Thesis

PREPARATION AND CHARACTERISATION OF

TRI-CALCIUM PHOSPHATE SCAFFOLDS WITH

TUNNEL-LIKE MACRO-PORES FOR BONE TISSUE

ENGINEERING

Wei Zheng

BE(Medical Engineering)

Principal Supervisor: Associate Professor Cheng Yan

Associate Supervisor: Associate Professor Clayton Adam

A Thesis submitted for Master of Engineering (Research)

Faculty of Built Environment and Engineering

Queensland University of Technology

2011

Key Words

Bioglass, bone formation, Compressive strength, Hydroxyapatite (HA),

Macro-tube Scaffold, MTT test, Phase transformation, Scanning electron

microscopy (SEM), Simulated body fluids (SBF), Sintering, Tissue engineering,

Tri-calcium phosphate (TCP).

II

Abstract

Calcium Phosphate ceramics have been widely used in tissue engineering due to

their excellent biocompatibility and biodegradability. In the physiological

environment, they are able to gradually degrade, absorbed and promote bone

growth. Ultimately, they are capable of replacing damaged bone with new tissue.

However, their low mechanical properties limit calcium phosphate ceramics in

load-bearing applications. To obtain sufficient mechanical properties as well as

high biocompatibility is one of the main focuses in biomaterials research.

Therefore, the current project focuses on the preparation and characterization of

porous tri-calcium phosphate (TCP) ceramic scaffolds. Hydroxapatite (HA) was

used as the raw material, and normal calcium phosphate bioglass was added to

adjust the ratio between calcium and phosphate. It was found that when 20%

bioglass was added to HA and sintered at 1400oC for 3 hours, the TCP scaffold

was obtained and this was confirmed by X-ray diffraction (XRD) analysis. Test

results have shown that by applying this method, TCP scaffolds have significantly

higher compressive strength (9.98MPa) than those made via TCP powder

(<3MPa). Moreover, in order to further increase the compressive strength of TCP

scaffolds, the samples were then coated with bioglass. For normal bioglass coated

TCP scaffold, compressive strength was 16.69±0.5MPa; the compressive strength

for single layer mesoporous bioglass coated scaffolds was 15.03±0.63MPa. In

addition, this project has also concentrated on sizes and shapes effects; it was

found that the cylinder scaffolds have more mechanical property than the club

ones. In addition, this project performed cell culture within scaffold to assess

biocompatibility. The cells were well distributed in the scaffold, and the

III

cytotoxicity test was performed by 3-(4,5)-dimethylthiahiazo(-z-y1)-3,5-di-

phenytetrazoliumromide (MTT) assay. The Alkaline Phosphatase (Alp) activity of

human bone marrow mesenchymal stem cell system (hBMSCs) seeded on

scaffold expressed higher in vitro than that in the positive control groups in

osteogenic medium, which indicated that the scaffolds were both osteoconductive

and osteoinductive, showing potential value in bone tissue engineering.

IV

Table of Contents

KEY WORDS ............................................................................................... I

ABSTRACT ................................................................................................. II

TABLE OF CONTENTS .......................................................................... IV

LISTS OF FIGURES ............................................................................. VIII

LISTS OF TABLES .................................................................................. XI

LISTS OF ABBREVIATIONS ................................................................ XII

GLOSSARY ............................................................................................ XIII

STATEMENT OF ORIGINAL AUTHORSHIP ................................... XV

ACKNOWLEDGEMENTS ................................................................... XVI

LISTS OF PUBLICATIONS ................................................................ XVII

CHAPTER 1 INTRODUCTION ................................................................. 1

1.0 INTRODUCTION ............................................................................................... 1

1.1 BACKGROUND ................................................................................................ 1

1.2 RESEARCH GOALS ........................................................................................... 4

1.3 INNOVATIONS ................................................................................................. 4

CHAPTER 2 LITERATURE REVIEW .................................................... 6

2.0 INTRODUCTION ............................................................................................... 6

2.1 THE ROLE OF SCAFFOLDS IN TISSUE ENGINEERING ......................................... 7

2.2 SOLID FREEFORM FABRICATION TECHNIQUES (SFFT) ................................... 10

2.3 TARGET MECHANICAL PROPERTIES FOR BONE SCAFFOLDS ............................ 16

2.4 CHOICE OF BIOMATERIALS ............................................................................ 18

2.4.1 CaP bioceramics for bone replacement and repair ............................. 19

2.4.2 Surface coating ..................................................................................... 21

2.4.3 Composite mechanics ........................................................................... 23

V

2.4.4 Bioglass ................................................................................................ 23

2.5 MATERIAL AND SCAFFOLD CHARACTERISTICS THAT INFLUENCE THE

MECHANICAL PROPERTIES OF CAPS .................................................................... 24

2.5.1 Porosity ................................................................................................ 24

2.5.2 Phase transformation ........................................................................... 25

2.5.3 Microstructure ..................................................................................... 26

2.5.4 Binder ................................................................................................... 26

2.6 TCP PREPARATION TECHNIQUE .................................................................... 27

2.7 APPLICATION OF BONE MARROW MESENCHYMAL CELLS IN BONE REPAIR ..... 29

2.8 COATING METHODS ...................................................................................... 31

2.8.1 Dip coating ........................................................................................... 31

2.8.2 Sol-gel method ...................................................................................... 31

2.9 SUMMARY .................................................................................................... 32

CHAPTER 3 METHODOLOGY AND SAMPLE CHARACTERIZATION

...................................................................................................................... 33

3.0 INTRODUCTION ............................................................................................. 33

3.1 METHODOLOGY ............................................................................................ 34

3.1.1 Preparation of TCP ceramic ................................................................ 34

3.1.1.1 Preparation of hydroxyapatite powder .......................................... 34

3.1.1.2 Preparation of bioactive glass powder .......................................... 35

3.1.1.3 Preparation of HA/bioglass composite material ........................... 35

3.1.2 Fabrication of porous TCP scaffolds ................................................... 36

3.1.2.1 Fabrication of sacrificial moulds ................................................... 36

3.1.2.2 Coating ABS scaffold with wax .................................................... 38

3.1.2.3 Sintering of the ceramic-coated scaffold....................................... 38

3.1.2.4 Coating the TCP scaffolds with normal bioglass .......................... 40

3.1.2.5 Coating TCP scaffold with sol-gel mesoporous bioactive glass ... 40

3.2 SAMPLE CHARACTERIZATION ....................................................................... 40

3.2.1 XRD (X-ray diffraction) analysis ......................................................... 40

3.2.2 Total porosity of the calcium phosphate ceramic scaffold .................. 41

3.2.3 Shrinkage ............................................................................................. 41

VI

3.2.4 Imaging the scaffold using SEM ........................................................... 41

3.2.5 Mechanical testing for scaffolds ........................................................... 42

3.2.6 Apatite-formation ability of microspheres in SBF ................................ 42

3.2.7 Cell seeding and culture ....................................................................... 43

3.2.8 Observation of cell attachment on scaffolds using SEM ...................... 44

3.2.9 Cytotoxicity test by MTT assay ............................................................. 44

3.2.10 Alkaline phosphatase activity ............................................................. 45

3.2.11 Statistical analysis .............................................................................. 45

3.3 SUMMARY..................................................................................................... 46

CHAPTER 4 MECHANICAL AND BIOLOGICAL TESTING ........... 47

4.0 INTRODUCTION ............................................................................................. 47

4.1 CONFIRMATION OF TCP FORMATION ............................................................ 47

4.2 SHRINKAGE ................................................................................................... 49

4.3 MICROSTRUCTURE OF CERAMIC COMPOSITES ............................................... 51

4.3.1 10% bioglass addition .......................................................................... 51



4.4 MECHANICAL PROPERTIES ............................................................................ 58




4.5 Porosity and compressive strength of the uncoated TCP scaffold .......... 61

4.6 POROSITY OF THE COATED SCAFFOLD (NORMAL BIOGLASS) .......................... 62

4.7 POROSITY OF THE COATED SCAFFOLD (MESOPOROUS BIOGLASS) .................. 64

4.8 MECHANICAL TESTING.................................................................................. 66

4.8.1 Shape and size effects ........................................................................... 66

4.8.2 Mechanical properties for bioglass coated scaffold............................. 70

4.9 APATITE-FORMATION ABILITY OF THE MACRO-TUBE SCAFFOLDS IN SBF ..... 71

4.10 CELL CULTURE............................................................................................ 72

4.10.1 SEM images of cells growth on uncoated scaffolds ........................... 72

4.10.2 SEM images of cell growth on bioglass coated scaffolds ................... 75

VII

4.10.3 Cell proliferation ................................................................................ 76

4.10.4 Alp (Alkaline Phosphatase) activity ................................................... 77

4.11 SUMMARY .................................................................................................. 79

5.1 CONCLUSIONS ............................................................................................ 81

5.2 FUTURE WORK .............................................................................................. 84

REFERENCES ........................................................................................... 85

VIII

Lists of Figures

Figure 1 Photograph of a scaffold produced by SFF techniques (machined

into a cylinder) .............................................................................. 11

Figure 2 The top row shows the rod diameter and the out-of-plane rod

spacing. The bottom row shows the in-place spacing. Cordell et.al [57]

...................................................................................................... 13

Figure 3 Dip coating processing [218] ................................................ 31

Figure 4 Flow chart showing the six main processing steps for fabricating

TCP scaffolds ............................................................................... 34

Figure 5 ABS macrotube scaffold templates (a) and schematics of in-plane

(top view of black box in (a)) (b) and out-of-plane (view of cross

section AA) geometry (c) ............................................................. 37

Figure 6 Sintering conditions............................................................... 39

Figure 7 (a) XRD pattern for a sintered porous calcium phosphate

(mainly-TCP) sample; (b) XRD patterns of β-TCP starting powder

and sintered body [204] ................................................................ 49

Figure 8 Linear shrinkage of TCP with respect to sintering temperature 50

Figure 9 SEM images of the fracture surfaces of HA with 10% bioglass

addition at (a) 1200oC (b) 1300

oC (c) 1400

oC (d) The porosity of 10%

bioglass addition ceramic scaffold sintered at three temperatures. Error

bars represent mean ± SD for n=5. ............................................... 52

Figure 10 SEM images of the fracture surfaces of HA with 15% bioglass


oC (c) 1400


IX


bars represent mean ± SD for n=5................................................ 54

Figure 11 SEM images of the fracture surface of HA with 20% bioglass


oC (c) 1400



bars represent mean ± SD for n=5................................................ 56

Figure 12 Compressive strength of 10% bioglass addition sintered at three

sintering temperatures. Error bars represent mean ± Standard

deviation(SD) for n=5 .................................................................. 58


sintering temperatures. Error bars represent mean ± SD for n=5 59


sintering temperatures. Error bars represent mean ± SD for n=5 60

Figure 15 SEM images of the fracture surfaces of the porous TCP showing

the presence of micro-pores: (a) made via sintering HA/bioglass and (b)

made via TCP powder [207] ........................................................ 61

Figure 16 SEM micrographs of a porous TCP scaffold with a normal

bioglass at layer 42X (a), 160X (b) and 3000X (c), as well as the surface

of the normal bioglass layer (d) ................................................... 63

Figure 17 SEM micrographs of a porous TCP scaffold with a mesoporous

bioglass layer 52X (a), and 200X (b) (a and b were single coatings);

1600X (c) 1600X(d) (c and d were double coatings) ................... 65

Figure 18 Compression stress–strain curves of the sintered porous scaffold

samples (a) made via pure TCP powder, (b) made via HA/Bioglass, (c)

X

TCP made via HA/Bioglass and coated with two different bioglasses.

(Error bars represent mean ± SD for n=5) .................................... 69

Figure 19 (a) SEM image of the scaffold after soaking in SBF for 21 days (b)

EDS spectra of the surfaces of TCP ............................................. 72

Figure 20 SEM micrographs show the attachment of the cells on TCP

ceramic strut surface 800X (a) and 2000X (b) ............................. 74

Figure 21 SEM image of TCP ceramic scaffold immersed into DMEM

without cells .................................................................................. 74

Figure 22 SEM micrographs show the attachment of the cells on bioglass

coated TCP ceramic strut surface 2000X (a) and 2500X (b) ....... 75

Figure 23 MTT assay for proliferation of hBMSCs and hBMSCs combined

with TCP scaffolds at different incubation periods under the same

culture condition. Error bars represent mean ± SD for n=3 ......... 76

Figure 24 Alp activity of hBMSCs after 7 days, 14 days of culture, and the

novel TCP scaffolds significantly increased the Alp activity to higher

level compared to positive control group after 14 days (a); Alp activity

of hBMSCs after culturing for 7 days on TCP scaffold and TCP scaffold

coated with normal bioglass (b). Error bars represent mean ± SD for

n=3 ................................................................................................ 78

XI

Lists of Tables

Table 1 Studies defining optimal pore size for bone generation ........... 9

Table 2 Summary of scaffold geometry, porosity and compressive properties

for the 3D porous scaffolds fabricated by SFFT .......................... 13

Table 3 Summary of mechanical properties and porosity of human bone 17

Table 4 Summary of advantages and disadvantages of each method for TCP

preparation.................................................................................... 28

Table 5 Sintering temperature and the bioglass additions ................... 36

Table 6 Parameters of the cube and cylinder ABS scaffold templates 38

Table 7 Ion concentrations of SBF and human blood plasma ............. 43

Table 8 XRD phase analysis data for composites of HA with the addition of

10, 15, and 20wt% bioglass.......................................................... 47

Table 9 Porosity and mechanical property for composites of HA with the

addition of 10, 15, 20wt% bioglass at three temperatures ........... 57

Table 10 Average porosity and compressive strength for the scaffolds67

XII

Lists of Abbreviations ABS Acrylontrile butadiene styrene.

