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Journal of Controlled Release 126 (2008) 139 – 148
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Long-term in vitro study of paclitaxel-eluting bioresorbable
core/shell fiber structures
Amir Kraitzer, Lia Ofek, Reut Schreiber, Meital Zilberman ⁎
Department of Biomedical Engineering, Faculty of Engineering, Tel-Aviv University, Tel-Aviv 69978, Israel
Received 10 October 2007; accepted 16 November 2007
Available online 28 November 2007
Abstract
Paclitaxel-eluting bioresorbable core/shell fiber structures for stent applications and local cancer treatment were developed and studied. These
structures were composed of a polyglyconate core and a porous PDLGA shell loaded with the anti-proliferative agent paclitaxel, prepared using
freeze drying of inverted emulsions. The investigation of these new composite fibers focused on the effects of the emulsion's composition
(formulation) and process kinetics on the long-term drug release from the fibers, in light of the shell's morphology and degradation profile.
Paclitaxel release from the porous shell was relatively slow due to its extremely hydrophobic nature. It exhibited three phases of release, which
corresponded to the degradation profile of the host PDLGA. We found that the effect of the emulsion formulation on the release profile is more
significant than the effect of the process kinetics. The copolymer composition had the most dominant effect on the drug release profile from the
composite fibers. The polymer content also affected the release profile, whereas the drug content and the organic:aqueous phase ratio resulted in
minor effects. Emulsions with a less hydrophobic nature are favorable for effective controlled release of the hydrophobic paclitaxel from the
porous shell.
© 2007 Elsevier B.V. All rights reserved.
Keywords: Paclitaxel; Drug-eluting fibers; Controlled release; Poly(DL-lactic-co-glycolic acid); Degradation
1. Introduction
Drug-loaded fibers have a high potential for various biomedical applications. Their advantages include ease of fabrication and a high surface area for controlled release and localized
delivery of bioactive agents to their target. Only few controlled
release fiber systems based on polymers have been investigated
to date [1–8]. The two basic types of drug-loaded fibers that
have been reported are monolithic fibers and reservoir fibers. In
systems that use monolithic fibers, the drug is dissolved or
dispersed throughout the polymer fiber. For example, curcumin,
paclitaxel and dexamethasone have been melt spun with poly
(L-lactic acid) (PLLA) to generate drug-loaded fibers [1] and
aqueous drugs have been solution spun with PLLA [2]. Various
steroid-loaded fiber systems have demonstrated expected firstorder release kinetics [3–5]. In systems that use hollow re⁎ Corresponding author. Tel.: +972 3 6405842; fax: + 972 3 6407939.
E-mail address: [email protected] (M. Zilberman).
0168-3659/$ - see front matter © 2007 Elsevier B.V. All rights reserved.
doi:10.1016/j.jconrel.2007.11.011
servoir fibers, drugs such as dexamethasone and methotrexane
have been added to the internal section of the fiber [6–8]. The
disadvantages of these systems include poor mechanical properties due to drug incorporation and limitations in drug loading.
Furthermore, many drugs and all proteins do not tolerate melt
processing and organic solvents.
In one of our recent studies we presented a new concept of
core/shell fiber structures which successfully meet these challenges [9]. These composite fibers combine a dense polymer
core fiber and a drug-loaded porous shell structure, i.e. the drug
or protein is located in a separate compartment (a “shell”)
around a melt spun “core” fiber. The shell is prepared using
mild processing conditions. This can result in good mechanical
properties as well as in the desired drug release profile. Our new
fibers are ideal for forming thin, delicate, biomedically important structures for various applications, such as fiber-based
bioresorbable endovascular stents that mechanically support
blood vessels while delivering drugs for preventing restenosis
directly to the blood vessel wall, or drug delivery systems for
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A. Kraitzer et al. / Journal of Controlled Release 126 (2008) 139–148
cancer treatment. The next two paragraphs focus on these two
applications.
