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Bioresorbable and bioactive polymer/Bioglass1 composites withtailored pore structure for tissue engineering applications

Aldo R. Boccaccinia,*, Veronique Maquetb

aDepartment of Materials and Centre for Tissue Engineering and Regenerative Medicine Imperial College London, London SW7 2BP, UK bCentre for Education and Research on Macromolecules (CERM), Interfacultary Centre for Biomaterials, University of Liege,

B-4000 Lie ge, Belgium

Received 17 December 2002; accepted 16 May 2003

Abstract

An overview about the development of porous bioresorbable composite materials for applications as scaffolds in tissue engi-

neering is presented. A thermally induced phase separation method was developed to fabricate porous foam-like structures of

poly(lactide-co-glycolide) (PLGA) containing bioactive glass particle additions (up to 50 wt.%) and exhibiting well-defined, orien-

ted and interconnected porosity. The in vitro bioactivity and the degradability of the composite foams were investigated in contact

with phosphate buffer saline (PBS). Weight loss, water absorption and molecular weight measurements were used to monitor the

polymer degradation after incubation periods of up to 7 weeks in PBS. It was found that the presence of bioactive glass retards the

polymer degradation rate for the time period investigated. The present results show a way of controlling the in vitro degradation

behaviour of PLGA porous composite scaffolds by tailoring the concentration of bioactive glass.

# 2003 Elsevier Ltd. All rights reserved.

Keywords: Bioresorbable polymers

1. Introduction

Tissue engineering presents an alternative approach to

the repair and regeneration of damaged human tissue,

avoiding the need for a permanent implant. The under-

lying principle involved is the regeneration of living tis-

sue, where a loss or damage has occurred as a result of

injury or disease [1, 2]. Tissue engineering can be there-

fore simply defined as the ‘‘science of persuading the

body to heal by its intrinsic repair mechanisms’’ [3]. The

scientific challenge encompasses understanding the cells

themselves, their mass transport requirements and bio-

logical environment as well as the development of sui-table scaffold materials, usually porous, that act as

templates for cell adhesion, growth and proliferation.

The ultimate goal is to return full biological and

mechanical functionality to a damaged tissue or organ.

Several physicochemical and biological requirements

have to be fulfilled by the scaffold, depending on

the particular tissue under consideration, which are

closely dependent on the scaffold porosity and porous

structure.

In a recent review paper, Hutmacher has summarized

the requirements for scaffolds intended for musculoske-

letal tissue engineering [4]. Ideally, scaffolds for this

application should have the following characteristics: (i)

three dimensional and highly porous with an inter-

connected pore network for cell growth and flow trans-

port of nutrients and metabolic waste, (ii) biocompatible

and bioresorbable with a controllable degradation and

resorption rate to match cell/tissue growth in vitro and/

or in vivo, (iii) suitable surface chemistry for cell

attachment, proliferation, and differentiation and (iv)mechanical properties to match those of the tissues at

the site of implantation. In particular for bone tissue

applications, it has also been reported [5], that a suitable

scaffold should bond to the host tissue without the for-

mation of scar tissue, i.e. it should exhibit bioactivity

and osteoconductivity. Moreover a suitable scaffold

should be made from versatile processing techniques

that can produce irregular, usually complex, shapes to

match that of the defect in the tissue of the patient.

A wide variety of both natural and synthetic materi-

als, or a combination of them, are being investigated for

0266-3538/$ - see front matter # 2003 Elsevier Ltd. All rights reserved.

doi:10.1016/S0266-3538(03)00275-6

Composites Science and Technology 63 (2003) 2417–2429

www.elsevier.com/locate/compscitech

* Corresponding author. Tel.: +44-207-5946731; fax: +44-207-

5843194.

E-mail address: [email protected] (A.R. Boccaccini).

http://www.elsevier.com/locate/compscitech/a4.3d

mailto:[email protected]

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design and construction of scaffold for tissue engineering.

These include naturally occurring polymers, e.g. hydro-

gels like gelatin, agar, fibrin or collagen [6–9], synthetic

bioresorbable polymers, e.g. poly(lactide acid), poly(-

glycolic acid), polycaprolactone and poly(propylene

fumarate) [3,10–15] as well as bioactive porous cera-

mics, e.g. synthetic foam-like bioactive glass and cal-cium phosphate structures [16–20] and naturally

occurring ceramics, such as coral [21]. Also metallic

foams are being considered for tissue engineering

scaffolds [22].

Bioresorbable synthetic polymers have attracted

increasing attention for their use as tissue engineering

scaffolds in the last ten years [3,10–15,23]. Many prac-

tical advantages arise when using synthetic scaffolds

because precise control of material composition and

micro- and macrostructure, including porosity, is possi-

ble. This allows adequate control of scaffold properties,

thus creating optimal conditions for cell survival, pro-

liferation, and subsequent tissue formation. Polyesterssuch as poly(lactic acid) (PLA), poly(glycolic acid)

(PGA) and poly(lactic acid-co-glycolic acid) (PLGA)

are being mainly considered for scaffold applications

[3,10–15,23]. These polymers have already demon-

strated promising results in clinical use, for example as

resorbable surgical sutures and meshes or in drug deliv-

ery systems, and they have the United States Federal

Food and Drug Administration (F.D.A.) approval for

clinical use. The specific use of synthetic bioresorbable

polymers in tissue engineering applications has been

reviewed by Hutmacher [4], Maquet and Jerome [24]

and Agrawal et al. [23].As mentioned above, porosity and pore structure are

key parameters determining the properties and the

applicability of scaffolds for tissue engineering. Fig. 1

shows a summary of the different functions related to

the pore structure in a tissue engineering scaffold. In

general, scaffold porosity, pore morphology and pore

orientation must be tailored to the particular tissue

under consideration.

Porous polymeric tissue engineering scaffolds of

3-dimensional (3-D) structure have been prepared by

numerous techniques, including solid–liquid and liquid–

liquid phase separation [25–31], solution casting [32], gel

casting [33], gas saturation [34], gas foaming [35,36],

fibre bonding, fabric forming and related textile based

processes [37–40], sintering of polymeric microspheres

[41], combined solvent casting and extrusion [42], emul-

sion freeze-drying [43], several rapid prototyping tech-niques, e.g. 3-D printing and fused deposition modelling

[4,6,44–48] and various solvent casting/particulate

leaching methods using different porogen additives [49–

55]. Overviews on fabrication methods of 3-D porous

polymeric scaffolds for tissue engineering are given by

Hutmacher [4], Agrawal et al. [23], Yang et al. [56] and

Thomson et al. [57]. In recent studies, advanced meth-

ods have been developed for the optimal designed of

porous scaffolds based on solid free form manufactur-

ing and conventional sponge scaffold fabrication [58].

