Advanced biomaterials for skeletal tissue regeneration...

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Materials Science and Engineering R 59 (2008) 38–71

Advanced biomaterials for skeletal tissue regeneration: Instructive

and smart functions

F. Barrere a,*, T.A. Mahmood b, K. de Groot a, C.A. van Blitterswijk a

a University of Twente, Department of Tissue Regeneration, 7500 AE Enschede, The Netherlandsb Amgen Inc., Division of Translational Sciences, One Amgen Center Drive, Thousand Oaks, CA 91320, United States

Available online 1 February 2008

Abstract

The past half century has seen explosive growth in the use of medical implants. Orthopedic, cardiac, oral, maxillofacial and plastic surgeons are

examples of medical specialists treating millions of patients each year by implanting devices varying from pace makers, artificial hip joints, breast

and dental implants, to implantable hearing aids. All such medical implants make use of special materials, known as biomaterials, defined as

‘‘materials intended to interface with biological systems to evaluate, treat, augment or replace any tissue, organ, or function of the body’’ [D.F.

Williams, The Williams Dictionnary of Biomaterials, Liverpool University Press, Liverpool, 1999]. While the priority for the first generation of

biomaterials was inertness with living tissues, the field is shifting towards biologically active systems in order to improve their performance and to

expand their use. Biomaterials can be combined as scaffolds with cells (i.e. tissue engineering), growth factors or genetic material in order to trigger

tissue regeneration. In addition, recent reports have shown the possibility to design biomaterials that can activate cellular processes and tissue

formation solely by their intrinsic physicochemical and three dimensional spatial properties. This article reviews the recent developments in the

design of biomaterials that integrate our understanding of cellular and molecular mechanisms with materials science. After an overview of the

physicochemical and biological processes occurring at the interface between the biomaterials and biological milieu, we will address the biological

principles contributing to the design and engineering of advanced biomaterials for application towards recent therapeutic strategies for tissue

regeneration. Finally, future directions for the design of advanced biomaterials will be discussed.

# 2007 Elsevier B.V. All rights reserved.

Keywords: Skeletal tissue; Medical implant; Biomaterials; Tissue Engineering

Contents

1. Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 39

2. Hard skeletal tissues. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 40

2.1. Structural and compositional organization . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 40

2.1.1. Physical and chemical component . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 40

2.1.2. Cellular components . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 41

2.1.3. Biological properties. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 42

2.2. Self repair of hard tissues: example of smart materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 43

3. Soft skeletal tissues . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 43

3.1. Structural and compositional organization of cartilage . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 43

3.1.1. Physical and chemical components . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 43

3.1.2. Cellular component. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 44

3.1.3. Biological properties. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 44

3.1.4. Intrinsic cartilage tissue repair . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 44

4. Current strategies to repair skeletal tissues . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 44

4.1. Biomaterials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 44

4.2. Cell-based therapies . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 45

* Corresponding author. Current address: Kuros Biosurgery AG, Technoparkstrasse 1, 8005 Zurich, Switzerland. Tel.: +41 44 200 56 52; fax: +41 44 200 57 52.

E-mail address: [email protected] (F. Barrere).

0927-796X/$ – see front matter # 2007 Elsevier B.V. All rights reserved.

doi:10.1016/j.mser.2007.12.001

mailto:[email protected]

http://dx.doi.org/10.1016/j.mser.2007.12.001

F. Barrere et al. / Materials Science and Engineering R 59 (2008) 38–71 39

4.3. Towards smart designs to repair tissues . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 47

4.3.1. Biomimetic materials for skeletal repair . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 47

4.3.2. Tissue engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 48

4.3.3. Materials for tissue engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 48

5. Instructing physico-chemical and biological processes at biomaterial interfaces. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 52

5.1. Osteoinductive biomaterials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 52

5.2. Orchestrating biomaterials degradation with new tissue formation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 54

5.2.1. Degradation mechanisms . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 54

5.2.2. Smart degradation designs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 54

5.3. The role of proteins in regulating biomaterial-induced biological response . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 55

5.3.1. Protein adsorption onto biomaterials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 56

5.3.2. Smart designs to control protein adsorption . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 56

5.4. Controlling cell–biomaterial interactions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 57

5.4.1. Cellular activities and functions on biomaterials. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 57

5.4.2. Effect of surface physico-chemistry . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 59

5.4.3. Effect of topography . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 59

5.4.4. Micro- and nanodesigns of biomaterials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 60

5.5. Multi-functional scaffolds . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 60

5.5.1. Multiple-component scaffolds . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 61

5.5.2. Multidrug delivery . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 61

5.6. Three-dimensional control of biomaterials on cells and tissues. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 62

5.7. Physical stimuli on cells . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 63

6. Concluding remarks . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 64

Acknowledgements . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 65

References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 65

1. Introduction

The past half century has seen explosive growth in the use of

medical implants. Orthopedic, cardiac, oral, maxillofacial and

plastic surgeons are only examples of medical specialists

treating millions of patients each year by implanting devices as

diverse as pace makers, artificial hip joints, breast implants, to

dental implants and implantable hearing aids. The cost of

treating diseases and problems caused by loss of tissue

function, now exceeds US$39 billion in North America alone

[2]. In addition, an aging population and increasing life

expectancy in the developed world now means that our tissues

and organs are pressed into service longer than they may be able

to independently withstand.

So what has been done to deal with these modalities? To

date, in cases where pharmacological treatments alone have

been insufficient, the principal approach has been the

transplantation of organs and tissues from other people

(allogeneic transplantation) or moving tissue from healthy

parts of the patient to diseased areas (autologous transplanta-

tion) [3,4]. The scarcity of donors compounded with issues of

transplant rejection and donor site morbidity stimulated early

research into replacing sophisticated tissues and organs with

artificial, man-made substitutes. To date, tens of millions of

individuals have had the quality of their lives enhanced for as

long as 25 years by the use of these man-made implants [5].

However, far from being long-term replacements, early devices

such as the Jarvik heart became temporary devices for

critically-ill patients waiting for heart transplants. The question

of how to solve the tissue and organ transplant problem

remained unanswered.

The lack of efficacy seen by using biomaterials alone has

been due to their poor responsiveness in comparison with the

flexibility and reactivity of natural tissues and organs. Materials

scientists alone cannot solve this complex biological issue as

these biomaterials by definition must interact with, and function

in living entities. Nowadays, the strategy to design smart

biomaterials lies in their capacity to instruct biological entities

to entirely regenerate tissues; in other words, to create a

synthetic twin tissue or organ that can function as its natural,

original tissue.

This strategy started in the late 1960s by Larry Hench.

Horrified by the amputation of limbs of thousands of Vietnam

War battlefied casualties, he started to design biomaterials ‘‘to

repair people, instead of making materials to destroy them’’

[5,6]. To do so, he hypothesized that an implant containing

calcium and phosphate in proportions similar to bone mineral

would not be rejected by the body. Indeed, a physicochemical

bond was observed between the biomaterials designed by

Hench (Bioglass1) and the hosting bone. To date, although

calcium phosphate biomaterials are relatively ‘‘old’’, their

ability to trigger bone formation is incomparable with other

biomaterials. However, these bone mineral-like substitutes

cannot substitute the bones’ mechanical function, illustrating

the as yet unmet medical need that is the driver behind our

ultimate goal of developing a synthetic entity that would

entirely substitute and regenerate a damaged tissue or organ.

Today, the way to deal with this issue remains similar to

Hench’s strategy. We aim at creating synthetics with the

appropriate and full responsiveness towards biological milieu,

i.e. smart systems. To do so, we continuously investigate

biological mechanisms occurring in tissues and organs, and at

F. Barrere et al. / Materials Science and Engineering R 59 (2008) 38–7140

biomaterial interfaces at the molecular, cellular and macro-

scopic levels in, order to pinpoint the critical phenomena that

can help us understand the fundamental processes that occur in

these systems. Knowledge of these mechanism is crucial to

create, a man-made substitute that could heal over the long-

term. Naturally, the biomaterials’ field is shifting towards

biologically active systems in order to improve their

performance and to expand their use. Biomaterials have

combined them as scaffold with autologous cells (i.e. tissue

engineering), to render tissue substitutes more ‘‘alive’’ and

more reactive towards biological environment. More recently,

there has been considerable interest in the development of

‘‘smart materials’’ that are able to instruct the behavior of

adhered or encapsulated cells by releasing bioactive molecules

into the local environment, or through extracellular protein/

peptide mimetics built into the delivery substrates [7,8].

However, the ability of materials to modulate downstream gene

response without exogenous growth factors, coatings or

complex ligand incorporation has the potential to greatly

facilitate the development of tissue engineering and cellular

therapies. As an illustration of this concept, a class of biphasic

calcium phosphate ceramic induced de novo bone formation at

non-osseous sites in vivo without requiring the delivery of cells

or biologic compounds [9], suggesting that the surface

chemistry of the ceramic allowed the selective adsorption of

morphogenetic proteins that trigger osteogenesis. It was also

demonstrated using polymer libraries that substrate chemistry

can influence the developmental lineages of embryonic stem

cells [7]. Although the use of embryonic stem cells lineages is

far from clinical applications in skeletal repair and will not be

discussed in this review, this pioneer study illustrates the

potential of biomaterials in directing cellular differentiation.

This article reviews recent developments and approaches in

the design of smart biomaterials that integrate cellular and

molecular biology in order instruct the biological milieu with

the ultimate goal of total tissue regeneration.

2. Hard skeletal tissues

Some 540 million years ago, within a period of a few million

years, a multitude of primarily multicellular organisms began to

produce mineralized structures that are widespread among the

mollusks, the vertebrates, echinoderms, plants and protoctists.

These mineralized structures were formed in order to fulfill

either specific or multipurpose functions [10]. In vertebrates,

besides the presence (negligible in weight) of tiny magnetite

minerals in human brains that are responsible for orientation,

navigation and homing skills [11], apatitic calcium phosphate

apatite is the main body’s mineral component. It is present in

hard tissues, i.e. bone and teeth, whose functions are structural

protection, motion and mastication. Bone is the component of

the skeletal system, which is involved in the protection, support

and motion of the body. Flexible and elastic, bones from the rib

cages protect heart, lungs and other organs which function

involves motion, expansion and contraction. Being stiff, bone

structurally supports the mechanical action of soft tissues, like

the contraction of muscles or expansion of lung. At a cellular

level, bone is a protective and production site for specialized

tissues such as bone marrow, which is a blood-forming system.

Finally, it is a mineral reservoir used by endocrine systems to

regulate the calcium and phosphate homeostasis in the

circulating body fluids.

Hard tissues are smart, labile and reactive towards their

environment. Their formation, their structure and adaptations

are induced by a highly complex physico-chemical and cellular

machinery.

2.1. Structural and compositional organization

2.1.1. Physical and chemical component

In shape and macrostructure, bones and teeth are affected by

genetic, metabolic and mechanical factors and functions. For

example, broad, flat plates, such as scapulae, anchor large

muscle masses, whereas hollow and thick-walled tubes, such as

the femur or radius, support weight. All bone consists of a basic

dual structure, for which their importance varies with the

function. An external layer, or cortex, covers the bone; it is

smooth, continuous and dense (approximately 1.80 g/cm3). In

the interior, cancellous bone is porous with an average porosity

of 75–95% and an average density of 0.2 g/cm3. However, bone

characteristics vary with age and site.

The mechanical properties of bone reconcile high stiffness

and high elasticity in a manner that is not yet possible with

synthetic materials. Cortical bone specimens have been found

to have tensile strength in the range of 78.8–151.0 MPa in

longitudinal direction and 51.0–56.0 MPa in transversal

direction. Bone’s elasticity is also important for its function

giving the ability to the skeleton to withstand impact. Estimates

of modulus of elasticity of bone samples are of the order of

17.0–20.0 GPa in longitudinal direction and of 6.0–13.0 GPa in

the transversal direction [12]. These remarkable mechanical

properties are due to the microstructure of bone combining an

organic matrix with mineral (calcium phosphate apatite)

crystals which are usually oriented in the longitudinal direction

of bone giving higher strength and stiffness in the longitudinal

axis than transversally.

By weight, bone mineral and dentin (inner part of the tooth)

contains approximately 60% mineral, 10% water and about

30% organic matrix (90% type I collagen and 10%

proteoglycans and numerous non collagenous proteins), while

mature enamel is composed of more than 90% mineral and less

than 10% organic matrix (mainly amelogenin). The structure of

human apatitic calcium phosphate biominerals (bone, dentin,

enamel) can be represented as:

Ca10�ðx�uÞðPO4Þ6�xðHPO4 or CO3ÞxðOH; F . . .Þ2�ðx�2uÞ

with 0 � x � 2 and 0 � 2u � x

in which (x � u) cationic vacancies and (x � 2u) monovalent

anionic vacancies coexist.

The u parameter appears generally very small and it can be

neglected. Two types of carbonate ions are present located on

the trivalent and monovalent anionic sites of the structure [13].


For bone and dentin, the proportion of cationic vacancies is

very high and close to the maximum, whereas it is very low for

dental enamel.

Bone and dentin apatites start to nucleate into the nanopores

present in the collagen fibrils [14] as the extracellular fluids are

supersaturated in calcium and phosphates ions. This hetero-

geneous nucleation is catalyzed by the presence of phosphated

esters groups [15] and carboxylate groups [16] at the surface of

collagen fibrils. Along these fibrils, crystallization takes place

to finally interconnect the entire collagen network. These

apatitic crystals are formed of thin plates of irregular shapes.

Their sizes range in length from 20 A for the smallest particles,

to 1100 A for the largest particles [17,18]. This results in a very

large surface area facing extracellular fluids. This property is

critically important for the rapid exchange of ions with these

fluids. The crystallization process of enamel is different than for

bone or dentin: amelogenin being hydrophobic self-assembles

into nanospheres that guide the growth of the ribbon-like dental

enamel crystals. Indeed, enamel crystals are tens of microns

long with an aspect ratio (length/width) of at least 1000 [19].