Restenosis is the re-narrowing of a blood vessel which causes
a reduction in the diameter of the lumen, consequently restricting
blood flow after an intravascular procedure. In-stent restenosis is
mainly characterized by an abnormal increase in tissue growth
(neointimal hyperplasia) of the tissue composing the blood
vessel [10,11]. Re-narrowing is the result of a complex series of
biological events in response to the initial injury of the vessel
caused by balloon expansion and the presence of the permanent
stent implant [12]. Further progress in the treatment of coronary
artery disease has been the development of drug-eluting stents
that dramatically reduce the incidence of in-stent restenosis
to less than 5% [13]. Several drug-coated metal stents have
received FDA approval to date. The most popular ones are
TAXUS™ (Boston Scientific) — a paclitaxel-eluting stent, and
Cypher™ (J&J) — a sirolimus eluting stent. These are metal
stents coated with a thin stable dense polymer layer in which the
drug is located. Bioresorbable stents are designed to support the
vascular wall during the vessel healing process, gradually
transferring the mechanical load to the vessel wall while the stent
mass and strength decrease over time. Restenosis usually occurs
within 3 to 6 months after coronary intervention, and rarely
thereafter. Interest in bioresorbable vascular stents, with or
without drug delivery, is therefore increasing. These stents are
usually made of fibers. For example, Tamai et al. [14,15]
described the Igaki/Tamai stent, a bioresorbable balloon expandable zigzag coil design based on a PLLA monofilament. We
have developed and studied a fiber-based expandable stent
design [16,17], which was prepared using a linear continuous
coil array principle, by which four furled lobes convert into a
single large lobe upon balloon expansion. This expandable stent
design is based on melt-spun bioresorbable fibers that were
made from a relatively high molecular weight PLLA. It
demonstrated excellent initial radial compression strength and
good in vitro degradation resistance for 20 weeks.
Cancer may be defined as a disease characterized by uncontrolled proliferation and the spread of abnormal forms of the
body's own cells in the body [18]. In cancer therapy, surgical
treatment is usually carried out for patients with a resectable
carcinoma. However, treatment failure due to local recurrence
of primary tumors or metastatic spread often occurs. Surgical
adjuvant chemotherapy employing an anticancer agent during
or immediately after surgery would therefore be beneficial [18].
A delivery system loaded with paclitaxel at the tumor resection
site will provide a high local concentration of the drug, which is
detrimental to malignant cells that may have survived surgery,
thus preventing tumor re-growth and metastasis. In glioblastoma, localized prolonged release of anti-proliferative drugs,
which induces fewer side effects, can be applied to regions
where systemic administration is not effective. Drug delivery systems explored for localized paclitaxel delivery to date
include microspheres, surgical pastes and implants [18]. An
example for a biodegradable implant is the Gliadel® wafer,
a biodegradable wafer used to deliver the anticancer drug
carmustine directly into a brain tumor site after the surgical
removal of the tumor. It has been FDA approved since 1996.
Paclitaxel is a potent cell proliferation inhibitor and is known
to be very effective in the treatment of cancer [18] as well as in
preventing restenosis [18,19]. However, it is potentially cytotoxic, and the paclitaxel dose that can be delivered is relatively
small [20,21]. Paclitaxel has been shown to markedly attenuate
stent-induced intimal thickening of the lumen [22,23]. Paclitaxel's anti-proliferative effect is reversible [24]. Its short cellular
residence time (1 h), along with the reversible anti-proliferative
activity, suggest that it should be formulated in sustainedrelease dosage form [25]. The TAXUS trials revealed significant
inhibition of coronary stenosis by paclitaxel [26–28]. Drugeluting polymer-coated stents have thus moved into the limelight as vehicles for local drug administration [29,30]. In addition to usage in bioresorbable drug-eluting stents, our novel
coating technique can be used for stable stents, where only the
drug-loaded coat degrades with time. Studies on paclitaxeleluting bioresorbable microspheres for cancer treatment [31,32]
have shown that its release in an aqueous medium is relatively
slow due to paclitaxel's high hydrophobicity. Since paclitaxel
has demonstrated effectiveness in both drug-eluting stents and
local cancer treatment applications, its controlled release in a
desired manner is highly important. Highly porous structures
that encapsulate the drug, such as our fiber's “shell”, may
therefore be advantageous for both applications.
The current study focuses on composite core/shell fiber
structures loaded with paclitaxel. The porous shell (drug-containing section) was prepared using the technique of freeze
drying an inverted emulsion. The effect of the emulsion's
composition (formulation) and process kinetics on the emulsion's stability and on the resulting shell's microstructure have
already been reported [9]. The current article focuses on the
long-term (9 months, relevant for cancer treatment) in vitro
release of paclitaxel and its mechanisms. It describes the effects
of the emulsion's formulation and process kinetics on the
paclitaxel release profile from the fibers, in light of the shell's
morphology and degradation profile. It should be noted that
these paclitaxel-eluting composite fiber structures can also
serve as a good model for composite fibers loaded with other
hydrophobic bioactive agents for a wide range of biomedical
applications.
2. Materials and methods
2.1. Materials
Maxon™ polyglyconate monofilament (3–0) sutures, with a
diameter of 0.20–0.25 mm, Syneture, USA were used as core
fibers.