Selected ceramics, such as hydroxyapatite (HA), tri-

calcium phosphate (TCP) and some compositions of

silicate and phosphate glasses and glass–ceramics, reactwith physiological fluids and form tenacious bonds to

hard and soft tissues through cellular activity [2,5].

These materials are therefore known as ‘‘bioactive’’ [59].

If biodegradability and bioactivity are to be combined

in an optimised tissue engineering scaffold for bone tis-

sue engineering, then the design of composite materials

offers an exceptional opportunity: by combining bior-

esorbable polymers and bioactive ceramic phases scaf-

folds with tailored physical, biological and mechanical

properties can be produced. Moreover the addition of a

ceramic or glass phase to a biodegradable polymer may

be exploited to favourable alter the in vitro and in vivopolymer degradation behaviour, which, in polylactides,

is strongly acidic. It has been proposed that bioactive

glass particles used as inclusions or coatings in biode-

gradable polylactides should lead to the rapid exchange

of protons in water for alkali in the glass which should

then provide a pH buffering effect at the polymer sur-

face, thus preventing acceleration of polymer degrada-

tion [60]. Moreover, as discussed elsewhere [61],

dissolution of the bioactive glass should result in

nucleation and growth of a crystalline HA layer on the

surface of the polymer scaffold, which should further

affect the polymer degradation behaviour in addition to

provide the required osteoconductivity. Thus, a combi-nation of porous bioresorbable polymers and bioactive

ceramic or glass phases is expected, for the following

reasons, to result in promising composite scaffolds for

(bone) tissue engineering [62]: (i) a better cell seeding

and growth environment can be achieved because of the

good osteoconductivity properties provided by the

bioactive phase, (ii) the acidic degradation by-products

from polyesters may be buffered, and (iii) the mechan-

ical properties may be improved by using the traditional

composite approach (inclusion of a stiffer particulate

ceramic phase in the polymeric matrix).Fig. 1. Schematic diagram showing the different functions of a tissue

engineering scaffold depending on its porosity and pore structure.

2418 A.R. Boccaccini, V. Maquet/ Composites Science and Technology 63 (2003) 2417–2429



In a broader sense, the microstructure of composite

materials, including porosity and pore structure, can be

engineered in such a way that the resorption rate of the

composite scaffold in the body can be designed to match

the formation rate of new tissue. In this regard, also

functionally graded materials with variable porosity and

graded bioactive phase content have been developed[63] in order to mimic the defect site in terms of biolo-

gical, mechanical and morphological properties.

A review covering the fabrication, properties and

applications of bioresorbable and bioactive porous

composites for bone tissue engineering scaffolds have

been presented elsewhere [64]. In the present paper, new

developments based on PLGA foams with tailored

porosity and bioactive glass particle inclusions are pre-

sented. Both the fabrication and in vitro characterisa-

tion of the foam composites are described in detail. Two

bioresorbable copolymers with different lactide/glyco-

lide ratios were investigated. This study complements

our recent work which was mainly based on poly(D,L-lactide) (PDLLA) foams and composites [61,65].

2. Experimental

2.1. Composites processing

Poly(lactide-co-glycolide) (PLGA) copolymers with

two different lactide:glycolide (LA:GA) ratios, i.e.

LA:GA=75:25 (i.v.=0.6) and LA:GA=50:50

(i.v.=0.2), Resomer RG 756 and 502, respectively, were

provided by Boehringer-Ingelheim (Germany). Thesecopolymers will be designated as PLGA75 and 50,

respectively. Dimethylcarbonate (DMC) of 99% purity

(Sigma Aldrich) was used as a solvent. Polymers and

solvent were used without further purification. The

bioactive material used was a bioactive glass powder

(Bioglass1 grade 45S5, US Biomaterials Co., Alachua,

FL, USA). The powder had a mean particle size <5

mm. The composition of the bioactive glass used was (in

weight percentage): 45% SiO2, 24.5% Na2O, 24.5%

CaO and 6% P2O5, which is the original composition of

the first bioactive glass developed by Hench and co-

workers [59].

A 50/50 mixture (wt./wt.) of PLGA75 and PLGA50was used as the polymer component. The Bioglass1

content was varied from 10 to 50 wt.%. For the sake of

comparison, neat (unfilled) polymer foams were also

prepared and characterised.

The preparation of Bioglass1-filled polymer foams,

referred here as composite foams, followed a thermally

induced phase separation process, also termed freeze-

drying, which has been described in detail elsewhere

[30,31]. The original process was conveniently modified

for the incorporation of Bioglass1 particles as follows.

The polymer was dissolved in DMC to produce a poly-

mer weight to solvent volume ratio of 5%. The mixture

was stirred overnight to obtain a homogeneous polymer

solution. Determined amounts of Bioglass1 powder,

calculated to result in final proportions of 10, 25 and 50

wt.% of Bioglass1 in the composites, were added into

the polymer solution, resulting in homogeneous poly-

mer-Bioglass

1

particles mixtures. Each mixture wasthen transferred into a 600 ml lyophilisation flask and

sonicated for 15 min in order to improve the dispersion

of the Bioglass1 particles into the polymer solution.

The flask was then rapidly immersed into liquid nitro-

gen and was maintained at 196 C for 2 h. The frozen

mixture was then transferred into an ethylenglycol bath

at 10 C and connected to a vacuum pump (102 Torr)

for solvent sublimation (at 10 C for 48 h, and then at

0 C for additional 48 h). The foam samples were sub-

sequently completely dried at room temperature in a

vacuum oven until reaching a constant weight, as

determined by using an electronic balance.