Alp Alkaline Phosphatase

BMPs Bone morphogenic proteins

hBMSCs Human bone marrow mesenchymal stem cell system

CaPs Calcium phosphate

CMCs Ceramic–matrix composites

CNT Carbon nanotubes

DMEM Dulbecco‘s Modified Eagle Medium

EDS Energy dispersive spectroscopy

FCS Fetal calf serum

MAPCs Multipotent adult progenitor cells

MTT (3-(4,5)-dimethylthiahiazo (-z-y1)-3,5-di- phenytetrazoliumromide)

HA Hydroxapatite

HSCs Hematopoietic stem cells

PBS Phosphate buffer saline

PLGA Poly (lactic-co-glycolic acid)

PLLA Polylactic acid

PS Polystyrene

PVC Polyvinyl chloride

pNPP pnitrophenylphosphate

SFFT Solid Freeform fabrication technique

SEM Scanning electron microscopy

SBF Simulated body fluids

TCP Tri-calcium phosphate

TEOS Tetraethyl orthosilicate

TEP Triethyl phosphate

XRD X-ray diffraction

XIII

Glossary

Adipocytes: also known as lipocytes and fat cells, are the cells that primarily

compose adipose tissue, specialized in storing energy as fat.

Alkaline phosphatase (ALP, ALKP): a hydrolase enzyme responsible for

removing phosphate groups from many types of molecules, including nucleotides,

proteins, and alkaloids.

Biodegradability: describes the ability of the scaffold material to degrade in vivo.

Bioglass: a commercially available family of bioactive glasses, composed of SiO2,

Na2O, CaO and P2O5 in specific proportions.

Chondrocytes: the only cells found in cartilage.

Cytotoxicity is the quality of being toxic to cells.

Dexamethasone: a potent synthetic member of the glucocorticoid class of steroid

drugs.

Furnace: a device used for heating

Glutaraldehyde: an organic compound with the formula CH2(CH2CHO)2.

Hepatocyte: a cell of the main tissue of the liver.

Mesenchymal stem cells: multipotent stem cells that can differentiate into a

variety of cell types, including: osteoblasts (bone cells), chondrocytes (cartilage

cells) and adipocytes (fat cells).

Mesoporous material is a material containing pores with diameters between 2

and 50 nm.

Neovascularization: a formation of functional micro-vascular networks with red

blood cell perfusion.

Osteoblasts: mononucleate cells that are responsible for bone formation

XIV

Osteoconductivity An ability of the scaffold to serve as a template for bone

formation by encouraging cells to adhere to the surface and to proliferate and

produce bone

Osteoinductivity of a scaffold refers to the ability to induce bone formation

without osteoinductive agents, such as bone morphogenic proteins (BMPs)

Osseointegration derives from the Greek osteon, bone, and the Latin integrare, to

make whole.

Phase transition: the transformation of a thermodynamic system from one phase

or state of matter to another

Scaffolding: a temporary structure used to support material in the construction or

repair of some large structures

Sintering: a method for making objects from powder, by heating the material in a

sintering furnace

Solid freeform fabrication techniques: a collection of techniques for

manufacturing solid objects by the sequential delivery of energy and/or material

to specified points in space to produce that solid.

Trypsinization: a process of using trypsin, a proteolytic enzyme which breaks

down proteins, to dissociate adherent cells from the vessel in which they are being

cultured.

Wollastonite is a calcium inosilicate mineral (CaSiO3) that may contain small

amounts of iron, magnesium, and manganese substituting for calcium.

XV

Statement of Original Authorship

The work contained in this thesis has not been previously submitted to meet

requirements for an award at this or any other higher education institution. To the best

of my knowledge and belief, the thesis contains no material previously published or

written by another person except where due reference is made.

2011/7/13

XVI

Acknowledgements

I would like to take this opportunity to express my gratitude and appreciation to

everyone that helped me in some way whilst completing this research. Special

recognition and thanks goes to my supervisor associate professor Cheng Yan for

his kind guidance, providing different solutions and inspiring motivation towards

my master thesis. He was also the person who suggested and introduced to me

how to work with my topic and how to enhance my skills. He has provided his

valuable time in discussing, going through my drafts, providing comments and

advising me on how to improve my work from time to time.

Also apart from my supervisor, I would like to thank some professional people

from different organization and institution for their suggestion, comments,

information, materials and other help. They are Simon Miao and Clayton Adam,

Greg Peterson, and Melissa Johnson.

XVII

Lists of Publications

1. Wei Zheng, Cheng Yan, Feng Lin, Wei Fan, Clayton Adam, AdeKunle

Oloyede. (2010) Preparation of porous tri-calcium phosphate ceramic scaffold

for bone tissue engineering. First International Conference on Cellular,

Molecular Biology, Biophysics and Bioengineering (CMBB 2010). 3: 300. Dec.

2010. Harbin, China

2. Feng Lin, Cheng Yan, Wei Zheng, Wei Fan, Clayton Adam, AdeKunle

Oloyede. (2010)Preparation of mesoporous bioglass coated zirconia scaffold

for bone tissue engineering. First International Conference on Cellular,

Molecular Biology, Biophysics and Bioengineering (CMBB 2010). 3: 330. Dec.

2010. Harbin, China

3. ―Preparation and characterisation of strong and bioactive tri-calcium

phosphate scaffolds with tunnel-like macropores for bone tissue engineering‖,

in preparation

4. ―Mechanical and biological properties of sol-gel derived zirconia scaffold

coated with mesoporous bioglass‖, in preparation

Chapter 1 Introduction

1.0 Introduction

In this chapter, the author‘s motivation for embarking upon research into the

specified area is introduced and explained. Background to the research topic is

presented first, followed by an introduction of the research goals and objectives of

this work.

1.1 Background

Tissue engineering applies methods from materials engineering and life sciences to

create artificial constructs for regeneration of new tissue [1]. Even though a range

of tissues has been studied, the translation of engineered tissues to clinical

applications has been limited. Bone tissue engineering has the potential to reach

millions annually through the repair of bone defects caused by disease, trauma or

congenital defects. In 2003 the potential market for tissue engineered products for

musculoskeletal applications totaled $23.8 billion in the USA and is expected to

rise to $39 billion by the year 2013 [2]. In 2004 alone there were $1.5million bone

graft procedures [3]. In the United States alone, at least eight million surgical

operations were carried out annually, requiring a total national healthcare cost

exceeding $400 billion per year [4, 5]. Autografts (from the patient) were

considered the gold standard for bone defect repair and allografts (from a donor)

were also commonly used. While multiple complications and risks were

associated with the use of both types of grafts [6–8], these remain appropriate

options for some simple and non-load-bearing defects that do not require a

2

significant amount of graft material (i.e. non-critical size defects). However, for

many defects the use of allogenic and autologous bone is not an option.

Researchers in bone tissue engineering are working to develop alternatives to

allogenic and autogolous bone grafts in order to address the growing needs of the

population, and much of the research is scaffold based. Thus the development of

interconnected porous scaffolds plays a significant role in bone tissue engineering,

especially in the restoration of large bone defects [9]. A three-dimensional (3D)

scaffold is usually used as an artificial matrix for bone regeneration and defects

repair or combined with cells and/or ―biologics‖, which are added to further

enhance bone regeneration [10]. However, the optimal scaffold ‗‗recipe‖,

including target mechanical properties, is still very much under debate. There are

several characteristics that are considered to be essential for bone scaffolds, such

as biocompatibility, osteoconductivity and interconnected porosity. Other

considerations in bone scaffold design and optimization include biodegradability,

permeability and mechanical integrity.

Calcium phosphate ceramics have been extensively used to produce porous

scaffolds due to their bone-like chemical composition as well as excellent

biological properties, including biocompatibility, osteoconductivity and

osteoinductivity [11]. In the calcium phosphate compounds, hydroxyapatite

(Ca10(PO4)6(OH)2, HA), the main inorganic component of natural bone, has been

widely investigated since calcium phosphates are recognized as ceramic materials

that simulate the composition and mineralogical structure of natural bone [12].

Nonetheless, because of the weak mechanical properties of HA ceramics, such

3

materials cannot be used as implant devices for replacing human bones [13]. It is

widely acknowledged that the incorporation of a ceramic reinforcement (i.e. fibers,

whiskers, platelets or particles) in a ceramic scaffold improves the mechanical

properties [14]. Thus, many studies have added some bioactive glass into the HA.

Bioglasses are popular biomaterials, because their compositions are similar to the

inorganic constituent of the mineral part of bone [15]. As a result, bioglass

addition greatly improves HA‘s compressive strength, largely due to improved

densification through the presence of a liquid phase during sintering [16]. When

dense HA ceramics are implanted, low resorption rates of HA hinder bone

ingrowth, resulting in chemical bonding only at the interface between the bone

and HA implant. This low biodegradability is the disadvantage of dense HA

ceramics, which limits the wide application of bulk HA [20, 21]. Compared to HA,

tri-calcium phosphate is generally considered as a resorbable bio-ceramic [20].

TCP ceramics display better biodegradability than HA and tend to be replaced by

bone as they degrade. Experimental studies have also showed that short sintering

time and limited glass addition such as 2.5–5wt%, can be used to control phase

transformation between HA and TCP [17-19]. It was reported that greater

amounts of phase transformation occurred with the addition of 5wt% bioglass [18].

However, to the author‘s knowledge there is no research been conducted to

investigate the processing parameters such as temperatures and addition of

bioglass into HA to obtain TCP. Therefore, in this project, the HA/bioglass

composite is sintered at high temperature to obtain the TCP ceramic, and the

novel macrotube scaffolds were designed to help the bone formation.

4

1.2 Research goals

The overall goal of this research is to create a macro-tube porous tri-calcium

phosphate scaffold through the use of composite hydroxyapatite and bioactive

glass, sintered at high temperature. The purpose of using macro-tubes is to

provide necessary oxygen and nutrients, which flow through the tube in order to

promote osteoblasts. The pores on the inner-walls of each macro-tube are used for

cell loading, ingrowth, and tissue-formation. The mechanical properties of

tri-calcium phosphate scaffolds depend on scaffold structure, the bioglass

additions and the sintering temperature. The biological properties are based on the

scaffold‘s structure and material properties.

1.3 Innovations

In the past decade, experiments were usually performed using

polyurethane-sponge [24-27] as a sacrificed scaffold, for the structure of the

sponge is fairly similar to trabecular bone. However, the high porosity sponge

ceramic scaffold has low mechanical property. As an improvement, the

innovations of this project were that the sacrificed scaffold would be designed as a

controlled macro-tube scaffold in order to control scaffold‘s porosity and

mechanical property. Furthermore, in past experiments [64, 66, 95], tri-calcium

phosphate ceramics were usually obtained by using more complex methods (see

section 2.6), however in this project, the TCP ceramic would be produced by

mixing HA and bioglass and sintering at high temperatures. It was known that

when HA reacts with bioglass at high temperatures, TCP and other phases

5

(depending on the bioglass) can be formed, which has a higher mechanical

property than conventional methods. Finally, the porous TCP ceramic scaffold

would be coated with bioglass to achieve greater mechanical property and

biocompatibility.

6

Chapter 2 Literature Review

2.0 Introduction

Chapter 2 reviews the literature relevant to this research field. This chapter is

divided into five components, (1) the importance of scaffolds in tissue engineering,

(2) scaffold design, (3) choice of materials, (4) current TCP preparation technique

and (5) cell culture. Through this chapter, the reader can understand the

importance of the role that porous ceramic scaffolds playing in tissue engineering.

They are required to be biocompatible, biodegradable and osteoinductive. The

scaffold can be fabricated using multiple techniques. However the solid freeform

fabrication technique (SFFT) is one of the most advanced one. By using this

technique, the porosity of the scaffold can be controlled, and consequently the

compressive strength of the scaffold can be controlled as well.

Various materials have been used for fabricating bone tissue engineering scaffolds.

Calcium- phosphate ceramics are the most common material used for the ceramic

scaffold, and polymers are commonly used for coating the ceramic scaffold.

However, there are some disadvantages for the polymer coating, thus the current

trend is towards the use of bioactive ceramic as the coating material. Moreover, in

recent studies, it was shown that in order to enhance the HA scaffold‘s

compressive strength, bioglass was added to HA. Some phases, including TCP

were formed when the composite material was sintered at high temperatures [22].

TCP is regarded as one of the most favorable bioactive materials for fabricating

porous ceramic scaffold [23]. Nevertheless, techniques for fabrication often vary.

In the TCP preparation section, some preparation techniques will be introduced,

7

and each technique‘s advantages and disadvantages will be listed. Unfortunately,

there have not been many studies on TCP fabrication using HA and bioglass

composites. The coating method will also be introduced at the end of this chapter.

