Drug-Eluting Medical Implants Meital Zilberman, Amir Kraitzer, Orly Grinberg, and Jonathan J. Elsner Contents 1 Drug-Eluting Vascular Stents . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 300 1.1 Introduction: Restenosis and Drug-Eluting Stents . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 300 1.2 The First Generation of Drug-Eluting Stents (DES-I) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 301 1.3 The Second Generation of Drug-Eluting Stents (DES-II) . . . . . . . . . . . . . . . . . . . . . . . . . . . 308 1.4 Biodegradable Stents . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 311 1.5 Novel Drug-Eluting Highly Porous Stent Coatings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 316 2 Drug-Eluting Wound Dressings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 318 2.1 Introduction: Infection, Wound Dressings and Local Antibiotic Release . . . . . . . . . . . 318 2.2 Wound Dressings Based on Synthetic Polymers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 320 2.3 Wound Dressings Based on Natural Polymers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 321 2.4 Composite Fiber Structures Loaded with Antibacterial Drugs for Wound Healing Applications . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 323 3 Protein-Eluting Scaffolds for Tissue Regeneration . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 330 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 332 Abstract Drug-eluting medical implants are actually active implants that induce healing effects, in addition to their regular task of support. This effect is achieved by controlled release of active pharmaceutical ingredients (API) into the surrounding tissue. In this chapter we focus on three types of drug-eluting devices: drugeluting vascular stents, drug-eluting wound dressings and protein-eluting scaffolds for tissue regeneration, thus describing both internal and external implants. Each of these drug-eluting devices also presents an approach for solving the drug release issue. Most drug-eluting vascular stents are loaded with water-insoluble antiproliferative agents, and their diffusion from the device to the surrounding tissue is relatively slow. In contrast, most drug-eluting wound dressings are loaded with M. Zilberman (*) Dept. of Biomedical Engineering, Faculty of Engineering, Tel-Aviv University, Tel-Aviv 69978, Israel e-mail: meitalz@eng.tau.ac.il M. Schäfer-Korting (ed.), Drug Delivery, Handbook of Experimental Pharmacology 197, DOI 10.1007/978-3-642-00477-3_11, # Springer-Verlag Berlin Heidelberg 2010 299 300 M. Zilberman et al. highly water-soluble antibacterial agents and the issue of fast release must therefore be addressed. Growth factor release from scaffolds for tissue regeneration offers a new approach of incorporating high-molecular-weight bioactive agents which are very sensitive to process conditions and preserve their activity during the preparation stage. The drug-eluting medical implants are described here in terms of matrix formats and polymers, incorporated drugs and their release profiles from the implants, and implant functioning. Basic elements, such as new composite core/ shell fibers and structured films, can be used to build new antibiotic-eluting devices. As presented in this chapter, the effect of the processing parameters on the microstructure and the resulting drug release profiles, mechanical and physical properties, and other relevant properties, must be elucidated in order to achieve the desired properties. Newly developed implants and novel modifications of previously developed approaches have enhanced the tools available for creating clinically important biomedical applications. Keywords Drug eluting stents
Stent coatings
Wound dressings
Composite structures
Scaffolds
Tissue regeneration 1 Drug-Eluting Vascular Stents 1.1 Introduction: Restenosis and Drug-Eluting Stents The re-narrowing, or restenosis, of a treated artery is the result of a complex series of biological events in response to the initial injury to the vessel which was caused by balloon expansion and the presence of a permanent stent implant. Restenosis is mainly characterized by intimal hyperplasia, i.e., an abnormal increase in the vascular smooth muscle cells (SMC) and vessel remodeling (Babapulle and Eisenberg 2002), causing a reduction in the lumen size and consequently restricting blood flow after an intravascular procedure. The process of in-stent restenosis peaks about the third month and reaches a plateau about six months after the procedure (Freed et al. 1996). See Table 1 for definitions and distinct measurements of clinical restenosis. Bare-metal coronary stents improve clinical and angiographic outcomes by reducing both restenosis and repetitive revascularization procedures compared to balloon angioplasty (Fischman et al. 1994; Serruys et al. 1994). However, despite their advantages, high in-stent restenosis (ISR) rates of 22% (Serruys et al. 1994) and 31.6% (Fischman et al. 1994) have been reported, with even higher rates (>50%) in cases of long and complex lesions, small vessels, multiple stenting, and diabetes mellitus (Elezi et al. 1998; Kastrati et al. 1997). The introduction of drug-eluting stents (DES) represents a breakthrough in the treatment of coronary artery disease owing to their ability to reduce the incidence Drug-Eluting Medical Implants 301 Table 1 Clinical restenosis definitions and distinct measurements Measurement Definition Target lesion The rate of reported re-intervention procedures inside the target revascularization (TLR) lesion Target vessel The rate of re-intervention procedures inside any lesion located in revascularization (TVR) the same coronary vessel of treatment Late lumen loss The resulting luminal length reduction during follow-up In-stent restenosis (ISR) Angiographic measurement during follow-up as stenosis in the treated segment >50% of the treated patients In-segment restenosis Angiographic measurement during follow-up as stenosis in the treated segment including the 5 mm segment distal and proximal to the stent edges >50% of the treated patients Major adverse cardiac Complications in cardiac trials such as death, Q-wave and non-Qevents (MACEs) wave infarction, and target lesion/vessel revascularizations Stent thrombosis Basically defined by the presence of angiographic thrombus in a stent during follow-up. However, it has variable definitions, such as probable or definite stent thrombosis. Recently, a set of definitions were developed by an academic research consortium (ARC) which included all unexplained deaths occurring early (<30 days), late (31–360 days), or very late (>360 days) after the procedure of in-stent restenosis to less than 5% (Colombo et al. 2003; Holmes et al. 2004), and to reduce target lesion revascularization (TLR) and major adverse cardiac events (MACE) by 70 and 50%, respectively (Stone et al. 2004), compared to bare-metal stents. Use of DES has therefore increased in the USA from 19.7% following their approval, to 78.2% of percutaneous procedures by the end of 2004 (Rao et al. 2006). DES were used in all lesion subsets, off-label usage, and highrisk patients where the results were less favorable (Beohar et al. 2007; Rao et al. 2006). Longer term studies in broader populations, which indicated the emergence of troubling clinical issues, led to a surge of manuscripts that tempered enthusiasm towards DES (Rao et al. 2008). However, recovery in the use of DES was noted in 2008, together with market approval of a new generation of DES. It should be noted that a new surge of manuscripts supports the safety and efficacy of DES in higher-risk patients (Turco et al. 2007). In the past five years of DES practice, manufacturers have studied their pitfalls and have enhanced their stent technology accordingly. We sought to learn from this rich experience, and distinguish between the first generation of drug-eluting stents (DES-I) and the second generation (DES-II). 1.2 The First Generation of Drug-Eluting Stents (DES-I) The first commercially available DES (in both the USA and Europe) and certainly the most studied are CypherTM (Cordis, J&J, sirolimus-eluting stent) and TaxusTM 302 M. Zilberman et al. (Boston Scientific, paclitaxel-eluting stent). These stents outperformed many DES of their generation that employed various technologies. DES consist of three components that can radically affect their safety and efficacy: the bioactive agent, the stent platform, and the controlled drug release mechanism. Most DES-I used a standard bare-metal stent platform and an immunosuppressant, anticoagulant, antiinflammatory or antiproliferative agent. The drug release platforms were either simply the drug itself, or a ceramic or polymer coating. 1.2.1 Agents The first generation of DES used a variety of agents for the prevention of restenosis, including anticoagulant, immunosuppressant, anti-inflammatory and antiproliferative agents. The effective and safe biological agents were those that prevented smooth muscle cell (SMC) proliferation but preserved vascular endothelial healing. The vascular endothelium is a key participant in the mechanism of post-procedure thrombosis prevention (Finn et al. 2007; Hofma et al. 2006). The active pharmaceutical ingredient (API) should have a wide therapeutic window and should not induce thrombosis or inflammation (Sousa et al. 2003a). Drug uptake into the vessel wall usually occurs by passive diffusion and convection and is facilitated by hydrophobic compounds that establish substantial partitioning and spatial gradients across the tissue (Chorny et al. 2000; Creel et al. 2000). Certain drugs failed to show antirestenotic properties, while others were effective and safe in reducing in-stent restenosis. This may also emphasize the differences in outcomes between animal and human models (Kutryk 2003). Cypher is a sirolimus-eluting stent which will be addressed in the following paragraphs in greater detail. Sirolimus, also known as rapamycin, is a macrolide that easily crosses the cell membrane and binds to an intracellular protein (FKBP12) which activates the protein mTOR (Gershlick 2005). This results in an inhibition of the cell cycle in the transition from G1 to S, thus blocking cell proliferation without inducing cell death (Moreno