Inter-bonded Fibrous Matrices for 3D Tissue Engineering

Inter-bonded Fibrous Matrices for 3D Tissue Engineering Scaffolds

by

Yanwei TANG(B.Eng, M.Eng)

Submitted in fulfilment of the requirements for the degree of

Doctor of Philosophy

Deakin University

January, 2011

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ACKNOWLEDGEMENTS

I would like to sincerely thank my supervisor A/Prof. Tong Lin, who have been

encouraging me to pursue such an exciting project, and who always direct me

in the new techniques and direction.

I am also deeply grateful to Prof. Xungai Wang for giving me the chance to

study in Deakin with AMCRC scholarship. Without this financial support, this

work could not have been completed.

I would like to acknowledge Dr Cynthia Wong for her invaluable advice and

suggestions in cell culture experiments.

I should also thank Dr. Xin Liu, Mr. Chris Hurren, Mr. Graeme Keating, and

Mrs. Elizabeth Laidlaw for their technical assistance, and all the friends in

CMFI for giving me a great time during my study.

Thanks should also been given to Barwon Biomedical Research (BBR) and

School of Life and Environmental Sciences (LES) at Deakin University for

kindly allowing me to use the biological labs throughout the project.

I thank my family for their love and support over the years, especially to my

husband, Zhaohuai Lin, who has always been there for me.

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PUBLICATIONS LIST

The work described in this thesis has led to the following publications:

1. Yanwei Tang, Cynthia Wong, Hongxia Wang, Alessandra Sutti, Mark

Kirkland, Xungai Wang, Tong Lin. Three Dimensional Tissue Scaffolds from

Inter-bonded PCL Fibrous Matrices with Controlled Porosity. Tissue

Eningeering.Part C, Volume 17, Number 2, 2010

2. Yan Zhao, Yanwei Tang, Xungai Wang, Tong Lin. Superhydrophobic

cotton fabric fabricated by electrostatic assembly of silica nanoparticles and its

remarkable buoyancy. Applied Surface Science, 256,6736-6742, 2010

3. Yanwei Tang, Cynthia Wong, Alessandra Sutti, Tong Lin, Xungai Wang.

3D Fibrous Tissue Scaffolds with Controlled Pore Structure and Self-sterilising

Function. 20th Annual ASBTE Conference (10-12 Feb. 2010, Brisbane,

Queensland, Australia), 2010

4. Yanwei Tang, Alessandra Sutti, Tong Lin. 3D Fibrous Matrices Having

Different Porosities and their CHO Cell Culture Performance.Internatiaonl

Conference on Materials for Advanced Technologies ICMAT 2009(28 Jun-3

Jul 2009, Singapore), A01811-03187 (2009)

5. Yanwei Tang, Cynthia Wong, Hongxia Wang, Alessandra Sutti, Tong Lin,

Xungai Wang, Mark Kirkland. 3D Fibrous Tissue Scaffolds: Influence of

Pore Structrue on the Cell Culture Performance. Conference Booklet of

ARNAM 2008 Annual Workshop(15-18 Dec 2008, Geelong, Australia), 74

(2008)

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Potential papers from this research are as listed below:

6. Proliferation and Differentiation of hFOB 1.19 on Nano-textured Surface in

3D Fibrous Scaffold

7. Response of Osteoblast Cells to Apatite Coated PCL Fibrous Scaffold by

Plasma Pre-treatment and SBF Soaking for Bone Tissue Engineering

8. Electrostatic Assembly of Anti-bacterial PHMB on PCL Fibrous Scaffolds

for Self-disinfection

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TABLE OF CONTENTS

ACKNOWLEDGEMENTS ............................................................................. I

PUBLICATIONS LIST ................................................................................ II

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

LIST OF FIGURES .................................................................................... IX

LIST OF TABLES .................................................................................... XXV

ABSTRACT ............................................................................................. XXVI

1. Introduction .................................................................................................. 1

1.1 Significance and research problems ............................................................. 1

1.2 Specific aims ................................................................................................. 4

1.3 Thesis outline ................................................................................................ 6

2. Literature Review ......................................................................................... 8

2.1 Tissue Engineering and Scaffolds ................................................................ 8

2.2 Current Developments in 3D Polymer Scaffolds ....................................... 10

2.2.1 Cell lines .............................................................................................. 11

2.2.2 Cell seeding and culture methods ........................................................ 11

2.2.3 Materials .............................................................................................. 15

2.2.4 Scaffold fabrications and structures .................................................... 19

2.3 Fibre-based Tissue Scaffolds ...................................................................... 33

2.3.1 Classification of fibre-based scaffolds ................................................ 33

2.3.2 Issues of fibre-based scaffolds ............................................................. 40

2.4 Role of nano-structured functional surfaces on cell performance of polymer

scaffolds ............................................................................................................ 40

3. Materials, Methods and Characterisation Technologies ........................ 56

3.1. Materials .................................................................................................... 56

3.2. Fabrication and surface modification of fibrous scaffolds ........................ 56

Synthesis of silica nanoparticles ................................................................... 57

Layer-by-layer electrostatic self-assembly of nanoparticles ........................ 57

Plasma treatment of PCL scaffolds .............................................................. 58

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Simulated Body Solution (SBF) soaking ....................................................... 58

Electroless plating ........................................................................................ 60

Layer-by-layer coating of PHMB on 3D fibrous scaffolds ........................... 60

3.3. Cell culture related technologies ............................................................... 61

3.3.1 Sterilisation .......................................................................................... 61

3.3.2 Cell culture media................................................................................ 61

3.3.3 Cell lines .............................................................................................. 62

3.3.4 Subculture of cells................................................................................ 62

3.3.5 Cell seeding ......................................................................................... 63

3.3.6 Cell culture .......................................................................................... 64

3.3.7 MTS assay ............................................................................................ 65

3.3.8 Cytotoxicity test ................................................................................... 65

3.3.8 Stain of samples for microscopic observation ..................................... 66

3.3.9 Stain of samples for SEM observation ................................................. 66

3.3.10 Laser scanning confocal microscopy................................................. 66

3.3.11 Alkaline Phosphatase (ALP) assay .................................................... 67

3.3.12 Alkaline Phosphatase (ALP) staining for microscopy observation ... 68

3.3.13Statistical analysis .............................................................................. 68

3.4 Bacterium related technologies................................................................... 69

3.4.1 Bacterial type ....................................................................................... 69

3.4.2 Bacterial culture media ....................................................................... 69

3.4.3 Antibacterial testing............................................................................. 69

3.5 Characterisation tools ................................................................................. 72

3.5.1 Photos .................................................................................................. 72

3.5.2 Matrix thickness measurement ............................................................ 72

3.5.3 Scanning electron microscopy ............................................................. 73

3.5.4 Micro-computed tomography .............................................................. 73

3.5.5 Processing of the micro-CT scanned series images ............................ 74

3.5.6 Mechanical property ............................................................................ 74

3.5.7 Water contact angle ............................................................................. 75

3.5.8 Water hydraulic permeability .............................................................. 76

3.5.9 Water binding ability ........................................................................... 76

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3.5.10 Fourier transform infrared spectroscopy .......................................... 77

3.5.11 Atomic force microscopy ................................................................... 77

3.5.12 Dynamic light scattering ................................................................... 77

3.5.13 Transmission electron microscopy .................................................... 78

3.5.14 UV-VIS spectroscopy ......................................................................... 78

3.5.15 Wide angle X-ray diffraction ............................................................. 79

3.5.15 X-ray photoelectron spectrometer ..................................................... 79

3.5.16 Inductively coupled plasma-atomic emission spectrometer .............. 79

3.5.17 Photo-spectroscopy............................................................................ 79

4. Inter-bonded 3D fibrous scaffolds and their cell culture performances 81

4.1 Experimental Procedure.............................................................................. 81

4.2 Results and Discussion ............................................................................... 82

4.2.1 Length and diameter distribution of short fibres ................................. 82

4.2.2 Size and distribution of sugar powders ............................................... 83

4.2.3 Formation of 3D fibrous structures ..................................................... 84

4.2.4 Optimisation of processing parameters ............................................... 85

4.2.5 Fibrous samples with different fibre/sugar ratios ............................... 89

4.2.6 Cytotoxicity test ................................................................................. 100

4.2.7 Comparison of dynamic and static seeding methods ......................... 100

4.2.8 Cell seeding efficiency ....................................................................... 107

4.2.9 Cell morphology ................................................................................ 108

4.2.10 Cells penetration inside the matrices .............................................. 115

4.2.11 Cell proliferation ............................................................................. 117

4.3 Conclusion ................................................................................................ 121

5. Nano-structured surface in 3D fibrous scaffolds and its cell culture

performances ................................................................................................. 123

5.1 Experimental Procedure............................................................................ 123

5.2 Results and Discussion ............................................................................. 123

5.2.1 Silica particles and size distribution ................................................. 123

5.2.2 Nanoparticle assembly and surface morphology .............................. 125

5.2.3 Surface chemical components............................................................ 126

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5.2.4 Atomic force microscope ................................................................... 129

5.2.5 Surface wettability ............................................................................. 129

5.2.6 Cytotoxicity ........................................................................................ 130

5.2.7 Osteoblast cell culture ....................................................................... 132

5.2.8 Cell morphology ................................................................................ 133

5.2.9 Alkaline Phosphatase activity ............................................................ 137

5.3 Conclusions .............................................................................................. 145

6. Deposition of apatite on inter-bonded 3D fibrous scaffolds and its effect

on cell performances ..................................................................................... 146



6.2.1 Plasma treatment ............................................................................... 147

6.2.2 Surface morphology ........................................................................... 150

6.2.3 Apatite formation ............................................................................... 152

6.2.4 Water contact angle ........................................................................... 154

6.2.5 Cytotoxicity study............................................................................... 156

6.2.6 Cell morphology, distribution and migration .................................... 158

6.2.7 Cell proliferation ............................................................................... 164

6.2.8 Alkaline phosphate activity ................................................................ 166

6.3 Conclusions .............................................................................................. 177

7. Inter-bonded 3D fibrous scaffolds with antibacterial surface coating and

cell growth performances ............................................................................. 178



7.2.1 Electroless plating of silver ............................................................... 179

7.2.2 Electrostatic assembly of PAA-PHMB .............................................. 187

7.2.3 Comparison ....................................................................................... 198

7.3 Conclusion ................................................................................................ 198

8. Summary and Outlook ............................................................................. 200

8.1 Conclusions .............................................................................................. 200

8.2 Suggestion for Future Work ..................................................................... 201

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References ...................................................................................................... 203

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LIST OF FIGURES

Fig.2.1. Schematic illustration of tissue engineering [2] ................................... 8

Fig.2.2. Schematic Diagram of basic dynamic seeding methods, a) spinner

flask, b) filtration method, c) rotary method, d) ultrasonic assisted method, e)

centrifugation force assisted method, f) vacuum assisted method and g)

magnetic force assisted method. ....................................................................... 13

Fig.2.3.Schematic diagram of bioreactor culture system, a) perfusion system,

b) pulsatile flow system.[32] ............................................................................ 14

Fig.2.4. Conventional methods to fabricate 3D scaffolds, a) salt leaching, b)

phase separation, c) gas foaming.[110] ............................................................ 20

Fig.2.5. Schematic diagrams of setups of different rapid prototyping techniques,

a) sterolithography, b) 3D printing, c) selective laser sintering, d) fused

deposition modelling. [152] .............................................................................. 23

Fig.2.6. Schematic representations of a) structural cell printing where both

scaffold and cells are printed simultaneously or serially, and b) conformal cell

printing where cells alone are printed onto thin layers of prefabricated scaffold.

[176] .................................................................................................................. 26

Fig.2.7. a) Flow chart for fabrication of a polymer tubular scaffold, b) a cross-

sectional view of the hybrid scaffolding construct. Scale bar is 100μm. [184] 27

Fig.2.8. a) A hybrid process combining direct polymer melt deposition and

electrospinning, b) Photograph of the overall 3D structure, c) the hybrid basic

unit layer composed of microfibres and nanofibres, d) and e) images in higher

magnifications. [185] ........................................................................................ 27

Fig.2.9. A setup combining 3D plotting and electrospinning systems [186] ... 28

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Fig.2.10. a) Schematic diagram of a hybrid electrospinning process, b) and c)

photographs of the hybrid structure from (b) front view and (c) side view, d)

and e) surface (d) and cross-section (e) morphologies of the hybrid fibrous

structure [187] ................................................................................................... 29

Fig.2.11. a) Schematic diagram of the hierarchical organisation of 3D

structures, b) 3D PCL mesh dried at room temperature, c) 3D PCL mesh

prepared by freeze drying in a cylinder with visible pores, d) SEM image of

dense yarn scaffolds, e) SEM image of 3D scaffold surface with distinctly

isolated yarns made out of aligned nanofibres. [188] ....................................... 29

Fig.2.12. a) Schematic illustration of three PLGA/collagen hybrid scaffolds.

Black; PLGA knitted mesh; Gray: type I collagen sponge. b) SEM observation

of THIN, SEMI, and SANDWICH PLGA/collagen hybrid scaffolds. A, B, C:

top view of the THIN scaffolds; D, G: top view of the SEMI and SANDWICH

scaffolds, respectively; E, H: bottom view of the SEMI and SANDWICH

scaffolds, respectively; F, I: cross-sectional view of the SEMI and

SANDWICH scaffolds, respectively. [189] ..................................................... 30

Fig.2.13. Schematic illustration of macro- and microporous 3D scaffolds a)

SEM image of the co-continuous blend after selective extraction of one

component, b) SEM image of the structure after gas foaming,c) SEM image of

the scaffold prepared by the selective extraction of one component from the

foamed structure. [190] ..................................................................................... 31

Fig.2.14. a) Fibre architecture of an orthogonally woven 3D structure. 3D

structures were woven by interlocking multiple layers of two perpendicularly

oriented sets of in-plane fibres (x- or warp direction, and y- or weft direction)

with a third set of fibres in the z-direction. a) Schematic diagram, b) surface

view of the X–Y plane (scanning electron microscope), c) cross-sectional view

of the Y–Z plane, d) cross-sectional view of the X–Z plane [121] .................. 34

Fig.2.15. a) A cross-section view of rapid prototyping structure and its

traditional 3D micro-CT structure of b) fibre and c) void. ............................... 35

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Fig.2.16. Illustration of a) weft knitting and b) warp knitting structures [224] 36

Fig.2.17. a) Traditional nonwoven fabric processings. An oriented

multifilament yarn was produced by polymer extrusion, then the yarns was

crimped, cut, carded into a lofty web, and further form nonwoven mesh. To

improve the dimensional stability, barbed needles, high pressure water, or heat

was used to entangle the fibres and lock them together; b) wet spinning

technique: polymer solution was spun into a coagulation bath to form

nonwoven structure scaffold simultaneously. This method was used to

fabricate starch-base nonwoven scaffolds [287]; c) melt spinning directly to

web. Polymer pellets were fed into the die, the polymer melt was then extruded

into thin polymer fibres under the heat and air flow. The thin fibres arrange

randomly to form nonwoven mat without further processing [285, 288]; d)

Electrospinning process, polymer solution was extruded into micro-fibre mat

though the force between high voltage and low voltage. By changing the

process condition, such as the concentration of polymer solution, the applied

voltage, the collect distance, various structures can be produced. ................... 39

Fig.2.18. Various nano-structured surfaces, a) nanopits [314], b) nanoparticles

[315], c) nanowires [316], d) nanogrooves, e) nanoridges, f) sharp-tip

nanoridges, g) nanocrystals [317], h) nanorods, and i) nanofibres [318]. ........ 42

Fig.2.19. PLLA film treated with different plasma voltages, a) plain surface;

b), c) and d) plasma treatment to form 70 nm to 100 nm nano-featured

surfaces. [13]..................................................................................................... 43

Fig.2.20. Scheme of template methods. a) Overturning method to form

nanopits, b) directly deposition method to form concaves [321] ..................... 44

Fig.2.21. Diagram of REBL nano-writer concept [322] .................................. 45

Fig.2.22. Chemical etching to get submicro-or nano-structured surfaces [308]

.......................................................................................................................... 46

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Fig.2.23. AFM images of islands formed using different polymer blends, a)

20% w/w poly(p-bromostyrene): 80% w/w poly(styrene), b) 60% w/w poly(p-

bromostyrene): 40% w/w poly(styrene,c) 80% w/w poly(p-bromostyrene):

20% w/w poly(styrene), d) AFM image of 45 nm high PS/PBrS islands,e)

AFM image of 50 nm high PS/PnBMA islands [331] ...................................... 47

Fig.2.24. SEM images of untreated cotton fabric (a and b) and cotton fabrics

assembled with (PAH/SiO2)n multilayers: (c) n = 1, (d) n = 3, (e) n = 5, and (f)

n=7 [334]. ......................................................................................................... 47

Fig.2.25. Apatite coatings with different morphologies, a) and b) apatite

precursor spheres, c) and d) plate-like apatite, e) and f) small plate-like apatite

formed, g) and h) conventional apatite formed after immersing in SBF for 1

and 14 days, respectively.[346] ........................................................................ 49

Fig.2.26. Structure of PHMB ........................................................................... 53

Fig.3.1. a) Vacuum plasma surface treatment unit, b) the plasma treatment

zone ................................................................................................................... 58

Fig.3.2. Cell morphology of fibroblasts (a, and a’) on TCP under low and high

magnification at 37ºC and CHO (b) on coverslip. ............................................ 63

Fig.3.3. Cell morphology of hFOB 1.19 (a, a’ and a”) and Saos-2 (b, b’ and b”)

on TCP under low and high magnification at 34 ºC and 37 ºC ........................ 63

Fig.3.4. Confocal microscope unit ................................................................... 67

Fig.3.5. Flow chart indicating static anti-bacterial study ................................. 70

Fig.3.6. Flow chart showing dynamic anti-bacterial study .............................. 72

Fig.3.7. MESDAN micrometer for fabric thickness measurement .................. 73

Fig.3.8. Mechanical testing machine ................................................................ 75

Fig.3.9. CAM101 KSV contact angle meter .................................................... 76

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Fig.3.10. Setup for water hydraulic permeability testing ................................. 76

Fig.3.11. Vertex 70 FTIR spectrophotometer .................................................. 77

Fig.3.12. Malvern Zetasizer 3000 dynamic light scattering ............................. 78

Fig.3.13. DH-2000-BAL photo-spectroscopy .................................................. 80

Fig.4.1. Flow chart for preparation of fibrous matrices ................................... 82

Fig.4.2. Distribution of the as-prepared short fibre length and diameter, a)

diameter distribution, b) length distribution ..................................................... 83

Fig.4.3. Particle size distribution of sucrose powders ...................................... 84

Fig.4.4. Scheme of fibrous structure formation ................................................ 85

Fig.4.5. DSC curve of PCL pellets, fibres and inter-bonded fibrous matrices . 85

Fig.4.6. Tensile stress-strain curve of scaffolds with different processing

temperatures ...................................................................................................... 86

Fig.4.7. Fibre morphology changes with processing temperatures .................. 87

Fig.4.8. Stress-strain curves of fibrous matrices made from PCL fibres with

different fibre diameters ................................................................................... 87

Fig.4.9. Stress-strain curves of fibrous matrices with different fibre–to-sugar

ratios ................................................................................................................. 88

Fig.4.10. Possible matrix breaking modes, matrix prepared by a) 1/5(wt/wt)

fibre/sugar, b) 1/10 (wt/wt) fibre/sugar, c) 1/20 (wt/wt) fibre/sugar, d) 1/30

(wt/wt) fibre/sugar ............................................................................................ 89

Fig.4.11. Digital photo of the PCL matrices produced from different

fibre/sugar ratios ............................................................................................... 90

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Fig.4.12. SEM images of PCL matrices produced at different fibre/sugar ratios

in the view of front and side, a) and a’) are front and side view of PCL-5, b)

and b’) are PCL-10, c) and c’) are PCL-20, d) and d’) are PCL-30. Scale bar =

250 μm .............................................................................................................. 91

Fig.4.13. μ-CT images of the four fibrous matrices (isometric view of whole,

isometric and front view of magnified central part), scale bar =250μm. ......... 93

Fig.4.14. 2D cross-section images of a) PCL-5, b) PCL-10, c) PCL-20 and d)

PCL-30 in the order of x-y, y-z, and x-z directions from left to right, scale

bar=500μm. ...................................................................................................... 94

Fig.4.15. Pore size, porosity, pore interconnectivity and surface-to-volume

ratio of the four fibrous matrices. ..................................................................... 95

Fig.4.16. Mechanical properties of four different fibrous scaffolds, a) in

elongation mode, and b) in compression mode ................................................ 96

Fig.4.17. Comparison of tensile stress-strain curves between the fibrous

matrices and normal nonwoven matrices, in which PCL-5(1) and PCL-5(2) are

in elongation mode tested in two different directions. PCL-30(1) and PCL-

30(2) are also tested in two different directions. Nonwoven 1 is needle-

punched nonwoven. Its tensile strength was also tested in two directions

(machine and width). Nonwoven 2 is lab-made non-bonded nonwoven fabric.

.......................................................................................................................... 97

Fig.4.18. Water contact angles, hydraulic permeability and water binding

capability of matrix samples (PCL-5, PCL-10, PCL-20 and PCL-30) ............. 99

Fig.4.19. Cell viability of fibroblasts growing in the extract of the scaffolding

materials (p>0.05) ........................................................................................... 100

Fig.4.20. Relationship between seeding volume and seeding efficiency ....... 101

Fig.4.21. Influence of shaking speed of orbital shaker on seeding efficiency 102

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Fig.4.22. Influence of rotating speed of blood roller on seeding efficiency .. 103

Fig.4.23. Influence of stirring speed of spinner flask on seeding efficiency .. 103

Fig.4.24. Confocal microscopic images of CHO cells seeded by static and

dynamic methods. From top to bottom, images were taken with different

magnifications. The bottom images were taken under white light. ................ 105

Fig.4.25. Morphology of dynamic seeded (a, c, e) and static seeded (b, d, f)

CHO cells, a & b are 3 days, c & d are 7 days, e& f are 14 days. .................. 106

Fig.4.26. Comparison of cell number in fibrous matrices between dynamic

seeding and static seeding methods ................................................................ 107

Fig.4.27. CHO cells in fibrous scaffolds (PCL-20) after a) 3 days, b) 7 days, c)

14 days, and d) 21 days of culture. e) CHO cells after 14 days of growing on

PCL film under the same culture condition. (Cells were fixed with OsO4, and

images were taken from front or back of samples. All scale bar = 50 μm. .... 110

Fig.4.28. Fibroblast cells on fibrous scaffolds (PCL-20) after a) 3days, b) 7

days, c) 14 days, and d) 21 days of culture. e) Fibroblast cells after 14 days of

growing on PCL film under the same culture condition. (Cells were fixed with

OsO4, and images were taken from front or back of samples. All scale bar = 50

μm. .................................................................................................................. 112

Fig.4.29. Confocal microscopic images of CHO cells on PCL fibrous matrix

(PCL-10). The cells were fluorescently stained in red for the actin filaments

and in blue for the nuclei. a) showed cells distribute randomly, b) showed cells

ranged orderly along the fibre, c) is the magnified images indicated clearly cell

shape, d) is image with long magnification verified cell distribution in the

matrices. .......................................................................................................... 113

Fig.4.30. Confocal microscopic images of fibroblasts on PCL fibrous matrix

(PCL-10). The cells were fluorescently stained in red for the actin filaments

and in blue for the nuclei. a)-a”) are images with high magnification. b)-b”) are

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images with low magnification, in which a) showing cell growing along the

fibre, a’) indicating cells can bridge between fibres, a”) demonstrated one

whole fibroblasts cell with clear shape; b) is image presenting nuclei only, b’)

is image expressing actin filaments only, b”) is images stacked by b) and b’)

scale bar=50μm............................................................................................... 115

Fig.4.31. Confocal microscopic cross-section slices of CHO cells on PCL-10 at

various z positions (the image were taken every 40 μm throughout the whole

scaffold thickness nearly 300 μm) (PCL-10 after 21 day culture of CHO cells)

Scale bar = 250 μm ......................................................................................... 116

Fig.4.32. Confocal microscopic cross-section slices of CHO cells on PCL-20 at

various z positions (the image were taken every 40 μm throughout the whole

scaffold thickness nearly 300 μm) (PCL-10 after 21 day culture of CHO cells)

Scale bar = 250 μm ......................................................................................... 116

Fig.4.33. micro-CT images of 3D scaffolds with and without cells inside .... 117

Fig.4.34. MTS assay of CHO and fibroblasts on fibrous matrices................. 117

Fig.4.35. CHO cells cultured on 2D plates for three days .............................. 119

Fig.5.1. a) SEM images of synthesised silica nanoparticles, b) histogram of

particle size distribution measured from SEM image, and c) histogram of

particle size distribution obtained by DLS. .................................................... 124

Fig.5.2 SEM images of control (a and a’), 1 bilayer (b and b’), 3 bilayers (c

and c’) and 5 bilayers (d and d’) self-assembled fibre samples. The scale bars

for images a-d are 10 μm, and for images a’-d’ are 200 nm. ......................... 126

Fig.5.3. ATR-FTIR spectra of control and surface modified samples ........... 126

Fig.5.4. XPS spectra of PCL fibrous samples a) non-treated, b) treated with

NaOH solution, c) treated with both NaOH and thin layer of nano-SiO2 ...... 127

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Fig.5.5. High resolution XPS C1s spectra of a) Pristine control sample, b)

NaOH treated sample, c) nano-SiO2 modified sample ................................... 129

Fig.5.6. AFM images of control and surface modified samples. a) pristine

sample, b) 1 bilayer modified sample,c) 3 bilayers modified sample, d) 5

bilayers modified sample. ............................................................................... 129

Fig.5.7. Surface water contact angle before and after surface coating, a) control

sample, b) 1 bilayer, c) 3 bilayers, and d) 5 bilayers coating samples; e) water

contact angle changing with contact time. ...................................................... 130

Fig.5.8. Cell viability of fibroblasts growing in the extract of the scaffolding

materials, bar chart indicates the mean cell number, and the line chart

displayed cell viability. Values shown are the means of three measurements on

individual specimens. Error bars show ± standard deviation (p>0.05). ......... 130

Fig.5.9. Cell morphology of fibroblasts growing in the extract of scaffolds after

culture for three days, a) control, b) 1 bilayer, c) 3 bilayers, and d) 5 bilayers;

The scale bar for images a-d are 500 μm, and for images a’-d’ are 100 μm. . 131

Fig.5.10. Mean cell number changes of hFOB 1.19 on different scaffolds as a

function of time. Values shown are means ± standard deviation. * p<0.05

versus 3DC at day 3, † p<0.05 versus 3DC at day 7. ..................................... 133

Fig.5.11. Confocal images of hFOB 1.19 on scaffolds for culture 1, 3, and 7

days (from left to right), a) control, b) 1 bilayer, c) 3 bilayers, d) 5 bilayers

(from top to bottom). The cells were fluorescently stained in red for the actin

filaments and in blue for the nuclei. Scale bar = 100 μm. .............................. 134

Fig.5.12. Cell morphology of hFOB 1.19 on scaffolds after 3 days and 7 days

of culture, a) control-3days, b) 1 bilayer-3days, c) 3 bilayers-3days, d) 5

bilayers-3days, e) control-7days, f) 1 bilayer-7days, g) 3 bilayers-7days, h) 5

bilayers-7days. Scale bar = 2 μm. ................................................................... 136

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Fig.5.13. Alkaline phosphate content of hFOB 1.19 on 1 bilayer coating and

control samples as a function of time (3 days, 7 days, 10 days, 14 days), *p

<0.05 versus 3D control samples and 2D TCP at corresponding day, † p<0.05

versus results on day 7. ................................................................................... 138

Fig.5.14. Formation of apatite on the silica nanoparticle coated PCL fibrous

matrices surface after immersing into 1×SBF for 7 days. a) & a’) pristine

control sample, b) & b’) silica treated samples just after self assembly, c) & c’)

silica treated samples after immersing in 1×SBF for 3 days, d) & d’) silica

treated samples after immersing in 1×SBF for 21days. .................................. 139

Fig.5.15. Stained ALP of hFOB 1.19 on 1 bilayer coating and control samples

at day 3, a) & a’) control sample with ALP stain only, b) & b’) silica treated

sample with ALP stain only, c) & c’) control samples with ALP and nucleus

stain together, d) & d’) silica treated samples with ALP and nucleus stain

together. .......................................................................................................... 141

Fig.5.16. Stained ALP images of hFOB 1.19 on 1 bilayer coating and control

samples at day 7, a) & a’) control sample with ALP stain only, b) & b’) silica

treated sample with ALP stain only, c) & c’) control samples with ALP and

nucleus stain together, d) & d’) silica treated samples with ALP and nucleus

stain together. .................................................................................................. 142





stain together. .................................................................................................. 143





stain together. .................................................................................................. 144

