1. Accelerating protein release from microparticles for regenerative
medicine applications
Lisa J. White ⁎, Giles T.S. Kirby, Helen C. Cox, Roozbeh Qodratnama, Omar Qutachi,
Felicity R.A.J. Rose, Kevin M. Shakesheff
School of Pharmacy, University of Nottingham, Nottingham NG7 2RD, UK
a b s t r a c ta r t i c l e i n f o
Article history:
Received 23 July 2012
Received in revised form 6 February 2013
Accepted 13 February 2013
Available online xxxx
Keywords:
Poly (D,L-lactic -co-glycolic acid) (PLGA)
Microparticles
Microspheres
Controlled release
Double emulsion solvent evaporation
Growth factor delivery
There is a need to control the spatio-temporal release kinetics of growth factors in order to mitigate current
usage of high doses. A novel delivery system, capable of providing both structural support and controlled re-
lease kinetics, has been developed from PLGA microparticles. The inclusion of a hydrophilic PLGA–PEG–PLGA
triblock copolymer altered release kinetics such that they were decoupled from polymer degradation. A quasi
zero order release profile over four weeks was produced using 10% w/w PLGA–PEG–PLGA with 50:50 PLGA
whereas complete and sustained release was achieved over ten days using 30% w/w PLGA–PEG–PLGA with
85:15 PLGA and over four days using 30% w/w PLGA–PEG–PLGA with 50:50 PLGA. These three formulations
are promising candidates for delivery of growth factors such as BMP-2, PDGF and VEGF. Release profiles were
also modified by mixing microparticles of two different formulations providing another route, not previously
reported, for controlling release kinetics. This system provides customisable, localised and controlled delivery
with adjustable release profiles, which will improve the efficacy and safety of recombinant growth factor
delivery.
© 2013 Elsevier B.V. All rights reserved.
1. Introduction
In bone and cartilage repair, a complex cascade of biological
events is controlled by growth factor signalling at injury sites,
promoting progenitors and inflammatory cells to migrate and trigger
healing processes [1]. Delivery of growth factors such as bone
morphogenetic proteins (BMPs), vascular endothelial growth factor
(VEGF) and transforming growth factor beta (TGF-β) can stimulate
cellular adhesion, proliferation and differentiation and form an
attractive therapeutic strategy to promote endogenous repair [2].
Bone morphogenetic proteins, in particular, have been extensively
studied [3,4] and two recombinant human bone morphogenetic
proteins (rhBMPs), rhBMP-2 (Infuse™, Medtronic Sofamor Danek
Inc.) and rhBMP-7 (OP-1™, Stryker Biotech) have been approved by
regulatory bodies for the treatment of non-union bone defects, open
tibial fractures and spinal fusion [5–9]. However, administration of
rhBMPs in orthopaedic applications is complicated by their short
biological half lives, localised action and rapid clearance [10,11].
This has resulted in the clinical use of supraphysiologic concentra-
tions of rhBMPs; currently a dose of ~60 μM BMP-2 is delivered (via
a purified type I collagen matrix) which greatly exceeds the effective
physiological range of 1–30 nM [12].
In 2008 the US Food and Drug Administration (FDA) issued a Public
Health Notification of life threatening complications associated with
rhBMP-2 use; these included complications associated with swelling
of neck and throat tissue in anterior cervical disectomy and fusion
(ACDF) procedures, as described by Perri and associates [13]. Further
to this, a recent critical review has outlined the risks associated with
clinical BMP-2 usage in spinal surgery and highlighted problems
encountered with high doses; an increased risk of malignancy was
indicated with the use of AMPLIFY in posterolateral spinal fusion [14].
Sustained delivery of rhBMP-2 could therefore address the need to mit-
igate these high doses and corresponding health and cost implications
[15]. These implications emphasise the need for strategies to improve
the efficacy and safety of recombinant growth factor delivery.
Control of delivery can be achieved by incorporating the growth
factor into a biomaterial carrier system; this strategy may also
prevent growth factor degradation as the biomaterial vehicle can
provide protection from proteolytic enzymes at the target site [16].
