring [4,5]. Alternative local delivery of antibiotics by topicaladministration, or even better by a local delivery device, addresses themajor disadvantages of the systemic approach by maintaining a highlocal antibiotic concentration for an extended duration of releasewithout exceeding systemic toxicity [6–8]. Antibiotics already incorpo-rated in controlled-release devices include vancomycin, tobramycin,cefamandol, cephalothin, carbenicillin, amoxicillin and gentamicin [4,9].

The effectiveness of such devices is strongly dependent on the rateand manner in which the drug is released [10]. These are determinedby the host matrix into which the antibiotic is loaded, the type of drugand its clearance rate. If the drug is released quickly, the entire drugcould be released before the infection is arrested. If release is delayed,infection may set in further, thus making it difficult to manage thewound. The release of antibiotics at levels below the minimuminhibitory concentration (MIC) may evoke bacterial resistance at therelease site and intensify infectious complications [11,12]. Further-more, certain bacterial species are able to attach to implant surfacesand form a protective bio-film layer which is extremely resistant toboth the immune system and antibiotics. These bio-films areconsidered the primary cause of implant-associated infection [13]. Ithas been found that killing bacteria in a bio-film sometimes requiresapproximately 1000 times the antibiotic dose necessary to achieve thesame results in a cell suspension [14].

A 6-h post implantation “decisive period” has been identifiedduring which prevention of bacterial adhesion is critical to the long-term success of an implant. Over this period, an implant is particularlysusceptible to surface colonization. At extended periods, certainspecies of adhered bacteria are capable of forming a bio-film at theimplant–tissue interface [13]. A local antibiotic release profiles shouldexhibit a high initial release rate in order to respond to the elevatedrisk of infection from bacteria introduced during the initial shock,followed by a sustained release at an effective level for inhibiting theoccurrence of latent infection. In the case of orthopedic-relateddevices it is important to combat the bacteria which were introducedduring the implantation and also those introduced systemicallyafterwards. Therefore the sustained release (second phase) isnecessary. In the case of wound dressings, the burn level and rate oftissue regeneration (depends on the patient's age and other

parameters) affect the wound healing process. Therefore, it is hardto describe an “ideal release profile” for the various applications ofantibiotic-eluting devices. Our activity in the field of “antibiotic-eluting medical devices” includes antibiotic-loaded bioresorbablefilms for orthopedic and periodontal applications, and bioresorbablefibers for wound healing applications. These projects are brieflydescribed in the current review, in addition to other research projects.In each project, various release profiles should be selected for in vivoexperiments, in order to choose the most suitable system.

The most extensively studied and earliest commercially availabledevice for controlled release of antibiotics was developed in the 1970'saccording toBuchholz and Engelbrecht's [15] innovative idea of releasingantibiotics from the newly introduced non-biodegradable polymethyl-methacrylate (PMMA) bone cement. This device is still widely acceptedas a means for reducing bone infection. However, it has severaldrawbacks: PMMA enables only a small fraction of the loaded drug todiffuse through the polymer pores [16–19] and may possibly shelterresistant bacteria, thus causing treatment failure.Moreover, PMMA is notbiodegradable, and when clinical failure occurs secondary surgery isnecessary to remove the PMMA before new bone can regenerate.

Various biodegradable devices from both natural and syntheticpolymers have been produced by different processes in recent years,for use as antibiotic carriers. Biodegradable polymers can release largerquantities of antibiotics and their degradation properties can be tailoredfor a specific application that will affect a range of processes such as cellgrowth, tissue regeneration, drug release and host response [20].Synthetic biodegradable polymers that have been reported for variousantibiotic-eluting devices include poly-(lactide-co-glycolide) copolymers[21–24], polycaprolactone [25,26] polyanhydrides [27–30], polyhydroxy-butyrate-co-hydroxyvalerate (PHBV) [31,32] and polyhydroxyalkanoates[33]. Natural polymers such as collagen [34–38] and chitosan [39–43]are attractive, since they exhibit superior biocompatibility and facilitatecell growth. They are also inexpensive and readily available. Table 1presents most of the antibacterial drugs that have been used incontrolled-release systems to date. The molecular weight of the drug, itswater solubility and its solubility in organic solvent, its meltingtemperature and its antibacterial spectrum must be known in orderto design an antibiotic-eluting system. These properties are presentedfor each drug mentioned in Table 1. This review presents the mostextensively studied systems of devices for the release of antibiotics forvarious applications.

2. Musculoskeletal and orthopedics-related devices

Bacterial infection remains a major limitation of the utility ofmedical implants, despite sterilization and aseptic procedures, withreported infection rates in the range of 0.5–5% for total jointarthroplasties [44,45]. Sources for infectious bacteria include theambient atmosphere of the operating room, surgical equipment,clothing worn by medical professionals, resident bacteria on thepatient's skin and bacteria already in the body [46]. In addition tohuman pain and suffering, direct medical costs associated with suchinfections are extremely high and often result in the removal of theorthopedic implants and the need for a follow-up operation. Theworldwide increase in the application of medical technology willeventually lead to greater demand for medical implants and anincrease in implant-associated infections [13].

Device-associated infections are the result of bacterial adhesionand subsequent bio-film formation at the implantation site. Inhibitingbacterial adhesion is often regarded as the most critical step inpreventing implant-associated infection. Competition between theintegration of the material into the surrounding tissue and adhesionof bacteria to the implant surface occurs upon implantation.Implantation will be successful only if tissue integration occurs priorto considerable bacterial adhesion, thus preventing colonization of theimplant [6]. The remarkable resistance of bio-films to conventional

Table 1Antibacterial drugs and their properties [136,137]

Class/drug Molecular weight(g/mol)

Water solubility(mg/ml)

pH induced in thesurrounding

Solubility in organicsolvents

Meltingtemperature (°C)

Antibacterial spectrum

AminoglycosidesAmikacin 585.6 Highly soluble

(185 mg/ml)Base Insoluble 220–230 (decomposes) Broad spectrum, many

Gram‐positive and Gram-negative bacteriaGentamicin 477.6 Highly soluble

(100 mg/ml)DMF, MeOH,EtOH, ether, CHCl3,acetone. Low solubility: DMSO

102–108(Hydrochloride 194-209)

Tobramycin 467.5 Highly soluble(538 mg/ml)

Low: EtOH 168

CefalosporinsCefazolin 454.5 Slightly soluble

(0.487 mg/ml)Weak acid DMF, pyridine, acetone.