2.2. Characterisation and in vitro studies

Neat polymer foams and polymer/Bioglass1 compo-

site samples were characterised by using scanning elec-

tron microscopy (SEM) in order to assess the porosity

structure. The apparent density of the foams (ra) was

determined by mercury pycnometry measurements as

follows: a sample of weight ws was placed in a pycn-

ometer, which was completely filled with mercury and

weighted to obtain wsl, then a was calculated according

to Eq. (1):

a ¼ ws

wl wsl þ ws

Hg ð1Þ

where wl is the weight of the pycnometer filled with

mercury, and Hg is the density of mercury (13.5 g

cm3). The density of the solid, non-porous polymer/

Bioglass1 skeleton, sk, was measured by helium pyc-

nometry using a AccuPyc 1330 pycnometer (Micro-

metrics Co.). The porosity of the foam (e) was

calculated according to Eq. (2):

" ¼sk a

sk100% ð2Þ

The compressive mechanical properties of the com-posite foams were measured with a rheometer (Ares,

Rheometric Scientific). Specimens of 10 mm diameter

and 4–5 mm height were compressed at a cross-head speed

of 1 mm/min. The compressive modulus (F) was deter-

mined from the initial linear region of the normal force

versus compression strain plot. The compressive mod-

ulus in Pa was calculated according to Eq. (3), where m

is the slope of the initial linear region of the curve:

F ¼m 104

16 ð3Þ

A.R. Boccaccini, V. Maquet/ Composites Science and Technology 63 (2003) 2417–2429 2419



For degradation studies in normal physiological con-

ditions, selected samples of polymer foams and compo-

sites were sterilised by UV exposure under a laminar

flow for 10 min and placed in sterile Falcon tubes con-

taining 50 ml of pre-filtered (0.22 mm porosity) phos-

phate buffer saline (PBS: 0. 13M, NaCl: 0.9%, NaN3:

0.02%, pH: 7.4). The samples were incubated underslow tangential agitation at 37 C and allowed to

degrade. The pH of the buffer was monitored during the

experiment. At each time point, 3–4 samples of each

foam composition were removed from the buffer, and

weighted wet after surface wiping. They were abun-

dantly rinsed with deionized distilled water (ddH2O) in

order to remove the soluble inorganic salt, and weighted

after freeze-drying. From the weight measurements,

water absorption (WA%) and weight loss (WL%) were

calculated according to Eqs. (4) and (5), respectively:

WA ¼ W

a W

oð Þ

Wo100% ð4Þ

WL ¼ Wo Wtð Þ

Wo

100% ð5Þ

where: Wo, Wa, and Wt are the samples’ weights before

immersion, after removal from the buffer and after

freeze-drying, respectively. Three to four samples of

each composition were measured and the results aver-

aged.

The changes in average molecular weight (Mw) were

determined by size exclusion chromatography (SEC)using a Helwett-Packard HP-1090 SEC apparatus

equipped with three Ultrastyragel columns (102 –105A ˚ ).

Tetrahydrofuran was used as an eluent (flow rate: 1 ml/

min) and calibration was performed using monodisperse

polystyrene standards (Polymer Laboratories Ltd.,

Shropshire, UK). After degradation for different time

periods in PBS, selected samples were characterised

using SEM, X-ray diffraction (XRD) analysis and

Raman spectroscopy.

Information on the elementary composition of Bio-

glass1 particles at the surface of the composite scaffolds

was obtained using environmental scanning electron

microscopy (ESEM) (Philips FEG XL-30) combinedwith energy dispersive X-ray analysis (EDXA). EDXA

was carried out to determine the Ca/P ratio of the Bio-

glass1 particles prior and during in vitro degradation

Fig. 2. SEM micrographs showing cross-sections of (a, b) an unfilled PLGA foam at two different magnifications; (c) a 100/25 PLGA/Bioglass1

composite at low magnification and (d) a 100/50 PLGA/Bioglass1 composite at high magnification.




and to monitor HA formation on the composite sur-

faces. The mean Ca/P ratios were determined from five

separate measurements in different areas of the compo-

site samples.

3. Results

Highly porous PLGA/Bioglass1 composites foams

have been prepared by solid-liquid phase separation and

subsequent solvent sublimation, as summarised inTable 1. The density of the composite increases with

Bioglass1 content. In parallel, the porosity decreases

when Bioglass1 content is increased. The porosity of

the composite foams was in general very high ( 90%)

even for the higher Bioglass1 content (50 wt.%). Figs. 2

(a–d) show SEM micrographs of unfilled and composite

foams. Unfilled PLGA foams are characterised by an

open, well-structured porous network composed of

interconnected tube-like macropores parallel to each

others and oriented along the heat-transfer direction

imposed by the unidirectional cooling process (Fig. 2a).

At a higher magnification (Fig. 2b) it is possible toobserve that the foam exhibits a ladder like substructure

(parallel microtubes with thin partitions). This kind of

pore architecture is typical of PLA and PLGA foams

prepared by uniaxial thermally induced phase separa-

tion process in solvents such as dioxane, benzene and

dimethylcarbonate at suitable polymer concentrations

(i.e. 5 wt.:v%) [24,30,31].

The pore morphology of the foams filled with Bio-

glass1 particles is slightly different to that of pure

PLGA foams. A typical SEM micrograph of a 100/25

PLGA/Bioglass1 composite foam shows a co-con-

tinuous structure of interconnected irregular pores of

size ranging from 10 to 100 mm (Fig. 2c). Bioglass1

particles can be easily identified on the polymer matrix

surface, more often assembled into aggregates. The dis-

persion was found to be more homogeneous for high

Bioglass1 contents, e.g. 100/50 PLGA/Bioglass1, as

seen in Fig. 2d. A qualitative good adhesion was foundbetween the PLGA matrix and Bioglass1 particles,

however no quantitative information about the inter-

facial bonding strength between polymer and Bioglass1

was obtained in this study.

Fig. 3. Compression modulus of the PLGA foam and 100/50 PLGA/

Bioglass1 composite foam.

Table 1

Density and porosity of PLGA/Bioglass1 composite foams

Composition(wt.%) Apparent density (g/cm3) Porosity (%)

PLGA/Bioglass1: 10 0/0 0 .073 0.001 94.3 0.4

PLGA/Bioglass1: 100/10 0.098 0.007 91.5 0.9

PLGA/Bioglass1: 100/25 0.124 0.024 90.4 1.8

PLGA/Bioglass1

: 100/50 0.130 0.020 89.9 0.4

Fig. 4. (a) Water absorption (WA), (b) weight loss (WL), and (c)

changes in pH of the incubation medium (PBS) versus incubation time

in PBS for PLGA/Bioglass1 composite foams with different PLGA/

Bioglass1 weight ratios: 100/0 (), 100/10(*), 100/25 (~), 100/50

(^).




Fig. 3 shows that the mechanical properties of the

PLGA/Bioglass1 composites in term of compression

modulus are significantly influenced by the incor-

poration of Bioglass1 particles: the compression mod-

ulus of the 100/50 PLGA/Bioglass1 composite foams is

three times higher than that of the pure PLGA foams.