Bioresorbable porous structures (the shell coating) were
made of the following polymers:
• 75/25 poly(DL-lactic-co-glycolic acid), inherent viscosity
(i.v.) = 0.65 dL/g (in CHCl3 at 30 °C, approximately
97,100 g/mol), Absorbable Polymer Technologies Inc.,
USA. This polymer is termed 75/25 PDLGA.
• 50/50 poly(DL-lactic-co-glycolic acid), inherent viscosity (i.v.) = 0.56 dL/g (in CHCl3 at 30 °C, approximately
A. Kraitzer et al. / Journal of Controlled Release 126 (2008) 139–148
80,000 g/mol), Absorbable Polymer Technologies Inc.,
USA. This polymer is termed 50/50 PDLGA.
Drug: Paclitaxel (Genexol™) was purchased from Sam Yang
Corp. (Seoul, Korea).
2.1.1. Reagents
1,1,1,3,3,3-Hexafluoro-2-propanol (H1008) was purchased
from Spectrum Chemical Mfg. Corp.
Chloroform (CHCl3) HPLC grade and methylene chloride
(CH2Cl2) HPLC grade were purchased from Frutarom, Israel.
Acetonitrile (CH3CN) HPLC grade was purchased from J.T.
Baker.
2.2. Preparation of core/shell fiber structures
2.2.1. Fiber surface treatment
The sutures were surface-treated in order to enhance the
adhesion between the core fiber and the coating. The polyglyconate fibers were slightly stretched on special holders and
dipped in 1,1,1,3,3,3-hexafluoro-2-propanol (hexafluorisopropanol) for 40 s. The fibers were then washed with ethanol and
dried at room temperature.
2.2.2. Emulsion formation
A known amount of 75/25 PDLGA was dissolved in chloroform to form an organic solution and paclitaxel was added
to the solution. Double-distilled water was then poured into
the organic phase (in a test tube) and homogenization of
the emulsion was performed using a homogenizer (Polytron
PT3100 Kinematica, 12 mm rotor) operating at 16,500 rpm
(medium rate) for 3 min, for most investigated samples. As a
reference sample we chose an emulsion formulation containing
17.5% w/v polymer in the organic solution, 1.43% w/w paclitaxel (relative to the polymer load), and an organic to aqueous
(O:A) phase ratio of 2:1 v/v. Other formulations included 22.5%
w/v polymer, 2.9% and 4.3% w/w paclitaxel, O:A phase ratios
of 4:1 and copolymer composition of 50/50 PDLGA. Certain
samples were prepared using homogenization rates of 5500 rpm
(low rate) or 25,000 rpm (high rate) in order to investigate the
effect of processing kinetics on the porous shell structure.
2.2.3. Core/shell fiber structure formation
The treated core polyglyconate fibers were dip-coated (while
placed on holders) in fresh emulsions and then frozen immediately in a liquid nitrogen bath. The holders + samples were
then placed in a pre-cooled (− 105 °C) freeze dryer (Virtis 101
equipped with a nitrogen trap) capable of working with organic
solvents (freezing temperature of the condenser was approximately − 105 °C) and freeze dried in order to preserve the
microstructure of the emulsion-based core/shell fiber structures.
Drying was performed in two stages:
1. The freeze dryer chamber pressure was reduced to 100 mTorr
while the temperature remained at − 105 °C.
2. The condenser was turned off and its plate temperature
slowly rose to room temperature while the pressure was
141
monitored between 100 mTorr and 700 mTorr. During this
step the liquid nitrogen trap condensed the excess water and
solvent vapors.
The samples were stored in desiccators until use.
2.2.4. Morphological characterization
The morphology of the composite core/shell fiber structures
(cryogenically fractured surfaces) was observed using a Jeol
JSM-6300 scanning electron microscope (SEM) at an accelerating voltage of 5 kV. The SEM samples were Au sputtered
prior to observation. The mean pore diameter of the observed
morphologies was analyzed using the Sigma Scan Pro software
and statistics were drawn using SPSS. Statistical significance
was determined using the ANOVA (Tukey–Kramer) method.
Statistical significance was determined at p b 0.05.
2.3. In vitro paclitaxel release studies
The composite core/shell fiber structures (triplicates) were
immersed in phosphate buffered saline (PBS) at 37 °C and
pH = 7.4 for 37 weeks in order to determine paclitaxel's release
kinetics from these structures. The release studies were conducted in closed glass vessels containing 3 ml PBS medium,
using a horizontal bath shaker operated at a constant rate of
130 rpm. The medium was removed (completely) periodically,
at certain sampling times, and fresh medium was introduced.
This experiment enabled simulating conditions similar to those
which exist in a blood vessel.