Low: EtOH, methanol198–200 (decomposes) Predominantly active

against Gram‐positivebacteria

Cefoperazone 645.7 Slightly soluble(0.286 mg/ml)

MeOH 169–171 Gram‐positive, withincreased activity againstGram‐negative bacteria

GlycopeptidesVancomycin 1449.3 Highly soluble

(N100 mg/ml)Amphoteric DMSO 185–188 Mainly Gram‐positive

bacteria. Mycobacteria

MacrolidesErythromycin 733.9 Slightly soluble

(1.44 mg/ml)Weak base High: MeOH, EtOH,

acetone, ACN, CHCl3,EtOAc, ether, DMF, DMSO.Low: hexane, toluene

191 Gram‐positive andfastidious Gram‐negativebacteria

NitromidazolesMetronidazole 171.2 Soluble (10 mg/ml) Lipophilic, low

ionizationHigh: EtOH. Low: ether,CHCl3

158–160 Most anaerobes

PenicillinsAmpicillin 349.4 Soluble (10.1 mg/ml) Acid High: MeOH, DMF.

EtOH, acetone, DMF.DMSO. Low: CHCl3

199–202 (decomposes) Gram‐positive and someGram‐negative bacteria.Broader spectrum than mostpenicillins.

PolypeptidesColistin(polymyxin E)

1155.4 Highly soluble(564 mg/ml)

Base MeOH, DMF, DMSO,Low: Dioxane

200–220 Mainly Gram‐negativebacteria

QuinolonesCiprofloxacin 331.4 Insoluble

(0.001 mg/ml)Amphoteric High: MeOH, DMF,

DMSO. Low: Dioxane255–257 (decomposes) Broad spectrum. Active

against both Gram‐positiveand Gram‐negative bacteria

Ofloxacin 361.4 Soluble(28.3 mg/ml)

CHCl3. Low: EtOH, MeOH 250–257 (decomposes) More effective againstGram‐negative than Gram-positive bacteria, but activeagainst several importantpathogens in both groups

RifamycinsRifampin/rifampicin

823.0 Slightly soluble(1.4 mg/ml)

Lipophilic, lowionization

DMSO, CHCl3, ethylacetate methanol,THF. Low: acetone

183–188 Gram‐positive andfastidious Gram‐negativebacteria. Mycobacteria

TetracyclinesDoxycycline 444.5 Slightly soluble

(0.63 mg/ml)Amphoteric High: MeOH, Dioxane, DMF 201 Broad spectrum, many

Gram‐positive and Gram-negative bacteria,mycoplasma. Doxycyclineand minocycline are moreactive against S. aureus andvarious streptococci thantetracycline.

Minocycline 457.5 Soluble (52 mg/ml) Low: EtOH N.ATetracycline 444.5 Slightly soluble

(0.23 mg/ml)High: Toluene, ether,EtOAc, acetone. Low:MeOH, EtOH, CHCl3,DMF, Dioxane

165

204 M. Zilberman, J.J. Elsner / Journal of Controlled Release 130 (2008) 202–215

antibiotic therapy has led to extensive research on synthetic surfacesand coatings that resist bacterial colonization.

2.1. Antibiotic-loaded bone cements and fillers

Non-degradable polymethylmethacrylate (PMMA) bone cementsand spacer beads loaded with antibiotics have been employed clinically

in various forms for nearly four decades in joint replacements and inprevention or treatmentof deep bone infections (osteomyelitis) [47–51].Such systems slowly release the soluble drug from the solidified PMMAbone cement surrounding the implant over time. Local release ofantibiotics at concentrations higher than can be achieved by systemictherapy is most desirable in treating osteomyelitis, since poor bonepenetration of many antibiotics is reported [52]. The non-degradable


PMMA matrices are produced by a polymerization reaction between asolid and a liquid component which are mixed together. The formertypically contains PMMA powder, an initiator, the drug and additivesand the latter contains methyl methacrylate monomers and otheradditives. Curing of the cement mixture occurs within minutes, thustrapping the drug within the dense glassy bulk. Incorporation ofantibiotics in this type of system is limited to antibacterial drugs that areable to withstand the heat generated by polymerization. Recordedpolymerization temperatures range between 70 °C and 120 °C [53].Loaded drugs are released through mechanisms of water porepenetration, soluble matrix dissolution and outward diffusion ofsolubilized drug via matrix imperfections (accessible pores and cracks).PMMA typically displays a biphasic release pattern characterized by aninitial burst release followed by a long tail of low, ineffective and largelyincomplete release that continues for days or months. A number ofstudies have revealed that less than10%of the trappeddrug is eventuallyreleased from the cement [16–19], with evidence of sub-therapeuticrelease of gentamicin 25 years after the primary operation [54].

Commercial acrylic antibiotic-impregnated bone cement productshave been sold in Europe for over 20 years. The antibiotics are eitherpremixed by the manufacturer or added by the surgeon in theoperating room. The American Food and Drug Administration (FDA)has recently approved the use of the following low-dose premixedcements: Cobalt™ G-HV (Biomet), Palacos® G (Biomet), DePuy 1(DePuy Orthpedics), Cemex® Genta (Exactech), VersaBond™ AB(Smith and Nephew) which contain gentamicin and Simplex® P(Stryker Orthpedics) which contains tobramycin. Nevertheless, FDAapproval of these low-dose antibiotic-loaded bone cements isrestricted to cases of joint revision following the elimination of anactive infection. These cements are therefore more appropriate as apreventative measure than for the treatment of an establishedinfection which still requires hand-mixing of higher dosages ofvarious antibiotics into the cement by the physician [55]. Surgeonshave been hand-mixing commonly used cements: Palacos® (Smith &Nephew), Simplex® (Howmedica), CMW (DePuy), and Zimmer(Zimmer) with antibiotics such as penicillin, erythromycin, colistin,cephalosporines, gentamicin, polymyxin, vancomycin, and tobramy-cin. This pattern of use has primarily been the result of antibioticselection based on identification of the infecting organism [55].

Díez-Peña et al. [56] reported that hand-mixing additionalantibiotics into the low-dose (2.89% wt. gentamicin) commercialbone cement CMW-1® (DePuy) to values roughly around 20%significantly improves the drug's release mechanism and enablesalmost complete release of the incorporated drug. It is thought thatwhere ‘reservoirs’ of gentamicin exist in close vicinity, water is able tocreate elution paths which enable more efficient release of gentamicinfrom the inner domains. Loading of additional drug may, however,lead to an undesirable weakening of the bone cement. This is a majorcompromise if the cement is used for implant fixation, wheremechanical strength is imperative [57]. A similar weakening effect isobserved following incorporation of various antibiotics into biode-gradable osteo-conductive calcium phosphate bone cements (CPC's)[58–61], due to an interaction between thewater-soluble drug and thesetting reaction of the cement after adsorption of the drug molecules[60,61]. Schnieders et al. [62] have reported that microencapsulationof gentamicin in biodegradable poly-(lactic-co-glycolic) microspheresprior to mixing of the cement can prevent the negative interaction ofantibiotic and cement and may also offer better control over drugrelease. This first example of a drug-eluting composite cement wasfound to be capable of up to 30% drug loading without compromisingmechanical properties and demonstrated both a low burst release andlinear release of gentamicin over a period of 100 days.

