m-biodegradable and bioactive porous polymer

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Biomaterials 27 (2006) 3413–3431

Review

Biodegradable and bioactive porous polymer/inorganic composite

scaffolds for bone tissue engineering

K. Rezwana, Q.Z. Chena, J.J. Blakera, Aldo Roberto Boccaccinia,b,Ã

aDepartment of Materials, Imperial College London, Prince Consort Road, London SW7 2BP, UK bCentre for Tissue Engineering and Regenerative Medicine, Imperial College London, London SW7 2AZ, UK

Received 6 December 2005; accepted 31 January 2006

Available online 28 February 2006

Abstract

Biodegradable polymers and bioactive ceramics are being combined in a variety of composite materials for tissue engineering scaffolds.

Materials and fabrication routes for three-dimensional (3D) scaffolds with interconnected high porosities suitable for bone tissue

engineering are reviewed. Different polymer and ceramic compositions applied and their impact on biodegradability and bioactivity of

the scaffolds are discussed, including in vitro and in vivo assessments. The mechanical properties of today’s available porous scaffolds are

analyzed in detail, revealing insufficient elastic stiffness and compressive strength compared to human bone. Further challenges in

scaffold fabrication for tissue engineering such as biomolecules incorporation, surface functionalization and 3D scaffold characterization

are discussed, giving possible solution strategies. Stem cell incorporation into scaffolds as a future trend is addressed shortly, highlighting

the immense potential for creating next-generation synthetic/living composite biomaterials that feature high adaptiveness to the

biological environment.

r 2006 Elsevier Ltd. All rights reserved.

Keywords: Scaffolds; Bioactivity; Bone tissue engineering; Composites; Porosity; Biodegradability

Contents

1. Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3414

2. Biodegradable polymer matrices . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3414

2.1. Saturated aliphatic polyesters (PLA, PGA and PCL). . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3415

2.2. Polypropylene fumarate (PPF). . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3416

2.3. Polyhydroxyalkanoates (PHB, PHBV, P4HB, PHBHHx, PHO) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3416

2.4. Surface bioeroding polymers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3417

3. Bioactive ceramic phases . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3417

3.1. Bioactive glasses and glass-ceramics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3418

3.1.1. Composition and bioactivity . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 34183.1.2. Mechanical properties. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3419

3.2. Calcium phosphates . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3419

3.2.1. Composition and bioactivity . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3419

3.2.2. Mechanical properties. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3419

3.3. Other bioactive ceramics. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3419

ARTICLE IN PRESS

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0142-9612/$ - see front matter r 2006 Elsevier Ltd. All rights reserved.

doi:10.1016/j.biomaterials.2006.01.039

ÃCorresponding author. Department of Materials, Imperial College London, Prince Consort Road, London SW7 2BP, UK. Tel.: +44 207 594 6731;

fax: +44207 5946757.

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

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4. Material processing strategies for composite scaffolds with interconnected pores . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3420

4.1. Thermally induced phase separation (TIPS) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3421

4.2. Solvent casting and particle leaching . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3423

4.3. Solid freeform fabrication techniques (SFFT) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3423

4.4. Microsphere sintering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3423

4.5. Coated scaffolds. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3424

5. Challenges and opportunities . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3424

5.1. Mechanical integrity of porous scaffolds. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 34245.2. Incorporation of biomolecules . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3424

5.3. Long-term characterization of porous composite scaffolds . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3425

5.4. In vitro and in vivo characterization . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3425

6. Summary of current status and future trends . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3425

6.1. Summary. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3425

6.2. Future trend: stem cells and ideal biomaterials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3426

Acknowledgment . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3426

References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3426

1. Introduction

Tissue engineering applies methods from materials

engineering and life sciences to create artificial constructs

for regeneration of new tissue [1]. One common approach

is to isolate specific cells through a small biopsy from a

patient to grow them on a three-dimensional (3D) scaffold

under controlled culture conditions. Subsequently, the

construct is delivered to the desired site in the patient’s

body with the aim to direct new tissue formation into the

scaffold that can be degraded over time [1–3]. An

alternative approach is to implant scaffolds for tissue

ingrowth directly in vivo with the purpose to stimulate and

to direct tissue formation in situ [2,4,5]. The advantage of

this approach is the reduced number of operations needed,resulting in a shorter recovery time for the patient.

Facing a complex biological and sensitive system as the

human body, the requirements of scaffold materials for

tissue engineering are manifold and extremely challenging.

First, biocompatibility of the substrate materials is

imperative; that is the material must not elicit an

unresolved inflammatory response nor demonstrate im-

munogenicity or cytotoxicity. In addition, the mechanical

properties of the scaffold must be sufficient and not

collapse during handling and during the patient’s normal

activities. As with all materials in contact with the human

body, tissue scaffolds must be easily sterilizable to prevent

infection [6]. This applies notably for bulk degradable

scaffolds, where both the surface and the bulk material

must be sterile. A further requirement for a scaffold

particularly in bone engineering is a controllable inter-

connected porosity to direct the cells to grow into the

desired physical form and to support vascularization of the

ingrown tissue. A typical porosity of 90% as well as a pore

diameter of at least 100 mm is known to be compulsory for

cell penetration and a proper vascularization of the

ingrown tissue [7–10]. Other highly desirable features

concerning the scaffold processing are near-net-shape

fabrication and scalability for cost-effective industrial

production.

Today, materials used for scaffolds are natural or

synthetic polymers such as polysaccharides, poly(a-hydro-

xy ester), hydrogels or thermoplastic elastomers [2,4,11,12].

Other important categories of materials are bioactive

ceramics such as calcium phosphates and bioactive glasses

or glass-ceramics [8,13,14]. Currently, composites of

polymers and ceramics are being developed with the aim

to increase the mechanical scaffold stability and to improve

tissue interaction [14–19]. In addition, efforts have also

been invested in developing scaffolds with a drug-delivery

capacity. These scaffolds can locally release growth factors

or antibiotics and enhance bone ingrowth to treat bone

defects and even support wound healing [14,20–23].

Aforementioned requirements for scaffold materials are

numerous. To fulfill as many requirements as possible,composite systems combining advantages of polymers and

ceramics seem to be a promising choice, in particular for

bone tissue engineering, as demonstrated by the increasing

research efforts worldwide [2,14–22,24–29]. This paper

reviews tissue engineering relevant biodegradable polymers

and bioactive ceramics, including strategies for fabrication

of composite scaffolds with interconnected pores. Micro-

structure and mechanical properties will be discussed and

compared, evaluating open challenges in this field of

biomedical materials research. In vitro and in vivo

characteristics of porous composite scaffolds, with focus

on bone regeneration, will be discussed as well as

summarizing the currently available literature and pointing

to research and development needs.

2. Biodegradable polymer matrices

There are two types of biodegradable polymers: The

natural-based materials are one category, including poly-

saccharides (starch, alginate, chitin/chitosan, hyaluronic

acid derivatives) or proteins (soy, collagen, fibrin gels, silk)

and, as reinforcement, a variety of biofibers such as

lignocellulosic natural fibers which are described in detailed

studies and reviews elsewhere [30–34].

