DOI: 10.1002/adem.201180088 INVITED REVIEW Controlled Release …meitalz/Articles/D5.pdf · 2016....

INVITED

REVIE

W
DOI: 10.1002/adem.201180088
Controlled Release of AntiproliferativeDrugs From Polymeric Systems for StentApplications and Local Cancer Treatment
By Amir Kraitzer and Meital Zilberman*
Restenosis (re-narrowing of the blood vessel wall) and cancer are two different pathologies that havedrawn extensive research attention over the years. Antiproliferative drugs such as paclitaxel inhibitcell proliferation and are therefore effective in the treatment of cancer as well as neointimal hyperplasia,which is known to be the main cause of restenosis. Drug-eluting stents (DES) significantly reduce theincidence of in-stent restenosis (ISR), which was once considered a major adverse outcome ofpercutaneous coronary stent implantations. Localized release of antiproliferative drugs interfereswith the pathological proliferation of vascular smooth muscle cells (VSMC), which is the main cause ofISR. Conventional approaches to treating cancer are mainly surgical excision, irradiation, andchemotherapy. In cancer therapy, surgical treatment is usually performed on patients with a resectablecarcinoma. An integrated therapeutic approach, such as the addition of a delivery system loaded withan antiproliferative drug at the tumor resection site, is desirable. This will provide a high localconcentration of a drug, that is, detrimental to malignant cells which may have survived surgery, thuspreventing re-growth and metastasis of the tumor. The present review describes recent advances insystems for controlled release of antiproliferative agents. It describes basic concepts in drug deliverysystems and antiproliferative drugs and then focuses on both types of systems: stents with controlledrelease of antiproliferative agents, and drug-eluting particles and implants for local cancer treatment.The last part of this article is dedicated to our novel drug-eluting composite fiber structures, which canbe used as basic stent elements as well as for local cancer treatment.

1. An Overview of Drug Release, BiodegradablePolymers and Antiproliferative Agents

1.1. Controlled Drug Delivery

Controlled delivery of drugs via polymeric systems is a

common technique, where drug release is regulated either by

erosion of or diffusion through a polymeric matrix in a

pre-designedmanner. Systemic dose systems (Figure 1, dotted

[*] Dr. A. Kraitzer, Prof. M. ZilbermanFaculty of Engineering,Department of Biomedical Engineering,Tel-Aviv University, Tel-Aviv 69978, IsraelE-mail: [email protected]

B294 wileyonlinelibrary.com � 2012 WILEY-VCH Verlag GmbH & Co

plot) cause a rapid rise in the drug blood concentration,

which then exponentially decays as the drug is excreted from

the body. Additional doses are then required at certain time

points in order to maintain the drug level above the minimum

effective level. However, providing the body with multiple

doses can be harmful if the drug concentrations reach the toxic

level, above which the drug produces undesirable side effects.

The difference between the toxic and the effective level is

known as the toxic-therapeutic window, or therapeutic index.

The disadvantages of systemic drug release eventually led to

the development of controlled release systems. Therapeutic

agents incorporated into controlled release systemsmaintain a

desired blood plasma level within the therapeutic index for

long periods of time, as presented in Figure 1. Controlled

release systems are expected to release drugs in predictable

. KGaA, Weinheim ADVANCED ENGINEERING MATERIALS 2012, 14, No. 6

INVITED

REVIE

W

A. Kraitzer and M. Zilberman/Controlled Release of Antiproliferative Drugs

Fig. 1. Drug concentration in the blood plasma as a function of time; ( ) systemicrelease, ( ) controlled release.

kinetics with minimal environmental influence and patient

variability. A further improvement is a feedback-controlled

device that releases the appropriate amount of drug in

response to a therapeutic marker.[1–3]

Polymeric devices may regulate and control the release

rate, thus maintaining therapeutic levels of the drug. In

addition to the type of polymer and release mechanism, the

geometry of the device, such as a three-dimensional matrix,

film, fiber, injectable gel, or micro/nanoparticles, also affects

the release profile. In many of the controlled release

formulations, the release profile is characterized by a burst

effect, a large initial release, followed by a decrease in release

rates with time. One of the explanations proposed for this

phenomenon is that part of the drug is entrapped on the

surface area of the polymeric matrix, especially if the initial

drug loading is high. High burst release should be avoided

since it may lead to drug concentrations near or above the

toxic level, especially when using toxic antiproliferative drugs

such as paclitaxel.

Controlled release polymeric systems may be classified

according to the mechanism of controlled drug release. There

are three main mechanisms of controlled drug release:

diffusion, chemical, and swelling. These mechanisms may

occur in a given release system alone or in combination.[1,3,4]

1.1.1. Diffusion-Controlled Systems

In systemswhich are controlled by diffusion, themotive for

drug release is a concentration gradient. Monolithic (matrix)

and membrane-controlled (reservoir) devices are the two

fundamentally different devices in which the rate of drug

release is controlled by diffusion. In monolithic devices the

drug is either dissolved or dispersed in a polymer matrix.

Diffusion occurs when the drug passes from the polymer

matrix into the external environment. With this type of

system, the release rate normally decreases with time since the

active agent has a progressively longer distance to travel and

therefore requires a longer diffusion time to release. Conse-

quently, these systems cannot yield a constant release rate.

As opposed to monolithic systems, membrane-controlled

ADVANCED ENGINEERING MATERIALS 2012, 14, No. 6 � 2012 WILEY-VCH Verla

systems can produce a fairly constant drug delivery rate. Such

a system consists of a drug-containing core and a thin

membrane made of a material that controls the release rate.

There are usually deviations from zero-order release kinetics

in such systems, mainly due to the fact that when exposed to a

release medium, initial release is rapid because the agent

diffuses from the saturated membrane.[1,3,4] The actual

difference between monolithic and membrane-controlled

systems is the location of the drug, i.e., either in the core of

the device or dispersed in the entire device.

1.1.2. Chemically Controlled Systems

These biodegradable polymer-based systems may be

classified into pendant chain and erodible mechanisms. In

pendant chain systems, the drug molecules are covalently

attached to the backbone of a biodegradable polymer and are

released by hydrolysis of these bonds. In order to ensure that

polymer fragments are not released with the drug, it is crucial

that the bonds between polymer monomers are less reactive

than the bond which attaches the drug to the polymer.

Bioerodible systems consist of either biodegradable reservoir

systems or a biodegradable matrix drug dispersed system.

The release of the drug from biodegradable reservoir systems

is similar to that described in the previous section and differs

only in the fact that the membrane surrounding the drug core

is biodegradable. The release rate from these types of systems

can be controlled by changing the nature of the bioerodible

membrane. Drug release from biodegradable matrix systems

can be either erosion or diffusion-controlled, where the drug is

released as the polymer degrades.[1,3]

1.1.3. Swelling-Controlled Systems

Swelling-controlled systems absorb water or other body

fluids when placed in the body. The swelling increases the

aqueous solvent content within the formulation as well as the

polymer mesh size, enabling the drug to diffuse through

the swollen network into the external environment. Most

swelling-controlled systems are based on hydrogels, polymers

which swell when placed in water or other biological fluids,

while the drug may be located in an internal section of the

system (reservoir system) or spread within the entire volume

(matrix system). These hydrogels can absorb large amounts of

fluids, and are typically comprised of 60–90% fluid and

10–30% polymer. The release rate can be controlled by altering

the surrounding environmental parameters, such as pH,

temperature, and ionic strength.[1,3]

1.2. Biodegradable Polymers

Biodegradable polymers are polymers that undergo a

chemical process resulting in the cleavage of the covalent

bonds that make up the polymer chain, producing shorter

polymer chains (oligomers), and polymer repeating units

(monomers).[1,5,6] Biodegradable polymers are attractive for

implant applications that require temporary presence, are

excellent for local and controlled drug release, and are free of

g GmbH & Co. KGaA, Weinheim http://www.aem-journal.com B295

INVITED

REVIE

W


Fig. 2. Monomer unit of (a) PGA and (b) PLA.

long-term biocompatibility issues. Biodegradable polymers

can be classified into natural and synthetic polymers.

Synthetic polymers may be tailored to the required mechan-

ical properties and degradation kinetics suitable for the

application.[7] Furthermore, since most synthetic polymers

undergo degradation by hydrolysis, the degradation rate

between individuals is almost identical, since water avail-

ability in biological tissues is constant and differs only slightly

between people.[1,7] Polyesters have been the most attractive

among the families of synthetic polymers. Their degradation

occurs at their ester bonds by hydrolysis and their degradation

products are resorbed via metabolic pathways.[8]

Poly(a-hydroxy acids) are the most widely investigated

and most commonly used synthetic biodegradable polymers

of the polyester family. They are also considered to be safe,

non-toxic, and biocompatible materials, since their degrada-

tion byproducts are eliminated from the body through the

Krebs cycle.[1,6,7] Poly(a-hydroxy acids) include polymers

such as poly(glycolic acid) (PGA), poly(lactic acid) (PLA), and

a range of their copolymers (PLGA) that have been approved

by the FDA. These materials have long history of use as

synthetic biodegradable materials in a number of clinical

applications such as resorbable sutures, plates, and fixtures for

fracture fixation devices and scaffolds for tissue engineering.[7–9]

Initial degradation of poly(a-hydroxy acids) occurs by

hydrolysis of the a carbon (the carbon belonging to the ester

bond), which comprises the backbone of these polymers,

resulting in many hydroxyl groups. This process occurs until

the molecular weight of the device is less than 5000Da, at

which point byproducts start leaving the device and it begins

to lose mass (erosion). At the final stage of polymer

degradation, the acidic degradation products are absorbed

by inflammatory cells (such as macrophages, lymphocytes,

and neutrophils) and are eliminated from the body through

the Krebs cycle as CO2 and H2O.