Eptacin™, a biodegradable polyanhydride implant in the form oflinked beads containing gentamicin for local delivery of the antibioticto infected bone, presents an alternative to the non-biodegradableacrylic fillers described thus far. The implant's capability of achieving a

high local drug concentration at the implantation site while limitingsystemic exposure to the drug has been shown in a safety studyconducted with patients [30]. The fabrication of this implant requiresthe melting of the polymer at 125 °C in order to produce a polymer-drug mixture. It is therefore limited to thermally stable drugs. Kraskoet al. [63] have recently developed an injectable biodegradablesynthetic polymeric device made of poly(sebacic-co-ricinoleic-ester-anhydride) for treatment of osteomyelitis which overcomes thisdrawback. The pasty hydrophobic copolymer is incorporated with 10–20% gentamicin by mixing the drug powder into the paste at roomtemperature, and gels in situ when exposed to aqueous surroundingsto form a hydrophobic protective environment for the entrapped drug.The polymer degrades mainly from its surface, releasing theentrapped drug. The safety and positive effect of the device wereconfirmed in vivo on established osteomyelitis induced in a rat model.

2.2. Antibiotic-loaded implant coatings

Antibiotic-loaded implant coatings present a straightforwardapproach for the prevention of implant-associated infections. Theycan provide an immediate response to the threat of implantcontamination but do not necessitate use of an additional carrier forthe antibacterial agent other than the orthopedic implant itself. This ismost relevant for ‘cementless’ implantation procedures that havegained popularity due to better early and intermediate-term results inyoung patients compared to cemented prostheses [64].

Unlike “passive” coating techniques that aim to reduce bacterialadhesion by altering the physiochemical properties of the substrate sothat bacteria-substrate interactions are not favorable, “active” coatingsare designed to temporarily release high fluxes of antibacterial agentsimmediately following the implantation [13]. High local doses ofantibiotics against specific pathogens associatedwith implant infectionscan thus be administeredwithout reaching systemic toxicity levels withenhanced efficacy and less probability for bacterial resistance. Recentstudies have also raised the possibility of incorporating growth factors inorder to promote tissue healing responses [65,66].

The utilization of a bioactive ceramic coating containing hydro-xyapatite (HA), calcium phosphate and other osteo-conductivematerials as antibiotic carriers offers the added value of providingthe physiochemical environment and structural scaffold required forbone-implant integration. In vitro release of antibiotics from hydro-xyapetite-coated implants has been reported for chlorhexidine,vancomycin, gentamicin, tobramycin and several other antibiotics[9,67–70] whose antibacterial efficacy was shown in vitro by theformation of inhibition zones in agar plate testing. The conventionalplasma spraying technique for HA-coating is associated with highprocessing temperatures and therefore does not enable the incorpora-tion of antibiotics in the process. Most reported work thereforefocuses on soaking antibiotics onto plasma-sprayed HA. Stigter et al.were the first to report the incorporation of tobramycin into HAcoatings using a “biomimetic” coating technology at a mild tempera-ture (37 °C) [71]. In short, a supersaturated solution of calciumphosphate containing approximately 3% w/w tobramycin was co-precipitated onto titanium alloy plates, forming an approximately40 μm thick carbonated hydroxyapatite layer. In their later work [9] itwas concluded that antibiotics containing carboxylic groups have abetter interaction with calcium, resulting in improved binding andhigher incorporation into the calcium phosphate coating. Alas, thelongest antibacterial effect achieved still does not exceed three days[72]. To date, the only in vivo examination of a hydroxyapetite-coatedimplant in a rabbit infectionmodel supports the concept by showing asignificant decrease in infection rates. However, further substantiationof its biocompatibility and osseo-integration must be carried out [73].

The study of biodegradable polymeric coatings made from poly-lactic acid and its copolymers with glycolic acid is more established.Release profiles last from several hours to 12 days after exposure to an

Fig. 1. In vitro cumulative gentamicin release from polymer/gentamicin films: (a) A-typefilms containing 10% w/w gentamicin, (b) B-type films containing 10% w/w gentamicin,(c) A-type films containing 30% w/w gentamicin, (d) B-type films containing 30% w/wgentamicin: — highMWPLLA; — lowMWPLLA; — highMWPDLGA; — lowMWPDLGA. The experiments were performed in triplicate and the results are presented asmeans±standard deviations.


aqueous environment [4,74–76]. An additional advantage of suchcoatings is the relative ease with which the polymer can be applied toboth alloys and plastics with polished, irregular or porous surfacesusing a simple dip-coating technique [74]. The implant can be dippedseveral times in a solution of polymer and antibiotics in an organicsolvent to achieve a dense or thick polymer coating. The promisingresults displayed in an animal model for this type of coating [75] weretaken a step further and its first use in humans was investigated forinternal fixation of open tibial fractures using gentamicin poly-(DL-lactic)-coated tibial nails (UTN, Synthes, Bochum, Germany) [65,76].Gentamicin was not detected in the serum and no adverse eventswere observed during a one-year follow-up.

Alternative biodegradable coating materials that have been studiedin recent years include natural rosin-based biopolymers and polyhy-droxyalkanoates. Rosin is a natural polymer obtained from pine treeswhich is composed of amixture of diterpene acids known as resin acidsand a smaller amount of other acidic andneutral bodies. It demonstratesexcellent film forming, coating, and microencapsulating properties. Itssuitability has been confirmed by Fulzele et al. [77] who demonstratedpermanent release of ciprofloxacin over a period of 90 days with 90% ofthe encapsulated drug released and good compatibility in vivo.Polyhydroxyalkanoates incorporatedwith Sulperazone® (cefoperazone)and Duocid® (ampicillin) in the form of rods have already shownpromising results in treating implant-related osteomyletis in rabbits[78]. In addition to their biodegradability and biocompatibility, they alsofeature piezoelectricity,which is claimed to induce bone growth in load-bearing areas [78]. Rossi et al. [32] have reported the coating of discs cutfrom a femoral hip implant with polyhydroxyalkanoates loaded withgentamicin. The coating was prepared by pouring dissolved polymerand gentamicin in chloroform ontometal specimens followed by dryingof themixture. The coating exhibited an initial burst release followed bycontinuous in vitro release of gentamicin over a period of 6 weeks withbacterial eradication within 24–48 h, depending on the copolymercomposition [32].

2.3. Structured antibiotic-loaded bioresorbable films

As mentioned above, bacterial adhesion to biomaterials and theability of many microorganisms to form bio-films on foreign bodiesare well-established as major contributors to the pathogenesis ofimplant-associated infections. Major problems in treating osteomye-litis include poor distribution of the antimicrobial agent at the site ofinfection due to limited blood circulation to infected skeletal tissue,and inability to directly address the bio-film pathogen scenario.Controlled antimicrobial release systems inside orthopedic devicesthus represent alternatives to conventional systemic treatments [79].In one of our recent studies [80] we developed and studiedgentamicin-loaded bioresorbable films that can be “bound” toorthopedic implants (by slightly dissolving their surface beforeattaching them to the implant surface) and prevent bacterialinfections by a gentamicin-controlled release phase for at least onemonth. These systems provide desired drug delivery profiles and donot require an additional implant.