ARTICLE IN PRESS

K. Rezwan et al. / Biomaterials 27 (2006) 3413–34313414



This review will focus on the second category, synthetic

biodegradable polymers. Synthetic polymers can be pro-

duced under controlled conditions and therefore exhibit in

general predictable and reproducible mechanical and

physical properties such as tensile strength, elastic modulus

and degradation rate. A further advantage is the control of

material impurities. Possible risks such as toxicity, im-

munogenicity and favoring of infections are lower for pure

synthetic polymers with constituent monomeric unitshaving a well-known and simple structure. Table 1 gives

an overview of the discussed polymers and their physical

properties [31,35–44].

2.1. Saturated aliphatic polyesters (PLA, PGA and PCL)

The most often utilized biodegradable synthetic poly-

mers for 3D scaffolds in tissue engineering are saturated

poly-a-hydroxy esters, including poly(lactic acid) (PLA)

and poly(glycolic acid) (PGA), as well as poly(lactic-co-

glycolide) (PLGA) copolymers [2,31,45,46]. PLA exists in

three forms: L-PLA (PLLA), D-PLA (PDLA), and racemic

mixture of D,L-PLA (PDLLA).

The chemical properties of these polymers allow hydro-

lytic degradation through de-esterification. Once degraded,

the monomeric components of each polymer are removed

by natural pathways. The body already contains highly

regulated mechanisms for completely removing monomeric

components of lactic and glycolic acids. PGA is converted

to metabolites or eliminated by other mechanisms, and

PLA can be cleared through the tricarboxylic acid cycle.

Due to these properties PLA and PGA have been used in

products and devices, such as degradable sutures which

have been approved by the US Food and Drug Adminis-

tration [2]. PLA and PGA can be processed easily and their

degradation rates, physical and mechanical properties are

adjustable over a wide range by using various molecular

weights and copolymers. However, these polymers undergo

a bulk erosion process such that they can cause scaffolds to

fail prematurely. In addition, abrupt release of these acidic

degradation products can cause a strong inflammatory

response [47,48].

In general, PGA degrades faster than PLA, as found inTable 1. Their degradation rates decrease in the following

order:

PGA4PDLLA4PLLA4PCL:

Biodegradable polyester degradation occurs by uptake

of water followed by the hydrolysis of ester bonds.

Different factors affect the degradation kinetics, such as:

chemical composition and configurational structure, pro-

cessing history, molar mass (Mw), polydispersity (Mw/

Mn), environmental conditions, stress and strain, crystal-

linity, device size, morphology (e.g. porosity) and chain

orientation, distribution of chemically reactive compounds

within the matrix, additives [49,50], presence of original

monomers and overall hydrophilicity. PLGA, for instance,

has a wide range of degradation rates, the degradation

kinetics being governed by both hydrophobic/hydrophilic

balance and crystallinity. Composition of chains (i.e.

contents in L-LA and D-LA and/or GA units) determines

the degradation rate of PLGA polymers. Blends containing

the greatest amount of PGA have been shown to degrade

faster [49]. Poly(e-caprolactone) (PCL) on the other hand,

can take several years to degrade in vivo [38,51].

Thick samples of these polymers can lead to hetero-

geneous degradation, faster inside than at the exterior.

ARTICLE IN PRESS

Table 1

Physical properties of synthetic, biocompatible, and biodegradable polymers used as scaffold materials

Polymer Melting point T m (1C) Glass transition point

T g (1C)

Biodegradation time

(months)

Compressive* or

tensile strength (MPa)

Modulus (GPa)

1. Bulk degradable polymers

PDLLA [36–38] Amorphous 55–60 12–16 Pellet: 35–150* Film or disk: 1.9–2.4

Film or disk: 29–35

PLLA [36,38] 173–178 60–65 424 Pellet: 40–120* Film or disk: 1.2–3.0

Film or disk: 28–50

Fibre: 870–2300 Fibre: 10–16

PGA [37–39] 225–230 35–40 6–12 Fibre: 340–920 Fibre: 7–14

PLGA [31] Amorphous 45–55 Adjustable: 1–12 41.4–55.2 1.4–2.8

PPF [31,40] Bulk 2–30*

PCL [41] 58 À72 424

PHA and blends [62] 120–177 À2 to 4 Bulk 20–43

2. Surface erodative polymers

Poly(anhydrides)

[31,40,41]

150–200 Surface 25–27 0.14–1.4

30–40*

Poly(ortho-esters)

[31,42]

30–100 Surface 4–16* 2.5–4.4

Polyphosphazene [43] À66 to 50 242 Surface

K. Rezwan et al. / Biomaterials 27 (2006) 3413–3431 3415



Heterogeneous degradation may be ascribed to two

phenomena [46]:

(i) easier diffusion of soluble oligomers from the

surface into the external medium than from inside,

and

(ii) neutralization of carboxylic end groups located atthe surface by the external buffer solution (in vitro

or in vivo).

These phenomena contribute to reduce the acidity at

the surface whereas, in the bulk, degradation rate is

enhanced by autocatalysis due to carboxylic end groups.

Hydrolysis of amorphous polymers, e.g. PDLLA, is

faster due to the lack of crystalline regions. In general,

the amount of absorbed water depends on diffusion

coefficients of chain fragments within the polymer matrix,

temperature, buffering capacity, pH, ionic strength, addi-

tions in the matrix, in the medium and processing history.

Different aliphatic polyesters can therefore exhibit quite

distinct degradation kinetics in aqueous solutions.

PGA, for example, is a stronger acid and is more

hydrophilic than PLA, which is hydrophobic due to its

methyl groups.

The stereochemistry influences the final properties; better

alignment of neighbors leads to higher crystallinity. In

general, the initial degree of crystallinity of polyesters

affects the rate of hydrolytic degradation, as the crystal

segments are chemically more stable than amorphous

segments and reduce water permeation into the matrix.

Degradation takes longer with the stereoisomers of the

polymer, e.g. PLA composed of L-lactic repeating unitstakes more than 5 years for total absorption, whereas only

about 1 year is needed for amorphous PLA (or PDLLA)

[51].

Of particular significance for applications in tissue

engineering are debris and crystalline by-products, as well

as particularly acidic degradation products of PLA, PGA,

PCL and their copolymers that have been implicated in

adverse tissue reactions [18,38]. Several groups have

incorporated basic compounds to stabilize the pH of the

environment surrounding the polymer and to control its

degradation. Bioactive glasses and calcium phosphates

have been used [49–51]. In fact, the possibility of counter-

acting the acidic degradation of biodegradable polymers is

another reason given for the use of composites [52] as

discussed further below.

PDLLA has been extensively investigated as a biomedi-

cal coating orthopedic material because of its excellent

features with respect to implant performance [31,53]. In

addition to its high mechanical stability [54], PDLLA also

shows excellent biocompatibility in vivo and a good

osteoconductive potential [55]. PDLLA of low molecular

weight can be combined with drugs like growth factors,

antibiotics, or thrombin inhibitor [56] to establish a locally

acting drug-delivery system. It is because of these desirable

features that much more attention has recently been paid to

PDLLA for applying it as a scaffold material for tissue

engineering.

PCL is also an important member of the aliphatic

polyester family. It has been used to effectively entrap

antibiotic drugs and thus a construct made with PCL can

be considered as a drug-delivery system, being used to

enhance bone ingrowth and regeneration in the treatmentof bone defects [57]. The degradation of PCL and its

copolymers involves similar mechanisms to that of PLA,

proceeding in two stages: random hydrolytic ester cleavage

and weight loss through the diffusion of oligometric species

from the bulk. It has been found that the degradation of

PCL system with a high molecular weight ( ¯ M n of 50,000) is

remarkably slow, requiring 3 years for complete removal

from the host body [58].