and Macaya 2005), allowing minimal vascular damage compared to other antirestenosis drugs. Taxus is a paclitaxel-eluting stent which will be addressed in the following paragraphs in greater detail. Paclitaxel is a potent antiproliferative agent, and is known to inhibit mitosis in dividing cells by binding to microtubules (Grube and Bullesfeld 2002), thus interfering with the pathological proliferation of SMC (Heldman et al. 2001). Paclitaxel provided profound inhibition of neointimal thickening depending on delivery duration and drug dosage under clinical investigation (Grube and Bullesfeld 2002). Studies (Grube and Bullesfeld 2002; Sousa et al. 2003b) indicate the need for a controlled drug release of paclitaxel due to the narrow therapeutic window and the high hydrophobicity of this compound. High paclitaxel dosages may lead to an inflammatory vessel response, medial thinning, and thrombosis, due to delayed re-endothelialization (Grube and Bullesfeld Drug-Eluting Medical Implants 303 2002). Dose control, delivery profile, and tissue pharmacokinetics are therefore essential. Several other APIs selected for localized release presented less favorable outcomes. Antinomycin D is an antineoplastic drug that prevents cell division. Multilink tetra-D polymer/antinomycin D-coated stents (Guidant) presented critical clinical safety performance which led to suspension of the use of this drug (Woods and Marks 2004). Batimastat is a matrix metalloproteinase inhibitor that acts to prevent cell migration and proliferation. Batimastat incorporated in phosphorylcholine-coated DES (Biocompatibles) was studied in the BRILLIANT trial. Early MACE rates led to suspension of the development of this stent (Salam et al. 2006). Anti-inflammatory and anticoagulant agents presented less favorable outcomes than antiproliferative drugs. Dexamethasone is a glucocorticoid that modifies protein synthesis, thereby inhibiting inflammatory responses (Liu et al. 2003) and has a rather low effect on endothelial and SMC proliferation (Daemen and Serruys 2007). DexametTM (Abbott Vascular) is a phosphorylcholine-eluting BiodivYsio (Biocompatibles) coronary stent that releases dexamethasone and was found safe and effective in the STRIDE trial (Liu et al. 2003); however, DexametTM was not found to have an antirestenosis effect (Ribichini et al. 2007). 1.2.2 Drug Release Mechanisms Vascular injury induces cellular and sub-cellular mediators of restenosis that can be found in the arterial wall within hours, and which may persist for days to weeks (Kraitzer et al. 2007). Therefore, a sufficient amount of API should be released with appropriate kinetics that are maintained for several weeks after the procedure in order to eventually eliminate in-stent restenosis while maintaining a confluent endothelial coverage that will suppress thrombosis. In general, DES-I offered three drug-release mechanisms: metal stent with API bound to the metal, porous metal, and metal-coated with durable polymers. QuaDDS-QP2 (Quanam) was the first antiproliferative polymer-coated DES that was developed in order to achieve controlled drug release. The stent used an existing platform, a Quest stent covered with thin rigid sleeves equally placed along the stent. The polymer sleeves were loaded with 7-hexanoyltaxol (QP2), a paclitaxel derivative microtubule inhibitor (Woods and Marks 2004). The SCORE clinical trial presented superiority over bare-metal stents in reduced in-stent restenosis. However, marked safety issues led to termination of both the trial and the stent program. The high surface area of the sleeves and the long duration of the drug elution were assessed as the causes of these outcomes (Grube and Bullesfeld 2002). We will return to DES coated with polymers in the following paragraphs. Paclitaxel-eluting non-polymer-coated stents were developed with the intention of avoiding the safety concerns related to polymeric coating as inferred from trials such as the SCORE trial. Dip-coated paclitaxel DES such as Supra G (Cook) and 304 M. Zilberman et al. V-Flex Plus (Cook) were found to be superior to bare-metal stents in the ASPECT (Grube and Bullesfeld 2002) and ELUTES (Gershlick et al. 2004) trials, respectively. ACHIEVE (Guidant) demonstrated only minor improvement over baremetal stents in the DELIVER trial (Lansky et al. 2004). A direct comparison of polymer and non-polymer-based paclitaxel-eluting stents demonstrated that the polymer-based paclitaxel-eluting stents resulted in superior outcomes compared to the non-polymer based paclitaxel-eluting stent (Iofina et al. 2006). Indirect comparison indicated higher late loss rates of non-polymer-coated stents compared to polymer-coated paclitaxel stents (Daemen and Serruys 2007; Iofina et al. 2006). Dip-coated stents are characterized by fast washout during implantation and limited control over the drug release profile (Halkin and Stone 2004; Iofina et al. 2006). This release mechanism might be suitable for drugs with a wider therapeutic index than paclitaxel. The importance of controlled release of paclitaxel was emphasized in a clinical trial where elution <30 days was associated with higher in-stent restenosis and MACE (Serruys et al. 2005). Durable polymer coatings were developed and applied on the Taxus and the Cypher DES in an attempt to achieve better control while minimizing stent/tissue interactions. Cypher was the first approved DES, receiving the CE Mark in 2002 and FDA approval in 2003 following extensive successful clinical trials (Woods and Marks 2004). Cypher is a tubular stainless steel Bx Velocity stent coated with a 5-mm-thick layer of non-erodible polymer (50:50 mixture of polyethylene-vinylacetate and poly-n-butyl methacrylate) containing 70–300 mg sirolimus and an additional topcoat cover over the drug polymer matrix to serve as a diffusion barrier (Duda et al. 2003; Venkatraman and Boey 2007). The API is released slowly over 4–6 weeks, where approximately 80% is released after four weeks, and 100% after 6 weeks (Woods and Marks 2004). Cypher drastically reduced the in-stent restenosis rate compared to bare-metal stents as confirmed in: RAVEL (Morice et al. 2002), SIRIUS (Moses et al. 2003), E-SIRIUS (Schofer et al. 2003), SES-SMART (Ardissino et al. 2004) in high-risk populations, and C-SIRIUS (Schampaert et al. 2004) studies (Fig. 1). TaxusTM was the second DES to be commercially approved following extensive successful clinical trials (Heldman et al. 2001). In its first version, Taxus was based on a 316L stainless steel NIRx stent coated with a poly(lactide-co-e-carpolactone) copolymer loaded with paclitaxel (Tanabe et al. 2003). The release profile was characterized by an initial burst release over the first 48 h followed by very slow release with ~2 mg being released within 15 days (Venkatraman and Boey 2007), and 92.5% of the drug remain in the matrix for a long period (Serruys et al. 2005). The NIRx platform was replaced with ExpressTM tandem architecture between the Taxus III and IV trials. This platform combines scaffolding with flexibility and deliverability which improve drug release homogeneity, especially in tortuous vessels as opposed to the closed cell structure (Halkin and Stone 2004). Taxus stents have demonstrated a reduced rate of restenosis and revascularization events in Taxus I (Halkin and Stone 2004), II (Colombo et al. 2003), IV (Halkin and Stone 2004), VI (Dawkins et al. 2005) (Fig. 2) and were not associated with increased risk of stent thrombosis, at least when dual antiplatelet treatment with aspirin and Drug-Eluting Medical Implants 305 Fig. 1 Selected Cypher pivotal studies: (a) angiographic in-segment restenosis, (b) target lesion revascularization (BMS, bare metal stent; SES, sirolimus eluting stent) clopidogrel was administered (Halkin and Stone 2004; Moreno and Macaya 2005; Venkatraman and Boey 2007). 1.2.3 Real World The clinical superiority of Cypher and Taxus over bare-metal stents has continued after four (Aoki et al. 2005a) and five (Corporation 2007; Morice et al. 2007) years of the pivotal randomized trials. However, continued neointimal formation over time is a phenomenon observed with DES, as opposed to bare-metal stents in which neointimal formation peaks at about six months and then regresses. Nonetheless, as more and more data emerge for complex lesions, longer-term follow-ups and a broader real world range of patients (Rogers 2004), DES outcomes appear to be less 306 M. Zilberman et al. Fig. 2 Selected Taxus pivotal studies: (a) angiographic in-stent restenosis (ISR), (b) target lesion revascularization (BMS, bare metal stent; PES, paclitaxel eluting stent) favorable (Venkatraman and Boey 2007). Between 50% (Beohar et al. 2007) and 75% (Planer et al. 2008) of all DES use occurs in off-label or untested settings (Beohar et al. 2007). Patients who receive DES for off-label and untested indications tend to present more severe clinical profiles (Beohar et al. 2007). A 3-year follow-up conducted by the Israeli arm of the e-Cypher registry showed that 87% of stent thromboses and 85% of all MACEs occurred in off-label indications (Planer et al. 2008). Figure 3 presents the increased rates of TLR, in-stent restenosis of such “real world” registries: world e-Cypher (Urban et al. 2006), Israeli arm e-Cypher (Planer et al. 2008), RESEARCH (Daemen et al. 2006), REALITY (Morice et al. 2006), REWARDS (Roy et al. 2008). No major differences were found in the clinical trials or angiographic outcomes that presented equivalent safety and efficacy, when comparing Taxus and Cypher in the real world (Morice et al. 2006; Roy et al. 2008). Drug-Eluting Medical Implants 307 Fig. 3 Real world Cypher and Taxus studies, target lesion revascularization results (BMS, bare metal stent; PES, paclitaxel eluting stent; SES, sirolimus eluting stent) Cypher and Taxus were associated with an increased rate of late stent thrombosis (LST) (Lagerqvist et al. 2007) and hypersensitivity reactions on a smaller scale (Virmani et al. 2004). DES implantation may increase death and myocardial rates at 6–18 months compared with bare-metal stents, particularly after discontinuation of anti-platelet therapy (Joner et al. 2006). Stent thrombosis is a low frequency event, with serious life-threatening consequences. A pooled 4-year data analysis