Drug-Eluting Medical Implants

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 surround-

ing tissue. In this chapter we focus on three types of drug-eluting devices: drug-

eluting 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 antiproli-

ferative 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: [email protected]

M. Schafer-Korting (ed.), Drug Delivery,Handbook of Experimental Pharmacology 197,

DOI 10.1007/978-3-642-00477-3_11, # Springer-Verlag Berlin Heidelberg 2010

299

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 prepara-

tion 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 micro-

structure 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 devel-

oped 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 vas-

cular 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

300 M. Zilberman et al.

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 high-

risk 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 genera-

tion (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

Table 1 Clinical restenosis definitions and distinct measurements

Measurement Definition

Target lesion

revascularization (TLR)

The rate of reported re-intervention procedures inside the target

lesion

Target vessel

revascularization (TVR)

The rate of re-intervention procedures inside any lesion located in

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

events (MACEs)

Complications in cardiac trials such as death, Q-wave and non-Q-

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

Drug-Eluting Medical Implants 301

(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, anti-

inflammatory 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 antiprolifera-

tive 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 pharma-

ceutical 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 Bulles-

feld 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 thin-

ning, and thrombosis, due to delayed re-endothelialization (Grube and Bullesfeld

302 M. Zilberman et al.

2002). Dose control, delivery profile, and tissue pharmacokinetics are therefore

essential.

Several other APIs selected for localized release presented less favorable out-

comes. Antinomycin D is an antineoplastic drug that prevents cell division. Multi-

link 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

Drug-Eluting Medical Implants 303

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, respec-

tively. ACHIEVE (Guidant) demonstrated only minor improvement over bare-

metal 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-vinyl-

acetate 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 resteno-

sis 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

304 M. Zilberman et al.

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

Fig. 1 Selected Cypher pivotal studies: (a) angiographic in-segment restenosis, (b) target lesion

revascularization (BMS, bare metal stent; SES, sirolimus eluting stent)

Drug-Eluting Medical Implants 305

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 indica-

tions 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 effi-

cacy, when comparing Taxus and Cypher in the real world (Morice et al. 2006;

Roy et al. 2008).

Fig. 2 Selected Taxus pivotal studies: (a) angiographic in-stent restenosis (ISR), (b) target lesion

revascularization (BMS, bare metal stent; PES, paclitaxel eluting stent)

306 M. Zilberman et al.

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 cover-

age characterized by persistent fibrin deposition and delayed re-endothelialization

(Finn et al. 2007; Joner et al. 2006). The mechanisms by which DES induce non-

uniform 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

Fig. 3 Real world Cypher and Taxus studies, target lesion revascularization results (BMS, bare

metal stent; PES, paclitaxel eluting stent; SES, sirolimus eluting stent)

Drug-Eluting Medical Implants 307

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 com-

plexity 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 reac-

tion 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 Liberte (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 Liberte 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 (ABT-

578), 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

308 M. Zilberman et al.

muscle cell selectivity (Daemen and Serruys 2007). Xience V (Abbott Vascular) is

an everolimus-eluting stent embedded in a durable fluoropolymer coating a thin-

strut 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 zotar-

olimus 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�sirolimus-

eluting 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: pimecro-

limus 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 investiga-

tion. Furthermore, a new pro-healing endothelial progenitor cell capture stent

consists of antibodies attached to a stainless steel stent which specifically tar-

get endothelial progenitor cells in the vascular circulation. The HEALING-FIM

(Aoki et al. 2005b) trial demonstrated that the endothelial progenitor cell cap-

ture 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. Endea-

vorTM 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

Drug-Eluting Medical Implants 309

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 asso-

ciated 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 biodegrad-

able 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 bioab-

sorbable polymer that gradually releases the drug over 6–9 months. BioMatrix was

tested in the STEALTH-1 (Daemen and Serruys 2007) trial, and exhibited supe-

riority 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). Paclitaxel-

eluting InfinniumTM DES (Sahajanand Medical Technologies) is a slotted-tube

stainless steel stent with lower strut thickness coated with three layers of biode-

gradable polymer: poly(DL-lactic acid-co-glycolic acid) (PDLGA) 50/50, PDLGA-co-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 sug-

gested 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 biode-

gradable polymer layer deposited in laser-cut holes embedded in the stent platform.