These physico-chemical differences between these human

apatites are related to their biological function. The collagen–

nanocrystals composite structure of bone meets the mechanical

requirements of the body. In addition, the high specific surface

area [20] and the numerous crystallographic vacancies of bone

crystals, are responsible of numerous ionic exchanges. In fact

bone acts as a calcium and phosphate reservoir for the entire

body. Enamel mineral content is higher than the one of bone,

and its crystalline structure is also more cohesive and stable in

order to resist to acid etching and abrasion often occurring

during mastication. The shape of enamel crystals is also

optimized for a maximal kinetic of growth that may facilitate

remineralization from saliva [13].

2.1.2. Cellular components

Skeletal tissues originate mostly from mesenchymal stem

cells during embryonic development. In the adult stage,

mesenchymal stem cells can be isolated from bone marrow,

adipose tissues, amniotic membrane or umbilical cord

perivascular tissue [21,22]. Mesenchymal stem cells are, by

definition, of self-renewal capacity to repopulate all the

Table 1

Bone cells and their associated functions and markers

Name Function

Osteorpogenitor High proliferative potential

Differentiation in bone lineage

Osteoblast Non proliferative cells

Highly differentiated

Bone mineralization

Osteoclast Bone resorption

Signaling osteoblasts to form new bone

appropriate cell lineages. They are multipotent cells that can

differentiate into osteoblastic, myoblastic, adipogenic, chon-

drogenic, endothelial and neurogenic lineage through a multi-

step differentiation sequence as follows: proliferation, commit-

ment, lineage progression, differentiation and maturation.

Regarding bone tissue, its formation takes place in an organism

during (i) embryonic development, (ii) growth, (iii) remodel-

ing, (iv) fracture healing, and (iv) after ectopic implantation of

osteoinductive matrices. The cells that are involved in

osteogenesis throughout life are summarized in Table 1 [23].

2.1.2.1. Osteoprogenitor cells. With regard to the osteogenic

lineage, mesenchymal stem cells sustain a cascade of

differentiation steps as described by the following sequence:

Mesenchymal stem cell! immature osteoprogenitor! ma-

mature osteoprogenitor! pre-osteoblast! mature osteo-

blast! osteocyte or lining cell. The later the differentiation

stage, the lower the capacity for self renewal and cell

proliferation [23]. In bone marrow, which to date has been

the most common source for harvesting mesenchymal stem

cells, osteoprogenitor cells represent a very small percentage

(less than 0.005%) of nucleated cell types in healthy adult bone

[24]. Differentiating osteoprogenitor cells express several bone

matrix macromolecules, namely alkaline phosphatase, collagen

type I, bone sialoprotein, osteocalcin, osteopontin [25].

2.1.2.2. Osteoblasts. Mature osteoblasts are non-migratory

and highly differentiated cells that can differ substantially in

their properties depending on their stage of development, from

which their function and phenotype can vary and be divided

into four categories: (i) active osteoblasts are cuboidal in shape,

mononuclear and rich in alkaline phosphatase activity. They

synthesize and secrete collagen type I and glycoproteins

(osteopontin, osteocalcin), cytokines and growth factors into a

region of unmineralized matrix (osteoid) between the cell body

and the mineralized matrix [26]. Osteoblasts also produce

calcium phosphate minerals extra- and intracellularly within

vesicles [27]. (ii) osteocytes are mature osteoblasts which have

become trapped within the bone matrix and are responsible for

bone maintenance and homeostasis. (iii) Bone lining cells are

found along the bone surfaces that undergo neither de novo

Markers Origin

Alkaline phosphatase activity Mesenchymal

Type 1 collagen

Bone sialopontin

Osteocalcin

Osteopontin

Alkaline phosphatase activity Mesenchymal

Calcium phosphate

Bone sialopontin

Osteocalcin

Osteopontin

Tartrate resistent acid phosphatase Hematopoietic


bone formation nor resorption. (iv) Inactive osteoblasts are

elongated cells, undistinguishable morphologically form the

bone lining cells. Only active osteoblasts and their precursors

contribute to bone formation (Fig. 1).

2.1.2.3. Osteoclasts. Osteoclasts are derived from hemato-

poietic stem cells that differentiate along the monocyte/

macrophage lineage. They are responsible for bone resorption

by acidification of bone mineral leading to its dissolution, and

by enzymatic degradation of demineralized extracellular bone

matrix. The mature osteoclast is a functionally polarized cell

that attaches via its apical pole to the mineralized bone matrix

by forming a tight ring-like zone of adhesion, known as the

sealing zone. In the resorbing compartment, situated under the

cell and delimited by the sealing zone, osteoclasts generate an

acidic milieu that results in the dissolution of bone mineral.

This osteoclastic acidification is mediated by the action of

carbonic anhydrase that produces carbon dioxide, water and

protons that are extruded across the cell membrane into the

resorbing compartment [26]. During bone remodeling, osteo-

clasts resorb old bone, and via local paracrine signaling

molecules, activate osteoblasts to form new bone.

2.1.3. Biological properties

Bone is a highly vascular, living and dynamic tissue, which

is remarkable for its hardness and regenerative capacity. In its

mineralized matrix, bone embeds osteocytes that constitute the

major cell type in mature bone. Vascular canals ramify within

Fig. 1. Diagram representing the structure of bone (a) and cartilage (b) and their a

apatitic nanocrystals. Osteoblasts derive from mesenchymal stem cells. They are resp

osteocytes. Osteoclasts are responsible for bone resorption, they initiate the mechanis

distribution at the various depths of tissue are shown. A: articulating surface, S:

subchondral bone, CB: cancellous bone (diagram adapted from Woodfield et al. [4

bone, providing its cells with metabolic support. The outer and

inner surfaces of bone are lined by periosteum and endosteum, a

fibrocellular layer, where osteoblasts and osteoclasts are

respectively located. In cancellous bone, the pores are filled

with blood vessels, as well as red and yellow marrow. Red bone

marrow includes cells which participate in the maintenance and

organization of bone, namely osteoprogenitor cells osteoblasts

and osteoclasts. The yellow bone marrow is composed of fat

cells.

In contrast with dental tissue, bone is a very dynamic tissue.

After the initial ossification of the embryonic skeleton,

osteoclasts and osteoblasts begin the modeling and remodeling

processes. In general, modeling refers to alteration in the shape;

whereas remodeling refers to turnover of bone that does not

alter the shape; however, the two processes often occur

simultaneously and the distinctions between them may not be

readily apparent. During skeletal growth, removal and

replacement of bone proceeds at a rapid pace. The rate of

turnover of the skeleton approaches 100% per year in the first

year of life, declining to about 10% per year in late childhood,

and then usually continues at approximately this rate or more

slowly throughout life. Much of the turnover of bone during

growth results from bone-modeling, but presumably at least

some remodeling also occurs. After the completion of skeletal

growth, the turnover of bone results primarily from remodeling.

Modeling and remodeling result from coordinated resorption

and formation of bone over extensive regions of the tissue, over

prolonged periods of time.

ssociation (c). (a) Bone is a composite material associating collagen fibers and

onsible for new matrix formation, in which they are later embedded and become

m of bone remodeling. (b) The orientation of the collagen fibers and chondrocyte

superficial zone, M: middle zone, D: deep zone, CC: calcified cartilage, SC:

8]).


Throughout life, physiological remodeling, removal, and

replacement of bone, at roughly the same location, occur

without affecting the shape or density of the bone, through a

sequence of events that include osteoclast activation, resorption

of bone, osteoblast activation, and formation of new bone at the

site of resorption [28]. Bone remodeling patterns change

dramatically with the age and are illustrated by diseases such as

postmenopausal osteoporosis, which is characterized by bone

mass loss that enhances bone fragility and fracture risks. This

loss of bone density is a result from an imbalance between bone

formation and bone resorption related to estrogen deprivation

[29]. Another example is the modification of bone shape and

density of astronauts, who are subjected to zero-gravity for long

periods of time [30].

2.2. Self repair of hard tissues: example of smart materials

Bone possesses self-regeneration capacity. In fact, in

mammals, the complete reconstitution of a pre-injured state

is a unique feature for bone; all other tissues, with exception of

embryonic tissues, heal with the formation of a scar [31]. The

mechanism of bone regeneration is different than the one of

modeling and remodeling as it involves a preliminary blood

invasion – up to 1 L – at the injured site. The bone-healing

mechanism involves several phases and mechanisms. First,

blood cells (red blood cells and platelets) and fluid proteins

escape the damaged tissue and vessels. A protein-based

network composed of fibrin is formed. Second, through the

emission of signals from the fibrin network and surrounding

cells, other cell types are migrating towards the wound in order

to initiate the formation of early tissues (granular tissues). This

granulation mechanism consists of three phenomena, namely

(i) cellular clearance involving macrophages that remove

damaged entities, (ii) multiplication and invasion of other cell

types that initiate neo vascularization of the wound, and (iii)

early synthesis of a matrix composed of a network of several

proteins such as fibronectin, proteoglycans, hyaluronic acid and

collagen. A callus is formed consisting of cartilage. This

Table 2

Different types of cartilage

Type Where Main fu

Hyaline cartilage Articulating ends of bones Reduce

Trachea Movem

Larynx Suppor

Tip of nose Longitu

Connection between ribs and

Breastbone

Epiphyseal growth plate

Fibrous cartilage Intervertebral disc Shock a

Meniscus Provide

Impedin

Deepen

Disloca

Elastic cartilage Lobe of the ear Maintai

Epiglottis Suppor

Parts of larynx

cartilage hypertrophies and will eventually replaced by

bone. . .. In summary, when bone is injured, several cellular

mechanisms involving signaling, sensing and migration are

activated to reconstitute functional bone tissue. After a

complete healing process, bone shape and mechanical proper-

ties are entirely restored.

There is a limit to the size of fractures and defects that can be

self-repaired by bone. The upper limit is called the critical size

defect, and is defined as a defect of a size that will not heal

during the lifetime of the animal [32]. For larger defects, human

interventions are necessary to completely restore the defect.

3. Soft skeletal tissues

3.1. Structural and compositional organization of cartilage

Normal articular cartilage is hyaline cartilage, which is one

of three types of cartilage found in humans (Table 2). It is

located at the articulating ends of long bones and in the septum,

whereas elastic cartilage and fibrocartilage are located in the

epiglottis and ear, and intervertebral discs and cartilage

menisci, respectively [33]. Articular cartilage structure consists

of four adjacent, interdigitating and organized zones. These are,

from the tissue exposed to the synovial fluid in the direction of

the subchondral bone, the superficial-, middle-, deep- and

calcified zone. Chondrocytes in each zone have different

morphologies and the specific zonal tissue organization and

orientation play distinct roles in cartilage function and

metabolic activities (Fig. 1) [34].

3.1.1. Physical and chemical components

When hydrated, hyaline articular cartilage consists of

approximately 30% extracellular matrix (ECM) proteins,

whereas approximately 70% is water [35]. The ECM consists

predominantly of cartilage-specific proteoglycan (aggrecan)

molecules with highly negatively-charged sulfated glycosami-

noglycan (GAG) side chains, as well as type II collagen fibrils.

Other structural and attachment proteins are also present in

nctions (depending on tissue) Typical markers

s friction at joints Collagen type 2

ent Aggrecan

t Glycosaminoglycans (GAG)

dinal bone growth Chondroitin

Collagen type 10

(hypertrophic chondrocytes only)

bsorbers Collagen type 2

s sturdiness without Collagen type 1

g movement

s sockets to prevent

tion of bones

n shape Collagen type 2

t Aggrecan

Glycosaminoglycans

Chondroitin


ECM in low quantities, such as collagens type IX and, X,

fibronectin and Tenascin-C [34,36,37]. Chondrocyte hypertro-

phy results in a modified pericellular matrix, with increased

secretion of collagen type X and lower levels of collagen type II.

There is a developmental relationship between bone and

cartilage, with cartilage playing a role in the formation of long

bones in vertebrates by ‘endochondral ossification’, a process in

which bones are preshaped in cartilage that is gradually

replaced by bone except for the cartilage found at the

articulating ends of the bone, hence the term ‘articular

cartilage’ [38].

3.1.2. Cellular component

Chondrocytes are the only cell type found in normal articular

cartilage, and are mesenchymal in origin [39]. Chondrocytes

share a common progenitor with osteoblasts: the osteochon-

droprogenitor cell. Chondrocytes contribute to less than 2% of

the wet weight in healthy adult tissue [34]. Under non-

pathological conditions, these cells have a spheroid morphol-

ogy, 10–20 mm in diameter, and are connected to a densely

GAG-rich pericellular matrix.

3.1.3. Biological properties

In close association with bone, articular cartilage insures

joint lubrication and movement. The GAG side chains of

aggrecan are able to bind water molecules, thereby sequestering

water and generating an internal swelling pressure within the

cartilage matrix. These hydrogel-like properties are essential

for the interstitial fluid flow patterns observed inside the matrix

during functional loading of cartilage, at which point water is

forced out of the tissue to an amount that allows the negatively

charged GAG chains to repel each other [40]. Upon release of

the compressive load, water is imbibed back into the tissue

matrix. The collagenous network, together with water-bound

GAG, enables articular cartilage to withstand large compressive

loads which gives the tissue its unique function in the skeletal

system.

3.1.4. Intrinsic cartilage tissue repair

It has been long known that chondral lesions have limited

capacity for self-repair [41]. Like other tissues, cartilage will

undergo an initial phase of necrosis in response to injury, but

inflammation and subsequent vascularization is largely lacking

[42]. Only when the injury reaches the subchondral bone self-

healing processes are initiated by the release of blood-born

factors and mesenchymal progenitor cells from bone marrow

into the wound site. The drilling through to subchondral bone,

allowing the migration of bone marrow in the cartilage lesion is

frequently used to treat chondral defects [3,36]. However, the

resulting tissue is mostly fibrous and not hyaline cartilage. Not

surprisingly, it lacks the mechanical characteristics of normal

articular cartilage and has been reported to begin to degrade

within a few months of the procedure [43]. Other techniques,

such as microfracture and abrasion chondroplasty, also attempt

to utilize the release of progenitor cells and growth factors in

blood into the lesion. They also result in the formation of

fibrous cartilage [3,36,44].

4. Current strategies to repair skeletal tissues

The performance of tissues are the result of millions of years

of evolution, while the performance of the substitution that

mankind has designed to repair tissues are only a few decades

old. The main current strategies to repair tissue are:

(i) G
rafts: a piece of viable tissue or a collection of viable
cells transferred from the donor site to a recipient site for

the purpose of reconstruction of the reconstruction site [1].