2.2. Polypropylene fumarate (PPF)

PPF is an unsaturated linear polyester. Like PLA and

PGA, the degradation products of PPF (i.e. propylene

glycol and fumaric acid) are biocompatible and readily

removed from the body. The double bond along the

backbone of the polymer permits cross-linking in situ,

which causes a moldable composite to harden within

10–15 min. Mechanical properties and degradation time of

the composite may be controlled by varying the PPF

molecular weight. Therefore, preservation of the double

bonds and control of molecular weight during PPF

synthesis are critical issues [59]. PPF has been suggested

for use as a scaffold for guided tissue regeneration, often as

part of an injectable bone replacement composite [60]. It

also has been used as a substrate for osteoblast cultures[61]. The development of composite materials combining

PPF and inorganic particles, e.g. hydroxyapatite (HA) or

bioactive glasses, has not been investigated to a large

extent, in comparison with the extensive research efforts

dedicated to PLGA- and PLA-based composites.

2.3. Polyhydroxyalkanoates (PHB, PHBV, P4HB,

PHBHHx, PHO)

Polyhydroxyalkanoates (PHA) are aliphatic polyesters

as well, but produced by microorganisms under unba-

lanced growth conditions [62,63]. They are generally

biodegradable (via hydrolysis) and thermoprocessable,

making them attractive as biomaterials for applications in

medical devices and tissue engineering. Over the past years,

PHA, particularly poly-3-hydroxybutyrate (PHB), copoly-

mers of 3-hydroxybutyrate and 3-hydroxyvalerate (PHBV),

poly-4-hydroxybutyrate (P4HB), copolymers of 3-hydro-

xybutyrate and 3-hydroxyhexanoate (PHBHHx) and poly-

3-hydroxyoctanoate (PHO) were demonstrated to be

suitable for tissue engineering and are reviewed in detail

in Ref. [35].

Dependent on the property requirement by different

applications, PHA polymers can be either blended, surface

modified or composed with other polymers; enzymes or

ARTICLE IN PRESS




inorganic materials to further adjust their mechanical

properties or biocompatibility. The blending among the

several PHA themselves can change dramatically the

material properties and biocompatibility [35,62].

PHB is of particular interest for bone tissue application

as it was demonstrated to produce a consistent favorable

bone tissue adaptation response with no evidence of anundesirable chronic inflammatory response after implanta-

tion periods of up to 12 months. Bone is formed close to

the material and subsequently becomes highly organized,

with up to 80% of the implant surface lying in direct

apposition to new bone. The materials showed no evidence

of extensive structural breakdown in vivo during the

implantation period of the study [64].

However, a drawback of some PHA polymers is their

limited availability and the time-consuming extraction

procedure from bacterial cultures that is required for

obtaining sufficient processing amounts as described in the

literature [35,65]. Therefore, the extraction process might

be a challenge to a cost-effective industrial upscale

production for large amounts of some PHA polymers.

2.4. Surface bioeroding polymers

There is a family of polymers that undergoes a

heterogeneous hydrolysis process which is predominantly

confined to the polymer–water interface. This property is

referred to as ‘‘surface eroding’’ as opposed to ‘‘bulk

degrading behavior’’. Polymers known to show this

property are poly(anhydrides), poly(ortho-esters) and

polyphosphazene. These surface bioeroding polymers have

been intensively investigated as drug-delivery vehicles. The

surface-eroding characteristics offer three key advantages

over bulk degradation when used as scaffold materials: (1)

retention of mechanical integrity over the degradative

lifetime of the device, owing to the maintenance of mass to

volume ratio, (2) minimal toxic effects (i.e. local acidity),

owing to lower solubility and concentration of degradation

products, and (3) significantly enhanced bone ingrowth

into the porous scaffolds, owing to the increment in pore

size as the erosion proceeds [66]. This group of polymers

can also be designed to be bulk degradable by introducing

a high surface to bulk ratio to the scaffold. Their properties

are summarized in Table 1.

3. Bioactive ceramic phases

A common characteristic of bioactive glasses and

ceramics is a time-dependent kinetic modification of thesurface that occurs upon implantation. The surface forms a

biologically active hydroxy carbonate apatite (HCA) layer

which provides the bonding interface with tissues. The

HCA phase that forms on bioactive implants is chemically

and structurally equivalent to the mineral phase in bone,

providing interfacial bonding [13,67]. The in vivo forma-

tion of an apatite layer on the surface of a bioactive

ceramic can be reproduced in a protein-free and acellular

simulated body fluid (SBF), which is prepared to have an

ARTICLE IN PRESS

Table 2Mechanical properties of dense and highly porous hydroxyapatite, 45S5 Bioglasss, A/W glass-ceramic, and human cortical bone

Ceramics Compressive strength

(MPa)

Tensile strength

(MPa)

Elastic modulus (GPa) Fracture toughness

(Mpaffiffiffiffi

mp

)

References

Hydroxyapatite (HA) 4400 $40 $100 $1.0 [70,71]

45S5 Bioglasss $500 42 35 0.5–1 [71,72,76]

Glass-ceramicA/W 1080 215 118 2.0 [73]

Porous bioactive glass70S30C

(82%)

2.25 — — — [105]

Porous Bioglasss-derived glass-

ceramic (490%)

0.2-0.4 — — — [69]

Porous HA (82–86%) 0.21–0.41 — 0.83–1.6Â 10À3 — [106]

Cortical bone 130–180 50–151 12–18 6–8 [31,74,75]

Cancellous bone 4–12 — 0.1–0.5 — [107,108]

Fig. 1. SEM micrographs illustrating the typical ‘‘cauliflower’’ morphol-

ogy of hydroxyapatite formed on the surface of a 45S5 Bioglasss-based

foam after immersion in simulated body fluid (SBF) for 28 days. The foam

was sintered at 1000 1C for 1 h. The magnification of the framed area

shown in the inlet picture reveals rod-shaped crystals of hydroxyapatite

(adapted from Ref. [69]).




ion concentration nearly equal to that of human blood

plasma. The composition of SBF can be found in Ref. [68].

As an example, the typical ‘‘cauliflower’’ morphologies of

HCA layers formed on a scaffold made of 45S5Bioglasss

immersed in SBF are illustrated by SEM micrographs in

Fig. 1 (adapted from Ref. [69]). Typical mechanical

properties of the bioactive ceramic phases discussed inthe following paragraphs can be found in Table 2,

compiled from available literature data [31,70–76].

3.1. Bioactive glasses and glass-ceramics

In 1969, Hench et al. discovered that certain glass

compositions had excellent biocompatibility as well as the

ability of bone bonding [77]. Through interfacial and cell-

mediated reactions, bioactive glass develops a calcium-

deficient, carbonated phosphate surface layer that allows it

to chemically bond to host bone. This bone-bonding

behavior is referred to as bioactivity and has been

associated with the formation of a carbonated hydroxya-

patite (HCA) layer on the glass surface when implanted

or in contact with biological fluids [13,71,76–80] (see also

Fig. 1).

