doi: 10.1002/adem.201180088 invited review controlled release …meitalz/articles/d5.pdf · 2016....
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DOI: 10.1002/adem.201180088Controlled 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]
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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
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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
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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
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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]
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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.
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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]
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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
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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.
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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
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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,
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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]
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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
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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
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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
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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
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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
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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
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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
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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]
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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 drugssuch 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
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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.
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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
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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 �
B306 http://www.aem-journal.com � 2012 WILEY-VCH Verlag GmbH & C
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]
o. KGaA, Weinheim ADVANCED ENGINEERING MATERIALS 2012, 14, No. 6
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REVIE
W
A. Kraitzer and M. Zilberman/Controlled Release of Antiproliferative Drugs
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
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