(ii) B
iomaterials: materials intended to interface with biolo-
gical systems to evaluate, treat, augment or replace any

tissue, organ or function of the body [1].

(iii) T
issue engineering: the persuasion of the body to heal
itself, through the delivery to the appropriate sites of

molecular signals, cells and supporting extracellular

structures [1].

Biomaterials and grafts are widely used in clinical

applications, while tissue engineering is still at its infancy.

The current strategy in tissue repair is to design hierarchical

constructs able to (i) be accepted by the living body, (ii) restore

the damaged function, and more challengingly, and (iii) react in

a controlled manner with the biological environment in order to

stimulate specific biological mechanisms.

4.1. Biomaterials

Replacing body parts, and specifically hard tissues, dates

back centuries by the use of natural or synthetic materials. For

instance, the Etruscans learned to substitute missing teeth with

bridges made from artificial teeth carved from the bones of

oxen, and in the 17th century a piece of dog skull was

successfully transplanted into the damaged skull of a Dutch

duke. The Chinese recorded the first use of dental amalgam to

repair decayed teeth in the year 659 AD, and the pre-Columbian

civilizations used gold sheets to heal cranial cavities following

trepanation. However, many other implantations failed as a

result of infection or lack of knowledge about toxicity of the

selected materials.

The safe use of materials to replace body parts did not come

into practice until the advance of aseptic surgical techniques at

the end of the 19th century. For decades, attempts have been

made to repair or to replace hard tissues (bone and teeth) by

various means. At first autologous bone was used, but grafting

usually requires a second surgical procedure. To overcome this

shortage, allogeneic bone was taken into consideration, but its

clinical performance is inferior as compared with autologous

bone. In addition in load bearing applications, such as teeth or

hip implants, bulk grafts cannot be used functionally. Instead,

metals and non-degradable ceramics have been used because of

their resistance to fatigue and high tensile strength [45]. Until

the 1960s, materials used to replace body parts were borrowed

from other industrial domains, and some of these are still

widely used. Since the 1960s, materials specifically designed

for body repair have been processed and used in clinical

settings.


Regardless of their composition or application, materials

used for body repair must meet both biofunctionality and

biocompatibility. Biofunctionality concerns the ability of the

implant to perform the purpose for which it was designated.

These requirements are: (i) mechanical properties such as

tensile strength, fracture toughness, elongation at fracture,

fatigue strength, Young’s modulus; (ii) physical properties such

as density in the case of orthopedic implants, or thermal

expansion in the case of bone cement; and (iii) surface

chemistry such as degradation resistance, oxidation, corrosion,

or bone bonding ability [45]. Biocompatibility is defined as the

ability of a material to perform with an appropriate host

response in a specific application [1].

Various types of synthetic substitutes have been developed

in order to comply with biofunctionality and biocompatibility.

They belong to the following main material classes:

(i) M
etals such as titanium, titanium alloys, stainless steel,
cobalt–chromium alloys.

(ii) C
eramics such as aluminum oxide, carbon, calcium
phosphates, glass–ceramics.

(iii) P
olymers such as silicon, poly(methyl methacrylate), poly
lactide, poly (urethane), ultra high molecular weight poly

ethylene.

(iv) C
omposites such as ceramic coating on metal implants, or
ceramic-reinforced polymers.

The choice of one material above another will depend on the

application and the type of function that needs replacement.

Unfortunately, none of the existing biomaterials can meet all of

the requirements. For example, in the case of load bearing

applications (dental or hip implants), the mechanical require-

ments are only met by metals. In particular, titanium alloys are

very promising in orthopedics due to their high specific

strength and low elastic modulus [46]. Titanium exhibits a

strong tendency towards passivation and rapidly forms an

oxide film in the presence of oxygen. Within a millisecond of

exposure to air or aqueous solution, an oxide layer of 1 nm in

thickness will be formed on the surface; and within 1 min, the

oxide film is about 5 nm. This oxide film is very adherent and

stable, and does not break down under normal physiological

conditions. Because of this fast spontaneous formation, there is

no direct contact between the titanium and the host tissue;

instead, they are separated by this thin layer of surface oxide.

The corrosion resistance of titanium towards biological

environment is provided by this protective oxide film [47].

However, titanium has a limited bone-bonding activity

compared to other materials such as calcium phosphates

which have strong bone-bonding capacity but unsatisfactory

mechanical properties (see Table 3).

4.2. Cell-based therapies

In contrast with bone, repair strategies for soft tissue repair

are mainly based mainly on grafts or cell therapies. In articular

cartilage, for example, the most common defects are either

contained within the layer of cartilage articulating the ends of

long bones, or penetrate deeper through the cartilage into the

subchondral bone [3,48]. They are formed due to joint diseases

such as osteoarthritis (OA), rheumatoid arthritis (RA), genetic

and metabolic conditions such as Paget’s disease or from

trauma to the joint [49,50]. Although all the origins of OA have

not completely been identified, it is known that lesions can

ultimately lead to OA. It is, however, not certain that lesions

always are the cause of OA. When OA does arise from

structural damage to the articular cartilage matrix, it does so by

the production of degradative enzymes including a series of

matrix metalloproteases (MMPs) by chondrocytes that results

in the release of proteoglycans from the ECM and erosion of the

collagenous network in cartilage due to an imbalance between

anabolic synthesis and catabolic proteolysis [3,51,52]. It has

also been demonstrated that chondral lesions left untreated for

extended periods (6 months, according to some clinical studies)

may result in joint instability causing the onset OA [50,53–55].

Autologous tissue transplantation can take the form of

‘‘mosaicplasty’’, in which osteochondral plugs from non- or

low-loading regions of the articulating surface are harvested

and transplanted to fill the defects [3,44,56]. However, lack of

suitable donor tissue and donor site morbidity limit the scope of

this technique. Mosaicplasty has also been performed with

allogeneic (cadaveric) osteochondral plugs, although there are

concerns about immunogenic response and the potential for

transfer of infectious diseases. Other tissue transplantation

procedures include the grafting of periosteum and perichon-

drium to defect sites in the articular tissue. In these techniques,

progenitor cells in the cambium are believed to be induced

towards chondrogenesis [36,57]. However, these procedures

fall short of satisfactory functional and histopathological repair

in long term studies due to a number of reasons, including the

lower chondrogenic potential of resident precursor cells in

older patients, who comprise most clinical cases [54,58].

However, it has also been reported that periosteum derived cells

maintain sufficient potential for chondrogenesis for all ages

assessed, if cultured in micromass with the inclusion of TGF-

b1 [59].

Although evaluated in rabbit models as early as 1971 [60],

the implantation of autologous expanded cells in suspension

has seen some growth since the first reported clinical study in

1994 [61–63]. In this procedure, chondrocytes isolated from

articular cartilage biopsies are expanded in vitro prior to being

resuspended in culture medium and injected into the defect site.

The defect is closed with a periosteal flap fixed with sutures to

prevent cells from escaping. However, in vivo studies have

yielded conflicting results. It is as yet unclear whether it is the

transplanted chondrocytes or progenitor cells in the periosteal

flap that are responsible for differences, if any, with respect to

empty defects [64]. Although cell persistence has been reported

[65], there is some uncertainty as to whether the cells remain

within the lesion, since the periosteal flaps have been shown to

detach within a few days after suturing [66]. Some clinical long

term follow-up studies have reported improved outcomes for

treated, as compared to untreated lesions, although in some

cases long recovery times have been cited as drawbacks [61–

63,67,68].

Table 3

Biomaterials clinically used in musculoskeletal repair in the US or Europe

Composition Type Origin Clinical applications Properties

Calcium phosphate,

i.e. hydroxapatite;

tricalcium phosphate;

octacalcium phosphate

Ceramic Synthetic Bone regeneration, non-loading

sites, bone void filler

(cements, granules, coatings)

Bone bonding (bioactivity),

biodegradable, tunability

of degradation

Silica-based calcium phosphate Glass ceramics Synthetic Bone regeneration, non-loading

sites, bone void filler

(granules, coatings)

Bone bonding (bioactivity),

Biodegradable

Alumina Ceramic Synthetic Joint replacement (knee, shoulder) Highest tensile strength,

resistance to fatigue,

non-bone bonding,

lubricating capacity

Titanium and alloys Metal Synthetic Bone replacement, load-bearing

sites, hip or dental prosthesis,

spinal cages

Bone bonding (bioactivity) in

some cases, non-corrosive,

resistance to fatigue, high

specific strength, low elasticity

modulus

Stainless steel Metal Synthetic Bone replacement Corrosive to long term

Cobalt chrome alloys Load-bearing sites, hip or dental

prosthesis, spinal cages, fixations

Polymethylmethacrylate Polymer Synthetic Bone replacement, load-bearing sites,

bone void filler (cement) fixation

of hip prostheses, vertebroplasty

Non-degradable

Polyesthers, i.e. poly lactide,

poly glycolic acid, poly

caprolactone, poly (urethane)

Polymer Synthetic Degradable bone fixation, soft tissue

suture, bone void filler, soft tissue

regeneration*, drug delivery*

Tunability by varying molecular

weight of degradation and

mechanical properties

Ultra high molecular weight

poly ethylene

Polymer Synthetic Articulating component for

orthopaedic prosthesis, load

bearing sites

Lubricating capacity

Poly poly ethylene oxide

terephtalate co

butylene terephtalate

Co-Polymer Synthetic Cement stopper, bone void filler, soft

tissue regeneration*, drug delivery*

Tunability by varying molecular

weight of degradation and

mechanical properties, bioactivity

Polyphosphazene Polymer Synthetic Drug delivery* Erosion degradation mechanism

favorable to long term stability

of the implant

Polyanhydride Polymer Synthetic Drug delivery, hard and soft

tissue repair*

Erosion degradation mechanism

favorable to long term stability

of the implant

Poly ortho esters Polymer Synthetic Food additive, drug delivery*,

hard and soft tissue repair*

Poly ethylene glycol Polymer Synthetic Drug and cosmetic excipient,

hard and soft tissue repair*

Injectable water gel,

degradable

Coral Mineral Natural (sea) Bone void filler High interconnectivity,

degradable

Bone Composite

mineral/proteins

Natural

(human, bovine)

Bone void filler Similar composition as the

host bone

Demineralized bone matrix Proteins Natural (human) Bone void filler, cartilage regeneration* Biodegradable, natural source

of osteoinductive

proteins (BMPs)

Collagen Protein Natural (bovine) Hard and soft tissue repair Biodegradable

Hyaluronic acid Polysaccharide Natural (human) Soft tissues repair Biodegradable, injectable

hydrogel, naturally abundant

in joint synovial fluids

Alginate Polysaccharide Natural (algae) Soft tissue repair Drug load, degradable

Agarose Polysaccharide Natural (algae) Soft tissue repair* Drug load, Degradable

Chitosan Polysaccharide Natural

(sea crustaceans)

Soft tissue repair Structurally similar to

glycosaminoglycans

(cartilage proteins)

Fibrin Protein Natural (human) Soft tissue healing, bone void filler Sealing capacity

(*) Indicates potential future clinical applications based on preclinical data.



To date, both the use of either biomaterials or cells alone is

not sufficient to comply with the highly complex and hierarchal

requirements of our body. We are forced to innovate new and

smarter ways to instruct bodies to heal.

4.3. Towards smart designs to repair tissues

4.3.1. Biomimetic materials for skeletal repair

Learning from nature, in order to synthesize its ‘‘products’’,

is a challenge. In tissue regeneration, nature offers examples of

smart, adaptive and ‘‘strong’’ materials, inspiring us to

synthesize similar performing products, commonly referred

to as biomimicry. The concept of biomimicry is quite novel in

the biomaterials field, but it has been used implicitly for much

longer, in different ways and for diverse purposes [69]:

(i) F
unctional biomimetism is the primary goal in designing
biomaterials where tissue function should be restored. For

example, a biomimetic hip prosthesis should restore the

function of a healthy hip with a very low friction

coefficient between the femoral head and the acetabulum

and with a high resistance to mechanical load (compres-

sion, elasticity).

(ii) M
aterial biomimetism, where restoration of the organ
function is assumed to be obtained if the tissues themselves

are imitated. For example, calcium phosphate ceramics

have been successfully proposed as bone substitutes

because of their chemical similarities with bone mineral

[69].

(iii) B
iological biomimetism, where restoration of the tissue is
synthetically stimulated by implanted cells or molecules

that are involved naturally in the mechanisms of tissue

formation or function.

4.3.1.1. Biomimetic biomaterials for hard tissues. Regarding

bone substitution, the biomimetic synthesis of biomaterials is

based on bone biomineralization that can be described as an

extracellular precipitation under physiological conditions of pH

and temperature. The nucleation and crystal growth are

controlled by (i) spatial delineation by supramolecular

assemblies, (ii) chemical regulation by transport processes,

and (iii) molecular recognition at inorganic–organic interfaces

[74]. It offers several advantageous alternatives in favor to bone

repair. As calcium phosphate ceramics are brittle [75], a bone-

like composite prepared under mild conditions, and composed

of an organic matrix and mineral crystal would open the

possibility to use these bioceramics in load-bearing applica-

tions. Additionally, a mineral phase synthesized under similar

physiological conditions to bone mineral might confer to the

ceramic a reactivity that could positively influence cells and

hosting tissues in favor to new bone formation. Indeed, the

biological apatitic crystals are highly reactive due to their

hydrated, poorly organized outer surface [20,76,77]. In

particular, the presence of non-apatitic phosphates in these

biological crystals affect their biological properties [76]. In that

sense, any calcium phosphate ceramic obtained by a classical

process involving high temperature cannot imitate this very

specific reactivity. The biomimetic synthesis of calcium

phosphate biomaterials under physiological conditions can

lead to new calcium phosphate ceramics with very different

physico-chemical properties, affecting their reactivity towards

the biological surroundings. The biomimetic method has been

explored and led to coatings on implants substrates or powdery

compounds.