Incorporation strategies include non-covalent mechanisms, such as
physical entrapment, surface adsorption and complexation, or cova-
lent immobilisation on or into the delivery vehicle. Spatio-temporal
control of growth factor delivery has been achieved via encapsulation
within polymeric microparticles as this can provide independent reg-
ulation of duration and availability of soluble factors whilst limiting
dose and undesirable side effects [17]. Additionally, scaffolds formed
from microparticles can host tissue formation whilst maintaining de-
sired local growth factor concentrations [18]. Homo and copolymers
Materials Science and Engineering C xxx (2013) xxx–xxx
⁎ Corresponding author. Tel.: +44 1158231232.
E-mail address: lisa.white@nottingham.ac.uk (L.J. White).
MSC-03881; No of Pages 6
0928-4931/$ – see front matter © 2013 Elsevier B.V. All rights reserved.
http://dx.doi.org/10.1016/j.msec.2013.02.020
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2. of lactide and glycolide have found particular application as con-
trolled delivery vehicles as they are FDA approved for various clinical
functions [19], degrade in vivo into natural products (lactic and
glycolic acid) that are processed by normal metabolic pathways
[20–22] and have tuneable physico-chemical properties [16].
Herein we describe the development of a novel growth factor de-
livery system with kinetics of release faster than scaffold degradation.
This system has the capability to provide both structural support and
the release of growth factors in a controlled spatio-temporal manner,
when injected or inserted into a defect site. The delivery system is
comprised of large (~100 μm) microparticles fabricated from poly
(D,L-lactic -co-glycolic acid) (PLGA) and an ‘in-house’ produced
triblock copolymer, containing poly(ethylene glycol) (PEG): PLGA–
PEG–PLGA [23]. Addition of the triblock copolymer accelerates
release kinetics such that the encapsulated growth factor can be
released prior to microparticle degradation. This localised growth
factor delivery system can also provide a supportive structure and
we anticipate that the size of the microparticles will ensure that the
void filled by the delivery system will confer mechanical support
whilst the pore space between the microparticles contributes to
enhanced cell infiltration and proliferation. We report on the charac-
terisation of the size, morphology and release kinetics of this novel de-
livery system with initial in vitro release studies undertaken using
lysozyme, a suitable model protein for rh-BMP2. This system provides
customisable, localised and controlled delivery with adjustable release
profiles, which we anticipate will improve the efficacy and safety of
therapeutic growth factor delivery.
2. Materials and methods
2.1. Materials
Poly(vinyl alcohol) (PVA, molecular weight: 13,000–23,000 Da,
87–89% hydrolysed), lysozyme from chicken egg white, human
serum albumin (HSA), polyethylene glycol (PEG, molecular weight:
1500 Da), stannous octoate/Tin(II) 2-ethylhexanoate (Sn(Oct)2),
dimethyl sulphoxide (DMSO), sodium hydroxide (NaOH), sodium
dodecyl sulphate (SDS) and Micrococcus lysodeikticus were obtained
from Sigma–Aldrich, Dorset, UK. Poly (D,L-lactide-co-glycolide)
(PLGA) polymers with lactide:glycolide ratios of 50:50 and 85:15
(PLGA 85:15 DLG 4A 56 kDa and PLGA 50:50 DLG 4.5A 59 kDa)
were purchased from Surmodics, Birmingham, USA. D,L-Lactide and
glycolide monomers were obtained from Lancaster Synthesis, Ward
Hill, MA, USA and PURAC, Gorinchem, Netherlands, respectively. The
Pierce Micro bicinchoninic acid (BCA) protein assay kit and HPLC
grade solvents (dichloromethane (DCM) and acetone) were pur-
chased from Fisher Scientific UK Ltd, Loughborough, UK. Recombinant
human BMP-2 (BMP-2) was purchased from Professor Walter Sebald
(University of Wurzburg, Germany).
2.2. Methods
2.2.1. Triblock copolymer synthesis, purification and characterisation
Synthesis of the PLGA–PEG–PLGA triblock copolymer occurred by
ring opening polymerisation of the D,L-lactide and glycolide monomers
in the presence of PEG and the Sn(Oct)2 catalyst under a dry nitrogen
atmosphere, as previously described [24,25]. Prior to polymerisation,
PEG was first dried in the reaction vessel under vacuum and stirring
at 120 °C for 3 h. The temperature was then raised to 150 °C, the
monomers added to the vessel and the reaction mixture allowed to
equilibrate under a dry nitrogen atmosphere. After 30 min, the
Sn(Oct)2 was added and the reaction was allowed to proceed for 8 h.