1.2.1. Poly(glycolic acid) (PGA)

PGA is the simplest linear aliphatic polyester. It degrades

within 6–12 months. It is highly crystalline (46–50%) and

therefore has a highmelting point (225 8C) and is not soluble in

most organic solvents. Due to its relatively hydrophilic

nature, PGA tends to lose its mechanical strength (in vivo)

rapidly, typically over a period of 2–4 weeks after implanta-

tion.[1,5,7,8,10]

1.2.2. Poly(lactic acid) (PLA)

PLA is a chiral molecule and therefore exists in two

steroisomeric forms: D-PLA and L-PLA and the racemic form

DL-PLA. Poly L-lactic acid (PLLA) degrades over a period of

more than 24 months (in vivo), and poly DL-lactic acid

(PDLLA) degrades within 12–16 months (in vivo). PDLLA is

amorphous, thus allowing homogeneous dispersion of the

active species within the carrier. It is usually used for drug

delivery, whereas the semicrystalline PLLA is preferred in

applications which require high mechanical strength and

toughness.[1,5–8,10]

B296 http://www.aem-journal.com � 2012 WILEY-VCH Verlag GmbH & C

Random copolymers of glycolic acid (GA) with the

morehydrophobic lactic acid (LA), known as poly(lactic-

co-glycolic acid) (PLGA) are presented at Figure 2.

PLGAs were investigated extensively in order to adapt the

material properties of PGA to awider range of applications. In

general, the hydrophobic nature of LA reduces the rate of

backbone hydrolysis compared to the homopolymer PGA,

since it limits water uptake. It is noteworthy that there is no

linear relationship between the ratio of GA and LA and the

physicomechanical properties of the corresponding copoly-

mers. PGA is highly crystalline, whereas crystallinity is lost in

copolymers of glycolic and LAs, resulting in an amorphous

PLGA. Thus, a copolymer of 50% GA and 50% LA degrades

more rapidly than either PGA or PLA. When comparing

different compositions of PLGA, lactide-rich PLGA copoly-

mers are more hydrophobic, absorb less water, and subse-

quently degrade more slowly. For example 50:50 PLGA

degrades within 2 months while 85:15 PLGA degrades within

5 months.[1,5,6,8,10] A copolymer of GA and DL-LA, known as

PDLGA, is used in the study presented in Section 4.

Recently, a water-soluble and biodegradable high mole-

cular weightN-(2-hydroxypropyl)methacrylamides (HPMAs)

were synthesized via RAFT polymerization and click

chemistry.[11,12] Former HPMAs that were used for drug

release lack biodegradability qualities hence, the molecular

weight of the polymer carrier was low, below the renal

threshold (�50 kDa). However, for cancer treatment, higher

molecular weights were found to increase drug accumulation

at target. The degradation of the multiblock polyHPMAs in

the presence of papain or lysosomal cathepsin B validated the

preparation strategy. This new approach provides a platform

for the design and preparation of biodegradable polyHPMA-

based drug carriers.

o. KGaA, Weinheim ADVANCED ENGINEERING MATERIALS 2012, 14, No. 6

INVITED

REVIE

W


Fig. 3. The chemical structure of (a) paclitaxel, (b) sirolimus, and (c) FTS.

The process of degradation describes the chain scission

during which polymer chains are cleaved to form oligomers

and finally monomers. A polymer chain can be chemically

degraded either by passive hydrolysis or by the active

enzymatic method. Hydrolytic degradation is caused by the

reaction of water with labile bonds, typically ester bonds, in

the polymer chain. Hydrophilic polymers take up large

quantities of water and degrade faster than hydrophobic

polymers.[5,13] The copolymer composition can alter the

degradation rate. For example, a PLGA copolymer with a

high percentage of GAwill degrade faster than one with a low

percentage of GA.

Polymeric erosion is the physical disintegration of a

polymer matrix following degradation, resulting in mass

loss. Post-chain scission, low molecular weight chains, i.e.,

oligomers and monomers, diffuse out to the surroundings.

There are two distinct erosion mechanisms described in the

literature: ‘‘bulk (or homogeneous) erosion’’ and ‘‘surface (or

heterogeneous) erosion’’ that differ in the rate of water

absorption into the polymer and in the degradation rate of the

polymer backbone.[13,14] The rate at which water imbibes into

poly a-hydroxy acids ismuch higher than the rate at which the

ester bonds are hydrolytically cleaved. Degradation thus

occurs through bulk erosion rather than throughout surface

erosion.

1.3. Antiproliferative Drugs

Antiproliferative drugs were originally developed and

used in cancer treatment since they directly inhibit cell

proliferation and migration. Additional studies suggested

their use in restenosis treatment, exploiting their ability to

reduce vascular smooth muscle cells (VSMC) growth. Indeed,

stents coated with such antiproliferative drugs were shown to

reduce neointimal growth in both animal and clinical studies.

1.3.1. Paclitaxel

Paclitaxel is a potent cell proliferation inhibitor and is

known to be very effective in the treatment of cancer as well as

neointimal hyperplasia, which is currently known as the main

cause of restenosis. Paclitaxel’s anti-tumor activity was

detected in 1967 by the US National Cancer Institute. It was

approved by the FDA for ovarian cancer in 1992, after which it

became a standard medication in oncology. It inhibits mitosis

Table 1. Chemical properties of antiproliferative drugs.

Properties Paclitaxel

Chemical formula C47H51NO14

Molecular weight 853.9 [g �mol�1]

Melting temperature 213–216 8CSolvents Dimethyl sulfoxide,

methanol, ethanol,

and acetonitrile, chloroform,

methylene chloride

Water solubility 5mg �mL�1[26]


in dividing cells by binding to microtubules and causes the

formation of extremely stable and non-functional microtu-

bules, thus preventing transition from theG2 to theMphase of

the mitotic cycle.[15] Paclitaxel’s structural formula is pre-

sented in Figure 3(a) and its chemical properties are presented

in Table 1.

The use of paclitaxel in DES inhibited VSMC proliferation,

migration, and secretion of extracellular matrix. Slow-release

paclitaxel applied perivascularly totally inhibits intimal

Sirolimus FTS

C47H51NO14 C22H30O2S

914.184 [g �mol�1] 358.5 [g �mol�1]

173–189 8C N/A

Dimethyl sulfoxide,

methanol

Dimethyl sulfoxide,

methanol, ethanol,

and acetonitrile, chloroform,

methylene chloride

2.6mg �mL�1[27] N/A


INVITED

REVIE

W


hyperplasia and prevents luminal narrowing following

balloon angioplasty and stent placement. The drug interacts

with arterial tissue elements as it moves under the forces of

diffusion and convection, and can establish substantial

partitioning and spatial gradients across the tissue. The

lipophilic nature of paclitaxel favors partitioning into

the organic phase of the encapsulating polymer. Studies

indicate the need for a controlled drug release of paclitaxel

due to its narrow toxic-therapeutic window and high

hydrophobic character.[16] Recent DES studies reported

delayed endothelialization within 90 days after treatment.

Thus, antiplatelet therapy with aspirin or Ticlopidine/

Clopidogrel is encouraged 1 year post-implantation of

DES.[17,18] Finally, the molecule is susceptible to solvolysis

of its ester bond, leading to loss of its cytotoxic activity with

maximum stability in the range of pH 3–5 at 37 8C.[19]

1.3.2. Sirolimus

Sirolimus was approved by the FDA in 1999 as Rapamune

(Wyeth, NJ), an immunosuppressive drug against transplant

rejection. Sirolimus is an immunosuppressive macrolide that

easily crosses the cell membrane and binds to an intracellular

protein (FKBP12) which activates the mTOR protein.[20]

Sirolimus inhibits the cell cycle in the transition from G1 to

S, blocking cell proliferation without inducing cell death,[21]

thus leading to cell reversion into a quiescent state.[22]

It has been found to have potent cell cycle inhibitory

activity and therefore inhibits SMC proliferation. Sirolimus

is non-specific. Sirolimus-coated DES may therefore

cause endothelial dysfunction[18] and a subsequent loss of

re-endothelialization[17] of the inner side of the metal stent.

Sirolimus’ structural formula is presented in Figure 3(b) and

its chemical properties are presented in Table 1.

1.3.3. Farnesylthiosalicylate (FTS)

The limitations of traditional modes of therapy with

anticancer drugs, namely non-specific distribution, systemic

toxicity, and rapid development of resistance, are forcing a

look at new modalities in cancer therapy and identification

of molecular targets which make the tumor cell highly

susceptible to such therapies. Target-directed cancer therapy

is the most prominent direction of cancer research today. The

best example of success in this field is the drug Gleevec, which

blocks the chimeric bcr-abl gene (Philadelphia chromosome)

product that causes chronic myeloid leukemia (CML).[23]

Gleevec is currently the standard care for CML. Many other

drugs that act on specific oncogenic proteins are now at

various stages of development. One of the most prominent

oncoproteins in human cancer is Ras. In its active GTP-bound

form, Ras promotes enhanced cell proliferation, tumor cell

resistance to drug-induced cell death, enhanced migration,

and invasion. Ras is therefore considered an important target

for cancer therapy as well as for therapy of other proliferation

diseases, including restenosis.

FTS (Salirasib) is a new rather specific non-toxic drug

which acts as a Ras antagonist, which was developed by Prof.


Yoel Kloog, Tel-Aviv University.[24,25] Its mechanism of action

is likely to be associated primarily with the dislodgment of the

mature protein from membrane domains that interact with

Ras, and with the subsequent accelerated degradation of the

dislodged Ras proteins. The apparent selectivity of FTS

toward active guanosine triphosphate (GTP-bound) Ras and

lack of toxic or adverse side effects in animal models[24] made

it a good candidate for cancer treatment. FTSwas found to be a

potent inhibitor of intimal thickening in the rat carotid artery

injury model which serves as a model for restenosis, while it

does not interfere with endothelial proliferation.[24] Thus, the

incorporation of FTS in a stent coating may overcome

incomplete healing and lack of endothelial coverage asso-

ciated with current DES. FTS’s structural formula is presented

in Figure 3(c) and its chemical properties are presented in

Table 1.

2. Drug-Eluting Stents

Restenosis is the re-narrowing of a blood vessel causing a

reduction in the lumen size, consequently restricting blood

flow after an intravascular procedure. Restenosis, which once

occurred in 20–55% of patients post-stent placement, has been

reduced to less than 5% in the DES era.[28–32] 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 is caused by balloon

expansion and the presence of a permanent stent implant.