Poly(L-lactic acid) (PLLA) and poly(DL-lactic-co-glycolic acid)(PDLGA) films containing gentamicin were prepared by solutionprocessing accompanied by a post-preparation isothermal heattreatment. In the process of film preparation, the solvent evaporationrate determines the kinetics of drug and polymer solidification andthus the drug dispersion/location in the film. The resulting drug-eluting systems are therefore termed “structured films”. In general,two types of polymer/gentamicin film structures were created andstudied for all matrix polymer types:

(a) A polymer film with drug particles located on its surface. Thisstructure, which is derived from a dilute solution, was obtainedusing a slow solvent evaporation rate which enables prior drug

nucleation and growth on the polymer solution surface. Thisskin formation is accompanied by a later polymer coreformation/solidification. This structure was named the “A-type”.

(b) A polymer film with most of the drug particles distributedwithin the bulk. This structure, which is derived from aconcentrated solution, was obtained using a fast solventevaporation rate and resulted from drug nucleation and

Fig. 2. Number of colony forming units (CFU) versus time for microbiological experiment:(a) Pseudomonas aeruginosa, (b) Staphylococcus epidermidis, (c) Staphylococcus aureus. Thereleasingfilms are:■—A-type PLLA film containing 30%w/wgentamicin; —B-type PLLAfilm containing 30% w/w gentamicin; □— B-Type PDLGA film containing 10% w/wgentamicin; — control – A-type PLLA filmwithout gentamicin.


segregation within a dense polymer solution. Solidification ofdrug and polymer occurred concomitantly. This structure wasnamed the “B-type”. Gentamicin is a water-soluble drug whichpractically does not dissolve in chloroform. Some of its particlesthus diffused out towards the surface during solvent evapora-tion. The drug concentration near the surface is thereforeprobably higher than in the center.

The effects of polymer type, initial molecular weight, filmmorphology (drug location/dispersion) and drug loading on thegentamicin release profile were examined. The cumulative gentamicinrelease profiles from films of various polymer/gentamicin systems arepresented in Fig. 1. All experiments were performed in triplicate andresults are presented as means±standard errors. All release profilesexhibited a burst release followed by a relatively slow release phase.Such profiles are desirable for applications such as fracture fixation,where a burst release is needed in order to prevent infection and killthe microorganisms found in the implant area before they settle andcreate a bio-film which antibiotics cannot easily penetrate. A secondphase of slow drug release is necessary to preventmicrobial infectionsat the implant site during healing. The burst effect is obtained due todiffusion of drug molecules located on the surface and in polymerlayers close to the surface, while the continuing release is obtaineddue to diffusion of drug molecules from the bulk and is affected by thehost polymer's degradation rate.

Gentamicin's therapeutic level in serum is 4–8 μg/ml and its toxiclevel is 12 μg/ml [81]. All studied films released gentamicin at levelshigher than the MIC. As expected, lower molecular weight polymersexhibited higher burst effects and higher release rates, due to a higherquantity of hydroxylic and carboxylic edge groups, which make itmore hydrophilic. Furthermore, a lower molecular weight results in alower glass transition temperature, which facilitates faster drugrelease from the polymer.

Processing conditions strongly affect the release profile throughmorphology. Thus, dilute solutions and slow evaporation ratesresulted in A-type films with the drug located on the surface. Thesefilms exhibited a relatively high burst effect followed by a slow releaserate. In contradistinction, concentrated solutions and fast evaporationrates resulted in B-type films, in which most of the drug is located inthe polymeric film and some is located on the surface. These filmsexhibited a relatively low burst effect followed by a lower release rate.Only the PDLGA with relatively low MW films exhibited similarrelease profiles for A and B-type films. This behavior is obtained due tothe polymer's relatively high degradation rate and gentamicin'sextremely hydrophilic nature. We concluded that the gentamicinrelease profiles from the various systems is determined by the hostpolymer, its initial molecular weight and the processing conditions,which affect the drug location/dispersion in the film. Drug loading hasa minor effect on the release profile.

Microbiological evaluation of the effect of gentamicin release onbacterial viability was performed. These experiments were carried outin order to monitor the effectiveness of various concentrations of theantibiotic released from the films in terms of the residual bacteriacompared with the initial bacterial concentration. Bacteria present inPBS only served as the control. In our experiments the bacteria wereadded at the beginning of the films' release, in order to simulatecontamination at the time of implantation. The results are presentedin Fig. 2. No bacteria were left after 1–3 d compared to the controlwhere all bacteria survived even after 7 d in the presence of a veryhigh concentration of the starter (1×108/ml CFU). All films exhibitedmarked gentamicin release, which was responsible for the dramaticdecrease in bacterial survival (103/ml CFU after 1 d). Moreover, thepolymer/gentamicin film preparation did not affect gentamicin'sactivity as an antimicrobial agent.

This study enabled in-depth understanding of gentamicin-loadedstructured films and as a result, the production of systems with the

desired controlled gentamicin release profiles, i.e. with the desiredburst effect and continuing release rate (within the therapeuticwindow) for several weeks. The developed systems can be applied onthe surface of any metallic or polymeric fracture fixation device, andcan therefore make a significant contribution to the field of orthopedicimplants.

3. Wound dressings

The skin is regarded as the largest organ of the body and has manydifferent functions. Wounds with tissue loss include burn wounds,wounds caused as a result of trauma, diabetic ulcers and pressuresores. The regeneration of damaged skin includes complex tissueinteractions between cells, extracellular matrix molecules and solublemediators in a manner that results in skin reconstruction. The moist,warm, and nutritious environment provided by wounds, togetherwith diminished immune functioning secondary to inadequatewoundperfusion, may allow a build-up of physical factors such as devitalized,ischemic, hypoxic, or necrotic tissue and foreign material, all of whichprovide an ideal environment for bacterial growth [82]. In burns,


infection is the major complication after the initial period of shock. Itis currently estimated that about 75% of the mortality following burninjuries is related to infections rather than to osmotic shock andhypovolemia [83].

Various wound dressings aim to restore the milieu required forskin regeneration and to protect the wound from environmentalthreats and penetration of bacteria. Although traditional gauzedressings offer some protection against bacteria, this protection islost when the outer surface of the dressing becomes moistened bywound exudates or external fluids. Furthermore, traditional gauzedressings exhibit low restriction of moisture evaporation which maylead to dehydration of the wound bed. This may lead to adherence ofthe dressing, particularly as wound fluid production diminishes, thuscausing pain and discomfort to the patient during removal.

Most modern dressings are designed according to the well-acceptedbilayer structural concept: an upper dense ‘skin’ layer to prevent bacterialpenetrationanda lower spongy layerdesigned toadsorbwoundexudatesand accommodate newly formed tissue. Unfortunately, dressingmaterialabsorbed with wound discharges provides conditions that are alsofavorable for bacterial growth. This has given rise to a new generation ofwound dressings with improved curative properties that provide anantimicrobial effect by eluting various germicidal compounds.