The resultant copolymer was dissolved and precipitated in water in
order to remove un-reacted monomers. The triblock copolymer was
then dried under vacuum to remove residual water and stored at
−20 °C until required.
Nuclear Magnetic Resonance (NMR) analysis of the copolymer was
undertaken using a Bruker DPX-300 Spectrometer (300 MHz) with
deuterated choloroform (CDCl3) as the solvent. A tetramethylsilane
(TMS) signal was taken as the zero chemical shift. As described by
Hou et al. [25], the composition of the copolymer was determined by
1
H NMR by integrating the signals pertaining to each monomer, i.e.
peaks from CH2 of ethylene glycol and glycolide and CH and CH3
from D,L-lactide. The molecular weight and polydispersity index of the
copolymer was determined using Gel Permeation Chromatography
(PL-GPC 120, Polymer Labs) with differential refractometer detection.
Tetrahydrofuran (THF) was employed as an eluent, with two columns
(30 cm, PolarGel-M) in series calibrated against polystyrene standards.
Characteristics of the triblock copolymer are presented in Table 1.
2.2.2. Microparticle preparation
Poly (D,L-lactide-co-glycolide) microparticles were formed using a
water-in-oil-in-water (w/o/w) emulsion method as previously de-
scribed [23]. Briefly, an aqueous solution of human serum albumin
(HSA) and lysozyme (or BMP-2) was added to a solution of PLGA and
PLGA–PEG–PLGA in dichloromethane. These phases were homogenised
for two minutes at 4000 rpm in a Silverson L5M homogeniser
(Silverson Machines, UK) to form the water-in-oil emulsion. This prima-
ry emulsion was transferred to 200 ml 0.3% (w/v) PVA solution and was
homogenised for a second time at 2000 rpm for two minutes. The resul-
tant double emulsion was stirred at 300 rpm on a Variomag 15-way
magnetic stirrer for a minimum of 4 h to facilitate DCM evaporation.
Microparticles were then filtered, washed and lyophilized (Edwards
Modulyo, IMA Edwards, UK) until dry.
The triblock copolymer was added to PLGA to provide weight per-
centages of 0, 10, and 30% (w/w) of the 1 g total mass (in 5 ml DCM).
Lysozyme (or BMP-2) and HSA were prepared at a ratio of 1:9 for a 1%
w/w loading in the microparticles. That is, for 1 g of polymer, 1 mg of
lysozyme (or BMP-2) and 9 mg of HSA were dissolved in 100 μl dis-
tilled water. Control particles were manufactured without protein,
using 100 μl distilled water in the primary emulsion.
Lysozyme was selected as an appropriate model for BMP-2 in this
work, due to the similarity of isoelectric points (lysozyme: 9, BMP-2:
>8.5) and molecular weights (lysozyme: 14 kD, BMP-2: 26 kD) [26].
2.2.3. Microparticle characterisation — size, morphology and protein
entrapment
Microparticles (50 mg/ml in double deionised water) were sized
using a laser diffraction method (Coulter LS230, Beckman Coulter, UK)
with agitation to prevent particle settling. To assess the morphology
of the microparticles, thin layers of freeze dried microparticles were
adhered to an adhesive stub and gold sputter coated for four minutes
at 30 mA (Balzers SCD 030 gold sputter coater, Balzers, Liechenstein).
Microparticles were imaged using a JSM 6060LV Scanning Electron
Microscope (SEM) (JEOL, Welwyn Garden City, UK) with the accelerat-
ing voltage set to 10 kV.
Measurement of the encapsulation efficiency of protein within the
microparticles was undertaken as previously reported [23]; the meth-
od was a modification of techniques proposed by Sah [27] and Morita
et al. [28]. Briefly, 10 mg of PLGA microspheres were added to 750 μl
DMSO and left at room temperature for one hour; 2150 μl of 0.02%
(w/v) SDS in 0.2 M NaOH was then added for further one hour
incubation. The Micro BCA protein assay kit was used to ascertain
the total protein content and compared against a standard curve of
HSA/lysozyme conducted at the same time. Sample (150 μl) and
Table 1
Triblock copolymer characteristics.
MN
a
% mole lactidea
% mole glycolidea
MN
b
MW
b
PDIb
1706–1500–1706 71 29 2442 4022 1.65
a
determined by 1
H NMR; b
determined by GPC.