ISR is mainly characterized by intimal hyperplasia, i.e., an

abnormal increase in the VSMC and vessel remodeling[33] that

causes a reduction in the lumen size. There are four phases of

ISR that generally occur post-stent implantation: (a) imme-

diately after stent placement, the injury causes platelet

aggregation and activation; (b) A variety of white cells gather

at the injury site over the next few days to weeks; (c) SMCs

migrate and proliferate to form the neointima and this process

decays after a month from the procedure; and (d) late

remodeling begins at about the third week. The process of ISR

peaks at about the third month and reaches a plateau at about

6 months after the procedure.[34]

The occurrence of late stent thrombosis (LST; > 30 days) in

DES is higher than with bare-metal stents (BMS).[35] DES’

occlusion is still low (0.5–3.1%) but unpredictable, and it often

involves fatal myocardial infarction occurring in up to 65% of

patients of thrombosis cases.[36] Angioscopic assessment in

humans 3–6months after stent deployment showed BMSwere

completely endothelialized, whereas 87% of DES were not,

and in 50% of the DES, thrombi were visible.[37] Newer DES

are now being designed to provide better stent deployment,

safety, and efficacy.

2.1. Drug-Eluting Metal-Coated Stents

The introduction of DES represents a breakthrough in the

treatment of coronary artery disease owing to their ability to

reduce the incidence of ISR to less than 5%.[31,32] Five FDA


INVITED

REVIE

W


Fig. 4. Plotted in vitro release data for sirolimus released from CypherTM, with SR andFR formulations.[22]

approved DES are currently available‘‘: CypherTM (Cordis

J&J, Sirolimus-eluting stent), TaxusTM Express2 (Boston

Scientific, Paclitaxel-eluting stent), TaxusTM Liberte (Boston

Scientific, paclitaxel-eluting stents), Endeavor (Medtronic,

Zotarolimus-eluting stent), and Xience V (Abbott Vascular,

Everolimus-eluting stent). Cordis Corporation recently

announced suspending the production line of the Cypher1

sirolimus-eluting stent by the end of 2011, and the abandon-

ment of any future plans to develop the Nevo stent: the first

sirolimus-eluting stent with a biodegradable polymer

scheduled for testing in the USA.[38] Their successors were

designed in light of past safety and efficacy concerns offering

an enhanced platform, release matrix, and more targeted

antiproliferative agents better deliverability, higher flexibil-

ity,[39] as well as drug release homogeneity and a low strut

profile.[40] DES consists of three components that can radically

affect their safety and efficacy: the bioactive agent, the stent

platform, and the controlled drug-release mechanism.

The clinical superiority of Cypher and Taxus over BMS has

been well documented over the last 5 years in randomized

trials.[41–43] However, continued neointimal formation over

time is often observed when DES is used, as opposed to BMS

in which neointimal formation peaks at about 6 months.

Long-term DES studies in broader populations, and in more

complex lesions,[44] presented less favorable outcomes[22] and

led to a surge of manuscripts that tempered enthusiasm

toward DES.[45] Cypher and Taxus were associated with an

increased rate of LST, a low frequency event with serious

life-threatening consequences,[46] and hypersensitivity reac-

tions on a smaller scale.[47] DES implantations were associated

with increased myocardial death rates at 6–18 months

post-implantation compared with BMS, particularly after

discontinuation of anti-platelet therapy.[36] Finn et al.[48] found

a correlation between stent struts that lack neointimal

coverage and the number of struts surrounded by

platelet-rich thrombi and determined that lack of enthothe-

lialization is currently the best predictor for LST. Impaired

re-enthothelialization is commonly caused by the antiproli-

ferative drug. This means that the drug’s inhibitory effect

abolishes the physiological vessel wall healing, leaving the

struts in direct contact with flowing blood and blood elements.

Hypersensitivity may be caused by the polymeric constituents

of the coatings.[47]

2.2. Current DES Coatings

The stent coating should have good mechanical properties

and should not elicit a negative tissue reaction. The stent

coating should hold high flexibility and long-lasting adher-

ence to the stent surface,[44] especially when the stent is

expanded to the required size during surgery. This expansion

can seriously affect the release of the drug from the polymer,

or worse, embolize the polymer.[49] Current stable polymer

coatings containing the drug may trigger local coronary

inflammation due to hypersensitivity reactions. For example,

the material from which the Cypher coating is constructed,


n-butyl methacrylate, was found to induce hypersensitivity in

rabbits[50] and chronic eosinophilic infiltration of the arterial

wall.[47] Endeavor is a zotarolimus-eluting stent on a thin-strut

cobalt–chromium platform coated with a non-degradable

phosphorylcholine coating.[51] This coatingmimics the natural

component of the cell membrane and as such is a good

example of an highly biocompatible DES coating.

Drug uptake into the vessel wall occurs by passive

diffusion and convection and is facilitated by the hydrophobic

nature of these antiproliferative drugs that establish

substantial partitioning and spatial gradients across the

tissue.[52,53] Localized release of antiproliferative drugs

interferes with the pathological proliferation of vascular

SMC, which is themain cause of ISR.[54] The first generation of

DES offered limited control over the drug release period,

drug load, and homogenous release in cases of complex

anatomical circumstances.[44,55] Cypher contains approxi-

mately 70–300mg sirolimus (140mg �mm�2), where 80% of

the drug is releasedwithin 28 days after stent implantation.[22]

Cypher’s release profile is obtained by coating a drug-free top

layer onto the drug reservoir layer. Without the top coating,

the entire amount of sirolimus is released within 15 days

compared to the 90 days required for complete release from

Cypher (Figure 4).

Taxus contains approximately 50–200mg (1mg �mm�2)

paclitaxel, where �2mg are released within 15 days[22] and

92.5% remain in the matrix for a long period.[56] Drug release

kinetics of these DES is determined by the drug/polymer ratio

and/or coating thickness[31] and is therefore far from the

optimum in terms of safety and efficacy. During trials,

three different release profiles have been reported for Taxus

(Figure 5), dictated by the paclitaxel concentration in the

polymer (8.8, 25, and 35%w/w). The hydrophobic nature of

the released antiproliferative drugs and the non-degradable

nature of the coatingmatrix, together with the low thickness of

current DES coating, suggest relatively poor control over the

drug release profile.[55]


INVITED

REVIE

W


Fig. 5. Plotted release data for paclitaxel from a coated TAXUSTM stent for SR,moderate release (MR), and FR.[22]

The cumulative DES knowledge emphasizes the impor-

tance of controlling drug release profiles. Studies[16,57] indicate

the need for controlled release of paclitaxel, due to the narrow

toxic-therapeutic window and high hydrophobic nature of

this compound. High paclitaxel dosages may lead to an

inflammatory vessel response, medial thinning, and throm-

bosis, due to delayed re-endothelialization.[57] A paclitaxel-

loaded stent study[56] reported that release durations shorter

than 10 days resulted in no improvement over a BMS, while

longer durations (over 30 days) presented superior restenosis

inhibition. Evaluated animal models suggested that sirolimu-

s-eluting stents reduced neointimal hyperplasia effectively in

30 days, but this effect disappeared at 90 days presumably due

to insufficient arterial drug level at 90 days.[58] The current

state of knowledge concerning optimal DES designs indicates

an interplay between drug selection and drug release

mechanisms, which determine the safety and optimizes the

local therapeutic benefit.[59,60] This could be summarized in

two statements: First, the stent should release a sufficient

amount of drug with appropriate kinetics, that is, maintained

for several weeks after the procedure in order to eventually

eliminate ISR. Second, the release profile should allow

confluent endothelial coverage that will suppress thrombosis.

Newer ‘‘-limus’’ drugs released from non-biodegradable

coatings demonstrated improved outcomes compared to

sirolumus and paclitaxel eluting stents. Xience V is an

everolimus-eluting stent embedded in a durable poly

vinylidene fluoride co-hexafluoropropylene coating a thin-

strut cobalt–chromium platform.[61] It proved rapid endothe-

lialization[62] while it releases about 80% of the drug within

30 days after implantation.[63] The COMPARE trial demon-

strated the improved safety and efficacy of the Everolimus

eluting stent compared to the second-generation paclitaxel

eluting stent in unselected patients.[64] Endeavor is a

zotarolimus-eluting stent on a thin-strut cobalt–chromium

platform coated with a non-degradable phosphorylcholine


coating[51] that releases zotarolimus in about 2 days.[39] The

ENDEAVOR IV trial demonstrated that this zotarolimus

eluting stent had similar effects in safety and efficacy to

paclitaxel eluting stent in both simple and medium complex-

ity single de novo coronary lesions with 12 months

follow-up.[65] From 3-year follow-up results of this trial, it

was indicated that it had similar anti-restenosis efficacy to

paclitaxel eluting stent, but the clinical safety had been

improved, for periprocedural and remote MIs had been

significantly reduced due to fewer incidents of very late

(>1 year) stent thrombosis.[66] Endeavor TM Resolute (Med-

tronic) was designed to release zotarolimus from a novel

BioLinxTM copolymer optimal for extended drug release.[67]

BioLinxTM is a unique blend of three different polymers: a

hydrophobic polymer for delayed drug release, a lipophilic

polymer for enhanced biocompatibility, and hydrophilic

polymer for release burst. Eighty five percent of its

zotarolimus content is release during the first 60 days and

the remainder in 180 days in vivo.[67]

More recent DES developments use biodegradable

matrices or do not use a polymer coating at all. Biodegradable

polymer-coated metal stents were first introduced in an

attempt to overcome the late risks associated with durable

polymers. A Biolimus-eluting BioMatrix stent (Biosensors

International) is coatedwith a poly-LA bioabsorbable polymer

that gradually releases drug over 6–9 months and exhibited

superiority over BMS.[39] Paclitaxel-eluting InfinniumTM DES

(Sahajanand Medical Technologies) is coated with three

layers of poly(DL-lactic acid-co-glycolic acid) (PDLGA) 50:50,

PDLGA-co-poly(capro lactone) (PDLGPCL) 75:25, and poly-

vinyl pyrrolidone[68] was found safe and effective.[39] Conor

Co-star DES[69] allows programmable and controlled pacli-

taxel release using different layers of poly(lactic-co-glycolic

acid) (PLGA) in different co-polymer ratios deposited in

laser-cut holes embedded in the stent’s struts. The PISCES

study proved that the Conor Medstent was safe and that the

duration of release had a greater impact on the inhibition of

in-stent neointimal hyperplasia than the dose.[56] More recent

studies[70,71] demonstrated that changing the hydrophobic/

hydrophilic ratio in the polymer matrices affected release

kinetics from the stents. Recently, a new pro-healing approach

involving endothelial progenitor cell capture (EPC) stent was

developed. The Genous EPC stent (Orbus Neich, Fort

Lauderdale, Florida) consists of antibodies attached to a

stainless steel stent which specifically target EPC in the

vascular circulation.[72] The EPC capture stent appears

effective in patients.[73,74]

2.3. Completely Biodegradable Drug-Eluting Stents

Completely biodegradable stents are now in an advanced

stage of research and development, and are considered the

next generation of DES.[75] Metal stents have thrombogenic

properties,[76] and may therefore cause permanent physical

irritation, with the risk of long-term endothelial dysfunction

or inflammation.[77] ISR commonly occurs within 3–6 months


INVITED

REVIE

W


after coronary intervention, and rarely thereafter. The clinical

need for stents is therefore limited after this period.[78]

Completely biodegradable stents may eliminate early and late

complications of BMS or DES placement by degrading into

non-toxic substances after maintaining luminal integrity.