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Fig.6.1. XPS survey spectra of a) pristine PCL matrix and b) after N2 plasma

treatment ......................................................................................................... 147

Fig.6.2. Curve-fitted XPS spectra of C1s and N1s, a) pristine control sample, b

& c) N2 plasma treated samples ...................................................................... 149

Fig.6.3. Surface morphology of PCL fibres after soaking in SBF solution for 3

days, a & a’) pristine control matrix, b & b’) N2 plasma treated sample. ...... 150

Fig.6.4. Surface morphology of PCL fibres after soaking in SBF solution for

21 days, a & a’) pristine control sample, b & b’) N2 plasma treated sample. 151

Fig.6.5. SEM images of the PCL matrices after N2 plasma treatment and

soaking in 10× SBF for 4 and 24 hours. a & a’) soaking for 4hrs, b & b’)

soaking for 24hrs. ........................................................................................... 152

Fig.6.6. SEM-EDX scanning results of the apatite treated PCL fibre samples, a)

soaking in 1×SBF solution, b) soaking in 10×SBF solution. ......................... 153

Fig.6.7. Ca and P ion concentration changes in 1×SBF and 10×SBF solutions

as a function of time ....................................................................................... 154

Fig.6.8. Surface water contact angles before and after N2 plasma treatment, a)

control sample, b) N2 plasma treated sample.................................................. 155

Fig.6.9. Water contact angles of the PCL matrices as a function of soaking

time, a) 1×SBF, b) 10×SBF. ........................................................................... 156

Fig.6.10. Cell cytotoxicity results using fibroblast cells, the bar shows mean

live/dead cell number after 3 day culture; the line indicates the mean cell

viability (p>0.05). ........................................................................................... 157

Fig.6.11. Fibroblast morphology after cultivation for 3 days in the extract

media of different samples, a) Control, b) O2 plasma treated, c) N2 plasma

treated. ............................................................................................................ 158

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Fig.6.12. Morphology of hFOB1.19 and Saos-2 cells grown on PCL fibrous

matrices after 7 days of culture, a &a’) hFOB1.19 on control, b & b’)

hFOB1.19 on apatite treated samples; c & c’) Saos-2 on control samples, d& d’)

Saos-2 on apatite treated samples. Cells were stained with 1% toluidine blue in

2% borax and 0.04% methylene blue. The scale bar for images a-d is 500 μm,

and for images a’-d’ is 100 μm. ...................................................................... 159

Fig.6.13. Morphologies of hFOB1.19 and Saos-2 cells in different culture

periods, a1~3) hFOB1.19 on apatite treated samples after 1 day, 3 days, 7 days

and 14 days of culturing; c) hFOB1.19 on control samples 14 days of culturing;

b-1~3) Saos-2 on apatite treated samples after 1 day, 3 days, 7 days and 14

days of culturing; d) Saos-2 on control samples after 14 days. ...................... 161

Fig.6.14. hFOB1.19 cultured on PCL fibrous samples at different culturing

times, a-1~4) on control samples after 1 day, 3 days and 7 days, and 14 days

of culturing; b-1~4) on the apatite treated samples after 1 day, 3 days, 7 days

and 14 days of culturing; scale bar = 10 μm.................................................. 163

Fig.6.15. Saos-2 cultured on PCL fibrous samples at different culturing times,

a-1~4) on control samples after 1 day, 3 days and 7 days, and 14 days of

culturing; b-1~4) on the apatite treated samples after 1 day, 3 days, 7 days and

14 days of culturing; scale bar = 10 μm ......................................................... 164

Fig.6.16. Proliferation of osteoblastic cells on PCL matrices with or without

apatite coating ................................................................................................. 165

Fig.6.17. ALP activity of Saos-2 and hFOB1.19 on different samples .......... 167

Fig.6.18. Images of hFOB1.19 stained with alkaline phosphate on PCL fibrous

samples at day 3, a) & a’) control sample with ALP stained only, b) & b’)

apatite treated sample with ALP stained only, c) & c’) control samples with

ALP and nucleus stained together, d) & d’) apatite treated samples with ALP

and nucleus stained together. .......................................................................... 169

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Fig.6.22. Images of Saos-2 stained with alkaline phosphate on PCL fibrous













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Fig.7.1. Photographs of PCL matrices with silver coated surfaces. The scale

bar=100μm. ..................................................................................................... 179

Fig.7.2. Photographic and SEM images of PCL matrices, a) uncoated PCL

matrix; b & c) PCL matrices coated with silver nanoparticles at different

plating times (low & high); d) untreated PCL matrix; e & f) Ag-coated PCL

matrices in low and high magnification.......................................................... 180

Fig.7.3. TEM images of Ag-plated PCL matrices. The silver coating layer is

around 100~120 nm in thickness. ................................................................... 181

Fig.7.4. Particle size distribution histograms tested by a) SEM image and b)

DLS (dynamic light scattering) ...................................................................... 182

Fig.7.5. XRD pattern of Ag coated PCL matrix ............................................. 182

Fig.7.6. UV-Vis spectra of Ag solution released from the coated fibrous

samples with single layer coating as a function of soaking time .................... 182

Fig.7.7. Antibacterial efficacy of different sample groups as a function of time

(Escherichia coli , E.coli)................................................................................ 183

Fig.7.8. Effects of nanoparticles on Escherichia coli (E.coli), a) after contact

with the Ag-coated PCL fibrous matrix for 7 hours, cell surface was

surrounded by Ag particles; b) E.coli cells had an irregular surface (cellular

shrinkage). ...................................................................................................... 184

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Fig.7.9. Cytotoxicity test results by culturing fibroblasts in extract liquids

from the Ag-plated fibrous samples and control samples (p>0.05) ............... 185

Fig.7.10. MTS results of fibroblasts cultured on PCL matrices and nAg-

anchored PCL matrices ................................................................................... 186

Fig.7.11. Visible spectra of the stained fibre samples with different PHMB

layers ............................................................................................................... 187

Fig.7.12. FTIR spectra of PCL fibrous samples before and after PHMB

assembling ...................................................................................................... 188

Fig.7.13. Water contact angles of different layers of PHMB coated films .... 189

Fig.7.14. Visible spectra of the PHMB-coated samples before and after

sterilising treatment, Inserted images were digital photos of the corresponding

samples. .......................................................................................................... 190

Fig.7.15. Bacterial growth and antibacterial rate of fibrous matrices having

different layers of PHMB treated samples after 4hrs contact ......................... 191

Fig.7.16. Bacterial growth as a function of time ............................................ 192

Fig.7.17. Cell viability results of the fibrous samples with different layers of

PHMB (p>0.05) .............................................................................................. 192

Fig.7.18. Cell growth morphology, a) & a’) on control samples, and on the

PHMB assembled fibrous samples with b) & b’) 1 layer PHMB coating, c) &

c’) 1 bilayer PHMB/PAA coating, d) & d’) 3 bilayers PHMB coating, e) & e’)

5 bilayers PHMB coating and f) & f’) 7 bilayers PHMB coating .................. 193

Fig.7.19. Mean cell number of fibroblasts on different samples after 7 days of

cell culturing ................................................................................................... 194

Fig.7.20. Confocal microscopic images of fibroblasts after 7 days of cell

culture, a) pristine fibre sample, and the fibrous sample with b) 1 PHMB layer,

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c) 1 PHMB/PAA bilayer, d) 3 PHMB/PAA bilayers, e) 5 PHMB/PAA bilayers,

f) 7 PHMB/PAA bilayers. Scale bar = 100 μm ........................................... 195

Fig.7.21. SEM images of fibroblasts after 7 days of culturing on fibrous

matrices, a) pristine fibre sample, and the fibrous sample with b) 1 PHMB

layer, c) 1 PHMB/PAA bilayer, d) 3 PHMB/PAA bilayers, e) 5 PHMB/PAA

bilayers, f) 7 PHMB/PAA bilayers. Scale bar = 100 μm .............................. 196

Fig.7.22. Fibroblast number on the control and the PHMB assembled (7

bilayers) samples. ........................................................................................... 197

Fig.7.23. SEM images of fibroblasts after 3 and 7 days of culture, a) pristine

fibre sample-3 days’ culture, b) pristine fibre sample-7 days’ culture, c) fibrous

sample with 7 PHMB/PAA bilayers-3 days’ culture, and d) fibrous sample

with 7 PHMB/PAA bilayers-3 days’ culture. Scale bar = 50 μm .................. 198

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LIST OF TABLES

Table 3.1. Ion Concentrations of the SBFs and Human Blood Plasma in Total

and Dissociated Amounts ................................................................................. 59

Table 4.1. Densities of different matrix samples ............................................. 90

Table 6.1. Elemental contents of the fibrous matrices before and after plasma

treatment ......................................................................................................... 148

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ABSTRACT

Scaffolds in tissue engineering play a crucial role in guiding cell growth and

tissue generation. Fibrous materials have been shown advantageous as tissue

engineering scaffolds because of the large surface area, high porosity and

excellent pore interconnectivity. However, fibrous materials are typically

unstable in pore size and morphology due to the lack of effective bonding

between fibres, and they also have difficulty in pore size controlling. This

remains a big challenge in scaffold fabrication. The functional surface of

scaffolds could further enhance cell performance. For example, nano-scaled

topography surface could assist in cell attachment, adhesion, proliferation and

even differentiation; scaffolds with a bioactive surface assists in cell-substrate

interaction; antibacterial agents could be employed in scaffolds to protect cells

and tissues from infection. So it is important to pursue a suitable surface for

successful scaffold in clinical application also.

This PhD project focuses on establishing a three-dimensional (3D) fibrous

tissue scaffold that has controlled pore size and inter-bonded fibrous structures,

and examining the effects of the 3D fibrous matrices and functional fibre

surfaces on the cell growth behaviour.

To realize it, the 3D fibrous matrices were prepared by combining melting

bonding of synthetic fibres with a template leaching technique. Short

polycaprolactone (PCL) fibres were used for decreasing fibre-fibre

entanglement and conferring isotropic structures. Sugar powder was used as

the template to create large pores in the fibrous structure. The as-prepared 3D

fibrous inter-bonded structure was proved to have desirable pore structures

combining small pores and large pores. Pore size and porosity could be

variable with different processing parameters, and the pores were well-

interconnected. The mechanical properties and stability also were superior to

normal nonwoven mat. Permeability of media inside the scaffolds was

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excellent as well. In-vitro cell culture experiments showed that Chinese

hamster ovary (CHO) and fibroblast cells healthily grew inside the scaffolds.

Nano-scaled surface was created on the 3D inter-bonded scaffolds for

improvement of cell functions. Layer-by-layer electrostatic self-assembly

technique was used to form a nano-structured surface on the scaffolds.

Experimental results revealed that the surface wettability increased after

coating silica nanoparticles on the PCL fibre surface, and the surface roughness

was also improved with the increase of coating layers. This nano-structured

surface dramatically enhanced cell attachment and cell proliferation. Alkaline

phosphatase (ALP) activity study also showed that the nano-structured surface

improved osteoconductivity and osteoindustivity.

To obtain a bioactive surface, apatite was applied to the 3D inter-bonded

fibrous scaffolds through pre-treating the scaffolds with vacuum plasma

followed by soaking them in simulated body fluid (SBF) solution. The plasma

treatment increased the surface reactivity, offering surface with good water

absorptive ability. Thus when the scaffold was soaked in SBF solution, the

plasma treatment was found to speed up the apatite nucleation and deposition.

The apatite-coated scaffolds showed improvement in cell attachment.

Osteoblastic cells were observed to proliferate and differentiate better in the

apatite-coated 3D fibrous scaffold compared to the untreated control.

In the last part, silver and polyhexamethylene biguanide hydrochloride (PHMB)

were chosen as antibacterial agents to increase the anti-infection ability. Using

electroless plating and layer-by-layer electrostatic self-assembling methods,

silver and PHMB could be respectively applied to the fibrous matrices. The

treated scaffolds showed ability to effectively kill bacteria. However, silver-

treated scaffolds were only good for short time cell growth. The PHMB treated

samples can support the normal growth of cells for 2 weeks.

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Overall, in this thesis, a novel fibrous scaffold was established for potential

tissue engineering application. The fibrous scaffold with a functionalised

surface enhanced cell performance.

-1-

C H A P T E R O N E

Introduction

1.1 Significance and research problems Tissue engineering is a modern tissue replacement and therapy technology

involving using scaffolds to support cell growth in vitro which is then

implanted into body for possible tissue replacement. The success of tissue

engineering is highly dependent on materials, cell culture, engineering, and

biological technologies. Tissue scaffolds as a key component of tissue

engineering must meet the structure (e.g. shape, pore size, porosity, pore

interconnectivity), mechanical (e.g. load bearing or non-load bearing), surface

and functional criteria so that the cells can grow as if they are in the native

environment.

The future trends of tissue engineering will be based on using biodegradable

polymer-based scaffolds because they can be biodegraded in the body while

the new tissue is formed, which avoids the immune-reaction in the long-term

implantation. They also show advantages in areas such as easy processing and

property optimization.

Recently, fibrous materials have been studied extensively in the tissue

engineering field. In comparison with other porous tissue scaffolds, fibrous

materials have unique features such as high surface area, large porosity, and

well-interconnected pore structures. The main issues for the fibrous scaffolds

are the lack of strength for load bearing tissue applications, and the insufficient

thickness for 3D scaffolding applications.

Nonwoven materials have good pore interconnectivity and porosity. Their

large surface area also assists in cell attachment. Moreover, the nonwoven

technique is cost-effective compared to other fibre processing techniques.

Some nonwovens have been used in tissue engineering therapy. However, the

existing nonwoven technique meets problems with controlling the pore size,

-2-

forming stable pore structure and having adequate mechanical strength in all

matrix directions.

The key to forming stable pores in nonwoven is to make fibres bond together.

This is relatively easy to be achieved through melt or solvent bonding of the

fibres. Conventionally, the pore size and porosity of a nonwoven can be

adjusted in a very small range through the control of fibre diameter and

nonwoven material density.

Particle leaching is another technology to prepare porous materials. The porous

structure in this technology is constructed through adding particulate materials

to a polymer, followed by leaching off the particles after the polymer material

is solidified. The pore size and porosity are controllable through particle size

and the ratio between particulate template and polymer. However, the main

issue for particle leaching technique is that the porous structures produced have

low pore-interactivity. The first question of this thesis is:

Research question 1: Whether the pore size, porosity and isotropic

mechanical strength of a nonwoven tissue scaffold can be controlled by

combining conventional nonwoven with a particle leaching technology?

Surface topography of a scaffold has close relationship with the cell attachment,

adhesion, proliferation and differentiation. Topography itself also has an

indirect influence on surface wettability, which affects protein absorption and

media accessibility. It is believed that nano-scaled structure can mimic native

extracellular matrix, which could be more effective in cell function. Nano-

scaled topography surface can improve not only the cell attachment and

adhesion, but also the cell spreading and sometimes proliferation. Nano-sized

patterns on the surfaces could direct the cell elongation and growth as well.

Many technologies have been used to modify the surface to understand the

influence of surface roughness and biochemistry characteristic on cell growth,

however, most of the studies are based on 2D films or membranes. The

understanding of nano-surface in 3D tissue scaffold structures is still limited.

-3-

Layer-by-layer electrostatic self-assembly has been used widely in the

biomedical field for incorporation of bio-molecules into scaffolds. This

technique offers the possibility to control the coating quantity and coating

thickness. The second question to be answered in this thesis is:

Research question 2: Can layer-by-layer self-assembly technique be used

to deposit nano-structured coating on 3D fibrous scaffold and how the self-

assembled coating influences the cell growth?

As mentioned, the surface chemistry plays a key role in determining the cell-

substrate interaction and associated cell attachment, adhesion and proliferation.

Surface biochemistry activity also affects the tissue integration into the target

body during the implantation. How to get a bioactive surface or improve the

scaffold bioactivity for 3D scaffold is still a hot topic in tissue engineering.

Hydroxyapatite is a bioactive component to enhance biocompatibility,

bioactivity and osteoconductivity. Introducing apatite to scaffold surface can

effectively enhance the integration between scaffold and host tissue. Among all

the techniques on introducing apatite to surface, soaking in a simulated body

fluid (SBF) solution is the best because it leads to amorphous apatite similar to

the natural bone. The problem for SBF soaking method is that it usually

requires a long soaking time to get a layer of sufficient apatite thickness on the

polymer surface.

Plasma technique can be used to modify surface properties. After plasma

treatment, the surface changes not only in chemical component, but also in

electrical charges. The surface wettability can also be improved. Plasma

treatment can introduce active groups onto polymer surface, which can be used

to further graft functional molecules. In this project, plasma technique was

used to accelerate the processing of apatite nucleation generation when dipping

into SBF solution.

Then it comes to the third research question:

-4-

Research question 3: Can vacuum plasma technique be used to speed up

the apatite deposition on 3D fibrous matrices? How does the plasma

induced apatite coating affect the growth of osteoblastic cells within 3D

fibrous matrices?

Infection is a fatal factor leading to failure of implantation. The importance of

bacterial infection has not been realised in the tissue engineering area. Normal

sterilising methods including 70% ethanol, UV irradiation or ionizing radiation

are either time consuming/costly, or not suitable for polymer scaffolds, and

they may not be able to prevent bacteria from invading the matrix.

Surface antibacterial treatment is a convenient method to prevent cells from

infection. For this purpose, antibacterial agents could be useful. So the last

research question of this PhD project is:

Research question 4: Could antibacterial agent be applied to 3D fibrous

scaffolds? How do the antibacterial agents influence the cell culture

performance?

In general, this PhD project aims to develop better 3D fibrous tissue scaffolds

and understand the interactions of 3D fibrous structure, surface structure and

activity and cell growth attributes. The ultimate aim is to find ideal artificial

scaffolds mimicking the native extracellular matrix in both structure and

function to best support cell growth and function.

1.2 Specific aims Soft fibres are difficult to mix uniformly with particles because the fibre-fibre

entanglement prevents the fibres from blending efficiently with particles. The

fibre-fibre entanglement can be reduced by shortening the fibre length. It has

been established that when the fibre length is reduced to less than 2 mm, the

fibre entanglement is diminished considerably. When short fibres are mixed

with particles, the fibres could also orient in any possible directions. If such a

short fibre-particle mixture is heated to a certain extent, the fibre component

-5-

could be melt, while the particle component still stays in a solid state. The

fibres will be bonded together only in the contacting parts, therefore forming

an inter-connected fibrous network. By dissolving the particles from the melt

bonded mixture, a nonwoven fibrous structure could be produced, and the pore

size and porosity could be controlled through the particle size and the

fibre/particle ratio. The first aim of this PhD research project is:

Specific aim 1: Development of pore-stabilised 3D nonwoven fibrous tissue

scaffolds by using short synthetic fibres and soluble particles as materials,

and controlling the pore size and porosity through particle size and

fibre/particle ratio.

In this project, Polycaprolactone (PCL) will be used as a model polymer to

build 3D fibrous scaffolds. Sugar powder will be used as template materials.

The reasons for using PCL and sugar as fibre and template materials are that

PCL is thermoplastic and biodegradable polymer, which allows melt

processing and being degraded for scaffolds fabrication and application. Sugar

is a non-toxic chemical, thus it wouldn’t bring any side-effect to scaffolds even

with incomplete removal.

The surface features from 3D fibrous matrices can be modified for cell

attachment. If a hydrophobic polymer is used, the water repellent porous

feature makes it difficult for nutrients to be introduced. Silica nanoparticles are

known to have very good water wettability. If silica nanoparticles are applied

on to the surface of hydorphobic fibres, surface wettability of the whole matrix

could be improved. Since silica nanoparticles can be coated on the fibre surface

by a layer-by-layer electrostatic self-assembly method, the coating thickness

and surface roughness could be controlled through the assembly layers. In this

way, the silica nanoparticles could be used to assemble nano-structured surface

on 3D fibrous matrix.

Specific aim 2: Application of silica nanoparticles on PCL fibrous matrices

via the electrostatic self-assembly technique, and studying the influence of

-6-

this treatment on osteoblast cell attachment, proliferation and

differentiation.

Apatite has been proved to be excellent in biocompatibility, bioactivity and

osteoconductivity. Scaffolds with this surface could be easily integrated with

human bone during implantation. However, for PCL fibre matrices, it may take

a long time to form apatite on the surface due to the lack of reactive groups on

the surfaces. To speed up the process, vacuum plasma treatment, followed by

forming apatite coating in SBF solution could be a simple and efficient option.

Specific aim 3: Using plasma treatment to activate the surface of fibres so

that a layer of apatite coating can be formed efficiently on the 3D fibrous

scaffold and demonstrating the existence of the apatite coating and its

benefit to the attachment, proliferation and differentiation of bone cells.

Silver, as the most effective antibacterial agent since ancient time, can kill over

650 disease-causing organisms in the body, even at low concentrations.

Electroless plating method is often used to form a thin layer of silver on

substrates.

Polyhexamethylene biguanide (PHMB) is another effective antibacterial agent

to gram-positive and gram-negative bacteria, fungi, yeast and even HIV-1.

What makes it particularly good for biomedical application is that it has a very

low tissue response and minimal patient discomfort. Thus, it is an ideal

antibacterial agent for tissue scaffold. In this project, these antibacterial agents

will be applied onto 3D tissue scaffolds and their effects on the cell growth

performance will be examined. The last aim of this research project is:

Specific aim 4: Applying Ag or PHMB to 3D fibrous matrix, and

examining the effects of the antibacterial coating on cytotoxicity and cell

culture performances.

1.3 Thesis outline This thesis is divided into eight chapters.

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Chapter 2 reviews existing literatures on 3D tissue scaffolds and their

requirements especially for fibrous scaffolds. The state-of-the-art in this area is

also discussed in terms of the fabrication methods, the scaffold structures and

their advantages and disadvantages for tissue engineering.

Chapter 3 describes the materials, scaffold fabrication and the characterisation

techniques.

In Chapter 4, the preparation method of inter-bonded 3D fibrous matrices is

introduced. Four kinds of matrices with different pore structures have been

successfully obtained and their properties including pore structure, surface

wettability, water permeability, water binding ability, mechanical properties

are examined. Furthermore, cell culture study has been undertaken to

investigate their potential for biological applications.

Chapter 5 presents a novel strategy to improve the surface property of PCL

fibrous scaffolds by layer-by-layer coating of silica nanoparticles on fibre

surface. This special surface treatment has very little influence on the nutrient

and waste exchange within the fibrous structure. The as-prepared scaffold with

novel surface has been tested for its cell compatibility, including cytotoxicity,

cell proliferation and cell differentiation.

In Chapter 6, bone-like apatite is introduced on the 3D fibrous PCL scaffolds

by pre-treating the PCL scaffolds with vacuum plasma and then immersing

them into simulated body fluid (SBF). The apatite formation has been

characterised and the possibility of using this scaffold for bone tissue

engineering is also explored.

Chapter 7 deals with 3D fibrous PCL scaffolds having antibacterial surface

coating (Ag or PHMB) and their antibacterial properties and cytotoxicity.

Chapter 8 finally summarises the entire work performed on the 3D inter-

bonded fibrous scaffolds. Suggestions for further research in this area are also

given.

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C H A P T E R 2

Literature Review

2.1 Tissue Engineering and Scaffolds

As first described by Langer and Vacanti [1], tissue engineering is an

interdisciplinary field that applies the principles of engineering and life

sciences towards the development of biological substitutes that restore,

maintain, or improve tissue function or a whole organ. It is aimed at the

development of biological alternatives for harvested tissues, implants and

prostheses. The basic idea for tissue engineering is to take cells from human

body and culture them on “scaffolds” into a tissue construct which is finally

transferred to human body for replacement or therapy purposes (Fig. 2.1).

Fig.2.1. Schematic illustration of tissue engineering [2]

Tissue scaffolds play a central role in tissue engineering. Scaffolds are

designed to provide anchorage for cells to attach, proliferate, differentiate, and

grow on, furthermore to form an extracellular matrix (ECM)” [3], from which

-9-

the important function of scaffold is outlined. They act as a supporting

template for cell growth. This is especially important for cells, which are

anchorage-dependent and require a specific environment to grow.

The success in tissue engineering requires not only an understanding of the

cells response towards the outer scaffolding environment, and designing

suitable scaffolds for a specific application, but also successful integration of

the scaffold construct into the human body without any immune reaction,

which are the focuses of research on scaffolds for tissue engineering.

The parameters influencing cell performance are complex and have not yet

been fully understood. It is believed that the scaffold structure, culture method

and medium component affect cell functions tremendously [4]. Parameters

concerning the matrix material itself, including mechanical strength (e.g.

stiffness, elasticity, tension), surface characteristic (e.g. surface chemistry,

topography, wettability, electricity) and pore architecture (e.g. pore size,

porosity, pore shape and interconnectivity), are important in regulating cell

growth and proliferation. Many research efforts have been devoted in the past

decades to understand these parameters. However, most of them were focused

on 2D structures, which cannot mimic the native tissue or cells in the body.

Since the cell growth in human body is in 3D environment, which is different

from that on 2D structure, 3D structures are expected to positively impact on

the cellular phenotype and functions, especially cell electrophysiological

properties, relative cell distribution, expression profiles and responsiveness to

exogenous signals [5].

It has been found that the criteria for an ideal scaffold vary depending on the

tissue type and location in the human body. Generally, the ideal 3D scaffolds

need to satisfy the requirements below [6]:

1) Biocompatibility;

2) Biodegradability and /or bioresorbability;

3) Suitable mechanical properties;

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4) Suitable pores and appropriate pore surface to induce desired cellular

activities and to guide 3D tissue regeneration;

5) Adequate physiochemical properties to direct cell-material interactions

matching the tissue to be replaced;

6) Ease in regaining the original shape of the damaged tissue and the

integration with the surrounding environments.

Other aspects such as hemocompatibility and anti-infection are also

prerequisites to any treatment therapy [7-8].

With the development of nanotechnology, nano-structured materials have been

taken into consideration for tissue scaffold design. It is usually believed that

nanostructure is good to interface with individual cell [9], because materials on

nano-scale are compatible to extra cellular matrix [10]. Scaffolds made of

nanostructure, or with nano-structured surface have been shown to improve

cell performance or activity somehow [11-13].

Because of the complex and tissue specific requirements, challenges still

remain in development of tissue scaffolding materials mimicking the structures

and functions of native scaffolds.

2.2 Current Developments in 3D Polymer Scaffolds

Three-dimensional scaffolds from metal, polymers and ceramics have been

used for surgical implantation over the past century. However, two main

drawbacks, non-biodegradablity and limited processing ability [14], have

prevented metal and ceramic scaffolds from wide applications in the tissue

engineering field. In contrast, polymer scaffolds are free of these issues. They

can be biodegradable and the degradability can be controlled easily through

tailoring the polymer structure or using polymer blend. Polymer materials can

be processed into 3D structures without using high temperature and expensive

-11-

machinery. In the following part, a general review will be given on the state of

the art in 3D polymer scaffolds.

2.2.1 Cell lines The human body comprises about 100 trillion cells with about 260 different

phenotypes, that divide, differentiate and self-assemble over time and space

into an integrated system of tissues and organs, and precise replication of a

functionally viable tissue is not a mean task [15]. However, not every kind of

cells can be separated from human body. There are two main kinds of cells

used now for tissue engineering: somatic cells and stem cells. They have

significantly different self-renewal and differentiation abilities. Somatic cells

are usually separated from special organ or tissue and are cultured in special

media for tissue/organ regeneration. They are often used in tissue engineering

study, therefore not limited to healthy patient. Stem cells provide a multi-

potential cell source for tissue engineering. Stem cells are usually divided into

hematopoietic stem cells (HSCs), which can give rise to all the blood cells, and

bone marrow stromal cells (mesenchymal stem cells, MSCs), which

differentiate into skeletal muscle cells. HSCs may differentiate into other three

major cells: brain cells, skeletal and cardiac muscle cells, and liver cells. MSCs

may also differentiate into osteoblasts, multilayered epidermis-like cells,

chondrocytes, adipocytes, and other myocytes like cardiac muscle cells [16].

The phenotype of the cells is stable throughout culture and there is no loss in

osteogenic, chondrogenic or aipogenic potential. Due to these advantages,

much research attention has been paid to stem cells, although there are

challenges on controlled differentiation [17] and debates as to whether the cells

can be differentiated and grown rapidly enough in vitro to provide sufficient

populations to meet clinic need due to their low division frequency [18].

2.2.2 Cell seeding and culture methods Seeding of cells into scaffolds might play a crucial role in determining the

progression of tissue formation [19-20]. The original cell distribution on the

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scaffolds and the fate of further tissue development are decided by seeding

methods.

