Accelerating protein release from microparticles for regenerative

medicine applications

Lisa J. White,1 Giles T. S. Kirby,1 Helen C. Cox,1 Roozbeh Qodratnama,1 Omar

Qutachi,1 Felicity R. A. J. Rose,1 Kevin M. Shakesheff1

1 School of Pharmacy, University of Nottingham, Nottingham, NG7 2RD, UK

ABSTRACT

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 release 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.

Keywords: Poly (DL-lactic acid-co-glycolic acid) (PLGA), microparticles, microspheres,

controlled release, double emulsion solvent evaporation, growth factor delivery

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 (InfuseTM, Medtronic

Sofamor Danek Inc.) and rhBMP-7 (OP-1TM, 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, localized action and rapid clearance [10, 11].

This has resulted in the clinical use of supraphysiologic concentrations of rhBMPs; currently

a dose of ~60µM 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 mitigate these high doses and corresponding

health and cost implications [15]. These implications emphasize 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 covalent 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 regulation 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 desired local growth factor concentrations [18]. Homo and copolymers of lactide

and glycolide have found particular application as controlled 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 delivery 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 (DL-lactic acid-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 localized 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

characterisation of the size, morphology and release kinetics of this novel delivery 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: 13000 – 23000 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 (DL-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. DL-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 purchased 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

2.2.1Triblock copolymer synthesis, purification and characterisation

Synthesis of the PLGA-PEG-PLGA triblock copolymer occurred by ring opening

polymerisation of the DL-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 hours. 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 minutes, the Sn(Oct)2 was added and the reaction allowed to proceed for 8 hours. 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 1H NMR

by integrating the signals pertaining to each monomer, i.e. peaks from CH2 of ethylene glycol

and glycolide and CH and CH3 from DL-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 (DL-lactide-co-glycolide) microparticles were formed using a water-in-oil-in-water

(w/o/w) emulsion method as previously described [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 4,000 rpm in a Silverson L5M homogeniser (Silverson Machines, UK) to form the

water-in-oil emulsion. This primary emulsion was transferred to 200 ml 0.3% (w/v) PVA

solution and was homogenised for a second time at 2,000 rpm for two minutes. The

resultant double emulsion was stirred at 300 rpm on a Variomag 15-way magnetic stirrer for

a minimum of 4 hours 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 percentages 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 distilled

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 particles 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 accelerating voltage set to 10 kV.

Measurement of the encapsulation efficiency of protein within the microparticles was

undertaken as previously reported [23]; the method 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; 2,150 µL of 0.02% (w/v)

SDS in 0.2M 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 BCA working

reagent (150 µL) were mixed and incubated for 2 hours 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

incubator 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 removed 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 a Micrococcus lysodeikticus solution when lysozyme lysed the 1,4

glycosidic bond within the cell wall of the bacteria, was measured. Micrococcus

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). Absorbance 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 copolymer (0%, 10% and 30%). The

size, morphology and in vitro release profile of microparticles from each formulation were

characterised.

Representative images of microparticle morphology are shown in Figure 1 B and D;

microparticles fabricated from each of the six formulations 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 (Figure 1). The mean diameter of the microparticles formed without protein

(Figure 1A) was 90.8 μm (SD ± 24.7) and the mean diameter of microparticles containing

protein (Figure 1C) was 94.0 μm (SD ± 27.6). Multiple batches formed under the same

conditions demonstrated repeatable distributions; in each case, the double emulsion process

produced a bell curve size distribution.

Incorporation of HSA with lysozyme improved overall protein encapsulation within

microparticles. Initial studies with lysozyme produced 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 formulations is shown in Figure

2. Formulations with no triblock (0% w/w PLGA-PEG-PLGA) exhibited typical triphasic release

profiles comprising: (i) an initial burst occurring during the first twenty-four hours, (ii) a lag

phase and (iii) a second burst at 15 days (Figure 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 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 addition of 30% w/w PLGA-PEG-PLGA (Figure 2B) and in both 50:50 and 85:15

formulations, more than 80% of the total protein was released (Figure 2A and B). A triphasic

release profile was exhibited by 85:15 PLGA with both 0% and 10 % w/w PLGA-PEG-PLGA

(Figure 2B).

Focusing in on the first 14 days of release (Figure 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 (approximately

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 Micrococcus 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 Figure 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 Figure 1). The mean diameter of 50:50

PLGA microparticles without triblock copolymer was 77.5 μm (SD ± 20.4) (Supplementary

Figure 1A) and the addition of 30% w/w PLGA-PEG-PLGA reduced the mean diameter to

64.6 μm (SD ± 15.6) (Supplementary Figure 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 (Figure 3).

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 (Figure 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 microparticles, 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 instability 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 protein 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 release 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 immediate 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

remodeling 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 regeneration 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 polymer 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 almost 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 release was observed for

the other four formulations, total protein release 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 address this phenomenon.

The kinetics of release were also modified by mixing microparticles of two different

formulations providing another route, not previously 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

emulsification [36, 48, 49]. However, in our study, good encapsulation efficiencies 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 similarity of isoelectric

points (lysozyme: 9, BMP-2: >8.5) and molecular weights (lysozyme: 14 kD, BMP-2: 26 kD)

[26]. The consistency shown in the release profiles of HSA/BMP-2 and HSA/lysozyme

confirms this. Previous work from our group with microspheres formulated 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 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 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

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. Good

encapsulation efficiencies 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 lysozyme is an appropriate model for BMP-2.

Acknowledgements

This work was supported by the Biotechnology and Biological Sciences 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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Figure and Table Captions

Figure 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).

Figure 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 copolymer and

(B) 85:15 PLGA with 30% (), 10% () and 0% () of a PLGA-PEG-PLGA copolymer.

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.

Figure 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.

Figure 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.

Supplementary Figure 1: Size distributions (left) and SEM images (right) of protein loaded

(1% w/w) microparticles formulated from 50:50 PLGA and a PLGA-PEG-PLGA copolymer.

Concentrations of the PLGA-PEG-PLGA component were varied: 0% (A and B), 10% (C

and D), 20% (E and F) and 30% (G and H).

Table 1: Triblock copolymer characteristics

adetermined by

1H NMR;

bdetermined by GPC

Table 2: Average daily release of HSA/lysozyme (1% w/w) from three microparticle

formulations; microparticle mass = 100 mgs.

Figure 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).

Figure 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 copolymer and

(B) 85:15 PLGA with 30% (), 10% () and 0% () of a PLGA-PEG-PLGA copolymer.

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.

Figure 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.

Figure 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.

Supplementary Figure 1: Size distributions (left) and SEM images (right) of protein loaded

(1% w/w) microparticles formulated from 50:50 PLGA and a PLGA-PEG-PLGA copolymer.

Concentrations of the PLGA-PEG-PLGA component were varied: 0% (A and B), 10% (C

and D), 20% (E and F) and 30% (G and H).

Table 1: Triblock copolymer characteristics

adetermined by

1H NMR;

bdetermined by GPC

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