2.3.1. Extraction procedure
Paclitaxel extraction from the medium was performed as
follows: the 3 ml medium was completely removed at each time
point and placed in a scintillation vial. 3 ml acetonitrile and 1 ml
methylene chloride were added. Methylene chloride evaporation was performed under a nitrogen stream and the paclitaxel
concentration was then estimated using HPLC. An extraction
factor was used for correction. Known weights of paclitaxel
were dissolved in 1 ml of methylene chloride and subjected to
the same extraction procedure in order to determine the recovery efficiency of the extraction procedure. Recoveries were
always greater than 75% and the value of the measured drug
was corrected accordingly.
The paclitaxel content of the medium samples was determined
using Jasco High Performance Liquid Chromatography (HPLC)
with a UV 2075 plus detector and a reverse phase column (Zorbax
ODS 5 µm, inner diameter d = 4.6 mm, length = 150 mm), and
was kept at 25 °C. The mobile phase consisted of a mixture of
acetonitrile and double-distilled water (55/45, v/v) at a flow rate
of 1 ml/min with a quaternary gradient pump (PU 2089 plus)
without gradient. 100 µl samples were injected with an autosampler (AS 2057 Plus). The column effluent was eluted for
15 min and detected at 227 nm. The area of each eluted peak was
integrated using the EZstart software version 3.1.7.
The cumulative release profiles were determined relative to
the initial amount of paclitaxel in the composite fibers (quantity
released during the incubation period + the residue remaining in
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the fibers). All experiments were performed in triplicates. Results are presented as means ± standard errors. The effects of the
emulsion's formulation on the release profile were studied by
examining the following parameters: polymer content in the
organic phase (%w/v), drug content relative to polymer content
(%w/w), organic:aqueous (O:A) phase ratio and copolymer
composition. The effect of the process kinetics (homogenization
rate) on the release profile was also studied.
2.4. Residual drug recovery from composite fibers
Residual paclitaxel recovery from the composite fibers was
measured as follows: the fibers were placed in 1 ml methylene
chloride and 3 ml of a 50/50 acetonitrile/water solution was
added. Methylene chloride evaporation was performed under a
nitrogen (99.999%) stream and the paclitaxel concentration was
then estimated using HPLC. An extraction factor was used for
correction. Five solutions with known concentrations of
paclitaxel were subjected to the same extraction procedure in
order to determine the recovery efficiency of the extraction
procedure. Recoveries were always greater than 75% and the
value of the residual drug was corrected accordingly.
2.5. Paclitaxel's chemical stability
Two milligrams of paclitaxel were placed in phosphate buffered
saline (PBS) at 37 °C for 160 days in order to determine paclitaxel's
stability throughout the release experiment. Closed scintillation
vials containing 3 ml PBS medium were placed in a horizontal bath
shaker operated at a constant rate of 130 rpm. The paclitaxel content
of a single vial was measured at each point of time.
Paclitaxel was extracted as follows: 3 ml methylene chloride
were added to the 3 ml PBS/paclitaxel mixture and stirred for
10 min in order to dissolve the drug. One milliliter was carefully
removed from the bottom of the vial (the organic phase) using a
pipette and placed in a new vial. Five milliliters of acetonitrile/
double-distilled water (the medium) were added to the vial and
then evaporated under nitrogen conditions. The content of each
vial was transferred to a test tube and diluted to 20 ml. The
paclitaxel content of the medium samples was determined using
HPLC, as described above.
permeation chromatography (GPC). The GPC (Waters 21515
isocratic pump, operating temperature 40 °C using a column
oven) was equipped with a refractive index detector (Waters
2414, operating temperature 40 °C) and calibrated with poly-Llactic acid MW kit standards (Polysciences Inc., USA). Data
were analyzed using the Breeze version 3.3 software. The
samples were dissolved in methylene chloride, filtered, and
eluted through 4 Styragel columns (model WAT044234 HR1
THF, WAT044237 HR2, WAT044225 HR4, WAT054460 HR5,
300 × 7.8 mm, 5 µm particle diameter) at a flow rate of 1 ml/
min. The elution medium was Baker analyzed HPLC grade
methylene chloride.
3. Results and discussion
The freeze drying technique is unique in being able to preserve the liquid structure in solids. We used this technique in
order to produce inverted emulsions in which the continuous
phase contained polymer and drug dissolved in a solvent, with
water being the dispersed phase. SEM fractographs showing the
bulk morphology of the reference specimen are presented in
Fig. 1. The diameter of the treated core fiber was in the range of
200–250 µm and a shell thickness of 30–70 µm was obtained
for most emulsion formulations. Relatively high contents of
hydrophobic components (such as PDLGA and paclitaxel) resulted in an increase in shell thickness, due to the emulsion's
2.6. In vitro degradation study of the porous PDLGA structure
Porous 75/25 PDLGA film structures especially prepared for
degradation trial were based on the reference emulsion formulation.