Simulated body fluids (SBF), supersaturated with respect to

apatite, have been developed and used to synthesize ‘‘bone-like

crystals’’ and other calcium phosphate phases that are

considered bone mineral precursors [14,78,79]. As with bone

mineral deposition, this biomimetic process requires an

heterogeneous substrate onto which nucleation and crystal-

lization will occur over time from fluids supersaturated with

respect to apatitic phase [80]. Some strategies have been

developed to synthesize the biomimetic biomaterials by:

(i) C
hemical regulation by transport processes from synthetic
extracellular fluids, such as simulated body fluids [81] and

derived solutions [82–84]. This synthesis led to the coating

of various biomaterials in order to enhance their bone-

bonding ability [85–87]. More interestingly, biomimetic

synthesis opens broad new possibilities for biomaterials

functionalization. First, the deposition of calcium phos-

phate layers on porous and/or thermosensitive scaffolds

triggering bone formation, ingrowth and contact with the

implant [87–90]; second, the co-precipitation of organic

molecules like the osteoinductive bone morphogenetic

protein (BMP-2) or antibiotics [91–93], third, a fine control

of the surface physicochemistry at the molecular and nano-

metric scale that can influence in vitro and in vivo the

biological response [78,94].

(ii) M
olecular recognition at inorganic–organic interfaces. The
grafting of specific functions onto biomaterials stimulating

self-assembling of ions have been explored [95,96]. In both

cases, nucleation and crystallization mechanisms were

enhanced by using biomimetic principles. Chemical

regulations by transport processes and spatial delineation

by supramolecular assemblies have been also exploited for

the synthesis of bulk biomaterials. As such, these strategies

led to the formation of powdery samples [97]. No bulk

biomimetic biomaterials could be produced in the size

range of ‘‘classical’’ bone biomaterials (from circa

0.125 cm3 and bigger). However, with the aid of modern

techniques such as rapid prototyping or free-form

manufacturing bulk bioceramics with complex shapes

and relevant sizes can be nowadays obtained [98,99].

4.3.1.2. Biomimetic biomaterials for soft tissues. Polymers

can also be fused with bioactive peptides to create biologically-

responsive hydrogels. This concept has been illustrated for the

repair of various tissues, including bone [100–104] and

cartilage [105]. This approach seeks to mimic the tissue

extracellular matrix with hydrogel matrices that incorporate

various biological signaling moieties, such as arginine–

glycine–aspartic acid (RGD) peptides or laminin-derived

peptide IKVAV, or linkers that are susceptible to cell-secreted


proteolytic factors, such as MMPs or plasmin [106–110]. These

quasi-ECMs allow for the spatial as well as temporal regulation

of cell migration and matrix degradation that is uniquely

determined by the cells contained within the gels. These are

man-made examples of bidirectional signaling between cells

and their ECM, that mimic (to a lesser extent) signaling from

the ECM to the cell (outside in) and vice versa (inside out), that

are crucial in maintaining cell viability, phenotype, migration

and production of ECM proteins. Specific examples include

poly(ethylene glycol)-bis-vinylsulfone polymers containing

peptide sequences with three cysteine residues [111]. In this

system, cell invasion in the hydrogel occurred when the

network was crosslinked by plasmin-sensitive peptides and

contained cell adhesion peptides. Conversely, hydrogel net-

works lacking either of these features did not allow cell

infiltration. PEG and N-(2-hydroxypropyl)-methacrylamide

(HPMA) are often used as physical or chemical crosslinkers,

with PEG impacting the physical properties of the hydrogel

[112].

The fibrillar architecture of the ECM has lead to the design

of polymers that mimic scale, topography and shape. These

structures can now be fabricated at the nanometer scale, which

is essential for providing signals to cells that mimic the natural

biological microenvironment. Techniques used for these

include electrospinning and molecular self assembly based

on non-covalent interaction between amphiphilic peptides,

synthetic polymers or oligonucleotides, and have been used to

demonstrate control of biologic function in chondrocytes, liver

hepatocytes and neural cells [113–117].

Biomimetic synthesis is an emerging promising field that

needs to demonstrate clinical benefit over ‘‘classical’’

approaches.

4.3.2. Tissue engineering

The last two decades have seen a surge in creative ideas and

technologies developed to tackle the problem of repairing or

replacing diseased and damaged tissues, leading to the

emergence of a new field in healthcare technology now

referred to as tissue engineering. It has been defined by some as

‘‘an interdisciplinary field that applies principles and methods

of engineering and the life sciences toward the development of

biological substitutes that restore, maintain and improve the

function of damaged tissues and organs’’ [118].

The earliest tissue engineering strategies were reported for

the delivery of encapsulated beta-islet cells in diabetic patients

(1975) [119], dermal regeneration by applying a glycosami-

noglycan-collagen composite matrix over burn sites (1980)

[120] and a collagen matrix impregnated with fibroblasts to

also aid in skin repair and regeneration (1983) [121]. This was

followed by the development of porous synthetic polymer

scaffolds to deliver cells and neo-tissue, a strategy that has

since then been evaluated for most tissues [122,123]. For

example, tissue engineering strategies for repairing different

tissues such as cartilage [3,50,52], bone [12,56,124], tendon

[125,126], liver [127], cardiac muscle [128–130] and neural

tissue [131,132] have been described and reviewed elsewhere

in detail.

The modern concept of tissue engineering draws on multiple

areas of expertise. It is a convergence of the fields of materials

science and engineering, cellular and molecular biology,

biochemistry, controlled release systems for biotherapeutics,

and surgery. Briefly, cells are removed from a patient via a

tissue biopsy performed by a surgeon or physician. Once, after

cell expansion, a sufficient number of cells have been reached

for the intended application, they are transplanted back into the

patient using either natural materials, synthetic polymers,

ceramics, or composites for cell delivery. These three-

dimensional structures can be formed in the shape of the

defect to be replaced, or be delivered as injectable gels to fill the

defect site in situ. They allow cells isolated from the biopsy to

produce the structural proteins found in the original tissue, such

as collagen. The neo-tissue can be grown in vitro in the scaffold,

eventually reaching a stage where it is suitable for implantation

back into the patient to replace damaged or diseased tissue or be

allowed to grow in situ within the defect. Fig. 2 schematically

illustrates an example of the tissue engineering process.

Recent advances in materials science, including photo-

polymerizable gels and ‘‘smart’’ materials that can change

surface chemistry based on the local electro- and biochemical

environment have greatly added to the options available to

tissue engineers [133,134]. On the other hand, rapid progress in

our understanding of the cellular and molecular events involved

in the natural development of tissues, such as those that occur

during the ultimate example of tissue engineering – the

embryonic development of the fetus – have opened the

possibility for therapy using pluripotent stem cells [79,135–

137].

In spite of the rapid developments taking place, the field of

tissue engineering is still in its infancy. The only true

commercially available products at this point are skin grafts

and cartilage repair procedures. However, this technology is

likely to benefit as further convergence of nanotechnology,

proteomics, gene therapy and stem cell research allows tissue

engineers to repair and perhaps regenerate tissues and organs,

with the potential to greatly improve the lives of people with

functional disabilities [138].

4.3.3. Materials for tissue engineering

Materials selected for tissue engineering must interact with

cells and culture media in vitro prior to their implantation. In

addition, these scaffolds and matrices, have to (i) host a

sufficient amount of cells and (ii) support their viability for

several weeks. Scaffolds refer to solid porous structures, while

matrices refer to gel-like structures.

4.3.3.1. Natural materials. Nature offers several potential

sources of materials for tissue regeneration scaffolds. Although

they have been shown to support cell viability and tissue

formation to various degrees, natural matrices such as

hyaluronan (hyaluronic acid) and fibrin are unable to endure

body’s loading conditions. Furthermore, the defect borders may

degrade if the defect is not filled with an implant of similar

mechanical properties [55], and esterified hyaluronan and fibrin

have sometimes demonstrated undesirable effects on the host


tissue and its repair [139–142]. Another group of natural

polymers that have been investigated for this purpose comprises

polysaccharides like alginate [143,144] and chitosan [145,146].

These materials have been studied for cartilage and skin

regeneration [147–149] and reached clinical applications for

the treatment of deep skin wounds and cartilage [150,151].

Chitosan has satisfactory biocompatibility, but mechanical

loading issues remain [152]. Agarose, alginate and collagen

hydrogels have been studied for their potential as injectable

materials in which chondrocytes or precursor cells can be

Fig. 2. Diagram representing current tissue engineering strategies. (a) Cells are ha

proliferate in vitro prior to (c) seeding in 3D scaffolds, and cultivated in bioreactor

prepare scaffold consists of design via 3D computer aided technique from a 3D-CT s

fibre deposition can realize bulk scaffolds (herewith an articular joint). (e) The final c

takes place.

mixed or embedded. Alginate consists of two repeating

monosaccharide units, L-gluronic and D-mannuronic acids,

which are water soluble and jellify when exposed to calcium

ions. Chitosan is structurally similar to glycosaminoglycan

(GAG) and is composed of b linked D-glucosamine residues. It

has been recently attracted more and more attention because of

its non-toxicity, bioresorbability, and wound healing abilities

[153]. Hyaluronic acid is also abundantly present in the human

body articulation, within the synovium fluid. However, in its

natural form this material lacks some desirable properties (too

rvested from tissue and isolated in vitro. (b) These cells can then be made to

conditions to produce viable constructs. (d) The most sophisticated strategy to

cans of the body part to be replaced. Solid free-form fabrication, including 3D-

ell-scaffold construct is implanted back into the patient, where tissue formation

Fig. 2. (Continued ).


high water solubility, fast resorption and tissue clearance times)

to consider it as a polymer for scaffold fabrication [154]. A

change of its chemical structure through an esterification

reaction allows the generation of a new set of biomaterials,

hyaluronan, with improved properties, increased biocompat-

ibility and fine-tunable degradation rates [155,147,156].

Although these materials allow tissue formation within the

gel, they too are limited with respect to the mechanical loads

they are able to withstand. Collagen has also been studied in

cross linked, hardened gel and lyophilized states, with some

success [157–159] and for various tissue regenerations [160–

164]. In particular crosslinked collagen type I and type II

scaffolds alone, or in combination with glycosaminoglycans

have been considered for bone and cartilage repair [163–166].

However, their gel nature seems to prevent the cell migration

within the matrix, reducing so the tissue repair [61]. A further

possibility is to use denatured collagen (gelatin) [161], fibrin

[167,168] or demineralized bone matrix (DBM) [169–171].

DBM is the organic phase of bone tissue after thorough

demineralization processing. The bioactive proteins, namely

bone morphogenetic proteins (BMPs) and some other growth

factors that are bound to the extracellular collagen matrix, are



made more available to the surrounding tissues [172]. However,

the demineralization process can affect strongly the proteins

conformation and also the organic matrix tri-dimensional

organization, affecting therefore the osteogenic potential of

these compounds [173]. Natural minerals, such as coral- and

bovine bone derived materials are also used in bone tissue

engineering; they provide an efficient porosity and intercon-

nectivity for cell and nutrient penetration in combination with

elevated mechanical properties.

4.3.3.2. Synthetic materials. Several materials and matrices

have been evaluated for tissue engineering purposes [3,48].

However, there are considerable issues associated with the

choice of carrier matrix, among which are biocompatibility of

the material, its degradation products and the ability of the

matrix to withstand mechanical loading. Biomaterial properties

that influence cell behavior are addressed in later this review.

Synthetic biomaterials containing calcium and phosphate

groups, such as calcium phosphate ceramics and calcium

phosphate silica glasses (or bioactive glasses) exhibit excellent

bone-bonding properties that are related to the surface

reactivity, via dissolution–precipitation mechanisms. In addi-

tion, they degrade naturally: the degradation products are

entirely metabolized in a natural way by our bodies [70,71].

These bone-bonding and degradation features are unique and

have contributed to their clinical success for 40 years [72,73].

The first clinical attempt to use calcium phosphate compound

was in the successful repair of a bony defect reported by Albee

in 1920 [73]. Since then, several calcium phosphate biomater-

ials have been developed and successfully applied in the clinic,

such as hydroxyapatite, tricalcium phosphate, bicalcium

phosphate, brushite and octacalcium phosphate. As porous

structure they are good candidates as scaffolds for bone tissue

engineering.

Synthetic polymer hydrogels have shown promise as

injectable gels and for embedding chondrocytes, although

their use too has been limited by biocompatibility issues and

often low mechanical stability under loading [3,48]. In

addition, hydrogels are not favorable to migratory cells often

used in bone tissue engineering [174]. But advances in polymer

technology have enabled these issues to be addressed, while

preserving their beneficial features [175]. Furthermore, these

advances have also enabled the generation of hydrogels that are

photopolymerizable, pH and thermoresponsive [133,176]. The

polymers most extensively tested for connective tissue repair

belong to the poly(a-hydroxy esters) family and include

poly(glycolic acid) (PGA), poly(lactic acid) (PLA) and their

copolymers (PLGA), with much of the focus on non-woven

PGA fiber meshes since this polymer was shown to support

improved cartilage matrix and tissue formation over PLA or

PLGA [177–183]. However, these polymers biodegrade to

produce acidic degradation products that can have adverse

effects [184,185]. Additionally, a key issue in tissue engineer-

ing is the ability of the tissue engineered construct to integrate

with the surrounding tissue at the defect site. This is an

especially relevant point in situations where relatively mature

cartilage tissue has been grown in vitro prior to construct

implantation. An approach to deal with this issue is to

transplant constructs with less mature tissue at an earlier stage,

with the reasoning that most of the tissue formation should

occur in situ and thus enable better integration with the

surrounding host tissue [186]. However, the lack of mechanical

stability of fibrous PGA as described earlier precludes their use

in applications in which chondrocytes, or progenitor cells, are

to be transplanted to defects at an early stage. Composite PLGA

scaffolds reinforced with PGA fibers have also been designed

and evaluated to compensate for shortcomings of the individual

components, with success in in vivo osteochondral defect


models [187,188]. A recent study using expanded chondrocytes

cultivated on hyaluronan scaffolds for the repair of articular

cartilage lesions demonstrated considerably positive results in

terms of histology and biological acceptability of neo-tissue as

well as functionality and patient quality of life [68].

Despite the apparent simplicity of this single cell-type

tissue, it is clear that current techniques for repairing skeletal

tissues are frequently unsatisfactory, requiring the develop-

ment and implementation of novel technologies and

procedures for achieving the desired clinical objectives.