3.1. Wound dressings based on synthetic polymers

A variety of dressings that contain and release antibiotic agents atthe wound surface have been introduced. These dressings aredesigned to provide controlled release of silver ions through a slowbut sustained release mechanism which helps avoid toxicity yetensures delivery of a therapeutic dose of silver ions to the wound [84].

A wide variety of semi-occlusive dressing formats, such as foams(Contreet® antimicrobial foam, Coloplast), hydrocolloids (Urgotul SSD,Urgo), alginates (Silvercel®, Johnson & Johnson) and hydrofibers®

(Aquacel, ConvaTec) are available. For instance, Acticoat® (Smith andNephew) is a 3-ply gauze dressing made of an absorbent rayonpolyester core, with upper and lower layers of nano-crystalline silver-coated high density polyethylene mesh [85]. It is applied wet and isthen moistened with water several times daily to allow the release ofthe silver ions so as to provide an antimicrobial effect for 3 days.Concerns have been raised by clinicians regarding the safety of thesilver ions included in most of these products. For example, it wasfound that a young personwith 30% mixed depth burns, who receivedone week of local treatment with Acticoat, had shown hepatotoxicityand argyria-like symptoms and the silver levels in his plasma andurine were clearly elevated, as well as the liver enzymes, during thetreatment period. Therefore the authors raised concern aboutpotential silver toxicity and suggested to monitor the silver levels inthe plasma and/or urine during the treatment [86].

In order to address this issue, the silver in Actisorb® (Johnson &Johnson) is impregnated into an activated charcoal cloth, after whichit is encased in a nylon sleeve which does not enable the silver in theproduct to be freely released at the wound surface but neverthelesseradicates bacteria that adsorb onto the activated charcoalcomponent.

Suzuki et al. [87,88] present a new concept for an antibioticdelivery system that releases gentamicin only in the presence ofwounds that are infected by Pseudomonas aeruginosawith a potentialfor an occlusive dressing. Gentamicin is bound to a polyvinyl alcoholderivative (PVA) hydrogel through a specially developed peptidelinker cleavable by a proteinase. This allows gentamicin to be releasedat specific times and locations, namely when and where P. aeruginosainfection occurs. PVA-(linker)-gentamicin demonstrated selectiverelease of gentamicin in P. aeruginosa-infected wound fluid, andcaused a significant reduction in its growth in vitro.

The substantial disadvantage of these temporary dressings is thefact that similarly to textile wound dressings, the necessary change of

dressings may be painful and increases the risk of secondarycontamination. Bioresorbable dressings successfully address thisshortcoming since they can easily be removed from the woundsurface once they have fulfilled their role. Film dressings made oflactide–caprolactone copolymers such as Topkin® (Biomet, Europe)and Oprafol® (Lohmann & Rauscher, Germany) are available forclinical use. Biodegradation of the film occurs via hydrolysis of thecopolymer into lactic acid and 6-hydroxycaproic acid. During thehydrolytic process the pH shifts towards the acid range, with pHvalues as low as 3.6 measured in vitro [89]. Although these twodressings do not contain antibiotic agents, it is claimed that the lowpH values induced by the polymer's degradation help reduce germgrowth [90] and also promote epithelization [91]. Furthermore, locallactate concentrations can stimulate local collagen synthesis [92]. Filmdressings are better suited for small wounds since they lack anabsorbing capacity and are impermeable to water vapors and gases,both of which cause accumulation of wound fluids on larger woundsurfaces.

Katti et al. [93] report the development of a biodegradable poly(lactide-co-glycolide) (PDLGA) nanofiber-based antibiotic deliverysystem. This system can serve as a biodegradable gauze dressingand an alternative to film dressings. This type of dressing is composedof continuous fibers that form a non-woven fiber mesh bymeans of anelectrospinning process. Briefly, the process of electrospinninginvolves use of a polymer solution that is contained in a syringe andheld at the end of the needle by its surface tension. Charge is inducedon the solution by an external electric field to overcome the surfacetension and form a charged jet of solution. As this jet travels throughair, it experiences instabilities and follows a spiral path. Evaporation ofthe solution leaves behind a charged polymer fiber that is collected ona grounded metal screen. Incorporation of antibiotics in the process ofelectrospinning requires solubility of the incorporated drug in thesolvents used in the process (a mixture of dimethylformamide andtetrahydrofuran). Cefazolin was chosen as complying with thisrequirement. The effects of orifice diameter, applied voltage as wellas polymer and drug solution concentrations were investigated and itwas demonstrated that drug-loaded fibers can be electrospun, andalthough they had larger diameters than the unloaded fibers, theywere still in the nanometer range. Antibiotic release from these fibershas not been reported to date.

3.2. Wound dressings based on natural polymers

Only a handful of natural materials have been investigated forwound dressing applications as either main or additional componentsto the dressing structure which are able to impact the local woundenvironment beyond moisture management and to elicit a cellularresponse. Collagen is the main structural protein of the extracellularmatrix (ECM), and was one of the first natural materials to be utilizedfor skin reconstruction and dressing applications. Collagen-basedproducts have been available commercially for over a decade. Theycome in a variety of set-ups ranging from gels, pastes and powders tomore elaborate sheets, sponges, and composite structures. Collagen'slimitations as a wound dressing ingredient are mainly due to its rapidbiodegradation by collagenase and its susceptibility to bacterialinvasion [94–97]. Drug-eluting collagen sponges have been founduseful in both partial-thickness and full-thickness burn wounds.Collatamp® (Innocoll GmbH, Germany), Syntacoll (AG, Switzerland),Sulmycin®-Implant (Schering-Plough, USA) and Septocoll® (BiometMerck, Germany) are several such products which have been found toaccelerate both granulation tissue formation and epithelialization, aswell as to protect the recovering tissue from potential infection or re-infection by eluting gentamicin. In vivo, drug is released by acombination of diffusion and natural enzymatic breakdown of thecollagen matrix [98]. A comprehensive clinical study of gentamicin-collagen sponges demonstrated their ability to induce high local

Fig. 3. SEM fractographs of a gentamicin-loaded composite fiber demonstrating theconcept of core/shell fiber structures.

Fig. 4. SEM fractographs of composite fibers showing the effect of the O:A phase ratio onthe shell microstructure of samples containing 20% w/v polymer and 20% w/wgentamicin: (a) 6:1, (b) 6:1 and 5% w/v albumin, (c) 12:1, (d) 12:1 and 5% w/v albumin.


concentrations of gentamicin (up to 9000 μg/ml) at the wound site forat least 72 h while serum levels remained well below the establishedtoxicity threshold of 10–12 μg/ml [5].