The stages that are involved in forming the bone bond of

bioactive glasses and bioactive glass-ceramics were sum-

marized by Hench [13,67]. Although some details remain

unknown, it is clearly recognized that for a bond with bone

tissue to occur, a layer of biologically active HCA must

form. This conclusion is based on the finding that HCA is

the only common characteristic of all the known bioactive

implant materials [81]. Bioactivity, however, is not an

exclusive property of bioactive glasses. HA and relatedcalcium phosphates also show an excellent ability to bond

to bone, as discussed further below. The capability of a

material to form a biological interface with surrounding

tissue is critical in elimination of scaffold loosening. Of

great importance and impact for applications in tissue

engineering is that bioactive glasses have also been found

to support enzyme activity [82–85]; vascularization [86,87];

foster osteoblast adhesion, growth, differentiation; and

induce the differentiation of mesenchymal cells into

osteoblasts [26,88–90].

A significant finding for the development of bone

engineering is that the dissolution products from bioactive

glasses, in particular the 45S5 Bioglasss composition,

upregulate the gene expression that control osteogenesis

and the production of growth factors [91]. Silicon has been

found to play a key role in the bone mineralization and

gene activation, which has led to an increased interest in

the substitution of silicon for calcium into synthetic HA.

Investigations in vivo have shown that bone ingrowth into

silicon-substituted HA granules was remarkably greater

than that into pure HA [92].

The above-mentioned advantages are the reasons why

45S5 Bioglasss is successfully used in clinical treatments of

periodontal disease (PerioglasTM) and as a bone filler

material (NovaboneTM) [13]. Bioglasss implants have also

been used to replace damaged middle ear bones, restoring

hearing to patients [79]. Bioactive glasses have gained new

attention recently as promising scaffold materials, either as

filler or coatings of polymer structures, or as porous

materials themselves, which involves melt-derived and

sol–gel-derived glasses [52,69,93–99].

3.1.1. Composition and bioactivity

The basic constituents of the most bioactive glasses are

SiO2, Na2O, CaO, and P2O5. 45S5 Bioglasss contains 45%

SiO2, 24.5% Na2O, 24.4% CaO and 6% P2O5, in weight

percent. An overview of different bioactive glass composi-

tions and their corresponding bioactivities are given in

Ref. [67]. Hench and coworkers have systematically studied

a series of glasses in the four-component systems with

a constant 6 wt% P2O5 content, as summarized in

Refs. [67,81] and they divided the compositions into three

regions according to their bioactivity. Bioactive glasses

(e.g. 45S5 Bioglasss) with compositions in the system

SiO2 –Na2O–CaO–P2O5, having o55% SiO2, exhibit a

high bioactivity index (in region A), and bond to both

soft and hard connective tissues. The bioactive glasses

(i.e. glasses in region A) are osteoproductive (bone

grows on material surfaces due to enhanced osteoblast

activity) and osteoconductive. Glasses of compositions

in region B exhibit only osteoconductivity. Compositions

in region C are resorbed within 10–30 days in tissue

[13].

It has been found that reactions on bioactive glass

surfaces can release critical concentrations of soluble Si,

Ca, P and Na ions, depending on the processing route and

particle size. The released ions induce intracellular andextracellular responses [91,100]. For example, a synchro-

nized sequence of genes is activated in osteoblasts that

undergo cell division and synthesize an extracellular

matrix, which mineralizes to become bone. In addition,

bioactive glass compositions doped with AgO2 have been

shown to elicit anti-bactericidal properties while maintain-

ing their bioactive function [101]. In recent investigations,

45S5 Bioglasss has also been shown to increase secretion

of vascular endothelial growth factor (VEGF) in vitro and

to enhance vascularization in vivo, suggesting scaffolds

containing controlled concentrations of Bioglasss might

stimulate neo-vascularization which is beneficial to large

tissue engineered constructs [86].

One key reason that makes bioactive glasses a relevant

scaffold material is the possibility of controlling a range of

chemical properties and thus the rate of bioresorption. The

structure and chemistry of glasses, in particular sol–gel-

derived glasses [78,79], can be tailored at a molecular level

by varying either composition, or thermal or environ-

mental processing history. It is possible to design glasses

with degradation properties specific to a particular

application of bone tissue engineering.

However, it was reported that crystallization of bioactive

glasses decreased the level of bioactivity [102] and even

turned a bioactive glass into an inert material [103]. This is

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a disadvantage that limits the application of bioactive

glasses as scaffold materials, as full crystallization happens

prior to densification by viscous flow sintering [104].

Extensive sintering however is necessary to densify the

struts of a scaffold to achieve the required mechanical

stability.

3.1.2. Mechanical properties

A drawback of bioactive glasses is their low fracture

toughness and mechanical strength, especially in a porous

form. Hence, bioactive glasses alone have limited applica-

tion in load-bearing situations. Table 2 gives compressive

strength, elastic modulus and some fracture toughness

values for dense and porous bioactive ceramics, compiled

from Refs. [31,69–75,105–108]. It can be seen that

particularly porous scaffolds needed for tissue engineering

exhibit very low mechanical properties compared to

cortical and cancellous bone.

Bioactive glass-ceramic materials can exhibit better

mechanical properties than amorphous glass and calcium

phosphate ceramics (Table 2). The high bending strength of

apatite-wollastonite (A/W) glass-ceramic is due to the

precipitation of the wollastonite as well as apatite phases

and attributed to the high fracture toughness of the

precipitation microstructure [109].

3.2. Calcium phosphates

3.2.1. Composition and bioactivity

Around 60 w t% of bone is made of HA Ca10(PO4)6(OH)2 and therefore it is evident why HA and

related calcium phosphates (e. g. a-TCP, b-TCP) have beenintensively investigated as the major component of scaffold

materials for bone tissue engineering [76,110–112]. As

expected, calcium phosphates have an excellent biocompat-

ibility due to their close chemical and crystal resemblance

to bone mineral [113]. Although they have not shown

osteoinductive ability, calcium phosphates certainly pos-

sess osteoconductive properties and may bind directly to

bone under certain conditions [114–116]. Numerous in vivo

and in vitro assessments have reported that calcium

phosphates, no matter of which form (bulk, coating,

powder, or porous) and of which phase (crystalline or

amorphous), always support the attachment, differentia-

tion, and proliferation of relevant cells (such as osteoblasts

and mesenchymal cells), with HA being possibly the most

efficient among them [117]. While the excellent biological

performance of HA and related crystalline calcium

phosphates has been well documented, their relatively slow

biodegradation and in particular low mechanical strength

limit their application in engineering of new bone tissue,

especially at load-bearing sites.

Crystalline calcium phosphates have long degradation

times in vivo, typically in the order of months or even

years. The dissolution rate of synthetic HA depends on the

type and concentration of the buffered or unbuffered

solutions, pH of the solution, degree of the saturation of

the solution, solid/solution ratio, and the composition and

crystallinity of the HA phase. In the case of crystalline HA,

the degree of micro- and macro-porosities, defect structure

and amount and type of other phases present have also a

significant influence [116]. Crystalline HA exhibits the

slowest degradation rate, compared with other calcium

phosphates. The dissolution rate decreases in the followingorder [118]:

Amorphous HAbaÀ TCPbbÀ TCPb crystalline HA:

3.2.2. Mechanical properties

In the body, the mechanical properties of natural bone

change with their biological location because the crystal-

linity, porosity, and composition of bone adjust to the

biological and biomechanical environment. The properties

of synthetic calcium phosphates vary significantly with

their crystallinity, grain size, porosity, and composition

(e.g. calcium deficiency) as well. In general, the mechanical

properties of synthetic calcium phosphates decrease sig-nificantly with increasing amorphous phase, microporosity

and grain size. High crystallinity, low porosity and small

grain size tend to give higher stiffness, compressive and

tensile strengths, and greater fracture toughness. It has

been reported that the flexural strength and fracture

toughness of dense HA are much lower in a dry condition

than in a wet condition [119].