Among the techniques under development for repairing

damaged tissues is the immobilization and transplantation of

cells on carrier scaffolds, or tissue that has been grown in

vitro, to fill defect sites. Cells (mostly autologous, although

allogeneic sources have also been studied) can then be seeded

either directly onto carrier scaffolds, or be proliferated in vitro

prior to seeding onto scaffolds, with the latter case being more

realistic due to the limited number of cells that can be

harvested from tissue biopsies. Depending on the length of

culture and cultivation conditions, tissues may be grown to

varying degrees within the carrier scaffolds prior to

implantation into defect sites [3,48,52,56,68,189]. To that

end, extensive research has been done in evaluating a range of

cell sources, materials for cell transplantation as well as

scaffold designs that can be used to immobilize cells and

support neo-tissue synthesis.

Among the strategies developed to improve tissue engineer-

ing therapies, selecting, designing and engineering scaffolds

and matrices are of utmost relevance. Depending on their

intrinsic properties, such as surface properties and porosity,

some biomaterials have shown interesting instructive properties

with regard to tissue formation.

5. Instructing physico-chemical and biological

processes at biomaterial interfaces

There is a great gap in the complexity, hierarchy and

intelligence between tissues and their potential substitutes. A

detailed understanding of the interactions between biomater-

Fig. 3. Physico-chemistry of the biomaterial in con

ials, cells, fluids and tissues is mandatory. With this knowledge,

we will be able to engineer smarter materials capable of

instructing their biological milieu for a complete tissue

regeneration.In vivo, the interactions between the implant

and its ‘‘biological surrounding’’ occur in non-equilibrium

conditions (Fig. 3). The amount of compounds playing a role in

these interactions remains undefined. Regardless the nature of

the biomaterial, its biological surrounding evolves with time.

For example, Dhert et al. have described the events taking place

between different biomaterial implants in contact with the host

bone. In the first 3 days, blood invaded all of the empty

interstitial space between the host bone and implant. At the end

of the first week of implantation, callus and mesenchymal

tissues entirely replaced blood, concomitant with host bone

resorption. Finally, between the second and fourth weeks of

implantation, callus, mesenchymal tissues and host bone

gradually disappear in favor of newly formed bone while bone

remodeling takes place [205].

The following sections will review the newest concepts on

how to design smart implants for hard and soft skeletal tissues.

5.1. Osteoinductive biomaterials

A first striking example is relative to osteoinductive

biomaterials. In non-bone sites, some synthetic materials have

shown the ability to induce bone formation (osteoinduction) as

illustrated in Fig. 4c. At a first glance, the ability of a material to

form bone in a non-bony environment may not be relevant in

bone tissue regeneration. However, it has been recently

demonstrated that osteoinductive biomaterials stimulate more

bone formation when used as tissue engineered constructs in

goat muscle [209] as well as in critical size orthopedic defect

models in goats [210]. Therefore understanding osteoinduction

mechanism by biomaterials could offer a useful tool to repair

large bone defect.

Polymeric, metallic and ceramic biomaterials have shown

osteoinductive properties. In 1969, polyhydroxyethylmethyl-

methacrylate (poly-HEMA) sponges were reported to induce to

form bone in soft tissues [211]. In the last decade, calcium

tact with biological milieu (in vitro or in vivo)

Fig. 4. Bone-instructive implants. (a, b) Back-scattering electronic microscopy picture of an calcium phosphate coating (OCP) coating on a metallic porous scaffold

implanted for 12 weeks in the femoral condyle (goat) at different magnification (scale bar (a) 50 mm and (b) 10 mm). Between the OCP and the newly formed bone, an

interfacial phase (arrow) that can be attributed to superficial phase transformation is clearly visible. (c) Light-micrographs of a macroporous osteoinductive scaffold

(BCP) implanted in goat muscle for 12 weeks representing ectopic bone (dark grey) (magnification �20) (courtesy P. Habibovic).


phosphates [85,210,212–222], alumina ceramics [223], tita-

nium [197,224] and glass ceramics [225] were also found

osteoinductive in soft tissues of different animals. Although the

mechanisms of osteoinduction by materials remain unclear,

critical physico-chemical factors of the biomaterials have been

identified: the macroporosity (pores larger than 10 mm), the

surface physico-chemistry, and the microporosity (pores

smaller than 10 mm) [210].

Regardless of their composition, all osteoinductive biomater-

ials exhibit macropores, whether they are initially present such as

macroporous structures [197,211,214,215,222,223,225,226] or

formed during the biomaterial’s degradation process (for

example, crevices at the surface of calcium phosphate cements)

[212,217]. Macropore geometry and tridimensional organization

are important criteria. Among several well-defined cavities

created on hydroxyapatite, only specific concavities were shown

to induce bone formation [226]. A comparative study between

two different macroporous structures of the same titanium

showed that only the complex tridimensional one was

osteoinductive while a regular fiber mesh did not induce bone

formation [197].

Osteoinduction has been found among various biomaterials

exhibiting very different surface properties. For example, the

calcium phosphates and bioactive silica glasses are known to

undergo surface transformation via a dissolution–reprecipita-

tion process favorable to bone-bonding. These ionic exchanges

properties of the calcium phosphate scaffolds with the

surrounding milieu have been pointed out as a relevant

parameter among others [210,215]. Titanium and poly (HEMA)

materials possess also the ability to nucleate a calcium

phosphate layer in vitro and in vivo [211,227–230]. This

calcification property has been proposed to be a precursor in the

osteoinduction mechanism [222]. However, alumina ceramics

that have been reported osteoinductive [223] are well-known

for their poor calcium phosphate nucleation ability [190,231].

They are used as knee joints because of their long-lasting

lubricant properties related to their resistance to calcification.

In addition, under physiological conditions of pH and

temperature, alumina surfaces are positively charged, while

titanium and calcium phosphate ceramics are slightly

negatively charged. These differences in surface charge will

also affect the interaction pattern with organic molecules

presents in the body fluids [232,233] and with cells. In view of

these large physico-chemical differences among the surfaces of

osteoinductive biomaterials, several different phenomena may

intervene in parallel.

By implanting intramuscularly in goats two macroporous

calcium phosphate scaffolds identical in composition, crystal-

linity and porosity but with different microporosities (pores

smaller than 10 mm), Habibovic have demonstrated that an

elevated microporosity was responsible for ectopic bone

formation [210]. This high microporosity is directly correlated

to roughness and exposed surface, and therefore reactivity.

Other types of osteoinductive biomaterials also exhibited an


elevated microporosity that may be responsible for triggering

cell differentiation towards the osteogenic lineage. In addition,

at specific surfaces, greater ionic and molecular exchanges

occur between the biomaterial and the biological milieu. A high

concentration of specific molecules (for example, osteoinduc-

tive BMPs), adsorbed directly at the biomaterial surface or co-

precipitated into the newly formed calcium phosphate layer,

favor bone formation [226].

Osteoinductive materials for bone demonstrate that bioma-

terials can specifically instruct the body to heal by itself.

Thorough studies have been conducted in order to study the

specific factors affecting these bone-instructive properties.

Osteoinduction mechanism remain unraveled but it has led to

the recognition of several biomaterials parameters that can be

tuned in order to guide the biological response.

The following section will detail the different parameters

affecting hard and soft skeletal tissue formation, and the smart

designs that have been derived from these.

5.2. Orchestrating biomaterials degradation with new

tissue formation

In the presence of fluids and cells, biomaterials undergo

degradation. In other words, their chemical structure, physical

properties, or appearance change [1]. The degradation

mechanisms depend on the nature of the biomaterials. While

calcium phosphate and bioactive glasses degrade by dissolution

and cellular (osteoclastic) mechanisms, polymers degrade by

chain scission, erosion and metabolization mechanisms [234].

5.2.1. Degradation mechanisms

5.2.1.1. Ceramics. In contact with biological fluids, calcium

phosphate ceramics degrade via dissolution–reprecipitation

mechanisms [203]. Ionic transfers occur from the solid phase to

the aqueous liquid via surface hydration of calcium, inorganic

phosphate species, and possible impurities like carbonate,

fluoride or chloride present in the biomaterial. Under

physiological conditions, this dissolution process is highly

dependent on the nature of the calcium phosphate substrate and

their thermodynamic stability, for example (in order of

increasing solubility), hydroxyapatite (HA) > tricalcium phos-

phate (TCP) > octacalcium phosphate (OCP) > bicalcium

phosphate dihydrate (DCPD) [190,194,199,235–238]. The

composition and supersaturation of the environment in vitro

[239–241], or the implantation site in vivo [242,243] also

influence the dissolution–reprecipitation mechanism. Ionic

transfers occur also from the surrounding fluids to the calcium

phosphate substrate in vitro and in vivo, as illustrated by the

formation of carbonated apatite nanocrystals as a result of

surface transformation [190–194,242,244] (Fig. 4a and b). The

presence of magnesium and carbonate contributes to the

formation of a poorly crystallized carbonated apatite that has

similar features with bone mineral phase [245–247]. In the

presence of proteins, this newly formed mineral phase is also

associated with organic compounds [190,192,195]. This phase

transformation occurs for all of the calcium phosphate

bioceramics, even the stable apatitic structures since they have

a strong ability to adapt to their environment by hosting foreign

ions and subsequently to undergo atomic rearrangements [20].

However crystalline hydroxyapatitic substrates are often too

stable to transform. The result of these ionic exchanges

favoring either phase transformation or dissolution follows

thermodynamic stability. This surface reactivity has pivotal

implications in the development of instructive functions,

namely osteoinduction and drug delivery systems. This will

be elaborated on later in this review.

5.2.1.2. Polymers. Biocompatible polymeric biomaterials

have also the potential to degrade in contact with biological

fluids [248]. Briefly, linear aliphatic polyesters such as

poly(lactic acid), poly(glycolic acid) and copolymers have

been broadly used, and their degradation rate can be tailored by

varying their copolymer ratio [249–253]. Their degradation

products (lactic and glycolic acids) obtained by hydrolysis are

normally present in the metabolic pathways of the human body.

However, their bulk degradation leads to the build-up of acidic

degradation products inside the matrix lowers the pH within the

polymeric matrix. This might result in local inflammation in

tissues if clearance of degradation products is insufficient [254].

Another family of thermoplastic polymers that has been

recently studied for bone and cartilage repair and regeneration

in view of tunable degradation properties is poly(ethylene

glycol)-terephthalate–co-poly(butylene terephthalate) (PEGT/

PBT). These polyether-ester multiblock copolymers belong to a

class of materials known as thermoplastic elastomers which

exhibit good physical properties like elasticity, toughness and

strength [92] which are critical for reconstructing load-bearing

tissues. By varying the molecular weight of the starting PEG

segments and the weight ratio of PEGT and PBT blocks it is

possible to tailor their biodegradation rate [255]. Being

polyether-esters, degradation occurs in aqueous media by

hydrolysis and oxidation, the rate of which varies from very low

(high PBT contents) to medium and high (larger contents of

PEGT and longer PEG segments) [92,255]. Among the

multitude of other synthetic polymers investigated for tissue

regeneration, interesting classes are polyphosphoesters [256],

polyphosphazenes [257–260], polyanhydrides [261] and poly-

ortho-esters [262].

5.2.2. Smart degradation designs

Controlling the degradation kinetics of biomaterials to

match tissue growth, to create space for the new tissues to grow

until full regeneration is reached remains a challenge in

biomaterial design. However, tools to tune the degradation

mechanism in favor to tissue regeneration are now becoming

available to researchers.

In large defects which cannot be healed naturally by bone,

adjusting the degradation kinetics of the calcium phosphate

bone filler to the kinetics of bone formation rate is possible by

changing the calcium phosphate phases of the biomaterials.

Mixing at various ratios a low soluble phase (HA) with a highly

soluble phase (amorphous, TCP) resulting in biphasic calcium

phosphates ceramics (BCP) [202,263,264], including additives

(magnesium, carbonate, fluoride) in a given crystalline phase


[190,199,265], or selecting different calcium phosphate phases

(amorphous, DCPD, OCP, HA, TCP) [190,265] are the options

to tailor the degradation kinetics of calcium ceramics, ranging

from weeks to years.

In view of the dissolution tunability of calcium phosphates,

several groups have used these ceramics as delivery systems for

gene [93] or drugs [91,92,266,267]. Drug association with

calcium phosphates can be performed by (i) adsorption on

powder followed by compaction, (ii) co-precipitation or (iii)

addition in the cement paste of compounds for bone

regeneration. With regard to the stimulation of bone regenera-

tion specific proteins have been administrated via calcium

phosphate carriers. Bone morphogenetic proteins (BMP,

especially BMP-2) adsorbed onto ceramics [267,268] or co-

precipitated with carbonated apatite coatings [91] induce more

bone formation than ceramics alone in vivo. Recently,

incorporation of silicate and zinc ions, respectively in

tricalcium phosphate and hydroxyapatite ceramics, were

reported to have a significant influence on osteogenesis in

vitro and in vivo [260–271]. However, the mechanisms of

release profile versus the nature of the calcium phosphate

remains poorly investigated.

Degradable polymeric biomaterials have been extensively

used for controlled drug delivery systems. The therapeutic

molecular cargo can either be incorporated directly in the

biomaterial matrix, or added to a prefabricated biomaterial by

mean of microspheres [272–274] or coatings [275]. PLA, PGA,

PLGA have already been studied for drug delivery [272,276–

282]. The release rate of incorporated proteins is linked to the

degradation rate of the polymer. However, it is possible for the

activity of the incorporated protein to decrease, due to its

denaturing in view the degradation mechanism of this class of

polymers [283–285]. For long term release, another linear

aliphatic polyester commonly used in tissue engineering is

poly(e-caprolactone) (PCL) is attractive as its degradation rate

is slow compared to other common biocompatible polymers

[174,256,286,287]. PEGT/PBT biomaterials allow the embed-

ding of proteins in their matrix [92]. The release mechanism is

due to a combination of protein diffusion and matrix

degradation, which allows zero-order release profiles over

long time period. A further modulation in degradation rate and

protein release profile can be achieved by substituting part or all

of the terephthalate groups with succinate blocks during the

copolymerization reaction [288–290].

In gels, bioactive agents can be easily incorporated [74,291].