2.1 The role of scaffolds in Tissue Engineering

Scaffolds for bone regeneration mainly serve an osteoconductive function, in

which new bone forms through creeping substitution from adjacent living bone

tissue [25]. Scaffolds can also act as three-dimensional vehicles for cell delivery

and tissue regeneration and to further enhance the regeneration process. A popular

approach is to expand bone progenitor stem cells in vitro and disperse them into

scaffolds to stimulate osteogenic differentiation, followed by implantation onto

the site of the bone defect. An ―ideal‖ scaffold for bone tissue engineering would

have certain characteristics, which are high biocompatibility and biodegradability

[221]. HA is an example of a biodegradable and biocompatible scaffold material,

as approved by the US Food and Drug Administration [222]. Other ideal

characteristics are osteoinductivity (actively inducing bone formation) or

osteoconductivity (guiding and supporting bone regeneration) [26-27]. Bioglass

and calcium phosphate ceramics are typical osteoconductive materials.

Bone that is damaged by accident or degenerated as a result of age and disease

often needs to be reconstructed by surgical treatments. Therefore, synthetic

bone-graft substitutes are in great demand [28]. The relationship between scaffold

structure and tissue compatibility has been studied previously [29-33]. Optimal

scaffolds for cell loading, ingrowth, and tissue-formation are basic problems in

8

synthetic bone-graft applications. Cellular composites are then seen as consisting

of three main structural components: (1) cells that are organized into functional

units, (2) extracellular matrix, and (3) scaffold architecture [223-224]. This

architecture is increasingly believed to contribute significantly to the development

of specific biological functions in tissues and thought to provide appropriate

nutritional conditions and spatial organization for cell growth. The regeneration of

specific tissues aided by synthetic materials has been indicated to be dependent on

the pore size and porosity of the supporting three-dimensional structure [34]. A

large surface area is necessary for cell attachment and growth. A large pore

volume is also required to accommodate and subsequently deliver a cell mass

sufficient for tissue repair. Highly porous biomaterials are also desirable for the

easy diffusion of nutrients to and waste products from the implant and for

vascularization, which are important requirements for the regeneration [36]. The

surface area/volume ratio of porous materials depends on the density and average

diameter of the pores. Nevertheless, the diameter of cells in suspension dictates

the minimum pore size, which varies from one cell type to another. Depending on

the envisioned applications, pore size must be carefully controlled. The effect of

implant pore size on tissue regeneration is emphasized by experiments

demonstrating optimum pore size of 5m for neovascularization, 5-15m for

fiberblast ingrowth, close to 20m for the ingrowth of hepatocytes, 20-125 m for

regeneration of adult mammalian skin, 40-100m for osteoid ingrowth [35], and

100-350m for regeneration of bone [36] (See Table 1). Fibrovascular tissues

appear to require pores sizes greater than 500m for rapid vascularization and for

the survival of transplanted cells [37]. While the increase in porosity and pore size

9

would favor bone ingrowth and nutrition supply, increase in porosity and pore

size would consequently decrease the biomechanical strength [38].

Table 1 Studies defining optimal pore size for bone generation

Reference Scaffold pore size

(mm) Porosity (%)

Mineralize tissue

ingrowth/comments

Klawitter et

al.[36] Type I: 2–6 m 33

No tissue ingrowth

(22weeks)

Type II: 15–40 m 46.2

No bone ingrowth,

fibrous tissue ingrowth

(22weeks)

Type III: 30–100

m(80% pores <100

m)

46.9

50 m of bone

ingrowth, osteoid and

fibrous 80% pores , 100

m tissue ingrowth

(22weeks)

Type IV: 50–100

m(63% pores <100

m)

46.9

20 m of bone ingrowth

by 11 weeks and 500

m of ingrowth by 22

weeks, osteoid and

fibrous tissue ingrowth

Type V: 60–100

m(37%<100 m) 48

600 m of bone

ingrowth by 11 weeks

and 1,500 m of

ingrowth by 22 weeks,

osteoid and fibrous

tissue ingrowth

Whang et al.[35] <100 m 35.3

Not statistically

different from untreated

controls

<200 m 51

Not statistically

different from untreated

controls

10

It is rather challenging to find a porous structure with a large pore size and a

suitable mechanical strength. A number of fabrication techniques have been

developed over the years for manufacturing porous bioceramics. These include,

amongst others, the replication of polymer foams by ceramic dip coating [39-40]

or impregnation, foaming of aqueous ceramic powder suspensions [41-42],

pyrolysis of pre-ceramic precursors [43] and firing of ceramic powder compacts

[44-45]. However, none of these methods can completely satisfy all the necessary

requirements. For instance, a controlled level of interconnected porosity combines

with good compressive strength (the strength in the region of 2-14MPa with a

porosity >50%) [46].

Trabecular bone can also be referred as spongy bone since the structure of the

Trabecular bone is quite similar with the sponge. Hence, sponges made from

Polyurethane (PU) are widely used as a template for making ceramic scaffolds.

Although PU sponge has very high volumes of porosity and excellent

interconnectivity levels, the mechanical properties of the sponge ceramic is poor

[40], and the pores of the sponge cannot be controlled.

2.2 Solid freeform fabrication techniques (SFFT)

SFFT, such as fused deposition modeling, have been employed to fabricate highly

<350 m 73.9

Statistically significant

more bone than all other

groups

11

reproducible scaffolds with fully interconnected porous networks [47-48] as

shown in Fig. 1. Using digital data produced by an imaging source such as

computer tomography or magnetic resonance imaging enables accurate design of

the scaffold structure [48]. Solid freeform (SFF) manufacturing combined with

conventional foam scaffold fabrication procedures such as sponge coating may be

used to develop scaffolds with controlled micro and macro-porous structures

[49-53]. This enables scaffold properties such as pore size, pore fraction, strength,

modulus and permeability to be varied systematically.

Figure 1 Photograph of a scaffold produced by SFF techniques (machined into a

cylinder)

Scaffolds fabricated using SFF technique is generally reported to have better

mechanical properties than those fabricated by conventional ceramic processing

routes [54]. Such biomimetic internal architectures may prove valuable for

multi-tissue and structural tissue interface engineering. To the authors‘ knowledge,

there is no literature available on degradable HA/bioactive glass made by SFF

Techniques. This technique has been only applied for composites containing

12

calcium phosphates as the bioactive phase [55-56]. For example, Xiong et al. [56]

fabricated composites of PLLA/TCP with porosities of up to 90% and mechanical

properties close to human cancellous bone by using low-temperature deposition

based on a layer-by-layer manufacturing method of SFF fabrication

(computer-driven by 3D digital models). PLLA was dissolved in dioxane and TCP

powder mixed to prepare slurry, which was formed into frozen scaffolds, and

subsequently freeze-dried. Alternate parallel layers formed macro-pores (400m

diameter) and sublimation of the solvent during freeze-drying formed micropores

(5m diameter). Taboas et al. [55] produced PLA scaffolds with computationally

designed pores (500-800m wide channels) and solvent-derived local pores

(50-100m wide voids or 5-10m length plates). Indirect fabrication using casting

in SFF moulds provided enhanced control over scaffold shape, porosity and pore

architecture, including size, geometry, orientation, branching and

interconnectivity. One of the major disadvantages of this method is to increase

scaffold fabrication time compared with direct methods, a temporary mould must

be made first. The macro-pore size and volume fraction are controlled by the

combination of the rod diameter and spacing in and out of plane. Details of

macro-porous scaffolds from several independent studies are shown in Table 2.

13

Figure 2 The top row shows the rod diameter and the out-of-plane rod spacing.

The bottom row shows the in-place spacing. Cordell et.al [57]

Table 2 Summary of scaffold geometry, porosity and compressive properties for

the 3D porous scaffolds fabricated by SFFT

(*NR, not report by the author)

Reference Cordell

[57]

Woodard

[58]

Dellinger

[59]

Miranda

[60]

Unit cell

geometry

(m)

Rod

diameter 394 394 415 415 570 570 220 220

Edge to edge

(out of

plane)

252 252 315 315 280 280 100 100

Edge to edge

(in plane) 359 359 315 315 270 270 80 80

14

The essential characteristics of scaffolds for using in bone generation and repair

are outlined below. Other characteristics that are usually considered in scaffold

design are also discussed. As no single study has addressed all these, many are

coupled and, therefore, difficult to assess.

1) Biocompatibility is among the most important scaffold characteristics. It is

defined broadly as the ability of a material or device ‗‗to perform its intended

function, including an appropriate degradation profile, without eliciting any

undesirable local or systemic effects in that host‖ [61]. Host responses, both

positive and negative, may include osteoblast/osteoclast response, prolonged

inflammation, microvascular changes, fibrous encapsulation, protein

adsorption and endothelial proliferation [62-63].

Porosity

Calculated

(or stated)

macro-pore

fraction

50% 50% 41% 41% 28% 28

% 28% 28%

Micro-pore

size (m) 1-10 2-15 2-8 2-8 1-30

1-3

0 NR* NR*

Strength in

Compressio

n

Compressive

Strength

(MPa)

7.8 8.3 101.5 53.9 302.3 10.4 47 14

Test Sample

Geometry

Test

geometry

Cylind

er Cylinder Cylinder Cube

Height (mm) 8.1 2 4.5 10

15

2) Osteoconductivity is the ability of the scaffold to serve as a template for bone

formation by encouraging cells to adhere to the surface and to proliferate and

produce bone.

3) Interconnected porosity is required for nutrient and waste transport throughout

and for bone growth. The minimum pore size for bone formation has been

quoted by many as 100 m [64-69]. However, more recently researchers have

shown bone formation in interconnected micropores less than 10m in size in

scaffolds that contained both macroporosities (>100m) and microporosities

(<10m) [70-71].

4) Biodegradability describes the ability of the scaffold material to degrade in

vivo. CaPs degrade relatively slowly by physiochemical, cell mediated or

mechanical degradation mechanisms [72-73].

5) Bioactivity is the tendency of the material to form a chemical bond with the

host bone. For CaPs this is postulated to occur through material dissolution

and precipitation of a carbonated apatite that is more similar to the mineral

phase of bone. Some factors that affect bioactivity are composition,

crystallinity, grain size and impurities.

6) Osteoinductivity of a scaffold refers to the ability to induce bone formation

without osteoinductive agents, such as bone morphogenic proteins (BMPs).

7) Permeability is the ease with which fluid can flow through a porous material.

It depends on the pore size, fraction and fenestration, which are all dependent

upon the scaffold pore architecture.

16

8) Manufacturability refers to the ease and flexibility of scaffold fabrication.

Solid free-form fabrication techniques are a common fabrication route

[74-75].

9) Mechanical integrity is a broad term that encompasses all mechanical

properties from post-manufacture through to complete healing.

2.3 Target mechanical properties for bone scaffolds

The need for scaffolds that can be used to repair load-bearing bone defects, which

are often also large defects, is apparent. However, the specific mechanical

requirements of the scaffolds being studied for the repair of such defects,

beginning from implantation through to complete healing, are yet to be

established. The most appropriate materials and their corresponding mechanical

properties are still under debate. The target properties such as strength and elastic

modulus that have been explicitly stated or implied in the literature, span several

orders of magnitude and several different approaches have been taken with regard

to the design for specific mechanical properties. For example, a number of

literatures have stated that bone scaffold properties should match those of natural

bone [76-80]. Table 3 summarizes the compressive, flexural and tensile strengths,

elastic moduli and porosities for both cortical and cancellous bone for referencing

purpose. Another study took a different approach and optimized scaffold pore

architecture such that the scaffold stiffness matched the stiffness of native bone

initially and the stiffness of the regenerated bone matched that of native bone at

the end point [79]. In a different paper the authors stated simply that ―the

mechanical properties of the scaffold must be sufficient and not collapse during

17

handling and during the patient‘s normal activities‖ [81]. In contrast to these

studies, another argued that scaffold strength should be greater than the bone it

will replace [82]. The same authors indicated a need for fixation devices to shield

scaffolds from loading. These examples from the literature demonstrate the wide

range of target mechanical properties for bone scaffolds.

Table 3 Summary of mechanical properties and porosity of human bone

*NR, not reported by author.

A final strategy is to design scaffolds such that the mechanical properties of the

composite of bone and scaffold are within some percentage of the mechanical

properties of the host bone.

To achieve this, the initial mechanical properties of the scaffold should account

for the change in properties with degradation and the change with the expected

bone in-growth. It was discovered that a high order for some systems whose

degradation rates are sensitive to processing parameters and the in vivo

environment or simply are not well characterized. On the other hand, degradation

18

of CaPs can be of the order of months to years, depending on porosity,

crystallinity and grain size as shown in the study by Grynpas [90].

CaPs is an excellent candidates for use in bone replacement and repair. CaPs can

have strengths and stiffnesses that are similar to those of cortical and cancellous

bone in some forms [91-99, 81, 65]. The real limiting factor for CaPs in

load-bearing applications may be in the inherent brittleness of this class of

materials. Unfortunately, few studies in the literature report properties such as

fracture toughness, reliability (i.e. Weibull modulus), or energy to failure for CaPs

or composite scaffolds. CaPs usually degrade slowly and in a controlled manner.

The stiffness and strength of the CaP bone composite is reasonably stable overall.

However, the issue of brittle behavior has yet to be resolved.