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3. BCA working reagent (150 μl) were mixed and incubated for 2 h at
37 °C and the absorbance at 562 nm measured using a plate reader
(Infinite M200, Tecan UK Ltd., Reading, UK).
2.2.4. Protein release from microparticles
Aliquots (100 mg) of the microparticles (triplicate samples from
each batch) were suspended in 3 ml phosphate buffered saline
(PBS, pH 7.4); samples were gently rocked on a 3-dimensional shaker
(Gyrotwister, Fisher Scientific UK Ltd) at 5 rpm in a humidified incu-
bator at 37 °C. At defined time intervals, the PBS was removed from
the microparticles and replaced with 3 ml fresh PBS; all liquid
above the particles was collected without removing particles. The re-
moved supernatants were stored frozen until required and were then
assayed for total protein content using the Micro BCA assay kit with a
standard curve of HSA/lysozyme (0–40 μg/ml).
The activity of lysozyme in collected supernatants was measured
using the method described by Bezemer et al. [29] and Sohier et al.
[30] with slight modifications. The change in turbidity, that occurred
in an M. lysodeikticus solution when lysozyme lysed the 1,4 glycosidic
bond within the cell wall of the bacteria, was measured. M. lysodeikticus
suspension (100 μl of 2.3 mg/ml) was added to 150 μl of the release
supernatant. The change in turbidity over a one minute period was
determined by evaluation of the absorbance at 450 nm, measured
using a plate reader (Infinite M200, Tecan UK Ltd., Reading, UK). Absor-
bance values were correlated to a standard curve of HSA/lysozyme
(0–10 μg/ml) in order to determine the activity of lysozyme.
3. Results
3.1. Size, morphology and protein entrapment
Polymer microparticles with encapsulated HSA and lysozyme
were fabricated using the double emulsion method described above.
Six different formulations were manufactured by varying the
lactide:glycolide ratio (50:50 and 85:15) and by incorporating three
different percentage weights of the PLGA–PEG–PLGA triblock copoly-
mer (0%, 10% and 30%). The size, morphology and in vitro release pro-
file of microparticles from each formulation were characterised.
Representative images of microparticle morphology are shown in
Fig. 1B and D; microparticles fabricated from each of the six formula-
tions were spherical with smooth, non-porous surfaces. Incorporation
of protein did not affect the morphology or size of the microparticles,
as shown in the representative example of 85:15 PLGA with 10% w/w
PLGA–PEG–PLGA (Fig. 1). The mean diameter of the microparticles
formed without protein (Fig. 1A) was 90.8 μm (SD±24.7) and the
mean diameter of microparticles containing protein (Fig. 1C) was
94.0 μm (SD±27.6). Multiple batches formed under the same condi-
tions demonstrated repeatable distributions; in each case, the double
emulsion process produced a bell curve size distribution.
Incorporation of HSA with lysozyme improved overall protein en-
capsulation within microparticles. Initial studies with lysozyme pro-
duced microparticles with low encapsulation efficiencies (average of
13%) whereas the co-addition of HSA with lysozyme increased the
encapsulation to an average of 62%.
3.2. Protein release
Cumulative protein release from microparticles of the six formula-
tions is shown in Fig. 2. Formulations with no triblock (0% w/w
PLGA–PEG–PLGA) exhibited typical triphasic release profiles compris-
ing: (i) an initial burst occurring during the first twenty-four hours,
(ii) a lag phase and (iii) a second burst at 15 days (Fig. 2A). For 50:50
PLGA formulations the addition of PLGA–PEG–PLGA modified the
release profile. A quasi zero order release profile was obtained for
50:50 PLGA 10% w/w PLGA–PEG–PLGA during the first four weeks of
release. Incorporation of 30% w/w PLGA–PEG–PLGA with 50:50 PLGA
Fig. 1. Morphology of microparticles formulated from 85:15 PLGA with 10% w/w PLGA–PEG–PLGA, with size distributions and SEM images of the blank (no protein) microparticles
(A and B) and protein loaded (1% w/w) microparticles (C and D).