However the bioactivity of these agents may be hampered when

they are combined with in situ polymerizable hydrogels due to

the exposure to ultraviolet light and crosslinking agents which

can induce protein denaturing or aggregation, and decrease the

activity of encapsulated proteins [292]. Natural polymers jellify

chemically, they can readily serve as drug delivery systems. For

example, collagen and polysaccharide matrices allow the

release of proteins and growth factors by diffusion, degradation

of the matrix [293–303,16,304].

In general, the incorporation of bioactive agents in

polymeric biomaterials can be achieved by dispersing the

protein in the polymer phase prior to scaffold processing, using

two main approaches: (i) adding the signaling molecule directly

to the polymer solution or powder [281] or (ii) a water phase

containing the protein can be mixed with a polymer dissolved in

an organic solvent to form a water-in-oil (w/o) emulsion

[273,276]. However, this may induce a loss of activity due to

protein denaturing [285]. The association of growth factors to

porous biomaterials is usually achieved by separating the

scaffold preparation step from the protein incorporation, to

reduce the detrimental effect of scaffold processing on protein.

Growth factor loaded microspheres or liposomes were

incorporated in hydrogels or prefabricated scaffolds

[168,301,305], polymer coating applied on compression

molding scaffolds [275] or, more often, prefabricated matrices

were soaked with growth factor solutions [304,306,307].

Although adsorption of the growth factors by soaking seems the

easiest and less harmful approach, it limits the possibility to

control the release of growth factors from the scaffolds. In

addition, it was demonstrated that adsorption could also result

in protein denaturing [308].

Studies have confirmed the potential of local controlled drug

delivery systems for tissue repair and regeneration. In vitro, the

sustained delivery of two growth factors, TGF-b1 and IGF-1,

supported cartilage repair and maintenance [274,300,301,305].

In vivo, the beneficial effect of growth factors sustained release

was as well demonstrated. In rabbit osteochondral defects, the

release of TGF-b1 over at least 5 days from alginate

microparticles [298] or a release of BMP-2 within 10 days

from collagen sponges were evaluated [309,310]. An improve-

ment of the tissue repair after 6, 12 or 24 weeks was measured,

in comparison to defects filled with unloaded matrix or left

empty. BMP-2 delivery showed similar cartilage restoration as

compared to the implantation of autologous chondrocytes in the

defect. This indicates the potency of the released growth factors

to differentiate progenitor cells present at the implant site,

which may eliminate the need of an extra cell source for

transplantation. However, recent studies in rabbit osteochondral

defects with scaffolds releasing TGF-b1 at similar concentra-

tions showed either only a limited improvement of cartilage

restoration [311] or no improvement when released over 12

days [312]. The same negative result was found in chondral

defects exposed to IGF-1-releasing liposomes, possibly

because of a wrong dosage or release rate which was not

evaluated, or to the lack of suitable progenitor cells [167].

Thorough systematic studies have shown the biological

sensitivity to degradation and drug release kinetics. However,

controlled and well-defined delivery systems to regenerate

tissues still have to be proven in clinics as tissue growth varies

between individuals, species and location of the tissues.

5.3. The role of proteins in regulating biomaterial-induced

biological response

Several properties have been proposed as potential

regulators of cell behavior including wettability (as measured

by water contact angle), surface chemistry, equilibrium water

content, surface flexibility and roughness [313–315]. However,

it is likely that a combination of these factors collectively


influences cell function, possibly by modulating the adsorption

of proteins onto attachment substrates. Although the natural

state of most proteins is in an aqueous environment, the contact

of a protein solution and a solid phase with which it is

immiscible leads to the accumulation of the protein between the

two phases [316]. The physicochemical properties of the

protein and chemistry of the solid phase would then determine

whether the protein adsorbs to the substrate. This is the scenario

when seeding and culturing cells on biomaterials in serum-

containing media conditions. Thus, in addition to direct cell–

material interaction, it is plausible that cells would also sense

the biochemical environment of the adsorbed proteins and that

this substrate-protein environment could markedly influence

cell behavior [317].

5.3.1. Protein adsorption onto biomaterials

Mammalian cells have long been cultivated on organic

polymers substrates for basic cell research [318–320] and the

advent of cell therapy and tissue engineering have reinforced

the need to understand cell–biomaterial substrate interactions

[3,48,123]. Among the techniques used to modulate protein

adsorption is the inclusion of specific molecules such as PEG

polymer chains at the surface of biomaterials. Incorporating

PEG increases hydrophilicity and mobility of surface

molecules [321,322]. It is achieved by grafting PEG to

hydrophobic surfaces, simple PEG adsorption or by synthesiz-

ing block copolymers containing PEG components [323].

When exposed to water, PEG chains become highly mobile to

form a large excluded volume in a liquid-like state [324]. PEG

length correlates directly with the volume of protein adsorp-

tion-inhibiting mobile water molecules present at the surface

[324].

The primary serum proteins believed to be involved in cell

attachment and growth on polymer substrates are fibronectin

(Fn) and vitronectin (Vn) and several reports have shown that

Vn, rather than Fn, preferentially adsorbs to some substrates

from serum [325–327]. It was also previously reported that Vn

adsorption was not inhibited by increasing PEG concentration

at a surface, regardless of substrate wettability [325] and that

Vn adsorbed equally well to untreated polystyrene (PS) as it did

to surface modified tissue culture polystyrene (TCPS) [328].

The presence of serum proteins has been demonstrated to be

more effective in inhibiting Fn adsorption to bacteriological

grade PS, as compared to TCPS [329]. Since the adsorption of

Fn to substrates has been shown to correlate with cell

attachment, this suggests that glow discharge treatment of

surfaces enhances cell adhesion by allowing increased Fn

adsorption from serum. However, TCPS is more hydrophilic

than bacteriological grade PS because treatment with glow

discharge increases wettability, which appears counter-intuitive

since it is also generally accepted that wettability and cell

adhesion are inversely related. However, this treatment also

forms charged oxygen-based functional groups at the surface

[320,330,331]. The charges endowed by these chemical groups

may dominate wettability, with respect to protein and cellular

interactions, and may be the reason for the difficulty in

correlating wettability and cell attachment over a wide range of

substrates [313,332–334]. However, the relationship between

Fn adsorption and cell attachment suggests a synergistic

relationship between Fn adsorption and chemical charges with

respect to cell adhesion [329].

At the surface of calcium phosphate ceramics, proteins

interfere with the ionic exchange mechanisms as observed in

vitro for calcium phosphate ceramics. These interactions

depend on the bioceramics’ characteristics (e.g. phase,

crystallinity, composition and texture) [335–337] and on the

properties of the proteins (e.g. conformation, isoelectric point),

concentration and whether they act in solution or on substrates

[193,338–341]. First, in suspension, proteins can inhibit or

support calcium phosphate nucleation and growth

[193,338,342]. Regarding bone proteins such as collagen,

osteopontin, osteonectin, bone sialoprotein or osteocalcin,

phosphorylated entities have demonstrated their ability to

nucleate and grow calcium phosphate crystals [342]. However,

not all of the phosphorylated proteins induce calcium phosphate

formation; osteopontin is a particularly strong crystallization

inhibitor [338]. Second, when proteins adsorb onto calcium

phosphate substrates, their charge, concentration and the

presence of calcium in the surrounding fluids influence the

surface coverage kinetics and pattern that can evolve with time

[343]. These adsorbed proteins can thereafter influence the new

formation of calcium phosphate crystals by blocking the

substrate’s nucleation sites [339–341], irrespective to the

protein’s isoelectric point [340]. In vivo, hundreds of proteins

are present in biological fluids and their global effect on

calcium phosphate reactivity is insufficiently understood. An in

vitro study conducted on a large panel of serum proteins

adsorbing on different inorganic materials, including calcium

phosphate ceramics did not show significant differences despite

the physical and chemical characteristics of these materials

[344]. It is clear however, that they play a significant role in

ionic exchanges and their subsequent effect on their biological

activity, since proteins are detected in close association with the

nanocrystalline carbonated apatite formed on the surface of

calcium phosphate bioceramics in vitro and in vivo

[190,192,193,196,204,345]. Consequently, the nature, quantity

and conformation of these proteins at the biomaterial surface

will determine cellular activity [335,337]. The specific pattern

of protein adsorption has been hypothesized as an influencing

factor in the osteoinductive properties of biomaterials.

5.3.2. Smart designs to control protein adsorption

The ability of materials to modulate downstream gene

response without exogenous growth factors, coatings or

complex ligand incorporation has the potential to greatly

facilitate the development of tissue engineering and cellular

therapies. As an illustration of this concept, a class of biphasic

calcium phosphate ceramic induced de novo bone formation at

non-osseous sites in vivo without requiring the delivery of cells

or biologic compounds [9], suggesting that the surface

chemistry of the ceramic allowed the selective adsorption of

morphogenetic proteins that trigger osteogenesis. It was also

demonstrated using polymer libraries that substrate chemistry

can influence the developmental lineages of embryonic stem


cells [7]. We have reported data that demonstrated that

chondrocytes cultivated on PEGT/PBT substrates can be

induced to express highly specific cell behavior, including

modulation of phenotype. This was evident at the cell-surface

receptor, focal adhesion complex, cytoskeletal, intracellular

signaling and gene expression levels, illustrating the versatility

of smart materials in designing the material–biologic interface

[346]. It was found the profiles of adsorbed proteins at the

material surface played a key role in the downstream cellular

response. In principle, this ability to ‘dial in’ the required

biological response from the substrate allows us to establish cell

function and fate by eliciting the interfacial conditions that

regulate signaling to and from the transplanted cells.

Molecular imprinting is also a sophisticated technique to

tether specific biological entities at biomaterial surfaces. This

strategy consists of constructing ligand selective recognition

sites onto synthetic materials where a template (atom, ion,

molecule, complex, molecular, ionic or macromolecular

assembly, including micro-organism) is employed in order to

recognize site formation during the covalent assembly of the

bulk phase by a polymerization or polycondensation process,

with subsequent removal of some or all template being

necessary for recognition to occur in the spaces vacated by the

templating species [347]. This technique is applied in

chromatography, membrane separations, solid phase extrac-

tion, immunoassays, synthesis, catalysis and sensors. Recently,

Alexander et al. reviewed the current therapeutic molecular

imprinting applications, imprinted materials have attracted

attention as vehicles for controlled release of drug, and as

screening tool in drug discovery. The use of molecular

imprinted sensors has enabled the thorough screening of

combinatorial libraries, and the synthesis of drug candidates in

the recognition sites of imprinted materials [347]. With regard

to tissue regeneration, imprinting sites onto a biomaterial in

order to attract molecules favorable to tissue formation could be

a smart way to instruct the body to heal by itself, avoiding

therefore exogenous entities.

5.4. Controlling cell–biomaterial interactions

In general cells/biomaterial interactions depend on the

surface characteristics such as topography, chemistry and

surface physics. As mentioned earlier, surface characteristics

determine ionic exchange dynamics and protein adsorption.

They also affect cellular activity, namely, cell attachment,

proliferation and differentiation.

5.4.1. Cellular activities and functions on biomaterials

In contrast with differentiated osteoblasts and chondrocytes,

mesenchymal stem cells are migratory, highly proliferative

cells and have greater differentiation potential. They can

migrate on a substrate by generating cycles of weak adhesion,

traction, movement and detachment. At the end of the

migration phase, mesenchymal/osteoprogenitor cells adhere

onto the substrate by developing strong focal adhesion with

substrates in order to start their differentiation phase, similar to

osteoblasts or chondrocytes. Cell migration and adhesion are

mediated via integrins which are transmembrane proteins

[348]. Fibronectin and vitronectin are serum proteins that are

ligands for integrins, and have been shown to mediate adhesion

to biomaterials in vitro [337,348]. In contact with the

biomaterial, the attached cells may further differentiate and

produce specific extracellular matrix components (see Fig. 5 for

a general overview of the mechanistic events that occur during

this process).

It has been shown that chondrocyte attachment and

spreading on two-dimensional surfaces leads to the loss of

expression of aggrecan and type II collagen, a phenomenon

known as ‘dedifferentiation’ [349–353]. Though not strictly a

movement along the cell’s differentiation lineage pathway to

yield a more primitive progenitor cell, it refers to

chondrocytes exhibiting fibroblast-like characteristics, such

as cell morphology and non-phenotypic protein expression

and synthesis. The ability of chondrocytes to rapidly alter

their differentiated phenotype and gene expression following

attachment makes them an interesting cell type to study

phenotype modulation by controlled variations in substrate

properties.

5.4.1.1. Cell surface receptors are the link between a cell’s

externally-induced signals and intracellular signaling

pathways. Ligand-receptor signaling is critical in determining

the fate of differentiated chondrocytes and osteoblasts, as well

as progenitor cells of these lineages. This is governed by

substrate/protein interactions with cell surface receptors known

as integrins (integral membrane receptors). Integrins are

heterodimers of transmembrane a and b subunit proteins that

are involved in cell–cell binding as well as cell–substrate

adhesion. Different combinations of a and b subunits are

possible, with the b1 subunit being most promiscuous in its

ability to form dimeric receptor complexes. More than 15 a

subunits and 8 b subunits have been identified with most

integrins also able to bind to multiple ligands at the integrin

headpiece [354].

Of the different integrins tested, chondrocytes in situ have

been shown to most abundantly express a5b1 and the cell

surface expression of this integrin is markedly increased during

in vitro proliferation in monolayer [355]. Other integrins found

at cell surfaces of adult articular chondrocytes include a1b1,

a3b1, avb5 and avb3 [356]. However, fetal chondrocytes

were shown to express a2b1 and a6b1, indicative of the

different roles of integrins during development and homeostasis

[357].

Certain ECM and adhesive proteins such as fibronectin and

vitronectin bind to integrins primarily by Arg-Gly-Asp (RGD)

peptide sequences contained within the proteins’ cell binding

domains. The RGD sequence is able to adopt variable

conformations in different proteins, and this conformational

flexibility has been suggested as a factor in integrin-ligand

specificity [358]. Studies examining the effects of RGD

peptides on chondrocyte behavior has revealed the suppression

of chondrocyte differentiation in the epiphyseal plate and the

inhibition of chondrogenesis in mouse limb bud cells

[359,360]. This demonstrates the importance of integrin-based


signaling in cartilage development, and is being leveraged by

stem cell and regenerative medicine scientists in decoding the

signals required to stimulate the differentiation of stem

cells towards the chondrogenic or osteogenic lineages

[149,361,362]. Recent reviews have elegantly elaborated on

signal transduction pathways and signal transmission to the

cytoskeleton by ligand binding to integrins [363,364].