The inverted emulsion was prepared as described earlier in the
paragraph about emulsion formation (in the Materials and methods
section describing preparation of core/shell fiber structures), poured
into an aluminum plate and freeze dried in liquid nitrogen. Each
dried film sample was then incubated in 20 ml phosphate buffered
saline (PBS) containing 0.05% (v/v) sodium azide (as preservative)
at 37 °C and pH = 7.4 for 17 weeks, under static conditions. Each
sample was taken out at weekly intervals and dried using a vacuum
oven (35 °C for 1 h), and stored in a dessicator.
The weight-average molecular weight and number-average
molecular weight of the samples were determined by gel
Fig. 1. SEM fractographs showing cross sections of the reference composite
fiber.
A. Kraitzer et al. / Journal of Controlled Release 126 (2008) 139–148
143
higher viscosity [9]. There are almost no gaps between the core
and the shell (Fig. 1b), indicating that the quality of the interface
between the fiber and the porous coating is high, i.e. the surface
treatment enabled good adhesion between core and shell. The
shell's porous structure in all studied specimens based on stable
emulsions contained round-shaped pores, usually within the 4–
10 µm range, with a porosity between 67 and 82%. The shell's
microstructure was uniform in each sample, probably due to
rapid quenching of the emulsion, which enabled preservation of
its microstructure [9]. The pores were partially interconnected
by smaller inner pores. The emulsions were stable only within a
certain formulation range. Only structures that were based on
unstable emulsion formulations demonstrated a pore size higher
than 10 µm. The effects of emulsion parameters (formulation)
and processing conditions on the shell morphology are described in detail in a previous publication [9]. The encapsulation
efficiency of all studied samples was in the 45–60% range.
The GPC curves of the porous 75/25 PDLGA reference
structure at various degradation times are presented in Fig. 2.
The curve of the initial structure (time 0) exhibits the expected single peak. As the structure degraded and low molecular
weight chains were produced, the peak shifted to the right
(higher release time, i.e. smaller molecular weight) and it became broader, indicating a larger molecular weight distribution.
The weight-average molecular weight (MW) of the porous
structure decreased with degradation time (Fig. 3a), as expected.
It is clear that the curve can be divided into two domains: (start–
week 10) and (week 10–week 19). The slope of the second
domain is double that of the first, indicating that the degradation
rate during the second domain is higher than during the first.
The cumulative release of paclitaxel from the reference
specimen for 37 weeks is presented in Fig. 3(b). Paclitaxel's
cumulative release exhibited the following three phases of
release:
Fig. 2. GPC curves of the 75/25 PDLGA porous structure at various times of
immersion in aqueous medium. The time is indicated next to each curve.
(a) The first phase of release (phase a) occurred during weeks
1–10, in which the drug was released in an exponential
manner, i.e. the rate of release decreased with time. Such a
release profile is typical of diffusion-controlled systems.
A minor initial burst release (less than 1%) was obtained
during the first day of release. Paclitaxel release from the
porous shell was relatively slow, mainly due to paclitaxel's extremely hydrophobic nature. Moreover, the release
rate decreased with time, since the drug had a progressively longer distance to pass and a lower driving force
for diffusion. During this phase of release the degradation
of the host polymer resulted in a decrease in MW from
60 kDa to 40 kDa (Fig. 3a).
(b) The second phase of release (phase b) occurred during
weeks 10–20, in which the drug was released at a constant, relatively high rate. The higher degradation rate
of the host polymer in this phase probably resulted in
polymer erosion and therefore enabled a faster rate of
drug release. During this phase of release the degradation
of the host polymer resulted in a decrease in MW from
40 kDa to 2 kDa (Fig. 3a).
(c) The third phase of release (phase c) occurred during
weeks 21–38, when the porous shell structure was already
destroyed due to intensive erosion. In fact, the drug was
released from oligomers and short chain polymers of less
than 2 kDa. Most of these shell remains are no longer
Fig. 3. (a) The weight-average molecular weight of the 75/25 PDLGA reference
as a function of degradation time. (b) Cumulative paclitaxel release profile from
the reference core/shell fiber specimen.
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A. Kraitzer et al. / Journal of Controlled Release 126 (2008) 139–148
be specific hydrophobic interactions between the paclitaxel and the polymer.
It is clear that the paclitaxel release profile during phases a
and b (Fig. 3b) corresponds to the degradation profile of the
porous host PDLGA shell (Fig. 3a). Intensive degradation of the
host polymer is necessary in order to obtain release of the highly
hydrophobic paclitaxel. Although a relatively small percentage
of the loaded drug was released, this is the relevant quantity for
the stent application [19]. Other investigators working on paclitaxel eluting systems also reported its relatively slow release
rate from various polymeric systems [33,34].