indicates that out of patients who had definite or probable stent thrombosis 30.9%s died and 83.8% underwent myocardial infraction (Mauri et al. 2007). The difficulty of uniquely defining thrombosis rates has posed an analysis bias. A set of definitions was recently developed by an academic research consortium (ARC) to serve as standard criteria for stent thrombosis in cases which were not taken into account in previous definitions (Planer et al. 2008). The incidence of definite or probable stent thrombosis according to any ARC thrombosis criteria was found to be three times higher than the incidence obtained by the previous definitions in a 4-year pooled data from RAVEL, SIRIUS, C, E-SIRIUS, Taxus I, II, IV, V (Mauri et al. 2007). Late stent thrombosis is mostly predicted by partial, heterogenic endothelial coverage characterized by persistent fibrin deposition and delayed re-endothelialization (Finn et al. 2007; Joner et al. 2006). The mechanisms by which DES induce nonuniform incomplete healing are not fully understood; however, the main suspects are lesion type, the antiproliferative drug itself, and its dose and distribution (Finn et al. 2007). Hofma et al. (2006) found that the sirolimus-eluting stent had an adverse effect on local endothelium-dependent vasomotor responses at six months. Late stent thrombosis may also result as a consequence of hypersensitive reactions caused by the durable polymer coating (Virmani et al. 2004). A summary of the cumulative experience with DES-I indicates a complex interplay between drug selection and drug-release mechanism which determines 308 M. Zilberman et al. the safety and optimizes the local therapeutic benefit (Duda et al. 2003; Sousa et al. 2003a). The DES era also poses new challenges as a result of the growing complexity of lesions that are treated percutaneously, which require better deliverability and higher flexibility (Daemen and Serruys 2007). DES-I offered limited control over the drug release period, drug content, and homogenous release in cases of complex anatomical circumstances (Rogers 2004; Wang et al. 2006). 1.3 The Second Generation of Drug-Eluting Stents (DES-II) DES-II were designed in light of safety and efficacy issues of DES-I, offering an enhanced platform, release matrix, and more targeted antiproliferative agents. The enhanced platform aims to increase clinical outcomes in high-risk populations and challenging anatomies. The offered release matrices improve the polymer/tissue reaction as well as drug release kinetics. Cypher and Taxus have set a standard for the development of DES-II and their clinical outcomes are frequently used for comparison. 1.3.1 Platforms DES-II platforms are designed for better deliverability, higher flexibility (Daemen and Serruys 2007), as well as drug release homogeneity and a low strut profile. Taxus Liberté (Turco et al. 2007), and Cypher SelectTM (Gao et al. 2008), the advanced versions of Taxus and Cypher, were designed for this purpose. Their platforms consist of a closed-cell, continuous structure, in order to provide both uniform scaffolding and homogenous drug distribution within the carrier. The safety and efficacy of Cypher Select and Taxus Liberté were found comparable to their predecessor in the CCSR (Gao et al. 2008) and ATLAS (Turco et al. 2007) studies, respectively. Most DES-II platforms hold thin flexible struts in order to reduce stent/tissue interactions and allow endothelial cells to bridge across the struts (Lewis et al. 2002). DES-II shifted from the 316L stainless steel to a chromium stent platform, since cobalt alloy is 45% stronger than stainless steel, thus allowing thinner struts while maintaining radial strength (Popma 2007). 1.3.2 New Agents Six Limus family agnets: sirolimus, everolimus, biolimus (A9), zotarolimus (ABT578), tacrolimus, and pimecrolimus are used extensively in DES-II. Sirolimus, everolimus, biolimus, and zotarolimus bind to the intracellular binding protein FKBP12, which subsequently binds to the mammalian target of rapamycin (mTOR) and blocks the cell cycle in the G1 to S phase (Daemen and Serruys 2007). Tacrolimus and pimecrolimus bind to FKBP506 but do not block the activation of mTOR. However, they result in inhibition of T-cell activation and lower smooth Drug-Eluting Medical Implants 309 muscle cell selectivity (Daemen and Serruys 2007). Xience V (Abbott Vascular) is an everolimus-eluting stent embedded in a durable fluoropolymer coating a thinstrut cobalt chromium platform (Stone et al. 2008). Everolimus proved rapid endothelialization with this stent compared to the sirolimus-eluting stent and the paclitaxel-eluting stent in rabbits (Joner et al. 2008). The everolimus-eluting stent was superior in performance compared to bare-metal stents in the SPIRIT I (Serruys 2005) trial, and non-inferior to the paclitaxel-eluting stent in terms of safety and effectiveness in the SPIRIT II (Serruys 2006) and III trials (Stone et al. 2008). Endeavor (Medtronic) is a zotarolimus-eluting stent on a Driver thin-strut cobalt chromium platform with a phosphorylcholine coating (Fajadet et al. 2006). The zotarolimus-eluting stent proved sustained clinical effectiveness up to four years in the ENDEAVOR I (Meredith et al. 2007a) trial, and superiority over bare-metal stents in the ENDEAVOR II (Fajadet et al. 2006) trial, but demonstrated lack of non-inferiority to the sirolimus-eluting stent in the ENDEAVOR III trial (Miyazawa et al. 2008). This may be explained by the fast drug release rate, since the zotarolimus is released within 2 days (Daemen and Serruys 2007). New agents, dual DES release, and a new pro-healing approach, are notable in DES-II. Sahajanand Medical Technologies designed a dual-layer heparin sirolimuseluting stent, which combines the antiproliferative action of sirolimus with the excellent biocompatibility and hemocompatibility of the heparin coating (Daemen and Serruys 2007). Pimecrolimus, primarily an anti-inflammatory agent, was loaded on a Conor stent and tested in a porcine model in two phases: pimecrolimus alone and combined with paclitaxel (Berg et al. 2007). The stent inhibited neointimal proliferation compared to bare-metal stents and proved the safety and efficacy of the dual concept. Genistein is a potent isoflavone, which possesses dose-dependent antiplatelet and antiproliferative properties (Daemen and Serruys 2007). A dual sirolimus genistein-eluting stent is currently under investigation. Furthermore, a new pro-healing endothelial progenitor cell capture stent consists of antibodies attached to a stainless steel stent which specifically target endothelial progenitor cells in the vascular circulation. The HEALING-FIM (Aoki et al. 2005b) trial demonstrated that the endothelial progenitor cell capture stent is safe and feasible. The HEALING-II trial reported 6-month in-stent restenosis rates of 17.2% with an associated in-stent late luminal loss of 0.78  0.39 mm that was reduced to 0.59  0.06 mm at 18 months (Daemen and Serruys 2007). 1.3.3 New Drug-Release Mechanisms The DES-II exhibited improved drug release matrix designs such as durable biodegradable polymers and layered release matrices for controlled release. EndeavorTM Resolute (Medtronic) is a new zotarolimus-eluting stent based on a novel BioLinxTM copolymer optimal for extended drug release (Meredith et al. 2007b). BioLinx is a unique blend of three different polymers: a hydrophobic polymer for delayed drug release, a lipophilic polymer for enhanced biocompatibility, and 310 M. Zilberman et al. a hydrophilic polymer for release burst. The Resolute stent elutes 85% of its zotarolimus content during the first 60 days and the remainder in 180 days in vivo (Meredith et al. 2007b). ZoMaxx (Abbott Laboratories) is another example of a smart matrix which elutes zotarolimus via a tri-layer coating consisting of a phosphorylcholine basecoat and a topcoat wrapped around a zotarolimus layer for slow elution (Abizaid et al. 2007). ZoMaxx also uses a novel Tri-Maxx stent platform which has a thin 3-layer tantalum sandwiched between two layers of stainless steel for enhanced fluoroscopic radiopacity. However, ZoMaxx was associated with a significantly high late loss and in-stent restenosis compared to Taxus in the ZoMaxx-I trial, which led Abbott to discontinue its program (Daemen and Serruys 2007). Biodegradable polymer-coated metal stents were first introduced in DES-II in an attempt to overcome the late risk associated with durable polymers. Guidant’s everolimus program consisted of the Biosensors Champion stent with a biodegradable polylactic acid coated S-stent with an elution profile of 70% drug release in 30 days (Grube et al. 2004), and 85% in 90 days (Tsuchiya et al. 2006). The FUTURE I (Grube et al. 2004), and II (Tsuchiya et al. 2006) trials reported superiority over bare-metal stents (Salam et al. 2006). However, the program was eventually discontinued by Guidant (Peterson 2004). The biolimus-eluting BioMatrix stent (Biosensors International) is coated with a poly-lactic acid bioabsorbable polymer that gradually releases the drug over 6–9 months. BioMatrix was tested in the STEALTH-1 (Daemen and Serruys 2007) trial, and exhibited superiority over bare-metal stents, non-inferiority with the paclitaxel-eluting stent (Chevalier et al. 2007) in the Nobori-I study, safety and efficacy in the NOBORI CORE trial (Ostojik et al. 2008) and non-inferiority with the sirolimus-eluting stent in the LEADERS trial in a 9-month follow-up (Windecker et al. 2008). Paclitaxeleluting InfinniumTM DES (Sahajanand Medical Technologies) is a slotted-tube stainless steel stent with lower strut thickness coated with three layers of biodegradable polymer: poly(DL-lactic acid-co-glycolic acid) (PDLGA) 50/50, PDLGAco-caprolactone (PDLGPCL) 75/25, and polyvinyl pyrrolidone (Vranckx et al. 2006). Paclitaxel is released in three release regimes: fast release during the first five days, followed by medium release in the next six days, and slow release until day 48. SIMPLE-I and II trials proved that InfinniumTM was safe and effective (Daemen and Serruys 2007). A classic approach offering an improved control over drug release was suggested by Conor Medsystems. Both their Conor Medstent (316L stainless steel based) (Serruys et al. 2005), and Conor Co-star (cobalt