Simple collagen sponge entrapment systems are characterized byhigh drug release upon the wetting of the sponge, typically within 1–2 h of application. Sripriya et al. [38] have suggested improving therelease profile of such systems by using succinylated collagen whichcan create ionic bonds with the cationic antibiotic ciprofloxacin so asto restrain its diffusion. It is claimed that in this way ciprofloxacinrelease corresponds to the nature of the wound in line with theamount of wound exudates absorbed in the sponge. Effective in vitrorelease from their systemwas found to last for 5 days, and was provensuccessful in controlling infection in rats. Other studies have aimed tobetter control drug release or improve wound healing properties bycombining collagen with other synthetic or natural biodegradableelements. Prabu et al. [36] focused on achieving a more sustainedrelease of the antimicrobial agent and describe a dressingmade from amixture of collagen and poly-(caprolactone) loaded with gentamicinand amikacin, whereas Shanmugasundaram et al. [37] chose toimpregnate collagen with alginate microspheres loaded with theantibacterial agent silver sulfadiazine (AgSD).

Other studies which focused on improving wound healingcapabilities tried to incorporate tobramycin, ciprofloxacin [35] andAgSD [34] into collagen–hyaluronan based dressings. The two latterstudies do not show conclusive evidence of improved healingproperties compared to their control. However, hyaluronan (HA), astructure-stabilizing component of the ECM, is thought to play a rolein several aspects of the healing process with hyaluronan-baseddressings, and exhibits promising results in the management ofchronic wounds such as venous leg ulcers [99,100].

A wide range of studies describe the employment of the poly-saccharides chitin and chitosan as structural materials analogous tocollagen for wound dressings. Both materials offer good woundprotection and have also been found to encourage wound healing

Fig. 5. Effect of the emulsion's parameters on the release profile of gentamicin fromcore/shell fiber structures containing 5% w/v albumin: (a) Effect of polymer content onfibers containing 20% w/w gentamicin and O:A phase ratio of 6:1. Polymer contentused: — 17.8% w/v, — 20% w/v, — 26.7% w/v. (b) Effect of gentamicin content onfibers containing 26.7% w/v polymer and O:A phase ratio of 6:1. Gentamicin contents:— 5% w/w, — 20% w/w. (c) Effect of the emulsion's O:A phase ratio on fibers

containing 20% w/v polymer and 20% w/w gentamicin. O:A phase ratio used: — 6:1,▲— 8:1, — 12:1. The experiments were performed in triplicate and the results arepresented as means±standard deviations.

Fig. 6. Gentamicin release profile from fiber samples which were used formicrobiological evaluation: — sample I (26.7% w/v polymer, 20% w/v drug, 6:1O:A phase ratio and 5% w/v albumin), ●— sample II (20% w/v polymer, 20% w/v drug,6:1 O:A), — sample III (26.7% w/v polymer, 5% w/v drug, 6:1 O:A phase ratio and5% w/v albumin). The experiments were performed in triplicate and the results arepresented as means±standard deviations.


without excessive granulation tissue and scar formation [40]. Chitosanhas also been documented as displaying considerable intrinsic anti-bacterial activity against a broad spectrum of bacteria [42]. Severalattempts to improve the dressing's antibacterial capabilities byincorporating various agents such as AgSD [41], chlorhexidine diacetate[43] and minocycline hydrochloride [39] have been reported.

3.3. Composite fiber structures loaded with antibacterial drugs forwound healing applications

Drug-eluting fibers can be used for various biomedical applica-tions. Few controlled-release fiber systems based on polymers have

been investigated to date. The two basic types of drug-loaded fibersthat have been reported are monolithic fibers in which the drug isdissolved or dispersed throughout the polymer fiber, and hollowreservoir fibers in which the drug is added to the internal section ofthe fiber. The advantages of drug-loaded fibers include ease offabrication, high surface area for controlled release and localizeddelivery of bioactive agents to their target. Disadvantages of mono-lithic and reservoir fibers include poor mechanical properties due todrug incorporation and limitations in drug loading. Furthermore,many drugs and all proteins do not tolerate melt processing andorganic solvents.

In one of our recent studies we presented a new concept of core/shell fiber structures which successfully meets these challenges [101].These composite fibers combine a dense polyglyconate core fiber anda drug-loaded porous PDLGA shell structure, i.e. the antibacterial druggentamicin is located in a separate compartment (a “shell”) aroundthe “core” fiber. The shell is prepared using freeze drying of invertedemulsions with mild processing conditions. These unique fibers aredesigned to be used as basic elements of bioresorbable burn and ulcerdressings. Their investigation focused on the effects of the emulsion'scomposition (formulation) on the shell microstructure, on the drugrelease profile from the fibers and on the resulting bacterial inhibition.

The freeze drying technique is unique in being able to preserve theliquid structure in solids. We used this technique in order to producethe shell from inverted emulsions in which the continuous phasecontained polymer dissolved in a solvent, with water and drugdissolved in it as the dispersed phase. A SEM fractograph showing thebulkmorphology of the reference specimen is presented in Fig. 3a. Thequality of the interface between the fiber and the porous coating ishigh (Fig. 3b), i.e. the preliminary surface treatment enabled goodadhesion between the core and the shell. The shell's porous structurecontains round-shaped pores with a diameter of 1.0–11.5 μm and aporosity of 8–68%. The shell's microstructure affects the drug releaseprofile and can also serve as a good measure of the emulsion'sstability. The following emulsion parameters were chosen in order toobtain a stable emulsion that will result in a homogeneous porousshell structure and a feasible release profile of the water-soluble drug:17.8–26.7% w/v polymer content in the organic phase, 5–20% w/wgentamicin (relative to the polymer) and an organic:aqueous (O:A)phase ratio in the 4:1–12:1 range. The release profiles generallyexhibited an initial burst effect accompanied by a decrease in releaserates with time. We found that the O:A phase ratio strongly affects theshell structure (Fig. 4). An increase in O:A phase ratio from 6:1 to 12:1

Fig. 7. Number of colony forming units (CFU) versus time, when an initial bacterialconcentration of 1×107 CFU/ml was used: a— Staphylococcus aureus, b— Staphylococcusepidermidis, c— Pseudomonas aeruginosa. The releasing fibers are: — sample I (26.7% w/vpolymer, 20% w/v drug, 6:1 O:A phase ratio and 5% w/v albumin), ■— sample II (20% w/vpolymer, 20% w/v drug, 6:1 O:A), — sample III (26.7% w/v polymer, 5% w/v drug, 6:1 O:A phase ratio and 5% w/v albumin), — control – reference fiber without gentamicin.


resulted in smaller pore size and thicker polymeric domains betweenpores. The polymer content demonstrated some effect on themicrostructure, whereas the drug content had practically no effect[101].

Albumin was found to be the most effective surfactant forstabilizing the inverted emulsions. As a surfactant, it is located atthe interface between the aqueous phase and the organic phase,reduces the interfacial tension between the two phases and thereforesignificantly decreases the pore size (Fig. 4). It also enables a highencapsulation efficiency and a relatively low burst release (33%)followed by a moderate release profile which enabled release of mostof the loaded gentamicin within two weeks. This behavior probablyresults from albumin's ability to bind the gentamicin through specificinteractions. The ability of albumin to bind drugs is well-known [102].Albumin can interact with acidic or basic drugs via van der Waalsdispersion forces, hydrogen bonds and ionic interactions. Based onthese results, we chose albumin as the preferred surfactant in oursystems, and most of the study focused on samples which werestabilized with albumin. All three formulation parameters had asignificant effect on gentamicin's release profile: an increase in thepolymer content and the O:A phase ratio or a decrease in the drugcontent resulted in a lower burst release and a more moderate releaseprofile (Fig. 5).