If we compare the properties of HA and related calcium

phosphates with those of cortical bone (Table 2), we find

that bone has a reasonably good compressive strength

though it is lower than that of HA. But bone has a

significantly higher fracture toughness than HA. Themechanical properties are even lower for porous HA

structures. The high tensile strength and fracture toughness

of bone are attributed to the tough and flexible collagen

fibers reinforced by HA crystals. Hence, calcium

phosphates alone cannot be used for load-bearing

scaffolds despite their good biocompatibility and osteo-

conductivity.

3.3. Other bioactive ceramics

Representative bioactive ceramics are, as mentioned,

Bioglasss, HA and glass-ceramics containing HA or its

components, such as CaO and P2O5. However, formation

of HA does not seem to be limited to those ceramics only.

As Kokubo et al. [68] reported, other materials can show

degrees of bioactivity after a simple chemical heat

treatment as well. Chemical treatment of metals and

ceramics, e.g. by NaOH and heat treatments of titanium

metal, titanium alloys, and tantalum metal, and by H3PO4

treatment of tetragonal zirconia, resulted in functional

graded surfaces that induced HA formation on the surface

in SBF studies. In vivo studies of NaOH and heat-treated

titanium metals implanted in a rabbit femur showed

detaching fracture loads of up to 4 times higher than the

untreated implants [68].

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It has been suggested that the Ti–OH, Zr–OH, Nb–OH,

and Ta–OH surface groups in anatase or the tetragonal/

monoclinic structures may provide effective epitaxial

nucleation sites for apatite crystals and that a negative

surface charge increases Ca2+ adsorption from the

solution. In addition, the chemical treatment induces the

incorporation of ions (e.g. Na+

) on the metal oxide surfaceand transforms the crystal phase to an amorphous phase

resulting in an increased ion solubility [120–124]. However,

the suggested models have not been investigated in detail

yet and more studies will be required to elucidate the

mechanisms.

4. Material processing strategies for composite scaffolds

with interconnected pores

Development of composite scaffold materials is attrac-

tive as advantageous properties of two or more types of

materials can be combined to suit better the mechanical

and physiological demands of the host tissue. By taking

advantage of the formability of polymers and including

controlled-volume fractions of a bioactive ceramic phase,

mechanical reinforcement of the fabricated scaffold can be

achieved [52,125]. At the same time, the poor bioactivity of

most polymers can be counteracted.

Probably the most important driving force behind the

development of polymer/bioactive glass composite scaf-

folds for bone tissue engineering is the need for conferring

bioactive behavior to the polymer matrix, which is

achieved by the bioactive inclusions or coatings. The

degree of bioactivity is adjustable by the volume fraction,

size, shape and arrangement of inclusions [24,52,126–137].It has been shown that increased volume fraction and

higher surface area to volume ratio of inclusions favor

higher bioactivity, hence in some applications the incor-

poration of fibers instead of particles is favored [27,138].

Addition of bioactive phases to bioresorbable polymers

can also alter the polymer degradation behavior, by

allowing rapid exchange of protons in water for alkali in

the glass or ceramic. This mechanism is suggested to

provide a pH buffering effect at the polymer surface,modifying the acidic polymer degradation [24,52,139].

Inclusion of bioactive glasses has been shown to modify

surface and bulk properties of composite scaffolds by

increasing the hydrophilicity and water absorption of the

hydrophobic polymer matrix, thus altering the scaffold

degradation kinetics. In particular, the inclusion of 45S5

Bioglasss particles was found to increase water absorption

compared to pure polymer foams of PDLLA [133] and

PLGA [24,52]. In related research, it has been reported that

polymer composites filled with HA particles hydrolyzed

homogeneously due to water penetrating the interfacial

regions [140]. Ideally, the degradation and resorption

kinetics of composite scaffolds are designed to allow cells

to proliferate and secrete their own extracellular matrix

while the scaffolds gradually vanish, leaving space for

new cell and tissue growth. The physical support provided

by the 3D scaffold should be maintained until the

engineered tissue has sufficient mechanical integrity to

support itself.

There are numerous foaming techniques including sol–gel

routes to obtain highly porous structures [141,142]. How-

ever, only relevant fabrication techniques leading to 3D

composite scaffolds with highly interconnected pores are

discussed in the following paragraphs and compared in

Table 3, compiled with data available from the literature[133,135,143–148]. A selection of dense and porous scaffold

composites including their physical properties is given in

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Table 3

Fabrication routes for 3D composite scaffolds with high pore interconnectivity and their advantages and disadvantages

Fabrication route Advantages Disadvantages

Thermally induced phase separation

(TIPS) [133,143]

High porosities ($95%) Long time to sublime solvent (48 hours)

Highly interconnected pore structures Shrinkage issues

Anisotropic and tubular pores possible Small scale production

Control of structure and pore size by varyingpreparation conditions

Use of organic solvents

Solvent casting/particle leaching [144,145] Controlled porosity Structures generally isotropic

Controlled interconnectivity (if particles are

sintered)

Use of organic solvents

Solid free-form [146,147] Porous structure can be tailored to host tissue Resolution needs to be improved to the micro-scale

Protein and cell encapsulation possible Some methods use organic solvents

Good interface with medical imaging

Microsphere sintering [148] Graded porosity structures possible Interconnectivity is an issue

Controlled porosity Use of organic solvents

Can be fabricated into complex shapes

Scaffold coating [135] Quick and easy Clogging of pores, sometimes organic solvents used,

coating adhesion to substrate can be too weak




Table 4, which represent typical systems reported in

the literature [15,17,25,28,29,31,74,75,93,107,108,127,132,134,136,148–159].

For comparison, the mechanical properties of human

cortical and cancellous bone are listed as well. Representa-

tive morphologies of the fabricated scaffolds are illustrated

in Fig. 2, taken from Refs. [147,148,160–162].

4.1. Thermally induced phase separation (TIPS)

3D resorbable polymer scaffolds with very high poros-

ities ($97%) can be produced using the TIPS technique to

give controlled macro- and microstructures suitable as

scaffolds for tissues such as nerve, muscle, tendon,

ligament, intestine, bone, and teeth [52,143,163]. The

obtained scaffolds are highly porous with anisotropic

tubular morphology and extensive pore interconnectivity.

Microporosity of TIPS produced foams, their pore

morphology, mechanical properties, bioactivity and degra-

dation rates can be controlled by varying the polymer

concentration in solution, volume fraction of the secondary

phase, quenching temperature and the polymer and solvent

used as discussed in a previous review paper [52].

Briefly, the polymer is dissolved in dimethylcarbonate

and stirred overnight to obtain a homogeneous polymer

solution. A given amount of glass or ceramic powder can

be added into the polymer solution. The mixture is

transferred into a flask and sonicated. Hereafter, the

flask is quenched in liquid nitrogen and maintainedat À196 1C for 2 h. The frozen mixture is then transferred

into a cooling bath at À10 1C and connected to a vacuum

pump. The solvent is sublimated at À10 1C for 48h and

then at 0 1C for 48h, followed by drying at room

temperature in a vacuum oven until reaching a constant

weight [24].