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4. increased protein release during the burst phase, with more than 60%
of protein released in the first twenty-four hours. The burst phase
from particles made with 85:15 PLGA was also modulated by the addi-
tion of 30% w/w PLGA–PEG–PLGA (Fig. 2B) and in both 50:50 and
85:15 formulations, more than 80% of the total protein was released
(Fig. 2A and B). A triphasic release profile was exhibited by 85:15
PLGA with both 0% and 10% w/w PLGA–PEG–PLGA (Fig. 2B).
Focusing in on the first 14 days of release (Fig. 2C) shows that
sustained release with a zero order profile was obtained from the
85:15 PLGA 30% w/w PLGA–PEG–PLGA formulation over this
timeframe. This was in contrast to both the rapid release (approxi-
mately three days) of protein from the 50:50 PLGA 30% w/w PLGA–
PEG–PLGA and the slow, steady release from 50:50 PLGA 10% w/w
PLGA–PEG–PLGA.
Analysis of the average daily release of protein from the various
formulations (Table 2) reveals that the 50:50 10% w/w PLGA–PEG–
PLGA provides a sustained total protein release of 1.8 μg per day in
the fifth week, corresponding to 180 ng/day of lysozyme. Lysozyme
release from the 85:15 PLGA 30% w/w PLGA–PEG–PLGA formulation
is similar (150 ng/day) whereas initial rapid protein release is evident
in the higher burst values and reduced lysozyme release (40 ng/day)
in the fifth week. The activity of released lysozyme was determined
using the M. lysodeikticus assay which confirmed that the released
lysozyme was active at all timepoints (data not shown).
3.3. Protein release from blended batches
Additional batches of microparticles were fabricated from 50:50
PLGA with 0%, 10%, 20% and 30% w/w PLGA–PEG–PLGA. Three batches
of each formulation were prepared and the size, morphology and in
vitro release profiles were characterised. Representative images and
size data are provided in Supplementary Fig. 1. The addition of the
triblock copolymer reduced the mean diameter of microparticles
and produced more narrow size distributions (as can be seen in the
left column of Supplementary Fig. 1). The mean diameter of 50:50
PLGA microparticles without triblock copolymer was 77.5 μm (SD±
20.4) (Supplementary Fig. 1A) and the addition of 30% w/w PLGA–
PEG–PLGA reduced the mean diameter to 64.6 μm (SD±15.6)
(Supplementary Fig. 1G). Multiple batches formed under the same
conditions demonstrated repeatable size distributions.
Microparticles from the 50:50 PLGA 10% and 30% w/w PLGA–PEG–
PLGA formulations were mixed in a 1:1 ratio to assess the effect of
batch blending upon the release profile. Mixing a slow releasing
formulation (containing 10% w/w PLGA–PEG–PLGA) with a more
rapid formulation (containing 30% w/w PLGA–PEG–PLGA) mitigated
the burst effect and provided a more overall sustained release (Fig. 3).
Fig. 2. Cumulative release of HSA/lysozyme (1% w/w) from microparticles formulated
from (A) 50:50 PLGA with 30% (■), 10% (▼) and 0% (●) of a PLGA–PEG–PLGA copoly-
mer and (B) 85:15 PLGA with 30% (□), 10% (▽) and 0% (○) of a PLGA–PEG–PLGA co-
polymer. Release profiles of 50:50 PLGA with 30% (■), 85:15 PLGA with 30% (□) and
50:50 PLGA with 10% (▼) over 14 days are shown in (C). Data represent mean±SD
for n=3.
Table 2
Average daily release of HSA/lysozyme (1% w/w) from three microparticle formulations;
microparticle mass=100 mgs.
Days Average daily release HSA/lysozyme (μg)
50:50 PLGA
10% PLGA–PEG–PLGA
85:15 PLGA
30% PLGA–PEG–PLGA
50:50 PLGA
30% PLGA–PEG–PLGA
1–7 27.1 104.5 116.3
8–14 9.4 8.2 0.3
15–21 10.7 4.6 0.0
22–28 10.9 2.0 0.2
29–35 1.8 1.5 0.4
Fig. 3. Cumulative release of HSA/lysozyme (1% w/w) from microparticles formulated
from 50:50 PLGA with 30% (■) and 10% (▼) of a PLGA–PEG–PLGA copolymer and
a mixture of these two formulations (1:1 ratio) (○). Data represent mean±SD for
n=9.