2.4 Choice of biomaterials

In tissue engineering, many types of materials have been applied for bone repair

and regeneration; these biomaterials in general can be divided into several groups:

metals, ceramics, polymers, and composites [100]. Ideally, materials and scaffolds

utilized in tissue engineering should meet some prerequisites including

mechanical properties and integration with host [101-102], enhanced

osteoinductive, osteoconductive properties and material biocompatibility [103]. In

addition, porosity, pore diameter, interconnectivity and microstructure should also

be considered for bone repair and regeneration as well [104]. Many studies have

demonstrated that hydroxyapatite (HA) or other calcium phosphate (CaP)

ceramics including tricalcium phosphate (TCP) and bioactive glasses can improve

19

the formation of a bone-like apatite layer on their surfaces [105]. The inorganic

ceramic phase can be combined with polymers or polymer precursors to produce

bioactive and biodegradable composites [106]. Therefore, the new strategies for

creating a CaP composition similar to bone apatite will be advantageous.

2.4.1 CaP bioceramics for bone replacement and repair

Calcium phosphate (CaP) is considered the most popular biomaterial for

exogenous bone grafts in bone reconstruction [107-109]. Calcium phosphates

were first considered for clinical application as a filler for bone defects in the

1920s and first incorporated in dentistry and orthopedics in the 1980s [33]. The

interest in CaP based ceramics for bone replacement and repair is well-deserved,

given that they have the requisite characteristics and many other attributes that

make them excellent candidates for such applications. CaPs are biocompatible,

have a composition and structure similar to the mineral phase of bone [91]. They

are osteoconductive [110-113] and they have been reported to be intrinsically

osteoinductive in some cases [110, 114, 64]. They are also bioactive.

The degradation products of CaPs are conducive to bone formation and strongly

linked to bioactivity; the dissolution process is followed by reprecipitation of a

carbonated apatite, which has a composition and chemical structure that is similar

to the mineral phase in bone [61, 76, 115]. The degradation rate for CaPs is

typically slow compared to that of many polymers [81] and even comparable with

bone growth [116]. This is advantageous because it addresses concerns about

balancing degradation and bone in-growth. A range of CaPs compositions have

been considered, hydroxyapatite (HA), with the chemical composition

Ca5(PO4)3OH and a calcium to phosphate ratio of 1.67 [117], and -tricalcium

20

phosphate (-TCP), with the chemical composition Ca3(PO4)2 and a calcium to

phosphate ratio of 1.5 [93]. Among these, HA is most commonly used in clinical

applications and has been used in bone cements for the repair of craniofacial

defects [118-119], for maxillary sinus floor augmentation [120] and in coatings

for hip replacements [121,122]. -TCP ceramics have been commercialized as

bioresorbable synthetic bone substitutes and are used in orthopedic and dental

applications [123], including augmentation of the alveolar ridge [124], sinus

reconstruction [125] and general bone reconstruction following injury or disease

[126]. Perhaps the most emphasized difference between the properties of HA and

-TCP is their relative degradation rates; HA is considered relatively

non-degrading while TCPs purportedly degrade readily [127-129], which has been

seen in previous studies. In Fang‘ clinical report, -TCP could help the bone

replacement in animal experiments [133].

TCP has three polymorphs: low temperature phase, β-TCP is presented below

1180°C, while α-TCP appears in the temperature range 1180-1400°C and α′-TCP

is observed above 1470°C [130-131]. The phase transformation is accompanied

by density changes for calcium phosphate materials. The density decreases with

phase transformation in the following sequence when heat-treated to higher

temperatures [132]

HA (3.14 g*cm-3

) >β-TCP (3.07 g*cm-3

) >α-TCP

(2.77 g*cm-3

) >α′-TCP (1)

Among the three allotropic forms of TCP, β-TCP is the preferred as a bio-ceramic

on account of its mechanical strength, good tissue compatibility, and ability to

bond directly to tissue to regenerate bone without any intermediate connective

21

tissue. In addition, fast bone regeneration and proper bioresorption rate are other

additional attributes of β-TCP [132]. It has been documented that the dissolution

rate of β-TCP is 3-12 times faster than HA [133]. In vitro studies reveal that

α-TCP exhibits a higher dissolution rate than β-TCP [132-134]. The order of

relative solubility is

α-TCP>β-TCP≫HA[132,134] (2)

Since tri-calcium phosphate exhibits higher solubility, it is expected that it can

potentially degrade upon implantation in the host and may be gradually replaced

by the newly formed regenerated bone.

2.4.2 Surface coating

A success of an implant also depends on its surface chemistry, which determines

the interactions at the implant material-tissue interface. To elicit desirable

material-host tissue interaction, the implant may need to be coated with suitable

coatings. Another important reason for a suitable surface coating is the wear of the

implant surface being in contact with the host tissues. For example, the hardness

of the bone leads to very heavy abrasion by fretting or direct wearing as soon as

the interfacial strains between the implant and the hard tissue occur. The final

reason for the coating is the coating material could improve the mechanical or

biological properties for the scaffold, and this reason is the most important for the

CaPs ceramic scaffold coating.

There are two main possibilities for kinds of the coating materials: polymer and

ceramic. As it was mentioned previously, CaPs make excellent candidates for

using in bone replacement, however, it has been reported that one of the limiting

factors for the CaP in load bearing applications is poor strength [117,135,136].

22

Polymer with excellent mechanical properties can improve CaPs scaffolds‘

strength, reaching to a level which is similar to the cancellous bone strength [137].

However, there are some disadvantages for the polymer coating. Poly

(lactic-co-glycolic acid), (also commonly known as PLGA) is a type of bioactive

polymer, which is used as the coating material and it has been approved by the US

Food and Drug Administration (FDA). Unfortunately, the chemical used to

dissolve PLGA may possess potential health and safety risks. When preparing

PLGA, it is usually required to be dissolved in dichloromethane (CH2Cl2)

solution first. Dichloromethane (CH2Cl2) is a toxic chemical and can

potentially damage the human nervous system. Recent studies have shown that

although the polymer coating can largely increase the compressive strength, it

decreases the osteoconductivity of the scaffold [138]. Another disadvantage is

that after applying polymer coating, the scaffold surface will become very

smooth. This could discourage cell growth as rough surfaces result in good

osseointegration as compared to smooth surfaces, it was shown that the number

of Alp (alkaline phosphatase) -positive cells on the rough surface was two to three

times higher than that on the smooth surface [139]. Another coating material is

bioactive ceramic. Many researchers have used bioactive ceramics (for

instance HA, bioglass) to improve scaffold‘s mechanical and biological

properties [140-146]. Normally, these kinds of scaffolds have high mechanical

properties, and the base materials for those scaffolds were bioinert ceramics or

metals. After the coating film is degraded, the base material will be connected

with the host-tissue, which can have detrimental effects on the bone

regeneration process. Therefore, bioactive ceramic base with bioactive

23

ceramic coating is the best way to improve scaffold‘s mechanical and

biological properties.

2.4.3 Composite mechanics

One of the main purposes of creating ceramic–matrix composites (CMCs) is to

improve the material‘s fracture toughness, or its ability to resist fracture when a

crack is present. Low strength in tension compared with compression is a major

problem with brittle ceramics. In tension, even the slightest flaw, such as a surface

or interior crack, internal pore, or grain corner, will amplify the applied stress.

Therefore, as the two go hand in hand, an improvement in toughness generally

leads to improvements in strength and stiffness as well [147]. The reinforcing

phase in a ceramic serves to prevent crack growth. This can be accomplished by a

number of methods, including deflecting crack tips, forming bridges across crack

faces, absorbing energy during pullout, and causing a redistribution of stresses in

regions adjacent to crack tips [148]. In addition to the reinforcing material‘s

mechanical properties, its morphology is critical.

2.4.4 Bioglass

Bioglass belongs to a family of bioactive glasses containing various proportions

of chemical constituents. When incorporated into a damaged part of the human

body, it is able to associate with bony structure, promoting repair and bringing it

to normal function. The difference of bioglass from traditional soda-lime glasses

is that the proportions of SiO2, Na2O, CaO and P2O5 are different. Bioglass has a

lower amount of silica, but more sodium and calcium, and high

calcium/phosphorus ratio. However, due to the weak mechanical strength of

24

bioglass, this material can only be used in applications subjected to a lower load.

Currently, its main use includes coating metal, increasing the biological

properties.

2.5 Material and scaffold characteristics that influence

the mechanical properties of CaPs

There are a number of main characteristics that influence the mechanical

properties of CaPs, as well as other ceramics. It is well known that the

characteristics that influence mechanical properties can be controlled by the

processing parameters [78], because this fabrication process requires high

temperatures and long heat treatments.

2.5.1 Porosity

Pore size, fraction, distribution and architecture have a strong influence on the

mechanical properties of CaPs and ceramics in general [64, 66, 95, 150-152]. One

of the most common representations of the relationship between porosity and

density is given by Eq. [153]

Total Porosity =1- M/ (V x ) (3)

In which M is weight of the sintered porous scaffold, V is volume of the sintered

sample porous scaffold,is the solid density of the ceramic. It is well accepted

that an increase in pore size leads to a decrease in strength. Pore geometry also

affects the strength [152] through the local stress concentration introduced by the

pores.

25

2.5.2 Phase transformation

The phases present in the CaPs bulk materials and scaffolds, determined by the

average composition as well as by the heat treatment [154], affect the mechanical

properties [155]. However, the trends are not well characterized and the

mechanisms are not well understood. The HA, TCP and bio-glass materials are

mainly consist of Ca, P, OH atoms, which are also the main constituents construct

main composition of human hard tissues. In general, sintering HA can lead to the

partial thermal decomposition of HA into tri-calcium phosphate (TCP). There are

two steps in the thermal decomposition i.e. dehydroxylation and decomposition.

Dehydroxylation to oxyhydroxyapatite proceeds at temperatures about 850°C to

900°C by the fully reversible reaction in accordance to equation 4 [156-157]:

Ca10 (PO4)6(OH)2 Ca10(PO4)6(OH)2-2xOx + xH2Ogas (4)

The decomposition to TCP and tetracalcium phosphate occurs at temperatures

greater than 900°C: According to the reaction given in equation 5 [156-157]:

Ca10 (PO4)6(OH)2 2Ca3(P04)2 (TCP)+Ca4P2O9+H2Ogas (5)

When bioglass is added to HA for the purpose of increasing HA‘s mechanical

properties, it is shown that HA can be reacted with bioglass at the critical

temperature, and this results in TCP and other phases (depends on the

composition in bioglass) [158-162]. For example, in Seema.k‘s experiment [22]

where P2O5, CaO, and Na2O were used for making bio-glass, and 10% bioglass

was added to HA at the temperature of 1250oC, the results showed the primary

26

ceramic present is hydroxyapatite, and β-TCP (Ca3(PO4)2) as the secondary phase.

No other phases like α-tricalcium phosphate (α-TCP) and calcium oxide (CaO)

were detected. In another experiment, A.Carr [161] mixed P2O5, CaO, Na2O,

Al2O3, and B2O3 to fabricate bioglass. When 25% bioglass was added at the

temperature of 1200oC, the amount of HA occupation decreased from 75% to

30%, and 12.6% of β-TCP was found in the final powder. The phases were

possibly some β-Na2Ca4(PO4)2SiO4 and wollastonite [163]. Therefore, it can be

concluded that phase transformation is an important factor for fabricating calcium

phosphate ceramic, and the phases formed depends on three parameters: 1) the

quantity of bioglass added into HA, 2) sintering temperature, and 3) the

composition of bioglass.

2.5.3 Microstructure

Crystallinity and grain size also affect the mechanical behavior of CaPs. A more

prominent crystalline phase, lower pore fraction, and smaller grain size are all

associated with increased stiffness, compressive and tensile strengths, and fracture

toughness [81]. In general, the strength for ceramics is proportional to the inverse

square root of the grain size [150], similar to the relationship to metals.

2.5.4 Binder

When a low ceramic content is employed to increase the porosity, cracks are

prone to occur in the sample, because of shrinkage during sintering, which,

consequently, causes a severe reduction in strength. So in most of the previous

experiments, use of binder is a requisite for sintering ceramic. Polymers are

usually used (i.e. Polystyrene (PS), Polyvinyl acetate (PVA)), as the binder, since

27

it would be expected to improve the strength of the body, which, in turn, would

prevent it from cracking [157]. The presence of small pores formed by the

removal of the polymer on the ceramic during sintering, the polymer could also be

used for increasing ceramic‘s porosity [164]. However, if polymer is over-added

into the ceramic, the mechanical properities would decrease instead. Thus, the

content of the binder in ceramic is one of the most important parameters. In

Yook‘s experiment, where a polystyrene (PS) polymer was used as the binder, he

pointed out that the compressive strength of the porous HA scaffolds was

significantly affected by the PS content, when increasing PS content from 0 to

20vol.%, the compressive strength of the sample was significantly increased.

However, a higher PS content of 30vol. % was observed to lead to a lower

compressive strength [165]. Safronova‘s report indicated that the presence of 0.25%

– 0.50% PVC (Polyvinyl chloride) strongly influences the mechanical properties

of the powder [166].

2.6 TCP preparation technique

TCP has been shown to be resorbable in vivo with new bone growth replacing the

implanted TCP [167]. This property imparts significant advantage onto TCP

compared to other biomedical materials that are not resorbed and replaced by

natural bone. Conventionally, β-TCP powders are synthesized via a solid-state

process [168-171] and wet-chemical method [172-176], as described by the

following equations

Wet-chemical methods

28

Ca(NO3)2+2(NH4)2HPO4+2NH4OHCa3(PO4)2+6NH4NO3+2H2O (6)

Ca(OH)2+H3PO4Ca3(PO4)2+2H2O (7)

Solid-state process

Ca(H2PO4)2+2Ca(OH)2Ca3(PO4)2+4H2O (sintered at 900oC) (8)

There are some other methods for fabricating TCP powder. Table 4 summaries

the advantage and disadvantage of each method.