Maquet et al. [24,133] developed highly porous PDLLA/

Bioglasss composite scaffolds prepared by TIPS with

bimodal and anisotropic pore structures composed of

tubular macropores of $100mm, interconnected with

micropores of 10–50 mm in diameter, as shown in Fig. 2a.

The pore volume was shown to decrease from 9.5 to

5.7 cm3/g after including 40 wt% Bioglasss, with little

change observed in the overall pore morphology [164]. In

vitro studies in phosphate-buffered saline at 37 1C showed

that addition of Bioglasss increased water absorption and

weight loss in comparison to pure polymer foams [24,133].

The molecular weight was found to decrease less within the

composite foams, possibly due to the dissolution of

alkaline ions from the Bioglasss providing a pH buffering

effect, as discussed above. Both the PDLLA/Bioglasss

composites and the neat PDLLA foams retained their

structural integrity until the end of the experiment (16

weeks), which means degradation was still in the early

stages.

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Table 4

Selection of scaffold composites designed for bone tissue engineering and their properties

Scaffold composite Percentage of

ceramic (%)

Porosity (%) Pore size (mm) Compressive (C),

tensile (T), flexural

(F) strength (MPa)

Modulus (MPa) Reference

Ceramic Polymer

1. Dense composites

HA fibre PDLLA 2–10.5 (vol.) — — 45 (F) 1.75–2.47Â 103 [149]

PLLA 10–70 (wt.) — — 50–60 (F) 6.4–12.8Â 103 [150]

HA PLGA 40–85 (vol.) — — 22 (F) 1.1Â 103 [29,151]

b-TCP PLLA-co-PEH 75 (wt.) — — 51 (F) 5.18Â 103 [152,153]

PPF 25 (wt.) — — 7.5–7.7 (C) 191–134

A/W PE 10–50 (vol.) — — 18–28 (B) 0.9–5.7Â 103 [154]

Cortical bone 50–150(T) 12–18Â 103 [31,74,75]

130–180 (C)

2. Porous composites

Amorphous CaP PLGA 28–75 (wt.) 75 4100 65 [28,155]

HA PLLA 50 (wt.) 85–96 100Â300 0.39 (C) 10–14 [127]

PLGA 60–75 (wt.) 81–91 800–1800 0.07–0.22 (C) 2–7.5 [156]

PLGA 30–40 110–150 337–1459 [157]

Bioglasss PLGA 75 (wt.) 43 89 0.42 (C) 51 [93,148,158]PLLA 20–50 (wt.) 77–80 $100 (macro) 1.5–3.9 (T) 137–260 [17]

$10 (micro)

PLGA 0.1–1 (wt.) 50–300 [15,148]

PDLLA 5–29 (wt.) 94 $100 (macro) 0.07–0.08 0.65–1.2 [132,134,136]

10–50 (micro)

Phosphate glass A/W P LA-PDLLA 40 (wt.) 93–97

PDLLA 20–40 (wt.) 85.5–95.2 98–154 0.017–0.020 (C) 0.075–0.12 [159,25]

Cancellous Bone 4–12 (C) 100–500 [107,108]




It is well documented [52,163,165] that due to auto-

catalysis non-porous biodegradable polylactides undergo

degradation more rapidly than the porous counterparts.

The reason for this effect lies in the fact that porous

materials are able to facilitate dissolving and spreading of

degradation products throughout the aqueous medium,

thereby preventing autocatalysis. PDLLA/Bioglasss com-

posites exhibit high bioactivity, assessed by the formation

of HA on the composite surfaces upon immersion in SBF

[133,134]. It has also been shown that the foams support

the migration, adhesion, spreading and viability of MG-63

cells (osteosarcoma cell line) [134].

The potential of these scaffolds in bone and soft-tissue

engineering has been demonstrated in vitro with optimized

concentrations of 45S5 Bioglasss added to PDLLA or

PLGA matrices [134,136]. Highly porous tubular scaffolds

with oriented porosity have also been fabricated by

exploiting the TIPS process [15,166]. These are candidate

materials for soft-tissue engineering with potential applica-

tion in the regeneration of tissues requiring tubular shapes

such as the intestine, trachea and blood vessels. TIPS

fabricated PDLLA foams with and without Bioglasss

additions have been shown to exhibit mechanical aniso-

tropy concomitant with the TIPS-induced pore architecture

[132].

Polymer matrix composite films containing nanosized

titania and other inorganic particulate inclusions have

demonstrated enhanced cell adhesion and a tendency

to increased Ca-containing mineral deposition [167].

Recently, 3D PDLLA foams containing both TiO2

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Fig. 2. Typical morphologies of porous polymer foams produced by different techniques and structure of cancellous bone. (a) Thermal induced phase

separation (TIPS, adapted from Ref. [160]), (b) solvent casting and particle leaching [161], (c) solid freeform fabrication technique [147], (d) microsphere-

sintering [148], (e) cancellous bone [162]. (Micrographs (b), (c) and (e) reprinted with permission of Elsevier Ltd. Micrograph (d) reprinted with permission

of John Wiley & Sons, Inc.)




nanoparticles and Bioglasss additions have been synthe-

sized by TIPS. These foams demonstrate enhanced

bioactivity and surface nanotopography [131].

4.2. Solvent casting and particle leaching

Solvent casting of biocomposite scaffolds involves thedissolution of the polymer in an organic solvent, mixing

with ceramic granules, and casting the solution into a

predefined 3D mould. The solvent is subsequently allowed

to evaporate. The main advantage of this processing

technique is the ease of fabrication without the need of

specialized equipment. The primary disadvantages of

solvent casting are: (1) the limitation in the shapes

(typically flat sheets and tubes are the only shapes that

can be formed); (2) the possible retention of toxic solvent

within the polymer; and (3) the denaturation of the

proteins and other molecules incorporated into the

polymer by the use of solvents. The use of organic solvents

to cast the polymer may decrease the activity of bioinduc-

tive molecules (e.g. protein). The detailed processing steps

have been described in the literature [93].

Polymer-ceramic constructs can also be fabricated by the

solvent aggregation method. The polymer microspheres are

first formed from traditional water oil/water emulsions.

Solvent-aggregated polymer-ceramic scaffolds can then be

constructed by mixing solvent, salt or sugar particles, ceramic

granules, and pre-hardened microspheres [131]. A 3D structure

of controlled porosity is formed based on this method

combined with particle leaching and microsphere packing.

Fig. 2b illustrates a typical pore morphology obtained by this

technique. The method shares similar advantages anddisadvantages with the solvent casting technique [93].

There has been little work done on producing bioactive

polymer-ceramic scaffolds using particle leaching. Cer-

tainly, a drawback of this technique is achieving pore

interconnectivity at low porogen (salt/sucrose) loadings, as

many of the porogen particles may remain trapped.

Nevertheless, composites based on calcium phosphate

inclusions with variable and graded porosity have been

produced using this route [168].

4.3. Solid freeform fabrication techniques (SFFT)

SFFT, such as fused deposition modeling, have been

employed to fabricate highly reproducible scaffolds with

fully interconnected porous networks [147,169] as shown in

Fig. 2c. Using digital data produced by an imaging source

such as computer tomography or magnetic resonance

imaging enables accurate design of the scaffold structure

[169]. Solid freeform (SFF) manufacturing coupled with

conventional foam scaffold fabrication procedures (phase

separation, emulsion-solvent diffusion or porogen leach-

ing) may be used to develop scaffolds with controlled

micro- and macroporous structures. Such biomimetic

internal architectures may prove valuable for multi-tissue

and structural tissue interface engineering.