Consequently, biodegradable stents may serve only during

the period of high risk restenosis,[79] while enabling safe

discontinuation of dual-antiplatelet therapy within several

months post-implantation.[75] Finally, biodegradable stents

have a higher capacity for drug incorporation, allowing

complex release kinetics by altering the biodegradation profile

of the polymer.[79] 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. It is also challenging to effectively incorporate

the drug during fabrication without damaging its activity.

Tamai et al.[78] were the first to study completely

bioresorbable stents in human trials. The Igaki-Tamai stent

is a PLLA zigzag helical coil design. It presented intimal

hyperplasia in rates comparable to BMS in 6 months, and 18%

trans-vessel revascularization (TVR) in a 4-year follow-up.[80]

Further designs exhibited high rates of inflammation due to

increased rates of polymer degradation and subsequent high

concentration of acidic degradation products. Lincoff et al.[81]

demonstrated that high molecular weight PLLA reduced

inflammatory reactions compared to low molecular weight

PLLA. Early drug-eluting biodegradable stent designs were

based on fiber or film structures. Yamawaki et al.[82]

incorporated Tranilast (ST638) into the Igaki-Tamai stent,

which presented less neointimal formation without signifi-

cantly reducing its mechanical properties. Uurto et al.[83]

presented acceptable results in a porcine model of a

monofilament-based stent made of a polymer consisting of

96% L-LA and 4% D-LA coated with a 50:50 ratio of two

bioactive agents: dexamethasone and simvastatin. Vogt

et al.[84] reported a paclitaxel-loaded PDLLA double-helical

stent exhibiting sufficient mechanical stability with a very

slow paclitaxel release pattern in a porcine model. Their

2-month evaluation demonstrated effective proliferation

inhibition, but also local inflammatory effects due to

polylactide resorption. Alexis et al.[27] incorporated paclitaxel

and rapamycin into PDLLA and poly DL-lactic-co-glycolic

(PDLGA) non-expandable helical stents prepared from film

strips exhibiting homogenous burst-free drug release. A

multiple lobe PLLA fiber-based stent[85] was coated with

drug-loaded microspheres in order to combine good mechan-

ical properties with the desired drug-release profile.[86] 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 content during 28 days and has a collapse pressure

similar to a stainless steel stent.[87] This stent was absorbed

after 2 years while vasomotion was restored and restenosis as

well as late thrombosis were prevented demonstrating clinical

safety and efficacy of the stent in simple lesions.[87,88] The BVS


stent has received a CE Mark approval in 2011, thereby

becoming the first approved drug-eluting bioresorbable

vascular stent. Preliminary animal studies using the IDEALTM

stent (Bioabsorbable Therapeutics Inc., Menlo Park, CA, USA)

were very encouraging.[87] This stent incorporates salicylic

acid into the backbone of a polyanhydride ester polymeric

stent and releases its entire sirolimus content in 30 days.[89]

The REVA stent is a paclitaxel-incorporated PLA-based

design (REVA medical, San Diego, Ca, USA) utilizing a

‘‘slide & lock’’ design rather than the usual material

deformation for deployment;[90] it releases 50% of the drug

in 10 days and 90% by 3 months.[75]

3. Local Treatment of Cancer

Conventional approaches to treating cancer are mainly

surgical excision, irradiation, and chemotherapy. In cancer

therapy, surgical treatment is usually performed on patients

with resectable carcinoma. However, treatment failure due to

local recurrence of primary tumors or metastatic spread often

occurs during management.[91] An integrated therapeutic

approach such as a delivery system loaded with an

antiproliferative drug at the tumor resection site is therefore

desirable. This will provide a high local concentration of a

drug detrimental to malignant cells which may have survived

surgery, thus preventing re-growth and metastasis of the

tumor. The current treatment of regional chemotherapy

through localized antiproliferative drug delivery is based

on the premise that anticancer agents display a steep

dose–response for both therapeutic effect and toxicity. The

narrow toxic-therapeutic window of the antiproliferative

drugs causes side effects and hypersensitivity reactions

during therapy. A local drug release technology that may

overcome the disadvantages of current systemic chemother-

apy treatments while attaining adequate drug levels at the

tumor site is of primary importance since inadequate tumor

cell drug-burden will lead to low cell kill and to the potential

for early development of resistance to the drug.[91]

Metastatic cancer is a clinical description for the spread of

cancer cells from the primary tumor site to distant organs,

establishing secondary tumor sites.[92] Detachment of cancer

cells from the primary tumor site and circulation in the

bloodstream allows the cells to arrest in organs such as the

lungs, liver, lymph nodes, skin, kidneys brain, colon, and

bones, where they can proliferate. Despite significant

increases in the understanding of metastatic cancer pathogen-

esis, early diagnosis, surgical methods, and irradiation

treatment, most cancer deaths are due to incurable metastases.

Reasons for this include resistance to treatments, difficulty in

accessing the tumor sites and removing all cancer cells during

surgery. Improving therapy of metastatic cancer is therefore

still a challenge, even thoughmultiple therapeutic approaches

are approved or in clinical development.[92]

Drug delivery systems explored so far for localized

paclitaxel delivery in cancer treatment include micro-

spheres,[93,94] surgical pastes,[95] and implants.[96] The major


INVITED

REVIE

W


limitations of these implants are attaining the required

amount of drug for a given amount of time and distributing

the antiproliferative drug, such as paclitaxel. The narrow

toxic-therapeutic window of these drugs causes side effects

and hypersensitivity reactions during therapy. The effective-

ness of the drug delivered by polymers depends on whether

drug molecules can be transported a sufficient distance from

the implantation site.[96] The high concentrations in the

vicinity of the implant and low concentrations maintained at

distant locations for a prolonged period after implantation

suggest that this deliverymodality may be effective in treating

multifocal tumors that recur at sites distant from the primary

tumor.

3.1. Glioblastoma Multiforme Treatment

Gliomas are a diverse set of primary brain tumors, which

are derived from normal glia cells that support the functions

of neurons in the brain.[97] About 30 000 patients in the United

States are diagnosed with a glioma each year.[98] Glioblastoma

multiforme is the most common and most aggressive primary

brain tumor, with amedian survival time fromdiagnosis of up

to one year.[99] Malignant gliomas are known to recur within a

few cm from the tumor excision site.[100] One of the problems

in treating glioblastoma is getting the drugs through the

blood–brain barrier (BBB). The physicochemical character-

istics of drugs largely determine the passive transport of drugs

across the BBB, such as hydrophilic paracellular–transcellular

transport, which is restricted by the tight junctions of the BBB

endothelial cells. This paracellular permeability is further

dependent on the charge of the molecules and the possibility

of forming hydrogen bonds.[101] As a general rule, drugs that

cross the BBB and produce appreciable brain drug levels are

either hydrophilic with MW� 160 or hydrophobic MW� 400.

One approach to overcome the BBB could be targeted drug

delivery to a particular site near the tumor.[91]

One of the first intracerebral delivery systems studied

for glioblastoma treatment were polymeric discs made

of poly(bis (p-carboxy phenoxy) propane-sebacic acid)

(PCPP-SA, 20:80) containing 20%w/w drug loading of

[1,3-bis-(2-chloroethyl)-1-nitrosourea] (BCNU), also known

as carmustine, which were surgically implanted next to the

tumor. This type of polymer is called polyanhydride and is a

surface-eroding bioresorbable polymer capable of releasing a

constant amount of drug per unit time. Paclitaxel incorporated

in PCPP-SA (20:80) polymer discs was evaluated in a rat

model of malignant glioma 5 days after tumor implantation.

The paclitaxel-loaded polymers doubled (38 days, 40%

paclitaxel loading, p< 0.02) to tripled (61.5 days, 20%

paclitaxel loading, p< 0.001) the median survival of tumor-

bearing rats relative to control rats (19.5 days).[102] This

approach was evaluated in a phase I-II study in 21 patients

with recurrent malignant glioma who were treated with

carmustine delivered from PCPP-SA (20:80) wafers and has

been shown to have promising initial activity and limited

toxicity.[103] This study eventually led to FDA approval of the


Gliadel1 wafer as the only interstitial chemotherapy treat-

ment currently approved for malignant glioma. These

PCPP-SA wafers containing 3.85% carmustine are placed in

the resection cavity during surgery.Malignant glioma patients

treated with Gliadel1 wafers at the time of initial surgery in

combination with radiation therapy demonstrated a survival

advantage of two and three years at follow-up compared

with placebo. The median survival of patients treated with

carmustine wafers was 13.8 months versus 11.6 months in

placebo-treated patients (p¼ 0.017).[104] The survival advan-

tage is derived without additional systemic side effects.[105]

However, this relatively minor achievement is due to the

resistance of many brain tumors to carmustine,[106] as well as

the low stability of the drug and its tendency to ionize at

physiological pH.[96]

3.2. Drug-Eluting Particles

The ability of nanoparticles to specifically target tumors

along with the controlled delivery of a therapeutic payload

provides powerful new ways to treat cancer which are only

starting to be realized.[92] The small size allows nanocarriers to

overcome biological barriers and achieve cellular uptake.