Methods to seed cells into scaffolds are normally divided into static and

dynamic seeding methods. Static seeding method was initially established for

2D scaffolds. It involves dropping cell/ media suspension manually onto

scaffolds. The cell distribution is dependent on the scaffold structure and the

operator. For a flat scaffold, the cell-media can spread evenly. For 3D

scaffolds, cell distribution depends on the locations of where the cell medium

reaches, and also most of the cells only attach on the superficial layers of the

3D scaffolds. Although it is easy to operate, it faces challenges to seed cells on

3D scaffolds of large volume.

Dynamic cell seeding has been used to overcome the drawbacks of the static

seeding method. Dynamic seeding can be subdivided into the following

methods:

1) Spinning flask (Fig.2.2-a) [21-22]

2) Filtration method (Fig.2.2-b) [23]

3) Rotary method (Fig.2.2-c) [24]

4) Ultrasonic assisted method (Fig.2.2-d) [25]

5) Centrifugation force assisted method (Fig.2.2-e) [26-27]

6) Vacuum assisted cell seeding method (Fig.2.2-f) [28]

7) Magnetic force assisted cell seeding method (Fig.2.2-g) [29]

8) Perfusion bioreactors (Fig.2.3) [32]

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Fig.2.2. Schematic Diagram of basic dynamic seeding methods, a) spinner flask, b) filtration method, c) rotary method, d) ultrasonic assisted method, e) centrifugation force assisted method, f) vacuum assisted method and g) magnetic force assisted method.

Bioreactor devices came into use since the formation of 3D scaffolds. It can

effectively seed a large number of cells into the scaffolds uniformly [30-31].

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As shown in Fig.2.3 [32], bioreactor devices have two basic models: perfusion

system and pulsatile flow reactor. Except being excellent for seeding cells into

3D scaffolds, bioreactors are good candidates for cell culture as they can

precisely control the environment and operation condition, such as pH value,

temperature, nutrient supply and waste removal. However, dynamic seeding

always needs large volume of fresh media, and an extra shear force might exist

on cells during the seeding time.

Fig.2.3.Schematic diagram of bioreactor culture system, a) perfusion system, b) pulsatile flow system.[32]

After being cultivated on the scaffolds, the cells are undertaken a period of

culturing. The culture environment may have a greater impact on cellular

behaviour than the intrinsic origin of the cells [33-34]. The culture can be

divided into static or dynamic modes. Static culture refers to that the seeded

scaffolds were statically cultured under a certain temperature and O2 and CO2

content. Dynamic culture usually deals with culturing the seeded scaffolds in a

reactor, as mentioned in Fig.2.3. Pressures make it possible to refresh media

and take out waste. It has been pointed out that by using a perfusion reactor,

the central oxygen concentration can sustain at a level near 4% after 5 days

compared to static culture, (0% for the same culture period), which prevents

cell from death and forms more homogeneous tissue eventually [35-36]. In

addition, the uniformity of cell distribution might affect cell cycle [28, 37].

Despite the great potential for dynamic culture to produce high-throughput

artificial tissues, the process has been a challenging one. For example, the

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flow-induced shear stress has an impact on cell growth. The collagen

synthesised from the fibroblasts cultured on 3D scaffolds is hardly retained

within the construct during culturing due to the severe extraction by the media

solution [38].

2.2.3 Materials In terms of scaffold itself, material is a basic element to determine the success

of tissue engineering. Polymers have their own characteristics like

bioresorbability, biodegradability, easy processing comparing to metals and

ceramics as mentioned previously. They can come from different sources such

as natural polymers and synthetic polymers. They can be classified into

biodegradable and non-degradable/biostable polymers according to the

degradation ability [39]. Although each polymer has its own feature, one

polymer usually cannot meet all requirements for making an ideal tissue

scaffolds. Synthesis of complex polymer with as many as possible necessary

properties and blending different polymers to get improved performance have

been the common strategies of developing scaffold materials over recent years

[40]. Reviews on polymeric biomaterials can be found in the literatures [39,

41].

Natural polymers

The commonly used natural polymers for tissue scaffolds are collagen,

chitosan, starch, alginate and elastin [6, 42-45]. They can be naturally isolated

from native tissues, thus easily mimic the native cellular milieu, but problems

related to variation between different batches, poor mechanical property, and

less controlled biodegradability still exist.

Collagen is an important natural biopolymer, which can be found from most of

the body organs like skin, bone, tendon, cartilage, and blood vessels. In human,

collagen comprises one third of the total proteins, accounting for three quarters

of the dry weight of skin, and is the most prevalent component of the

extracellular matrix (ECM) [46-47]. It has effective biological properties for

tissue engineering applications and is widely used as a biomaterial already [48-

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49]. The concerns on using collagen are antigenicity and immunogenicity,

potential pathogen transmission and immune reactions [39].

Chitosan is a copolymer of glucosamine and N-acetylglucosamine units

obtained by deacetylation of chitin with a percentage of N-acetyl-glucosamine

lower than 50% [50-51]. It is often used for wound healing, drug delivery

system, and tissue engineering because of its unique nontoxic and anti-

microbial properties [52]. However, it is only soluble in acetic acid solution

hindering the effective fabrication of 3D scaffolds [6, 53].

Starch emerges as a frequently applied material since it has great processing

versatibility [54]. It is a polysaccharide produced by mostly higher order plants

as a means of storing energy, and it consists of two polymers of D-glucose:

amylase (lightly branched polymer) and amylopectin (highly branched

polymer), which make its property variable and tunable [55], conferring the

scaffolds with different properties for different applications [44, 56-57].

Alginate, a natural anionic polysaccharide composed of blocks, (1-4)-linked β-

de-mannuronic acid and α-l-guluronic acid, which can be extracted from brown

algae[45], has gained popularity in biomedical applications because of the

gelling ability in the presence of divalent cations (such as calcium and barium),

stability and viscosity in aqueous solutions [58]. However, the low mechanical

property and slow degradability limit their wide applications [59].

Elastin is the main protein component found in extracellular matrix. It is

synthesised during the growth of tissues. Elastin is composed of cross-linked

insoluble tropoelastin molecules that make it stable for a long time. Elastin is

usually applied in biomaterials because of the combination in both mechanical

strength and flexibility [42]. However, the immune responses and the

insolubility limit its applications in some tissue engineering fields [60].

Synthetic polymers

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Synthetic polymers can be synthesised on large scales and have a repeatable

property. This overcomes the aforementioned problems associated with natural

polymers and offers wide uses as tissue scaffolding materials. Many

biodegradable polymers such as polylactic acid (PLA), polyglycolic acid

(PGA) and polycaprolactone (PCL), have been approved by Food and Drug

Administration (FDA) in United States.

PGA, obtained from polymerisation of glycolide, is a highly crystalline

polymer with crystallinity ranging of 35-75%. It has relatively low melting

temperature, about 185 °C to 225°C [61], and exhibits good mechanical

properties with low solubility in most organic solvents. Its melting properties

make it possible to process scaffolds using thermal techniques. Many different

structures of tissue scaffolds have been developed from this material [62-64].

This polymer also has very good degradable properties. It is known to lose

strength after implantation in body over several weeks, and break down into

glycine that can be absorbed in about 4-6 months. The high degradation rate

has been a concern to some specific applications [65].

PLA is very similar to PGA but the presence of a pendant methyl group on the

alpha carbon. It has a chirality structure of L, D, and DL isomer for the alpha

carbon. According to the molecular weight and polymer processing parameters,

distribution of L- and D-lactide units could be variable. There are two

dramatically different crystallinities for PLLA and PDLLA. PLLA is a semi-

crystalline polymer (~35%) with a melting temperature of 170~180°C [66]. It

is a slow degrading polymer, capable of degrading between 2 and 6 years

depending on the molecular weight [67]. But its high mechanical strength

gives it possible for usage in many fields [68-70]. Poly (dl-lactide) (PDLLA) is

an amorphous polymer. Its amorphous nature permits it losing mass in 12~16

months, but gives its very low mechanical strength [65]. The possibility to

polymerise copolymers having different crystallinities, molecular weights

through adjusting the composition allows for optimising this polymer for

specific tissue engineering applications [71-73].

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Polycaprolactone (PCL) is another semicrystalline polymer and having a

number of advantages to be used as tissue engineering materials. It can be

easily polymerised by a cheap monomer, ε-caprolactone through ring-opening

polymerisation reaction. The melting temperature of this polymer is near

58~60 °C, which has low requirement for thermal processing techniques, and

other active chemical could also be processed without much concern of

degradation during the scaffold fabrications. Also, it can be solved in most of

organic solvents, which is fantastic for processing. The semicrystallinity

makes it degradable and water and carbon dioxide are the main degradation

products, which is nontoxic to cells, tissues/organs, and even the whole body.

PCL also has reasonable mechanical strength, especially high elongation at

breakage [61]. More importantly, such a polymer has excellent

biocompatibility. Except the degradation rate is relatively slow (2-3 years),

PCL has no reactive groups, which is often necessary for scaffolds application.

However, it is still an ideal material for tissue scaffolds especially for bone

regeneration and other long term implanting applications [74-78].

Polyhydoxyalkanoates (PHA), poly 3-hydroxybutyrate (PHB), copolymers of 3-

hydroxybutyrate and 3-hydroxyvalerate (PHBV), poly 4-hydroxybutyrate

(P4HB), copolymers of 3-hydroxybutyrate and 3-hydroxyhexanoate (PHBHHx)

and poly 3-hydroxyoctanoate (PHO) are termed as bacterial polymers because

they are produced by microorganisms under unbalanced growth conditions [79-

80]. Similar to other biopolymers, they are biodegradable and thermo-

processable. But their drawbacks are limited availability and time-consuming

extraction procedure from bacterial cultures [81-82]. As a model polymer,

PHB in this series is introduced below.

PHB is firstly produced in 1920 by bacteria Bacillus megaterium [83-84]. The

polymer has a melting temperature between 160~180 ˚C. It is soluble in a wide

range of solvents. PHB is also a piezoelectric material, which could be a nice

property for scaffolding application because cells are known to react with

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electricity. PHB has good toughness and the mechanical property can be

adjusted by copolymer, like PHBV. PHB has a very low degradation rate due

to its high crystallinity, but its copolymer PHBV has a low crystallinity

resulting in much better degradation behaviour[85].

Polyurethane is usually synthesised by polycondensation of diisocyanates with

alcohols and/or amines, which is regarded as toxic components to human body.

It is biostable and biocompatible, thus has been investigated as long-term

medical implants. Biodegradable polyurethanes have also been developed

using biocompatible aliphatic diisocyanates. Degradable poly (ester urethanes)

were developed by reacting lysine diisocyanate with polyester diols or triols

based on D,L-lactide, caprolactone and other co-polymers with a wide range

of properties [86]. Either biostable polyurethane or biodegradable poly (ester

urethanes) has been widely used in tissue engineering field recently [87-91].

Poly (propylene fumarate) (PPF) is an unsaturated linear polymer, and it is

injectable and biodegradable for orthopaedic applications [92-94]. Due to the

unsaturated fumarate groups on the polymer backbone, the polymer property

can be improved by further crosslinking [95-97]. The degradation products of

this polymer are biocompatible and readily removed from the body [98]. It has

already been used in bone tissue engineering [99-102].

With the well development of polymer science, a huge number of polymers

have been emerged for applications as biomaterials. Here only some commonly

appeared polymers are described. Intensive reviews on biopolymers can be

found in the literatures [65, 103]. Although the properties of synthetic

biopolymers can be tailored to meet specific applications, pre-inflammatory

could be a problem affecting the fate of neo-tissue form in tissue engineering

[104-109].

2.2.4 Scaffold fabrications and structures Many 3D scaffold structures have been developed, by using either traditional

technique (e.g. template-leaching, gas foaming, fibre bonding, phase

separation) (Fig. 2.4), or model processing technique such as solid freeform

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fabrication (SFF) (e.g. stereolithography, selective laser sintering, fused

deposition modeling, 3D printing). However, every technique has its own pros

and cons. It is not easy for one technique to meet all requirements of 3D

scaffolds. Therefore, two or more techniques were often combined together to

take the advantages of each technique. Here the main techniques are

summarized on basis of literature reports.

Fig.2.4. Conventional methods to fabricate 3D scaffolds, a) salt leaching, b) phase separation, c) gas foaming.[110]

Salt leaching technique: this method usually works together with solvent

casting or freeze-drying method. Polymer solution containing porogen

materials is castor freeze-dried to shape into 3D objects and get rid of the

solvent. The porogen particles are then leached away to form a 3D scaffold

[111-113]. The porogen material plays functions in forming pores inside the

scaffolds. The pore size and shape are dependent on the porogen. This method

is easy to undertake, but it is limited to produce very thin film or membrane

samples [114-117]. Although the scaffold thickness can be increased by using

multiple layers, the use of solvent in the processing is still a problem. If the

solvent is not removed completely, cytotoxicity could occur [110].

Fibre related method: this technique uses fibres to build different 3D structures.

Textile technologies of braided, nonwoven, weaving or knitting are normally

the main fibre construction methods [68-69, 112, 118-126]. New emerging

techniques such as electro-spinning, is also employed. Textile techniques are

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matured in producing different fibrous structures. The pores in the fibre

scaffolds are in irregular shapes, which are much different from those in

particle leaching ones. The pore size can be controlled through fibre diameter,

fibrous structure and processing parameters. Also, the pores in fibrous

materials have excellent interconnectivity. Furthermore, the large surface area

provided by fibrous materials makes fibrous scaffolds an ideal candidate for

cell attachment and expansion. The issues for fibrous scaffolds are their weak

mechanical property [127].

Electrospun nanofibre mesh has been used as tissue engineering scaffolds

because it has a similar fibrous morphology to the native extracellular matrix

(ECM) and can provide fibrous scaffolds with instructive surface properties to

direct cell faith [128-129]. Its powerful ability to produce different structures,

such as side-by-side [130], core-sheath [131-132] and aligned nanofibre mats

[133], surface-nanofibre tubes [134-135], makes it one of the most promising

techniques in this century. However, electrospun nanofibre mats are typically

regarded as 2D scaffolds since cells cannot grow and penetrate easily into the

scaffolds [136-137]. To solve these problems, multiple-layered electrospun

mat, or mixing electrospun mat and dual-porosity nanofibrous mat have been

prepared [138-139]. Cell seeding has also been combined with electrospinning

to directly form layered 3D scaffolds seeded with living cells [140].

Thermal induced phase separation (TIPS) applies the mechanism of lowering

system free energy to separate a homogenous multi-component polymer

system into several phases: polymer-rich phase and polymer-lean phase. Once

the solvent is removed, the pores can be generated [141-142]. This technique

generates pores with different morphologies [143-144]. Phase separation can

be divided into solid-liquid, and liquid-liquid phase separations. For solid-

liquid phase separation, the solvent for the polymer solution is crystallized

under certain temperature. After removing the crystals, pores are formed. The

solid phase is discontinuous, except the liquid phase [145-146]. Unlike the

solid-liquid phase separation, liquid-liquid phase separation has an upper

critical and a lower critical solution temperature. The two phases (polymer-rich

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and polymer-lean) are both continuous. An open pore structure can be

produced when the solvent is removed [144]. Phase separation technique

involves the use of solvent, which may be a reason for the limited applications.

Also, solid-liquid method usually produces pores up to tens of microns, which

limits its applications in some specific biological fields [147].

Gas foaming technique is an alternative technique to eliminate using organic

solvent in the 3D scaffold preparation, which is thought to be an improvement

in scaffold fabrication technique. CO2 gas from high pressure to low pressure

changes results in the formation of porous scaffolds [148]. Pores in such

scaffolds are usually interconnected and are hundreds of microns in size [149].

The honeycomb pores also can be prepared from this method [150]. The high

pressure required in this technique limits many biopolymers’ applications in

3D scaffolds [151].

The above conventional scaffold fabricating methods face challenges in

fabricating 3D tissue constructs resembling the native tissue microvasculature

and micro-architectures. New techniques like rapid prototyping, 3D bioprinting

or bioplotting, self-assembling, hydrogel technique, as well as the hybrid

techniques have appeared in recent years.

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Fig.2.5. Schematic diagrams of setups of different rapid prototyping techniques, a) sterolithography, b) 3D printing, c) selective laser sintering, d) fused deposition modelling. [152]

Rapid prototyping (RP) techniques, also termed solid freeform fabrication

(SFF), is the most promising technique to fabricate scaffolds because it can

effectively control the scaffold structure, and design the exact shape of

complex scaffold for tissue replacement via computer-aided design (CAD). It

is also named as 3-D printing (3DP), fused deposition modelling (FDM) and

selective laser sintering (SLS), 3D fibre deposition, and so on. Fig. 2.5 shows

some of the setups. For every technique, it has its own distinct advantages and

limitations, which have been reviewed in several papers [152-156]. 3DP

technique is the most commonly used SFF technique for scaffold fabrication. It

uses powder as raw material to form 3D structure via ink jet printing method,

which allows processing of a wide range of powders from polymers, metals to

ceramics. Additionally, the unprocessed powders are easily trapped in the

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small inner holes and affect the functions of 3D scaffolds. FDM technique uses

polymer melt as the material to form a series of material roads deposited layer-

by-layer by a nozzle with a small orifice. The deposition direction, the filament

thickness and road intervals can be easily modified to obtain different

structures. Due to the requirement of melting state of polymer, this technique

narrows its application in tissue scaffolds. The feedstock materials of 3D fibre

deposition technique can be in pellet or granule forms, and the material flow

can be regulated by applying pressure to the syringe. The 3D structure can be

regulated by fibre thickness and orientation. SLS system also uses powders as

raw materials. However CO2 laser is used to selectively sinter polymer

powders into thin-layered structures. During the SLS processing, the laser

selectively scans the powders, making the powder temperature rise to glass

transition and then fusing them together. 3D structure is built up by layer-by-

layer deposition of the polymer melt. This processing route must use powders

as material, and the exposure of the powder material to a high energy laser

beam might degrade the t polymer chains. Stereo-lithography (SLA) has also

been exploited for 3D scaffold fabrication. A UV laser vector is used to scan

over the top of a bath of a photo polymerisable liquid polymer material. Once

the polymerisation is initiated, the laser beam can create the first solid layer.

Again, a layer-by-layer 3D structure can be formed by tracking the laser beam.

After the model is complete, it is cured in an UV oven to smoothen the surface.

Undoubtedly, the raw materials need to be UV photo-polymerisable, and

compatible to UV curing. A relatively new 3D printing technique, modified by

using cryogenic plotting system, can impart fibres with a rough surface and

hydrophilic properties [43]. Overall, the RP techniques make it possible to

produce 3D structure automatically, rapidly, and accurately. But they still have

limitations in terms of raw material selection, processing conditions, and

expensive machinery.

Indirect 3D printing is based on solid freeform fabrication technique. It uses

the 3D rapid prototyping technique to get a negative mould, and then polymer

solution is casted into the mould to get a 3D scaffold after the mould removal

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[157-158]. This technique can overcome the limitations in polymer selection,

solvent use, and incorporation of localised pores [157, 159-160].

3D hydrogel technique

3D hydrogel porous structure has a high content of water (> 30%) and is

capable of incorporating cells homogenously throughout the gel network. Its

fantastic 3D structure allows the nutrient transport and the injectable gel state

broadens its application in tissue engineering [161-162]. So far, hydrogel has

become one of focuses in tissue scaffolding research. The methods to prepare

hydrogel scaffolds are mostly based on bottom-up technique, including direct-

write assembly [163], protein self-assembly [60], photo-initiated cross-linking

[164-166], thermal cross-linking [166], and physical or chemical cross-linking

of hydrophilic polymers [167-168]. The application varies from muscle cell

culture, cartilage tissue replacement, space filling scaffold, cell-delivery

scaffold, to bioactive molecule delivery scaffolds [169-171]. There are still

many aspects for 3D hydrogel scaffolds to improve, such as the weak

mechanical property [172-173], and sometimes the use of toxic cross-linker

[174-175].

3D cell printing is based on the concept of building tissue or organ layer by

layer from individual cell-containing components. There are usually two cell

printing methods, one is called structural cell printing, and the other is

conformal cell printing. They both build 3D cell constructs with heterogeneous

cell and biomolecular distribution inside the 3D structure. The difference is

that the ink for 3D structural cell printing concludes scaffolding, cells and

biomolecules simultaneously, but the ink for 3D conformal cell printing

contains cell only. In this way, cells are printed into thin layers of prefabricated

cell scaffold (See Fig.2.6) [176]. Most of the 3D printing technique can be

utilized for 3D structural cell printing [177-179].

But the application of harsh or organic solvents prevents many good choices

of polymer materials in the live cell deposition process. Many non-toxic

“bioinks” (hydrogels of gels) have been found, but they often lack rigidity and

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structural integrity [180-181]. The “contradiction” between building up 3D

structure and the deposition of individual cells is another issue. On the

contrary, the 3D conformal cell printing has no issue related to toxic solvent,

and the printing resolution can also reach a single-cell level [182-183]. The

shining point of 3D cell printing to tissue scaffolds is that cells can be

distributed uniformly inside and outside the scaffold structure, and therefore

can grow into homogeneous tissue or organs. Since the whole process deals

with cells, a sterilised environment is necessary.

Fig.2.6. Schematic representations of a) structural cell printing where both scaffold and cells are printed simultaneously or serially, and b) conformal cell printing where cells alone are printed onto thin layers of prefabricated scaffold. [176]

Hybrid technology for fabricating hybrid scaffolding structures is an appealing

method in tissue engineering. Making use of the pros of different techniques is

critical to make hybrid scaffolds. Here some recent hybrid techniques are

introduced.

The electron beam and photolithography methods have been utilized together

to get a tubular scaffold with micro- and nano-structure in the 3D scaffolds to

enhance the scaffolding performance. The nano-pits or nano-pillars are created

using electron beam lithography, while micro-pattern structures are generated

by photolithography (Fig. 2.7) [184].

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Fig.2.7. a) Flow chart for fabrication of a polymer tubular scaffold, b) a cross-sectional view of the hybrid scaffolding construct. Scale bar is 100μm. [184]

A direct polymer melt deposition method has been combined with

electrospinning to produce a 3D structure with dual scaled pores. The

microfibre diameter can be accurately controlled by computer-aided design,

and the electrospun nanofibre mat has been coated on the microfibre surface

and between microfibres to enhance the cell attachment and adhesion for cell

growth (Fig. 2.8) [185].

Fig.2.8. a) A hybrid process combining direct polymer melt deposition and electrospinning, b) Photograph of the overall 3D structure, c) the hybrid basic unit layer composed of microfibres and nanofibres, d) and e) images in higher magnifications. [185]

Other rapid prototyping techniques also can be combined with electrospinning

to get a dual pore size 3D structure. The 3D microstructure enables cell

penetration and nutrient transport, while the electrospun mat provides a nano-

scaled surface, which enhances cell attachment, proliferation and

differentiation. Another example is given in Fig. 2.9 [186].

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Fig.2.9. A setup combining 3D plotting and electrospinning systems [186]

The 3D prototyping technique has also been used to produce a mould for 3D

structure, followed by pouring a polymer solution into it to form a nanofibrous

structure through thermal phase separation. Thus this kind of scaffold ensures

both 3D and nanofibrous structures. The scaffolds having complex geometry

have been proved to be a good candidate for bone tissue engineering [160].

Melt electrospinning and wet electrospinning techniques have been applied

together to produce a 3D scaffold with microfibres as support and nanofibres

as filling for cell attachments. This nano-/micro-fibre scaffolds with a porosity

of 91.2% and a mean pore size of 93.9 μm have shown great potential for 3D

organisation and guidance of cells (Fig.2.10 ) [187].

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Fig.2.10. a) Schematic diagram of a hybrid electrospinning process, b) and c) photographs of the hybrid structure from (b) front view and (c) side view, d) and e) surface (d) and cross-section (e) morphologies of the hybrid fibrous structure [187]

Electrospun nanofibres have been deposited into water and then freeze-dried to

obtain a hybrid 3D structure, in which nanofibres can either assemble into a

loose 3D structure or form a dense 3D structure, dependent on the freeze-

drying condition. This method results in a structure consisting of nanofibres

completely (Fig. 2.11) [188].

Fig.2.11. a) Schematic diagram of the hierarchical organisation of 3D structures, b) 3D PCL mesh dried at room temperature, c) 3D PCL mesh prepared by freeze drying in a cylinder with visible pores, d) SEM image of dense yarn scaffolds, e) SEM image of 3D scaffold surface with distinctly isolated yarns made out of aligned nanofibres. [188]

A hybrid structure combining the advantages of collagen micro-sponge for its

cell affinity with the good mechanical property from knitted PLGA mesh has

been developed by dipping knitted fabric into collagen solution, followed by a

freeze drying procedure. The hybrid structure has a much better cell

distribution and extra cellular matrix deposition ability (Fig. 2.12) [189].

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Fig.2.12. a) Schematic illustration of three PLGA/collagen hybrid scaffolds. Black; PLGA knitted mesh; Gray: type I collagen sponge. b) SEM observation of THIN, SEMI, and SANDWICH PLGA/collagen hybrid scaffolds. A, B, C: top view of the THIN scaffolds; D, G: top view of the SEMI and SANDWICH scaffolds, respectively; E, H: bottom view of the SEMI and SANDWICH scaffolds, respectively; F, I: cross-sectional view of the SEMI and SANDWICH scaffolds, respectively. [189]

Porous scaffolds with macro- and micro-pores have been prepared by a

combination of gas foaming (GF) and selective polymer extraction (PE) from

co-continuous blends. This special scaffold has both macro-porosity suitable

for rapid cell colonisation and tissue in-growth, and micro-porosity serving as

preferential route for fluid transport (Fig. 2. 13) [190].

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Fig.2.13. Schematic illustration of macro- and microporous 3D scaffolds a) SEM image of the co-continuous blend after selective extraction of one component, b) SEM image of the structure after gas foaming,c) SEM image of the scaffold prepared by the selective extraction of one component from the foamed structure. [190]

Most hybrid techniques used recently for tissue engineering have been tried to

get ideal scaffold structures with macro-, micro-, and nano-scales inside. These

structures combine at least two levels of scale characteristics to achieve

enhanced properties. Macro-scale pores can assist tissue in-growth during the

implantation, the micro-pores are effective for mass transport of nutrients and

waste, and the nano-scaled pores are perfect for cell attachment, extra cellular

matrix formation, and cell–matrix interaction. Hybrid techniques are likely to

be the focus of future research in this area.

2.2.5 Applications So far, some commercial tissue engineering products, for example, artificial

bone, cartilage, man-made skin have appeared [191]. Tissue scaffold can be

put into as many possible uses as one can imagine: ligament [192-193], liver

[194-195], tendon [196-197], muscle [168, 198], nerve [137, 199], joints [200-

201], bladder [202-203], vascular [36, 204], blood vessels [205-207], cornea

[208-209], skeleton [210-211], orthopaedic [11, 212], trachea [213], and so on.

The requirement for every specific application is quite different. In this thesis,

the focus is on bone tissue scaffolds.

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2.2.6 Complications

To make successful tissue engineering therapy, the engineered 3D scaffolds

should meet the requirements as previously mentioned in Section 2.1. There

are many attempts to get scaffolds with these characteristics. However, many

difficulties still exist, some of which are given below:

1) Controlling the biodegradation rate so it is about the same as neo-tissue

formation rate. The scaffold has to serve as a supporting construct in the early

stage, with the time, it needs to be degraded and totally absorbed by human

body without any toxic reaction.

2) Controlling the structure with good mechanical property, and appropriate

pore size and porosity to guide cellular activities and 3D tissue formation.

3) Low-cost, easy-processing, repeatable and controllable technique to make

3D scaffolds. The technique also needs to avoid using toxic solvent.

4) Uniform cell distribution, easily cell penetration, and homogeneous tissue

regeneration.

5) Adequate surface physiochemical properties to direct cell-material

interactions.

6) Ease for scaffold to integrate with the surrounding environment.

7) Disinfection characteristic after implantation into the body.

Once these aspects/difficulties are solved, the 3D scaffolds will constitute an

ideal environment to foster the formation of new tissue and to integrate with

the host environment.

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2.3 Fibre-based Tissue Scaffolds

Fibre based structures have been applied in tissue engineering field due to their

many unique features compared to other types of scaffolds. For example, they

have high surface area, high porosity, interconnected pore structures, and

controlled pore size as well. The huge choices of fibre and different processing

routes (eg. woven, knitting, and nonwoven) provide for a wide range of

materials and structures.

2.3.1 Classification of fibre-based scaffolds

The structures from traditional textile processing techniques, including fibres,

yarns, ribbons, woven, knitted and non-woven fabrics, have already been

studied extensively for tissue scaffolding applications for several decades.