To the authors’ knowledge, there is no literature

available on degradable polymer/bioactive glass compo-

sites made by SFFT. This technique has been only applied

for composites containing calcium phosphates as the

bioactive phase [147,170]. For example, Xiong et al. [170]

fabricated composites of PLLA/TCP with porosities of up

to 90% and mechanical properties close to humancancellous bone by using low-temperature deposition based

on a layer-by-layer manufacturing method of SFF fabrica-

tion (computer-driven by 3D digital models). PLLA was

dissolved in dioxane and TCP powder mixed to prepare a

slurry, which was formed into frozen scaffolds, and

subsequently freeze-dried. Alternate parallel layers formed

macropores (400mm diameter) and sublimation of the

solvent during freeze-drying formed micropores (5mm

diameter). Taboas et al. [147] produced PLA scaffolds

with computationally designed pores (500–800mm wide

channels) and solvent-derived local pores (50–100mm wide

voids or 5–10 mm length plates). Indirect fabrication using

casting in SFF moulds provided enhanced control over

scaffold shape, porosity and pore architecture, including

size, geometry, orientation, branching and interconnectivity.

A shortcoming of this route is increased scaffold fabrication

time compared with direct methods, as a temporary mould

must be made first.

4.4. Microsphere sintering

In this process, microspheres of a ceramic and polymer

composite are synthesized first, using emulsion/solvent

evaporation technique. Sintering the composite micro-spheres yields a 3D, porous scaffold [28,155]. 3D compo-

sites of degradable polymers and bioactive glass have been

produced by sintering composite microspheres by Lu et al.

[148]. Starting materials were PLAGA-Bioglasss compo-

site microspheres obtained through a water–oil–water

emulsion technique. Sintering of the microspheres into

cylindrical shapes resulted in a well-integrated intercon-

nected porous structure, with the microspheres joined at

the contact necks. Average porosity was 40% with pore

diameters of 90mm, and mechanical properties close to

cancellous bone. The composites were shown to be

bioactive as a calcium phosphate layer formed on the

surface of the composite on immersion in SBF for 7 days.

Moreover, Bioglasss reinforcement gave a two-fold

increase in compressive strength. The scaffolds were shown

to support the adhesion, growth and mineralization of

human osteoblast-like cells in vitro. Over a 3-week period,

cultures with PLAGA/Bioglasss maintained pH variations

within physiological ranges. More recently, Yao et al. [19]

synthesized PLGA/bioactive glass microspheres by emulsi-

fication and heated them in moulds to fabricate porous 3D

scaffolds. They demonstrated the bioactivity of the

composites and their ability to promote osteogenesis of

marrow stromal cells. A typical microsphere-sintered 3D

structure fabricated by Lu et al. [148] is given by Fig. 2d.

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4.5. Coated scaffolds

Bioceramic-coated porous scaffolds have been produced

either as foams [135], fibrous bodies [171] or meshes [86,158]

by slurry dipping or electrophoretic deposition (EPD).

Roether et al. [135] were the first to develop composites of

macroporous PDLLA foams coated with Bioglasss

parti-cles (grade 45S5 with particle sizeo5 mm) by slurry dipping.

A stable and homogeneous coating on the foam surface and

infiltration of Bioglasss particles throughout the porous

network were achieved. A stable slurry of 42wt%

Bioglasss in deionized water gave relatively dense and

uniform adherent coatings. EPD was investigated as an

alternative route whereby charged Bioglasss particles in

aqueous suspension infiltrated the foam with its tubular

macropores oriented perpendicularly to the larger dimen-

sion of the electrodes [135]. The slurry dipping technique

was found to be more suitable than the EPD route; the

latter caused frequently sealing of the interconnected pores

by Bioglasss particles. EPD, however, is an attractive

method to incorporate nanoparticles into porous structures

with potential use as tissue engineering scaffolds [172].

Composites tested in vitro in acellular SBF exhibited

increasing formation of HA (layers of 10 mm were formed

after 28 days). In addition, changes in pore morphology as

a result of polymer degradation with increasing immersion

time were observed. The investigation of the in vitro

behavior of osteoblast-like cells demonstrated that cells

were able to migrate through the porous network and

colonize the inner sections of the foams. Also, after 24 h a

higher cell density was observed in the Bioglasss coated

foams compared to the pure PDLLA foams [90].A related approach, but using PHA woven meshes as

substrate, was followed by Olsen et al. [173]. In this case, a

45S5 Bioglasss aqueous slurry was used to deposit

micrometer-sized Bioglasss particles on the surfaces of

PHA fibers. The process was optimized to coat individual

fibers maintaining the pore structure of the woven mesh.

There are however no published results on the in vivo

behavior of these Bioglasss-PHA composites.

Another promising method to coat polymer surfaces with

bone like calcium phosphates is using a biomimetic approach

as demonstrated in the literature [174–176]. Moreover,

sodium silicate gel [174] has been used to nucleate CaP

coatings on polymer surfaces that were immersed in SBF. A

review on biomimetic formation of calcium phosphate

coatings has been published recently [177].

5. Challenges and opportunities

5.1. Mechanical integrity of porous scaffolds

Comparison of the mechanical properties of today’s

available porous scaffolds with relevant properties of bone

reveals the insufficient mechanical integrity of the man-

made scaffolds. In Fig. 3 the elastic modulus and the

compressive strength of dense bioactive ceramic, biode-

gradable polymers, cancellous and cortical bone are

compared with porous monophasic scaffolds and compo-

sites thereof. Mechanical data for porous bioactiveceramics and for polymer foams were taken from Refs.

[105,106,132]. It can be seen that some dense polymers

match cancellous bone properties and approach cortical

bone properties. Moreover, the bioactive ceramics region is

close to the properties of cortical bone as well. Porous

scaffolds however are at least one order of magnitude

weaker than cancellous bone and orders of magnitude

weaker than cortical bone.

Interestingly, the stiffness achieved by up-to-date fabri-

cated porous bioactive ceramics is less than the stiffness of

most porous biodegradable polymer scaffolds. By compar-

ing the mechanical properties of the porous composites tothose of porous polymer scaffolds, a slight increase of

mechanical properties is revealed. But the increase in

stiffness and strength is certainly below expectations; most

probably this can be attributed to the lack of interfacial

bonding strength between the ceramic phase and the

polymer matrix, which has been neglected in most studies.

The ceramic phases are in general very hydrophilic whereas

the polymers are hydrophobic. The increase of bonding at

the interface is a challenge and might be achieved by using

surfactants chemisorbed on the particle surface prior to

composite processing. Using surface functionalized parti-

cles in the nanosize range—featuring a higher specific

surface and thus a higher interface area—might even

increase the interfacial bonding strength, and thus the

overall mechanical properties of the composite scaffold

could be effectively enhanced. However, the increase of

interfacial bonding and introduction of surfactants are

likely to have an impact on degradation kinetics and

cytotoxicity of the composite. These effects are largely

unknown and remain to be investigated.

5.2. Incorporation of biomolecules

There is a significant scope in the application of surface

modifications, through the use of protein adsorption or

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Fig. 3. Elastic modulus vs. compressive strength of biodegradable

polymers, bioactive ceramics and composites reviewed in this paper.