chromium based) (Krucoff et al. 2008) allow programmable and controlled paclitaxel release using a biodegradable polymer layer deposited in laser-cut holes embedded in the stent platform. Paclitaxel distribution and kinetics is obtained by using a poly-lactic-co-glycolic acid (PLGA) copolymer deposited within the different layers (Krucoff et al. 2008). The PISCES study proved that the Conor Medstent was safe and that the duration of release had a greater impact on the inhibition of the in-stent neointimal hyperplasia than the dose (Serruys et al. 2005). COSTAR I evaluated three dose-release formulations and reported that the low-dose slow-release formulation (10 mg Drug-Eluting Medical Implants 311 Fig. 4 DES-II selected clinical target vessel revascularization outcomes (EES everolimus-eluting stent, ZES – zotarolimus-eluting stent, BES – biolimus-eluting stent, COSTAR paclitaxel-eluting stent, SES – sirolimus-eluting stent (DES-I), PES – paclitaxel-eluting stent (DES-I), BMS – baremetal stent) paclitaxel release over 30 days) achieved the best clinical and angiographic outcomes (Kau 2007). The EUROSTAR (Dawkins et al. 2007) trial confirmed Co-star safety and efficacy in the low-dose slow-release formulation compared to the highdose slow-release formulation. Nonetheless, the COSTAR II study could not conclude that the CoStar DES is non-inferior in clinical and angiographic performance compared to the paclitaxel-eluting stent (Krucoff et al. 2008), leading to discontinuation of the program (Medsystems 2007). The promising Xtent (Biosensors International) biolimus-eluting cobalt chromium, polylactic acid coated stent combines all of the above-stated DES-II characteristics. It is tailored for long/multiple lesions and multivessel disease and allows in situ customization of stent length. The CUSTOM-I evaluation study reports 12-month safety and efficacy with Xtent (Grube 2008). The CUSTOM-II trial presented favorable 1-year follow-up results in high-risk multi-vessel-treated patients demonstrating safe in situ customization (Stella et al. 2008). Figure 4 presents clinical outcomes of selected second-generation stents and Fig. 5 illustrates selected drug-eluting matrix concepts of current DES. Table 2 summarizes selected metal drug-eluting stent designs. 1.4 Biodegradable Stents Biodegradable stents are now in an advanced stage of research and development, and are considered by many as the next generation of DES. Metal stents have thrombogenic properties (Tepe et al. 2002), and may therefore cause permanent physical irritation, with the risk of long-term endothelial dysfunction or 312 M. Zilberman et al. Fig. 5 Selected drug-eluting matrix illustrations (source: Kutryk 2003). (1) Boston Scientific; polystyrene-polyisobutylene polystyrene copolymer intermediate layer. (2) Cook; non-polymerized albumin coating. (3) Guidant; ethylene-vinyl acetate intermediate layer. (4) Conor-Medsystems: PDLGA copolymer intermediate layer. (5) Cordis: ethylene-vinyl acetate and poly(n-butyl methacrylate) polymer intermediate layers. (6) Abbot/Jomed: ethylene-vinyl acetate intermediate layer. (7) Sorin: no coating. (8) Jomed: nanoporous ceramic. (9) Igaki–Tamai: poly(L-lactic acid) stent. (10) BiodivYsio Matrix LO (for drugs with a molecular weight of <1200 Da); phosphorylcholine coating. (11) BiodivYsio Matrix HI (for drugs with a molecular weight of <1200 Da); phosphorylcholine coating inflammation (Farb et al. 1999). Permanent stents also pose a barrier for possible future bypass surgery (Waksman 2006) and noninvasive imaging. These disadvantages may be avoided by using a temporary scaffold, considering that the need for a permanent prosthesis decreases dramatically six months post-implantation (Tamai et al. 2000). Completely biodegradable stents may eliminate early and late complications of bare-metal stents and DES implantation by degrading into non-toxic substances after maintaining luminal integrity only during the period of high risk restenosis in the first 9–12 months (Commandeur et al. 2006). Finally, biodegradable stents have a higher capacity for drug incorporation, and allow more complex release kinetics by altering the biodegradation profile of the polymer (Commandeur et al. 2006). The main challenge in designing a biodegradable DES is overcoming the trade-off between mechanical properties and drug loading, since the radial compression strength of the stent is dramatically affected by the drug load. A second challenge is to effectively incorporate the drug during fabrication without damaging the molecules. The early bioresorbable stents (1980s) were designed as simple scaffolds and did not contain drugs. The bioresorbable materials used were PDLGA, polycaprolactone (PCL), poly(hydroxybutyrate hydroxyvalerate), and polyorthoester. All were associated with a significant inflammatory response and neointimal proliferation (Saito et al. 2002; Stack et al. 1988; Zidar et al. 1999). Tamai et al. (1999), Tamai et al. (2000) were the first to study a bioresorbable stent in clinical trials. The Igaki– Tamai stent is a poly L-lactic acid (PLLA) zig-zag helical coil design, and presented Endeavor Resolute/Medtronic PROMUSTM/Boston Scientific Cypher SelectTM Plus/Cordis (J&J) Infinnium/Sahajanand Medical Technologies CoStarTM/Conor Medsystems BioMatrix1/Biosensors International NOBORITM TERUMO ACHIEVETM/Cook-Guidant V Flex Plus PTXTM/ Cook JANUS/Sorin Biomedica GenousTM/OrbusNeich DexametTM/Abbott vascular XTENT1/Biosensors International Supralimus-CoreTM/Sahajanand Medical Technology Platform BX Velocity Matrix type Durable polymer Express 2 Durable polymer Driver1 Phosphorylcholine Multi-Link Durable polymer Vision 1 BioLinxTM Zotarolimus Cobalt chromium Driver1 A private-label version of Abbott’s XIENCETM V Everolimus Eluting Coronary Stent System Sirolimus 316L Stainless steel Undisclosed Durable hydrophilic coating (Rapamycin) Biodegradeable polymers Paclitaxel 316L Stainless steel Millennium Matrix1 Paclitaxel Cobalt chromium CoStarTM Biodegradable PDLGA layers Biolimus1 316L Stainless steel S-StentTM Biodegradable PLA NOBORITM is marketed with licensing agreement with biosensor under its technology. Paclitaxel 316L Stainless steel MULTI LINK Polymer-free PENTATM Paclitaxel 316L Stainless steel V-Flex Plus Polymer-free Tacrolimus 316L Stainless steel Carbostent Nanoporous ceramic coating Endothelial 316L Stainless steel R-Stent Polysaccharide, covalently progenitor cell coupled to stent surface capture Dexamethasone 316L Stainless steel Biodiv YsioTM Phosphorylcholine 1 Biolimus Cobalt chromium Custom NX1 Biodegradable PLA Sirolimus 316L Stainless steel Millennium Biodegradable Polymers Matrix1 Status FDA, CE Mark FDA, CE Mark FDA, CE Mark FDA, CE Mark CE Mark FDA CE Mark Drug-Eluting Medical Implants Table 2 Selected Coated Metal Drug Eluting Stent (DES-I and DES-II) DES/manufacturer Drug Stent material 316L Stainless steel Cypher/Cordis (J&J) Sirolimus (Rapamycin) Paclitaxel 316L Stainless steel Taxus LibertéTM/Boston Scientific Zotarolimus Cobalt chromium Endeavor1/Medtronic Everolimus Cobalt chromium XIENCETM V/Abbott vascular CE Mark CE Mark CE Mark CE Mark CE Mark CE] Mark CE Mark CE Mark approved in Taiwan CE Mark Trial Trial (continued) 313 Drug Zotarolimus Stent material Stainless steel/ tantalum/ stainless steel 316L Stainless steel Platform Tri-Max Matrix type Phosphorylcholine Status Suspended BiodivYsioTM Phosphorylcholine Suspended QueST Batimastet BiodivYsio (Biocompatibles and British Biotech) QuaDDS-QP2/Quanam Medical and Boston Scientific Antinomycin D stent/ Guidant Batimastet 7-hexanoyltaxol (QP2) Antinomycin D 316L Stainless steel Jomed Flexmaster/Jomed Tacrolimus 316L Stainless steel 316L Stainless steel Rigid polymer/drug sleeve cover MULTILINK Durable polymer (Ethylene tetra-D stent vinyl acetate) Jomed Nanoporous ceramic layer Flexmaster 314 Table 2 (continued) DES/manufacturer Zomaxx/Abbott Labratories Suspended Suspended Suspended M. Zilberman et al. Drug-Eluting Medical Implants 315 intimal hyperplasia in rates comparable to bare-metal stents in six months, and 18% TVR in a 4-year follow-up (Waksman 2006). Lincoff et al. (1997) demonstrated that high-molecular-weight PLLA performance was more favorable than lowmolecular-weight PLLA in terms of an inflammatory reaction. Biodegradation of PLLA is achieved by hydrolytically unstable ester linkages in the backbone of the polymer, which result in chain scission into oligomers and eventually erosion with an overall degradation time which depends on the initial molecular weight (Commandeur et al. 2006). The oligomers are then broken down into lactic acid, and are completely metabolized via the Krebs cycle (Ormiston et al. 2008) Early biodegradable drug-eluting stent designs were based on fiber or film structures. Yamawaki et al. (1998) incorporated Tranilast (ST638) or ST494 (an inactive metabolite of ST638) agents into the Igaki–Tamai stent. The stent presented significantly less neointimal formation and geometric remodeling with ST638 than with ST494. Uurto et al. (2005) evaluated a monofilamentbased stent made of a polymer consisting of 96% L-lactic acid and 4% D-lactic acid coated with a 50/50 ratio of two bioactive agents: dexamethasone and simvastatin. The stent presented acceptable results in a porcine model. Vogt et al. (2004) reported paclitaxel-loaded poly(D,L-lactic acid) (PDLLA) doublehelical stent exhibiting sufficient mechanical stability with a very slow release pattern of paclitaxel in a porcine model. Their 2-month evaluation demonstrated effective proliferation inhibition, but also local inflammatory effects due to polylactide resorption. Alexis et al. (2004) incorporated paclitaxel and rapamycin into PDLLA and PDLGA non-expandable helical stents prepared from film strips exhibiting a homogenous, burst-free drug release. Ye et al. (1998) demonstrated the successful transfer and expression of a nuclear-localizing b-Gal reporter gene in cells in the arterial wall of rabbits after the implantation of biodegradable stents made of PLLA/PCL films. These stents were made of a porous tubular structure impregnated with a recombinant adenovirus carrying that gene and demonstrated an exciting possibility for restenosis prevention. The multiple lobe PLLA fiber-based stent (Nguyen et al. 