Similar structures from emerging techniques, such as electrospinning, self-

assembling, and rapid prototyping also are appealing candidates. In this part, a

review on different fibrous scaffolds is given.

Fibres/Yarns/Braided/Twisted structures

Fibres are usually braided or twisted into yarns for suture or anterior cruciate

ligament applications. For example, a commercial PCL suture was made on a

computer-aided embroidery machine and was coated with collagen I for end

use [214]. PLLA microfibre scaffolds with tunable mechanical properties were

fabricated by braiding or braiding-twist technique for anterior cruciate

ligaments (ACL) [215].

Woven structure

Woven structures are produced on a weaving machine or loom, using warp and

weft yarns. The three basic weaves are plain weave, satin weave, and twill

weave [216]. Woven structures have been used in liver cell culture [194],

cartilage constructs [217], rotator cuff repair [218-219], heart repair [220], and

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tendon regeneration [221]. Because they are usually in plain fabric form with a

thickness less than two yarn diameters, simple woven structures cannot meet

complex shape requirements. To overcome this limitation, 3D woven structures

have been developed, initially for applications in the composite area [222-223].

The fibres are not only in the warp (x) or weft direction (y), but also in a third

direction (z) to form a real 3D structure. 3D fibrous woven structure was

proved suitable for tissue engineering for the first time in 2007 by Franklin et

al [121] (Fig. 2.14).

Fig.2.14. a) Fibre architecture of an orthogonally woven 3D structure. 3D structures were woven by interlocking multiple layers of two perpendicularly oriented sets of in-plane fibres (x- or warp direction, and y- or weft direction) with a third set of fibres in the z-direction. a) Schematic diagram, b) surface view of the X–Y plane (scanning electron microscope), c) cross-sectional view of the Y–Z plane, d) cross-sectional view of the X–Z plane [121]

3D rapid prototyping technique described previously can form 3D structure by

a computer-aided design. The structure is very similar to that of woven

structure except that in woven structure the warp and weft yarns are woven

together, but here the fibres are simply deposited together. Usually yarns

consist of many thin fibres. For 3D fuse deposition the fibre is very thick

relatively. Fig. 2.15 shows a 3D rapid prototyping structure in the cross-

section, and 3D micro-CT images for the fibre parts and the void parts.

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Fig.2.15. a) A cross-section view of rapid prototyping structure and its traditional 3D micro-CT structure of b) fibre and c) void.

Knitted structure

Knitting is different from weaving in that the intermeshing of yarn loops forms

its fabric. As shown in Fig. 2.16, there are two knitting methods: weft knitting

and warp knitting [216]. Knitted structures tend to have very good elastic

property, which makes them particularly suitable for bladder reconstruction,

artery replacement, and tendon or ligament tissue regeneration. Some of these

applications are listed below.

Knitting fabrics can be used as human urinary bladder reconstruction material

because of its various advantages, such as high elasticity, porosity, good

mechanical property and drape. Poly (ethylene terephthalate) (PET) knitting

fabrics have already been studied extensively. Biodegradabe PLA/PLLA/PGA

knitting fabrics have been used to engineer artificial bladders [224].

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Fig.2.16. Illustration of a) weft knitting and b) warp knitting structures [224]

The stability under dynamic stress and the good elasticity are also important

for tissue-engineered artery. Poly (L/D) lactide 96/4 warp knitted fabric has

also been used for artery replacement [225].

Knitted PLGA scaffolds also give better tendon tissue regeneration than

braided fibre scaffolds in terms of the new tissue in-growth and the mechanical

property [196].

Knitted PLGA and PLLA scaffolds can be used to reconstruct anterior cruciate

ligament (ACL) [226], and the hybrid structure based on silk knitted fabric has

also been used in ligament tissue engineering. The hybrid contains silk fibres

for providing mechanical strength, a knitted structure aiming to improve the

internal connective space and a collagen matrix assisting and modulating

neoligament tissue in-growth [192, 227]. To avoid the cell-gel separation from

composite scaffolds of fibrin gel/PLGA knitted structure [228-229] , a bMSCs

sheet was assembled directly into a knitted PLLA scaffold for ligament

engineering [226]. Silk warp knitted fabric cultured with mesenchymal stem

cells has also been investigated for treatment of stress urinary incontinence,

with promising results from animal trials [230-231].

Another hybrid scaffolds have been fabricated by using PLA warp-knit textile

to enhance a gel, in which the gel can act as a cell carrier, but the textile can

provide the necessary mechanical property for the hybrid scaffold for

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cardiovascular tissue engineering [232]. Other application for vascular tissue

engineering also uses gel to fill a knitted valve structure, like using fibrin gel to

fill PCL knitting structure [233].

Hybrid structures based on knitted fabrics have been surface coated with

electrospun nanofibres to get micro-nano scaffold system for ligament/ tendon

tissue engineering [196], and with hydroxyapatite to form hybrid structure for

bone scaffold [234].

Nonwoven structure

Fibres in a nonwoven structure are neither woven nor knitted. In industry,

nonwoven structures are normally made from continuous filaments or from

staple fibre webs strengthened by bonding using various techniques. The

bonding techniques include adhesive bonding, mechanical interlocking by

needling or fluid jet entanglement, thermal bonding and stitch bonding.

Electrospinning usually produces a nonwoven structure also. In Fig. 2.17,

conventional nonwoven processing, as well as some new fibre-mat formation

methods are introduced.

Nonwoven structure was proved to be compatible to blood in 1969, and

became an alternative to woven and knitted fabrics in tissue engineering [235].

It was initially used as cartilage scaffold by Langer and Vacanti [236]. Since

then, nonwoven scaffolds have been developed for a wide range of medical

applications, such as wound dressing [237-238], dermal equivalents [126],

bone tissue engineering (femoral shaft [239], cartilage [240-242]), ligament

tissue [243], smooth muscle tissue [244], blood compatible interfaces [245],

intestinal mucosa [246], and heart valve [247-248] .

Commercial fibres (usually nonbiodegradable) from polymer materials such as

poly(ethyleneterephthalate) (PET) [249], Polypropylene (PP) [250], nylon

[251], polyester (polytetramethyleneterephthalate) [245, 252-253], and

cellulose (regenerated, oxidized) [254-256] have been used in nonwovens.

Fibres specially designed for biomedical uses (usually biodegradable) such as

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from PGA [112, 239-242, 257], poly(L-lactic acid) (PLLA)[258-259],

Poly(L/D) lactide 96/4 [260], PGA and PLLA 50/50 or PGA/PLA 90/10 [261],

chitosan(CS)/chitin or composite of chitosan and PLGA [262-263],

starch/polycaprolactone (SPCL) [264], silk firoin [126], hyaluronan esters

(Hyafft derivatives) [265-267], alginate [268], extracted PEA-derived [269-

270], poly (3-hydroxybutyrate(P3HB) [271], and benzyl ester of hyaluronic

acid (Hyaff-11) [272] have also been employed.

To meet the requirements of wettability, tomography or bio-signals for cell

growth, studies on nonwoven scaffolds have been focused on the surface

modification. To change surface energy and get a hydrophilic surface, surface

treatments include chemical methods (e.g. boiling in base [273-275] or strong

acid [276-278] solution, inducing diethyllaminoethyl (DEAE) groups [249]),

physical methods (e.g. γ-ray or UV lightradiation [237], low-temperature

plasma [260], microwave discharge treatment [245] or direct-current pulsed

plasma treatment [279]), combined methods (e.g. plasma followed by grafting

thermally with acrylic acid (AA) [279], grafted with AA or NIPAAm and then

immobilised with collagen and chitosan [237-238, 279]), coating with

bioactive molecules such as insulin, RGD, heparin, hydroxycarbonated apatite,

or growth factor [253, 277-278, 280-282]. These treatments help to improve

the surface biocompatibility, so the scaffold/cell composites can easily

integrate with the surrounding environments.

Schmelzisen’s group has succeeded in using PLGA nonwoven scaffold to

tissue-engineer bone chips in low-load bearing areas [283-284]. Many

problems related to nonwoven tissue engineering have still not been solved

effectively. Nonwoven structures are not as stable as woven or knitted

structures. Deformation, delaminating or cranking happens at macro levels,

while the fibre slippage within nonwoven scaffold occurs at micro-levels.

These have a deadly influence on the cell fate. The control of porosity, pore

size between the fibres can be achieved by changing the fibre diameter, and

web-laying speed [252, 285]. But a highly porous structure is still hard to

obtain. For electrospun nonwoven fabrics, which have been used in scaffold

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fabrication [128-129], the key issue is still poor cell penetration [136-137].

There have been many attempts to address this issue [125, 127]. Fibres can be

disentangled from nonwoven with tweezers manually [252], which is difficult

to obtain uniform matrices. Hydroentangled nonwoven was cultured with cells

in each layer, and these layers were then assembled into a 3D structure to

overcome the cell penetration problems or cell in-growth limitation [286]. New

nonwoven techniques to modulate the pore structures, including porosity, pore

size and pore interconnectivity are needed.

Fig.2.17. a) Traditional nonwoven fabric processings. An oriented multifilament yarn was produced by polymer extrusion, then the yarns was crimped, cut, carded into a lofty web, and further form nonwoven mesh. To improve the dimensional stability, barbed needles, high pressure water, or heat was used to entangle the fibres and lock them together; b) wet spinning technique: polymer solution was spun into a coagulation bath to form nonwoven structure scaffold simultaneously. This method was used to fabricate starch-base nonwoven scaffolds [287]; c) melt spinning directly to web. Polymer pellets were fed into the die, the polymer melt was then extruded into thin polymer fibres under the heat and air flow. The thin fibres arrange randomly to form nonwoven mat without further processing [285, 288]; d) Electrospinning process, polymer solution was extruded into micro-fibre mat though the force between high voltage and low voltage. By changing the process condition, such as the concentration of polymer solution, the applied voltage, the collect distance, various structures can be produced.

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2.3.2 Issues of fibre-based scaffolds

Fibrous scaffolds are usually not strong enough for load bearing tissue

regeneration. The chemical property of polymer fibres may affect cell

attachment, even though the fibres themselves are usually nontoxic to cells.

Further surface modification is necessary. For fibre bundles, yarns, braided or

twisted structures, their morphologies limit their applications in sutures,

ligaments or tendons. Circular knitted tube/solid structure is suitable for

specific applications like artery, ligament and tendon. Plain woven or knitted

structures are usually thin and flat and cannot be used as 3D tissue engineering

structures. Nonwoven fabrics face a similar challenge. Further developments in

3D woven, knitted, and nonwoven structures are needed. The nonwoven

process is particularly attractive due to its relatively low production cost.

2.4 Role of nano-structured functional surfaces on cell performance of polymer scaffolds

Surface design of polymer scaffolds that can promote tissue regeneration has

become a focus of tissue engineering research. Surface architecture

(morphology and topography) and surface chemistry may have great effects on

cellular responses in the initial stage of cell attachment, and further cell

proliferation, differentiation, and even neo-tissue integration with human body.

Surface topography patterns at different scales can direct cell growth [289].

Patterns at the single-cell scale (0.1-100μm) can provide contact guidance,

which limits cell adhesion to specific locations and influence cell shape and

orientation [290-293]. Patterns in nano-scales are important for cell attachment

[294]. Surface topography is also directly related to ECM formation [295-296].

Surface topography can indirectly affect the surface wettability, thus controls

protein absorption and cell attachment [297].

Surface chemistry is playing a profound role in cell growth and gene

expression. Numerous surface treatment or functionalisation techniques have

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been developed to confer specific surface for cell-related scaffold therapy.

Surface can be loaded with bioactive components for better cell-surface

interaction [298]. Surface can be modified to enhance cell attachments [299].

Conductive property can also be imparted to the surface for controlling neural,

muscle or ligament cells orientation [300]. Surface can have anti-bacterial

property for reducing infection during the implantation [301].

In the rest of this part, nano-structured surface for improvement of cell

performance, bioactivity, antibacterial property of 3D polymer scaffolds is

introduced.

2.4.1 Nano-structured surface for enhancement of cell

performance Cell interaction with the environment has been examined for a long time. It has

already been proved that material surface topology in micrometer-range

features such as grooves, ridges, and wells have a strong influence on the cell

attachment, adhesion and proliferation [10]. Cells even can align, grow and

migrate along a defined topography orientation. This was explained by contact

guidance [302]. It was found that micrometre surface features might influence

cells in morphology, orientation, cytoskeleton and focal contact arrangement,

motility and gene expression [303-305]. The nano-scaled surface architecture

can mimic and provide physiological conditions similar to native extracellular

matrix (ECM) for proper growth of engineered tissue [306]. The increase in

nano-roughness can alter the surface energy, and then increase protein

absorption and bioactivity. Subsequently, the cells attachment, adhesion,

proliferation and long term extracellular matrix will be enhanced. Cell response

to nano-patterned surface has been reviewed [306-307]. Cell performance on

the nano- or micro-structured surface depends on cell types, because different

cell lines might favour different surface roughness. For example, endothelial

cells and osteoblast cells preferentially adhere onto both nano-structured and

sub-micron structured titanium surfaces rather than smooth titanium surfaces

[308]. Macrophage-like cells respond to smaller surface features 30 nm in

height, and endothelial cells react to larger surface features such as those with a

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height greater than 100 nm [309-310]. More recently, it was proposed that

disordered nano-patterns might increase cellular adhesion better, and

osteoblastic/osteospecific differentiations as well in contrast to ordered nano-

patterns [311-313].

The techniques for modification of surface roughness can be classified into

three categories: physical method, chemical method and their combinations.

The surface can be featured as ordered or disordered patterns. The nano-scale

features can be nano-rods, nano-ridges, nano-wires, nano-pits, nano-dots, nano-

particles, nano-crystals, and nano-fibres (See in Fig. 2.18).

Fig.2.18. Various nano-structured surfaces, a) nanopits [314], b) nanoparticles [315], c) nanowires [316], d) nanogrooves, e) nanoridges, f) sharp-tip nanoridges, g) nanocrystals [317], h) nanorods, and i) nanofibres [318].

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Plasma etching: high energy plasma can etch the substrate surface, causing the

formation of rough surfaces. It was used to produce disordered nano features

on scaffold surface many years ago. For example, atmospheric pressure plasma

treatment on PLLA can induce formation of nano-patterned surface for cell

growth (Fig. 2.19) [13]. Plasma etching method is mainly suitable for thin

films or membranes, because the plasma cannot penetrate into the thick

scaffolds [319-320]. For this reason, it is not quite suitable for surface

modification on 3D scaffolds.

Fig.2.19. PLLA film treated with different plasma voltages, a) plain surface; b), c) and d) plasma treatment to form 70 nm to 100 nm nano-featured surfaces. [13]

Template method: The procedure of using template to make a nano/micro-

structured surface is very simple, cost effective and there is no need for

expensive facility. This method was widely used in tissue engineering (Fig.

2.20) [321]. Silicon and AAO are good template candidates due to their porous

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structure, but they are not suitable for 3D porous scaffold surface

modifications.

Fig.2.20. Scheme of template methods, a) Overturning method to form nanopits, b) directly deposition method to form concaves [321]

Lithography: includes photo-nanolithography, colloidal lithography,

nanoimprint lithography, electro-beam lithography, focused-ion beam

lithography, dip-pen lithography, transfer lithography and interference

lithography. This series of methods are modern tool to modify surface. Fig.

2.21 illustrates one of the novel lithography systems, reflective electron beam

lithography. It can greatly improve the throughout per hour [322]. Photo-

nanolithography needs photo-masks and the polymer should be photo-active,

which narrows the selection of polymers [323]. Ion implantation method can

introduce a severe compositional and chemical modification to the polymer

surface. But the shallow penetration depth and the relatively high machine cost

prevent its wide use in 3D polymer scaffolds [324-325]. Other lithography face

similar challenges: it requires costly special equipments, and the processing

procedure is time-consuming. This treatment is limited to small area.

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Fig.2.21. Diagram of REBL nano-writer concept [322]

Electrospinning: can generate porous fibrous structure for better cell growth.

The weak points of this technique are that it is not able to provide thick 3D

structure, and the surface nano-features cannot be controlled exactly [17].

Although novel fabrication method can make electrospun mat form 3D tissue

scaffolds, the nanofibre mat is actually situated between the micro-pores, but

not on the surface of individual pore [186].

Wet chemistry methods including chemical etching, polymer demixing, LBL

coating, sputter coating and sandblasting may solve the problem of poor

penetration. It also can modify surfaces throughout the whole scaffolds [326-

327].

Chemical etching: is the basic method to get nano-textured surface. Acid or

base etching method is usually used in metal, ceramic scaffold [328-329], and

sometimes polymer surfaces [330]. For example, PU and PLGA were etched

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by HNO3 and NaOH separately to get nano-structured surfaces (Fig.2.22) [12].

However, strong acid and base are not suitable for polymer-related scaffolds,

because they might de-composite polymer and influence the mechanical

property of scaffolds.

Fig.2.22. Chemical etching to get submicro-or nano-structured surfaces [308]

Polymer demixing method: for nano-patterned surfaces relies on the

spontaneous phase separation of polymer blends (uniform in one solvent).

Polymer demixing can get island structure having the exact same component as

the body polymer, and this method can also control the island height from 10-

100 nm (Fig. 2.23) [331].

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Fig.2.23. AFM images of islands formed using different polymer blends, a) 20% w/w poly(p-bromostyrene): 80% w/w poly(styrene), b) 60% w/w poly(p-bromostyrene): 40% w/w poly(styrene,c) 80% w/w poly(p-bromostyrene): 20% w/w poly(styrene), d) AFM image of 45 nm high PS/PBrS islands,e) AFM image of 50 nm high PS/PnBMA islands [331]

Layer by layer electsostatic self-assembly (ESA): is often used in biomedical

field for incorporation of specific biomacromolecules or bioactive molecules

onto scaffold surface, due to its ability to control the coating quantity and

coating thickness [332-333]. Layer by layer coating can be used to get nano-

patterned surface by coating nanoparticles or nanorods on the surface. Silica

nanoparticles have been used to coat fabrics to obtain buoyancy properties by

this method (Fig. 2.24) [334].

Fig.2.24. SEM images of untreated cotton fabric (a and b) and cotton fabrics assembled with (PAH/SiO2)n multilayers: (c) n = 1, (d) n = 3, (e) n = 5, and (f) n=7 [334].

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Issues on nano-patterned surfaces

1) It is often difficult to separate the influences between the surface

chemistry and the surface topography, because the surface chemistry

usually changes with the modification of surface topography.

2) Nanoparticles modified surface may induce allergic or inflammatory

reactions or influence hemolysis and blood coagulation [335-338].

3) Whist many methods are available for achieving nano-surface features,

achieving fully ordered and controlled surface structures for 3D

scaffolds remains a challenge.

2.4.2 Bioactivity improvement for bone tissue engineering The natural bone is a composite consisting of inorganic apatite nanocrystals

and collagen fibres, and the inorganic apatite nanocrystals deposit on collagen

fibres to form a three dimensional structure [339]. Apatite is proved to be

excellent in biocompatibility, bioactivity and osteoconductivity [340]. It can

reduce the interface energy between human bone mineral and the surface

apatite layer on scaffold when implantation. Considerable efforts have been

spent on enhancing bioactivity by coating with bone-like apatite, calcium

phosphate ceramic or hydroxyapatite on the surface of metal or polymer

through plasma spraying [341], ion sputtering [342], laser deposition [343],

sol-gel deposition [340], and dip coating sintering [344]. They usually required

high temperature or specific pH condition. The prepared apatite is not similar

to that in human body, and lacks bioactivity. It was believed that via soaking

in SBF solution, the apatite formed a low crystallinity and nano-crystal size,

which is more similar to natural bone and easy to degrade [345]. Fig. 2.25

shows various CaP surface structures [346]. Usually the polymer surface was

modified to have functional groups like carboxylic acid, hydroxyl groups, Si-

OH, Ti-OH, Zr-OH, Nb-OH and Ta-OH which were effective to induce apatite

nucleation by capturing calcium and phosphate ions from SBF solution [347-

348]. With time, bone-like apatite layer was easily formed on polymer surface.

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Fig.2.25. Apatite coatings with different morphologies, a) and b) apatite precursor spheres, c) and d) plate-like apatite, e) and f) small plate-like apatite formed, g) and h) conventional apatite formed after immersing in SBF for 1 and 14 days, respectively.[346]

The common problem for those methods is that the deposition is time-

consuming. For example it took 60 days to form apatite on PCL film surface.

So to speed up the apatite formation on polymer surface, many methods have

been tried.

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1) Pre-dipping in Ca2+ and PO43- solution to achieve nucleation of CaP on the

surface, and then soaking in SBF solution. These procedures were repeated

several times for better nucleation [349].

2) Surface was functionalised with special groups for effectively controlling

the nucleation procedure. NaOH treatment of PCL was used to induce the

formation of –COOH groups for speeding up the nucleation procedure [350].

NaOH is a strong base, which may cause decomposition of polymer bone

scaffold.

Plasma technology was used to modify the surface chemical properties in an

environment-friendly way. It could offer polymer surface with active groups

such as carboxylic acid, hydroxyl groups by simple atmosphere treatment

[319]. O2 or argon plasma pre-treatment to introduce functional groups on

polymer surface has exhibited significant improvement in CaP formation on

polymer surface [351-352].

3) Soaking in high concentration SBF

Soaking can be done in a solution of high SBF concentration, for example

using 2×, 3×, 5×, or 10× times of SBF. The 10×SBF is very unstable, and hard

to prepare, CO2 gas was used to improve the stability of 10×SBF [353-354].

Another problem for the high concentration SBF is that apatite formed from a

high concentration SBF may have different component to that of human bone.

This may affect the vivo experiment or implantation [355].

Among these methods, the nucleation induced by surface functionalisation is

the most effective, and the time for the formation of a sufficient CaP layer can

be shortened significantly.

2.4.3 Antibacterial scaffolds

The main failure of tissue implantation is caused by infection [356], like many

cases from dental implantation [357-358] and in vitro fertilization cycles [359].

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Such an issue has not received much attention, presumably because most of the

studies are focused on how to improve cell attachment, growth and even

proliferation.

The usually used disinfection method includes UV irradiation, autoclave,

dry/wet heat, ionizing radiation or 75% ethanol disinfection, which is either

time-consuming or not cost-effective. More importantly, they have limitation

in certain scaffolds and might decrease the mechanical properties. Alcohol

sterilisation is not always effective for bacteria. After the implantation, the

tissue scaffold in the body is also facing a great risk of infection. Moreover, the

tissue implanted often resists antibiotics resulting in failure of implantation

[360].

If scaffolds were coated with one antibacterial layer, it would be able to stop

bacterial infection. Actually, there are already many applications of

antibacterial coatings in the biomedical area, such as valves [361], pins [362],

vascular [363], catheter [364] and prostheses [365-366]. The antibacterial

agents that have been considered for biomedical uses range from metal ions,

metal compounds and organic agents. They can kill microorganisms but are

somewhat toxic too. One useful article defined a biocompatibility index to

compare the safety of antibacterial agents in biomedical use [367]. Results

indicated that silver nitrate (AgNO3) is the most toxic one to murine fibroblast,

followed by octenidine dihy- drochloride (OCT), silver (I) sulfadiazine (SSD),

benzalkonium chloride (BAC), cetylpyri- dinium chloride (CPC),

chlorhexidine digluconate (CHX), triclosan (TRI), polyhexamethylene

biguanide (PHMB), mild silver protein (MSP), iodine in solution [PVP-I(s)]

and povidone iodine in ointment [PVP-I(o)]. When taking the antibacterial

properties into account, the OCT and PHMB are the best candidates for cell-

related antibacterial agents.

Nano-Silver:Silver has been recognised for its excellent antibacterial activity

since ancient time against a broad range of microorganisms, over 650 disease-

causing organisms in the body, even at low concentrations [368]. More

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recently, it was identified to interrupt HIV cellular replication [369]. With the

development of nanotechnology, Ag is favoured due to its small size and large

surface to volume ratio. This can increase the contact chance to microbes and

then vastly improve its antimicrobial activity [370]. It differs from bulk

materials in many aspects, like possessing high extinction coefficient, high

surface Plasmon resonance and anti-microbial properties that are less toxic

than bulk form [371-372]. It also has been proved to kill several pathogenic

bacteria that have resistance to various antibiotics [373]. Silver has been

applied as biocide widely in daily life, such as bandage [374], food packing

[375], water or air treatment [376-377]. In biomedical area, Ag based wound

dressing [372, 378] and Ag coated medical devices have been extensively used

as well [362-363, 365].

To date, the mechanism of Ag killing bacteria is explained as: 1) Ag caused

leakage of the ROS outside the cell leading to apoptosis [186]; 2) Ag interacted

with thiol groups to prevent the cell proteins from forming correctly and DNA

replication [379]; 3) Attachment to the cell surface, altering the function of the

cell membrane [380].The antibacterial property of Ag depends on the size,

shape and charge. For example, size ranging from 1 to 10 nm has better

bactericidal activity because of more chances to contact with bacteria [381].

Truncated triangular shaped is the best among other shapes like tubes, rods,

triangles, wires and aligned wires [380, 382].

Although many researchers have raised concern that Ag might have potentially

negative influence to health, some cell culture experiments have shown Ag

nanoparticles are cytotoxic to rat liver cell line [383], lung epithelial cells [384],

keratinocytes and fibroblast cell lines [385], and PC-12 neuroendocrine cell

line [386] and brain ultrastructrue [387]. Ag in most studies is suggested to be

non-toxic to mammalian systems [388-389]. In comparison to the surgery

failure due to the infection, Ag is much less toxic to human body. If only the

concentration, size and status for optimum antimicrobial activity is known, Ag

would be perfect for disinfection of scaffolds for implantation applications

[390].

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Till now, the Ag coating on bio-devices or fabrics for antibacterial function has

been through wet chemical and physical methods. Generally, it is to make the

silver salt to reduce into nAg particles. Most of the physical methods (e.g.

microwave irradiation [391], UV irradiation, γ-irradiation [392-393], ultrasonic

[394-395] and plasma coating) are best for film surface treatment and need

expensive devices. However, wet chemical modification provides easy,

convenient treatment to both film and 3D structure, and no expensive devices

are necessary. Sol-gel method has been widely applied in the textile industry.

A disadvantage of this method is that it is hard to get uniformly distributed Ag

nanoparticles throughout the whole fabric [396-397]. Layer by layer coating

method makes it possible to coat 3D matrix evenly, but it involves multi-step

and complicated treatment [398-399].

PHMB:Polyhexamethylene biguanide (PHMB), also known as polyhexanide

and polyaminopropyl biguanide, was firstly found to have good antibacterial

property in mid 1970s. As another very popular biocide,it is a mixture of

polymeric biguanides (Fig. 2.26, where n=2–40), with a mean size of n=11,

giving a molecular mass range of approximately 400–8000, with various

combinations of amino, guanide or cyanoguanide as end-groups. PHMB can

be synthesised using the following simple method. Firstly,

hexamethylenediamine reacts with sodium dicyandiamide to get

hexamethylene biscyanoguanidine. Secondly, the resultant products react neat

with hexamethylenediamine hydrochloride at 160-185 °C for 4 hours under

constant stirring to form oligomeric molecules of PHMB [400-401].

Fig.2.26. Structure of PHMB

PHMB is effective to a broad spectrum of gram-positive and gram-negative

bacteria, fungi, yeast, and even HIV-1 [402]. Many researches focus on the

effectiveness of PHMB on bacteria from ophthalmic area (Acanthamoeba

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polyphaga, A castellanii, and A hatchetti.), water systems (Legionella

pneumophila.), biofilms (E coli and S epidermidis), and fungal infection

(Trichophyton mentagrophytes).

PHMB has been found diverse applications. For example, it was used to

prepare antimicrobial fabrics in textile industry [403-404], food disinfection in

food industry [405], and in many wound dressing products (e.g. AMD). It was

used to avoid wound infection both in film and foam [406]. It was also widely

used for contact lens disinfection solution in ophthalmology [407]. In hospital,

it has been used as skin antiseptics clinically [408], commercial swimming

pool cleaning agents also contain PHMB components [409].

The action of PHMB on microbial is usually regarded as contact-killing.

PHMB can attach itself to the bacterial cell membrane, disrupting microbial to

kill the bacteria. PHMB interacts with the bacteria surface initially and then is

transferred to the cytoplasm and cytoplasmic membranes. Another explanation

is that the cationic biguanide functional group can react with negatively

charged bacterial outer membrane, causing the increase in fluidity and

permeability. As a result, the lipopolysaccharides are released from the outer

membrane, and organism is killed. Some researchers believe that PHMB can

cause the formation of the acidic phospholipids of the cytoplasmic membrane,

resulting in the organism death eventually.