Porosities of the porous scaffolds are 475% and mostly interconnected.




plasma treatment, to provide more cues to cell attachment

and response [3,178]. The possibility of incorporating

growth factors into composites formed by biodegradable

polymers and bioactive glasses or HA inclusions has

started to be explored [179–182].

Integrins, laminin and RGD proteins were shown to be

essential for cell attachment to materials surfaces[183–186]. The immobilization of these proteins should

not only promote cell adhesion and proliferation but also

increase wettability of hydrophobic polymers such as

PDLLA. To control protein adhesion and release kinetics,

different protein immobilization routes can be used as

demonstrated for polymer surfaces [187] and ceramic

surfaces [188,189]. Certain growth factors were shown in

in vivo studies to be osteoinductive. Possible growth

factors include bone morphogenetic proteins, transforming

growth factor beta, VEGF, and insulin-like growth factor

as reviewed in Refs. [190–192]. Immobilizing these growth

factors on the scaffold surface might significantly shorten

the bone healing process and reduce patient recovery time.

The incorporation of biomolecules does not allow

extreme temperature ranges (470 1C) or extremely aggres-

sive chemical conditions during processing, being a

challenge to the scaffold fabrication process. ‘‘Soft’’

material routes like sol–gel processing might be a strategy

to incorporate biomolecules during scaffold fabrication. To

the authors’ knowledge, however, sol–gel-derived bioactive

organic/inorganic hybrids have not yet been formed into

highly interconnected porous structures, which is essential

for application of these composites as scaffolds. Another

related challenge is the elucidation of the local impact of

growth factors on the cell and tissue systems, includinglong-term effects.

5.3. Long-term characterization of porous composite

scaffolds

There is a lack of current understanding in the literature

regarding the long-term in vitro and in vivo characteriza-

tion of the porous 3D scaffold composites discussed here,

specifically regarding the long-term effect of the incorpora-

tion of inorganic bioactive phases on the degradation and

ion release kinetics of these highly porous systems. In this

regard, the development of appropriate characterization

techniques coupled with predictive analytical models is

mandatory in order to be able to comprehensively assess

the degradation of these systems with respect to pore

structure, scaffolds geometry, fluid flow and the influence

of the bioactive additions. Here, the use of X-ray

microtomography as a reliable tool for 3D pore structure

quantification is likely to gain increased impetus allowing

resolutions down to 1 mm [193]. Combining image analysisand impedance spectroscopy is another possible approach

to characterize pore interconnectivities of scaffolds as

shown recently, being less straightforward however [164].

5.4. In vitro and in vivo characterization

While a good number of in vitro and in vivo studies exist

for biodegradable polymers and bioactive ceramics alone,

in vitro studies for polymer/ceramic composites have just

started [15,19,86,134,136,137,148,194,195]. Table 5 gives

examples of the types of composite scaffolds investigated in

vitro and the applied cell cultures. Very few compositesystems have been investigated in vivo up to date. More

research needs also to be directed at assessing the

suitability of the reviewed bioactive composite scaffolds

in soft-tissue engineering strategies, including further

investigations of the effect of dissolution products from

the bioactive phase on vascularization and in vivo new

tissue growth.

6. Summary of current status and future trends

6.1. Summary

The synthetic and biodegradable, polymer/inorganic

bioactive phase composites reviewed in this article are

particularly attractive as tissue engineering scaffolds due to

their shapability, bioactive behavior and adjustable biode-

gradation kinetics. Conventional materials processing

methods have been adapted and extended for incorpora-

tion of inorganic bioactive phases into porous and

interconnected 3D polymer networks.

From the materials science perspective, the present

challenge in tissue engineering is to design and fabricate

reproducible bioactive and bioresorbable 3D scaffolds of

tailored porosity and pore structure, which are able to

maintain their structure and integrity for predictable times,

even under load-bearing conditions. As reviewed here, the

ARTICLE IN PRESS

Table 5

Overview of in vitro investigated biodegradable polymer/inorganic phase composites available in the literature

Composite PLGA/45S5 Bioglasss PDLLA/45S5 Bioglasss PLGA/HA PCL/HA

Cell culture Mouse fibroblasts (L929) [15],

marrow stromal cells [19],

mouse fibroblasts (208F) [86],

human osteosarcoma cells

(SaOS-2) [148]

Mouse fibroblasts (208F) [137],

human osteosarcoma cell line

(MG-63) [134], human lung

carcinoma (A549) [134,136]

Rat calvarial osteoblasts [194] Human osteosarcoma cells

(SaOS-2), osteoblasts from

human trabecular bone [195]




mechanical integrity of man-made composite scaffolds is

still at least one order of magnitude lower than that of

cancellous or cortical bone. Achieving the mechanical

properties of bone might also allow replacing bigger parts

of damaged bone tissue than what is possible today.

The incorporation of biomolecules such as growth

factors with the aim to accelerate local bone healing ispromising and currently under extensive research. Incor-

porating biomolecules during scaffold processing however

is not simple as biomolecules are sensitive to elevated

temperatures and extreme chemical conditions. A promis-

ing strategy is the immobilization of proteins and growth

factors in the post-processing phase via surface functiona-

lization of the scaffold.

Hardly any in vitro and particularly in vivo studies exist

for the composite scaffolds reviewed in this article.

However, in order to target clinical applications, in vitro

and in vivo studies are inevitable and the need for more

studies in biological systems is imperative.

6.2. Future trend: stem cells and ideal biomaterials

The application of biomaterials in hard tissue repair

started with bioinert approaches, which involved the

development and application of bioinert materials. These

materials are applied in most permanent bioimplants in

today’s clinical use such as, for example, hip–joint

replacements. The subsequent development of biomaterials

focused on bone-bonding properties of bioactive glasses

and ceramics [81]. This period was soon followed by the

development of biodegradable materials for bone tissue

engineering scaffolds that can stimulate specific cellularresponses at the molecular level [23].

The composite scaffolds reviewed in this article combine

the features of the biomaterials in the second and third

periods: they possess bioactivity, degradability and the

possibility of biomolecule incorporation. Over the years,

the developed biomaterials have addressed biological

aspects of increased complexity, starting on the level of

ion interactions and moving then to growth factor

incorporation. The biomaterials were extended from purely

synthetic materials to material/biologic hybrids, engineer-

ing at the same time bioactivity and biodegradability.

While current research is still focused on the interaction

between stromal cells and biomaterials, the fundaments for

biomaterials seem to originate from introducing stem cells.

Scaffolds seeded with stem cells allow local cell function

adaptation by differentiation of stem cells as demonstrated

by Levenberg et al. in 2003 [8,196]. This new approach

enables the scaffold surface to mimic complex local

biological functions and may lead in near future to in

vitro and in vivo growth of tissues and organs. The

interface of stem cells and scaffolds are at the moment in

the center of attention, issuing growth factor incorporation

and cell adhesion [8]. In this approach, we anticipate that

engineered composite scaffolds made by biodegradable

polymer matrices with bioactive inorganic phases, as

reviewed here, will play a vital role and perhaps they will

be the ‘‘scaffolds of choice’’ in combination with stem cell

seeding.

Acknowledgment

Stimulating discussions with Prof. L.L. Hench, Prof. J.Polak and Dr. A. Bishop (TERM-Centre, Imperial College

London) and with Dr. F. Filser (ETH Zurich) are

acknowledged.

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