Polymer–drug conjugates, generally below 20nm, are among

the most extensively investigated types of nanocarriers and

are currently in clinical trials as advanced as phase III.

Polymer–drug conjugates are formed through side-chain

grafting of drugs to polymer chains, allowing them to deliver

high doses of chemotherapeutic drugs. These technologies

include polymeric nanoparticles, dendrimers, nanoshells,

liposomes, inorganic/metallic nanoparticles, hybrid nanopar-

ticles, micelles, and magnetic and bacterial nanoparticles.[92]

Degradation and drug release kinetics can be precisely

controlled by the physicochemical properties of the polymer,

such as molecular weight, dispersity index, hydrophobicity,

and crystallinity. The disadvantages of nanoparticles include

tuning the drug release rate, since small changes in the

polymer–drug conjugation may significantly modify the

pharmacokinetic parameters and tissue biodistribution.

Non-targeted nanoparticles circulating in the blood have

been shown to significantly improve drug bioavailability and

accumulation in tumors through the enhanced permeability

and retention effect.[92] This effect allows the passive targeting

of nanoparticles to tumors due to pathological abnormalities

in the tumor vasculature. Interendothelial gap defects increase

vascular permeability in tumors, allowing extravasation of

nanoparticles up to 400 nm. Accumulation of nanoparticles is

further enhanced due to poor lymphatic drainage in tumors.

The local release of anti-cancer drugs from nanocarriers in the

extravascular space results in an increased intra-tumoral drug

concentration. In general, hydrophobic drugs released extra-

cellularly will diffuse and be taken up by cancer cells, leading

to enhanced tumor cytotoxicity.

Several paclitaxel-loaded bioresorbable structures, such as

micro and nanospheres, thin films and polymeric micelles,

were studied extensively. These systems have been found to


INVITED

REVIE

W


Fig. 6. In vitro release profiles of 10 and 30% paclitaxel-loaded small and large PLLAmicrospheres.[108]

be attractive for cancer treatment. In most cases the challenge

was to increase the release rate of paclitaxel. Small size range

paclitaxel-loaded microspheres (1–30mm) composed of 50:50

PLA/EVA blend or PLLA have been shown to increase the

drug release rate when compared to large microspheres

(35–100mm), as shown in Figure 6.[107,108] It was observed that

as the drug loading percentage increased, the burst effect and

total drug release rates also increased.[108] Paclitaxe-

l-incorporated nanospheres were studied in an attempt to

Fig. 7. Release of carmustine in PBS. Each symbol represents the cumulative amoupaclitaxel-impregnated (c) polymer disc. The cumulative mass released was plotted versonly; insets of panels a and b). The slopes of the solid lines in the insets were determined


improve release kinetics. Nanospheres (mean diame-

ter¼ 133 nm) composed of 50:50 PDLGA (MW¼ 14,500Da)

exhibited a biphasic pattern characterized by an initial fast

release (FR) during the first 24 h, followed by a slower release

that reached about 80% in 9 days.[94] Another parameter

which was broadly studied was the effect of surfactant

incorporation into bioresorbable paclitaxel-loaded spheres.

Incorporation of hydrophilic surfactants such as PVA or PEG

resulted in faster release rates compared to spheres’

incorporation with hydrophobic surfactants or control speci-

mens which did not include any surfactants at all.[15,109,110]

Paclitaxel-loaded copolymer films fabricated from

poly(CPP-SA) and poly(FAD-SA) exhibited relatively slow

release (SR) profiles. However, release rates increased as the

polymer matrices were made more hydrophilic by increasing

the SA content.[102,111] The drug molecule itself has a strong

effect on the release profile. The release profiles of 20%-loaded

biodegradable PCPP-SA discs containing 20% carmustine

[1.3-bis(2-chloroethyl)-l-nitrosourea], 4-HC, or paclitaxel mea-

sured in PBS are shown in Figure 7. Release of carmustine and

4-HCwas linear with respect to the square root of time during

the first 4 days, suggesting that diffusion through the polymer

matrix controlled the rate of release. The rates of release for

both carmustine and 4-HC dropped exponentially over time

(from �2mg �day�1 initially to 1–10mg �day�1 after 30 days),

whereas the rate of release for paclitaxel remained constant

over the 30-day period (�3mg �day�1).

3.3. Drug-Eluting Fibers

Drug-eluting fibers may efficiently deliver antiproliferative

drugs locally at the tumor resection site or a few cm from the

nt of drug released from a carmustine-impregnated (a). 4-HCimpregnated (b), orus time (for all three agents) and the square root of time (for carmustine and 4-HCby linear regression.[96]


INVITED

REVIE

W


tumor to help target tumor metastases. The advantages of

fibers include ease of fabrication, high surface area, wide

range of possible physical structures, and localized delivery of

the bioactive agent to the target. Two basic types of

drug-eluting fibers have been reported: monolithic fibers

and reservoir fibers.[112–119]

(1) M

Fig. 8. Drug-eluting bioresorbable core/shell fiber platform.

B30

onolithic fibers: In these systems the drug is dissolved or

dispersed throughout the polymer fiber. For example:

curcumin, paclitaxel, and dexamethasone were melt spun

with PLLA to generate drug-loaded fibers[112] and aqu-

eous drugs were solution spun with PLLA.[113] Various

steroid-loaded fiber systems have demonstrated the

expected first-order release kinetics.[115,116]

(2) R
eservoir fibers: These are hollow fibers, where drugs
such as dexamethasone and methotrexane were added

to the internal section of the fiber post-melt extru-

sion.[117–119]

The main disadvantage of monolithic fibers is poor

mechanical properties, due to drug incorporation in the fiber.

Furthermore, many drugs and all proteins cannot tolerate the

high temperatures involved in the fabrication process of

monolithic fibers. Reservoir fibers also do not exhibit good

mechanical properties.

One of the current methods for producing drug-eluting

fibers involves electrospinning. Paclitaxel-loaded PLGAmicro

and nanofibers (diameters from around 30nm to 10mm) were

recently fabricated by electrospinning[106] to treat C6 glioma.

Cell viability test results suggested that the paclitaxel-loaded

PLGA nanofibers were effective for 72 h incubation. Ranga-

nath and Wang[120] developed paclitaxel-incorporated poly(-

D,L-lactide-co-glycolide) (PDLGA) implants in the form of

microfiber discs and sheets. Paclitaxel was released from the

PDLGA co-polymer implants (85:15 PDLGA and 50:50

PDLGA) for 80 days. An animal study confirmed brain tumor

growth inhibition after 24 days using these fibers. Coaxial

electrospinning has been reported as a method for preparing

core–shell structured nanofibers,[121] in which two compo-

nents can be coaxially and simultaneously electrospun

through different feeding capillary channels and where drug

may be incorporated.

4. Novel Antiproliferative Agent ReleaseSystems Applicable for Stent and for LocalCancer Treatment

4.1. The Concept of Core/Shell Fiber Structures

In the last section of this review article we chose to present

our composite fibers, which were designed to be used in both

applications, stents and local cancer treatment. The general

goal of our study was to develop and investigate a novel

drug-eluting bioresorbable core/shell fiber platform that will

successfully serve as a basic element for medical implants. The

concept of core/shell fibers is based on location of the drug

molecules in a separate compartment (‘‘shell’’) around a melt

4 http://www.aem-journal.com � 2012 WILEY-VCH Verlag GmbH & C

spun ‘‘core’’ fiber (Figure 8). This fiber platform is designed to

combine good mechanical properties with the desired drug

release profile.[122,123]

The shell (coating) is highly porous shell and is designed to

provide a large surface area for diffusion and thus control the

antiproliferative drug release. Most antiproliferative drugs

are hydrophobic and are therefore released slowly in an

aqueous environment. Furthermore, most antiproliferative

drugs are highly cytotoxic. Therefore, maintaining the drug

concentration between the effective and the toxic levels, in a

single dosage, is a complex task when incorporating

hydrophobic/cytotoxic drugs. Preparation of the porous

coating was based on the freeze-drying of water in oil

(inverted) emulsions technique.[122,123]

Our new fibers are designed for two purposes. The first is

use as basic elements of endovascular stents in order to

mechanically support blood vessels while delivering drugs

directly to the blood vessel wall for prevention of restenosis.

The second application offers local treatment of cancer

post-tumor resection in conjunction with standard treatment.

These systems, derived from drug-loaded emulsions, present

a new approach in the field of polymeric biomaterials and

controlled drug release.

Current drug-eluting biodegradable or biostable stent

coatings exhibit side effects due to delayed or incomplete

healing and are far from optimal in terms of controlled release

of drugs within the therapeutic range. Biodegradable stents

may overcome current DES endothelial related limitations

and suggest a larger drug reservoir if they could provide

mechanical stability along the healing period. Nevertheless,

these stents cannot carry enough drug because of the trade-off

between the mechanical properties and drug loading.

Although both types of DES have long been studied, there

is still no such drug release device, biodegradable or stable,

that can provide controlled release of a drug within the

therapeutic dosage with safe healing of the tissue. We present

a new approach for the basic elements of biodegradable

endovascular stents that mechanically support the blood

vessels while delivering drugs for prevention of restenosis

directly to the blood vessel wall. Our novel fiber systems,

derived from drug-loaded emulsions, may provide targeted


INVITED

REVIE

W


Fig. 9. The effect of the copolymer composition on the cumulative drug release profilefrom core/shell fiber structures (~, 50:50 PDLGA; �, 75:25 PDLGA): (a) Paclitaxelrelease and (b) FTS release. Plots of dMt/dt versus sqrt (1/t) for the first 5 weeks of release(in the small frames) indicate diffusion controlled region.[125]

Fig. 10. Schematic representation of a qualitative model describing the affect of theemulsion’s formulation and process kinetics on the drug release profile, through variousmechanisms.[125]

and controlled drug release without interfering with the

mechanical properties of the device. The highly porous

coating can also be applied successfully on metal stents.