Since PHMB has many active groups in the backbone, it is easy to bond

PHMB to cotton fibre by multiple cationic groups, providing the stability

between PHMB and cotton fibres. It is also easily applied to cotton fibre by a

simple pad-dry-cure procedure. The PHMB can be exhausted on the fibre

surface effectively. For synthetic fibre, self-crosslinkable resin and catalyst can

assist the reaction between PHMB and fibres. For wound dressing application,

PHMB can be introduced to the dressing fabric to protect against the

development of wound infection by decreasing the bacterial load in the

dressing and bacterial penetration through the dressing. PHMB is also

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incorporated into the foam dressing during processing while the dense open-

cell foam matrix develops.

When assessing the suitability of PHMB for use in the biomedical field, its

cytotoxicity to mammalian cells must be taken into consideration. It has a low

toxicity to eye cells in very low concentration (about 0.0001%), and low tissue

response and minimal patient discomfort [410]. When used for wound

dressing, it didn’t elicit any skin reactions. As to other cells such as HaCaT

cells, Human keratinocytes and fibroblasts [411-412], it showed very low and

nearly no toxicity.

Several researchers [210, 411] have studied the co-culture of bacterial and

human mammalian cells on biocide-containing matrices. Many exciting results

have been obtained. For example, PHMB has been proved to have ability to

inhibit HIV-1 virus bonding to human health cells [402]. PHMB treatment can

be useful for removing bacteria from cartilage if a defined predetermined

concentration is applied [413]. PHMB has also been tested for protecting

human keratinocytes from bacterial damages [411]. Nanosilver containing

PHBV scaffolds has been proved to effectively inhibit the proliferation of

Staphylococcus aureus (Gram-positive) and Klebsiella pneumoniae (Gram-

negative) bacteria, whereas the scaffolds do not show any in vitro cell

cytotoxicity [301]. But there are also some opposite examples. The nano-silver

contained ceramic scaffolds were proved to negatively affect osteoblast-like

cell proliferation, meanwhile it can kill 99.9% Staphylococcus aureus [414]

PHMB can result in severe damage to osteoblast cells hFOB [415].

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C H A P T E R T H R E E

Materials, Methods and Characterisation Technologies

The materials, preparation and characterisation techniques used for all the

experiments in this PhD study are listed in this chapter.

3.1. Materials

Poly(ε-caprolactone) (purity >99% Mn=80,000, Aldrich), anhydrous ethanol (

>99.5%, Aldrich), Tetraethoxysilane (TEOS), G418 solution, food-level

ground sugar (sucrose, local supermarket), Dulbecco’s phosphate buffer saline

powder (DPBS, Gibco), phenol-red-free media (RPMI, Invitrogen),

glutaraldehyde (2.1%) in phosphate-buffered saline (PBS 10 mM, pH 7.4)

(Whiteley medical, aidal plus), Osmium tetroxide (OsO4) (SPI suppliers, 2% in

water), tannic acid (Aldrich), MTS-kit bioassay (Promega), phenazine

methosulphate (PMS, Aldrich), sodium chloride (≥99.5%, Aldrich) and

phosphate-buffered saline (PBS 10 mM, pH 7.4) were used as received. Milli

Q water was always used in experiments.

3.2. Fabrication and surface modification of fibrous

scaffolds

PCL filaments were extruded on a Wayne Bantam single screw extrusion line

with a 12.5 mm screw diameter and 20:1 length to diameter ratio. The screw

used had a 4:1 compression ratio. The screw speed was set at 25 rpm. An 18-

filament die was used for the extrusion. This extrusion unit has a 5-zone heated

unit set at 93.3 ºC in the first zone ramping to 121 ºC in the final die zone. The

extruded fibres were drawn from the spinneret die at 100 metres per minute.

The PCL fibres were cut to 2 mm length prior to use.

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Sucrose sugar powders were sieved (sieves 48 mesh and 28 mesh, Cole-

Parmer) to narrow the size in the range between 300 μm and 600 μm. The

sieved sugar powders were then mixed with short PCL fibres in different ratios.

To improve the sugar-fibre adhesion, a small volume of deionized water (1 ml

per 10g of sugar-fibre mixture) was added during blending.

The mixture was then cast into a capped metallic mould and placed into a 62

ºC oven for 30 minutes to allow the fibres to bond with each other. After

cooling down to room temperature, the heated fibre-sugar blend was washed

extensively with deionized water to leach off the sugar. The matrices were then

dried at room temperature for 24 hours. The thickness of the matrices was

controlled to near 2 mm. In order to match the size of tissue culture wells, the

matrices were cut into a thin disk shape (diameter 1.4 cm).

Synthesis of silica nanoparticles

Silica nanoparticles were synthesised according to the Stober method [416].

Briefly, 4.5 ml TEOS were dropwise added in a mixed solution with 3 ml

water, 150 ml ethanol and 9 ml ammonia, and then stirred overnight, followed

by centrifuging at 6000 rpm for 30 mins to collect the particle pallets. Then the

nanoparticles could be washed twice with water and keep in distilled water for

further use.

Layer-by-layer electrostatic self-assembly of nanoparticles

Electrostatic layer-by-layer coating method was conducted according to the

following procedure: 1) PCL matrix was treated by 1M NaOH for 2h to form –

COOH on the surface, 2) then it was dipped into silica nanoparticle solution for

5 mins to get a thick silica coating, 3) following it, the samples were dipped

into 1% PAA solution to get a negative charge layer, 4) the samples were

dipped into silica solution to get the first bilayer treated samples. Between

every step, the samples were thoroughly washed in distilled water and allowed

getting rid of the extra components on the surfaces. Through repeat dipping

into silica and PAA solution, multilayer silica can be applied on the surface.

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With this method, the thickness of silica nanoparticles on the surface could be

controlled effectively and the surface nano-morphology could also be

manipulated easily.

Plasma treatment of PCL scaffolds

A lab-scale vacuum plasma system (Fig. 3.1) was used for plasma surface

treatments. The energy used was 20W. Both the front and back surfaces of

each specimen were treated for 3 min, separately.

Fig.3.1. a) Vacuum plasma surface treatment unit, b) the plasma treatment zone

Simulated Body Solution (SBF) soaking

1×SBF solution was prepared following Kokubo’s methods [417]. Briefly, the

SBF was prepared by dissolving NaCl, NaHCO3, KCl, K2HPO4•3H2O,

MgCl2•6H2O, 1M HCl, CaCl2, and Na2SO4 in ultra-pure water and buffering to

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pH 7.25 at 36.5 °C with tris(hydroxymethyl) aminomethane (final

concentration of 50mM) and aqueous 1M HCl solution. The final solutions

were filtered to eliminate impurities and stored at 4 ºC for several weeks

without precipitation.

For 10 times concentrated SBF (10×SBF), the preparation was according to

document [418]. NaCl, KCl, CaCl2•2H2O, MgCl2•6H2O, Na2HPO4 were

completely dissolved in MilliQ water in the given order at room temperature.

The solution was then adjusted to pH near 4.35 and kept at room temperature

for several months. NaHCO3 was added in to improve pH near 6.50 just before

coating samples. The calculated ion concentrations of 1×SBF and 10×SBF

solutions in comparison with those in human blood plasma (HBP) are listed in

Table 3.1.

Table 3.1. Ion Concentrations of the SBFs and Human Blood Plasma in Total and Dissociated Amounts

Concentration/mM

Ion Blood

Plasma 1×SBF 10×SBF

Na+ 142.0 142.0 1020

K+ 5.0 5.0 5.0

Mg 2+ 1.5 1.5 5.0

Ca2+ 2.5 2.5 25

Cl- 103.0 147.8 1065

HCO3- 27.0 4.2 10

HPO42- 1.0 1.0 10

SO42- 0.5 0.5 N/A

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Samples after the vacuum plasma treatment were immediately immersed into

1×SBF or 10×SBF at physiological pH of 7.4 and temperature of 37 °C for a

certain time to deposit one layer of apatite. After that, the matrix disks were

thoroughly washed by MilliQ water and dried overnight at room temperature

for further use.

Electroless plating

Electroless plating was chosen to coat nano-Ag to the surface of 3D fibrous

scaffolds. Two-part electroless Ag plating was prepared as follows:

Solution A: 30g of NaOH were added to an AgNO3 aqueous solution

containing 70g AgNO3 and 1200 ml water. Precipitation occurred instantly

during the addition of the NaOH. Concentrated ammonia (35 wt %) was

dropped into the Ag solution until it became clear. 5mg of L-cystine was then

added into this solution to give the Solution A.

Solution B: 90 g of D-glucose and 8 g of citric acid were dissolved into 200 ml

ethanol and 2L water to give the Solution B.

The cut circular discs were put into reaction beaker, and the two solutions were

added into it together for a certain time (from 10 mins to 40 mins were

examined in this study)

Layer-by-layer coating of PHMB on 3D fibrous scaffolds

3D PCL fibrous matrices were firstly treated with 1M NaOH for 2h to induce

the formation of –COOH groups on the surface, and then they were dipped

into 1% PHMB solution (pH 8.0) to get a full negatively charged surface,

following which was a continuous cycle of dipping between 1% PAA solution

and 1% PHMB solution, both at pH 8.0. It was marked as 1 layer PHMB for

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the first dipping, and then 1bilayer PHMB for the first cycle, 3 bilayers PHMB

for 3 cycles, and so on.

3.3. Cell culture related technologies

3.3.1 Sterilisation

All the samples in this thesis used for biological experiments were sterilised by

rinsing with 70% ethanol, saturated sodium chloride aqueous solution and

MilliQ water (R=18.2MΩ). They were then placed individually in the wells of

tissue culture plates and washed with sterile 70% ethanol for 30 min ~1.5 hr on

an orbital shaker. The 3D scaffolds were rinsed with fresh 70% ethanol for 5

minutes, dried in biohazard hood, and kept in clean and dry containers for

further uses.

3.3.2 Cell culture media

The culture media for fibroblasts was prepared as mixture of 500 ml fibroblast

media (RPMI + Glutamax) (Invitrogen) supplemented with 40 ml foetal bovine

serum (FBS, Bovogen), 5 ml of penicillin and streptomycin solutions (10000

U/ml penicillin, 10000 g/ml streptomycin, Sigma-Aldrich) and 2 ml amphostat

B (Thermo Scientific).

The culture media for CHO cells was prepared as mixture of 500 ml of CHO

media (GIBCO + Glutamax) (Invitrogen) supplemented with 40 ml foetal

bovine serum (FBS, Bovogen), 5 ml of penicillin and streptomycin solutions

(10000 U/ml penicillin, 10000 g/ml streptomycin, Sigma-Aldrich) and 2 ml

amphostat B (Thermo Scientific).

The culture media for Saos-2 cells was prepared as mixture of 500 ml of Saos-

2 media (MEM+Glutamax TM-1) (Invitrogen) supplemented with 50 ml foetal

bovine serum (FBS, Bovogen), 5 ml of penicillin and streptomycin solutions

(10000 U/ml penicillin, 10000 g/ml streptomycin, Sigma-Aldrich) and 5 ml

non-essential amino acid solution (NEAA) (Sigma-Aldrich).

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The culture media for Hfob 1.19 cells were prepared as mixture of 500 ml of

hFOB media (DMEM: F-12 Medium without phenol red) (Invitrogen)

supplemented with 50 ml foetal bovine serum (FBS, Bovogen), 150 mg of

G418 (Sigma-Aldrich).

3.3.3 Cell lines

Rat skin fibroblasts, Chinese hamster ovary (CHO) cells and Saos-2 osteoblast

like cells were donated by Barwon Health hospital in Australia.

Human Osteoblast cells (Cell line number CRL-11372) were purchased from

American Type Culture Collection (ATCC).

3.3.4 Subculture of cells

The obtained rat skin fibroblast and CHO cells were subcultured firstly in 37

°C incubators to get a steady growth for further cell culture study usage. Fig.

3.2 indicated the healthy state of fibroblasts and CHO cells. For fibroblasts,

when growing on the flat tissue culture plates (TCP), they maintained a flat

multipolygon shape. In contrast, CHO cells were spindle shaped and also grew

nicely on the TCP surface.

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Fig.3.2. Cell morphology of fibroblasts (a, and a’) on TCP under low and high magnification at 37ºC and CHO (b) on coverslip.

The original hFOB1.19 and Saos-2 cells were subcultured in 34 °C and 37 °C

incubators separately, and they were ensured to grow healthily before doing

any experiment further. The morphologies of hFOB 1.19 and Saos-2 cells used

were shown in Fig. 3.3. From the lower and the higher magnification of

images, the cells attach to the TCP surface nicely and grow very well.

Fig.3.3. Cell morphology of hFOB 1.19 (a, a’ and a”) and Saos-2 (b, b’ and b”) on TCP under low and high magnification at 34 ºC and 37 ºC

3.3.5 Cell seeding

In this thesis, different seeding methods were utilised.

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1) Static seeding method: Briefly, 50 μl cell/media suspension was dropped

onto the scaffold surface as uniformly as possible using a pipette. After 1.5 hr,

other 950 μl media was topped up. On the second day of seeding, the matrices

were washed with PBS and transferred to a new culture well. The number of

cells attached on the matrices also was calculated by MTS assay described

later.

The seeding efficiency was calculated according to the equation (3.1):

(3.1)

2) Dynamic seeding method using spinner flask: A certain volume of media

was put into the spinner flasks, and it was magnetically stirred under a certain

speed for a pre-set time. In detail, 110 ml of seeding media was put into the

spinner flasks, and the speed of spinner flasks was set in the range between 50

rpm and 80 rpm. A seeding density of 1×106 cells/ml was used.

3) Dynamic seeding method using small rotary vial: A certain volume of media

was put into the small rotary vial, and it was put into the blood roller machine

to rotate overnight to seed the cell uniformly onto the scaffolds. The seeding

volume was only 2 ml, the rotating speed was between 23 rpm to 52 rpm, and

the seeding density was 5×104 cells/matrix.

4) Dynamic seeding method using orbital shaker: Different seeding volumes

ranging from 50, 100, 200, 500, to 1000μl were put into the well of tissue

culture plates, and it was put onto orbital shaker to shake overnight at speed of

0, 50, 75, 100, or 125 rpm. The seeding density was controlled at 5×104

cells/matrix.

3.3.6 Cell culture

All the samples after cell seeding experienced static cell culture. They were put

into the standard incubator (37 ºC, 5% CO2). In the culture period, the media

were first replenished after the first seven days and then every three days.

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3.3.7 MTS assay

MTS assay is a straightforward and simple method to detect the cell

proliferation. The principle of the assay is that MTS tetrazolium compound is

reduced in cells into a coloured formazan compound that is soluble in tissue

culture media. This conversion occurs via dehydrogenase enzymes that are

present in live cells. The colour then could be read in density to decide the live

cell number [419].

During the culture period, matrices at different time points were chosen for

MTS assay. The matrices were rinsed with PBS, to remove non-viable cells,

and were transferred to new wells with 0.5 ml phenol-red-free RPMI media

(Invitrogen). The original seeded well was also topped up with 0.5 ml phenol-

red-free media. Wells and matrices were let to react with 50 μl of 5 mg/ml

MTS solution in phosphate buffered saline (PBS, Sigma Aldrich) for 4 hrs in

37 C incubator. As colour-development solution, 375 μl phenol red free media

and 125 μl MTS and PMS in PBS solution (0.046 mg/ml PMS and 2 mg/ml

MTS) were added. Samples were incubated in a 37 C incubator for 1 hr, after

which 100 μl of the solution was withdrawn to measure the absorption at 490

nm on a microplate reader within the hour.

MTS standard curves were obtained in a similar way by using a series of

samples containing a known number of viable cells. The absorption was

adjusted by subtracting the readings of the zero wells from all of the other

readings. The cell numbers from the wells, the matrix or the sum of both were

calculated based on the standard curve. All the cell culture related results were

collected in triplicate and were presented in form of average and standard

deviation.

3.3.8 Cytotoxicity test

A typical cytotoxicity test was implemented for 3 days [DS/EN ISO10993-5].

Sample scaffolds were seeded with rat skin fibroblasts (donated by Geelong

Hospital, VIC, Australia) via static seeding into a 24-well tissue culture plates

(Greiner, Interpath), using a calculated density of 1 × 104 cells/well. Cells were

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cultivated in 1 ml of fibroblast media under standard culture conditions (i.e. at

37 ºC, in a humidified atmosphere containing 5% CO2 and 95% air). The cell-

seeded scaffolds were maintained for 4 hrs in a 37 °C incubator prior to the

addition of 1 ml culture medium into each well. The samples were then

cultivated at 37 °C, in a humidified atmosphere containing 5% CO2 and 95%

air. The cell number was determined by haemocytometer using a Trypan blue

dye (Sigma-Aldrich) to exclude nonviable cells from the count.

3.3.8 Stain of samples for microscopic observation

After cell culture, cells on matrices were collected from the culture plates and

washed three times with sterilized PBS, then immersed in a 2% para-

formaldehyde in PBS for 10 minutes at room temperature. After rinsing with

PBS for three times, cells were stained with 1% toluidine blue in 2% borax and

0.04% methylene blue. The stained samples now were ready for observation

under microscope.

3.3.9 Stain of samples for SEM observation

Before SEM observation, cells on the scaffold samples were fixed by

immersing the cell-carrying scaffolds in 2.1% wt/vol glutaraldehyde in

phosphate-buffered saline (PBS 10mM, pH 7.4) for 4 hrs at room temperature.

After rinsing with sequential dilutions of PBS in deionized water, the matrices

were immersed in a 2% wt/vol solution of OsO4 in water for 20 minutes. The

scaffolds were then immersed in a 1% wt/vol aqueous solution of tannic acid

for 10 minutes, and then rinsed with 20% ethanol/water solution. The scaffolds

were then freeze-dried in a freeze-drier (Labconco Freezone 2.0)

3.3.10 Laser scanning confocal microscopy

Laser scanning confocal microscopy (LSCM, Leica TCS SP5, Germany)

equipped with Argon (458nm, 476nm, 488nm, 496nm and 514nm) and 405

Diode (405 nm) lasers was used to observed all the cell-contained samples.

After cell culture, the cell-carrying matrices were collected from the culture

plates and washed three times with sterilised PBS to remove medium and

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unviable cells. The matrices were then immersed in a 2% para-formaldehyde in

PBS for 10 minutes at room temperature. After rinsing with PBS for three

times, a permeabilisation solution (0.2% Triton X-100 in PBS, Sigma-Aldrich)

was applied for 10 minutes at room temperature, and the matrices were rinsed

again with fresh PBS for three times. Staining with DAPI (dilution ratio of

1:100, Invitrogen) and Phalloidin Alexa 568 (dilution ratio of 1:100,

invitrogen) was performed overnight at 4 °C in dark. After rinsing with PBS

for three times to remove residuary fluorescent dye, the matrices were imaged

under the confocal microscope (Fig. 3.4). The excitation wavelengths for DAPI

and Phalloidin are 358 nm and 578 nm, respectively. The emission

wavelengths for DAPI and Phalloidin are 461 nm and 600 nm, respectively. A

series of slices were obtained as well by focusing different depth at certain step

throughout thickness in the Z-axis.

Fig.3.4. Confocal microscope unit

3.3.11 Alkaline Phosphatase (ALP) assay

ALP assay can quantitatively determine the alkaline phosphatase formed in

cells. The principle of this method is that alkaline phosphatase catalyses the

hydrolysis of p-nitrophenyl phosphate and liberating p-nitrophenol and

phosphate. The rate of p-nitrophenol formation could be measured

photometrically, which is proportional to the catalytic concentration of alkaline

phosphatase [420].

To determine the amount of ALP produced by the cells seeded on the

scaffolds, scaffolds/cells constructs were washed with PBS for three times,

moved to a new well, and then lysated by 1% Triton-100 under 37 °C for 1 hr

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after pre-decided culture time. p-Nitrophenyl phosphate (pNPP) was added to

the supernatant and incubated at room temperature for 1 hr. The enzyme

reaction (conversion from p-nitrophenyl phosphate (pNPP) into p-nitrophenol

(pNP)) was then stopped by a solution containing 2 M NaOH in distilled water.

The absorption of p-nitrophenol (pNP) formed was determined at 410 nm. A

standard curve was made using pNP values ranging from 0 to 300 U/l. The

results were expressed as Units of pNP produced/l/hr. Each experiment was

done in triplicate.

3.3.12 Alkaline Phosphatase (ALP) staining for microscopy

observation

After a certain culture time, cell/scaffold composites were flushed with PBS

and moved to a new well. They were then stained with 0.5% (v/v) Naphthol

AS-MX Phosphate/N,N Dimetylformamide (DMF) solution and 0.6 mg/ml

diazonium salt (fast red violet LB salts) in tris-hydrochloric acid buffer (pH

8.74) and water solution at 37 °C for 30 minutes. The outcome of this is that

enzyme activities were stain in colour from pink to red. After taking images,

the cell/scaffold composites were rinsed in distilled water for three times, and

Mayer’s hematoxylin was utilised to stain nucleus of hFOB cells for 1 minute,

then taking images again to get figures with both nuclei and ALP stain.

3.3.13Statistical analysis

All experiments were conducted in triplicates and data are expressed as mean ±

standard deviation. The cytotoxicity, MTS and ALP data were analysed on

SPSS (SPSS statistics 17.0) by using ANOVA and post hoc multiple

comparison tests. A p-value of less than 0.05 was considered statistically

significant.

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3.4 Bacterium related technologies

3.4.1 Bacterial type

Escherichia coli bacteria were purchased from American Type Culture

Collection ( ATCC,Cat. No.35218).

3.4.2 Bacterial culture media

Modified tryptone soya broth (CM 0989B, approximate formular: Pancreatic

digest of casein, 17 g/L, Papaic digest of soybean meal,3 g/L) and tryptone

soya agars (CM 0131B, Pancreatic digest of casein, 15 g/L, Enzymatic digest

of soya bean, 5 g/L) were obtained from Oxoid Australia Pty Ltd.

The culture media were prepared following the instruction of ATCC. For E.

Coli: 16.5 g nutrient broth powders were dissolved in 500 ml distilled water by

gently warming and stirring in a 1 L capped bottle. The broth was sterilised by

autoclaving at 121 °C for 15 min. Then it was cooled down and kept in an

incubator at 37 °C with constant shaking for further use. The culture agar for E.

coli was prepared similarly instead by using 20 g agar powder in 500 ml

distilled water. After cooling down to near 50 °C, the agar was mixed by hand

shaking, and then directly poured into the plain plate to get agar plate with

almost equal agar thickness (2.5mm). It took 25-35 min to solidify the agar

medium. These agar plates can be kept stably in 4 °C for further use.

3.4.3 Antibacterial testing

The antibacterial assessment was tested based on a standard methodology: 1)

static direct contact method and 2) dynamic shaking flask method.

1) Static direct contact method. This was used to observe the antibacterial

property of samples directly by adding bacterial solution to sample surface.

Prepare the bacterial working solution by measuring the overnight bacterial

culture at 550 nm after 10× dilution in nutrient broth, calculating the

concentration of the culture according the OD value and diluting the solution

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into 4 × 107 cells/mL. Seeding 107 cells onto the four stacked samples, and

these samples with cells were kept in a sterile bottle for a certain time (e.g. 30

min, 1 hr, 3hr, 5hr, 7hr, 24hr). After that, distilled water was added into the

bottle and these bottles were vigorously shaken to elute bacteria. Plates were

prepared for a series dilution of the elution and cultured at 37 ºC overnight.

The second day, the number of colonies on the plates were counted and

calculated for appropriate dilutions. In the experiments, samples without any

anti-bacterial treatments were used as control. Then the reduction percentage of

bacterial can be calculated by the equation below.

%100(%)A

ABreduction (3.2)

Where A is the colonies after incubation on the untreated samples, B is the

colonies after incubation on the anti-bacterial treatment samples. If B>A, no

reduction happens.

The entire handling and testing procedure was done in front of a flame of a

spirit lamp.

Fig.3.5. Flow chart indicating static anti-bacterial study

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2) Dynamic shaking flask method. This can well mimic the possible

environment in future human body. Samples were in direct contact with

bacteria solution under vigorous shaking for a desired time. Briefly, working

solution was prepared by culture of E. coli from agar plates in sterilised

nutrient broth for one day. Then the Ag-PCL matrices were introduced into 8

mL of working solution in sterile tubes, which contains about 1.5x106 colony

forming units (CFU) of E. coli. The mixtures were cultured at 37 °C in a

shaking incubator (HD Scientific Suplies Pty Ltd) at 200 rpm for 1~8 hrs

separately. Pure sterile broth with non-treated PCL matrix was also tested as a

blank control. After the culture, the samples were vortexed for 1 min and

diluted serially in sterile distilled water. 50 μl of each dilution was seeded onto

LB agar using a surface spread plate technique duplicate. These plates were

incubated at 37 °C static for 24 hrs. Then the numbers of bacterial colonies

forming units (CFU) were counted. The counts were used to calculate the

surviving number of bacteria. The antibacterial efficacy (Ab%) of the specimen

was calculated according to the following equation:

%100)1((%)original

survive

CFUCFU

Ab (3.3)

Where CFUsurvive represents the colonies recovered from the samples; and

CFUoriginal is colonies recovered from original working solution. The entire

handling and testing procedure was done in front of a flame of a spirit lamp.

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Fig.3.6. Flow chart showing dynamic anti-bacterial study

3.5 Characterisation tools

3.5.1 Photos

Matrix disks and as-prepared other samples were imaged under a digital

microscopy (Dino-lite plus, AnMo Electronics, China).

3.5.2 Matrix thickness measurement

The matrix thickness was measured under a load of 0.5N by digital MESDAN

micrometer showed in Fig. 3. 7.

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Fig.3.7. MESDAN micrometer for fabric thickness measurement

3.5.3 Scanning electron microscopy

Scanning electron microscopy was carried out with a Supra 55VP SEM (Carl

Zeiss, Germany) microscope. It was used to investigate the surface

morphology of PCL fibrous matrices, the surface treated-samples, and cell-

bearing samples. Samples were cut into small circle (1.4cm in diameter) and

placed on SEM holders. Pictures were taken after sputter coating with gold

using a Bal-tec SCD50 sputter coater at 40 mA for 60-90 seconds. All images

were taken at a voltage of 10 keV and the magnification varied between 100

times up to 10000 times.

Electron microscopic images and energy-dispersive X-ray (EDX) mapping

were taken on a scanning electron microscope (SEM) Leica S440 and analysed

by ISIS software respectively.

3.5.4 Micro-computed tomography

The Micro XCT XRadia machine (XRadia Ine. USA) was used to scan samples

to get X-ray micro-computed tomography images. It has a cone beam geometry

that permits magnification regulation by adjusting the ratio of the distance

between the X-ray source and the sample (Z1) to the distance between the X-

ray source and the detector (Z2). At the same time the magnification also can

be controlled by magnification lens which are limited from 0.5 to 40. The

limiting resolution is between 1 and 30 m. Detector system: CCD camera

(Andor) combined with scintillara system to transfer X-ray to visible light. The

camera is capable of acquiring X-ray images of 2048 2048 pixels with a

depth of 16 bits per pixel.

The scan procotol of this experiment included: X-ray source voltage was

standerised at 40 kV and beam current at 150 A from tungshen as a target

material. The distances were set to 125 mm for Z1 and 20 mm for Z2. The

magnification factor from lens was 20. The samples were scanned at a

rotation step of 0.25 degrees over 180 degrees. It took about 13 hours for every

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sample to get a 10243 voxel tomogram of 1.0 G data and capture 721

projections of 6.0 G data with a spatial resolution of 1.155 m and exposure

time of 60s/projection.

3.5.5 Processing of the micro-CT scanned series images

3D reconstruction was performed using Mimics software (Materilise,

Belgium). 100 sequential 200 × 200 pixel images were cropped from the serial

images from the centre of each sample. Imported into mimics, these serial core

images were reconstructed into 3D volumetric models. Thresholds were

inverted to allow measurement of pore characteristics within the model.

Subsequently, a region growing operation was performed to create a mask

consisting only of interconnected pore spaces. Volume for this region-grown

mask was determined and the ratio of region grown volume to the total volume

was calculated. The percentage of this ratio is defined as the degree of

interconnectivity.

3.5.6 Mechanical property

Uniaxial tensile testing of fibrous matrix was carried out at the ambient

temperature with reference of ASTM D638-98. The quadrate strip sample with

a size of 50 10 2mm was cut from fibrous mat and tested on an Tensile

Tester (LR30K, Lloyd Instruments, UK) equipped with serrated vise grips and

a 100 N load cell at a crosshead speed of 10 cm/min, gauge of 3 cm and no pre-

load. Maximum strength was determined as the strength values corresponding

to 50% strain if, at this stage, the samples were not damaged. The thickness of

every sample was pre-tested by a micrometer under pressure of 0.5 N.