In addition, differentiated chondrocytes also express CD44,

a transmembrane proteoglycan receptor that is the principal

receptor involved in binding to hyaluronan [365,366]. It

enables the retention of the hydrated aggrecan aggregates at

chondrocyte surfaces, which can be released from the surface

by the addition of small HA oligosaccharides or CD44

Fig. 5. Cellular events at a biomaterial surface. (a) Colonization and attachment;

antisense nucleotides [141,366]. The extracellular domain of

this receptor contains strong sequence similarities with HA

binding regions in link protein and chondroitin sulfate side

chains [358]. Chondrocyte delivery substrates for tissue

engineering applications that enable cell–substrate binding

via CD44 are able to maintain the differentiated cell

phenotype. The intracellular domain of CD44 associates with

cytoskeletal proteins and can regulate signaling in this manner

by the coordinated activities of actin network modification and

protein phosphorylation, primarily through interaction via

tyrosine kinases, ankyrin and ezrin/radixin/moesin (ERM),

leading to differential states of phenotypic expression

[364,367].

(b) proliferation; (c) migration; and (d) differentiation and tissue formation.



5.4.2. Effect of surface physico-chemistry

Chondrocyte function is known to depend on the adsorption

of specific ECM proteins on the culture substrate: it has been

shown that vitronectin is required for chondrocyte adhesion to

methacrylate-based polymer substrates [327] and fibronectin

enhances chondrocyte dedifferentiation in monolayer culture

[368,369]. Although wettability has been shown to influence

biomaterial–cell interactions and is often used as a parameter

by which to characterize and compare biomaterial surfaces

[313,370], conclusive correlations linking wettability, protein

adsorption and cell function have yet to be made [332,333,371–

374]. However, these assays have usually been performed using

a limited range of substrates and do not address material–

protein interaction over a wider range of material properties.

Understanding such interactions is of considerable importance

in cell transplantation therapies since the resulting cell–

material constructs are implanted into the defect site with the

desired goal to mimic the important properties of cartilage,

thereby enabling improved functionality [122]. To elucidate the

effects of selecting certain materials for such cell transplanta-

tion therapies, it is crucial to understand the cell–material

interactions that occur during this process, from cell attachment

to downstream signal transduction events.

In the bone environment, some biomaterials have an

intrinsic ability to bind with newly formed bone with a

physico-chemical continuity. Calcium phosphates and bioac-

tive silica glasses, are the first bioceramics that have been

specifically developed for bone repair. The idea behind the

development of these bioceramics was that making synthetic

materials with composition similar to bone mineral would

improve their biocompatibility and acceptance by host bone.

These bioceramics indeed exhibit excellent bone-bonding

properties that are related to the surface reactivity, via

dissolution–precipitation mechanisms, creating an interfacial

mineralized layer between the implant and bone tissue that

insures their cohesion. Structurally, this layer is comparable to

the films grown in vitro by dissolution–precipitation mechan-

isms, i.e. nanocrystals of carbonated apatite [19,190–194], in

simulated body fluids that mimic the mineral composition of

blood plasma. When formed in the presence of osteogenic cells

experiments, this mineralized layer is comparable to the cement

lines synthesized in vivo [31,195]. In vivo (osseous and non-

osseous environment), physicochemical and crystallographic

continuity are observed between the calcium phosphate implant

and the newly formed mineralized layer [19,196,197]. Its

occurrence and thickness are related to the reactivity

(dissolution–precipitation) of the calcium phosphate substrate

[197], also referred to as ‘bioactivity’ [198]. This mineralized

interface ensures a physicochemical and mechanical cohesion

between the implant and the host bone (Fig. 4a and b). It is

particularly relevant for load-bearing applications, such as hip

metallic prostheses coated with calcium phosphate which layer

improve undoubtedly the mechanical stability of the implant by

augmenting and accelerating the bone apposition [199–201].

The bone-bonding ability of these bioactive biomaterials is also

exploited in bone tissue engineering [202].

Fig. 4 shows bone formation at the surface of metal and

apatitic implant biomaterials.

5.4.3. Effect of topography

Surface roughness is generally known to influence cell

attachment in vitro, including attachment of bone cells. Rough

apatitic surfaces appear to enhance osteoclastic attachment

compared with smooth ones [375]. Grooved surfaces influence

osteoblast guidance, as does the groove profile and topography,

independent of the chemical nature of the substrate [376]. On

the other hand, on micro- and macroporous calcium phosphate

ceramics, osteoblasts sense the surface microporosity and can

bridge even large pores many times larger than fully spread

osteoblasts [377]. With regard to osteoblasts differentiation,

Chou et al. demonstrated in vitro that osteoblastic cells were

sensitive to crystal shape: large apatite crystals induced more


bone sialoprotein and osteocalcin expression after 3 weeks of

culture [94]. Finally, Redey et al. have shown that surface

energy strongly affected initial osteoblast and osteoclast

activity. In vitro, early osteoblast proliferation and function

and osteoclast adhesion are affected by the polar component of

the calcium phosphate surface. However, at a later culturing

stage, surface energy does influence anymore both osteoblast

and osteoclast activity [378,379]. Thus, adjusting the under-

lying micro- and nanotopography is also a smart way to trigger

and modulate specific cellular functions.

5.4.4. Micro- and nanodesigns of biomaterials

Cell dimensions generally range from 5 to 100 mm.

Regardless their type, cells can react to micrometer and

nanometer scales. The realization that the physical features of

the substratum on which cells are growing can affect their

morphology and migration dates to the 1930s, but it was not

until the development of micro- and nano-fabrication methods

that a wide range of physical features could be made on the

micro- and nanometer scale. Surface topography can affect

various cellular reactions, namely cell orientation, adhesion,

movement, gene expression, activation of phagocytosis and

orientation of the cytoskeleton in vitro [380]. In vitro micro-

features with specific shapes can influence the cellular

activities, including osteogenic differentiation [381]. In vivo,

the surface microtopography can significantly affect tissue

neoformation. For example, the initial surface roughness of the

titanium prostheses greatly influences early bone formation and

contact with the implant [382,383]. More recently, the micro-

texturing of macroporous biomaterials induced bone formation

in non-osseous environment [210], and micro-grooved hollow

fibers induced tendon regeneration [384].

Webster et al. have demonstrated that the presence of

nanophases (smaller than 100 nm) in titanium oxide, alumina

and hydroxyapatite substrate could have consequences on bone

cells activity and functions [385–387]. Independent of ceramic

composition, osteoblasts and osteoclasts favored the introduc-

tion of nanophases into the ceramic [385,386]. Examination of

the underlying mechanism(s) of cell adhesion on nanophase

ceramics revealed that these ceramics adsorbed significantly

greater quantities of vitronectin (an adhesion cellular protein),

which, subsequently, may have contributed to the observed

select enhanced adhesion of osteoblasts. Select enhanced

osteoblast adhesion was independent of surface chemistry and

material phase but was dependent on the surface topography

(specifically on grain and pore size) of nanophase ceramics

[388].

Several techniques, namely electron beam lithography,

colloidal resists, self-assembling systems, casting, micro-

contact printing, masters made by one of the above techniques

and particle synthesis, are nowadays available to create

nanotopography on organic and inorganic surfaces [380]. It

has been demonstrated that cells can react in vitro to objects as

small as 5 nm, which are 1000–5000-fold smaller in size [389].

Curtis et al. have shown a relation between symmetry and

regularities of nano-objects and cellular adhesion. With cliffs,

adhesion is enhanced at the cliff concave edge, while pits or

pillars in ordered arrays diminish adhesion. The results

implicate ordered topography and possibly symmetry effects

in the adhesion of cells. More precisely, the asymmetry and the

presence of concavities may increase the wettability of the

substrate, and therefore enhance cell adhesion [390]. Other

reports on the cellular interactions with specific nano-patterned

substrates of various composition have shown that nano-shaped

holes can also (i) control cell life and death [391] and (ii) orient

cell commitment towards osteogenic lineage [149,392]. The

influence of surface nanotopography on cell behavior is

mediated via changes in the orientation and conformation of

proteins that interact with the nanotextured substrate

[386,393,394].

Micro- and nano-patterning can include modification of

surface chemistry or topography. Nano-patterning can consist

as well at grafting chemical functions onto biomaterials. In the

field of tissue engineering, the grafting of RGD peptides has

been the focus of much attention [395,396]. Their presence at

biomaterials surfaces improved cell adhesion. Indeed this

peptide sequence is present in various extracellular matrix and

plasma proteins, and it constitutes a major recognition site of a

large number of adhesive extracellular matrix, blood and cell

surface proteins [358]. Other peptides and organic molecules

have been used to stimulate cell adhesion [236,397,399]. Zhang

et al. could design adhesive arrays and patterns for cells by

combining two types of oligopeptides on a flat substrate [398].

Other patterned substrates containing two different peptides

could specifically and locally induce the adhesion of two

different cell types relevant in the field of bone tissue

engineering (osteoblasts and fibroblasts) [397]. Small changes

in the graft can induce significant differences in the efficiency

of the cellular adhesion [236,399]. For example, alkylsilanes

were also grafted on model surfaces. The nature of their

termination (epoxide, carboxyl, amine, or methyl) strongly

affected cellular adhesion [236]. The grafting of specific

function does not concern solely cellular interactions, but aims

also at the immobilization of bioactive molecules. Puleo et al.

could immobilize an osteoinductive protein (BMP-4) after

plasma polymerization of allyl amine. Cells cultured on this

BMP-4 immobilized biomaterials differentiate towards osteo-

genic lineage. This indicated that the bioactivity of the BMP-4

was retained trough the immobilization process [400].

However, similar titanium surface treatment favored also the

immobilization of another protein. No evidence was shown that

this immobilization technique was specific to a particular

bioactive molecule.

Although micro-patterning has proven a biological effect in

vitro and in vivo, long term in vivo consequences of nano-

patterning remain to be demonstrated.

5.5. Multi-functional scaffolds

Tissues are multi-component and multi-functional, while

current biomaterials are designed to replace only one

predominant function. The natural tissue repair process

involves multiple signaling molecules, in a time and

concentration-dependent fashion, as is clearly established for


bone repair [401–403]. In bone tissue engineering, the main

focus has been on bone forming biomaterials and bone forming

cells. However, the poor viability of newly formed tissues has

underlined the need for their early vascularization [404]. The

design of biomaterials and cell systems is focusing increasingly

on the simultaneous development of bone tissue and vascular

network. Trauma and joint diseases frequently involve both the

articular cartilage surface and underlying subchondral bone.

Making a mechanically stable construct, that can bear the

articular loads and would support both bone and cartilage

formation, is an important goal in tissue engineering.

5.5.1. Multiple-component scaffolds

A more recent strategy consists of combining biomaterials

with intrinsic properties that stimulate the formation of specific

tissues into a multi-phasic construct. For example, a poly-L-

lactic acid/hydroxyapatite composite was produced after

image-based design and solid-free form fabrication with the

aim to generate a biomaterial matching the load-bearing and

articular geometry [405,406]. In vivo, the chondral zone seeded

with chondrocytes and the osseous zone loaded with BMP-7

transfected fibroblasts showed the formation of cartilage and

bone, respectively [405]. A more clinically relevant approach

consists of using one cell source, such as mesenchymal stem

cells, that differentiate into cartilage and bone in well-defined

sites. Towards this end, triphasic scaffolds have also been

fabricated for osteochondral defects using an osteoinductive

Fig. 6. Examples of multi-functional scaffolds designed based on 3D-fibres deposit

with osteoinductive BCP inserts (a), the cartilage and bone compartments are tight

polymeric scaffold: (c) the fibres are composed of two different polymers (outer an

calcium phosphate ceramic in combination with a copolymer

composed of poly(ethylene glycol) (PEGT) and terephthalate-

co-poly(butylene terephthalate) (PBT), with different mole-

cular weight of the starting PEG blocks, and different weight

ratios of the PEOT and PBT blocks. The copolymer was

selected among others as it has shown better results in terms of

mechanical properties for cartilage repair applications, and cell

attachment for bone tissue engineering [407,408]. The porous

triphasic structure has been obtained by three-dimensional fiber

deposition technique, with mechanical integrity achieved by the

deposition of layers of circular and concentric fibres of the two

polymer compositions (Fig. 6a). Another example of a multi-

functional scaffold consists of embedding hollow fibres within

a bulk scaffold in order to favor the formation of neo-

vasculature within the scaffold that would integrate with the

native host vascular network. In a representative study, Moroni

et al. demonstrated the encapsulation of fibers within a scaffold,

exploiting the differential melting temperatures and viscosities

of two polymers, as shown in Fig. 6b [275].

5.5.2. Multidrug delivery

This technique consists of the smart combination of signals

able to stimiulate in parallel several cell types and tissues. This

approach can be obtained by multi-drug delivery systems or by

employing multiple scaffold components.

Multi-drug delivery attempts have been made using

biomaterials in the shape of rods [409], hydrogels [305], or

ion. (a, b) triphasic osteochondral construct combining two polymeric scaffolds

ly glued together with a mixture of the two polymers (b); (c, d) hollow-fibres

d inner shell), and (d) after dissolution of the inner shell (courtesy L. Moroni).


gelatin layers [410]. Porous scaffolds as reservoirs for multiple

proteins were obtained by assembly and fusion with micro-

spheres [278,279] or by associating them with pre-existing

porous structures [300]. Regariding cartilage repair, the

opportunity to combine the release of different growth factors

was investigated by mixing IGF-1 and TGF-b1-containing

PLGA microspheres in a hydrogel [305]. This approach

appeared promising as the two growth factors had synergistic

effects on the enhancement of chondrocyte proliferation and

maintenance of their phenotype. In a broader view, it is likely

that the release from scaffolds of different growth factors with

different release profiles would be therapeutically beneficial.