Table 4 Summary of advantages and disadvantages of each method for TCP

preparation

Preparation

techniques

Average

particle sizes Advantages Disadvantages

Wet-chemical method

[172-176] 0.64m

Simple and low

cost

Takes a long time

Particle sizes are

not even

solid-state process

[170-171] 0.1~1.0m

Good crystalline

Particle sizes are

not even

Other phases exist

Hydrothermal

[177] ~1.0m

Particle sizes are

bigger Not easy to control

Mechanical–Chemical

Synthesis

[178]

0.13~0.14m

Simple and could

be completed in a

short time, also the

particles sizes are

even

Should be

conducted in strict

conditions

Should be done in

manually

29

2.7 Application of bone marrow mesenchymal cells in

bone repair

Bone marrow cells consist of progenitor cells: the hematopoietic cell system and

non-hematopoietic human bone marrow mesenchymal stem cell system (hBMSCs)

[181-182]. Hematopoietic stem cells (HSCs) in bone marrow are the reservoir of

various blood cells, such as erythrocytes, leukocytes, macrophages or platelets

[183-184]. This mixture of multipotent progenitor cells can differentiate into

various other mesodermal cells, such as osteoblasts, chondrocytes, fibroblasts or

adipocytes [185]. When these cells are cultured in vitro, hBMSCs quickly adhere

and can be easily separated from the nonadherent hematopoietic cells by repeated

washing. hBMSCs form colonies and are therefore defined as Colony Forming

Units—Fibroblasts (CFU-Fs), each colony derives from a single proliferating

progenitor [186]. An approach to distinguish the subsets of hBMSCs with the

most active replication and individual differentiation potential would certainly be

very important for both theoretical and applicative reasons. Some laboratories

have developed monoclonal antibodies in order to identify one or more markers

suitable for hBMSCs identification and sorting [187-189]. As a mixture of

Sol-gel processing

[179] Nano size

Particle sizes are

small

Easily introduce

other impurities

Chemical Synthesis

[180] 50nm

Average particle

sizes are smaller

Easily introduce

other impurities

30

multipotent cells, bone marrow cells have the latent ability to differentiate into

different cell lineages depending on the cell culture environment, a fact that makes

applying bone marrow cells as a cell source for various tissue engineering

applications possible, especially in bone tissue engineering. As multipotent adult

progenitor cells (MAPCs) and a pool of cells are committed to various lineages

and stages of differentiation, the hBMSCs contain potentially osteogenic colonies.

These colonies can be referred to as ―skeletal stem cells‖ meaning they are

ancestors to mesodermal supportive tissues such as bone, cartilage, fibrous tissues,

and so on [190]. These ―skeletal stem cells‖, however, stay in different stages of

differentiation. Some are tripotential and are able to differentiate into osteoblasts,

chondrocytes, and adipocytes in vitro; some are only able to bipotentially

differentiate into a chondrogenic-osteogenic phenotype, whereas the others

differentiable only into osteogenic clones. Interestingly, over a prolonged time

cultured in vitro, the differentiation potential is reduced and the tripotential clones

progressively lose adipogenic, then chondrogenic phenotypes, and finally

osteogenic potential, suggesting a preferential commitment of the progenitors

towards the osteogenic phenotype [191]. Many studies have shown the osteogenic

abilities of hBMSCs, both in vitro and in vivo, and applied it to bone defect

regeneration [192]. It is undoubted that under certain conditions, hBMSCs are

able to differentiate into osteogenic progenitor cells and finally osteoblasts,

thereby able to regenerate tissue of bone defects by direct orthotopic placement in

conjunction with appropriate scaffolds, most commonly those containing

hydroxyapatite/tricalcium phosphate (HA/TCP).

31

2.8 Coating methods

2.8.1 Dip coating

Dip coating refers to the immersing of a substrate into a tank containing coating

material, removing the piece from the tank, and allowing it to drain. The coated

piece can then be dried by force-drying or baking. It is an easy and economic way

to coat scaffolds. Fig.3 shows the dip coating processing.

Figure 3 Dip coating processing [218]

2.8.2 Sol-gel method

Many different coating methods have been used to combine mechanical and

biological durability of coating films for the best performance of the implants.

Lately, the sol-gel method has come to prominence because of its significant

advantages, such as low crystallization temperature, enhanced substrate adhesion,

low cost, controllable microstructure, and excellent chemical homogeneity due to

atomic level mixing [142,193,194]. The milder thin film synthesis conditions of

sol-gel methods eliminate most of the defects that originate during plasma

spraying and lead to better structural integrity [195].

32

2.9 Summary

The main conclusions of this chapter are that porous ceramic scaffolds play an

important role in tissue engineering. For scaffold manufacture, solid freeform

fabrication technique is one of the most advanced methods in the world, so the

controlled porous scaffold could be fabricated by using this technique. Due to

some disadvantages using the polymer coating, calcium phosphate ceramics

could preferring be used for both base material and coating material. Furthermore,

many studies report that TCP could be formed by sintering HA and bioglass

composite, thus porous TCP scaffolds could be fabricated by this novel technique.

33

Chapter 3 Methodology and Sample

characterization

3.0 Introduction

Chapter 3 was focused on the materials and methodology of the project. HA and

bioglass were used as the raw materials to fabricate TCP. In this study, SFFT was

firstly used to fabricate a sacrificial scaffold. The sacrificial scaffold was then

coated with HA and bioglass composite slurry, the TCP scaffold was obtained

when the scaffold was sintered at 1400oC for 3hrs. In order to further increase the

scaffold's compressive strength, the porous TCP scaffolds were then coated with

bioglass. Results from the in vitro biological tests performed in this research will

be presented in Chapter 4. The sample characterization techniques used were

XRD analysis, total porosity of TCP scaffold, imaging of the scaffold by scanning

electron microscopy, mechanical testing and in vitro biological test (including

image of the cell, MTT test and Alkaline phosphatase activity). This study

examined effect of the macro-tube scaffold size on the mechanical properties of

tri-calcium phosphate scaffolds. Fig.4 is a flow chart illustrating the fabrication

route for the porous TCP scaffolds, and the corresponding mechanical and

biological tests conducted on each scaffold.

34

Figure 4 Flow chart showing the six main processing steps for fabricating TCP

scaffolds

3.1 Methodology

3.1.1 Preparation of TCP ceramic

3.1.1.1 Preparation of hydroxyapatite powder

Synthetic hydroxyapatite was prepared by reaction of orthophosphoric acid in

aqueous solution and calcium hydroxide [196], according to the following

equation:

35

6H3PO4(aq)+10Ca(OH)2(aq)Ca10(PO4)6(OH)2(s)+18H2O(l) (9)

The Ca(OH)2 (>=95%) and H3PO4 (85%) were both supplied by in Sigma-Aldrich,

AUS. The reaction was carried out at 85oC, with the pH maintained above 9.5 by

the addition of ammonia solution. X-ray diffractometry (XRD) indicated that

when sintered Synthetic hydroxyapatite in air at temperatures between 900oC and

1350oC, there was no secondary phase and was stable up to at least 1350

oC.

3.1.1.2 Preparation of bioactive glass powder

The method of preparation of bioglass followed previous publications [197].

Typically, 33.5g of tetraethyl orthosilicate (TEOS, 98%), 7g of Ca(NO3)2.4H2O,

3.65g of triethyl phosphate (TEP, 99.8%) and 5g of 37% HCl were dissolved in

300 g of ethanol (Si/Ca/P 80:15:5, molar ratio) and stirred at room temperature for

1 day. After the composite powders were completely dry, they were milled

(8000M Mixer/Mill) to a fine powder. Finally, the powders were sintered at 700oC

for 5 hours.

3.1.1.3 Preparation of HA/bioglass composite material

HA/bioglass composite powders containing 10, 15 and 20 wt% glass, were

prepared by wet ball-milling for 1 hour. The resulting paste was then dried at

100oC for 24 hours, then the calcined ceramic pieces were milled using a ball mill

under a wet condition for 2 hours. Finally, polyvinyl alcohol and 3 drops of a

dispersant were added to produce the final slurry. After the ceramic was

completely dry, it was slowly sintered to the controlled temperature for 2 hours

and furnace cooled. The sintering temperature and bioglass additions were

detailed in Table 5. The powder was then examined by X-ray diffraction (XRD),

36

it was found that 20% bioglass additions at the 1400oC for 2h, HA/bioglass were

completely transferred to calcium phosphate, mainly -TCP, little -TCP. And

these results will be observed and discussed in Chapter 4.

Table 5 Sintering temperature and the bioglass additions

3.1.2 Fabrication of porous TCP scaffolds

3.1.2.1 Fabrication of sacrificial moulds

In this study, a scaffold template with macro-tubes was prepared using a 3D

printer which can produce a 3D structure with designed architectures. The

commonly used material for the 3D printing was acrylontrile butadiene styrene

(ABS). Fig.5 shows the scaffold of different sizes and shapes prepared by 3D

printer. Table 6 summarizes all details of the scaffolds template design. The ABS

scaffolds were cooled in the water after removal from the 3D printer, and then

dried for 24 hours.

37

Figure 5 ABS macrotube scaffold templates (a) and schematics of in-plane (top

view of black box in (a)) (b) and out-of-plane (view of cross section AA)

geometry (c)

38

Table 6 Parameters of the cube and cylinder ABS scaffold templates

Shapes of the scaffold Top

area(mm2)

rods

diameter(mm

)

Spacing

between two

rods(mm)

Height(m

m)

Cube 676 1 1.5 15

169 1 1.5 15

56.25 1 1.5 15

Cylinder 1 1.5 15

1 1.5 15

3.1.2.2 Coating ABS scaffold with wax

In this project, wax was used to take precautions against ceramic scaffold cracking.

Basically, during the sintering process, the wax was burned away and provided

enough space for ABS expansion. The wax (Sigma–Aldrich, AUS) was prepared

for the ABS scaffold. The wax was heated in an oven at 80oC for 10 minutes.

After the wax was fully melted, the ABS scaffold was dip coated with liquid wax.

This process should be completed quickly since wax can easily become solidified

in room temperature, and it may block the gaps. After coating, the scaffold was

kept at room temperature for at least 12 hours.

3.1.2.3 Sintering of the ceramic-coated scaffold

During the sintering process there were several parameters involved that need to

be optimized: the concentration of slurry, the maximum temperature, and the

dwelling time at the maximum temperature. The heating rate can also affect the

final density and phase purity of the material. Higher ramp rates can give a higher

final density, but a rate greater than 10oC/min has been shown to result in

decomposition [198]. In this step, two heating rates were used for sintering the

39

scaffold. The sintering was carried out in an electric tube furnace, using a heating

rate of 1oC/min up to sintering temperature, which was held for 3 hours, followed

by cooling rate 5oC/min to room temperature. The reason for setting 1

oC/min as

the heating rate was that the ABS would expand during the melting process, and if

the heating rate was too high, more cracks may be created. There were five stages

in the schedule (Fig.6), including:

(1) Heating from room temperature to 700oC with a heating rate 1

oC/min to

prevent the mould from cracking (the ABS and other organic additives will be

burned out at this step);

(2) Holding this temperature (700oC) for 3hours;

(3) Increasing the temperature from 700oC to 1400

oC at a rate of 5

oC/min;

(4) Holding this temperature for 3 hours;

(5) Cooling the furnace down to room temperature at a rate of 5oC/min. The

HA/Bioglass scaffolds were then removed from the furnace after it had cooled

down.

Figure 6 Sintering conditions

Time (hrs)

1

2

3

4

5

40

3.1.2.4 Coating the TCP scaffolds with normal bioglass

The sintered TCP scaffold was immersed into the normal bioglass slurry, and then

the vacuum pump was used to allow full infiltration. The soaked scaffolds were

then placed in a centrifuge running at 700rpm for 30sec to remove the excess

bioglass solution from the scaffolds. The scaffold was then removed from the

centrifuge tube and left to dry in a fume hood overnight. After the scaffolds were

completely dry, they were sintered at 700oC for 5hours to obtain the normal

bioglass coated TCP scaffold.

3.1.2.5 Coating TCP scaffold with sol-gel mesoporous bioactive glass

To prepare mesoporous bioglass scaffold required P123 (EO20–PO70–EO20) is

required to create nano pores [197]. Thus, 20g of P123, 33.5g of tetraethyl

orthosilicate (TEOS, 98%), 7g of Ca(NO3)2.4H2O, 3.65g of triethyl phosphate

(TEP, 99.8%) and 5 g of 37% HCl were dissolved in 300 g of ethanol (Si/Ca/P

80:15:5, molar ratio) and stirred at room temperature for 1day and the sol-gel

bioglass solution was prepared. Then the prepared TCP scaffold was put into this

sol-gel and centrifuged at 700rpm for 30 sec. After the scaffolds were completely

dry and sintered at 700oC for 5hours, the mesoporous bioglass coated TCP

scaffold was obtained.

3.2 Sample characterization

3.2.1 XRD (X-ray diffraction) analysis

XRD was used to identify the crystallographic phases of the reaction products

such as the TCP powder. For the XRD analysis, the samples were ground into fine

powders and powder was mounted in a specimen holder for the diffractometer

41

(6000 Shimadzu). Cu Kα1 ray (λ=1.5406 Å) scanning was conducted using a 2θ

angle of from 20° to 45°. The scan rate and the step size were 2.0° min−1

and 0.02°,

respectively.