Measurements were done three times for one matrix in exactly the same setting

to avoid errors from sample slippage and to get the average value and the

standard deviation. Compression strength was tested with the same machine

with sample in diameter of 37.55 mm.

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Fig.3.8. Mechanical testing machine

3.5.7 Water contact angle

The water contact angles were measured using a KSV Model CAM101 Contact

Angle Meter (KSV Instruments Ltd, Finland, Fig. 3.9). The instrument

includes a video camera, a LED light source and electronics interface for the

unit. Samples were cut into 1.4 cm diameter circle disk and stick on a glass

slide. Then liquid drops of 1.20 mm diameter were placed on each sample

surface with a syringe (diameter: 0.71 mm) and the image of each drop was

captured at a rate of 1 per second after deposition onto sample surface and until

the water drop penetrated to the sample. The final angle was the average of five

testing results.

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Fig.3.9. CAM101 KSV contact angle meter

3.5.8 Water hydraulic permeability

Hydraulic permeability was measured by a simple lab setup (Fig. 3.10), and the

hydraulic permeability of the porous materials were calculated according to

Darcy’s law [421]

Px

AQ

k (3.4)

Where k is the hydraulic permeability of the porous material, A is the flushing

area of the porous scaffold, Δx is the thickness of the scaffold, ΔP is the

pressure difference driving the fluid flow across the scaffold.

Fig.3.10. Setup for water hydraulic permeability testing

3.5.9 Water binding ability

The water binding ability of matrices was measured by weighing the dry mass

and wet mass after immersing in 5 ml water for 24 hrs at room temperature.

The binding ability was calculated by equation (3.5).

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%100(%)wet

wetdry

Mass

MassMassabilitybindingWater

(3.5)

3.5.10 Fourier transform infrared spectroscopy

Fourier transform infrared (FTIR) spectra were recorded on FTIR

spectrophotometer (Bruker Vertex 70, Bruker Optik GmbH, Ettlingen,

Germany) coupled to a Golden Gate single reflection ATR unit for PCL matrix

and PCL surface treated samples. Routinely, background due to air was

subtracted. The absorption spectra of the surfaces were obtained for each

specimen and a total of 32 scans were accumulated at a resolution of 4 cm-1.

All the tests were carried out in a controlled environment, at 20 ± 2 ºC

temperature and 65 ± 2% relative humidity. The data was analysed using

OPUS 5.5 software.

Fig.3.11. Vertex 70 FTIR spectrophotometer

3.5.11 Atomic force microscopy

Atomic force microscope (AFM) images were taken by a DualScope DS 45-

40scanner (Danish Micro Engineering, Denmark) with a scan size of 5 × 5 μm.

Applied load during testing was 0.15 nN. The scanning speeds on nanofibre

membranes and polyester fibres were 100 and 5μm/s, respectively. The surface

roughness was calculated using DualScope software.

3.5.12 Dynamic light scattering

Dynamic light scattering (DLS, Zetasizer 3000, Malvern Instruments) was

utilised to measure the particle size distribution and zeta-potential of

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nanoparticles. The Ag-PCL samples were dissolved into dichloromethane, and

then tested. By measuring the intensity of light scattered from the samples, the

particle size distribution was obtained.

Fig.3.12. Malvern Zetasizer 3000 dynamic light scattering

3.5.13 Transmission electron microscopy

Transmission electron microscopy (TEM, JEOL JEM-2100) was used to

confirm component layer formed outside the surface of PCL fibres after

surface treatment. TEM samples were prepared by embedding the sample into

epoxy resin followed by drying in a vacuum oven at 40 °C. The samples were

then microtomed using a diamond knife on ultramictrotome machine. The thin

slices (90 nm in thickness) were placed on 400 mesh copper grids for TEM

observation. Images were obtained at an acceleration voltage of 200 kV.

3.5.14 UV-VIS spectroscopy

Absorption curves were measured on a UV-visible spectrophotometer (Varian

cary 3) at room temperature using a quartz cell (10mm patch) for samples.

Basically the Ag-PCL samples were dispersed in dichloromethane and scanned

in the wavelength ranging from 400 nm to 800 nm. The spectra were

subtracted by the background UV-vis spectra of PCL solvent mixture. For the

measurement of release of silver nanoparticles, samples were cut into a disc

shape with 1.5 cm of diameter, and then eight pieces of samples were

immersed into 50 ml of distilled water and stirred at 200 rpm for 1 day at 37

°C. During this period, 1 ml of solution was drawn at pre-set time.

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3.5.15 Wide angle X-ray diffraction

X-ray diffraction (XRD) analyses were performed with a Philips Glancing

Angle and Powder Diffractometer 1140/90 using CuK α radiation (λ=1.54056

A) running at 40 kV and 40 mA. To determine the crystalline phases of the

specimens, a 2θ scanning range from 20 to 60 degree was used with a scanning

speed of 4 degrees per min.

3.5.15 X-ray photoelectron spectrometer

X-ray photoelectron spectra (XPS) were collected on a VG ESCALAB 220-

iXL spectrometer with a monochromated Al Kα source (1486.6 eV) using

samples of ca. 3 mm2 in size. The X-ray beam incidence angle is 0 ° with

respect to the surface normal, which corresponds to a sampling depth of ca. 3

mm2. The obtained XPS spectra were analysed by the CasaXPS software.

3.5.16 Inductively coupled plasma-atomic emission

spectrometer

After soaking the samples in SBF for certain days, 10 ml simulated body fluid

(SBF) post-soaking solution was collected for inductively coupled plasma-

atomic emission spectrometer (ICP-AES) testing.

3.5.17 Photo-spectroscopy

Photo-spectroscopy was used to detect the colour change of the matrix in a

very sensitive way. The spectrum could reflect the absorption peak of the

specific colour.

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Fig.3.13. DH-2000-BAL photo-spectroscopy

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C H A P T E R F O U R

Inter-bonded 3D fibrous scaffolds and their

cell culture performances

In this chapter, the preparation of inter-bonded 3D fibrous matrices was

introduced. The processing parameters influencing the formation of 3D

structures were discussed. Four matrices with different structures were

successfully prepared by controlling the ratio of fibre to sugar particles, and

their properties including pore structure, surface wettability, water

permeability, water binding ability and mechanical properties were examined.

Rat skin fibroblast and Chinese hamster ovary (CHO) cells were cultured on

the fibrous matrices. The cell growth performance was discussed in terms of

the cell morphology and cell proliferation. Two cell seeding methods, dynamic

and static seeding, were also compared.

4.1 Experimental Procedure

The fibrous matrices were prepared by the following procedure:

Step 1: melt extrusion of PCL filaments;

Step 2: cutting the long filaments into short fibres ( 1.5 mm);

Step 3: mixing the short fibres with particles of 300 μm 600 μm in size;

Step 4: putting the mixture in a mould and heating in an oven;

Step 5: leaching off the particles to form fibrous scaffolds.

The whole process is illustrated in Fig. 4.1.

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Fig.4.1. Flow chart for preparation of fibrous matrices

The fibrous scaffolds were then used to grow cells. Details on the cell culture

experiments including cell subculture, cell seeding, cytotoxicity test, MTS

experiment, cell morphology observation have been given in Chapter 3.

4.2 Results and Discussion

Many parameters influence the preparation of fibrous scaffolds, such as

temperature, time, fibre length, fibre diameter, particle size, and the ratio

between the fibre and sugar. Here, the process temperature in the range from

60 ºC to 70 ºC, fibre diameter from 50 μm to 200 μm, fibre to sugar ratio from

1:5 to 1:30 (w/w) were investigated. Fibre length was fixed at 1.5 mm or so.

4.2.1 Length and diameter distribution of short fibres

The fibre diameter was controlled by fibre extrusion conditions. Fig. 4.2a

shows the fibre diameter and diameter distribution. The average fibre diameter

is about 50 μm. In this thesis, the fibres with diameter of 100, 120, 150, 200

μm have also been used.

After obtaining the PCL filaments, staple fibres were obtained by manually

cutting these filaments. As shown in Fig. 4.2b, the mean fibre length of the cut

fibres is 1.5 mm. Due to the limitation of cutting machine, the fibre length in

this thesis was fixed at 1.5 ± 0.3 mm.

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Fig.4.2. Distribution of the as-prepared short fibre length and diameter, a) diameter distribution, b) length distribution

4.2.2 Size and distribution of sugar powders

The granular size of sucrose powders was regulated by sieving through a series

of meshes with different sizes. The particles of size between 300 μm to 600 μm

were collected. The size and distribution are shown in Fig.4.3.

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Fig.4.3. Particle size distribution of sucrose powders

4.2.3 Formation of 3D fibrous structures

The main purpose of using this method to make fibrous matrices is to

overcome disadvantages in the normal nonwoven structures. The basic

fabrication process includes melt bonding synthetic fibres in the presence of

particulate template (Fig.4.4). The short fibre makes it easy to mix with the

powdery template, with very low fibre-fibre entanglement. Sugar template

regulates the pore size and ensures the fibre bonding just at the initial contact

points. When the melting temperature is near 60 ºC, the fibre contact points can

melt together, but the sugar still stays in the solid state. The pressure can help

to improve fibre-fibre contact points.

One significant advantage of this method is using sugar as the template

material, which is non-hazardous to cells [422]. It causes no cytotoxic effects

even with incomplete removal from the matrix during the leaching process.

Template makes it possible to form larger pores besides smaller pores formed

by fibre-fibre interaction.

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Fig.4.4. Scheme of fibrous structure formation

4.2.4 Optimisation of processing parameters

4.2.4.1 Influence of temperature

Fig.4.5 shows the DSC curve of PCL materials. When the temperature is

higher than 50 ºC, PCL starts melting, and at 55.10 ºC it reaches the peak. The

shoulder of melting point of PCL is between 50 ºC to 62 ºC. The only

difference between PCL pellet, fibre and matrix is the width of shoulder, which

is in the order: pellet >fibre >matrix.

This suggests that in this temperature range, the polymer molecules move

quickly and can rearrange their structure easily. The melt bonding temperature

was thus chosen in the range between 50 ºC and 70 ºC.

Fig.4.5. DSC curve of PCL pellets, fibres and inter-bonded fibrous matrices

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Fig.4.6 shows the effect of processing temperature on the tensile strength of

end matrices. All the matrices were fabricated under the same conditions (1.5

mm short fibres, 1/15 (wt/wt) fibre/sugar ratio, 50 μm fibre diameter). At 60

ºC, fibres in the matrix still remained loose. The matrix almost had no tensile

strength. With increasing processing temperature, the matrix strength

increased. At 65 ºC, the ultimate breaking strength reached the highest,

however the matrix also showed the highest stiffness, which might not be good

for the further cell culture work. At 62 ºC, the matrix had certain bonding

between the fibres providing certain strength, but it was not too stiff. This

temperature was chosen for further study.

Fig.4.6. Tensile stress-strain curve of scaffolds with different processing temperatures

With the increase in temperature, fibres tend to melt together on the contact

positions. But when the temperature is very high, the short fibres would melt

completely and become small powders, which no longer maintain the fibrous

structure (Fig.4.7).

As proved by microscope images, when the processing temperature was 70 ºC,

many PCL particles were formed inside the matrix. Mode c (Fig. 4.7) was

chosen as ideal fibre-bonding structure.

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Fig.4.7. Fibre morphology changes with processing temperatures

4.2.1.2 Influence of fibre diameter

Fibre diameter played an important role in the formation of fibrous matrices.

When fibre/sugar and processing temperature were fixed at1/7(wt/wt) and 62

ºC, respectively, different fibre diameters resulted in different matrix tensile

strengths. As shown in Fig.4.8, with decreasing the fibre diameter, the stiffness

increased, and so did the ultimate stress.

Fig.4.8. Stress-strain curves of fibrous matrices made from PCL fibres with different fibre diameters

Under the fixed fibre length (1.5 mm), finer fibres have larger aspect ratio.

With higher aspect ratio, the fibres cannot easily mix well with powders due to

fibre-fibre entanglement.

4.2.1.3 Influence of fibre to sugar ratio

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By changing the fibre/sugar ratio, fibrous matrices with different densities were

obtained. In the experiment, fibre to sugar ratios of 1/5, 1/7, 1/10, 1/15, 1/20,

1/30 (wt/wt) were chosen. The tensile stress-strain curves indicated that this

parameter had larger effect on the tensile strength than other parameters such

as fibre diameter and processing temperature. The fibre-to-sugar ratio 1/5,

which had the highest fibre content, resulted in the largest stiffness, and the

highest ultimate strength. On the contrary, the ratio 1/30 gave the lowest

ultimate strength (Fig.4.9).

Fig.4.9. Stress-strain curves of fibrous matrices with different fibre–to-sugar ratios

The fibre in the matrix started breaking during extension, followed by the

bonding points when the tensile strength at the bonding points was larger than

that of the fibre. Otherwise, the break would occur firstly at the bonding points.

During the tensile test, the two phenomena could happen in the same samples

(Fig.4.10).

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Fig.4.10. Possible matrix breaking modes, matrix prepared by a) 1/5(wt/wt) fibre/sugar, b) 1/10 (wt/wt) fibre/sugar, c) 1/20 (wt/wt) fibre/sugar, d) 1/30 (wt/wt) fibre/sugar

By optimising the processing parameters, the optimum conditions to prepare

the fibrous matrices were obtained: Fibre length: 1.5 mm; fibre diameter: 50

μm; processing temperature: 62 ºC; heating time: 30 mins. Matrices with

different fibre-to-sugar ratios showed different densities and pore sizes.

4.2.5 Fibrous samples with different fibre/sugar ratios

Four matrices were prepared under the fibre/sugar ratio of 1:5, 1:10, 1:20 and

1:30 (wt/wt), which were respectively marked as PCL-5, PCL-10, PCL-20 and

PCL-30.

4.2.5.1 Densities of the matrices

The samples density was calculated based on the mass and the dimension of

dry samples. The density results were shown in Table 4.1.

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Table 4.1. Densities of different matrix samples

Samples Density (g/cm3)

PCL-5 0.0208±0.0007

PCL-10 0.0126±0.0008

PCL-20 0.0074±0.0005

PCL-30 0.00600±0.0006

4.2.5.1 Morphology of fibrous matrices

4.2.5.1.1 Optical microscopy

The optical microscopy images of the four matrices with different densities are

shown in Fig.4.11. PCL films (thickness 0.85 ± 0.03 mm) were prepared by

hot pressing. PCL-5 was the densest one, and PCL-30 was the least dense one.

The pores in PCL-10, PCL-20, PCL-30 were easily observed.

Fig.4.11. Digital photo of the PCL matrices produced from different fibre/sugar ratios

4.2.5.1.2 Scanning electron microscopy

As evidenced in the scanning electron microscopy (SEM) images in Fig.4.12,

the fibres in the matrices appeared uniformly and randomly distributed

showing no preferential orientation. They were bonded with each other at the

contact points. Such an inter-fibre bonding stabilises the fibrous architecture

and the porous network.

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Fig.4.12. SEM images of PCL matrices produced at different fibre/sugar ratios in the view of front and side, a) and a’) are front and side view of PCL-5, b) and b’) are PCL-10, c) and c’) are PCL-20, d) and d’) are PCL-30. Scale bar = 250 μm

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Among the four different matrices, PCL-5 was the densest one. Both small and

large pores existed in the matrix structure. Similar surface morphology was

also found for the sample PCL-10. For PCL-20 and PCL-30, matrices were

more or less loose. Fibre end could be found even in the cross-section of PCL-

30, indicating weak bonding between fibres.

4.2.5.1.2 Micro-computed tomography

X-ray micro-computed tomography (μ-CT) is a non-invasive and non-

destructive technique to obtain precise quantitative and qualitative information

about 3D micro-architectures. Here this technique was employed to explore the

3D structure of the fibrous matrices. The μ-CT images indicated that the matrix

contained two types of pores: randomly oriented small pores arising from fibre

accumulation and larger pores due to the sugar template. All fibrous matrices

had an excellent porous structure (Fig.4.13). All pores in the fibrous matrices

were well interconnected, without any dead-end pores. Such an excellent pore-

interconnectivity may facilitate the transport of nutrients and the formation of

neo-tissues.

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Fig.4.13. μ-CT images of the four fibrous matrices (isometric view of whole, isometric and front view of magnified central part), scale bar =250μm.

Unlike the conventional nonwovens in which the pores are oriented along the

X-Y plane (matrix surface), the pores in the prepared matrices showed no

preferential orientation, as proved by the randomly obtained 2D images of

micro-CT results in different directions (i.e. x-y, x-z, and y-z), which are

shown in Fig.4.14. This is possibly because of the use of short fibres in the

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matrix preparation which enables the fibres to orient in different directions.

Short fibre with small aspect ratio reduced the fibre-fibre entanglement.

Fig.4.14. 2D cross-section images of a) PCL-5, b) PCL-10, c) PCL-20 and d) PCL-30 in the order of x-y, y-z, and x-z directions from left to right, scale bar=500μm.

4.2.5.2 Pore structures

By changing the fibre/sugar ratio for the matrix preparation, the pore structures

of fibrous matrices were changed in a controlled way. As shown in Fig.4.15,

with a change of the fibre/sugar ratios from 1:5 to 1:30 (wt/wt), the average

pore size increased from 150 μm to 1000 μm. This increase in pore size led to

an increase of porosity from 73% to 99%, but a decrease in surface-to-volume

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ratio, from 4.1 cm-1 to 0.4 cm-1. These four matrices all had very good pore

interconnectivity, above 95%. These porous characteristics constitute the main

advantage of this template-aided fibre-bonding technique over the conventional

nonwoven and template leaching techniques. The conventional nonwovens

have a lower degree of bonding across the fabric thickness direction and very

limited space to adjust the pore size and porosity, while the conventional

template-leaching techniques typically result in low pore-interconnectivity.

Fig.4.15. Pore size, porosity, pore interconnectivity and surface-to-volume ratio of the four fibrous matrices.

4.2.5.3 Stability and mechanical properties

To prove the pore stability, the fibrous matrices were immersed in water and

stirred mechanically for 1 day. No obvious structural change was observed

after the treatment. However, for lab-made needle-punched nonwoven fabric

and non-bonded nonwovens having a similar porosity, lots of fibres were

detached from the matrices during the test, indicating the weak ability to

maintain the porous structure.

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The mechanical property of as-prepared matrices was shown in Fig.4.16. The

fibre/sugar ratio had an influence on both tensile and compression strengths.

PCL-5 has the largest ultimate tensile strength (6.7 MPa) and the smallest

compression strength (0.13 MPa). PCL-30 with the highest porosity showed

the smallest tensile strength (0.75 MPa) and the highest compression strength

(0.64 MPa). Higher porosity could give larger space for compression, but

reduces the tensile strength.

Fig.4.16. Mechanical properties of four different fibrous scaffolds, a) in elongation mode, and b) in compression mode

In comparison, the tensile strength of needle-punched nonwoven fabric was

direction dependence (Fig.4.17). In the machine direction, the nonwoven web

had good strength, but the tensile strength along the width direction was very

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weak, much lower than the fibrous matrices prepared in this work. The needle-

punched nonwoven had a high bulk density (0.27g/cm3), similar to PCL-5

(0.208 g/cm3), but the strength was much lower than PCL-5. The lab-made

nonwoven had similar density and porosity to PCL-30, but not inter-bonded,

had hardly any strength.

Fig.4.17. Comparison of tensile stress-strain curves between the fibrous matrices and normal nonwoven matrices, in which PCL-5(1) and PCL-5(2) are in elongation mode tested in two different directions. PCL-30(1) and PCL-30(2) are also tested in two different directions. Nonwoven 1 is needle-punched nonwoven. Its tensile strength was also tested in two directions (machine and width). Nonwoven 2 is lab-made non-bonded nonwoven fabric.

4.2.5.4 Water contact angle, hydraulic permeability and water binding ability

Water contact angle of the matrices was measured to detect the hydrophilicity

or hydrophobicity of the matrices. Increasing the density of fibrous matrices

reduced the water contact angle (Fig.4.18 a). This is probably due to the

wicking function existing in the low density matrices. Hydraulic permeability

of matrices is an important characteristic for tissue engineering, because an

excellent permeability can allow the exchange of nutrient and metabolic

wastes, and hence regulate cell migration deep into the fibrous scaffolds. For

the different fibrous matrices, the hydraulic permeability values were in the

range of 2×10-8~10×10-8 m4N-1s-1 (Fig.4.18b), which were between the

reported woven tissue scaffold (~1×10-15 m4N-1s-1) [121] and nonwoven tissue

scaffold (~5×10-6 m4N-1s-1) [421]. This suggests that the as-prepared matrices

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should be good for nutrient and waste exchanges inside the matrices, and thus

good for cell in-growth. Water bonding ability is also crucial for tissue

engineering applications. With good water bonding ability, the matrices can

hold enough nutrients to support cell growth. Fig.4.18 c shows that these

fibrous scaffolds have high water binding capacity. They can hold as high as

1200 weight percentage of water comparing to their original weight.

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Fig.4.18. Water contact angles, hydraulic permeability and water binding capability of matrix samples (PCL-5, PCL-10, PCL-20 and PCL-30)

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4.2.6 Cytotoxicity test

The cytotoxicity was determined by growing fibroblasts in the extract of the

scaffolding materials. The cell viability obtained was higher than 90% for all

the samples, including PCL pellets (raw material), as-extruded PCL fibres,

PCL fibrous matrices and PCL 2D films, and there was no significant

difference among the samples tested (p>0.05) (Fig.4.19), indicating that all the

tested PCL samples were non-toxic to cells.

This suggests that in the processing of fibrous matrices, from fibre extrusion

and fibre cutting to 3D structure formation, no extra toxic components are

introduced. In other words, all the as-prepared matrices are suitable for further

biological study.

Fig.4.19. Cell viability of fibroblasts growing in the extract of the scaffolding materials (p>0.05)

4.2.7 Comparison of dynamic and static seeding methods

On the base of the equipment availability, the cell seeding methods, such as

shaking seeding using orbital shaker, rotating seeding using blood roller and

small vials and stirring seeding using spinner flask, have been trialed. In the

experiment, the influence of speed on the seeding efficiency was examined.

Cells used in all the related experiments were CHO cells.

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4.2.7.1 Shaking seeding method using orbital shaker

The speed of orbital shaker was set at 0, 50, 75, 100, or 125 rpm. Seeding

density is 5×104 cells/matrix (because at low seeding density like 1×104,

seeding efficiency couldn’t be detected correctly, which may result in large

experimental errors for MTS assay since there was almost no colour developed

when the cell number was low. Another reason for the high seeding density is

that the high seeding density is often required in many circumstances, and it is

believed that static seeding methods are mainly limited to high seeding density.

The seeding volumes varied at 50, 100, 200, 500, and 1000μl.

The seeding efficiency was calculated using equation (4.1).

(E4.1)

It changed when different seeding volumes or different shaking speeds were

applied.

As shown in Fig.4.20, the influence of seeding volume on the seeding

efficiency is not apparent. The volume 200 μl gave the least fluctuating

condition. But the effect of shaking speed on the seeding efficiency was great

(Fig.4.21). At speed 50 rpm, the seeding efficiency was the highest, which

proved that lower speed led to better cell adhesion. This is similar to that was

concluded from the stirring seeding method.

Fig.4.20. Relationship between seeding volume and seeding efficiency

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Fig.4.21. Influence of shaking speed of orbital shaker on seeding efficiency

4.2.7.2 Rotating seeding method using blood roller

Here, roller having different speeds was firstly trialled and seeding density of

5×104 cells/matrix was used. The volume in one 25 ml centrifuge tube is 2 ml,

which can just cover the whole matrices.

The cell number was measured by MTS, and seeding efficiency was then

calculated according equation (4.1).

The seeding efficiency varied with the speed, and the best seeding efficiency

was found at the speed of 37 rpm (Fig.4.22). Lowering the speed improved cell

attachment.

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Fig.4.22. Influence of rotating speed of blood roller on seeding efficiency

4.2.7.3 Stirring seeding method using spinner flask

The speed of spinner flasks was set in the range between 50 rpm and 80 rpm

based on the previous results on shaker and blood roller which indicated lower

speed was better for cell attachment. Seeding density of 1×106 cells/ml was

used. Because this method requires a huge volume of media to ensure the

sufficient contact between media and samples, if seeding density was lower,

e.g. 5×104 cells/matrix, the actually cell density within the seeding media was

just 5×103 cells/ml. With such a low cell density, the number of cell attaching

on the surface might be very small. Therefore, a cell density of 1×106 was

used. The seeding volume was fixed at 110ml. With such a medium volume,

the matrix can be merged into the media easily.

Fig.4.23 indicates the seeding efficiency of the stirring seeding method. With

a high seeding density (1×106), the seeding efficiency was not higher than 5%.

This seeding method could give a good cell distribution. Because this method

consumes a huge volume of media (110 ml), it was not considered a good

seeding method to seed cells into 3D porous scaffolds.

Fig.4.23. Influence of stirring speed of spinner flask on seeding efficiency

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4.2.7.4 Comparison between static and dynamic seeding methods

Among these dynamic seeding methods, the shake seeding using blood roller

gave the highest seeding efficiency. Therefore, rotate seeding was chosen to

further compare with static seeding method. PCL-10 samples were seeded with

CHO cells by the standard static seeding method and the rotate seeding method

at the optimised conditions. The cell morphology and cell number at various

time-courses were recorded.

The confocal images of CHO on matrices after the static seeding and the

dynamic seeding were shown in Fig.4.24. CHO cells looked healthy in the

fibrous matrix after either static or dynamic seeding. The cells distributed on

the fibre surfaces also looked uniform. Cells were growing on the fibre surface

along the fibre direction.

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Fig.4.24. Confocal microscopic images of CHO cells seeded by static and dynamic methods. From top to bottom, images were taken with different magnifications. The bottom images were taken under white light.

SEM images in Fig.4.25 also reveal the healthy growth state of the cell seeded.

There is no much difference observed among the SEM images.

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Fig.4.25. Morphology of dynamic seeded (a, c, e) and static seeded (b, d, f) CHO cells, a & b are 3 days, c & d are 7 days, e& f are 14 days.

The difference between static and dynamic seeding methods can be found after

14 days of cell growth. After two weeks of cell culture, the MTS assay was

conducted on the fibrous matrices and the results are shown in Fig.4.26.

Initially, the cells number showed no obvious difference. With longer time of

cell growth, cells seeded by the static method became confluent, and then died

in some parts, while for dynamic seeding, the cells still kept growing because

of more uniform distribution within the fibrous matrix. After 7 days of culture,

the cells number for the dynamically seeded samples was higher. This trend

was maintained until 14 days of cell culture, and on the 14th day, the cell

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number in the dynamically seeded was larger than that by static seeding

method.

Fig.4.26. Comparison of cell number in fibrous matrices between dynamic seeding and static seeding methods

According to the above experiment, dynamic seeding is good for long-term

culture study due to uniform cell distribution, and static seeding is much easier

to conduct, although it has problems of cell arrangements in the scaffolds. Due

to the limitation in the equipment availability, the consumption of the culture

media and the possibility of scale-up study (rotary vials) for dynamic seeding

methods, however, classical static seeding method was mainly used in the

following part.

4.2.8 Cell seeding efficiency

Seeding efficiency reveals the ability of cells to attach on a scaffold material.

The seeding efficiencies of both cell lines on the fibrous matrices are listed in

Table 4.2. The PCL-5 had the largest seeding efficiency for both fibroblast and

CHO cells. The matrices having a smaller mean pore size showed larger cell

population. More cells were trapped by fibres, and less penetrated into the

wells. The seeding efficiency for fibroblasts was larger than that of the CHO

cells. This might be because fibroblasts were larger than CHO cells in size.

This assisted in the cell attachment on the fibres, but not on the wells. 2D PCL

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film showed lower seeding efficiency than the commercial 2D culture plate.

Because 2D culture plate was surface-treated by plasma better for cell culture

study, while 2D PCL film used here received no surface treatment.

Table 4.2 Seeding efficiency of CHO and fibroblast cells on different fibrous

matrices

Seeding efficiency on the Petri dish is 93.6% for CHO and 93.9% for

fibroblast

4.2.9 Cell morphology

At different time intervals, the cell-matrix composites were taken out, and

treated for SEM observation and confocal microscopy observation.

4.2.9.1 SEM

The morphologies of CHO cells growing in the PCL fibrous matrices are

shown in Fig.4.27. The cells adhered and spread well on the surface of PCL

fibres. Rounded cells, indicative of non-viable or unhealthy cells were not

observed, further demonstrating the conducive environment for cell culture.