2004) was coated with drug-loaded microspheres in order to combine good mechanical properties with the desired drug release profile (Zilberman et al. 2004). These microsphere reservoirs, which were prepared by the double-emulsion technique, can be loaded with biologically active aqueous or non-aqueous molecules. Since mild materials and processing steps are used, these microspheres can be loaded with all drugs, proteins and gene transfer vectors. Current drug-eluting biodegradable stents in the stage of clinical trials are described below. The everolimus bioasbsorbable stent (BVS, Abbott Vascular) consists of a bioabsorbable PLLA base coated with a more rapidly degrading PDLLA coating and releases 80% of its drug in 28 days and is kept at 20 C in order to extend its shelf-life (Ormiston et al. 2008). The first in-men ABSORB trial (n ¼ 30) demonstrated high procedural success, with a collapse pressure similar to a stainless steel stent, with 0% TLR, 0% stent thrombosis, one patient had a myocardial infraction, and late loss of 0.44  0.35 mm in one year (Ormiston et al. 2008). The REVA bioresorbable stent uses a tyrosine-derived polycarbonate 316 M. Zilberman et al. platform and utilizes a slide and lock mechanism rather than deforming during usage. REVA Medical announced the enrollment of first-in-man RESORB study for paclitaxel-coated REVA (REVA Medical 2007) backed by promising animal trials (Waksman 2006). An interesting new approach was presented by Bioabsorbable Therapeutics, Inc. (BTI): a sirolimus-eluting biodegradable stent composed of salicylic acid (active ingredient in aspirin) chemically incorporated into polyanhydride layers (Buchbinder 2007). 1.5 Novel Drug-Eluting Highly Porous Stent Coatings As mentioned above, one of the issues that should be addressed in the field of DES is the very slow release rate of highly hydrophobic antiproliferative drugs, such as paclitaxel. We developed and studied paclitaxel-eluting porous coatings (shell) for both metal and polymeric (biodegradable) stents in order to address this issue (Kraitzer and Zilberman 2007). The coating preparation is based on the freezedrying of inverted emulsions technique. The investigation of these new coatings focused on the effects of the emulsion’s composition (formulation) and process kinetics on the long-term drug release from fibers, in light of the shell’s morphology and degradation profile. The paclitaxel release exhibited three phases, which corresponded to the degradation profile of the host PDLGA. We found that the effect of the emulsion formulation on the release profile is more significant than the effect of the process kinetics. The copolymer composition had the most dominant effect on the drug release profile from the composite fibers. An increase in the glycolic acid content of the copolymer (or decrease in lactic acid content) resulted in a tremendous increase in the release rate during the second phase, which was attributed mainly to the increased degradation rate and decreased drug attachment to the host polymer (Fig. 6a). The paclitaxel release profile was improved and we concluded that emulsions with a less hydrophobic nature are favorable for effective controlled release of the hydrophobic paclitaxel from the porous shell. Farnesylthiosalicylate (FTS, Salirasib) is a rather specific non-toxic new agent which was recently developed at the Tel-Aviv University, Israel (Kloog and Cox 2004). It acts as a Ras antagonist (George et al. 2004; Kloog and Cox 2004) and has a mild hydrophobic nature. In its active form (GTP-bound) Ras promotes enhanced cell proliferation, tumor cell resistance to drug-induced cell death, cell migration and invasion. Ras is therefore considered an important target for cancer therapy as well as for therapy of other proliferation diseases, including restenosis. The apparent selectivity of FTS towards active (GTP-bound) Ras and absence of toxic or adverse side-effects was proven in animal models (George et al. 2004) and in human trials (phase I performed at MD Anderson Cancer Center). FTS was found to be a potent inhibitor of intimal thickening in the rat carotid artery injury model which serves as a model for restenosis where it does not interfere with endothelial proliferation (George et al. 2004). The incorporation of the new drug FTS in a stent Drug-Eluting Medical Implants 317 Fig. 6 The effect of copolymer composition on the cumulative drug release profile from core/shell fiber structures (~ – PDLGA 50/50, ● – PDLGA 75/25): (a) paclitaxel release, (b) FTS release. Source (a) Kraitzer et al. 2008, (b) Kraitzer et al. 2009 coating may overcome the incomplete healing and lack of endothelial coverage associated with current DES. Our novel porous coatings were loaded with the new agent FTS (Kraitzer et al. 2009). Our results show that the most important parameter affecting release in this system was again the copolymer composition. An increase in the glycolic acid content of the PDLGA copolymer enhanced the burst effect and release rate during the first two weeks, mainly due to higher water uptake and swelling but also due to a higher degradation rate of the host polymer (Fig. 6b). The FTS release from our highly porous coatings is faster, more adjustable and totally different from that of 318 M. Zilberman et al. paclitaxel. Paclitaxel is more hydrophobic than FTS and creates more specific interactions with the host polyesters. Therefore, paclitaxel’s diffusion through the host polymer is much slower and all changes in formulation parameters affect its release profile mainly after 10 weeks of degradation. In paclitaxel-eluting systems the emulsion formulation influences diffusion by producing binding regions for the drug as more hydrophobic materials are introduced into the emulsion, thus delaying the drug molecules. The release profile of FTS from our composite fibers is therefore more suitable for the stent application. Furthermore, since FTS is less toxic, some burst release can be tolerated and may be beneficial. In summary, in this section we described the two vascular drug-eluting metal stent generations, DES-I and DES-II, in terms of stent platform, incorporated APIs, drug-release profile and stent functioning. We also discussed biodegradable drugeluting stents and our novel drug-eluting porous coatings, which can be applied on both metal and biodegradable stents. It can be concluded that the field of drugeluting vascular stents has progressed significantly over the past several years. The combination of the above-described new coatings and new APIs that have recently been developed will further improve this life-saving device. 2 Drug-Eluting Wound Dressings 2.1 Introduction: Infection, Wound Dressings and Local Antibiotic Release The skin is regarded as the largest organ of the body and has many different functions. Wounds with tissue loss include burn wounds, wounds caused as a result of trauma, diabetic ulcers and pressure sores. The regeneration of damaged skin includes complex tissue interactions between cells, extracellular matrix (ECM) molecules and soluble mediators in a manner that results in skin reconstruction. The moist, warm, and nutritious environment provided by wounds, together with diminished immune functioning secondary to inadequate wound perfusion, may allow build-up of physical factors such as devitalized, ischemic, hypoxic, or necrotic tissue and foreign material, all of which provide an ideal environment for bacterial growth (Jones et al. 2004). In burns, infection is the major complication after the initial period of shock. It is currently estimated that about 75% of the mortality following burn injuries is related to infections rather than to osmotic shock and hypovolemia (Revathi et al. 1998). Infection is defined as a homeostatic imbalance between the host tissue and the presence of microorganisms at concentrations that exceeds 105 organisms per gram of tissue or the presence of beta-hemolytic streptococci (Sussman and Bates-Jensen 2001; Xu et al. 2004). The main goal of treating the various types of wound infections should be to reduce the bacterial load in the wound to a level at which wound healing processes can take place. Otherwise, the formation of an infection Drug-Eluting Medical Implants 319 can seriously limit the wound healing process, can interfere with wound closure and may even lead to bacteremia, sepsis and multi-system failure. Evidence of bacterial resistance is on the rise, and complications associated with infections are therefore expected to increase in the general population. Various wound dressings aim to restore the milieu required for skin regeneration and to protect the wound from environmental threats and penetration of bacteria. Although traditional gauze dressings offer some protection against bacteria, this protection is lost when the outer surface of the dressing becomes moistened by wound exudates or external fluids. Furthermore, traditional gauze dressings exhibit low restriction of moisture evaporation which may lead to dehydration of the wound bed. This may lead to adherence of the dressing, particularly as wound fluid production diminishes, causing pain and discomfort to the patient during removal. Most modern dressings are designed according to the well-accepted bilayer structural concept: an upper dense “skin” layer to prevent bacterial penetration and a lower spongy layer designed to adsorb wound exudates and accommodate newly formed tissue. Unfortunately, dressing material adsorbed with wound discharges provides conditions that are also favorable for bacterial growth. This has given rise to a new generation of wound dressings with improved curative properties that provide a local antimicrobial effect by eluting various germicidal compounds. Local delivery of antibiotics and disinfectants addresses the major disadvantages of the systemic approach, namely poor penetration into ischemic and necrotic tissue typical of post-traumatic and postoperative tissue, renal and liver complications, and need for hospitalized monitoring (Price et al. 1996; Ruszczak and Friess 2003) by maintaining a high local antibiotic concentration for an extended duration of release without causing systemic toxicity (Gristina 1987; Springer et al. 2004; Zalavras et al. 2004). The effectiveness of such devices is strongly dependent on the rate and manner in which the drug is released (Wu and Grainger 2006). These are determined by the host matrix into which the antibiotic is loaded, the type of drug/disinfectant and its clearance rate. If the agent is released quickly, the entire drug could be released before the infection is arrested. If release is delayed, infection may set in further, thus making it difficult to manage the wound. The release of antibiotics at levels below the minimum inhibitory concentration (MIC) may lead to bacterial resistance at the release site and intensify infectious complications (Gold and Moellering 1996; Gransden 1997). A local antibiotic release profile should therefore generally exhibit a considerable initial release rate in order to respond to the elevated risk of infection from bacteria introduced during the initial shock, followed by a sustained release of antibiotics at an effective level, long enough to inhibit latent infection (Ruszczak and Friess 2003). The location, size and degree of injury as well as the rate of tissue regeneration affect the wound healing process, and are reported to typically last 3–7 weeks (Roenigk and Roenigk 1989). This section describes the main features of wound dressings based on both synthetic and natural polymers. 