We performed microbiological experiments in order to monitorthe effectiveness of various concentrations of the antibiotic releasedfrom the fibers in terms of the residual bacteria compared with theinitial number of bacteria. Bacteria in PBS only served as the control.We chose the following three types of fibers with different releaseprofiles (Fig. 6):

I. Fibers with a shell based on 26.7% w/v polymer, 20% w/v drug,6:1 O:A phase ratio and albumin as surfactant. These fibersdemonstrated a moderate burst release of 32% followed by amoderate release profile.

II. Fibers with a shell based on 20%w/v polymer, 20%w/v drug, 6:1O:A. These fibers demonstrated a high burst release ofapproximately 60%.

III. Fibers with a shell based on 26.7% w/v polymer, 5% w/v drug,6:1 O:A phase ratio and albumin as surfactant. These fibersdemonstrated a low burst release of 13% during the first dayand 60% within three days. After 3 days the release pattern wassimilar to that of sample II.

The bacterial strains used in this study were Staphylococcus aureus,Staphylococcus epidermidis and P. aeruginosa. Their minimal inhibitoryconcentration (MIC) values are 2.5, 5 and 6.3 μg/ml, respectively. Allthree strains were clinically isolated. These strains were chosenbecause they are prevalent in wound infections, especially S. aureusand P. aeruginosa. The third strain, S. epidermidis, usually comprisesthe normal flora of the skin. However, under grave conditions it cancause wound infections. Moreover, these bacteria can produce bio-films, which prevent antibiotics from reaching the target, thereforecausing resistance. The bacteria were added at the beginning of thefibers' release in order to simulate contamination at the time ofimplantation. The results for all three bacterial types when using arelatively high initial bacterial concentration of 1×107 CFU/ml arepresented in Fig. 7. The released gentamicin resulted in a significantdecrease in bacterial viability, and practically no bacteria survivedafter 2 days. The fiber preparation did not affect gentamicin's potencyas an antibacterial agent. Our new fiber structures are thus effectiveagainst the relevant bacterial strains and can be used as basicelements of bioresorbable drug-eluting wound dressings.

4. Periodontal devices

Periodontal disease is a localized inflammatory response due toinfection of a periodontal pocket arising from the accumulation of

subgingival plaque and is one of the world's most prevalent chronicdiseases. It is estimated that 36.8% of American adults have thedisease. Periodontal disease is thus more prevalent than cancer, heartdisease, arthritis, obesity, AIDS and many other diseases [103]. Topicaladministration of antibacterial agents in the form of mouthwashes isineffective in controlling disease progression since only a limitedamount of drug actually accesses the periodontal pocket. Moreover,the drug is constantly flushed due to a very high fluid clearance rate(an estimated 40 replacements of the fluid an hour within a 5 mmpocket) [104].

Local drug delivery to the pocket in the form of subgingivally-placed systems has numerous advantages. Periodontal diseases arelocalized in the immediate environment of the pocket, which is easilyaccessible for the insertion of a delivery device using a syringe ortweezers depending on the physical form of the delivery system. Thecritical period of exposure of the pocket to the antibacterial drug isbetween 7–10 days [105]. Maintenance of a sustained high drugconcentration can be achieved by correct planning, taking intoaccount the high fluid clearance rate. Sustained release devices inthe form of fibers, powders, strips, pastes, gels and ointments have

Fig. 8. The effect of copolymer type on metronidazole release profile from films loadedwith 10 %wt drug. Host polymers: — 50/50 PDLGA, — 75/25 PDLGA, —PDLLA. Theexperiments were performed in triplicate and the results are presented as means±standard deviations.


been reported. Some systems undergo a phase change from a liquid toan in situ forming solid. These systems have the advantage ofsyringeable delivery and an implant with good retention. Intra-pocketdelivery systems can be divided into degradable and non-degradablesystems. Non-biodegradable systems must be removed or dischargedfrom the pocket subsequent to the accomplishment of their drugrelease function.

4.1. Non-degradable devices

Non-degradable cellulose acetate fibers loaded with tetracy-cline were first reported by Goodson et al. in 1983 [106]. Variousstudies on this system showed that it was unable to deliversustainable levels of tetracycline [106,107], chlorhexidine [108] ormetronidazole [109] to become clinically useful, although iteventually led to the development of the commercially availableActisite™ (Alza Corp. Palo Alto CA) delivery system which iscomposed of a monolithic ethylene vinyl acetate fiber loaded with25% tetracycline. The fiber can be placed around the circumferenceof the tooth to the depth of the pocket and folded upon itself tocompletely fill the pocket. The drug concentration in the pocketwas found constant until the removal of the fiber 10 days afterinsertion [104]. This system's main disadvantages include areported 23% risk of extrusion of the fiber from the pocket [110]and need for removal.

Non-degradable film or slab-based devices made from PMMA andethylcellulose have been reported. PMMA slabs have been formed bymixing various antibiotic agents (tetracycline, metronidazole orchlorhexidine) during self-polymerization of PMMA similar to thatdescribed previously for bone cements, only cured as sheets underhigh pressure and cut in the desired shape. In vitro studies showedthat therapeutic levels of all three drugs may be achieved for a periodof two weeks and that these depend on the nature of the drug and itsinitial concentration [111]. Clinical studies have shown variousdegrees of efficacy, although they did not evolve to clinical use[112]. The second system is created by dissolving ethylcellulose anddrug (chlorhexidine [113,114], metronidazole [115] or minocycline[116]) in either ethanol or chloroform followed by solvent evaporationand cutting of the films to shape. The most extensively studiedsystems which contain chlorhexidine have shown promising clinicalresults in the maintenance of periodontal pockets over a two yearperiod [108].

4.2. Degradable devices

Biodegradable systems are usually polymeric or protein innature and undergo natural degradation following exposure togingival fluid components. Various film-based devices have beendescribed. The first degradable systems to be developed werebased on hydoxypropylcellulose loaded with various agents:tetracycline, chlorhexidime and ofloxacin. Similarly to otherdegradable applications described above, fast release of the drugoccurs from the film within 2 h, followed by maintenance oftetracycline within the pocket for 24 h after insertion. Severalmodifications have been made to address the rapid degradationand short duration of drug release. For example, incorporation ofmethacrylic acid copolymer particles into the film has beenreported to prolong the release of ofloxacin in vitro and in vivofor 7 days [117,118]. Polyhydroxybuteric films loaded with 25%tetracycline or metronidazole have been used clinically anddemonstrated an improvement in clinical and microbiologicalparameters, although they suffer from rapid degradation in theirmechanical properties and therefore required several consecutiveplacements every 4 days during the trial [119].