Live CHO cells mostly maintained a spindle-shape even though some cells

with elliptical and polygonal morphologies were observed. The cells spread

and proliferated randomly around the fibres. They wrapped around the fibre

Samples CHO (%) Fibroblast (%)

Matrix Well Media Matrix Well Media

PCL-5 18.3 32.0 49.7 47.8 37.7 14.5

PCL-10 13.5 71.1 15.4 43.7 44.8 11.6

PCL-20 9.6 46.5 43.8 26.5 63.0 10.5

PCL-30 3.7 47.9 48.4 14.7 75.1 10.3

PCL film 3.9 39.2 57.0 12.8 78.5 6.1

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surface and grew along the fibre axis without directional preference. From the

very beginning, some of CHO cells were seeded inside the scaffolds due to the

fabulous pore structures, cells distributed not only on the surface of scaffolds,

but also inside it. After some time, cells on the surfaces could also grow along

the fibre length direction to fibres inside the matrices. This phenomenon

resulted in the random distribution of CHO throughout the whole matrices.

The cell morphology on the fibrous scaffolds varied over time. After 3 days of

seeding, CHO cells cultured on PCL-20 proliferated throughout the matrix,

indicating that the matrices are suitable for the growth of CHO cells (Fig.

4.27a and 4.27a’). On the 7th day of culture, the cells covered more surfaces

of the fibres and they thickened to anticipate rounding up for mitosis (Fig.

4.27b and 4.27b’), persumbly because of the ability of CHO cells to change in

response to contact with other cells in the population [423-424]. With an

increase in the culture time, the cell population increased dramatically (Fig.

4.27c&4.27c’, 4.27d-4.27d”). As a consequence of the adaptive processes,

cells shape changed. Viable cells were obviously in a highly asymmetric

(bipolar) form, toward spindle-like shaped and orientated along the fibre

direction due to a process called contact guidance [295, 303, 425]. Cell growth

on other fibrous matrices showed a similar tendency.

For comparison, CHO cells were also cultured on 2D culture plates in the same

conditions and their morphology on the 14th day of culture is shown in Fig.

4.27e. CHO cells appeared flattened, well spread, but not as elongated as those

cultured on 3D PCL matrices. Such differences in cell morphology and cell

spread should derive from the apparently different culture environment

between 2D film and 3D fibrous structure [426-427].

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Fig.4.27. CHO cells in fibrous scaffolds (PCL-20) after a) 3 days, b) 7 days, c) 14 days, and d) 21 days of culture. e) CHO cells after 14 days of growing on PCL film under the same culture

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condition. (Cells were fixed with OsO4, and images were taken from front or back of samples. All scale bar = 50 μm.

Fig.4.28 displays morphology of fibroblasts cultured on PCL-20. The arrows

point to the cells attached to multiple fibres close to the inter-fibre bonding

sites or where fibres came in close proximity (overlapping fibres). Dehydration

and gold coating before the SEM observation are the main causes of the cracks

visible in the SEM images, especially for fibroblasts layers. Within 7 days of

the culture, fibroblasts adhered well and showed a healthy morphology.

Especially, cells were observed growing along the fibre axis and bridging

between adjacent fibres (Fig.4.28b & 4.28b’). At day 14, the number of

bridging cells increased. At the 21st day, layered fibroblasts were observed.

These results suggest that fibroblasts attached, spread and grew very well on

the 3D fibrous scaffolds during the culture period and that they could form

layers spanning the entire scaffold surface with time. By comparison, no

layered cell morphology could be seen on the PCL film.

Fibroblasts grew on the fibre surfaces when there were enough spaces at the

early culture day. With the time increasing, fibroblasts could bridge between

fibres while the distance was short enough for their branch. This kind of branch

of fibroblast made it possible to form an entity where pores were small and

also allowed it to grow inside the matrices.

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Fig.4.28. Fibroblast cells on fibrous scaffolds (PCL-20) after a) 3days, b) 7 days, c) 14 days, and d) 21 days of culture. e) Fibroblast cells after 14 days of growing on PCL film under the

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same culture condition. (Cells were fixed with OsO4, and images were taken from front or back of samples. All scale bar = 50 μm.

4.2.9.3 Confocal microscopic images

Fig.4.29 shows the confocal microscopy images of the CHO cells growing in

the fibrous matrices. They either grew randomly on the fibre surface

(Fig.4.29a), or they grew orderly along the fibres (Fig.4.29b), all of them

maintained elliptical, polygonal or spindle-like shape showing their healthy

state. In Fig.4.29d, an image with lower magnification proved the cell

distributed uniformly on the fibres in the matrices.

Fig.4.29. Confocal microscopic images of CHO cells on PCL fibrous matrix (PCL-10). The cells were fluorescently stained in red for the actin filaments and in blue for the nuclei. a) showed cells distribute randomly, b) showed cells ranged orderly along the fibre, c) is the magnified images indicated clearly cell shape, d) is image with long magnification verified cell distribution in the matrices.

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Fibroblast grew nicely on the PCL matrices were also be proved by confocal

microscopic images in Fig.4.30. Fibroblasts wrapped the fibre surface and

grew along the length direction, which resulted in cell migration happening

(Fig.4.30a). Fibroblasts also bridged when the distances between fibres were

compatible to their sizes (Fig.4.30a’), similar conclusions could also be drawn

from SEM images.

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Fig.4.30. Confocal microscopic images of fibroblasts on PCL fibrous matrix (PCL-10). The cells were fluorescently stained in red for the actin filaments and in blue for the nuclei. a)-a”) are images with high magnification. b)-b”) are images with low magnification, in which a) showing cell growing along the fibre, a’) indicating cells can bridge between fibres, a”) demonstrated one whole fibroblasts cell with clear shape; b) is image presenting nuclei only, b’) is image expressing actin filaments only, b”) is images stacked by b) and b’) scale bar=50μm

4.2.10 Cells penetration inside the matrices

4.2.10.1 Confocal microscopic slices

Laser confocal microscopy makes it possible to focus on different z position of

samples, to scrutinise cell migration inside the samples. As shown in Fig.4.31

for CHO cells, and Fig.4.32 for fibroblasts, throughout the thickness of 320

μm, there were red colours and green colours at different layers indicating the

presence of cells on different level. Since low magnification was used to give a

general view of the cell distribution, cell cytoskeleton and nuclei are not well

distinguishable in these images.

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Fig.4.31. Confocal microscopic cross-section slices of CHO cells on PCL-10 at various z positions (the image were taken every 40 μm throughout the whole scaffold thickness nearly 300 μm) (PCL-10 after 21 day culture of CHO cells) Scale bar = 250 μm

Fig.4.32. Confocal microscopic cross-section slices of CHO cells on PCL-20 at various z positions (the image were taken every 40 μm throughout the whole scaffold thickness nearly 300 μm) (PCL-10 after 21 day culture of CHO cells) Scale bar = 250 μm

4.2.10.2 micro-CT

Confocal microscopy has a limitation to collect clear images over 300 μm in

depth. To confirm cell growth throughout the whole thickness of the matrix

samples, micro-CT was used. As evident in the reconstructed micro-CT image

in Fig.4.33a, many OsO4-stained cells (red spots) were observed on the

surface of fibres (green) inside the 3D scaffolds. In contrast, a micro-CT image

of the PCL fibrous matrix without containing any cells was also included (Fig.

4.33b), but no red spots can be seen in the matrix.

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Fig.4.33. micro-CT images of 3D scaffolds with and without cells inside

4.2.11 Cell proliferation

Fig.4.34. MTS assay of CHO and fibroblasts on fibrous matrices

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Fig.4.34a summarizes the mean number of CHO cells growing on the fibrous

matrices over 21 days of culture. Because of the low seeding efficiency, the

culture number initially on 2D control should be higher than that on the 3D

matrices. This is true for the fibroblasts. The fibroblast cells on the 2D control

after 3 days of culture showed a higher number than those in the 3D matrices

(Fig.4.34b). However, for the CHO cells, it showed a reverse trend. The low

CHO number on the 2D control in the early days came from the uneven

distribution of the CHO cells in the culture plate. Despite the high seeding

efficiency for tissue culture plate (TCP), most of the CHO cells preferred to

stay on the edge area of the TCP well (see Fig.4.35). This led to the occurrence

of cell confluence in this area just within three days of culture.

The higher cell numbers in the 3D matrices than on the 2D control was

maintained for the whole culture period. The CHO cells in the matrices grew at

a similar rate during the culture period. However, the cell number on the 2D

PCL film and the culture plate was stabilized after 2 weeks of culture,

indicating that the cell growth rate slowed down from day 14 of culture.

Because of the contact-inhibitive nature, the CHO cell growth is surface area

dependent. Larger surface area in scaffold should result in more cells to grow.

This could be the main reason as to why the 2D controls accommodated fewer

cells than the 3D fibrous matrices.

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Fig.4.35. CHO cells cultured on 2D plates for three days

The sample PCL-5 had the highest surface area (265.7 cm2), the lowest

porosity (73.19%) and the smallest pores (119.6 μm) among the fibrous

matrices. It also showed higher seeding efficiency and larger cell number in 3

days of cell culture than the other matrices. The results that longer culture time

resulted in decreased cell growth rate could be attributed to two reasons: the

growing cell population reduced the pore size, which adversely led to inhabited

mass transfers of nutrients and waste, and the cells were confluent in some

areas of the matrix. Scale bar=50μm

By comparison with PCL-5, PCL-10 provided larger pores (size above 200

μm) for the cell growth. This also facilitated the nutrient and waste exchange,

allowing cell growth at a steady rate. In the first week, the cell number in PCL-

5 was higher than that of in PCL-10. However, the cell number in both

matrices on the day 14 became very similar. Longer culture resulted in higher

cell number in PCL-10 than PCL-5. This suggested that smaller pores could

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provide more opportunities for the fibroblasts to attach to scaffold, but larger

pore could provide larger space for the cells to grow over time. PCL-20 and

PCL-30 contained larger pores, respectively about 800 μm and 1000 μm. The

mean cell number on PCL-30 was much less than that on PCL-20 after 21 days

of culture, which could come from the reduced overall surface area of the

fibrous matrix. These results indicated that pore size played an important role

in the cell growth.

The MTS assays performed on the fibrous matrices revealed fibroblasts

proliferation on the 3D PCL matrices (Fig.4.34 b). Within the 21 days of the

culture period, the number of fibroblasts continuously increased in the fibrous

matrices. Comparing with the 2D tissue culture plates, the 3D matrices had less

cell numbers at day 3 and day 7, due to the low seeding efficiency. However,

the cell population on the matrices increased continuously with time and

became comparable to 2D tissue culture plates in two weeks, and then

dramatically larger by day 21. The 2D PCL film showed the lowest cell

number among the sample groups after 21 days of culture. This can be

explained in that the 3D fibrous matrices have a larger surface area than 2D

films or well, allowing the cells to spread on the fibre surface. The 3D porous

structure also facilitates the cell uptake and exchange of nutrients and disposal

of waste.

Again, the porosity, pore size and surface area of the fibrous matrices affected

the fibroblasts growth. The matrix PCL-5 yielded the highest cell number. This

trend was maintained for the whole three-week culture period. This is different

from what was observed for the CHO cells, in which the cell growth rate

levelled off from day 14. Fibroblasts are large flat cells with the size about

30~100 μm and the cell thickness about 3~4 μm. In this case, cells growing on

fibre surface had relatively smaller influence on pore volume hence the media

and waste exchange compared to the CHO cells. As a result, sensibly larger

cell population can be achieved on the matrix with a larger surface area, and

PCL-5 disk has the largest cell population after three weeks of culture. It is

reasonable to predict that longer culture period would lead to more fibroblasts

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growing in the fibrous matrices, and the inhibited mass transfer will still

happen once the cells take up certain pore volume.

4.3 Conclusion

Stable fibrous matrices with different pore porosities, and pore sizes were

successfully produced by the combined hot-press and particle-leaching method.

The influences of bonding temperature, fibre diameter, fibre/sugar ratio on the

formation of fibrous structures were examined. Through controlling the

fibre/sugar ratio, four matrices with different structures were obtained.

The SEM and micro-CT observations revealed that these fibrous matrices had

excellent porous structures with small pores formed by fibre assembly, and

large pores due to template particles. These structures were highly porous (over

80% porosity). The pores inside the matrices were well interconnected and no

dead pores were found even in the dense areas. Moreover, the pore structure

had no preferred orientation.

These fibrous matrices also had good mechanical properties compared to

nonwoven fabrics. They were stable due to the inter-bonded fibrous

characteristic. They also had excellent permeability to culture media showing

their potential for use in tissue engineering.

Both static seeding and dynamic seeding methods were used to seed CHO cells

on the matrices. It was found that cells could grow on the fibrous porous

matrices by either static seeding or dynamic seeding method. Although

dynamic seeding method provided high seeding efficiency and high level of

uniformity on cell distribution, it is limited to large scales of seeding. Static

seeding has been used instead in this study.

The as-prepared fibrous matrices were biocompatible and conducive to cell

growth, which was proved by culturing CHO and fibroblast cells on them. By

using micro-CT and confocal microscopy, the cell migration and growth inside

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the fibrous matrices were clearly proved. It was also found that the pore

characteristics and surface area played an important role in mediating the

growth of CHO cells and fibroblasts inside the matrices differently. The

integration of large surface area and smaller pores was suitable for the growth

of fibroblasts. Smaller pores could provide more opportunities for the

fibroblasts to attach to scaffold, but larger pore could provide larger space for

the cells to grow on a long term basis. The scaffold-making approach

developed allows the choice of different fibres and the control of pore size and

porosity, and fibrous scaffolds with different pore characteristics can be

produced to suit different cell lines for many different biomedical applications.

Since the developed fibrous matrices show better mechanical performance and

more stable pore structure than the conventional nonwoven scaffolds, they

should have better performance in tissue engineering applications.

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C H A P T E R F I V E

Nano-structured surface in 3D fibrous

scaffolds and its cell culture performances

In this chapter, the influences of nano-structured fibre surface in the 3D fibrous

matrices on the cell culture performances are presented. An electrostatic self-

assembly method was used to load silica nanoparticles on PCL fibrous

matrices. Osteoblasts were seeded onto the scaffolds to examine cell

proliferation and cell differentiation performance.


In this chapter, only matrix PCL-10 (porosity >90%, mean pore size 200 μm)

was used. Silica nanoparticles were synthesised from TEOS according to the

Stober method. The detailed method was introduced in Chapter 3 already. Then

these obtained silica nanoparticles were coated on the PCL fibrous scaffold

surfaces by electrostatic layer-by-layer self-assembling method. After the

NaOH surface treatment, PCL surface could form abundant carboxylic groups,

with a negative charge. Since silica nanoparticles had a positive charge,

electrostatic assembly was possible. PAA/SiO2 was used in the following

double electric layers to control the content of nano-SiO2. The silica

nanoparticle coatings were controlled in 0 layer, 1 bilayer, 3 bilayers and 5

bilayers.


5.2.1 Silica particles and size distribution

Fig. 5.1 shows a typical SEM image of silica nanoparticles synthesised and the

histogram of the particle size distribution. The synthesised silica nanoparticles

ranged from 50~70 nm in size. Due to the dynamic light scattering (DLS)

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measurement used water as solvent, and the reflection of particles in water

resulted in relatively larger size. Sizes obtained based on SEM images also

were a little larger than the real size due to the gold coating.

Fig.5.1. a) SEM images of synthesised silica nanoparticles, b) histogram of particle size distribution measured from SEM image, and c) histogram of particle size distribution obtained by DLS.

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5.2.2 Nanoparticle assembly and surface morphology The fibre morphologies before and after self-assembly with silica nanoparticles

are presented in Fig.5.2. Before self-assembly, there were no particular

structures found on the fibre surface (control sample, PCL fibrous matrix,

Fig.5.2a). When assembling with one layer of silica nanoparticles, particles

can be observed clearly around the fibre surface (Fig.5.2b). After coating with

three and five layers of nano-silica, the nanoparticles covered the entire fibre

surface homogeneously, and some parts even formed an aggregated structure

(Fig.5.2c and 5.2d).

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Fig.5.2 SEM images of control (a and a’), 1 bilayer (b and b’), 3 bilayers (c and c’) and 5 bilayers (d and d’) self-assembled fibre samples. The scale bars for images a-d are 10 μm, and for images a’-d’ are 200 nm.

5.2.3 Surface chemical components To confirm the surface components, ATR FTIR spectra were recorded and

shown in Fig.5.3. Most of the main band characteristics of PCL were observed

in the spectra, suggesting a very thin layer of silica nanoparticles was deposited

on the PCL substrate. For the fibre coated with 1 layer silica, the difference

was not so distinct. After silica nanoparticle surface modification, the band

peaks at 800 cm-1, 1101 cm-1, 3360 cm-1 arose from Si-O-Si symmetric

stretching vibration, Si-O-Si asymmetric stretching vibrations, and Si-OH

stretching [428-429]. With increasing silica nanoparticle layers, the peak

intensity increased indicating more silica nanoparticles coated on the PCL

surface.

Fig.5.3. ATR-FTIR spectra of control and surface modified samples

Surface chemical components of the treated fibrous matrix were detected by X-

ray photoelectron spectroscopy (XPS) technology. As shown in Fig.5.4, the

survey spectra reveal that silica nanoparticles are successfully assembled on

the surface. For comparison, XPS spectrum of the pristine control samples is

also included, but only has O and C peaks (Fig.5.4a). When the fibre surface

was treated by NaOH solution, a new peak with the binding energy of 490 eV

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appeared, corresponding to Na 2s (Fig.5.4b). When a thin layer of PAA/SiO2

was assembled on the fibre surface, new peaks with binding energies of 153

and 102 eV, attributed to Si 2s and Si 2p respectively, appeared which

confirmed the formation of SiO2 on the surface (Fig.5.4c).

Fig.5.4. XPS spectra of PCL fibrous samples a) non-treated, b) treated with NaOH solution, c) treated with both NaOH and thin layer of nano-SiO2

The high resolution XPS C1s spectra of the fibrous matrices are shown in Fig.

5.5. Curve fitting the spectrum revealed that the pristine control samples had

four components (Fig.5.5a), -CH2-(284.56 eV), -CH2-C=O (285.05 6eV), -

CH2-O (285.936 eV), and -C=O (288.598 eV). After NaOH treatment, the

group –O-C=O was changed to –COOH and -CH2-OH, as indicated by the

binding energy shift in the range of 287~289 eV. When PAA/SiO2 layers were

assembled to the surface, a huge increase in –C-OH (285.993eV) was detected.

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Fig.5.5. High resolution XPS C1s spectra of a) Pristine control sample, b) NaOH treated sample, c) nano-SiO2 modified sample

5.2.4 Atomic force microscope After surface modification, the surface roughness of PCL fibres was measured

by AFM. Only one layer of silica presented a roughness value (RMS) at 24 nm,

and with increasing bilayer number, surface roughness increased to 33 nm and

37 nm. This rough surface compared to the smooth surface of pristine samples

(3 nm) would improve osteoblast cell attachment, and might assist in further

cell growth performance.

Fig.5.6. AFM images of control and surface modified samples. a) pristine sample, b) 1 bilayer modified sample,c) 3 bilayers modified sample, d) 5 bilayers modified sample.

5.2.5 Surface wettability As known from literature, cell adhesion, spreading and growth require the

scaffold surface having suitable hydrophilicity [430]. Since silica nanoparticles

are hydrophilic, applying silica nanoparticles on PCL fibre surface could

improve the surface wettability. With an increase in the silica nanoparticle

layer, from 0, to 1, 3 and 5, the water contact angle decreased from 78, to 52,

37, and 28 degree, responsively (Fig.5.7a~5.7d), suggesting the enhanced

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water wettability. Fig.5.7e shows the contact angle could be stable in a very

short time, no more than 0.3 s.

Fig.5.7. Surface water contact angle before and after surface coating, a) control sample, b) 1 bilayer, c) 3 bilayers, and d) 5 bilayers coating samples; e) water contact angle changing with contact time.

5.2.6 Cytotoxicity Cytotoxicity study was conducted on all the as-prepared samples. Results were

given in Fig.5.8. All the samples provided over 90% cell viability, no matter

how many layers were applied. This indicates that silica surface treatment is

not toxic to cells, and thus suitable for further cell culture study.

Fig.5.8. Cell viability of fibroblasts growing in the extract of the scaffolding materials, bar chart indicates the mean cell number, and the line chart displayed cell viability. Values shown are the means of three measurements on individual specimens. Error bars show ± standard deviation (p>0.05).

Fibroblast cells attached and spread very nicely on the tissue culture plates

(TCP) for all tested samples (Fig.5.9). These images suggested that even in the

presence of extract of silica nanoparticles solution, fibroblasts can still sustain

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a thinly spread healthy state, illuminating the nontoxic feature of PCL matrices

and surface silica nanoparticles coated PCL matrices.

Fig.5.9. Cell morphology of fibroblasts growing in the extract of scaffolds after culture for three days, a) control, b) 1 bilayer, c) 3 bilayers, and d) 5 bilayers; The scale bar for images a-d are 500 μm, and for images a’-d’ are 100 μm.

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5.2.7 Osteoblast cell culture

MTS assay was applied to measure the mean cell number at different culture

time (1 day, 3 days, 7 days), referring to the cell proliferation on different

samples (Fig.5.10). At the very early stage, i.e. day 1, the mean cell number

was in the order of 5 bilayers> 3 bilayers> 1 bilayers> 3DC, which was an

evidence that surface roughness affected cell attachment. The higher the

surface roughness, the higher was the number of adhered cells. After three

days of culture, cell number on the three treated samples was dramatically

higher than that on the control pristine sample. But the difference is not

statistically significant (at the 5% level) among the three treated samples. A

further 4 days culture offered a clear trend of variability among different

samples. Samples with one layer of silica enhanced cell division greatly. Cell

number was nearly two times more than other samples. Surface nano-textured

structure was believed to enhance cell attachment and adhesion, but probably

would prohibit the cell proliferation [431]. That could be the reason why

samples with three and five bilayers had relative fewer cells on day 3 and day

7. If only the cell attachment is considered, samples with five layers of silica

nanoparticles are better. For cell proliferation, samples with one bilayer silica

coating were better than others.

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Fig.5.10. Mean cell number changes of hFOB 1.19 on different scaffolds as a function of time. Values shown are means ± standard deviation. * p<0.05 versus 3DC at day 3, † p<0.05 versus 3DC at day 7.

5.2.8 Cell morphology

Cell morphology was recorded by laser scanning confocal microscopy (LSCM)

after a certain culture period. Cells were fluorescently stained with nuclei

shown in blue colour and actin filaments in red colour in these images. The

osteoblast cells maintained a polygon shape on control samples, but on silica

treated samples, at day 1, osteoblasts tended to become round, and their size

was a little smaller than that on the control samples. At day 3 and day 7, more

spread osteoblast appeared, as shown in Fig.5.11. The changes in surface

topography and roughness may hold vital clues to cell activity, especially cell

attachment [308]. Nano-structured surface can enhance the cell elongation and

spread because the cells can recognise the nano-scale topography [138, 302,

432]. In addition, there were fewer cells on the control pristine samples than on

silica nanoparticle coated samples.

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Fig.5.11. Confocal images of hFOB 1.19 on scaffolds for culture 1, 3, and 7 days (from left to right), a) control, b) 1 bilayer, c) 3 bilayers, d) 5 bilayers (from top to bottom). The cells were fluorescently stained in red for the actin filaments and in blue for the nuclei. Scale bar = 100 μm.

SEM images of cells on different substrates confirmed the results obtained

from confocal microscopy. PCL pristine surface has no reactive groups on the

surface for osteoblast attachment (Fig.5.12a and 5.12e). At day 3, the cells still

kept a round shape as just seeded. At day 7, cells attached to the surface, but no

actin filament was observed. In contrast, the surface assembled with silica

nanoparticles gave cells nano-scaled rough surface to attach and spread

(Fig.5.12b, 5.12c, 5.12d). The actin filaments are exhibited in Fig.5.12c’ and

5.12d’, which is another evidence showing the biocompatibility of silica

nanoparticles coating. With 7 days of culturing, the osteoblast cells tended to

be elongated comparing to 3 days of culturing (Fig.5.12 f-h).

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Fig.5.12. Cell morphology of hFOB 1.19 on scaffolds after 3 days and 7 days of culture, a) control-3days, b) 1 bilayer-3days, c) 3 bilayers-3days, d) 5 bilayers-3days, e) control-7days, f) 1 bilayer-7days, g) 3 bilayers-7days, h) 5 bilayers-7days. Scale bar = 2 μm.

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5.2.9 Alkaline Phosphatase activity

The synthesis of alkaline phosphatise during cell cultures was also assayed,

which was an early indication to bone formation. ALP activity also reflects the

overall osteoindustivity and cell differentiation. To measure the

osteoindustivity of the silica nanoparticles coated PCL fibrous scaffolds, ALP

activity of the osteoblasts that were cultured on those samples was quantified.

Samples with 1 layer of silica nanoparticles were tested, and the 2D tissue

culture plate (TCP) and the 3D fibrous matrix without the surface modification

were used as controls.

After three days of culture, the osteoblasts growing on the silica treated

samples showed significant increase in the ALP activity, as evidenced in

Fig.5.13. By comparison, the ALP activity of osteoblasts on 3D and 2D

controls was lower. Longer time of culture leads to lower ALP activity. At day

7, both silica treated samples and 2D/3D control samples had a low ALP

activity, while silica treated samples were still much more than that on control

samples. Another peak appeared near 10 day and 14 days. This tendency for

ALP activity was also described by other researchers [433]. ALP could reach

the first peak just after seeding, but with time, its quantity reduced until another

peak appeared. The nearly 3-fold increase in ALP activity observed in the

osteoblasts growing on the silica-treated samples after 14 days of culture was

an indication that this special surface-treated scaffold had osteoinductivity in

vitro.

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Fig.5.13. Alkaline phosphate content of hFOB 1.19 on 1 bilayer coating and control samples as a function of time (3 days, 7 days, 10 days, 14 days), *p <0.05 versus 3D control samples and 2D TCP at corresponding day, † p<0.05 versus results on day 7.

The higher ALP activity for the silica nanoparticle treated fibrous matrices

could be explained by three factors. First, 3D environment could provide better

nutrient and waste exchange for osteoblasts, which might help with increasing

the ALP activity. Although the ALP activity here is similar to 2D control, the

cell number on 3D control was dramatically different to that on 2D control.

The seeding efficiency was 17.32% of cells attached on the 3D control, but

93.52% for 2D tissue culture plates. Second, the hydrophilic and nano-scale

surface characteristics enhance the cell attachment, thus the cells initially

attached to the silica treated samples were much higher in number than that on

3D control samples. It was tested that the cell number attached on the treated

scaffold was 26.4% of the original seeded cells, while only 17.32% of original

cells attached on the 3D control. The large cell numbers might have higher

ALP activity in the same environment. Third, the silica nanoparticles

assembled can enhance the osteoinductivity[434]. Si is one of the components

composing bone, Si-O-Si, and Si-OH groups existing in SiO2 nanoparticles

could help the formation of bioactive CaP on the substrate surface. In addition,

PCL fibrous matrices with self-assembled silica nanoparticle surface coating

could form uniform hydroxyapatite layers after immersing in simulated body

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fluid (SBF) for 7 days, and the hydroxyapatite layer became thick and

continuous after exposing to SBF for 21 days (Fig.5.14).

Fig.5.14. Formation of apatite on the silica nanoparticle coated PCL fibrous matrices surface after immersing into 1×SBF for 7 days. a) & a’) pristine control sample, b) & b’) silica treated samples just after self assembly, c) & c’) silica treated samples after immersing in 1×SBF for 3 days, d) & d’) silica treated samples after immersing in 1×SBF for 21days.

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ALP activity was also demonstrated visually by staining the ALP by Naphthol

AS-MX phosphate and diazonium salt. All the ALP generated was stained to

pink to red colour. The cells can further be nucleus stained by Mayer’s

hematoxylin, showing blue colour in the images.

Although some osteoblasts with dark nucleus stained existed, stained ALP was

really hard to be found in 3D control samples, even at day 3 (Fig. 5.15), at

which point it was proved previously to have the largest ALP content for 3D

control samples. As comparison, the stained ALP could be observed for

nanoparticle treated samples at every time point: day 3 (Fig.5.15), day 7

(Fig.5.16), day 10 (Fig.5.17), day 14 (Fig.5.18). The pink colour with a darker

centre distributed evenly throughout the whole samples. Quantitative

differences among different time intervals at least indicated much more ALP

activity in silica-treated samples than in 3D control samples.

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Fig.5.15. Stained ALP of hFOB 1.19 on 1 bilayer coating and control samples at day 3, a) & a’) control sample with ALP stain only, b) & b’) silica treated sample with ALP stain only, c) & c’) control samples with ALP and nucleus stain together, d) & d’) silica treated samples with ALP and nucleus stain together.

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Fig.5.16. Stained ALP images of hFOB 1.19 on 1 bilayer coating and control samples at day 7, a) & a’) control sample with ALP stain only, b) & b’) silica treated sample with ALP stain only, c) & c’) control samples with ALP and nucleus stain together, d) & d’) silica treated samples with ALP and nucleus stain together.