320 2.2 M. Zilberman et al. Wound Dressings Based on Synthetic Polymers A variety of dressings that contain and release antibiotic/disinfectant agents at the wound surface have been introduced to the market. Most of these dressings have been designed to provide controlled release of silver ions through a slow but sustained release mechanism which helps avoid toxicity yet ensures delivery of a therapeutic dose of silver ions to the wound (Heggers et al. 2005). Various dressing formats, such as foams (Contreet1 antimicrobial foam, Coloplast), membranes (PolyMem Silver, Ferris), hydrocolloids (Urgotul SSD, Urgo), alginates (Silvercel1, Johnson & Johnson), and hydrofibers (Aquacel, ConvaTec) are available. For instance, Acticoat1 (Smith and Nephew) is a 3-ply gauze dressing made of an absorbent rayon polyester core, with upper and lower layers of a nanocrystalline silver-coated high-density polyethylene mesh (Fraser et al. 2004). It is applied wet and is then moistened with water several times daily to allow the release of the silver ions so as to provide an antimicrobial effect for three days. Concerns have been raised by clinicians regarding the safety of the silver ions included in most of these products. For example, it was found that a young person with 30% mixed depth burns who received one week of local treatment with Acticoat1 exhibited hepatotoxicity and argyria-like symptoms and the silver levels in his plasma and urine as well as the liver enzymes were clearly elevated during the treatment period. The authors therefore raised concern about potential silver toxicity and suggested monitoring the silver levels in the plasma or urine during treatment (Trop 2006). In order to address this issue, the silver in Actisorb1 (Johnson & Johnson) is impregnated into an activated charcoal cloth, after which it is encased in a nylon sleeve which does not enable the silver in the product to be freely released at the wound surface but nevertheless eradicates bacteria that adsorb onto the activated charcoal component. Suzuki et al. (Suzuki et al. 1997, 1998) presented a new concept for an antibiotic delivery system that releases gentamicin only in the presence of wounds infected by Pseudomonas aeruginosa. Gentamicin is bound to a polyvinyl alcohol derivative (PVA) hydrogel through a specially developed peptide linker cleavable by a proteinase. This allows gentamicin to be released at specific times and locations, namely when and where P. aeruginosa infection occurs. PVA-(linker)-gentamicin demonstrated selective release of gentamicin in P. aeruginosa-infected wound fluid, and caused a significant reduction in its growth in vitro. The substantial disadvantage of the majority of the available synthetic wound dressings is the fact that like textile wound dressings, the necessary change of dressings may be painful and increases the risk of secondary contamination. Bioresorbable dressings successfully address this shortcoming since they degrade from the wound surface once they have fulfilled their role. Film dressings made of lactide caprolactone copolymers such as Topkin1 (Biomet, Europe) and Oprafol1 (Lohmann & Rauscher, Germany) are currently available (Jürgens et al. 2006). Biodegradation of the film occurs via hydrolysis of the copolymer into lactic Drug-Eluting Medical Implants 321 acid and 6-hydroxycaproic acid. During the hydrolytic process the pH shifts towards the acidic range, with pH values as low as 3.6 measured in vitro (Jürgens et al. 2006). Although these two dressings do not contain antibiotic agents, it is claimed that the low pH values induced by the polymer’s degradation help reduce bacterial growth (Varghese et al. 1986) and also promote epithelialization (Eisinger et al. 1979). Furthermore, local lactate concentrations can stimulate local collagen synthesis (Hutchinson and Furr 1985). Film dressings are better suited for small wounds, since they lack an absorbing capacity and are impermeable to water vapors and gasses, which may cause accumulation of wound fluids on larger wound surfaces. Fiber-based wound dressings with antibiotic delivery offer a good alternative to the previously described films by providing a high surface area for controlled release and improved absorbency and pliability. Antibiotics are typically incorporated into these fibers during the process of fiber spinning (e.g., electrospinning or solution spinning). Katti et al. (2004) focused on the development of a biodegradable non-woven PDLGA fiber mesh dressing, made by means of an electrospinning process. Briefly, the process of electrospinning involves use of a polymer solution that is contained in a syringe and held at the end of the needle by its surface tension. Charge is induced on the solution by an external electric field to overcome the surface tension and form a charged jet of solution. As this jet travels through air, it experiences instabilities and follows a spiral path. Evaporation of the solution leaves behind a charged polymer fiber that is collected on a grounded metal screen. It was shown that the antibiotic cefazolin can be successfully incorporated into the fibers in this way, and even though its release from these fibers has not yet been reported, the effects of process parameters such as orifice diameter, applied voltage and polymer and drug solution concentrations were investigated. Chang et al. (2008) created gentamicin-eluting fibers by gravity spinning of a PCL with gentamicin. The in vitro release of gentamicin from their fibers lasted 50 days (Chang et al. 2008). 2.3 Wound Dressings Based on Natural Polymers Natural polymers such as collagen (Lee et al. 2002; Park et al. 2004; Prabu et al. 2006; Shanmugasundaram et al. 2006; Sripriya et al. 2004), chitosan (Aoyagi et al. 2007; Chung et al. 1994; Mi et al. 2002; Muzzarelli et al. 1990; Rossi et al. 2007) and alginate (Knill et al. 2004) have been investigated in various forms for wound dressing applications as either main or additional components of the dressing structure which are able to impact the local wound environment beyond moisture management and elicit a cellular response. Collagen is the main structural protein of the ECM, and was one of the first natural materials to be utilized for skin reconstruction and dressing applications. Collagen-based products have been available commercially for over a decade, ranging from gels, pastes and powders to more elaborate sheets, sponges, and composite structures. Collagen’s limitations as a 322 M. Zilberman et al. wound dressing ingredient are mainly due to its rapid biodegradation by collagenase and its susceptibility to bacterial invasion (Cairns et al. 1993; Maruguchi et al. 1994; Pruitt and Levine 1984; Trafny et al. 1998). Drug-eluting collagen sponges have been found useful in both partial-thickness and full-thickness burn wounds. Collatamp1 (Innocoll GmbH, Germany), Syntacoll (AG, Switzerland), Sulmycin1Implant (Schering-Plough, USA) and Septocoll1 (Biomet Merck, Germany) are several such products which have been found to accelerate both granulation tissue formation and epithelialization, as well as to protect the recovering tissue from potential infection or re-infection by eluting gentamicin. In vitro, the drug is released by a combination of diffusion and natural enzymatic breakdown of the collagen matrix (Radu et al. 2002). A comprehensive clinical study of gentamicin collagen sponges demonstrated their ability to induce high local concentrations of gentamicin (up to 9,000 mg mL 1) at the wound site for at least 72 h while serum levels remained well below the established toxicity threshold of 10–12 mg mL 1 (Ruszczak and Friess 2003). Simple collagen sponge entrapment systems are characterized by high drug release upon the wetting of the sponge, typically within 1–2 h of application. Sripriya et al. (2004) have suggested improving the release profile of such systems by using succinylated collagen which can create ionic bonds with the cationic antibiotic ciprofloxacin so as to restrain its diffusion. It is claimed that in this way ciprofloxacin release corresponds to the nature of the wound in line with the amount of wound exudates absorbed in the sponge. Effective in vitro release from their system was found to last five days, and was proven successful in controlling infection in rats. Other studies have aimed to better control drug release or improve wound healing properties by combining collagen with other synthetic or natural biodegradable elements. Prabu et al. (2006) focused on achieving a more sustained release of the antimicrobial agent and described a dressing made from a mixture of collagen and PCL loaded with gentamicin and amikacin, whereas Shanmugasundaram et al. (2006) chose to impregnate collagen with alginate microspheres loaded with the antibacterial agent silver sulfadiazine (AgSD). Other studies which focused on improving wound healing capabilities tried to incorporate tobramycin, ciprofloxacin (Park et al. 2004) and AgSD (Lee et al. 