The concept of drug-eluting devices for cancer treatment

has been studied extensively, and systems explored so far for

localized antiproliferative drug delivery in cancer treatment

include wafers, microspheres, and fibers. However, current

solutions include non-selectivity of the drug, sub-optimal

control over drug release, and problems in drug incorpora-

tion. Our delicate fibers are designed to combine good

strength with flexibility and can therefore be handled easily

and implanted in the desired location during and post-

surgery. Since these fibers are very delicate, they may also be

used stereotactically, obviating the need for surgery. Themain

advantages of our composite drug-loaded fibers include ease

of fabrication and high surface area for controlled release.

Furthermore, an integrated therapeutic approach for cancer

treatment may be highly advantageous and may provide high

local concentrations of antiproliferative drugs at the tumor

resection site in a controlled manner. This method could

prevent re-growth and metastasis of tumors and may enable

passage of drugs directly through the BBB, which is crucial in

cases of glioblastoma, a pathology for which there is still no

effective treatment.

4.2. Release Profiles of Antiproliferative Drugs and

Qualitative Model

The dense core of our composite fibers enables obtaining

the desired mechanical properties and the drug is located in a

porous shell so as not to affect the mechanical properties. The

shell is highly porous so as to enable release of the relatively

hydrophobic antiproliferative drugs in a desired manner. In

order to characterize the drug-eluting core/shell fiber plat-

form, the FTS and paclitaxel release from the fibers were

studied in light of the shells’ morphology and degradation

and weight loss profiles. The main results are presented in

Figure 9 and the whole study is described in details

elsewhere.[124,125]

A qualitative model describing the process! structure!drug release profile effects in our fibers with controlled release

of antiproliferative agents can be described as follows

(Figure 10): There are two routes by which the process affects

the drug-release profile: direct and indirect.[125]

4.2.1. Direct Route

The emulsion formulation (especially the host polymer)

affects the water uptake and swelling of the structure and,

therefore also the burst release of antiproliferative drugs such

as FTS (early mechanism). In such cases degradation of the

host polymer affects the release rate at a later stage. When a

relatively large and extremely hydrophobic drug such as

paclitaxel is incorporated into the shell, its diffusion through

the host polymer is much slower and massive degradation

and erosion of the host polymer must occur in order to

enable it.


4.2.2. Indirect Route

The effect of the process on the microstructure occurs also

via an emulsion stability mechanism. The emulsion stability

determines the surface area for diffusion through the

microstructure, e.g., the surface area increases when porosity

is high and pore size is low. For example, the 50:50 PDLGA’s

structure is finer than that of the 75:25 PDLGA. This affects

both the burst release and later release.

The most important parameter which affects the release

behavior in this system is the copolymer composition. It

affects the water uptake and swelling and therefore the FTS


INVITED

REVIE

W


release profile (earlymechanism). The copolymer composition

affects the degradation rate of the polymer and therefore also

the paclitaxel release profile (late mechanism). The copolymer

composition thus plays a very important role in the drug

release profile via the direct route.

4.3. Biological Performance

In order to study the effect of the drug release profile on

cancer cells a set of experiments in which cells were directly

exposed to the FTS-loaded core/shell fibers was performed.

Both slow and FR fibers (Figure 9) were used to allow a

relatively slow and fast accumulation of FTS in the wells. A

Fig. 11. FTS loaded core/shell fiber structures inhibit growth or induce cell death of gliobladensity of 8� 103 cells/well in a 24-well plate in tetraplicate (n¼ 2). One day after plating, c(two fibers, each with a length of 1 cm, as described in the materials and methods). (a) Imagesthe fiber (magnification �100). (b) A single well edge to edge panoramic view shows a gradi(magnification �100, cells were counted using an image analysis software). (c) Images wereviability while the well containing the fast FTS fiber exhibited cell death (magnification �


relatively SR rate was obtained with shells based on 75:25

PDLGA, while a relatively fast FTS release rate was obtained

with shells based on 50:50 PDLGA. U87, A549, and EJ cells

were used in these experiments in order to document

inhibition of cell growth and induction of cell death. The

results indicate that FTS-eluting composite fibers can

effectively induce growth inhibition or cell death by a

gradient effect and a dose-dependent manner (Figure 11).

Hence, the combined effect of the targeted mechanism of FTS

as a Ras inhibitor together with the localized and controlled

release characteristics of the fiber is an advantageous

antiproliferative quality. These results were described I

details elsewhere.[126]

stoma cells by a gradient effect and dose-dependent manner. U87 cells were plated at aontrol fibers (not loaded with FTS), slow or fast FTS release fibers were added to each welltaken after 5 days of incubation with FTS fibers, in locations which are near and distant toential increase in the cell concentration with the increase in the distance from a SR fibertaken after 7 days of incubation with FTS fibers; the control fiber well presents high cell100). Note that the dark shape is the actual fiber.[126]


INVITED

REVIE

W


5. Conclusions

In summary, in this review article we describe various

systems for controlled release of antiproliferative agents,

designed to be used for stent applications and local cancer

treatment. DES are discussed in details with emphasis on

challenges and strategies, especially in stent coatings and

release mechanisms of the antiproliferative drugs. Drug-

eluting systems for local cancer treatment were described in

terms of matrix materials, processing techniques, system

formats (particles and fibers mainly), and drug-release

profiles. Recent advances in the fields of biodegradable

polymers and drug delivery aim to achieve better functioning

in these two applications, save life and enable improvements

in the patient’s quality of life. These will be achieved through

development of new drug delivery systems, that will enable

better control of the release profile of the highly hydrophobic

antiproliferative drugs, and new suitable polymeric systems

that will serve as matrix. An example for such systems is

described in the last section, which focuses on our new

composite fibers structures, for both applications. These

enable desired release profile due to structuring effects in

the highly porous biodegradable matrix. The understanding

of the relationships between processing, degradation beha-

vior, materials microstructure, and the resulting controlled

release mechanisms of the antiproliferative drugs, are

expected to lead to new designs that will advance

the therapeutic fields of DES and local cancer treatment.

Received: September 16, 2011

Final Version: April 3, 2012

Published online: May 16, 2012

[1] B. D. Ratner, Biomaterials Science: An Introduction to

Materials in Medicine, Academic Press, San Diego 1996.

[2] K. Park, Controlled Drug Delivery: Challenges and Strat-

egies, American Chemical Society, Washington, DC

1997.

[3] L. Brannon-Peppas, in Medical plastics and biomater-

ials magazine, 1997.

[4] S. Dumitriu, Polymeric Biomaterials, Marcel Dekker,

New York 1994.

[5] U. Edlund, A. C. Albertsson, Adv. Polym. Sci. 2002,

157, 67.

[6] R. A. Jain, Biomaterials 2000, 21, 2475.

[7] P. A. Gunatillake, R. Adhikari, Eur. Cells Mater. 2003,

5, 1.

[8] M. Chasin, R. S. Langer, Biodegradable Polymers as Drug

Delivery Systems, Marcel Dekker, New York 1990.

[9] A. G. Mikos, J. S. Temenoff, Electron. J. Biotechnol. 2000,

3, 114.

[10] R. P. Lanza, R. S. Langer, J. Vacanti, in Principles of

Tissue Engineering, Academic Press, San Diego 2000,

251.


[11] K. Luo, J. Yang, P. Kopeckova, J. Kopecek, Macromol-

ecules 2011, 44, 2481.

[12] J. Yang, K. Luo, H. Pan, P. Kopeckova, J. Kopecek,

React. Funct. Polym. 2010, 71, 294.

[13] A. Gopferich, Biomaterials 1996, 17, 103.

[14] F. von Burkersroda, L. Schedl, A. Gopferich, Biomater-

ials 2002, 23, 4221.

[15] S. Feng, G. Huang, J. Controlled Release 2001, 71, 53.

[16] J. E. Sousa, P. W. Serruys, M. A. Costa, Circulation 2003,

107, 2383.

[17] Y. Li, N. Fukuda, S. Kunimoto, S. Yokoyama,

K. Hagikura, T. Kawano, T. Takayama, J. Honye,

N. Kobayashi, H. Mugishima, S. Saito, K. Serie,

J. Cardiovasc. Pharmacol. 2006, 48, 184.

[18] S. H. Hofma, W. J. van der Giessen, B. M. van Dalen,

P. A. Lemos, E. P. McFadden, G. Sianos, J. M. Ligthart,

D. van Essen, P. J. de Feyter, P. W. Serruys, Eur. Heart J

2006, 27, 166.

[19] S. K. Dordunoo, H.M. Burt, Int. J. Pharm. 1996, 133, 191.

[20] A. H. Gershlick, Heart (British Cardiac Society) 2005, 91

(Suppl 3), iii24.

[21] R. Moreno, C. Macaya, Curr. Med. Chem. 2005, 3, 221.

[22] S. Venkatraman, F. Boey, J. Controlled Release 2007, 120,

149.

[23] N. von Bubnoff, F. Schneller, C. Peschel, J. Duyster,

Lancet 2002, 359, 487.

[24] J. George, J. Sack, I. Barshack, P. Keren, I. Goldberg,

R. Haklai, G. Elad-Sfadia, Y. Kloog, G. Keren, Arter-

ioscler. Thromb. Vasc. Biol. 2004, 24, 363.

[25] Y. Kloog, A. D. Cox, Sem. Cancer Biol. 2004, 14, 253.

[26] A. Shikanov, B. Vaisman, M. Y. Krasko, A. Nyska,

A. J. Domb, J. Biomed. Mater. Res. 2004, 69, 47.

[27] F. Alexis, S. S. Venkatraman, S. K. Rath, F. Boey, J.

Controlled Release 2004, 98, 67.

[28] D. L. Fischman, M. B. Leon, D. S. Baim, R. A. Schatz,

M. P. Savage, I. Penn, K. Detre, L. Veltri, D. Ricci,

M. Nobuyoshi, M. Cleman, R. Heuser, D. Almond,

P. S. Teirstein, R. D. Fish, A. Colombo, J. Brinker,

J. Moses, A. Shaknovish, J. Hirshfeld, S. Bailey,

S. Ellis, R. Rake, S. Goldberg, N. Engl. J. Med. 1994,

331, 496.