3.2.2 Total porosity of the calcium phosphate ceramic scaffold

The pores in the calcium phosphate ceramic scaffold which were made by the HA

and the macro-tubes in the scaffold were responsible for the total porosity. In this

study, the total porosity of the sintered calcium phosphate ceramic scaffolds was

determined using the following equations [153]:

Total Porosity =1- M/ (V x ) (10)

Where

M= mass of the sintered sample

V= volume of the sintered sample

= Density of the sintered ceramic is 3.07 g/cm3

Specimen bulk density () was measured by application of Archimedes' principle.

The dimensions and the weight of each sample were measured and recorded using

a vernier calliper and an electronic balance.

3.2.3 Shrinkage

The shrinkage was determined using the following equations

Shrinkage=(O-F)/O (11)

O= the original length of the scaffold

F= the final length of the scaffold

3.2.4 Imaging the scaffold using SEM

Scanning electron microscopy (SEM) was used to determine the pore size

42

distribution in the calcium phosphate ceramic macro-tubes. The scaffold was

sectioned with a knife in the longitudinal and transverse planes to show the best

overview of the porous structure to verify the pore homogeneity and

interconnectivity. The scaffold samples were mounted onto aluminium stubs with

carbon tape and coated with gold using a sputter Coater (BioRad SC500) and

imaged using a FEI QUANTA 200 scanning microscope.

3.2.5 Mechanical testing for scaffolds

The compressive strength was main parameter for the mechanical testing. An

Instron 3000 mechanical tester with 10kN load cells was used for the compression

mechanical tests. The crosshead speed was set at 0.5mm/min, and the load was

applied until the scaffold was crushed completely. Rubber pads (1mm thick) were

placed on the top and bottom surfaces of the sample to ensure an evenly

distributed load on the sample. Five samples of each type were tested for

mechanical properties.

3.2.6 Apatite-formation ability of microspheres in SBF

The SBF (simulated body fluids) solution was prepared according to the

procedure described by Kokubo [199] and Table 7 shows the ion concentrations

of the SBF solution and human blood plasma.

43

Table 7 Ion concentrations of SBF and human blood plasma

TCP scaffold was soaked in SBF at 37ºC for 21 days with refreshing SBF every

week, and the ratio of the sample weight to the SBF volume (mg/ml) as 3:5. After

soaking, samples were removed from the SBF, gently washed with deionized

water, dried at room temperature, and characterized by SEM.

3.2.7 Cell seeding and culture

Human bone marrow was sourced from patients in Orthopaedic Department of

Prince Charles Hospital with informed consent and ethics approval from the

Ethics Committee of Queensland University of Technology. The human bone

marrow stromal cells (hBMSCs) for this study were isolated by density gradient

centrifugation over Lymphoprep (Axis-shield PoC AS, Oslo, Norway) according

to the manufacturer‘s protocol. The hBMSCs were seeded in culture flasks in

Dulbecco‘s Modified Eagle Medium (DMEM; Invitrogen Pty Ltd., Mt Waverley,

VIC, Australia) containing 10% fetal calf serum (FCS; InVitro Technology, Noble

Park, VIC, Australia) and 1% penicillin/streptomycin (Invitrogen) at 37°C, 5%

CO2. The medium was changed twice weekly to wash out all non-adherent cells.

After the cells reached 80% confluence, the cells were trypsinized and

44

re-suspended in DMEM. 1×105 cells were seeded onto each scaffold and cultured

for one week in DMEM at 37°C, 5% CO2.

3.2.8 Observation of cell attachment on scaffolds using SEM

The cell culture medium in each well was pipetted out, and immediately replaced

with phosphate buffered saline (PBS). The rinsing was repeated three times for

each sample, and then the scaffolds were fixed with a 3% glutaraldehyde solution.

The scaffolds were then processed by two changes of cacodylate buffer for

20mins each; and then soaking them in an osmium tetroxide solution for 1 hour,

finally they were dehydrated through a series of ethanol solutions with graded

concentrations, followed by two changes of 100% amyl acetate for 15mins each.

The scaffolds were then dried using a supercritical point dryer before observation

using SEM.

3.2.9 Cytotoxicity test by MTT assay

To evaluate hBMSCs proliferation with the existence of different materials,

hBMSCs were seeded at a density of 1×104

cells/well into 24-well plate and

incubated for 4 hours. 20mg of TCP was added to the culture plate. Cells were

then incubated at 37ºC in 5% CO2 for 1, 3 and 6 days. Then, 40μL of 0.5 mg/ml

MTT solution (Sigma, Aldrich) was added in each well and incubated for 4 hours

at 37ºC. The reaction was terminated by the addition of 100μL dimethyl sulfoxide.

The absorbance of the formazan was read at 495 nm using an Enzyme-linked

immunosorbent assay (ELISA) plate reader (Bio-Rad Laboratories, Pty., Ltd,

Gladesville, New South Wales, Australia). The MTT assay was to assess cell

viability and growth based upon the conversion of MTT to formazan. Results

45

were expressed as absorbance readings from each well. For the control, hBMSCs

proliferation in normal cell culture media was evaluated by the same procedure.

3.2.10 Alkaline phosphatase activity

Hunman bone marrow mesenchymal stem cell system (hBMSCs) were

respectively seeding into 6-well cell culture plates and TCP scaffolds for the

induction of hBMSCS at a number of 1×104 cells in the osteogenic medium

(DMEM supplemented with 10% FBS, 100nM dexamethasone, 50 mg/ml

ascorbic acid, and 10mM b-glycerophosphate). Cells were cultured in osteogenic

differentiation media for 7, 14 days. Then alkaline phosphatase (Alp) activity was

determined using pNPP assay (p-nitrophenyl phosphate liquid substrate,

Sigma-Aldrich). Briefly, hBMSCs were washed with PBS, then were lysed in

0.5ml PBS containing 0.1M glycine, 1 mM MgCl2 and 0.05% Triton X-100 for

10min at 4ºC. The lysate was incubated with pnitrophenylphosphate (pNPP)

solution at 37oC for 30mins, and then subjected to a spectrophotometer on which

the absorbance at 405 nm was measured and recorded to indicate Alp activity

[200].

3.2.11 Statistical analysis

Three completely independent experiments for cell culture and induction were

performed for every assay and the results were expressed as means ± standard

deviations. Statistical significance was calculated using one-way analysis of

variance (one-way ANOVA). Comparison between the two means was performed

using Turkey test and the significance was determined by p<0.05

46

3.3 Summary

In this chapter, the materials and methodology used were discussed. HA and

bioglass were used as the raw material to fabricate the TCP and SFFT was used

for fabricate the sacrificial scaffold template from ABS. TCP scaffold was

obtained after sintered HA/Bioglass composite at 1400oC for 3 hours. In order to

further increase the scaffold‘s mechanical property, the porous TCP scaffolds

were then coated with bioglass. Several biological tests were done. Sample

characterizations, such as XRD analysis, total porosity of TCP scaffold, imaging

of the scaffold by using scanning electron microscopy, mechanical testing and

biological testing (including SBF, intro cell culture using hBMSCs , MTT test and

Alkaline phosphatase activity) were also completed.

47

Chapter 4 Mechanical and biological

testing

4.0 Introduction

This chapter presents the results of the mechanical and biological tests and

discusses these in light of the suitability of scaffolds for application in bone tissue

engineering. All porosity and compressive strength for each HA/bioglass ceramic

scaffold was shown in Table 9.

4.1 Confirmation of TCP formation

In all cases, reaction occurred between the HA and the bioglass addition, resulting

in a reduction of HA content, and the formation of additional phases such as

-TCP and -TCP. The XRD results were summarized in Table 8.

Table 8 XRD phase analysis data for composites of HA with the addition of 10,

15, and 20wt% bioglass

Bioglass(wt%) Tsinter(oC) HA (wt%) TCP (wt%)

10 1200 53.3 29.1

1300 42.6 23.2

1400 26.2 33.5

15 1200 37.3 50.6

1300 20.6 63.2

1400 N.P* 87.2

48

20 1200 26.7 60.1

1300 12.4 75.3

1400 N.P* 100

*N.P not present

The data presented in Table 8 shows that at the same content of bioglass, with

increasing temperature, the content of the HA was decreased, and the weight

percentage of TCP was increased; at the same temperature, the weight percentage

of TCP was increased when more bioglass was added to HA. When 20% bioglass

was added to HA sintered at 1400oC, there was no HA phase present, which

meant that the HA and bioglass were totally transformed to TCP crystalline phase.

The phases formed in the scaffolds during the sintering process depended on two

factors, i.e. the sintering time and the sintering temperature. When the temperature

was above 1200°C, HA became unstable and could potentially eliminate OH

groups to form decomposed products of additional phases such as -TCP and

β-TCP. When the temperature was above 1400oC, -TCP phase was formed.

After the temperature was decreased, most of the -TCP was transformed to

β-TCP [201-202]. Miao et. al., suggested that some -TCP phase could be

retained due to the elastic strain constraint from the surrounding matrix [203]. In

addition, according to the HA decomposition equations shown above, there were a

number of undesirable phases detected during the sintering process. It was then

decided to add bioglass in order to adjust the Ca/P ratio, thus ensuring

HA/bioglass would be completely transformed into TCP. As a result, the phases

achieved in the sample should contain both α-TCP and β-TCP (mainly -TCP).

This was confirmed by the XRD analysis (Fig 7a). The XRD patterns in Fig. 7a

49

mostly matched with the pure β-TCP patterns in Fig 7b, thus, it was believed that

the TCP scaffold made via HA/bioglass was successfully obtained.

Figure 7 (a) XRD pattern for a sintered porous calcium phosphate (mainly-TCP)

sample; (b) XRD patterns of β-TCP starting powder and sintered body [204]

4.2 Shrinkage

Porous structures were characterized for their physical, mechanical and biological

properties to understand the influence of porosity parameters. Shrinkage was

(o)

(a)

(b)

50

measured as a part of physical property measurement to determine the sintering

behavior of these ceramics. None uniform shrinkage tends to cause cracking and

warping in the final products. Moreover, controlled shrinkage information can be

utilized in the design stage to make necessary design modifications. Shrinkage

during sintering is dependent on sintering temperature and the densification.

Shrinkage also varies as a function of volume fraction porosity in these samples in

which higher volume fraction porosity showed a lower amount of shrinkage [205].

For the porous TCP samples, linear shrinkage varied from 27% to 30%, and the

shrinkage variations were primarily due to the variations in the temperature. One

study indicated that, in a HA specimen, shrinkage began at about 800oC, from the

initially expanded state, and a discontinuity was observed at 1160oC [206]. Linear

shrinkage in percentage with respect to sintering temperature was shown in Fig.8.

The results indicated that TCP sintered at 1400oC exhibited highest shrinkage.

Figure 8 Linear shrinkage of TCP with respect to sintering temperature

25

26

27

28

29

30

12001300

1400

Lin

ear

sh

rin

kage

(%)

Temperature(oC)

51

4.3 Microstructure of ceramic composites

4.3.1 10% bioglass addition

Fig. 8 shows SEM images of HA with 10% bioglass addition at different

temperatures. With increasing sintering temperature (from 1200oC (Fig. 9a) to

1400oC (Fig. 9c)), increased amounts of pores were formed in the scaffold,

resulting in increased porosity values (Fig. 9d) from 61.15±1.3% to 70.23±0.8%.

The possible reason for this was that the higher temperature and higher shrinkage

caused the formation of some micro-cracks and pores, as well as large grains.

52

Figure 9 SEM images of the fracture surfaces of HA with 10% bioglass addition

at (a) 1200oC (b) 1300

oC (c) 1400

oC (d) The porosity of 10% bioglass addition

ceramic scaffold sintered at three temperatures. Error bars represent mean ± SD

for n=5.

60

62

64

66

68

70

72

1200 1300 1400

Po

rosi

ty(%

)

Temperature(oC)

(d)

53


The samples containing 15w% bioglass, sintered at 1200oC (a), 1300

oC (b),

1400oC(c) are shown in Fig. 10, and the pictures show a high degree of fine

interconnecting micro-porosity.

54

Figure 10 SEM images of the fracture surfaces of HA with 15% bioglass addition

at (a) 1200oC (b) 1300

oC (c) 1400



for n=5

Fig. 10d indicates that higher sintering temperature results in increased porosity.

As mentioned before, the HA and bioglass would react at high temperature, and

some other phases may formed. Consequently, in this condition, beside HA and

60

63

66

69

72

75

1200 1300 1400

Po

rosi

ty(%

)

Temperature(oC)

(d)

55

TCP, some other phases existed in the scaffold and influenced its structure and

porosity. It could also be noted that, at the lowest sintering temperature of 1200oC,

the structure was unusual, as some crystal-like features were formed, and this was

because a locally dense glassy matrix was formed on the samples.


Fig. 11shows the change in fracture surface for the 20wt% bioglass series between

1200oC (Fig.11a), 1300

oC (Fig.11b) and 1400

oC (Fig.11c). Different from the

10wt% and 15wt% bioglass additions, in this condition, the porosity of the

samples was decreased with increasing temperature, and the amounts of pores

were reduced.