2002) into collagen hyaluronan-based dressings. The two latter studies did not show conclusive evidence of improved healing properties compared to their control. However, hyaluronan, a structure-stabilizing component of the ECM, is thought to play a role in several aspects of the healing process with hyaluronan-based dressings, and exhibited promising results in the management of chronic wounds such as venous leg ulcers (Colletta et al. 2003; Taddeucci et al. 2004). A wide range of studies describe the employment of the polysaccharide chitosan and its partially deacetylated derivative chitin as structural materials analogous to collagen for wound dressings. Both materials offer good wound protection and have also been found to promote wound healing without excessive granulation tissue and scar formation (Chung et al. 1994). Chitosan has also been documented as displaying considerable intrinsic antibacterial activity against a broad spectrum of bacteria (Muzzarelli et al. 1990). Ignatova et al. (2006) reported the electrospinning of Drug-Eluting Medical Implants 323 chitosan with PVA into non-woven nanofiber mats with good in vitro bactericidal activity against Staphylococcus aureus and Escherichia coli (Ignatova et al. 2006). Another interesting fibrous form which combines the polysaccharides chitosan and alginate was reported by Knill et al. (2004), who developed a composite structure of calcium alginate filaments coated with chitosan, utilizing the cationic interaction of chitosan with the anionic nature of alginate to bond the two together. It has been suggested that the core alginate fiber may manage excess exudates whereas chitosan would provide antibacterial, hemostatic and wound healing properties. In this case too, antibacterial testing of the fibers demonstrated an antibacterial effect. Several attempts to improve the chitosan dressing’s antibacterial capabilities by incorporating various agents such as AgSD (Mi et al. 2002), chlorhexidine diacetate (Rossi et al. 2007) and minocycline hydrochloride (Aoyagi et al. 2007) have been reported. 2.4 Composite Fiber Structures Loaded with Antibacterial Drugs for Wound Healing Applications Drug-eluting fibers can be used for various biomedical applications. Few controlledrelease fiber systems based on polymers have been investigated to date. The two basic types of drug-loaded fibers that have been reported are monolithic fibers in which the agent is dissolved or dispersed throughout the polymer fiber, and hollow reservoir fibers in which the agent is added to the internal section of the fiber. The advantages of drug-loaded fibers include ease of fabrication, high surface area for controlled release and localized delivery of bioactive agents to their target. Disadvantages of monolithic and reservoir fibers include poor mechanical properties due to drug incorporation and limitations in drug loading. Furthermore, many small molecules and all proteins do not tolerate melt processing and organic solvents. In one of our recent studies we presented a new concept of core/shell fiber structures which successfully meet these challenges (Zilberman et al. 2009). These composite fibers combine a dense polyglyconate core fiber and a drug-loaded porous 75/25 PDLGA shell structure, i.e., the antibacterial drug molecules are located in a separate compartment (a “shell”) around the “core” fiber. A general view of our composite fibers and a schematic representation showing the core/shell concept are presented in Fig. 7a, b, respectively. The shell is prepared using freezedrying of inverted emulsions with mild processing conditions. These unique fibers are designed to be used as basic elements of bioresorbable burn and ulcer dressings. The main goal of these studies was to investigate core/shell fiber structures loaded with the antibiotic drugs gentamicin sulfate, ceftazidime pentahydrate and mafenide acetate. The first two antibacterial drugs are broad spectrum antibiotics which can be used systemically or locally, whereas the third is typically used in burn dressings. Their investigation focused on the effects of the emulsion’s composition 324 M. Zilberman et al. Fig. 7 The structure of the composite core/shell fiber structures: (a) general view (photograph) of the fiber, (b) schematic representation showing the core dense fiber and the porous drug-loaded shell, (c) and (d) SEM fractographs of part of a core fiber coated with drug-loaded porous PDLGA shell (cross section). Good adhesion between core and shell is demonstrated (Elsner and Zilberman 2008) (formulation) on the shell microstructure, on the drug release profile from the fibers and on the resulting bacterial inhibition. The freeze-drying technique is unique in being able to preserve the liquid structure in solids. We used this technique in order to produce the shell from inverted emulsions in which the continuous phase contained polymer dissolved in a solvent, with water and the antibiotics dissolved in it as the dispersed phase. SEM fractographs showing the bulk morphology of the reference specimen are presented in Fig. 7c, d. The quality of the interface between the fiber and the porous coating is high (Fig. 7d), i.e., the preliminary surface treatment enabled good adhesion between the core and the shell. The shell’s microstructure affects the drug release profile and can also serve as a good measure of the emulsion’s stability. The shell’s porous structure contains round pores with a diameter of 0.5–5 mm and a porosity of 16–82%. The release profiles commonly exhibited an initial burst effect accompanied by a decrease in release rates with time over periods ranging from several days to 50 days, depending on the formulation (Elsner and Zilberman 2009). The effects of the emulsion’s parameters on the shell’s microstructure and on the drug release profile are demonstrated on ceftazidime-loaded fibers and presented in Figs. 8 and 9, respectively. Higher organic: aqueous (O:A) phase ratios, polymer content and molecular weight (MW) reduced the burst release of antibiotics from the fibers and prolonged their release due to changes in the shell’s structure. A higher MW and polymer content Drug-Eluting Medical Implants 325 Fig. 8 SEM fractographs of core/shell fibers showing the effect of change in certain formulation parameters on shell microstructure. (a) reference formulation 5% (w/w) ceftazidime, 15% (w/v) polymer (75/25 PDLGA, MW 100 KDa), O:A phase ratio of 6:1, (b) O:A modified to 12:1, (c) polymer content modified to 20% (w/v) polymer, (d) higher polymer MW (modified to 240 kDa) (Elsner and Zilberman 2008) resulted in a larger effect on the microstructure and release profile than the O:A phase ratio (Elsner and Zilberman 2008). Albumin was found to be the most effective surfactant for stabilizing the inverted emulsions (Elsner and Zilberman 2008; Zilberman et al. 2009). As a surfactant, it is located at the interface between the aqueous phase and the organic phase, reduces the interfacial tension between the two phases and therefore significantly decreases the pore size. It also enables a high encapsulation efficiency and a relatively low burst release followed by a moderate release profile which enables prolonged drug release. This behavior probably results from albumin’s ability to bind the drug molecules (especially mafenide acetate) through specific interactions. The effect of albumin on the release profile of mafenide acetate from the fibers is presented in Fig. 10(a) and the release profiles of the three investigated antibiotics from fibers with shells that contain albumin are presented in Fig. 10(b). The ability of albumin to bind APIs is well-known (Foye et al. 1998). Albumin can interact with acidic or basic agents via van der Waals dispersion forces, hydrogen bonds and ionic interactions. Based on these results, we chose albumin as the preferred surfactant in our systems. We also performed microbiological experiments in order to monitor the effectiveness of various concentrations of the antibiotic released from the fibers in terms of the residual bacteria compared with the initial number of bacteria. Bacteria in PBS only served as the control. We chose the following three types of gentamicineluting fibers with different release profiles (Fig. 11): 326 M. Zilberman et al. Fig. 9 In vitro release of ceftazidime from core/shell fibers demonstrating the effect of change in certain formulation parameters compared to the reference formulation: (a) effect of O:A phase ratio: ■ – 6:1, ● – 8:1, ~ – 10:1, r – 12:1, (b) effect of polymer content: ■ – 15% (w/v), ● – 17.5 (w/v), r – 20% (w/v), (c) effect of polymer MW: ■ – 100 kDa, ● – 240 kDa (Elsner and Zilberman 2008) Drug-Eluting Medical Implants 327 Fig. 10 Drug release profiles from fibers containing albumin as surfactant: (a) mafenide acetate release profiles from core/shell fiber structures containing 5% (w/w) drug, 15% (w/v) polymer and O:A phase ratio of 6:1: ~ no surfactant r 0.5% (w/v) surfactant; ■ 1% (w/v) surfactant. (b) Antibiotic release from fibers containing 15% (w/w) polymer, O:A phase ratio of 6:1, 5% (w/w) drug and 1% (w/v) albumin: ■ mafenide, ~ – ceftazidime, ● – gentamicin (Elsner and Zilberman 2008) 1. Fibers with a shell based on 26.7% (w/v) polymer, 20% (w/v) gentamicin, 6:1 O: A phase ratio and albumin as surfactant. These fibers demonstrated a moderate burst release of 32% followed by a moderate release profile. 2. Fibers with a shell based on 20% (w/v) polymer, 20% (w/v) gentamicin, 6:1 O: A. These fibers demonstrated a high burst release of approximately 60%. 