Such improvement of the elastic modulus has been oftenobserved in degradable polymer composites reinforced

by HA and other ceramic particles [64], confirming the

expected effect of adding a rigid, inorganic particulate

phase (Bioglass1) to the polymer matrix. The mechan-

ical properties of similar foams with different polymer

matrices (PDLLA, PLGA) and increasing content of

Bioglass1 have been measured by compression test and

discussed in a previous study [66]. It was confirmed that

the compression modulus of the composite scaffolds can

be significantly enhanced by addition of Bioglass1,

especially for amorphous and though poly(D,L-lactide)

and poly(D,L-lactide-co-glycolide).

For in vitro degradation studies, the PLGA/Bio-

glass1 composites were incubated into phosphate buffer

saline in normal physiological conditions (at pH=7.4

and 37

C). Fig. 4a shows that the PLGA/Bioglass

1

composites absorbed a high amount of water during the

first week of incubation. However WA reached an

equilibrium value and started to decrease after 21 days

in all composites. In the neat PLGA foam, WA was

lower than in the composites and increased slightly over

the entire incubation period.

According to weight loss (WL) vs incubation time

plots (Fig. 4b), WL in the composites increased during

the whole incubation period and proportionally to the

Bioglass1 content. The weight of the composites con-

taining 525 wt.% Bioglass1 decreased largely during

the first week of incubation (WL = 15%) and con-

tinued to decrease proportionally to the Bioglass1 con-tent. At the end of the incubation period (35 days), WL

was around 12, 16, and 25% in the composites contain-

ing 10, 25 and 50 wt.% of Bioglass1, respectively. The

weight of the neat PLGA foams did not change to a

large extent during the whole incubation period

(WL47%).

The pH variation patterns of the media containing the

composite foams are shown in Fig. 4c. The pH of the

incubation medium decreased from the initial value

Fig. 5. SEC chromatograms for 50–50 PLGA75–PLGA50 mixtures:

(a) not incubated, (b) after 35 days of incubation in PBS.

Fig. 6. Changes in molecular weight versus incubation time in PBS for unfilled and Bioglass1 filled PLGA75 and 50 foam samples of the indicated

compositions.




(7.4) in the unfilled PLGA foams and in the composite

filled with 10 wt.% Bioglass1 after the first week of

incubation. On the contrary, the pH of the medium

slightly increased to values 57.4 for the 25 wt% and 50

wt.% Bioglass1 composites. Such difference may be

correlated to dissolution of alkaline ions from the Bio-

glass

1

particles that locally compensate for the acid-ification of the medium due to acidic products of the

polymer degradation. Such a buffering effect of Bio-

glass1 was reported elsewhere [60], and it has been

considered to be another benefit of using Bioglass1

particles in composite scaffolds with the aim to avoid

possible inflammatory response due to acidic degrada-

tion of the polymers.

The SEC chromatogram for a 50/50 mixture of

PLGA75 and 50 shows the presence of two peaks that

overlap each other, as shown in Fig. 5a. For this reason

the value of Mw at the maximum of the distribution

peak that represents the most significant fraction of the

polymer chains present in the samples has been con-sidered. This Mw value at the top of the peak has then

be used to follow the evolution of the molar mass of the

two copolymers during the incubation period. Fig. 6

shows the evolution of the Mw of both PLGA75 and

50, as measured by SEC, by using Mw at the top of each

peak as the indicative value for polymer degradation. At

the end of the incubation time, the peak for PLGA75

appears like a shoulder in the larger peak for PLGA50

and the SEC program gave only one value of Mw at thetop of the peak which was assigned to PLGA50, as

shown in Fig. 5b.

Fig. 6 shows that the Mw of PLGA75 rapidly

decreased during the first three weeks of incubation for

all Bioglass1 particle contents, but with a slower rate

for the composites containing 50 wt.% Bioglass1 as

compared to the unfilled PLGA foam. At day 21, the

Mw of PLGA75 was higher in the 100/50 PLGA/Bio-

glass1 composite in comparison to the others. The Mw

of PLGA50 did not change during the first three weeks

of incubation. For longer incubation times, the Mw of

PLGA50 started to slightly decrease. The results repor-

ted here seem to indicate that the presence of Bioglass1

particles retards the degradation rate of the polymer, at

least for PLGA75. The phenomenon was enhanced for

Fig. 7. SEM micrographs of PLGA/Bioglass1 (100/50) composite samples after incubation time of three weeks in PBS. View of the surface of the

composite sample at (a) low and (b) high magnification.

Fig. 8. SEM micrographs of PLGA/Bioglass1 (100/25) composite samples after five weeks incubation time in PBS. View of a longitudinal section of

the composite sample at (a) low and (b) high magnification.




longer degradation times, in agreement with previous

results on poly(D,L-lactide) (PDLLA)/ Bioglass1 foams

[66].

The in vitro bioactivity of the samples was assessed by

investigating the formation of HA on their surfaces

during immersion in PBS under normal physiological

conditions. After three weeks, a large number of micro-particles with diameter 1–2 mm was formed on the

surfaces of the pore walls, as shown in Fig. 7 for a 100/

50 PLGA/Bioglass1 composite. These particles were

assembled to form larger aggregates (5 mm). Their mor-

phology is typical of carbonated HA, which has been

also observed to form on Bioglass1 coated PDLLA

foams after incubating in simulated body fluid (SBF)

[61]. HA crystals were also formed on the interior of

foams as observed on longitudinal sections of composite

samples after five weeks on incubation in PBS, as shownin Fig. 8. The formation of HA on the surfaces of com-

posite foams after immersion in PBS was confirmed by

ESEM observations coupled with EDX analysis. Fig. 9

Fig. 9. (a) ESEM micrograph and (b) EDX analysis of a PLGA/Bioglass1 (100/50) composite sample after incubation time of three weeks in PBS,

confirming HA formation.




shows the morphology of a 100/50 PLGA/Bioglass1

composite after three weeks of immersion in PBS. HA

particles appear on the surface of the polymer foam,

and high Ca and P peaks belonging to HA were detec-

ted by EDX analysis (Fig. 9b). EDS analyses carried out

on different regions of the same sample showed a var-

iation in the Ca and P peak heights, which are related to

different thickness of the HA clusters formed. However,

the average Ca/P ratio of a given sample was overall

remarkably constant. The Ca/P ratio of the as-receivedBioglass1 particles was around 5. After incubation in

PBS for one week, this ratio decreased to 1.3 to reach a

value of 2 after seven weeks of incubation. This Ca/P

ratio is close to that of carbonated HA which confirms

the transformation of Bioglass1 into HA following dis-

solution mechanisms as reported in the literature [59].