56

Figure 11 SEM images of the fracture surface of HA with 20% bioglass addition

at (a) 1200oC (b) 1300

oC (c) 1400



for n=5

At the sintering temperature of 1300oC, there was an evidence of structural

deterioration at the surface of the specimens, possibly due to localized melting of

the composite. Fracture surfaces for higher sintering temperatures were totally

different; the samples sintered at 1400oC showed a film covering the surface, and

connecting all the grains and this obviously decreased the amount of the pores as

well as the pore sizes.

62

64

66

68

70

1200 1300 1400

Po

rosi

ty(%

)(d)

Temperature (oC)

57

Table 9 Porosity and mechanical property for composites of HA with the addition

of 10, 15, 20wt% bioglass at three temperatures

Bioglass(wt%) Tsinter(oC) Porosity (%)

Compressive

strength (MPa)

10 1200 61.15±1.3% 2.7±0.3

1300 63.05±1.2% 2.15±0.4

1400 70.23±0.8% 1.83±0.3

15 1200 60.63±0.4 2.97±0.5

1300 68.47±1.23% 1.75±0.2

1400 72.64±1.2% 1.69±0.2

20 1200 67.08±0.4% 1.92±0.8

1300 64.42±0.6% 1.87±1.1

1400 62.99±1% 9.98±0.6

58

4.4 Mechanical properties


Variations of the compressive strength depending on sintering temperature are

shown in Fig.12. The series containing 10 wt% bioglass sintered at 1200oC,

1300oC, 1400

oC produced cracks on the specimens upon sintering. The pore size

was greater at higher temperature, and the cracking was more severe at the higher

temperature, thus the compressive strength was decreased from 2.7±0.3MPa at

1200oC to 1.83±0.3MPa at 1400

oC.


sintering temperatures. Error bars represent mean ± Standard deviation(SD) for

n=5

0

0.5

1

1.5

2

2.5

3

1200 1300 1400

Co

mp

ress

ive

St

ren

gth

(MP

a)

Temperature(oC)

59


Fig.13 indicates that the higher the sintering temperature, the lower the

compressive strength. The values were decreased from 2.97±0.5MPa (1200oC) to

1.69±0.2MPa (1400oC). In the graph, the compressive strength between 1300

oC

and 1400oC were similar, but much higher compressive strength was shown at

1200oC. Compared to the previous study with 10% bioglass addition, the trends

were the same i.e. with the increase of the temperature, the compressive strength

is decreased.


sintering temperatures. Error bars represent mean ± SD for n=5


Compression test results of 20wt% bioglass composites at 1400oC were

substantially higher strength values compared to those of the 1200oC and 1300

oC

sintered bodies (Fig.14). While strengths have been observed at 9.98±0.6MPa for

0

1

2

3

1200 1300 1400

Co

mp

ress

ive

St

ren

gth

(MP

a)

Temperature(oC)

60

20wt% bioglass addition sintered at 1400oC, there were interesting values of

1.92±0.8MPa at 1200oC, and 1.87±1.1MPa at 1300

oC as can be seen in Table 9.

These materials had strengths that were influenced by the sintering temperature,

with the exception in the trend compared to the other two bioglass additions.


sintering temperatures. Error bars represent mean ± SD for n=5

From Fig.12-14, it could be easily found that at 10% and 15% bioglass additions,

the values of compressive strength were decreased with increased temperature.

However, 20% bioglass addition was different from the others, i.e. the higher

temperature, the higher compressive strength. At 1400oC the values reached the

highest which was 9.98MPa and compressive modulus was 1.24GPa (higher than

the Trabecular bone‘s compressive modulus), it was because all the HA and

bioglass were transformed to TCP.

0

2

4

6

8

10

1200 1300 1400

Co

mp

ress

ive

St

ren

gth

(MP

a)

Temperature(oC)

61

4.5 Porosity and compressive strength of the uncoated TCP scaffold

The total porosity for uncoated TCP scaffolds was approximately 62.99%±1.9%,

because the macro-tubes were mainly responsible for the porosity. Under the same

condition (i.e. temperature), with increase of the ceramics‘ porosity, the

compressive strength would decrease. For the same template ABS scaffold, the

scaffold made from the TCP powder did not provide sufficient mechanical

stability due to higher porosity.

Figure 15 SEM images of the fracture surfaces of the porous TCP showing the

presence of micro-pores: (a) made via sintering HA/bioglass and (b) made via

TCP powder [207]

62

The pore morphology and distribution of macroporous TCP scaffold can be

visualized by the SEM photomicrographs presented in Fig. 15. It could be

observed that some interconnected macropores ranging from 5m to50m were

present in the 3D hybrid scaffold, and these pores may be produced by foaming a

mixture of PVA aqueous solution and water vapor. Compared with the hybrid

scaffold, the TCP powder scaffold (Fig.15 b) had more pores and the pore size

distrobution between 15m-600m [28]. Such a difference was because the

melting point for the bioglass was 700oC, and when the temperature reached

700oC, the bioglass transferred from the solid to the fluid, and filled some pores in

the scaffold, i.e. the bioglass acted binder, connecting all particles and resulting in

increasing of the scaffold‘s compressive strength.

4.6 Porosity of the coated scaffold (normal bioglass)

The sintered porous TCP scaffold mechanical properties nearly match those of

real bone but the still not strong enough, and this was due to the high porosity and

the intrinsically low mechanical strength of the calcium phosphates. One way to

strengthen and toughen the porous TCP was to coat the porous struts with a

bioactive ceramic. Thus, bioglass was used again in the project, and there were

two kinds of bioglass used (i.e. normal bioglass and sol-gel mesoporous bioglass).

63

Figure 16 SEM micrographs of a porous TCP scaffold with a normal bioglass at

layer 42X (a), 160X (b) and 3000X (c), as well as the surface of the normal

bioglass layer (d)

Normal bioglass slurry was used to infiltrate and coat the sintered porous TCP, as

shown in Fig.16 and the porosity of the TCP scaffold which coated with normal

bioglass was 59%±3%, a little lower than the TCP scaffold (63%). The

micro-pores in the coated scaffolds (Fig.16b) were less than in uncoated scaffold

(Fig.15a). Also it was observed that the TCP scaffolds had more crack-like defects

on and within the ceramic struts than those coated with bioglass.

Several factors affected the coating process, namely, the viscosity of the slurry,

64

and the particle sizes of the bioactive glass. The coating quality was also

controlled by the details involved in the dipping coating process. Some bioglass

had become a layer covering the struts. Some bioactive glass particles had stayed

on the surface of the bioglass layer (Fig.16d).

4.7 Porosity of the coated scaffold (mesoporous bioglass)

The total porosity of mesoporous bioglass coated scaffold was 61%±2%. The

macroporosity of the scaffolds after infiltration and coating with the bioglass was

slightly decreased due to the observed thin layer coating present on the strut

surfaces.

65

Figure 17 SEM micrographs of a porous TCP scaffold with a mesoporous

bioglass layer 52X (a), and 200X (b) (a and b were single coatings); 1600X (c)

1600X(d) (c and d were double coatings)

In a similar study [208], Jun et al. had shown an about 2% reduction in porosity as

a result of the bioglass coating on the ceramic scaffolds. Different from the

normal bioglass, the mesoporous bioglass needed to be coated first, and then

sintered. Fig. 17a shows that TCP scaffold was coated with mesoporous bioglass

and it was found that the mesoporous bioglass still did not cover the entire

scaffold; the reason for this was the viscosity of the meso-bioglass as well as the

sintering temperature. This problem was resolved in this study by a secondary

coating as well as sintering process, Fig.17c shows that bioglass was laid over the

scaffold. After the initial coating process, the TCP sample was coated again

(double coating) and had gone through the sintering process. It was discovered

that each piece of bioglass had been connected with each other. Fig.17d shows the

surface of the meso-porous bioglass, it was observed that some nano-size pores

were present in the meso-porous bioglass. The significance of the mesoporous

66

bioglass was that the nano-size pores imparted the scaffold with better

biocompatibility [217].

4.8 Mechanical testing

4.8.1 Shape and size effects

The sizes and shapes of porous scaffolds could influence their mechanical

strength, permeability and the presence of structure defects. Table 10 shows the

mechanical strength and porosity of TCP scaffolds prepared with different ABS

templates (Error bars represent mean ± SD for n=5). In Table 10, it was found that

that scaffolds with similar sizes (i.e. length and height), cylindrical structure have

higher mechanical strength than cube ones. In a recent study, Alsayed reported

that for comparison between compressive cube strength and compressive cylinder

strength, cube strength= 0.8*cylinder strength [220]. In this project, results

showed that compressive strength for cube scaffold was 9.98MPa (mid-size), and

for cylinder strength was 12.13MPa. The compressive strength of the cubic

scaffold was approximately 80% of the cylindrical scaffold compressive strength,

which was very similar to the results obtained from the literature.

In these kinds of scaffolds, macro-tubes were mainly responsible for the porosity,

and the average porosity of the scaffolds ranged from 40%±0.5% to 72%±1.3%.

The larger surface area, the higher scaffold porosity, and thus both favour cell

adhesion to the scaffold and promote bone tissue regeneration. As the surface area

was increasing, more cracks and pores were observed. This would be mainly

attributed to the difference in thermal expansion coefficient between the ABS

67

template and the TCP. As a result, the large scaffold exhibits an increasingly

brittle response towards the end. For a decrease of surface area from 676 to

56.25mm2, the compressive strength increased from 1.52±0.65 to 14.25±0.3MPa.

For the purpose of comparison, in this work the scaffold was also made via TCP

powder directly by using 169mm2 top area ABS templates (cube). A higher

porosity was observed in these scaffolds (66.21±1.20%) resulting in a lower

mechanical strength (2.8±1.3MPa) (Fig.18a) and the compressive modulus for this

scaffold was 0.021GPa.

Table 10 Average porosity and compressive strength for the scaffolds

Shapes of the

scaffold

Top area

(mm2)

Average

porosity (%)

Compressive

Strength(MPa)

(HA/bioglass)

Compressive

Strength(MPa)

(TCP powder)

Cube 676 72.40±1.3 1.52±0.65

169 63.34±1.5 9.98±1.40 (Fig.17b) 2.8±1.3MPa

(Fig.17a)

56.25 40.65±0.5 14.25±0.3

Cylinder* 169 64.26±0.7 12.13±1.2

56.25 41.13±0.8 17.68±0.2

*In this project, all scaffolds were fabricated as cube and the top area was 169mm2, the

cylinder and other sizes of scaffolds just for comparison in this chapter.

The difference in morphology between the scaffolds made via sintering

HA/bioglass and those via TCP powder was shown in Fig.15. The micropores in

HA/bioglass TCP scaffold were less than those in TCP powder scaffold, which

was because the bioglass was melted at 700oC, and the melted bioglass could fill

up some pores and cracks. In addition, when the temperature was increased to

68

1400oC, the bioglass would react with HA and form a continuous layer of TCP on

the surface of the scaffold, resulting in increase of the compressive strength.

0

0.5

1

1.5

2

2.5

3

0 20 40 60 80

Co

mp

ress

ive

Str

en

gth

(Mp

a)

Compressive Strain(%)

0

2

4

6

8

10

12

0 20 40

Co

mp

ress

ive

Str

en

gth

(Mp

a)

Compression strain(%)

(a)

(b)

69

Figure 18 Compression stress–strain curves of the sintered porous scaffold

samples (a) made via pure TCP powder, (b) made via HA/Bioglass, (c) TCP made

via HA/Bioglass and coated with two different bioglasses. (Error bars represent

mean ± SD for n=5)

According to data from Table 10, for the cubic shape of TCP scaffold samples, in

terms of the average porosity, the scaffold sample with the top area of 676 mm2

was higher than the scaffold sample with the top area of 56.25mm2. However, the

compressive strength had decreased significantly as the top area of the scaffold

sample increased. The key objective was to achieve a harmonious balance

between the average porosity and the compressive strength. Based on this theory,

it was believed that the scaffold sample with the top area of 169mm2 was the best

result achieved among the three sample sizes.

0

3

6

9

12

15

18

0 10 20 30 40

Co

mp

ress

ive

Str

en

gth

(Mp

a)

Compression strain(%)

Normal BG

Meso-porous BG

(c)

70

4.8.2 Mechanical properties for bioglass coated scaffold

Fig.18b shows a stress-strain curve of the scaffold sample made via sintering of

HA/Bioglass. A higher compressive strength, i.e., 9.98±1.4MPa was observed. On

the other hand, Fig.18c observes a stress-strain curve of the sintered porous

HA/Bioglass interpenetrating composite that was infiltrated with the bioglass.

Two different kinds of bioglasses were used for enhancing scaffolds‘ mechanical

and biological properties. The compressive strength for normal bioglass coated

scaffold was 16.69±0.5MPa, and for single mesoporous bioglass coated scaffold

was 15.03±0.63MPa, the mechanical data were comparable to those of spongy

bones, which show a compressive strength of 2-20MPa [209-210], respectively.

Thus it could be said that scaffold after bioglass coating, the compressive strength

was increased about 40%, which was almost identical to results found in Jun et

al‘s report [24]. According to Jun et al‘s report [24], HA scaffold was coated with

apatite-wollastonite glass (i.e. a type of bioglass), and compressive strength was

increased from 0.58MPa to 0.95MPa, also nearly increased 40%. When the TCP

scaffold for the secondary coating, its porosity decreased, and its compressive

strength was measured to be about 22MPa. The strength of the scaffold was

remarkably increased by coating scaffold for twice instead of the single coating.