Paclitaxel distribution and kinetics is obtained by using a poly-lactic-co-glycolicacid (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

310 M. Zilberman et al.

paclitaxel release over 30 days) achieved the best clinical and angiographic out-

comes (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 high-

dose slow-release formulation. Nonetheless, the COSTAR II study could not

conclude that the CoStar DES is non-inferior in clinical and angiographic perfor-

mance 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

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 – bare-

metal stent)

Drug-Eluting Medical Implants 311

inflammation (Farb et al. 1999). Permanent stents also pose a barrier for possible

future bypass surgery (Waksman 2006) and noninvasive imaging. These disadvan-

tages 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 complica-

tions of bare-metal stents and DES implantation by degrading into non-toxic sub-

stances 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, polycaprolac-

tone (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

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 meth-

acrylate) 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); phosphoryl-

choline coating

312 M. Zilberman et al.

Table

2SelectedCoated

Metal

DrugElutingStent(D

ES-IandDES-II)

DES/m

anufacturer

Drug

Stentmaterial

Platform

Matrixtype

Status

Cypher/Cordis(J&J)

Sirolimus

(Rapam

ycin)

316LStainless

steel

BX

Velocity

Durable

polymer

FDA,CEMark

TaxusLiberteTM/BostonScientific

Paclitaxel

316LStainless

steel

Express

2Durable

polymer

FDA,CEMark

Endeavor1

/Medtronic

Zotarolimus

Cobaltchromium

Driver

1Phosphorylcholine

FDA,CEMark

XIENCETMV/Abbottvascular

Everolimus

Cobaltchromium

Multi-Link

Vision

1Durable

polymer

FDA,CEMark

EndeavorResolute/M

edtronic

Zotarolimus

Cobaltchromium

Driver

1BioLinxTM

CEMark

PROMUSTM/BostonScientific

Aprivate-label

versionofAbbott’sXIENCETMVEverolimusElutingCoronaryStentSystem

FDA

Cypher

SelectT

MPlus/Cordis(J&J)

Sirolimus

(Rapam

ycin)

316LStainless

steel

Undisclosed

Durable

hydrophilic

coating

CEMark

Infinnium/SahajanandMedical

Technologies

Paclitaxel

316LStainless

steel

Millennium

Matrix1

Biodegradeable

polymers

CEMark

CoStarT

M/ConorMedsystem

sPaclitaxel

Cobaltchromium

CoStarT

MBiodegradable

PDLGAlayers

CEMark

BioMatrix1/Biosensors

International

Biolimus1

316LStainless

steel

S-StentTM

Biodegradable

PLA

CEMark

NOBORIT

MTERUMO

NOBORIT

Mismarketed

withlicensingagreem

entwithbiosensorunder

itstechnology.

CEMark

ACHIEVETM/Cook-G

uidant

Paclitaxel

316LStainless

steel

MULTILIN

K

PENTATM

Polymer-free

CEMark

VFlexPlusPTXTM/Cook

Paclitaxel

316LStainless

steel

V-FlexPlus

Polymer-free

CE]Mark

JANUS/SorinBiomedica

Tacrolimus

316LStainless

steel

Carbostent

Nanoporousceramic

coating

CEMark

GenousT

M/OrbusN

eich

Endothelial

progenitorcell

capture

316LStainless

steel

R-Stent

Polysaccharide,covalently

coupledto

stentsurface

CEMark

approved

in

Taiwan

Dexam

etTM/Abbottvascular

Dexam

ethasone

316LStainless

steel

Biodiv

YsioTM

Phosphorylcholine

CEMark

XTENT1/Biosensors

International

Biolimus1

Cobaltchromium

Custom

NX1

Biodegradable

PLA

Trial

Supralimus-Core

TM/Sahajanand

Medical

Technology

Sirolimus

316LStainless

steel

Millennium

Matrix1

Biodegradable

Polymers

Trial

(con

tinu

ed)

Drug-Eluting Medical Implants 313

Table

2(continued)

DES/m

anufacturer

Drug

Stentmaterial

Platform

Matrixtype

Status

Zomaxx/AbbottLabratories

Zotarolimus

Stainless

steel/

tantalum/stainless

steel

Tri-M

axPhosphorylcholine

Suspended

Batim

astetBiodivYsio

(BiocompatiblesandBritish

Biotech)

Batim

astet

316LStainless

steel

BiodivYsioTM

Phosphorylcholine

Suspended

QuaD

DS-Q

P2/Quanam

Medical

andBostonScientific

7-hexanoyltaxol

(QP2)