Other methods have also been considered for this purpose. For

example, mixing two populations of gelatin microparticles

releasing IGF-1 and TGF-b1 within an hydrogel or adsorbing

TGF-b1 to the hydrogel directly, allowed independent control

of the release profiles of the two proteins [411]. Another

approach consists of applying multiple gelatin coatings

containing BMP-2 and IGF-1 on flat surfaces to control the

release of each growth factor independently by diffusion

through the superposed layers [410,412]. Similarly, the

successive coating of PEGT/PBT copolymers containing

different model proteins on prefabricated compression-molded

scaffolds allowed a tailored and independent release [274].

However these techniques are yet preliminary and still have to

be tested in relevant articular cartilage defect models in vivo.

The incorporation of different signaling molecules has also

been studied in order to induce simultaneous bone formation

and neo-vascularization. The bi-functionalization of (poly-

lactic-co-glycolic-acid) porous scaffolds was achieved by two

different protocols and investigated in vivo. In one experi-

mental group, an angiogenic factor (vascular endothelial

growth factor, VEGF) and an osteogenic factor (plasmid-

DNA encoding for BMP-4) were aggregated at the scaffold

surface [413]. These bi-functional scaffolds associated with

mesenchymal stem cells in vivo, and showed significant

increase in bone regeneration and blood vessel formation. In

another group, a calcium phosphate layer was combined with

VEGF containing porous scaffold before implantation. The

presence of both VEGF and calcium phosphate layer in a single

scaffold induced more blood vessel formation than either

component alone. However, bone tissue formation was not

significantly higher than the mono-functionalized scaffolds

[282]. In both studies, it was not clear whether a connection was

established between the engineered vascular system and the

host vasculature. The perfusion of the engineered vascular

network is essential for the viability of the tissue engineered

construct, as it was shown in a model system for muscle tissue

engineering [414]. Truly smart systems will connect the

vasculature between the implant and host tissue.

5.6. Three-dimensional control of biomaterials on cells

and tissues

In tissue regeneration, the three-dimensional features of the

biomaterial are of high relevance as it affects the activity of

various cell types, the subsequent tissue neoformation and

viability. Spatial effects of biomaterials are evident in vivo from

the centimeter to nanometer scales. For example, the angle of

curvature of a hip stem can significantly affect the prosthesis

integration. Biomaterials, used to fill tissue defects, act as an

exchange platform between the hosting tissues and the implant.

Their three-dimensional organization and features will play a

critical role in these exchanges, namely, (i) the selective

penetration of cells, nutrients and oxygen, and tissues, (ii)

clearance of metabolic products, (iii) proliferation and

differentiation of cells, (iv) neo-formation and viability of

desired tissues, and (v) degradation of the biomaterial. The

biomaterials options supporting these exchanges are either

porous solids or hydrogels.

With respect to bone biomaterials, porous solid structures

are preferred because osteogenic differentiation is positively

influenced by strong adhesion onto surfaces [174]. In osseous

environments, scaffolds having higher porosity (percentage of

void space in a solid) and pore size stimulate greater bone

ingrowth. However, this trend results in diminished mechanical

properties, thereby defining a practical upper limit for pore size

and porosity. A recent review established that the minimum

requirement for pore size is considered to be approximately

100 mm due to cell size, migration requirements and nutrient

transport. However, pore sizes >300 mm are recommended, to

achieve enhanced new bone formation and the formation of

blood capillaries. The vascularization of the porous biomaterial

is desired as early as possible, as it allows a constant and local

delivery of oxygen and nutrients during tissue morphogenesis.

This efficient transport of oxygen and nutrient are necessary for

both transplanted cells and newly formed bone survival inside

the scaffold. Interestingly, small pores favor hypoxic conditions

(low oxygen concentration) and induced cartilage formation

before bone [415]. Therefore, regardless of biomaterial

composition, it is possible to direct either bone or cartilage

formation via the control of the pore size. In principle, by

preparing a gradient in pore size, one could simultaneously

stimulate bone and cartilage formation at desired locations in a

single scaffold.

In addition to the porosity and pore size, the shape of the

pores can also play a pivotal role. We reported earlier, that, for a

given biomaterial, only macropores with specific geometry can

induce bone formation in non-osseous sites [197,226]. Pore

confinement has been also pointed out as a critical factor [88].

In addition, the size of the macroporous implant influences as

well its osteoinductive abilities. In goat muscles, Habibovic

et al. observed that the larger the implant, the more bone

formation is induced [219]. The three-dimensional character-

istics of the biomaterials clearly have marked effects on cell

behavior, and therefore on tissue formation. However, the

complete picture of processes involved in biomaterial regula-

tion of biological mechanisms remains unclear.

To date, unorganized porosity seems to play a positive role in

tissues. It is a challenge to ascertain exactly which parameters

play dominating roles in instructing the biological response,

improved understanding of which would help in designing

better biomaterials. Designing, controlling and characterizing

porous structures from the macro- down to the nanometer scale


still demand tremendous technical improvements in multiple

areas.

Polymer hydrogels have also been designed for responsive-

ness to their local biological environment. These can be based

on both synthetic or natural polymers, and have been defined as

hydrophilic polymers that swell by taking up water in the range

of 10% to �103-fold their dry weight [416,417]. In general,

these materials can be classified as either chemical or physical

hydrogels. Chemical hydrogels are crosslinked, with nodes of

high crosslinking density in a mainly low crosslinked structure.

Conversely, physical hydrogels are supported without covalent

interaction, and contain both hydrophilic and hydrophobic

domains. Water uptake occurs sequentially at the hydrophilic,

hydrophobic and interstitial regions to cause swelling.

Hydrogels have been made to be responsive to a number of

stimuli, including: pH, Ca2+, Mg2+, organic solvents, tempera-

ture (including sol–gel transition), external magnetic fields,

electric potential, UV, IR and ultrasound [112,419–431].

Among the frequently studied hydrogel systems is poly(N-

isopropylacrylamide) [432–436]. This family of thermosensi-

tive hydrogels has the ability to collapse above a specific, pre-

defined temperature (lower critical solution temperature—

LCST). Below the LCST, the hydrogel is dehydrated and

hydrophobic, whereas above the LCST, it is hydrated and

hydrophilic. An additional degree of responsiveness can be

attained with the inclusion of methacrylic acid, which makes

the structures responsive to both temperature and pH. Other

combinations include responsiveness to light and temperature,

or Ca2+ and temperature. The potential of such materials is

derived from their ability to transplant cells or locally

administer growth factors based on specific biochemical cues,

ex vivo or in vivo.

5.7. Physical stimuli on cells

It has been known for more than a hundred years that bone and

cartilage are sensitive to mechanical loading: cartilage is able to

remodel only when mechanical load is applied and bone density

patterns are governed by the distribution of stress. Nowadays, we

now that physical stimuli are converted into biochemical stimuli

that affect cellular functions and activities. This conversion can

be divided in four steps: (1) mechanocoupling is the conversion

of the applied physical force to secondary forces or physical

phenomena detected by the cells; (2) mechanotransduction is the

conversion of either the primary or secondary physical stimulus

into an electrical, chemical or biochemical response; (3) the

signal transduction entails the conversion of one biochemical

signal to another; and (4) final step completes the conversion

from initial stimulus to final tissue-level response. The four steps

outlined above suggest methods by which the mechanical or

biochemical environment may be modified to control the

development of engineered tissue. For instance, one might expect

tissue-engineered bone to increase its density when subjected to

mechanical conditions resembling those known to stimulate net

bone formation in vivo [252].

In vivo, mechanical forces arise from diverse sources such as

muscular contraction resulting in stress and strain of muscle

and tendon, and locomotion generating small amplitude cyclic

compression of bone and cartilage. Blood flow exerts shear

stress on the endothelium, and pressure and cyclic strain is

experienced by the endothelium and vessel wall. Interstitial

fluid flow applies shear stress to bone cells and growth which

stresses skin [252]. These naturally occurring mechanical

stimuli have been mimicked in vitro in order to investigate their

effect on cells [437–439]. It has been reported that the

production of shear stress on rat calvarial osteoblasts cultured

on a collagen scaffold increased the production of osteogenic

markers. In other words, mechanical stimulation stimulated

osteogenic differentiation [438]. Conversely, a pulsatile flow

induced osteogenic differentiation of similar cells [439]. Under

cyclic pressures, cellular functions can be positively or

negatively affected depending on the frequency and duration

of the cycles. Differential effects were reported for osteoblasts,

endothelial cells and fibroblasts [437].

Other types of ‘‘artificial’’ physical stimuli on cellular

behavior have been considered as tools to contribute to tissue

repair and healing [307,440–454], namely electrical

[441,447,451,452] or magnetic field [307,445,454], laser

irradiation [440,449,450], heat shock [449] and ultrasound

[449]. These physical stimuli have been proven to influence cell

activity in vitro, and tissue formation in vivo [257,450,452].

Some of these physical stimuli are applied in clinics to treat

tissue defects [441,449]. Pulsating electrical fields, consisting

at applying a voltage inducing a magnetic field, have been

widely documented in vitro and in vivo for hard and soft tissue

repair [441,448,453,455–457]. The cellular responses, namely

proliferation, differentiation, orientation, depend on the

pulsatile frequencies, total energy of the applied field, and

the maturation stage of the cells [447,449,457,458]. Although

the stimulation mechanisms of these biophysical factors are not

well understood, they influence the cellular production of

proteins [447,453,456,457], growth factors [459], minerals

[460] and free radicals [461]. These pulsating electrical fields

have also been proposed as a therapy against osteoporosis by

influencing osteoclast activity [455,456].

These physical stimuli have been also investigated as factors

contributing to tissue healing after biomaterial implantation,

cell proliferation in vitro [440,451,462,445,463], and in vivo

[257,464]. In vivo, these physical stimuli can, in particular,

affect the orientation of the newly formed and growing tissues

[445], as well as their mechanical resistance to load [464].

Developing biomaterials able to generate physical stimuli

can be of great interests for cellular and tissue stimulations.

Piezoelectric materials are able to generate transient surface

charges under minute mechanical strain. Biocompatible

piezoelectric ceramics and polymers have been developed

for various tissue repair purposes [465–468]. The piezoelectric

barium titanate was investigated as a bone tissue growth

enhancer in the 1980s and 1990s [466,467]. The piezoelectric

features of barium titanate can be detected up to 86 days after

implantation in femurs [467]. The same investigators did not

find any significant difference the polarized and the electrically

neutral barium titanate-tissue interfaces [466]. However, in

another study, Feng et al. reported that the implantation of


polarized hydroxyapatite-barium titanate ceramics in canine

jaws promoted the growth and repair of the bone significantly

compared with hydroxyapatite. The growth of newly formed

tissue around the piezoelectric ceramic was found to be

direction-dependent [465]. With regard to nerve regeneration,

piezoelectric polymers, such as polyvinylidene fluoride

(PVDF) have also shown a beneficial effect on nerve cells in

vitro [468]. Within 98 h, polarized PVDF substrates exhibited

significantly greater levels of process outgrowth and neurite

lengths at all time periods compared with non-polarized

substrates. The polarization of PVDF did not induce its physic-

chemistry. Therefore, the stimulating effect on cells could

solely be attributed to the piezoelectric stimulus [468]. The

main drawback of piezoelectric material remains the mechan-

ical deformation required to generate charges.

On the contrary, electrically conductive materials have the

advantage of a non-invasive (external) control over the level

and duration of stimulation [469]. Of the electric conducting

polymers, oxidized polypyrolle (PP) has been the most

thoroughly investigated for use in biological systems

[470,471]. In vivo, nerve cells cultured on PP films and

subjected to an electrical stimulus through the film showed a

significant increase in neurite lengths compared with ones that

were not subjected to electrical stimulation through the film and

tissue culture polystyrene controls [470]. Modified PP has been

more recently developed with the purpose to improve the

degradation properties of the conducting polymers [472].

Another approach to produce conducting biomaterials consists

of blending polylactic acid and carbon nanotubes. When

osteoblasts cultured on the surfaces of these nanocomposites

were exposed to electric stimulation for various periods of time,

there was an increase in cell proliferation and differentiation.

These results provide evidence that electrical stimulation

Fig. 7. The convergence of a multitude of technologies and scientific disciplines h

regeneration.

delivered through current-conducting biomaterials promotes

osteoblast functions that are responsible for the chemical

composition of the organic and inorganic phases of bone [473].

The positive effect of physical stimuli in vitro is clear and

promising, and has been shown in vivo.

6. Concluding remarks

Scientists working in the tissue regeneration field are faced

with complex biological systems. From a biological point of

view, significant improvements have been realized with regards

to investigative research tools, and in our understanding of the

mechanisms involved in tissue and organ regeneration. In

parallel, the materials design and fabrication capabilities and

processing technologies have become increasingly sophisti-

cated. These recent developments are mainly due to the fusion

of several fields (Fig. 7). A challenge that has emerged as a

result of these synergistic combinations is the exponential

increase of possibilities that have to be tested and validated.

Smart paths have to be taken to evaluate all of these smart

materials. High-throughput screening technologies can now be

combined with biological and polymer systems for rapid

evaluation of cellular effects caused by underlying polymer

substrates. For example, Anderson et al. have demonstrated the

ability to screen 3500 polymer compositions for their biological

effects on chondrocytes and neural cells, as well as 1700

compositions for their biological control effects on human

embryonic stem cells [7,249]. As technological advancement

continues in high throughput screening, these tools will enable

rapid assessment of interactions between biomaterials and cells

at the cellular, protein and gene levels, offering new windows

into the regulatory signals transmitted by polymers to the

biological environment. In vivo, ‘‘smart’’ screening models and

as enabled the advancement of novel smart and instructive strategies for tissue


biomarkers have been developed to better select biomaterials

from a plethora of options [474]. The acquisition of all of the

subsequent biological data will require smart software and data

processing technologies. Closer to mimicking the complexity

of natural tissues, ‘‘smart’’ solutions to regenerate damaged

tissues have emerged, and will further evolve due to the

continued convergence of various fields of expertise.

Acknowledgements

Much of the work presented here is based on the valuable

contribution of the many researchers at the University of

Twente, Department of Tissue Regeneration, including Pamela

Habibovic, Clayton Wilson, Tim Woodfield, Jos Malda,

Shihong Li, Du Chang, Joost de Wijn, Jerome Sohier, Lorenzo

Moroni and Marcel Karperien.

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