3. Fibers with a shell based on 26.7% (w/v) polymer, 5% (w/v) gentamicin, 6:1 O: A phase ratio and albumin as surfactant. These fibers demonstrated a low burst release of 13% during the first day and 60% within three days. After three days the release pattern was similar to that of sample II. 328 M. Zilberman et al. Fig. 11 Gentamicin’s release profile from fiber samples which were used for microbiological evaluation: ■ – sample I (26.7% (w/v) polymer, 20% (w/v) drug, 6:1 O: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 and 5% (w/v) albumin) (Zilberman et al. 2009) The bacterial strains used in this study were S. aureus, S. epidermidis and Pseudomonas aeruginosa. Their minimal inhibitory concentration (MIC) values are 2.5, 5 and 6.3 mg mL 1, respectively. All three strains were clinically isolated. These strains were chosen because they are prevalent in wound infections, especially S. aureus and P. aeruginosa. The third strain, S. epidermidis, usually comprises the normal flora of the skin. However, under grave conditions it can cause wound infections. Moreover, these bacteria can produce bio-films, which prevent antibiotics from reaching the target, therefore causing resistance. The bacteria were added at the beginning of the fibers’ release in order to simulate contamination at the time of implantation. The results for all three bacterial types when using a relatively high initial bacterial concentration of 1107 CFU mL 1 are presented in Fig. 12. The released gentamicin significantly decreased bacterial viability, and practically no bacteria survived after two days. The fiber preparation did not affect gentamicin’s potency as an antibacterial agent. Our new fiber structures are thus effective against the relevant bacterial strains and can be used as basic elements of bioresorbable drug-eluting wound dressings. In practice, a wound dressing can be woven from a combination of several types of fibers to create a resultant release profile which is the product of several release profiles or drug types. The diverse profiles achieved in this study with higher and Drug-Eluting Medical Implants 329 Fig. 12 Number of colony forming units (CFU) versus time, when an initial bacterial concentration of 1107 CFU/mL was used: A – Staphylococcus aureus, B – Staphylococcus epidermidis, C – Pseudomonas aeruginosa. The releasing fibers are: – sample I (26.7% (w/v) polymer, 20% (w/v) drug, 6:1 O: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 and 5% (w/v) albumin), – control – reference fiber without gentamicin (Zilberman et al. 2009) lower burst release rates and with varying elution spans may serve as a good basis for further in vivo examination of the fibers in order to create the ideal profile for a particular wound-healing application. 330 M. Zilberman et al. In conclusion, this section described wound dressings with emphasis on approaches for local drug release for prevention of bacterial infections that accompany the initial period of shock, due to wounds with tissue loss. Drug-eluting wound dressings based on both synthetic and natural polymers were described in terms of matrix material, processing techniques, dressing format and drug release profile. Our novel composite antibiotic-eluting composite fibers which are designed to be used as basic elements of wound dressings were described. Recent advances in the field of wound dressings aim to enable improvements in the patient’s quality of life due to higher biocompatibility and controlled release of bioactive agents that enhances the healing process and eliminate the need for change of dressing, which is painful and increases the risk for secondary contamination. Wound dressings that combine controlled release of antibiotics with controlled release of growth factors will further advance the field of wound healing. 3 Protein-Eluting Scaffolds for Tissue Regeneration Organ or tissue deficiency or loss is one of the most frequent and devastating problems in human healthcare. Tissue engineering is described as “an interdisciplinary field that applies the principles of engineering and life sciences towards the development of biological substitutes that restore, maintain, or improve tissue function or a whole organ” (Langer and Vacanti 1993). One of the major approaches in tissue engineering is scaffolds that elute bioactive agents. Upon implantation of such scaffolds, cells from the body are recruited to the site and a tissue can be formed (Howard et al. 2008). Growth factors are essential for promoting cell proliferation and differentiation; however, their direct administration is problematic due to their poor in vivo stability (Babensee et al. 2000; Chen and Mooney 2003). It is therefore necessary to develop scaffolds with controlled delivery of bioactive agents that can achieve prolonged availability as well as protection of these bioactive agents which might otherwise undergo rapid proteolysis (Buket Basmanav et al. 2008; Zhu et al. 2008). The main obstacle to successful incorporation and delivery of small molecules as well as proteins from scaffolds is their inactivation during the process of scaffold manufacture due to exposure to high temperatures or harsh chemical environments, which is why it is necessary to develop a method that minimizes protein inactivation. Growth factors can be attached onto scaffolds by adsorption to the surface of the scaffolds. For example, Elcin and Elcin (2006) formed sponges made of PLGA using a solvent casting and particulate leaching method. The sponges were immersed overnight in a solution containing vascular endothelial growth factor (VEGF) at a temperature of 4 C and then freeze-dried. VEGF release from the sponges was examined in vivo in Wistar rats. Rapid release of VEGF was observed in the first two days, followed by a slow release for two weeks. Furthermore, the released VEGF retained its bioactivity. Drug-Eluting Medical Implants 331 Growth factors can also be incorporated into scaffolds directly during the fabrication process. Liu et al. (2003) fabricated thin films using the solvent casting and particulate leaching technique. A solution containing bone morphogenetic protein 7 (BMP-7) was added to the polymer solution. The final solution was exposed to 20 C and the solvent was slowly evaporated. Cells from a muscle tissue of New Zealand white rabbits were seeded on the BMP-7 eluted films and differentiated into osteoblast-like cells. Whang et al. (1995, 1998) presented an important method for incorporating bioactive molecules during the scaffold-production processing, based on the technique of freeze-drying an inverted emulsion. This method enables the production of highly porous scaffolds with a controllable pore size and the ability to incorporate and release proteins in a controlled manner. Since the bioactive molecules are incorporated into the aqueous solution of the emulsion, very sensitive bioactive molecules can be incorporated without exposing them to organic solvents, thus preserving their activity. A recombinant human bone morphogenetic protein (rhBMP-2) was incorporated and released in vivo in a rat ectopic bone induction assay from a scaffold produced by freeze-drying an inverted emulsion. The rhBMP2 incorporated scaffold-induced bone formation, which confirmed the viability of the rhBMP-2 released from the scaffold. Contact radiography, radiomorphometry, histology, and histomorphometry revealed significantly more bone in the rhBMP2 implants than in the controls. In one of our recent studies we investigated highly porous scaffolds that were produced using the technique of freeze-drying inverted emulsions (Grinberg 2007). The enzyme horseradish peroxidase (HRP) was used as a protein source in order to examine the effect of the emulsion’s formulation on the film microstructure and on the resulting cumulative protein release. A porous structure with 72–93% porosity and partially interconnected pores was obtained. The released profiles usually exhibited an initial burst effect, accompanied by a decrease in release rate with time, as is typical of diffusion-controlled systems. The release profile was mostly affected by the polymer’s initial molecular weight and the enzyme. An increase in the polymer’s initial molecular weight resulted in a decrease in the initial burst effect and in the release rate due to changes in the microstructure. An increase in the enzyme content resulted in an increase in the initial burst effect and in the release rate due to a higher driving force for diffusion. The other emulsion parameters had minor effects on the release profile in the examined range. We demonstrated that appropriate selection of the inverted emulsion’s parameters results in the desired release profile which can be adjusted to the specific application. We observed cell adherence and growth with fibroblast characteristics on the film. The studied films are thus biocompatible and can be used as a controlled release system for bioactive agents in tissue regeneration applications. Another approach for encapsulation of bioactive molecules in scaffolds is composite scaffold/microsphere structures. The growth factors are initially encapsulated into the microspheres, and the microspheres can then be integrated into scaffolds. Zhu et al. (2008) demonstrated encapsulation and controlled release of hepatocyte growth factor (HGF) from such scaffolds. A mixture of poly(3-hydroxybutyrate- 332 M. Zilberman et al. co-3-hydroxyvalerate) (PHBV) and PLGA was chosen as the microsphere/polymer matrix. The PLGA/PHBV microspheres had a core/shell structure and the released HGF preserved its bioactivity for at least 40 days. Moreover, the three-dimensional microsphere scaffolds were more effective in maintaining the viability and phenotype of the primary hepatocytes. Wei et al. (2006) reported that microspheres can also be incorporated into scaffolds using the post-seeding method. PLGA microspheres consisting of platelet-derived growth factor-BB (PDGF-BB) were incorporated into a poly(L-lactic acid) (PLLA) nanofiber scaffold. The released PDGF-BB from the microspheres and from the microsphere-incorporated scaffolds remained bioactive. Furthermore, the burst effect was significantly lower in the microsphere-incorporated scaffolds than in the microspheres. In conclusion, the field of active scaffolds for tissue regeneration is relatively new and only a small number of such scaffolds have been developed to date. Three approaches to protein (growth factor) incorporation into bioresorbable scaffolds were presented in this section: adsorption onto the surface of the scaffold, freezedrying of inverted emulsions for incorporation of the bioactive agents into the scaffold during its formation process, and composite scaffold/microsphere structures. It is important to note that incorporation of protein in all three approaches is carried out such that it preserves the activity, i.e., without exposure to elevated temperatures or organic solvents. 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