The crystallinity of HA formed on the surface of

PLGA/Bioglass1 composites was confirmed by XRD,

as shown in Fig. 10 for a 50 wt.% Bioglass1 composite

sample incubated for seven and 35 days in PBS. In the

same figure the XRD pattern of the unfilled PLGA

foam is shown for comparison The HA peaks can

be seen at diffraction angle 2-theta=32

. Finally,Raman spectroscopy was used to provide an indepen-

dent confirmation of HA formation. Typical spectra are

shown in Fig. 11. A more detailed study on the structure

and further characteristics of the HA formed and its

effect on composite biocompatibility, e.g. using osteo-

blast cell culture experiments, is the focus of current

investigations.

4. Discussion

As mentioned in the Introduction several methods

have been reported in the literature for production of bioresorbable porous scaffolds for tissue engineering

applications. All the methods exhibit relative advan-

tages and disadvantages, depending on the material

used and porous microstructure required. Thermally

induced phase separation (TIPS) offers, in particular,

many advantages [24]: (i) optimal control over pore

volume fraction, (ii) possibility of designing pore shape,

orientation and size, (iii) amenable to be applied to any

polymer soluble in a suitable solvent, (iii) reproduci-

bility, and (iv) possibility to incorporate bioactive sub-

stances or growth factors into the polymer matrix. In

Fig. 10. XRD diagrams of PLGA/Bioglass1 composite samples (100/50) incubated for (a) 7 and (b) 35 days in PBS showing the development of

crystalline HA. The pattern of the neat PLGA foam (c) is also shown for comparison.

Fig. 11. Raman spectra of PLGA/Bioglass1 composite samples (100/

50) as function of incubation time in PBS for (a) 0 days, (b) 7 days, (c)

35 days. The formation of HA is indicated by the P-O peaks at 962

cm1.




the present experiments using TIPS, ultrasonication of

the polymer/Bioglass1 mixtures before freezing of the

polymer solution allows for an optimal dispersion of the

Bioglass1 particles throughout the polymer matrix. The

composition of the foams was controlled by changing

the polymer/Bioglass1 weight ratio while foam density

was mainly governed by the concentration of the poly-mer solution. In this study, dimethylcarbonate was

advantageously used as solvent instead of the more fre-

quently used dioxane, which is suspected to be carcino-

gen for humans [67].

Foams produced by the TIPS process are char-

acterised by a high porosity (>90%), comprising two

distinct pores sizes: i) macropores of average diameter

> 100 mm, and ii) micropores with an average dia-

meter of 20–30 mm, which form an interconnected net-

work. The tubular macropores are seen to be highly

oriented and parallel to each other as a result of the

unidirectional cooling process.

It is well known that for bone regeneration the idealpore size of the scaffold should approximate that of

bone [68]. In particular, scaffold macroporosity in the

range 100–250 mm of mean pore diameter is ideal for

bone-cell colonization. This is typically the range of

macroporosity that can be generated using the TIPS

process (freeze-drying). In previous studies the influence

of the processing parameters (cooling rate) and for-

mulation conditions (nature of the solvent, polymer

concentration) on pore size distribution has been pre-

sented [24,30,31]. These parameters can be conveniently

chosen for tailoring the scaffold porosity.

Pore interconnection is another relevant property thatmust be achieved in an optimised scaffold in order to

supply body fluid circulation [68]. In foams produced by

TIPS, the walls of the tubular macropores are porous

and open forming a highly interconnected microporous

network (see Fig. 2a–d). The macropores are thus con-

nected to each other through the microporous network.

This desired high pore interconnectivity achieved by the

TIPS method is not possible to be obtained by other

techniques, for example salt-leaching or gas foaming

methods. In previous studies, the transport properties of

3D porous foams made by TIPS, which are in close

relation to the pore interconnectivity, have been inves-

tigated by impedance spectroscopy, which relies uponthe measurement of the ionic conduction of water-satu-

rated foams [69,70].

Analysis of the available literature has indicated that

the application of bioactive glasses in resorbable bio-

composites for tissue engineering scaffolds is not as

widespread as that involving synthetic HA or related

calcium phosphate ceramics [64]. However, it is well

known that bioactive glasses possess a higher index of

bioactivity than hydroxyapatite [59]. In particular, the

glass used here, 45S5 Bioglass1, exhibits ‘‘Class A’’

bioactivity, i.e. it shows properties of osteoconduction

and osteoproduction, while HA exhibits ‘‘Class B’’

bioactivity, i.e. only osteoconductive behaviour [5,17,

59]. Thus, bioactive glass provides an ideal environment

for colonization, proliferation and differentiation of

osteoblasts to form new bone that has a mechanically

strong bond to the implant surface [5,59]. It has been

recently shown that this response is genetically con-trolled, with seven families of genes up regulated within

48 hours of the exposure of primary human osteoblasts

to the ionic dissolution products of bioactive glasses

[71]. Some compositions of bioactive glasses also show

strong interaction with soft tissues [2]. Pilot studies in

our laboratory (not published) have shown that rat

fibroblasts cultured with 45S5 Bioglass1 secrete

increased amounts of VEGF. In addition, in recent in

vivo studies, loosely woven meshes of polyglycolide acid

fibres (Dexon1 mesh) coated with 45S5 Bioglass1 have

been placed subcutaneously into rats leading to

increased vascularization compared with uncoated

meshes [72, 73].Thus, the incorporation of Bioglass1 into resorbable

polymer scaffolds is seen as a convenient way towards

tissue engineering scaffolds for both hard and soft tissue

applications. The principle of incorporation of Bio-

glass1 in a biodegradable polymer for tissue engineer-

ing scaffolds has been patented [74]. In our own

previous research, Bioglass1 particles have been applied

both as coating and fillers in PDLLA foams for

enhanced bioactivity [61,65,66]. Osteoblast cell culturing

studies on Bioglass1 coated PDLLA foams have

demonstrated the positive effect of the bioactive coating

in promoting cell adhesion and spreading even afteronly few minutes in culture [65]. The degradation beha-

viour of PDLLA foams filled with Bioglass1 particles

has been studied by immersion in PBS [66]. Since the

conditions of the test were the same as those of the

present study on PLGA foams, the results are compar-

able and the relative effect of Bioglass1 additions on

degradation of PDLLA and PLGA can be assessed.