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5. 3.4. Comparative release of lysozyme and BMP-2
Microparticles containing HSA/BMP-2 were fabricated from 50:50
PLGA with 10% w/w PLGA–PEG–PLGA. The release profiles of HSA/
BMP-2 and HSA/lysozyme from this formulation were in accord
(Fig. 4) indicating that lysozyme was a suitable model for BMP-2 in
this work.
4. Discussion
Successful scaffolds for growth factor delivery must provide both
structural support and the controlled spatio-temporal release of
growth factors required to host tissue formation. Although spatio-
temporal control of growth factor and protein delivery has been
achieved previously via encapsulation within polymeric microparti-
cles, providing appropriate release profiles for the protein or growth
factor of interest remains a challenge [16]. Protein release profiles
consisting of an initial burst and subsequent slow and incomplete
release (that does not match polymer degradation rate) have often
been observed and are deemed inappropriate for clinical use
[31–35]. In addition, incomplete release has been linked to protein in-
stability issues that may occur during formulation, storage and
release of protein from loaded microspheres [31,36,37].
Thus, the objective of this study was to develop a novel delivery
system, capable of providing structural support, with controlled and
tuneable release kinetics. Our strategy was to include a hydrophilic
polymer within microparticle formulations. Protein delivery systems
based on hydrophilic–hydrophobic block copolymers have been
utilised as controlled release systems since the degradation rate,
and hence release kinetics, can be modulated by adjusting copolymer
compositions [24,25,38–40]. In particular, copolymers of PLGA and
PEG are more compatible with proteins, reduce protein adsorption
and favour homogeneous protein distribution within the polymer
matrix as well as increasing water uptake in microspheres [41].
In this study we show, for the first time, that protein release
kinetics can be altered and controlled by the inclusion of the PLGA–
PEG–PLGA triblock copolymer and that these release kinetics can be
decoupled from polymer degradation rate. Subsequently, we have
developed a novel delivery system with capability for delivering
multiple proteins with different and independent release kinetics.
The addition of the PLGA–PEG–PLGA copolymer modulated pro-
tein release by accelerating water ingress into the matrix. Increased
water uptake can lead to a combined mechanism of swelling and
structural erosion, which may produce a hydrogel-like structure in
the microparticles, facilitating controlled release of the protein.
Addition of 10% w/w PLGA–PEG–PLGA to 50:50 PLGA modulated the
typical triphasic profile and instead produced a quasi zero order re-
lease profile over a four week duration. For both 85:15 and 50:50
PLGA, almost complete protein release was achieved with 30% w/w
triblock copolymer included in the formulation; sustained release
over ten days was achieved with 85:15 PLGA 30% w/w PLGA–PEG–
PLGA. These three formulations provide diverse and independent
release profiles that could be used for the delivery of growth factors.
For example, the rapid release from 50:50 PLGA 30% w/w PLGA–
PEG–PLGA may be suitable for delivering VEGF (to stimulate immedi-
ate blood vessel formation) whereas the more sustained release of
the 85:15 PLGA 30% w/w PLGA–PEG–PLGA formulation may be
appropriate for PDGF (to promote maturation of the newly formed
blood vessels). The sustained duration of release of HSA/lysozyme
from 50:50 PLGA 10% w/w PLGA–PEG–PLGA renders this an ideal
delivery vehicle as BMP-2 is typically expressed throughout the
fracture healing process until remodelling occurs (21 days) [42].
The kinetics of release of rhBMP-2 have recently been shown to
have profound effects on bone regeneration; an initial burst followed
by sustained release significantly improved healing and bone regen-
eration in rat femurs [43,44]. Furthermore, the average daily release
values of lysozyme, reported in Table 2, correspond to values reported
for BMP-2 delivery for effective bone repair [45–47].
For microparticles fabricated with 50:50 PLGA, the onset of poly-
mer degradation occurred at approximately day 15. This resulted in
a second burst of protein release in the triphasic release profile of
the 50:50 PLGA 0% w/w PLGA–PEG–PLGA microparticles. With
50:50 PLGA 30% w/w PLGA–PEG–PLGA, we were able to achieve al-
most complete release in five days, i.e. release occurred prior to the
onset of polymer degradation. Analogous behaviour was observed
with 85:15 PLGA 30% w/w PLGA–PEG–PLGA with kinetics of release
decoupled from polymer degradation rate. Although incomplete re-
lease was observed for the other four formulations, total protein re-
lease was higher in the 10% w/w PLGA–PEG–PLGA formulations
compared to those fabricated from PLGA alone. This suggests that
the presence of the triblock copolymer may provide a means to ad-
dress this phenomenon.