The shell's microstructure may affect the drug release profile, which is determined by both thermodynamic and kinetic
parameters. The thermodynamic parameters are actually the
emulsion's formulation, i.e. the polymer and drug contents, the
O:A ratio and the copolymer composition (relative quantities
of lactic acid and glycolic acid in PDLGA), whereas the kinetic considerations are actually the processing conditions,
which include the rate of homogenization. The effects of these
parameters on the paclitaxel's controlled release profile from
the fibers are described below (Fig. 4) with reference to the
shell's microstructure (Fig. 5) and its degradation profile.
3.1. Effect of polymer content
The effect of the polymer content of the organic phase on the
drug release from the porous shell is presented in Fig. 4a. In
both release phases, a and b, the release rate and quantity from
the formulation containing 17.5% w/v polymer were higher
Fig. 4. Effect of the emulsion formulation on the release profile of paclitaxel
from core/shell fiber structures: (a) effect of polymer content: — 17.5% w/v,
— 22.5% w/v, (b) effect of paclitaxel content: — 1.43% w/w, — 2.9% w/w,
— 4.3% w/w, (c) effect of O:A phase ratio: — 2:1 v/v, — 4:1, (d) copolymer
composition: — 75/25 PDLGA, — 50/50 PDLGA.
attached to the core fiber and the core fiber also undergoes
erosion. We would expect release of the entire loaded
drug during phases a and b. However, in our system the
highly hydrophobic paclitaxel is probably attached to the
surface of the hydrophobic 75/25 PDLGA's remains even
when their molecular weight is extremely low. There may
Fig. 5. SEM fractographs of composite fibers showing the effect of the
emulsion's formulation on the shell's microstructure: (a) The reference
specimen containing 17.5% w/v polymer, 1.43% w/w paclitaxel and phase
ratio of 2:1, (b) formulation with 22.5% w/v polymer, (c) formulation with 4.3%
w/w paclitaxel, (d) formulation with O:A ratio of 4:1, (e) formulation with
copolymer composition of 50/50 PDLGA.
A. Kraitzer et al. / Journal of Controlled Release 126 (2008) 139–148
than the release rate and quantity from the formulation containing 22.5% w/v polymer. Since larger drug quantities are
released during phase b, the differences between these two
formulations are more significant during this phase of release
than during phase a. It is suggested that since the pore size and
porosity remained almost unchanged with the emulsion's polymer content (Fig. 5a,b and Table 1) but denser “polymeric
walls” were probably created between adjacent pores of the
formulation with 22.5% w/v polymer, a relatively high polymer
content increases the interactions between the matrix and paclitaxel, resulting in a lower diffusion coefficient which results in
slower drug release. During the third phase of release, in which
the drug was released from shell remains with a very low
molecular weight, the polymer content had almost no effect on
the rate of drug release, as expected.
3.2. Effect of drug content
The effect of the paclitaxel content on its release profile is
presented in Fig. 4b. Since the drug content strongly affects the
emulsion's stability [9], a relatively low range of drug contents
(1.4% w/w–4.3% w/w) was used. The highly hydrophobic
paclitaxel diffuses out very slowly from the shell. Therefore, the
drug content did not show a significant effect on the release
profile during the first phase of release. During the second phase
of release the release rate and the amount of released drug
decreased slightly with the increase in the drug content from
2.8% w/w to 4.3% w/w (Fig. 4b). We would expect an increase
in the release rate with the increase in drug content, due to a
higher driving force for diffusion. However, paclitaxel is a
hydrophobic drug and a higher paclitaxel content in the organic
phase of the emulsion therefore results in a higher interfacial
tension (difference between the surface tension of the organic
and aqueous phases), leading to a less stable emulsion with
a larger pore size (Fig. 5a,c and Table 1). Larger pores are
expected to reduce the release rate when the porosity and
interconnectivity do not change too much. It can therefore
be concluded that in this specific range of drug contents, the
morphological changes do not enable the driving force for
Table 1
The examined specimens and the structural characteristics of their shells
Process parameters
Reference
Polymer content [% w/v]
Paclitaxel content [% w/w]
a
22.5
2.86
4.286
4:1
Mean pore size ±
SD [µm]
Mean porosity ±
SD [%]
6.42 ± 2.32
5.34 ± 1.64 ( b)
9.22 ± 4.54 ( b)
29.96 ± 3.1 ( b)
5.96 ± 2.41 ( b)
69 ± 6
72 ± 8
73 ± 3
82 ± 9
80 ± 2
Organic to aqueous phase
ratio [v/v]
Homogenization rate [rpm] 5500
8.23 ± 3.15 ( b)
77 ± 8
25,000 5.98 ± 1.6 (P = 0.053) 75 ± 10
PDLGA copolymer
50/50 4.1 ± 1.35 ( b)
67 ± 6
composition [w/w]
a
The reference sample was prepared using the following conditions:
emulsion-based on 75/25 PDLGA, 17.5% w/v polymer, 1.43% w/w paclitaxel,
O:A = 2:1 v/v, homogenization rate = 16,500 rpm.
b
Significant compared to the reference.