[29] J. J. Popma, R. M. Califf, E. J. Topol, Circulation 1991, 84,

1426.

[30] P. W. Serruys, P. de Jaegere, F. Kiemeneij, C. Macaya,

W. Rutsch, G. Heyndrickx, H. Emanuelsson, J. Marco,

V. Legrand, P. Materne, J. Belardi, U. Sigwart,

A. Colombo, J. J. Goy, P. van den Heuvel, J. Delcan,

M. A. Morel, N. Engl. J. Med. 1994, 331, 489.

[31] A. Colombo, J. Drzewiecki, A. Banning, E. Grube,

K. Hauptmann, S. Silber, D. Dudek, S. Fort,

F. Schiele, K. Zmudka, G. Guagliumi, M. E. Russell,

Circulation 2003, 108, 788.

[32] D. R. Holmes, Jr, M. B. Leon, J. W. Moses, J. J. Popma,

D. Cutlip, P. J. Fitzgerald, C. Brown, T. Fischell,

S. C.Wong,M.Midei, D. Snead, R. E. Kuntz, Circulation

2004, 109, 634.


INVITED

REVIE

W


[33] M. N. Babapulle, M. J. Eisenberg, Circulation 2002, 106,

2859.

[34] M. Freed, R. D. Safian, C. Grines, Manual of Interven-

tional Cardiology, Physicians’ Press, Birmingham, MI

1996.

[35] G. W. Stone, J. W. Moses, S. G. Ellis, J. Schofer,

K. D. Dawkins, M. C. Morice, A. Colombo,

E. Schampaert, E. Grube, A. J. Kirtane, D. E. Cutlip,

M. Fahy, S. J. Pocock, R. Mehran, M. B. Leon,N. Engl. J.

Med. 2007, 356, 998.

[36] M. Joner, A. V. Finn, A. Farb, E. K.Mont, F. D. Kolodgie,

E. Ladich, R. Kutys, K. Skorija, H. K. Gold, R. Virmani,

J. Am. Coll. Cardiol. 2006, 48, 193.

[37] J. Kotani, M. Awata, S. Nanto, M. Uematsu, F. Oshima,

H. Minamiguchi, G. S. Mintz, S. Nagata, J. Am. Coll.

Cardiol. 2006, 47, 2108.

[38] W. Nammas, Future Cardiol. 2012, 8, 1.

[39] J. Daemen, P. W. Serruys, Circulation 2007, 116, 316.

[40] M. Zilberman, A. Kraitzer, O. Grinberg, J. J. Elsner,

Handbook of Experimental Pharmacology, Springer-

Verlag, Berlin, Heidelberg 2010, p. 299.

[41] J. Aoki, A. C. Abizaid, P. W. Serruys, A. T. Ong,

E. Boersma, J. E. Sousa, N. Bruining, J. Am. Coll. Cardiol.

2005, 46, 1670.

[42] B.S. Corporation, 2007.

[43] M. C. Morice, P. W. Serruys, P. Barragan, C. Bode,

G. A. Van Es, H. P. Stoll, D. Snead, L. Mauri,

D. E. Cutlip, E. Sousa, J. Am. Coll. Cardiol. 2007, 50,

1299.

[44] C. D. Rogers, Rev. Cardiovasc. Med. 2004, 5 (Suppl 2), S9.

[45] S. V. Rao, R. M. Califf, J. M. Kramer, E. D. Peterson,

T. P. Gross, C. J. Pepine, D. O. Williams, D. Donohoe,

R. Waksman, R. Mehran, M. W. Krucoff, Am. Heart J.

2008, 156, 201.

[46] B. Lagerqvist, S. K. James, U. Stenestrand, J. Lindback,

T. Nilsson, L.Wallentin,N. Engl. J. Med. 2007, 356, 1009.

[47] R. Virmani, G. Guagliumi, A. Farb, G. Musumeci,

N. Grieco, T. Motta, L. Mihalcsik, M. Tespili,

O. Valsecchi, F. D. Kolodgie, Circulation 2004, 109, 701.

[48] A. V. Finn, M. Joner, G. Nakazawa, F. Kolodgie,

J. Newell, M. C. John, H. K. Gold, R. Virmani, Circula-

tion 2007, 115, 2435.

[49] A. Finkelstein, D. McClean, S. Kar, K. Takizawa,

K. Varghese, N. Baek, K. Park, M. C. Fishbein,

R. Makkar, F. Litvack, N. L. Eigler, Circulation 2003,

107, 777.

[50] P. A. Revell, M. Braden, M. A. Freeman, Biomaterials

1998, 19, 1579.

[51] J. Fajadet, W. Wijns, G. J. Laarman, K. H. Kuck,

J. Ormiston, T. Munzel, J. J. Popma, P. J. Fitzgerald,

R. Bonan, R. E. Kuntz, Circulation 2006, 114, 798.

[52] J. K. Jackson, J. Smith, K. Letchford, K. A. Babiuk,

L. Machan, P. Signore, W. L. Hunter, K. Wang,

H. M. Burt, Int. J. Pharm. 2004, 283, 97.

[53] C. J. Creel, M. A. Lovich, E. R. Edelman, Circ. Res. 2000,

86, 879.


[54] A. W. Heldman, L. Cheng, G. M. Jenkins, P. F. Heller,

D. W. Kim, M. Ware, Jr, C. Nater, R. H. Hruban,

B. Rezai, B. S. Abella, K. E. Bunge, J. L. Kinsella,

S. J. Sollott, E. G. Lakatta, J. A. Brinker, W. L. Hunter,

J. P. Froehlich, Circulation 2001, 103, 2289.

[55] X. Wang, S. S. Venkatraman, F. Y. Boey, J. S. Loo,

L. P. Tan, Biomaterials 2006, 27, 5588.

[56] P. W. Serruys, G. Sianos, A. Abizaid, J. Aoki, P. den

Heijer, H. Bonnier, P. Smits, D. McClean, S. Verheye,

J. Belardi, J. Condado, M. Pieper, L. Gambone,

M. Bressers, J. Symons, E. Sousa, F. Litvack, J. Am. Coll.

Cardiol. 2005, 46, 253.

[57] E. Grube, L. Bullesfeld, J. Interv. Cardiol. 2002, 15,

471.

[58] A. J. Carter,M. Aggarwal, G. A. Kopia, F. Tio, P. S. Tsao,

R. Kolata, A. C. Yeung, G. Llanos, J. Dooley, R. Falotico,

Cardiovasc. Res. 2004, 63, 617.

[59] S. H. Duda, T. C. Poerner, B.Wiesinger, J. H. Rundback,

G. Tepe, J. Wiskirchen, K. K. Haase, J. Vasc. Interv.

Radiol. 2003, 14, 291.

[60] J. E. Sousa, P. W. Serruys, M. A. Costa, Circulation 2003,

107, 2274.

[61] G. W. Stone, M. Midei, W. Newman, M. Sanz,

J. B. Hermiller, J. Williams, N. Farhat, K. W. Mahaffey,

D. E. Cutlip, P. J. Fitzgerald, P. Sood, X. Su, A. J. Lansky,

JAMA 2008, 299, 1903.

[62] M. Joner, G. Nakazawa, A. V. Finn, S. C. Quee,

L. Coleman, E. Acampado, P. S. Wilson, K. Skorija,

Q. Cheng, X. Xu, H. K. Gold, F. D. Kolodgie, R. Virmani,

J. Am. Coll. Cardiol. 2008, 52, 333.

[63] G. W. Stone, M. Midei, W. Newman, M. Sanz,

J. B. Hermiller, J. Williams, N. Farhat, R. Caputo,

N. Xenopoulos, R. Applegate, P. Gordon, R. M. White,

K. Sudhir, D. E. Cutlip, J. L. Petersen, Circulation 2009,

119, 680.

[64] E. Kedhi, K. S. Joesoef, E. McFadden, J. Wassing, C. van

Mieghem, D. Goedhart, P. C. Smits, Lancet 2010, 375,

201.

[65] M. B. Leon, L. Mauri, J. J. Popma, D. E. Cutlip, E.

Nikolsky, C. O’Shaughnessy, P. A. Overlie, B. T.

McLaurin, S. L. Solomon, J. S. Douglas, Jr, M. W. Ball,

R. P. Caputo, A. Jain, T. R. Tolleson, B. M. Reen, III, A. J.

Kirtane, P. J. Fitzgerald, K. Thompson, D. E. Kandzari,

J. Am. Coll. Cardiol. 2010, 55, 543.

[66] M. B. Leon, E. Nikolsky, D. E. Cutlip, L. Mauri, H.

Liberman, H. Wilson, J. Patterson, J. Moses, D. E.

Kandzari, JACC 2010, 3, 1043.

[67] I. Meredith, S. Worthley, R. Whitbourn, D. Walters,

J. Popma, D. Cutlip, P. Fitzgerald, Am. J. Cardiol. 2007,

100, 155l.

[68] P. Vranckx, P. W. Serruys, S. Gambhir, S. J. Eduardo,

A. Abizaid, P. Lemos, E. Ribeiro, S. I. Dani, J. J. Dalal,

V. Mehan, A. Dhar, A. L. Dutta, K. N. Reddy, R. Chand,

A. Ray, Eur. Interv. 2006, 2, 310.

[69] M. W. Krucoff, D. J. Kereiakes, J. L. Petersen,

R. Mehran, V. Hasselblad, A. J. Lansky, P. J. Fitzgerald,


INVITED

REVIE

W


J. Garg, M. A. Turco, C. A. Simonton, III, S. Verheye,

C. L. Dubois, R. Gammon, W. B. Batchelor,

C. D. O’Shaughnessy, J. B. Hermiller, Jr, J. Schofer,

M. Buchbinder, W. Wijns, J. Am. Coll. Cardiol. 2008,

51, 1543.

[70] U. Westedt, M. Wittmar, M. Hellwig, P. Hanefeld,

A. Greiner, A. K. Schaper, T. Kissel, J. Controlled Release

2006, 111, 235.