This enhanced compressive strength was attributed not only to the elimination of

the defects present on the scaffold but also to reinforcement the scaffold by using

a strong glass ceramic phase as the framework (each piece of bioglass had been

connected with each other(Fig.17c)).

71

4.9 Apatite-formation ability of the macro-tube scaffolds

in SBF

SEM images of the scaffold after soaking in SBF solution for 21 days were shown

in Fig.19a. At a high magnification, it was noted that a rough deposit layer was

formed on the TCP scaffold, and some crystal clusters were also found on the

layer surface. To determine the chemical composition of the crystal deposits, the

surfaces of the scaffold were further characterized by EDS (Energy dispersive

spectroscopy). The EDS spectra of the TCP scaffold after incubation in SBF

solution for 21 days were shown in Fig. 19b.

72

Figure 19 (a) SEM image of the scaffold after soaking in SBF for 21 days (b)

EDS spectra of the surfaces of TCP

The HA granules showed high peaks of P, Ca, and the atom ratio of Ca to P was

about 1.65, which was close to that of carbonated apatite. These characteristics

could be found in HA. Since HA was capable of stimulating the osteo-induction

of stem cells in vivo with some ability of osteo-conductivity [211], it was believed

that these TCP scaffolds were highly bioactive.

4.10 Cell culture

4.10.1 SEM images of cells growth on uncoated scaffolds

Bone marrow stromal stem cells were seeded into the scaffolds by adding drops of

the cell suspension. As a result, the cells were homogeneously seeded across the

surface of the scaffolds. The penetration of the cells into the scaffolds was

evaluated using the cross-sections of the scaffolds. SEM imaging (Fig. 20) has

revealed that the cells spread and adhere well on the surface. Interestingly, it was

73

noted that a significant amount of HA granules were formed on the scaffold

surfaces and distributed among the cells, which was advantageous as far as

bioactivity was concerned. In general, HA granules can be formed in a simulated

body fluid (SBF) solution when a biomaterial has good bioactivity (Fig. 20a).

However, in this experiment, the HA granules were also formed when the TCP

scaffold was immersed into the DMEM solution. There were two possible reasons.

The first one was the solution contained bovine serum, which possessed some

ions; at the same time some P and Ca ions could be slowly released into the

solution from TCP, thus the solution in the medium was similar to a SBF solution.

Naturally, some HA granules were formed on the surface of the scaffold after 9

days (Fig. 20a). Fig.21 shows that the formation of HA granules (confirm by EDS)

when living cells were not present, indicating the formation of HA was

independent of the cell environment.

74

Figure 20 SEM micrographs show the attachment of the cells on TCP ceramic

strut surface 800X (a) and 2000X (b)

Figure 21 SEM image of TCP ceramic scaffold immersed into DMEM without

cells

75

4.10.2 SEM images of cell growth on bioglass coated scaffolds

Figure 22 SEM micrographs show the attachment of the cells on bioglass coated

TCP ceramic strut surface 2000X (a) and 2500X (b)

The bone marrow stromal stem cells were well spread on the surface of bioglass

coated TCP ceramic. This coated scaffold was different from the non-coated

scaffold, because no HA granules were found in the coated scaffold as shown in

Fig. 22. This occurs because more bioglass in the medium means more Ca and P

irons found in the medium, thus having negative impact on the balance in SBF.

76

4.10.3 Cell proliferation

MTT assay is an important method to evaluate the cytotoxicity of the scaffold and

its slowly released components in an aqueous environment. Fig. 23 shows the

result of MTT assay for proliferation status of hBMSCs as the negative control

group and as the experimental group of hBMSCs combining with TCP scaffold

during 9 days after the same culture condition.

Figure 23 MTT assay for proliferation of hBMSCs and hBMSCs combined with

TCP scaffolds at different incubation periods under the same culture condition.

Error bars represent mean ±SD for n=3

Cell proliferation increases with the culture time in both groups, but the cell

growth rate of the experimental group was much higher the negative control

groups at 1, 3 and 6 days respectively. At 3, 6 and 9 days, the cell proliferation

0

0.3

0.6

0.9

1.2

1 4

Ab

sorb

ance

at

49

5n

m

Time/(days)

BMSCs

BMSCs in TCP scaffold

93 6

77

between the negative control group and the experimental group has no significant

difference. Thus, the TCP scaffold was proven non-cytotoxic and has good

biocompatibility in vitro.

4.10.4 Alp (Alkaline Phosphatase) activity

To study the bioactivity of the TCP scaffold, the Alp activity was measured. As

one of the cell membrane associated enzymes, Alp is known to be closely

associated with osteoblast differentiation. Alp regulates phosphate metabolism

and locally down-regulates inhibitors of apatite crystal growth [219]. Therefore, it

is used as a marker that appears early during osteoblast differentiation. A high Alp

activity was detected in positive TCP scaffold groups in both day 7 and 14 in

comparison to pure hBMSCs‘ differentiation in osteogenic medium (Fig. 24a).

0

0.5

1

1.5

2

2.5

3

3.5

4

7 days 14 days

Alp

Act

ivit

y(O

D/m

in/m

g p

rote

in)

Culture time (days)

BMSCs induction

BMSCs induction with TCP scaffold

(a)

78

Figure 24 Alp activity of hBMSCs after 7 days, 14 days of culture, and the novel

TCP scaffolds significantly increased the Alp activity to higher level compared to

positive control group after 14 days (a); Alp activity of hBMSCs after culturing

for 7 days on TCP scaffold and TCP scaffold coated with normal bioglass (b).

Error bars represent mean ± SD for n=3

Bioactivity is thought to be an important issue in the chemical interactions

between the implanting materials and the bone tissue, and ultimately affects the in

vivo success of the bone grafting materials [212]. In this study, compared to the

positive differentiation group, the new scaffold has an enhanced effect on Alp

activity of hBMSCs after 14 days, which can be attributed to the higher

concentration of Ca and Si ions that were slowly released from bioglass, and more

stable pH environment [213] bioglass led to in cell culture medium than the

general in vitro environment. The proved reasons were that bioglass incorporated

0

0.5

1

1.5

2

2.5

3

TCP scaffold TCP scaffold coated with normal bioglass

Alp

Act

ivit

y(O

D/m

in/m

g p

rote

in)

(b)

79

in the TCP scaffolds can release Ca and Si ions which can stimulate cell response

[214-215]. The differentiation characteristics of the cells were evaluated by

measuring the Alp activity of hBMSCs‘ after culturing 7 days. The cells on TCP

scaffold coated with bioglass expressed Alp activity at a significantly higher level

than those TCP scaffold, as shown in Fig.24b. The higher Alp activity of the cells

cultured on the scaffold coated with bioglass compared with that the cultured on

TCP scaffold strongly suggested that the bioglass coating on the TCP scaffold is

effective in improving the bioactivity of the porous scaffold.

4.11 Summary

A 3D macrotube porous TCP scaffold was successfully synthesized by the

sintering HA and 20% bioglass at 1400oC. The novel scaffold exhibits macro-tube

(0.8mm) and as well as a hierarchical structure with interconnected macropores.

The resultant scaffold exhibits good mechanical properties (i.e. 9.98MPa for

compressive strength and 1.24GPa for compressive modulus: cube, 169 mm2

top-area) that was significantly influenced by the bioglass melting at 1400oC and

the fluid bioglass covering on the surface of the scaffold. The mechanical

properties of the obtained TCP scaffold had values that were still not sufficient for

load-bearing applications, in order to further improve scaffold‘s mechanical and

biological properties, the TCP scaffold was further coated with bioglass, and

compressive strength was largely enhanced (more than 15MPa). These kinds of

scaffolds could be considered for the implantation in bone defects for the bone

formation. Also, the cells were well spread in the scaffold, and the MTT test

showed that there was no toxicity in this material. Alp activities tests indicated

80

that TCP scaffold has good bioactivity and improved bioactivity when the TCP

scaffold was coated with bioglass. This new class of material combines macrotube

scaffold possessing excellent physicochemical, biological properties, hence

indicating their potential application for bone tissue engineering. This scaffold

solves some existing problems in clinical grafts by providing strength and

controlled released bioactive molecules to prompt hBMSCs‘ osteogenesis.

81

Chapter 5 Conclusions and future work

5.1 Conclusions

The objective of this research was to develop a macro-tube porous tri-calcium

phosphate (TCP) scaffold through the use of composite HA (Hydroxyapatite) and

bioactive glass sintered at high temperatures. To achieve this objective, a number

of key steps were taken throughout the project. The first key step was to develop

and manufacture an acrylonitrile butadiene styrene (ABS) sacrificed scaffold

using solid freeform fabrication techniques (SFFT) as a template. SFFT has been

identified as an advanced technique for controlling porous scaffold morphology.

The second key step was to fabricate TCP scaffold by sintering 80% HA and 20%

bioglass at 1400oC for three hours. Following that, appropriate coating materials

for coating the TCP scaffold were selected. Poly (lactic-co-glycolic acid) (also

commonly known as PLGA) and calcium phosphate ceramics were initially

chosen. Unfortunately, it was discovered that the chemical used to dissolve PLGA

may possess potential health and safety risks. In addition, after applying polymer

coating, the scaffold surface will become smooth, which can discourage cell

growth. Therefore, calcium phosphate ceramics were the most favorable materials

to be used for both base and coating purposes. Cell culture was also conducted to

study the bioactivity of the TCP scaffold.

A number of conclusions can be drawn from this project:

The method of manufacturing TCP was explored throughout the project. TCP

82

was made by mixing two basic ingredients: HA and Bioglass together. By

varying the amount of HA and bioglass added into the mixture and altering

the sintering temperature, it was discovered that the most favorable method of

making TCP was when 20wt% bioglass was added into the mixture and the

sintering temperature was 1400oC.

In terms of the size and the shape of the scaffold and their associated effect

on both mechanical properties and porosity, it was discovered that as the top

area of the scaffolds increased, their mechanical properties were becoming

lower. The effect on the porosity was also studied. It was found that the

scaffolds with larger top areas had higher porosity. Taking both factors into

account, it was concluded that scaffolds with mid-size top areas were the best

candidate. This conclusion was supported by the relevant tests conducted

throughout the project. For instance, the compressive strength of the scaffold

with cubic cross-section area had a high reading of 9.98MPa, and high

compressive modulus reading of 1.24GPa. The porosity was measured to be

63%, which was also relatively high. The shape of the scaffold had also

impacted on the compressive strength of the scaffold. With identical size of

the top area, the cylindrical scaffolds had higher compressive strength then

the cubic scaffold.

In order to further improve the scaffold‘s mechanical properties, TCP

scaffold was coated with bioglass. Scaffold‘s compressive strength was

largely improved from 9.98MPa to above 15MPa. Two types of bioglass

(normal and single mesoporous bioglass) were used for enhancing the

83

mechanical properties of the TCP scaffold. It was discovered that the TCP

scaffold coated with normal bioglass had slight higher compressive strength

(16.69±0.5MPa) then the one coated with single mesoporous bioglass

(15.03±0.63MPa). These results were comparable with those of spongy bones.

Therefore, these scaffolds could be potentially used as implantation materials

for bone formation.

In terms of analyzing the bioactivity of the TCP scaffold, the first major study

was to evaluate the cytotoxicity of the TCP scaffold and its slowly released

components in an aqueous environment. By using MTT assay, the TCP

scaffold was proven to be non-cytotoxic. It would also have good

biocompatibility in vitro.

The second major study was to determine the bioactivity of the TCP scaffold.

This was based on the study of osteoblast differentiation in the TCP scaffold.

The Alp activity of human bone marrow mesenchymal stem cell systems

(hBMSCs) after culturing seven days and fourteen days were measured as

Alp is known to be closely associated with osteoblast differentiation. It was

discovered that hBMSCs induction with TCP scaffold had higher Alp activity

after seven days as well as after fourteen days. The higher Alp activity of the

cells cultured on the scaffold coated with bioglass compared with that the

cultured on TCP scaffold strongly suggested that bioglass coating on the TCP

scaffold was effective in improving the bioactivity of the porous scaffold.

84

5.2 Future work

It is recommended that the finding of this study be utilized to help the bone

formation. In future work, there are three aspects which can be improved. In terms

of fabrication method, as mentioned before, some cracks were created in the

scaffold due to the fact that ABS polymer expanded at high temperature, thus, the

ABS polymer could be instead replaced by other materials. In the mechanical part,

carbon nanotubes (CNT) could be considered as the fiber reinforcement within the

ceramic matrix for the scaffold fabricating. CNT have aroused increasing interest

due to their remarkable tensile strength, high resilience, flexibility and other

unique structural, mechanical, electrical and physicochemical properties [214,

215]. Attempts have been made to develop advanced engineering materials with

improved or novel properties through the incorporation of CNT in selected

matrices such as polymers, metals and ceramic. It is expected that the inclusion of

CNT in a ceramic matrix will produce composites possessing high stiffness and

improved mechanical properties compared with a single phase ceramic material.

According to the latest paper [216], the results showed that CoCl2 pre-treated

hBMSCs induced higher degree of vascularization and enhanced osteogenesis

within the implants in both ectopic and orthotopic areas, thus, in the future work

the CoCl2 could be added into the bioglass coating and the Co ions can be

released in the body help the blood and bone formation.

85

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