The results of the degradation studies in PBS for the

PLGA/Bioglass1 composites presented here show that

the introduction of a bioactive filler in the polymer

foams increases their capacity to absorb water during

the initial incubation period (Fig. 4a). An equilibrium in

WA was rapidly reached, i.e. for three weeks on incu-bation, after which WA started to decrease. However

this decrease in WA can be correlated to a significant

weight loss, especially for composites with a high Bio-

glass1 content. Comparing with the results collected on

PDLLA/Bioglass1 composites treated in the same way

[66], it becomes evident that weight loss was much

higher in the PLGA composites than in the PDLLA

ones, all the other conditions being the same (Bioglass1

content and time of incubation). This means that

degradation was higher for PLGA composites, which is

also confirmed by the rapid decrease of the PLGA75




molecular weight during the 35 days of incubation

(Fig. 6). A quantitative study on the mechanisms of

degradation of both polymers and on the effect of Bio-

glass1 on their degradation rate is however beyond the

scope of the present work. Nevertheless the experi-

mental results presented here may serve as a suitable

and reliable data base for verification of models and forcomparison of the behaviour of different polymer/Bio-

glass1 systems.

The presence of Bioglass1 particles on the pore walls

both on the outer and internal surfaces of the foams

(Fig. 2c, d). may encourage both bone and soft tissue in-

growth from the implant/tissue interface to the interior

of the scaffold. The bioactivity of the composite scaf-

folds, determined by the rapid formation of carbonated

HA crystals on the sample surfaces during immersion in

PBS, was confirmed by electron microscopy, XRD

analyses and Raman spectroscopy.

It is envisaged that using Bioglass1 particles both as

filler and coatings in porous resorbable polymer scaf-folds will add to the possibilities of tailoring the

mechanical properties and the rate of in vivo resorption

of the composite scaffolds for the required application.

Further logical steps to the optimisation of scaffolds for

tissue engineering should focus on tailoring the micro-

structure of the composite foams for determined appli-

cations, including the development of graded porosity

and graded bioactive glass coatings, this being the focus

of current developments.

5. Conclusions

The present challenge for the progress of tissue engi-

neering is to design and fabricate reproducible bior-

esorbable 3-dimensional scaffolds, which are able to

function for a certain period of time in the body also

under load-bearing conditions. In particular for bone

and cartilage tissue engineering, numerous porous bior-

esorbable and bioactive composite systems are being

currently considered, mainly based on synthetic bior-

esorbable polymers and bioactive glass or calcium

phosphate ceramic phases. Current research and devel-

opment activities focus on choosing the adequate com-

ponents of the composites and on optimising theprocessing routes to produce the required composite

architecture, including porosity, for the envisaged

application. The composites presented in this paper

exhibit a very attractive combination of bioresorption

and bioactivity attributes in a highly porous PLGA

scaffold. They are therefore potential materials for fab-

rication of hard and soft tissue engineering scaffolds. In

particular, the use Bioglass1 instead of HA or other

calcium phosphates yields composites of class A bioac-

tivity with the possibility to be used also in soft tissue

engineering applications. To the authors‘ knowledge,

this is the first experimental report assessing the effects

of Bioglass1 particles addition on the in vitro degrada-

tion of PLGA. Future research should focus on tailoring

novel microstructures for the envisaged applications,

including the development of composite foams with

graded porosity and ‘‘engineered’’ Bioglass1 coating

microstructures.

Acknowledgements

Helpful discussions with Professor Larry L. Hench

(Imperial College London) are appreciated. The authors

acknowledge experimental assistance of I. Notingher, J.

A. Roether and J. Blaker (Imperial College London)

and of L. Pravata (University of Liege, Belgium). VM is

‘‘Postdoctoral Researcher’’ by the ‘‘Fonds National de

la Recherche Scientifique’’ (F.N.R.S). CERM is

indebted to the ‘‘Services Fe ´ de ´ raux des Affaires Scienti-

fiques, Techniques et Culturelles’’ for financial supportin the frame of the ‘‘Poles d’Attraction Inter-

universitaires: PAI 4/11’’.

References

[1] Langer R, Vacanti J. Tissue engineering. Science 1993;260:920–6.

[2] Hench LL, Polak JM. Third-generation biomedical materials.

Science 2002;295:1014–7.

[3] Agrawal CM, Ray RB. Biodegradable polymeric scaffolds for

musculoskeletal tissue engineering. J Biomed Mater Res 2001;55:

141–50.

[4] Hutmacher DW. Scaffolds in tissue engineering bone and carti-

lage. Biomaterials 2000;21:2529–43.

[5] Jones JR, Hench LL. Biomedical materials for new millenium:

perspective on the future. Mater Sci Tech 2001;17:891–900.

[6] Landers R, Huebner U, Schmelzeisen R, Muelhaupt R. Rapid

prototyping of scaffolds derived from thermoreversible hydrogels

and tailored for applications in tissue engineering. Biomaterials

2002;23:4437–47.

[7] Matthew HWT. Polymers for tissue engineering scaffolds. In:

Dumitriu S, editor. Polymeric biomaterials, 2nd ed. New York:

Marcel Dekker; 2002. p. 167–86.

[8] Weng J, Wang M. Producing chitin scaffolds with controlled pore

size and interconnectivity for tissue engineering. J Mat Sci Lett

2001;20:1401–3.

[9] Zhang Y, Zhang M. Synthesis and characterization of macro-

porous chitosan/calcium phosphate composite scaffolds for tissueengineering. J Biomed Mater Res 2001;55:304–12.

[10] Slivka MA, Leatherbury NC, Kieswetter K, Niederauer G. Por-

ous, resorbable, fibre-reinforced scaffolds tailored for articular

cartilage repair. Tissue Eng 2001;7:767–80.

[11] Seal BL, Otero TC, Panitch A. Polymeric Biomaterials for Tissue

and Organ Regeneration. Mat Sci Eng R 2001;34:147–230.

[12] Griffith LG. Polymeric biomaterials. Acta Mat 2000;48:263–77.

[13] Murphy WL, Kohn DH, Mooney DJ. Growth of continuous

bonelike mineral within porous poly(lactide-co-glycolide) scaf-

folds in vitro. J Biomed Mat Res 2000;50:50–8.

[14] Behrend D, Schmitz K-P, Haubold A. Bioresorbable polymer

materials for implant technology. Adv Eng Mat 2000;3:123–5.

[15] Chaignaud BE, Langer R, Vacanti JP. The history of tissue

engineering using synthetic biodegradable polymer scaffolds and




cells. In: Atala A, Mooney DJ, editors. Synthetic biodegradable

polymer scaffolds. Boston: Birkhauser; 1997. p. 1–14.