[71] A. Breitenbach, D. Mohr, T. Kissel, J. Controlled Release

2000, 63, 53.

[72] J. Aoki, P. W. Serruys, H. van Beusekom, A. T. Ong,

E. P. McFadden, G. Sianos, W. J. van der Giessen,

E. Regar, P. J. de Feyter, H. R. Davis, S. Rowland,

M. J. Kutryk, J. Am. Coll. Cardiol. 2005, 45, 1574.

[73] H. J. Duckers, S. Silber, R. de Winter, P. den Heijer,

B. Rensing, M. Rau, H. Mudra, E. Benit, S. Verheye,

W. Wijns, P. W. Serruys, EuroIntervention 2007, 3,

67.

[74] H. J. Duckers, T. Soullie, P. den Heijer, B. Rensing,

R. J. de Winter, M. Rau, H. Mudra, S. Silber, E. Benit,

S. Verheye, W. Wijns, P. W. Serruys, EuroIntervention

2007, 3, 350.

[75] D. Brendan, R. H. David, Med. Devices: Evid. Res. 2009,

2, 47.

[76] G. Tepe, H. P. Wendel, S. Khorchidi, J. Schmehl,

J. Wiskirchen, B. Pusich, C. D. Claussen, S. H. Duda,

J. Vasc. Interv. Radiol. 2002, 13, 1029.

[77] A. Farb, G. Sangiorgi, A. J. Carter, V. M. Walley,

W. D. Edwards, R. S. Schwartz, R. Virmani, Circulation

1999, 99, 44.

[78] H. Tamai, K. Igaki, E. Kyo, K. Kosuga, A. Kawashima,

S. Matsui, H. Komori, T. Tsuji, S. Motohara, H. Uehata,

Circulation 2000, 102, 399.

[79] S. Commandeur, H. M. van Beusekom, W. J. van der

Giessen, J. Interv. Cardiol. 2006, 19, 500.

[80] R. Waksman, J. Interv. Cardiol. 2006, 19, 414.

[81] A. M. Lincoff, J. G. Furst, S. G. Ellis, R. J. Tuch,

E. J. Topol, J. Am. Coll. Cardiol. 1997, 29, 808.

[82] T. Yamawaki, H. Shimokawa, T. Kozai, K. Miyata,

T. Higo, E. Tanaka, K. Egashira, T. Shiraishi,

H. Tamai, K. Igaki, A. Takeshita, J. Am. Coll. Cardiol.

1998, 32, 780.

[83] I. Uurto, J. Mikkonen, J. Parkkinen, L. Keski-Nisula,

T. Nevalainen, M. Kellomaki, P. Tormala, J. P. Salenius,

J. Endovasc. Ther. 2005, 12, 371.

[84] F. Vogt, A. Stein, G. Rettemeier, N. Krott, R. Hoffmann,

J. vom Dahl, A. K. Bosserhoff, W. Michaeli, P. Hanrath,

C. Weber, R. Blindt, Eur. Heart J. 2004, 25, 1330.

[85] K. T. Nguyen, S. H. Su, M. Zilberman, P. Bohluli, P.

Frenkel, L. Tang, R. C. Eberhart, in Tissue Engineering

and Novel Delivery Systems (Eds: M. Yaszemski, D.

Rantolo, K. U. Lewandrowski, H. V, A. D, W. D),

Marcel Dekker, New York 2004, 107.

[86] M. Zilberman, N. D. Schwade, R. C. Eberhart, J. Biomed.

Mater. Res., Part B 2004, 69, 1.


[87] J. A. Ormiston, P. W. Serruys, E. Regar, D. Dudek,

L. Thuesen, M. W. Webster, Y. Onuma, H. M. Garcia-

Garcia, R. McGreevy, S. Veldhof, Lancet 2008, 371, 899.

[88] P. W. Serruys, J. A. Ormiston, Y. Onuma, E. Regar,

N.Gonzalo,H.M.Garcia-Garcia, K.Nieman,N. Bruining,

C. Dorange, K. Miquel-Hebert, S. Veldhof, M. Webster,

L. Thuesen, D. Dudek, Lancet 2009, 373, 897.

[89] M. Buchbinder, Foundation for Cardiovascular Medicine,

2007.

[90] N. Kukreja, Y. Onuma, J. Daemen, P. W. Serruys,

Pharmacol. Res. 2008, 57, 171.

[91] A. B. Dhanikula, R. Panchagnula, Int. J. Pharm. 1999,

183, 85.

[92] F. Alexis, E. P. Pridgen, R. Langer, O. C. Farokhzad, in

Handbook of Experimental Pharmacology (Drug Delivery),

Vol. 197 (Ed: M. Schafer-Korting), Springer-Verlag,

Berlin, Heidelberg 2009.

[93] L. Mu, S. S. Feng, J. Controlled Release 2003, 86, 33.

[94] C. Fonseca, S. Simoes, R. Gaspar, J. Controlled Release

2002, 83, 273.

[95] C. I. Winternitz, J. K. Jackson, A.M. Oktaba, H.M. Burt,

Pharm. Res. 1996, 13, 368.

[96] L. K. Fung, M. G. Ewend, A. Sills, E. P. Sipos,

R. Thompson, M. Watts, O. M. Colvin, H. Brem,

W. M. Saltzman, Cancer Res. 1998, 58, 672.

[97] E. C. Holland, Nat. Rev. Genet. 2001, 2, 120.

[98] H. B. Newton, Am. Fam. Physician 1994, 49, 787.

[99] L. Goldberg, Y. Kloog, Cancer Res. 2006, 66, 11709.

[100] F. H. Hochberg, A. Pruitt, Neurology 1980, 30, 907.

[101] A. G. de Boer, I. C. van der Sandt, P. J. Gaillard, Annu.

Rev. Pharm. Toxicol. 2003, 43, 629.

[102] K. A. Walter, M. A. Cahan, A. Gur, B. Tyler, J. Hilton,

O. M. Colvin, P. C. Burger, A. Domb, H. Brem, Cancer

Res. 1994, 54, 2207.

[103] H. Brem, M. S. Mahaley, Jr, N. A. Vick, K. L. Black, S. C.

Schold, Jr, P. C. Burger, A. H. Friedman, I. S. Ciric, T.W.

Eller, J. W. Cozzens, J. N. Kenealy, J. Neurosurg. 1991,

74, 441.

[104] M.Westphal, Z. Ram, V. Riddle, D. Hilt, E. Bortey, Acta

Neurochir. 2006, 148, 269.

[105] S. A. Limentani, A. Asher, M. Heafner, J. W. Kim,

R. Fraser, J. Neurooncol. 2005, 72, 241.

[106] J. Xie, C. H. Wang, Pharm. Res. 2006, 23, 1817.

[107] H. M. Burt, J. K. Jackson, S. K. Bains, R. T. Liggins,

A. M. Oktaba, A. L. Arsenault, W. L. Hunter, Cancer

Lett. 1995, 88, 73.

[108] R. T. Liggins, H. M. Burt, Int. J. Pharm. 2004, 281, 103.

[109] Y. Dong, S. S. Feng, Biomaterials 2004, 25, 2843.

[110] G. Ruan, S. S. Feng, Biomaterials 2003, 24, 5037.

[111] A. H. Shikani, A. J. Domb, Laryngoscope 2000, 110, 907.

[112] S. H. Su, R. Y. Chao, C. L. Landau, K. D. Nelson,

R. B. Timmons, R. S. Meidell, R. C. Eberhart, Ann.

Biomed. Eng. 2003, 31, 667.

[113] N. Alikacem, T. Yoshizawa, K. D. Nelson, C. A.Wilson,

Invest. Ophthalmol. Vis. Sci. 2000, 41, 1561.


INVITED

REVIE

W


[114] R. L. Dunn, D. H. Lewis, G. JM, ‘‘Monolithic fibers for

controlled delivery of tetracycline’’, presented at Proc.

Int. Symp. Control. Rel. Bioact. Mater, 1982.

[115] R. L. Dunn, J. P. English, W. C. Stoner, A. G. Potter,

B. H. Perkins, ‘‘Biodegradable fibers for the controlled

release of tetracycline in treatment of peridontal

disease’’, presented at Proc. Int. Symp. Control. Rel.

Bioact. Mater, 1987.

[116] R. L. Dunn, D. H. Lewis, L. R. Beck, in Controlled Release

of Pesticides and Pharmaceuticals, Plenum press, New

York 1981, 125.

[117] M. J. D. Eenink, J. Feijen, J. Oligslanger, J. H. M.

Albers, J. C. Rieke, G. PJ, J. Controlled Release 1987,

6, 225.

[118] L. Lazzeri, M. G. Cascone, L. Quiriconi, L. Morabito,

P. Giusti, Polym. Int. 2005, 54, 101.


[119] G. Polacco, M. G. Cascone, L. Lazzeri, S. Ferrara,

P. Giusti, Polym. Int. 2002, 51, 1464.

[120] S. H. Ranganath, C. H. Wang, Biomaterials 2008, 29,

2996.

[121] S. Yan, L. Xiaoqiang, L. Shuiping, M. Xiumei,

S. Ramakrishna, Colloids Surf. 2009, 73, 376.

[122] A. Kraitzer, M. Zilberman, J. Biomed. Mater. Res. 2007,

81, 427.

[123] M. Zilberman, A. Kraitzer, J. Biomed. Mater. Res. 2008,

84, 313.

[124] A. Kraitzer, L. Ofek, R. Schreiber, M. Zilberman,

J. Controlled Release 2008, 126, 139.

[125] A. Kraitzer, Y. Kloog, M. Zilberman, Eur. J. Pharm. Sci.

2009, 37, 351.

[126] A. Kraitzer, Y. Kloog, R. Haklai, M. Zilberman,

J. Pharm. Sci. 2011, 100, 133.


Date post:	12-Sep-2020
Category:	Documents
Upload:	others
View:	0 times
Download:	0 times

DOI: 10.1002/adem.201180088 INVITED REVIEW Controlled Release …meitalz/Articles/D5.pdf · 2016....

Documents