316LStainless

steel

QueST

Rigid

polymer/drugsleeve

cover

Suspended

AntinomycinDstent/Guidant

AntinomycinD

316LStainless

steel

MULTILIN

K

tetra-D

stent

Durable

polymer

(Ethylene

vinylacetate)

Suspended

Jomed

Flexmaster/Jomed

Tacrolimus

316LStainless

steel

Jomed Flexmaster

Nanoporousceramic

layer

Suspended

314 M. Zilberman et al.

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 low-

molecular-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 monofilament-

based 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) double-

helical 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 genein 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 dem-

onstrated an exciting possibility for restenosis prevention. The multiple lobe

PLLA fiber-based stent (Nguyen et al. 2004) was coated with drug-loaded micro-

spheres 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

Drug-Eluting Medical Implants 315

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 Bioabsorb-

able Therapeutics, Inc. (BTI): a sirolimus-eluting biodegradable stent composed of

salicylic acid (active ingredient in aspirin) chemically incorporated into polyanhy-

dride 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 freeze-

drying 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 domi-

nant 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 appar-

ent 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

316 M. Zilberman et al.

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

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

Drug-Eluting Medical Implants 317

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 drug-

eluting 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 drug-

eluting 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 LocalAntibiotic 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 complica-

tion 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

318 M. Zilberman et al.

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 dis-

charges provides conditions that are also favorable for bacterial growth. This has

given rise to a new generation of wound dressings with improved curative proper-

ties that provide a local antimicrobial effect by eluting various germicidal com-

pounds.

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 compli-

cations (Gold and Moellering 1996; Gransden 1997).

A local antibiotic release profile should therefore generally exhibit a consider-

able 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.

Drug-Eluting Medical Implants 319

2.2 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), mem-

branes (PolyMem Silver, Ferris), hydrocolloids (Urgotul SSD, Urgo), alginates

(Silvercel1, Johnson & Johnson), and hydrofibers (Aquacel, ConvaTec) are avail-

able. 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 nanocrystal-

line 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 Opra-

fol1 (Lohmann & Rauscher, Germany) are currently available (Jurgens et al.

2006). Biodegradation of the film occurs via hydrolysis of the copolymer into lactic

320 M. Zilberman et al.

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 (Jurgens

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., electrospin-

ning 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 recon-

struction 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

Drug-Eluting Medical Implants 321

wound dressing ingredient are mainly due to its rapid biodegradation by collage-

nase 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), Sulmycin1-

Implant (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 display-

ing considerable intrinsic antibacterial activity against a broad spectrum of bacteria

(Muzzarelli et al. 1990). Ignatova et al. (2006) reported the electrospinning of

322 M. Zilberman et al.

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 chit-

osan 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 AntibacterialDrugs for Wound Healing Applications

Drug-eluting fibers can be used for various biomedical applications. Few controlled-

release 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 proper-

ties 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 freeze-

drying 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 mafe-

nide 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

Drug-Eluting Medical Implants 323

(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, depend-

ing 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 demon-

strated 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

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)

324 M. Zilberman et al.

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 signifi-

cantly 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 effec-

tiveness 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 gentamicin-

eluting fibers with different release profiles (Fig. 11):

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)

Drug-Eluting Medical Implants 325

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)

326 M. Zilberman et al.

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.

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 surfactantr� 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)

Drug-Eluting Medical Implants 327

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, espe-

cially S. aureus and P. aeruginosa. The third strain, S. epidermidis, usually com-

prises 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 1�107 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

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)

328 M. Zilberman et al.

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.

Fig. 12 Number of colony forming units (CFU) versus time, when an initial bacterial concentra-

tion of 1�107 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)

Drug-Eluting Medical Implants 329

In conclusion, this section described wound dressings with emphasis on

approaches for local drug release for prevention of bacterial infections that accom-

pany 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 dres-

sings 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 interdisci-

plinary 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 administra-

tion 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 proteol-

ysis (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.

330 M. Zilberman et al.

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 tech-

nique 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 rhBMP-

2 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 rhBMP-

2 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 com-

posite 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-

Drug-Eluting Medical Implants 331

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 phe-

notype 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, freeze-

drying of inverted emulsions for incorporation of the bioactive agents into the

scaffold during its formation process, and composite scaffold/microsphere struc-

tures. 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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