The kinetics of release were also modified by mixing microparti-
cles of two different formulations providing another route, not previ-
ously reported, for controlling release kinetics. Mixing a slow
releasing formulation (containing 10% w/w PLGA–PEG–PLGA) with
a more rapid formulation (containing 30% w/w PLGA–PEG–PLGA)
mitigated the burst effect and provided a sustained release profile.
In this study, protein entrapment occurred via a double emulsion
process. The creation of a water/organic solvent interface during the
primary emulsion stage has been well documented in the literature
as being primarily responsible for protein instability during emulsifi-
cation [36,48,49]. However, in our study, good encapsulation efficien-
cies and retention of protein activity was observed. This is likely to be
due to our approach of adding an excipient, (HSA in our case) to the
inner aqueous phase that can compete with the therapeutic protein
at the water/organic solvent interface. We chose HSA for our work
based on publications where serum albumins in particular have
been shown to limit emulsification induced aggregation of several
different proteins, including lysozyme, ovalbumin and recombinant
human erythropoietin [49–51]. HSA was co-encapsulated with
lysozyme (in a ratio of 9:1 HSA:lysozyme) which did indeed appear
to shield lysozyme from interface induced aggregation evidenced by
retention of enzymatic activity but also increased the encapsulation
efficiency. A similar effect was observed by Srinivasan et al. [51]
with rat serum albumin.
Lysozyme was used as a model protein in this work due to the
breadth of detailed characterisation of this protein present in published
literature [48,49,52–56] coupled with the ability to easily measure
biological activity [57]. An additional motivation for using lysozyme
was that it is considered an appropriate model for BMP-2 due to the sim-
ilarity of isoelectric points (lysozyme: 9, BMP-2: >8.5) and molecular
Fig. 4. Cumulative release of HSA/lysozyme (1% w/w) (▼) and HSA/BMP-2 (1% w/w)
(▽) from microparticles formulated from 50:50 PLGA with 10% w/w PLGA–PEG–
PLGA copolymer. Data represent mean±SD for n=3.
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6. weights (lysozyme: 14 kD, BMP-2: 26 kD) [26]. The consistency
shown in the release profiles of HSA/BMP-2 and HSA/lysozyme con-
firms this. Previous work from our group with microspheres formu-
lated as described above verified the biological activity of released
BMP-2 from PLGA microparticles formulated with PLGA–PEG–PLGA
[23].
5. Conclusions
A new delivery system, capable of providing controlled release ki-
netics, has been developed from PLGA microparticles. The inclusion of
a hydrophilic PLGA–PEG–PLGA triblock copolymer altered release ki-
netics such that they were decoupled from polymer degradation. A
quasi zero order release profile over four weeks was produced using
10% w/w PLGA–PEG–PLGA with 50:50 PLGA whereas complete and
sustained release over was achieved over ten days using 30% w/w
PLGA–PEG–PLGA with 85:15 PLGA and over four days using 30% w/w
PLGA–PEG–PLGA with 50:50 PLGA. These three formulations are prom-
ising candidates for delivery of growth factors such as BMP-2, PDGF
and VEGF. Release profiles were also modified by mixing microparticles
of two different formulations providing another route, not previously
reported, for controlling release kinetics. Good encapsulation efficien-
cies and retention of protein activity were achieved via the co-
addition of HSA with lysozyme. Comparative release of HSA/lysozyme
and HSA/BMP-2 showed excellent agreement confirming that lyso-
zyme is an appropriate model for BMP-2.
Supplementary data to this article can be found online at http://
dx.doi.org/10.1016/j.msec.2013.02.020.
Acknowledgements
This work was supported by the Biotechnology and Biological Sci-
ences Research Council (BBSRC), UK through a LoLa grant, grant no.
BB/G010617/1. The authors express their appreciation to Miss
Natasha Birkin for her assistance with characterisation of the triblock
copolymer.
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6 L.J. White et al. / Materials Science and Engineering C xxx (2013) xxx–xxx
Please cite this article as: L.J. White, et al., Mater. Sci. Eng., C (2013), http://dx.doi.org/10.1016/j.msec.2013.02.020