145
diffusion to affect paclitaxel's release profile. During the third
phase of release (20–38 weeks) the release rate from the remaining short polymers and oligomers of the formulations with
2.9% w/w and 4.3% w/w paclitaxel is significantly higher than
that obtained from the formulation with 1.4% w/w paclitaxel
(Fig. 5a). The driving force for diffusion plays a major role at this
stage, after intensive polymer erosion when the morphological
features of the initial microstructure no longer exist.
3.3. Effect of O:A phase ratio
The effect of the O:A phase ratio on paclitaxel's release
profile from the composite fibers and on the shell's microstructure are presented in Figs. 4c and 5a,d, respectively. The
release profile during release phases a and b (20 weeks) as well
as the pore size (Table 1) exhibited almost no change in the O:A
range of 2:1 to 4:1. It should be mentioned that the relatively
narrow O:A range chosen for this study resulted from the
emulsion's stability considerations. The porosity of samples
derived from emulsions with O:A ratios higher than 4:1 may not
be high enough to enable effective release of water-insoluble
agents such as paclitaxel. On the other hand, samples derived
from emulsions with O:A ratios less than 2:1 are not stable.
During the third phase of release (20–40 weeks) the release rate
from the 4:1 phase ratio was higher than from the 2:1 phase
ratio. A relatively high organic phase content provides more
binding sites for the hydrophobic paclitaxel. This higher drug
content enables a higher release rate only when the drug diffuses
out from very low molecular weight components which remain
in the third phase of release.
3.4. Effect of copolymer composition
The effect of the copolymer composition, i.e. the relative
quantities of lactic acid (LA) and glycolic acid (GA) in the
copolymer, on the drug release profile and on the shell's
microstructure is presented in Figs. 4d and 5a,e, respectively.
The rate of release from the 50/50 PDLGA formulation is higher
than that obtained from the 75/25 PDLGA formulation
(reference) during phases a and b (Fig. 4d). It is important to
note that the differences in the release rate and quantity are low
during the first phase of release but very high during the second
phase of release. After 20 weeks of degradation the 50/50
formulation released 70% of the drug, whereas the 75/25 formulation released only 13%. These differences are attributed
mainly to differences in the degradation rate of these two copolymers and differences in their hydrophilic/hydrophobic nature. The 50/50 copolymer degrades faster than the 75/25
copolymer and therefore during intensive erosion (phase b) it
releases the drug at a faster rate. Furthermore, the less hydrophobic nature of the 50/50 copolymer relative to the 75/25
copolymer results in less drug attachment to its surface and
therefore also contributes to a higher paclitaxel release rate.
The mean pore diameter increased with the increase in the
copolymer's LA content. The 50/50 composition exhibited a
mean pore diameter of 4.1 µm, whereas the 75/25 composition
(reference) exhibited a mean pore diameter of 6.4 µm (Table 1).
146
A. Kraitzer et al. / Journal of Controlled Release 126 (2008) 139–148
An increase in the copolymer's LA content actually increases
the surface tension, prompting the water droplets to flocculate
and coalescence. Therefore, increasing the mean pore diameter
affects the emulsion in the same manner as addition of paclitaxel. The smaller pore size of the 50/50 PDLGA relative to
the 75/25 PDLGA enables a higher surface area for diffusion.
Since the pores are partially interconnected by smaller inner
pores and the shell structure is highly porous, the smaller pore
size also contributes to the higher rate of paclitaxel release
during phases a and b.
It can be concluded that a higher glycolic acid content in the
copolymer, i.e. a less hydrophobic copolymer, enables higher
paclitaxel release from the fibers due to a combination of 3
phenomena: higher degradation rate of the host polymer, less
drug attachment to its surface and finer porous microstructure.
The degradation rates achieved during the second phase of
release in our study are actually higher than those reported in
other studies on paclitaxel release from bioresorbable polymers.
3.5. Effect of process kinetics
The effects of the emulsion's homogenization rate (prepared
during a 180 s homogenization) on the drug release from the
porous shell and on the microstructure are presented in Fig. 6.