[16] Livingston T, Ducheyne P, Garino J. In vivo evaluation of a

bioactive scaffold for bone tissue engineering. J Biomed Mater

Res 2002;62:1–13.

[17] Sepulveda P, Jones JR, Hench LL. Bioactive sol-gel foams for

tissue repair. J Biomed Mater Res 2002;59:340–8.

[18] Domingues RZ, Clark AE, Brennan AB. A sol-gel derived

bioactive fibrous mesh. J Biomed Mater Res 2001;55:468–74.

[19] Sepulveda P, Bressiani AH, Bressiani JC, Meseguer L, Koenig Jr

B. In vivo evaluation of hydroxyapatite foams. J Biomed Mater

Res 2002;62:587–92.

[20] Barralet JE, Grover L, Gaunt T, Wright AJ, Gibson IR. Pre-

paration of macroporous calcium phosphate cement tissue engi-

neering scaffold. Biomaterials 2002;23:3063–72.

[21] Petite H, Viateau V, Bensaid W, Meunier A, de Pollak C, Bour-

guignon M, et al. Tissue-engineered bone regeneration. Nature

Biotechnology 2000;18:959–63.

[22] Wen CE, Yamada Y, Shimojima K, Chino Y, Hosokawa H,

Mabuchi M. Novel titanium foam for bone tissue engineering. J

Mat Res 2002;17:2633–7.

[23] Agrawal CM, Athanasiou KA, Heckman JD. Biodegradable

PLA–PGA polymers for tissue engineering in orthopadics. Mat

Sci Forum 1997;250:115–28.

[24] Maquet V, Jerome R. Design of macroporous biodegradable

polymer scaffold for cell transplantation. Mat Sci Forum 1997;

250:15–42.

[25] Nam YS, Park TG. Porous biodegradable polymeric scaffolds

prepared by thermally induced phase separation. J Biomed Mater

Res 1999;47:8–17.

[26] Zhang R, Ma P. Porous Poly(L-lactic acid)/apatite composites

created by biomimetic process. J Biomed Mater Res 1999;45:285–

93.

[27] Hua FJ, Kim GE, Lee JD, Son YK, Lee DS. Macroporous

poly(L-lactide) scaffold 1. preparation of a macroporous scaffold

by liquid–liquid phase separation of PLLA-dioxane-water sys-

tem. J Biomed Mater Res (Appl Biomater) 2002;63:161–7.

[28] Zmora S, Glicklis R, Cohen S. Tailoring the pore architecture in

3-D alginate scaffolds by controlling the freezing regime during

fabrication. Biomaterials 2002;23:4087–94.

[29] Ma PX, Zhang R. Microtubular architecture of biodegradable

polymer scaffolds. J Biomed Mater Res 2001;56:469–77.

[30] Schugens C, Maquet V, Grandfils C, Je ´ rome R, Teyssie ´ P. Bio-

degradable and macroporous polylactide implants for cell trans-

plantation: 1. Preparation of macroporous polylactide supports

by solid-liquid phase separation. Polymer 1996;37:1027–38.

[31] Schugens C, Maquet V, Grandfils C, Je ´ rome R, T eyssie ´ P. Poly-

lactide macroporous biodegradable implants for cell transplanta-

tion. II. Preparation of polylactide foams by liquid–liquid phase

separation. J Biomed Mater Res 1996;30:449–61.

[32] Schmitz JP, Hollinger JO. A preliminary study of the osteogenic

potential of a biodegradable alloplastic-osteoconductive alloim-

plant. Clinical Orthopaedics and Related Research 1988;237:245– 55.

[33] Coombes AGA, Heckman JD. Gel casting of resorbable poly-

mers I. Processing and applications. Biomaterials 1993;14:217–

24.

[34] Mikos AG, Bao Y, Linda LG. Preparation of poly(glycolic acid)

bonded fibre structures for cell attachment and transplantation. J

Biomed Mater Res 1993;27:183–9.

[35] Mooney DJ, Baldwin DF, Suh NP, Vacanti JP, Langer R. Novel

approach to fabricate porous sponges of poly(D,L-lactic-co-gly-

colic acid) without the use of organic solvents. Biomaterials 1996;

17:1417–22.

[36] Harris LD, Kim BS, Mooney DJ. Open pore biodegradable

matrices formed with gas foaming. J Biomed Mater Res 1998;42:

396–402.

[37] Wintermantel E, Mayer J, Blum J, Eckert K-L, Luescher P,

Mathey M. Tissue enginering scaffolds using superstructures.

Biomaterials 1996;17:83–91.

[38] Li W-J, Laurencin CT, Caterson EJ, Tuan RS, Ko FK. Electro-

spun nanofibrous structure: a novel scaffold for tissue engineer-

ing. J Biomed Mater Res 2002;60:613–21.

[39] Grande DA, Halberstadt C, Naughton G, Schwatz R, Manji R.

Evaluation of matrix scaffolds for tissue engineering of articular

cartilage grafts. J Biomed Mater Res 1997;34:211–20.

[40] Sittinger M, Reitzel D, Dauner M, Hierlemann H, Hammer C,

Kastenbauer E, et al. resorbable polyesters in cartilage engineer-

ing: affinity and biocompatibility of polymer fibre structures to

chondrocytes. J Biomed Mater Res Appl Biomat 1996;33:57–63.

[41] Borden M, El-Amin SF, Attawia M, Laurencin CT. Structural

and human cellular assessment of a novel microsphere-based tis-

sue engineered scaffold for bone repair. Biomaterials 2003;24:

597–609.

[42] Widmer MS, Gupta PK, Lu L, Meszlenyi RK, Evans GRD,

Brandt K, et al. manufacture of porous biodegradable polymer

conduits by an extrusion process for guided tissue regeneration.


[43] Whang K, Thomas CH, Healy KE, Nuber G. A novel method to

fabricate bioresorbable scaffolds. Polymers 1995;36:837–42.

[44] Landers R, Pfister A, Huebner U, John H, Schmelzeisen R,

Muelhaupt R. Fabrication of soft tissue engineering scaffolds by

means of rapid prototyping techniques. J Mat Sci 2002;37:3107–16.

[45] Zein I, Hutmacher DW, Tan KC, Teoh SH. Fused deposition

modeling of novel scaffold architectures for tissue engineering

applications. Biomaterials 2002;23:1169–85.

[46] Xiong Z, Yan Y, Zhang R, Sun L. Fabrication of porous poly(L-

lactic acid) scaffolds for bone tissue engineering via precise

extrusion. Scipta Mat 2001;45:773–9.