A degradable device based on hydrolyzed gelatin cross-linked byformaldehyde as reported by Steinberg et al. [120] has evolved into the

commercial Perio-chip™ (Perio Products Ltd., Jerusalem Israel). Adifferent commercially available system, Elysol (Dumex, CopenhagenDenmark), is based on a water-free mixture of melted glycerol mono-oleate and metronidazole to which sesame oil was added to improveits flow properties in the syringe. The gel flows deep into theperiodontal pocket and readily adapts to root morphology. When itcomes in contact with water it sets in a liquid crystalline state. Thematrix is degraded as a result of neutrophil and bacterial activitywithin the pocket [121]. Effective doses of metronidazole within thepocket are maintained for 24–36 h. Another antibiotic gel, Atridox™(Block Drug Corporation Inc., Jersey City, NJ USA) has a solutionformulation that is composed of two separate syringes that arecoupled together. One syringe contains 8.5% w/w doxycycline hyclateand the other 37% w/w poly DL-lactide (PLA). They are dissolved in abiocompatible carrier of 63% w/w N-methyl-2-pyrrolidone, whichquickly hardens into a wax-like substance upon contact with thecervicular fluid. The system slowly releases doxycycline into thesurrounding tissue for seven days. This system has been approved bythe FDA.

In one of our recent studies we developed and studiedmetronidazole-loaded 50/50 PDLGA, 75/25 PDLGA and PLLA films.These structured films were prepared using the solution castingtechnique, as described earlier for polymer/gentamicin films(Chapter 2.3). Concentrated solutions and high solvent evaporationrates were used in order to obtain most of the drug within the bulk.These films are designed to be inserted into periodontal pocketsand treat infections during the metronidazole controlled-releasephase, for at least one month. The effects of copolymer compositionand drug content on the release profile, on cell growth and onbacterial inhibition were investigated. The metronidazole releaseprofiles from films containing 10% drug are presented in Fig. 8.Although the 50/50 PDLGA film degrades faster than the 75/25PDLGA and PDLLA films, the rate of drug release from the latter twofilms loaded with 10% metronidazole was faster than from theformer, due to differences in drug location/dispersion within thefilm. The drug crystals appear to be located mainly on the surface ofthe PDLLA and 75/25 PDLGA films, whereas in the 50/50 PDLGAfilms the drug was located in the bulk and also on the surface. Ourresults indicate that the copolymer composition affects the releaseprofile, while the drug content did not show any significant effecton the shape of the release curves. Human gingival cells and ratmesenchymal bone marrow cells have demonstrated normal invitro growth on the drug-eluting films. The released drug alsoexhibited effectiveness against Bacteroides fragilis. Our microbio-logical inhibition kinetics showed that metronidazole cumulativerelease during 3 days succeeded in totally inhibiting bacterialgrowth after two days [122].


5. Intravascular devices and vascular grafts

Infection of intravascular devices for vascular access and vascularprostheses for the replacement or bypass of damaged arteries is a rarebut serious event. The infection of a vascular graft is a rarecomplication, with an estimated incidence of 0.5 to 2.5% of bypassprocedures. However, the mortality and morbidity rates due to thiscomplication are high (25 to 75%) [123], especially when the aorta isinvolved [124]. Once a prosthetic graft is infected, it almost alwaysnecessitates excision and replacement with a new prosthetic bypass.The development of infection-resistant vascular prostheses maytherefore contribute to the prevention and treatment of thiscomplication.

PET (polyethyleneterephtalate, Dacron™) and ePTFE (expandedpolytetrafluorethylene) vascular prostheses soaked in an antibioticsolution produce a wash-out release of antibiotics within minutesafter placement [125,126]. Several approaches have been proposed forextending release over days and weeks. Antibiotics have been‘bonded’ by soaking collagen [127,128], albumin [129], and gelatin[130–132] sealed grafts to produce extended antibacterial activity. Acomprehensive study on the effect of sealant matter and type ofantibiotic used has been reported by Galdbart et al. [129]. Theantibiotic release ratewas found to varywith the type of antibiotic andprotein support. Excess antibiotic unable to bind to the proteinsealants was released immediately after soaking the graft in water,reaching up to 50%. Albumin and gelatin-sealed grafts displayedrelatively longer elution periods, especially for rifampicin, althoughnone of the combinations displayed quantifiable amounts of anti-biotics for periods exceeding 48 h. Succinylation of gelatin-sealedgrafts has been used to improve matrix–drug bonding via ionicreactions between the drug and the matrix [132]. Overall, a prosthesissoaking in antibiotic (passive adsorption) provides immediatepreventive protection of the graft as the drug reservoir is depletedwithin 4–7 days after implantation. Ginalska et al. have recentlyreported an attempt to covalently immobilize gentamicin [133] andamikacin [134] to a gelatin-sealed PET graft via glutaraldehydeactivation. They found that the antibiotic was bound in mixed-typeway via three types of interactions, predominately strong covalentbonds but also weak interactions: physical adsorption and ionicbonds. Only 3 and 15% of the total drug amount was released in vitrowithin 7 days for gentamicin and amikacin, respectively, and theremaining drug was bound to the biomaterial surface at highconcentrations for at least 30 days. During this period the prosthesesexhibited growth inhibition of several bacterial strains at lowinoculum concentrations. They may thus offer better protectionagainst bacterial infection and bio-film formation than previouslydescribed [134]. The mode of action of very firmly bound antibioticsagainst bacteria remains unknown, but it is possible that they alterbacterial adherence to the prosthesis without being released as freemolecules [124].

An alternative to modified gelatin binding is offered by Blanchem-ain et al. [126,135] who demonstrate the feasibility of coatingcyclodextrins (CDs) on vascular Dacron grafts. CDs are truncatedtorus-shaped cyclic oligosaccharides that have a hydrophobic internalcavity and a hydrophilic external wall, and are able to capture variousactive molecules and progressively release them unmodified. Dacronfibers are coated by a polycondensation reaction between CDs andcitric acid as a cross-linking agent at 90 °C to form a polymer networkof cross-linked CDs which physically adhere to the Dacron fibers. An invitro drug release study of coated grafts demonstrated a linear releaseof vancomycin over 50 days [126].

6. Conclusions

This review article presents an extensive overview of publishedstudies on antibiotic-eluting medical devices for various applications.

These include bone cements, fillers and coatings for orthopedicapplications, wound dressings based on synthetic and naturalpolymers, intravascular devices, vascular grafts and periodontaldevices. Basic elements, such as our new composite core/shell fibersand structured films, can be used to build new antibiotic-elutingdevices. The effect of the processing parameters on themicrostructureand on the resulting antibiotic release profiles, mechanical andphysical properties, and bacterial inhibition, must be elucidated inorder to achieve the desired properties. New designs based on novelnano- and micro-structured basic elements may thus advance thetherapeutic field of antibiotic-eluting devices.

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Antibiotic-eluting medical devices for various applications