The homogenization rate had practically no effect on the release
profile during the first phase of release (10 weeks), but had
some effect on the release profile during the second phase of
release (Fig. 6a) due to small changes in the shell's porous
structure (Fig. 6b,c and Table 1). An increase in homogenization
rate from 5500 rpm to 25,000 rpm resulted in a slight decrease
in mean pore size from 8.2 µm to 6.0 µm with almost no change
Fig. 6. (a) The effect of the stirring rate on the release profile of paclitaxel from
core/shell fiber structures: — low rate, — high rate, (b) SEM fractograph of
a fiber prepared using the low stirring rate (5500 rpm), (c) SEM fractograph of a
fiber prepared using the high stirring rate (25,000 rpm).
Fig. 7. A schematic representation of the qualitative model describing the
dependence of the drug release profile on the emulsion's formulation and
process kinetics.
in the porosity. The higher surface area for diffusion enabled
some increase in drug release rate and quantity.
In conclusion, a qualitative model describing the process →
structure → property effects in our system is presented in Fig. 7.
Diffusion of the drug molecules is controlled chemically, by the
emulsion formulation and also geometrically, through the surface area. The emulsion formulation directly influences the
diffusion by producing binding regions for the drug as more
hydrophobic materials are introduced into the emulsion, thus
delaying the drug molecules. The emulsion's formulation also
affects its stability during processing, and as a result determines
the freeze dried microstructure. The microstructure determines
the surface area available for diffusion, which affects the drug
release rate. In general, a high diffusion rate is achieved when
high porosity is combined with small pore size. In our system,
the process kinetics (rate of homogenization) showed a
considerably smaller effect on the microstructure and on the
paclitaxel release profile. It should be mentioned that degradation of the host polymer also affects the diffusion rate, as in
other bioresorbable drug delivery systems.
Finally, since long-term drug release was examined in this
study, the chemical stability of paclitaxel was measured, as
described in the experimental section. An example of an HPLC
run is presented in Fig. 8a. Two peaks are shown: a major peak
Fig. 8. (a) HPLC run showing the peaks of paclitaxel and its derivative, (b) the
derivative percentage as a function of immersion time in the aqueous medium.
A. Kraitzer et al. / Journal of Controlled Release 126 (2008) 139–148
of paclitaxel after an elution time of 4–5 min, and a minor
peak of its derivative after an elution time of 7 min. The
derivative percentage (derivative content divided by paclitaxel content) as a function of immersion time in the aqueous
medium is presented in Fig. 8b. As expected, the derivative
percentage increases with time, but even after 23 weeks it is
lower than 5%, thus demonstrating that 95% of the paclitaxel
remained stable.
4. Summary and conclusions
Bioresorbable core/shell fiber structures for biomedical applications were developed and studied. These structures were
composed of a polyglyconate core and a porous PDLGA shell
loaded with the anti-proliferative agent paclitaxel, prepared
using freeze drying of inverted emulsions. A long-term in vitro
release study was performed. The investigation of these new
composite fibers focused on the effects of the emulsion's composition (formulation) and process kinetics on the drug release
profile from the fibers, in light of the shell's morphology and
degradation profile. These unique fiber structures can be used as
basic elements of vascular stents and also for local cancer
treatment.
In general, porous “shell” structures (porosity of 67–82%
and pore size of 4–10 µm) were obtained with good adhesion to
the core fiber. Paclitaxel release from the porous shell was
relatively slow due to its extremely hydrophobic nature. It
exhibited three phases of release. It is clear that the first phase
(1–10 weeks) was governed by diffusion, whereas the second
phase (10–20 weeks) was governed by degradation. The release
profile of paclitaxel from the porous shell corresponds to the
degradation profile of the host 75/25 PDLGA.
The copolymer composition had the most dominant effect on
the drug release profile from the composite fibers. An increase
in the glycolic acid content (or decrease in lactic acid content)
resulted in a tremendous increase in the release rate during the
second phase, which was attributed mainly to the increased
degradation rate and decreased drug attachment to the host
polymer. A decrease in polymer content also resulted in an
increase in the release rate during the second phase of release,
whereas the drug content and the O:A phase ratio in the studied
range affected the release profile only during the third phase of
release.
The effect of the emulsion formulation on the release profile
is more significant than the effect of the process kinetics. Furthermore, emulsions with a less hydrophobic nature are favorable for effective controlled release of the hydrophobic
paclitaxel from the porous shell, since they do not provide many
binding regions for the drug and their higher stability enables
obtaining highly porous shell structures with a large surface area
for diffusion after freeze drying.
Acknowledgements
The authors are grateful to the RAMOT (Horowitz) Foundation and to the Slezak Foundation, Tel-Aviv University, for
supporting this research.
147
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