[47] Xiong Z, Yan Y, Wang S, Zhang R, Zhang C. Fabrication of

porous scaffolds for bone tissue engineering via low-temperature

deposition. Scripta Mat 2002;46:771–6.

[48] Park A, Wu B, Griffith L. Integration of surface modification and

3D fabrication techniques to prepare patterned poly(L-lactide)

substrates allowing regionally selective cell adhesion. J Biomater

Sci Polym Ed 1998;9:89–110.

[49] Devin JE, Attawia MA, Laurencin CT. Three-dimensional

degradable porous polymer-ceramic matrices for use in bone

repair. J Biomater Sci Polymer Edn 1996;7:661–9.

[50] Lin H-R, Kuo C-J, Yang CY, Shaw S-Y, Wu Y-J. Preparation of

macroporous biodegradable PLGA scaffolds for cell attachment

with the use of mixed salts as porogen additives. J Biomed Mater

Res (Appl Biomater) 2002;63:271–9.

[51] Nam YS, Yoon JJ, Park TG. A novel fabrication method of

macroporous biodegradable polymer scaffolds using gas foaming

salt as a pororgen. J Biomed Mater Res (Appl Biomater) 2000;53:

1–7.

[52] Durucan C, Brown PW. Calcium-deficient hydroxyapatite-PLGA

composites: mechanical and microstructural investigation. J

Biomed Mater Res 2000;51:726–34.[53] Sherwood JK, Riley SL, Palazzolo R, Brown SC, Monkhouse

DC, Coates M, et al. A three-dimensional osteochondral compo-

site scaffold for articular cartilage repair. Biomaterials 2002;23:

4739–51.

[54] Lin ASP, Barrows TH, Cartmell SH, Guldberg RE. Micro-

architectural and mechanical characterisation of oriented porous

polymer scaffolds. Biomaterials 2003;24:481–9.

[55] Mikos AG, Saranikos G, Vacanti JP, Langer RS, Cima LG.

Biocompatible polymer membranes and methods of preparation

of three dimensional membrane structures. US Patent No.

5,514,378, 1996.

[56] Yang S, Leong K-F, Du Z, Chua C-K. The design of scaffolds for

use in tissue engineering. Part I. Traditional factors. Tissue Eng

2001;7:679–89.




[57] Thomson RC, Yaszemski MJ, Mikos AG. Polymer scaffold pro-

cessing. In: Lanza RP, Langer R, Chick WL, editors. Principles

of tissue engineering. Austin, TX: R.G. Landes; 1997. p. 263–71.

[58] Taboas JM, Maddox RD, Krebsbach PH, Hollister SJ. Indirect

solid free form fabrication of local and global porous, biomimetic

and composite 3D polymer-ceramic scaffolds. Biomaterials 2003;

24:181–94.

[59] Hench LL. Bioceramics. J Am Ceram Soc 1998;81:1705–28.

[60] Stamboulis AG, Hench LL, Boccaccini AR. Mechanical proper-

ties of biodegradable polymer sutures coated with bioactive glass.

J Mat Sci: Mat Med 2002;13:843–8.

[61] Roether JA, Boccaccini AR, Hench LL, Maquet V, Gautier S,

Jerome R. Development and in vitro characterisation of novel

bioresorbable and bioactive composite materials based on poly-

lactide foams and Bioglass1 for tissue engineering applications.


[62] Laurencin CT, Lu HH. Polymer–ceramic composites for bone-

tissue engineering. In: Davies JE, editor. Bone engineering. Tor-

onto, Canada: em squared incorporated; 2000. p. 462–72.

[63] Schiller C, Siedler M, Peters F, Epple M. Functionally graded

materials of biodegradable polyesters and bone-like calcium

phosphates for bone replacement. Ceram Trans 2001;114:97–108.

[64] Boccaccini AR, Roether JA, Hench LL, Maquet V, Jerome R. A

composites approach to tissue engineering. Ceram Eng Sci Proc

2002;23(4):805–16.

[65] Roether JA, Gough JE, Boccaccini AR, Hench LL, Maquet V,

Je ´ rome R. Novel bioresorbable and bioactive composite based on

bioactive glass and polylactide foams for bone tissue engineering.

J Mater Sci Mater Med 2002;13:1207–14.

[66] Maquet V, Boccaccini AR, Pravata L, Notingher I, Jerome R.

Preparation, characterisation and in vitro degradation of bior-

esorbable and bioactive composites based on bioglass1-filled

polylactide foams. J Biomed Mat Res 2003;66A:333–46.

[67] Materials Safety Data Sheets on http://fisher.com and S Buda-

vari, The Merck Index. Withouse Station, NJ: Merck Research

Laboratory, Division of Merck & Co., Inc., 12th ed; 1996.

[68] Legeros RZ, Lin S, Rohanizadeh R, Mijares D, Legeros JP.

Diphasic calcium phosphate bioceramics: preparation, properties

and applications. J Mat Sci Mat Med 2003;14:201–9.

[69] Maquet V, Blacher S, Pirard R, Pirard J-P, Je ´ rome R. Char-

acterization of porous polylactide foams by image analysis and

impedance spectroscopy. Langmuir 2000;16:10463–70.

[70] Maquet V, Blacher S, Pirard R, Pirard J-P, Vyakarnam M, Je ´ r-

ome R. Preparation of macroporous biodegradable poly(L-lac-

tide-co-e-caprolactone) foams and characterization by mercury

intrusion porosimetry, image analysis, and impedance spectro-

scopy. J Biomed Mat Res 2003;66A:199–213.

[71] Xynos ID, Edgar AJ, Buttery LD, Hench LL, Polak JM. Gene-

expression profiling of human osteoblasts following treatment

with the ionic products of Bioglass1 45S5 dissolution. J Biomed

Mater Res 2001;55:151–7.

[72] Day R, Boccaccini AR, Roether JA, Surey S, Forbes A, Hench

LL, et al. The effect of Bioglass1 on epithelial cell and fibroblast

proliferation and incorporation into a PGA matrix. Gastro-

enterology 2002;122(4):T875.

[73] Day R, Surey S, Boccaccini AR, Roether J, Forbes A, Hench

L, et al. Assessment of polyglycolic acid mesh and bioactive

glass for soft tissue engineering scaffolds. Biomaterials 2003: in

press.

[74] Boyan BD, Niederauer G, Kieswetter K, Leatherbury NC,

Greenspan DC. biodegradable implant material comprising

bioactive ceramic, US Patent nr. 5,977, 204. November 2, 1999.


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