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5.3 Conclusions

Silica nanoparticles were assembled on PCL fibrous porous scaffolds. The

surface morphology and wettability could be controlled by the coating layers.

Cytotoxicity study using fibroblasts demonstrated that all the samples

including the pristine control and the silica-treated samples were nontoxic to

cells and suitable for tissue engineering application.

MTS results and SEM images suggested that the nano-structured surface can

effectively enhance the cell attachments, and cell proliferation was much more

active for the silica-treated 3D PCL fibrous scaffold compared to the non-

treated 3D control samples. ALP activity in these silica-treated 3D PCL fibrous

scaffolds was also higher than that on control samples, revealing better

osteoconductivity and osteoindustivity.

These results suggested that nano-structured silica coating on PCL surfaces

exhibited much improved cell attachments, cell division and proliferation and

cell differentiation. These fibrous scaffolds with thin nano-structured surface

had potential applications in bone tissue engineering.

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C H A P T E R S I X

Deposition of apatite on inter-bonded 3D

fibrous scaffolds and its effect on cell

performances

Bone-like apatite was deposited on the fibre surface of 3D inter-bonded fibrous

matrices via immersing in the simulated body fluid (SBF). Oxygen or nitrogen

plasma was used to treat PCL surface before the deposition, which was found

to speed up the CaP nucleation and the formation of bioactive apatite. Two

osteoblast cell lines: human fetal osteoblast cells (hFOB1.19) and osteoblast-

like cells (Saos-2) were used to examine the influence of apatite coating on

osteoblastic cell responses. The cell morphology, cell proliferation and ALP

enzyme activity for cell differentiation were found to be affected by the apatite

coating dramatically.


In this chapter, only PCL-10 (porosity >90%, mean pore size 200 μm) was

used as matrix sample.

The PCL-10 samples were firstly pre-treated by vacuum plasma to activate the

fibre surface, followed by immediately immersing them into 1×SBF or

10×SBF at physiological pH condition (7.4) and 37 ° C for apatite deposition.

After that, the fibrous sample were thoroughly washed with MilliQ water and

dried overnight at room temperature. hFOB 1.19 and Saos-2 cells were then

seeded and cultured on the matrix.

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6.2.1 Plasma treatment

X-ray photoelectron spectroscopy (XPS) measurement was used to identify the

surface component changes before and after plasma treatment. The survey

spectra are shown in Fig.6.1. Before plasma treatment, PCL matrix only had C,

O and Si peaks. When the sample was treated with N2 plasma, a weak N peak

appeared.

Fig.6.1. XPS survey spectra of a) pristine PCL matrix and b) after N2 plasma treatment

The atomic contents of the fibrous matrices are listed in Table 6.1. For the

pristine PCL, it contained 73.99% carbon (Table 6.1). The carbon content

decreased after the plasma treatment, and the C/O ratio decreased for plasma

treatment. N2 plasma introduced N onto the surface, and the surface N content

increased a lot after N2 plasma treatment.

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Table 6.1. Elemental contents of the fibrous matrices before and after plasma treatment

Elements Untreated After treatment

C 73.99 70.98

O 23.54 24.72

N 0.602 1.568

Si 1.861 2.728

C/O 3.143 2.871

Fig.6.2 shows high resolution C1s and N1s spectra, in which all the studied

fibrous samples exhibit three main components corresponding to the aliphatic

carbon bonds (–C–C– or –C–H), carbon single bonded to oxygen (–C–OH or –

C–O–), and carbonyl functional groups (–C–O-) located at approximately

285eV, 286eV, and 289eV, respectively. The samples showed similar peaks,

suggesting that the plasma treatments had a very small influence on the surface

chemical components.

The N1s peaks at 399.90eV and 402.26eV indicated the formation of –

CH3CONH2 groups after the N2 plasma treatment.

Thus, the XPS analyses confirmed that the plasma treatments lead to slightly

different surface chemistry, which may affect the wettability and deposition of

CaP.

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Fig.6.2. Curve-fitted XPS spectra of C1s and N1s, a) pristine control sample, b & c) N2 plasma treated samples

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6.2.2 Surface morphology

As shown in Fig.6.3, without plasma treatment, the fibre surface shows no

apatite-like particles after soaking in SBF for three days. For the plasma treated

samples however, particle like structure formed on the fibre surface after

immersing in the SBF solution at the same condition. This suggested that the

plasma activated PCL surface showed improvement in the apatite nuclei

forming and deposition ability.

Fig.6.3. Surface morphology of PCL fibres after soaking in SBF solution for 3 days, a & a’) pristine control matrix, b & b’) N2 plasma treated sample.

Fig.6.4 shows the surface morphology of the PCL fibrous matrices after 21

days soaking in SBF. For the untreated pristine PCL samples, the surface

looked similar to that of the samples soaked for 3 days, and there were no

apatite-like particles found on the surface. For the nitrogen plasma treated PCL

fibre matrix, thicker particles were observed, and the particles were hard to be

recognised due to the intensive agglomeration.

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Fig.6.4. Surface morphology of PCL fibres after soaking in SBF solution for 21 days, a & a’) pristine control sample, b & b’) N2 plasma treated sample.

From the above experiments and results, N2 plasma could give a good apatite

deposition. Therefore, in the following study on using 10×SBF solution to

deposit apatite, N2 plasma was still used to pre-treat the PCL fibre matrices.

10×SBF has higher ion concentration that can speed up the formation of apatite

on the substrate surface. As shown in Fig.6.5, within 4 hours of immersing the

fibre matrix in the solution, apatite particles could be observed on the surface,

the level of which was similar to that of samples soaked in 1×SBF for three

days. When the plasma treated fibrous matrix was immersed in the solution for

one day, a thick apatite layer was generated on the fibre surface, which was

even better than soaking in 1×SBF for 21 days. The high concentration SBF

was effective in formation of apatite layers.

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Fig.6.5. SEM images of the PCL matrices after N2 plasma treatment and soaking in 10× SBF for 4 and 24 hours. a & a’) soaking for 4hrs, b & b’) soaking for 24hrs.

6.2.3 Apatite formation

The SEM EDX curves of the PCL fibre matrix after socking in the SBF

solutions are shown in Fig.6.6. The peaks Ca and P clearly appeared in the

curves, which was a proof of the CaP formation. For the 1×SBF treated

matrix, the peaks for both Ca and P were very sharp and clear, similar to the

results from literature on chitosan substrates [435]. While for 10×SBF, peaks

showed a little difference. This was because 10×SBF usually can not exactly

form bone-like apatite, which has been proposed as an issue in the literatures

[436]. This result also suggested that that apatite formed from 1×SBF was

dramatically different from that formed 10×SBF.

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Fig.6.6. SEM-EDX scanning results of the apatite treated PCL fibre samples, a) soaking in 1×SBF solution, b) soaking in 10×SBF solution.

The CaP on PCL fibre surface was also characterised through calculating the

ion concentration changes in the SBF solution. Fig.6.7 provides the calcium

and phosphate ions concentrations after immersing the fibrous samples. Both

concentrations decreased greatly with time for the treatment using 10×SBF.

However, the trend for the 1×SBF treatment was not obvious. This could be

due to the low ion concentration, and the uneven fibrous structures.

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Fig.6.7. Ca and P ion concentration changes in 1×SBF and 10×SBF solutions as a function of time

6.2.4 Water contact angle

Because the matrices had macro-pores insides them, water drops might be kept

in the pores or on the fibres. To eliminate the difference, films with exactly the

same treatment were used for water contact angle (WCA) testing after plasma

treatment. Water contact angle changed from about 80º to nearly 45º after

plasma treatment (Fig.6.8), which indicated that surface had a good water

wettability. This would assist in the whole surface contact and reaction with

SBF solution.

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Fig.6.8. Surface water contact angles before and after N2 plasma treatment, a) control sample, b) N2 plasma treated sample.

Water contact angle results for matrix samples undertaken different soaking

times in 1×SBF or 10×SBF are exhibited in Fig.6.9. After 21 days of soaking

in SBF, a layer of apatite was deposited on the fibre surface, resulting in

decrease in the water contact angle, from 106 to 55º. For the samples soaking

in 10×SBF (24 hours), a decrease trend was also obtained. Different contents

of apatite formation on the surface resulting from variable soaking times and

soaking concentrations made the initial variety of WCA among these samples.

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Fig.6.9. Water contact angles of the PCL matrices as a function of soaking time, a) 1×SBF, b) 10×SBF.

6.2.5 Cytotoxicity study

The cytotoxicity results are shown in Fig.6.10. All the samples had a very good

compatibility to fibroblast cells. 3 days’ culture resulted in the cell number

rising from the original number (1×104) to above 3×104 with the cell viability

higher than 95%. These positive results confirmed that the samples with apatite

on the surface were non-toxic to cells, and were suitable for tissue engineering

application.

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Fig.6.10. Cell cytotoxicity results using fibroblast cells, the bar shows mean live/dead cell number after 3 day culture; the line indicates the mean cell viability (p>0.05).

Fig.6.11 further shows the healthy state of fibroblast cultured on the extract of

different samples. From the images of low magnification, it can be seen that

the cells spread uniformly on the surface of TCP. From higher magnification

images, the cells are observed to interact with each other and maintain a

polygon shape, indicating their happy growing state.

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Fig.6.11. Fibroblast morphology after cultivation for 3 days in the extract media of different samples, a) Control, b) O2 plasma treated, c) N2 plasma treated.

6.2.6 Cell morphology, distribution and migration

6.2.6.1 Light microscopy

Under normal light microscopy, the cells can be distinguished from the

polymer when the samples were stained with dyes, and the low magnification

also provides information on the cell distribution and migration. As

demonstrated in Fig.6.12, the influence of apatite on the hFOB1.19 and Saos-2

become apparent only after 7 days of culture. Without the apatite layer, the

hFOB1.19 was hard to attach to the PCL fibre surface. There was very small

number of cells found on the matrix. With the help of a thin layer of apatite,

more hFOB1.19 could be observed. The hFOB1.19 cells on both the control

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and the apatite coated samples looked very healthy in terms of their shapes. In

contrast, the difference between apatite and no-apatite coated samples for Saos-

2 was very small. This is mainly because Saos-2 cells have much higher

growth and proliferation rate than the hFOB1.19 cells. Saos-2 cells grew not

only on the matrix surface, but also inside the matrix. The cells had no

preference growth orientation along the fibres, and they all looked very healthy

as well.

Fig.6.12. Morphology of hFOB1.19 and Saos-2 cells grown on PCL fibrous matrices after 7 days of culture, a &a’) hFOB1.19 on control, b & b’) hFOB1.19 on apatite treated samples; c & c’) Saos-2 on control samples, d& d’) Saos-2 on apatite treated samples. Cells were stained with 1% toluidine blue in 2% borax and 0.04% methylene blue. The scale bar for images a-d is 500 μm, and for images a’-d’ is 100 μm.

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6.2.6.2 Laser confocal microscopy (CLMS)

Fig.6.13 shows the morphology of the cells growing on the PCL fibrous

matrices. In the early days, both hFOB1.19 cells and Saos-2 cells were not easy

to be found. The cells also had small sizes but no elongation (Fig.6.13 a-1 &

b-1). At day 3, both hFOB1.19 and Saos-2 cells were found to start elongating

along the fibre length direction (Fig.6.13 a-2& b-2). This phenomenon became

clearer at day 7 (Fig.6.13 a-3). For the area with many bonded fibres, the cells

elongated less (Fig.6.13 b-3). Longer culture time led to huge amounts of cells

covering on the fibre surface (Fig.6.13 a-4 and b-4). By comparison, less

hFOB1.19 cells were found on the control sample. For Saos-2 cells, the cells

covered most of fibre surfaces. The enhancement of Saos-2 cells growth with

apatite was not observed visually.

Cells-hFOB1.19 have a similar shape to the Saos-2 cells. In biologics, they

both are classified as osteoblasts. Saos-2 cells have exactly the same

characteristic as osteoblasts and they are often called osteoblast-like cells.

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Fig.6.13. Morphologies of hFOB1.19 and Saos-2 cells in different culture periods, a1~3) hFOB1.19 on apatite treated samples after 1 day, 3 days, 7 days and 14 days of culturing; c)

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hFOB1.19 on control samples 14 days of culturing; b-1~3) Saos-2 on apatite treated samples after 1 day, 3 days, 7 days and 14 days of culturing; d) Saos-2 on control samples after 14 days.

6.2.6.3 Scanning electron microscopy (SEM)

SEM images in Fig.6.14 and Fig.6.15 exhibit the cell morphology of

hFOB1.19 and Saos-2 on PCL fibrous samples as a function of culturing time.

For hFOB1.19 cells (Fig.6.14), at day 1, the cells seemed to dislike growing on

the surface of PCL fibre, they just attached to the surface, showing similar

phenomenon on apatite coated surface. At day 3 and day 7, cells could spread

on the surface, but they tended to crouch and the cells were still small in

number. In comparison, more cells were observed to spread on the apatite

coated surface, and the cells were more easily found. With two weeks of

growth, more hFOB1.19 cells were found on the surface, showing healthy

cells-matrix interaction.

For Saos-2 cells (Fig.6.15), at day 1 and day 3, the cells were found to scatter

on the fibre surface. The cells attaching to the untreated samples were very low

in number. One week later, more Saos-2 cells distributed on the fibre surface.

Saos-2 cells had a tendency to orient along the fibre length direction on the

apatite treated samples, but not on the untreated control samples. After two

weeks of growth, Saos-2 cells covered most of the fibre surfaces, cells on the

apatite coated surface grew with an orientation along the fibre length direction.

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Fig.6.14. hFOB1.19 cultured on PCL fibrous samples at different culturing times, a-1~4) on control samples after 1 day, 3 days and 7 days, and 14 days of culturing; b-1~4) on the apatite treated samples after 1 day, 3 days, 7 days and 14 days of culturing; scale bar = 10 μm

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Fig.6.15. Saos-2 cultured on PCL fibrous samples at different culturing times, a-1~4) on control samples after 1 day, 3 days and 7 days, and 14 days of culturing; b-1~4) on the apatite treated samples after 1 day, 3 days, 7 days and 14 days of culturing; scale bar = 10 μm

6.2.7 Cell proliferation The cell proliferation was measured by MTS assay. The culture time was set at

day 1, 3, 7, and 14. The changes of live cell number at different time periods

are shown in Fig.6.16.

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When Saos-2 and hFOB1.19 were firstly cultured on these scaffolds, more

hFOB1.19 cells stayed on the matrix surface than Saos-2 cells, and the apatite

coated surface contained more cells than the untreated control samples. At day

3, cell number did not change much for the Saos-2. For hFOB1.19, however,

the cell number on the apatite-coated surface was nearly double that on non-

treated pristine sample (Fig.6.16b). This tendency sustained for at least two

weeks. The enhancement effect of the apatite layer to Saos-2 cell growth was

clear for a long time. As seen in Fig.6.16a, at day 14, the cell number on the

apatite-coated samples was much higher than that on the non-treated pristine

one. Saos-2 cells were more populous than hFOB1.19, mainly because of their

high growth speed.

Fig.6.16. Proliferation of osteoblastic cells on PCL matrices with or without apatite coating

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6.2.8 Alkaline phosphate activity

ALP activity can be calculated by ALP experiment results shown in Fig.6.17

quantitatively and Fig.6.18-Fig. 6.25 qualitatively.

Cells hFOB1.19 were cultured at 34 ºC for the first three days to obtain enough

cells. Then these hFOB1.19 cells-scaffold composites were moved to the 39 ºC

incubator for best cell differentiation, at which time, the day was set as zero.

To make the same, for Saos-2 cells, although the beginning temperature was 37

ºC, and the cell proliferation temperature was unchanged, they also allowed

proliferating for three days. For the following days, it was counted at 3 days, 7

days, 10 days, and 14 days, and at every interval, the ALP content was

measured and calculated.

As indicated in Fig.6.17 a, ALP activity for hFOB1.19 is variable with

culturing time. At the first three days, the ALP activity was high, and then

decreased at day 7. At day 10, the content increased again and reached the

highest value. All scaffold samples showed a similar trend.

Comparing to the control samples, the apatite coated fibrous samples expressed

an improved ALP activity at day 3, 10, 14. At day 7, there was no distinct

difference between the control samples and the apatite coated 3D samples. In

this study, there was not much difference between 2D control samples and 3D

control samples in terms of ALP activity.

Fig.6.17b shows the ALP results for Saos-2 cells, which are quite different

from those obtained on hFOB1.19 cells. The ALP contents of Saos-2 were

much higher than that on the hFOB1.19, even for 1 day culture, the content

was 100 times higher. Another difference for Saos-2 was that the peaks of ALP

activity appeared at day 7. Only one peak existed in all the tested groups. There

was no statistically significant difference between the apatite coated and the

non-coated samples at day 7, day 10 and day 14. The apatite improvement of

ALP activity was only found at day 3. High proliferation rate of Saos-2 was the

main reason for this phenomenon.

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Fig.6.17. ALP activity of Saos-2 and hFOB1.19 on different samples

By staining the cells with dye, ALP enzyme could be observed visually. This

assisted in understanding the difference between sample groups.

Figs.6.18 to 6.25 show all stained ALP images. Samples were firstly stained by

naphthol AS-MX phosphate/N, N dimetylformamide (DMF) solution and 0.6

mg/ml diazonium salt for ALP enzyme, then all the samples were further

stained by Mayer’s hematoxylin for cell nuclei. Pink to red colour represented

ALP generated by cells. The dark blue colour was from cell nuclei.

At day 3, there was already some ALP observed in the apatite coated samples

grown with hFOB1.19. When looking at the images of control samples

(Fig.6.18 c and c’), some cells were observed on the matrices, but no ALP

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(Fig.6.18a and a’). These results were in accordance with that from ALP

content mentioned previously: apatite might have a good effect on the cell

differentiation at the early stage.

For the Saos-2 cells at day 3, the influence of apatite on ALP activity was

evident (Fig.6.22). Cells or ALP enzyme was found to distribute throughout

the matrix, which was different to the untreated pristine samples.

At day 7, the ALP activity of hFOB1.19 was very low. There were only several

red spots in the matrices, for both control and the apatite coated samples

(Fig.6.19).

For Saos-2, the ALP activity was much higher than that of hFOB1.19. The

matrices had red colour on some fibres. Such a good ALP activity maintained

for all the tested time intervals (Fig.6.23 & Fig.6.24), and nearly the whole

matrices looked red after 14 days of culture (Fig.6.25)

Longer culturing time gave an easy observation of hFOB1.19 cells on the

matrices. The ALP activity, however, was still very low and not much red

colour was shown (Fig.6.20& Fig.6.21). This could be attributed to two

factors: hFOB1.19 had really low cell proliferation rate, even after 14 culturing

days, therefore not enough cells were grown on the matrices, and hFOB1.19

cells were hard to differentiate at 39 ºC.

The results on ALP contents and stained ALP images suggested that ALP

activity could be enhanced by apatite coating all the time for hFOB1.19 cells,

but only at the early stages for Saos-2 cells.

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Fig.6.18. Images of hFOB1.19 stained with alkaline phosphate on PCL fibrous samples at day 3, a) & a’) control sample with ALP stained only, b) & b’) apatite treated sample with ALP stained only, c) & c’) control samples with ALP and nucleus stained together, d) & d’) apatite treated samples with ALP and nucleus stained together.

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Fig.6.22. Images of Saos-2 stained with alkaline phosphate on PCL fibrous samples at day 3, a) & a’) control sample with ALP stained only, b) & b’) apatite treated sample with ALP stained only, c) & c’) control samples with ALP and nucleus stained together, d) & d’) apatite treated samples with ALP and nucleus stained together.

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6.3 Conclusions

A thin layer of apatite was deposited on the PCL fibre surface by pre-treatment

of fibre with N2 plasma and soaking in SBF solution. The plasma pre-treatment

sped up the apatite formation. The apatite formation was confirmed by SEM

and EDX, and the apatite coating made the fibre surface more hydrophilic.

These apatite-coated samples were proved to be biocompatible to fibroblast

cells with no cytotoxicity. They also improved the cell attachments because of

the increased surface roughness and wettability. The apatite coating improved

the cell proliferation, because the apatite surface was more bioactive for cells

to grow than pristine PCL surface. This good cell differentiation was also

indicated by the increase in the ALP activity. In summary, apatite was

successfully applied on PCL 3D fibrous matrices, and this coating was proved

to effectively enhance the adhesion, proliferation and differentiation of

osteoblast cells.

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C H A P T E R S E V E N

Inter-bonded 3D fibrous scaffolds with

antibacterial surface coating and cell growth

performances

In this chapter, 3D fibrous scaffolds were surface coated with antibacterial

agents, silver and polyhexamethylene biguanide hydrochloride (PHMB). The

cytotoxicity of the coated fibrous matrices was studied. The antibacterial

activity was confirmed, and the compatibility to fibroblast cell growth was

examined.


PCL fibrous matrices PCL-10 (porosity >90%, mean pore size 200 μm) were

used. The antibacterial agents (Ag nanoparticles) were coated on the surface by

an electroless plating method. Briefly, two solutions separately containing Ag+

source and the reducing agent were mixed together to initiate the silver-mirror

reaction in the presence of the fibrous matrices. The silver generated was

deposited onto the fibre surface. Separately, PHMB was coated on fibre surface

by the layer-by-layer electrostatic assembly technology. To do this, PAA

abundant with –COOH groups was applied onto the fibre surface, followed by

assembling positively charged PHMB under pH 8.0. Then another negatively

charged PAA was introduced, followed by the PHMB again.

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7.2.1 Electroless plating of silver

7.2.1.1 Electroless plating The plating time was studied from 2 minutes to 40 minutes and the optical

microscopy images of the coated fibrous matrices are shown in Fig.7.1. It was

found that 10 minutes of plating was enough to finish the reaction and also

avoided the aggregation of silver particles on the fibre surface, although these

precipitations could be washed away by ethanol or distilled water.

Fig.7.1. Photographs of PCL matrices with silver coated surfaces. The scale bar=100μm.

Silver nanoparticles show distinct surface Plasmon absorption from the visible

to the near infrared region [437]. As shown in Fig 7.2 a~c, the white fibrous

sample after 2 minute treatment becomes yellow. Longer plating time made the

sample become brown and even grey eventually. The dark brown colour

indicated the formation of silver nanoparticles. The SEM images shown in Fig

7.2 d~f indicates that a fine nano-silver coating evenly covers the entire fibre

surface. Higher magnification image illustrated that the deposited Ag was in

the form of nanoparticles.

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Fig.7.2. Photographic and SEM images of PCL matrices, a) uncoated PCL matrix; b & c) PCL matrices coated with silver nanoparticles at different plating times (low & high); d) untreated PCL matrix; e & f) Ag-coated PCL matrices in low and high magnification.

TEM images in Fig.7.3 indicate that a very thin layer of Ag coating is formed

on the PCL fibre surface. The thickness of the Ag coating was around 100~120

nm, after 5 minute Ag plating treatment.

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Fig.7.3. TEM images of Ag-plated PCL matrices. The silver coating layer is around 100~120 nm in thickness.

7.2.1.2 Particle-size and distribution Based on the SEM images, the Ag particle size was calculated. Fig.7.4a

provides the particle size result measured from SEM images. The mean particle

size was about 75.04 nm with a standard deviation of 25.74 nm. The

distribution of Ag nanoparticle size fitted the Gaussian distribution (R2 =

0.97).

The particle size was verified by DLS. Samples were dissolved in

dichloromethane, and the solution was then tested. Fig.7.4b shows a typical

testing result. The mean particle size measured by this method was near 200

nm, which was much larger than that measured from SEM images.

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Fig.7.4. Particle size distribution histograms tested by a) SEM image and b) DLS (dynamic light scattering)

7.2.1.3 Silver crystalline X-ray diffraction (XRD) patterns of the electrolessly plated Ag nanoparticles

on the PCL matrix are shown in Fig. 7.5. The peaks at 2θ = 38.32°, 44.04°,

64.64° and 77.42° were assigned to the (111), (200), (220), and (311) reflection

lines of Ag particles, respectively [438-439].

Fig.7.5. XRD pattern of Ag coated PCL matrix

7.2.1.4 Ultraviolet visible (UV-Vis) absorption

Fig.7.6. UV-Vis spectra of Ag solution released from the coated fibrous samples with single layer coating as a function of soaking time

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Fig.7.6 shows the UV-Vis spectra of Ag nanoparticles released from the coated

PCL fibrous matrix, with the maximum absorption peak at about 420 nm,

which is consistent with the absorption of silver nanoparticles [440]. Since the

release of silver nanoparticles is time dependent, the peak increased gradually

with the increase in the immersion time.

7.2.1.5 Antibacterial properties Fig.7.7 gives the antibacterial ratio of different samples. All the samples

demonstrated strong antibacterial properties. For the sample with 10 min Ag

plating and washed with ethanol and water before use, more than 99% bacteria

were killed within 3hr. With increasing the plating time, a high percentage of

bacteria were killed.

At the same treatment condition (10 mins double layers coating), the unwashed

samples had a high antibacterial ratio compared to the washed samples because

washing removed some silver nanoparticles off the matrix, which decreased the

antibacterial property of the matrix. Similarly, when increasing the plating time

from 10 mins to 30 mins, more particles were assembled on the fibre surface,

leading to stronger antibacterial effect. Plating more layers of Ag also could

improve the bacterial inhibition function, because more Ag ions would be

released from the increased surface area.

Fig.7.7. Antibacterial efficacy of different sample groups as a function of time (Escherichia coli, E.coli)

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The Ag particle coating changed the E.coli surface from smooth to irregular as

indicated in literatures [441-442]. Here the Ag particles on the PCL fibre

surface damaged the normal biological function (Fig.7.8).

Fig.7.8. Effects of nanoparticles on Escherichia coli (E.coli), a) after contact with the Ag-coated PCL fibrous matrix for 7 hours, cell surface was surrounded by Ag particles; b) E.coli cells had an irregular surface (cellular shrinkage).

7.2.1.6 Cytotoxicity Fig.7.9 shows the cytotoxicity testing results of the Ag-plated fibrous matrices.

Without any sterilization treatment, the scaffolds showed more than 90%

fibroblast cell viability after 5 days of cell culture. This result was quite similar

to that of the control matrix. This suggests that the silver coating has very low

toxicity to human mammalian cells and the silver nanoparticles coated fibrous

matrices can be used for tissue scaffolding applications.

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Fig.7.9. Cytotoxicity test results by culturing fibroblasts in extract liquids from the Ag-plated fibrous samples and control samples (p>0.05)

7.2.1.7 Culture of fibroblasts Cell culture test was conducted on the Ag coated PCL matrix. As shown in

Fig.7.10a, after 1 day of cell culturing, although the number of cells on the

matrix is slightly lower than 2D control due to the static seeding limitation, the

mean viable cell number on the Ag-coated matrix is equivalent to the 3D PCL

control matrix. After 3 days and 7 days of culture, the cell number in the Ag-

coated fibrous matrix nearly had no changes with time, which was abnormal

when comparing to the 3D control and 2D control samples (Fig.7.10b&c).

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Fig.7.10. MTS results of fibroblasts cultured on PCL matrices and nAg-anchored PCL matrices

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This result suggested that the Ag-coated reduced the cell growth. To improve

the cell growth on the Ag-coated fibrous matrix, the sample was pre-coated

with a thin layer of cellulose acetate using a dip coating method. The cell

culture result was also included in Fig.7.10. It was shown that a thin layer of

polymer was still not able to improve the cell growth much.

7.2.2 Electrostatic assembly of PAA-PHMB 7.2.2.1 Confirmation of PHMB To verify the assembly of PHMB on the 3D PCL scaffolds, the samples were

stained with a dye (Eosin). Due to the reaction between PHMB and the dye,

red colour would be developed when the dye molecules interacted with PHMB.

More dye-PHMB interaction led to deeper colour. Fig.7.11 inset image shows

the PCL fibrous matrices having different layers of PHMB stained by Eosin

dye. Without PHMB layer, no absorption peak appeared in the range of

500~570 nm. With increasing the PHMB assembling layers, the matrices had

stronger absorption peak at 540 nm.

Fig.7.11. Visible spectra of the stained fibre samples with different PHMB layers

PHMB coatings were confirmed by FTIR. As shown in Fig.7.12 the vibration

peaks at 3375 cm-1, 1625 cm-1, and 1575cm-1 are corresponding to NH, NH2

and NH4+, when 5 layers or more layers of PHMB are assembled onto PCL

fibres.

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Fig.7.12. FTIR spectra of PCL fibrous samples before and after PHMB assembling

7.2.2.2 Water contact angle Since the matrices had macro-pores, water drops might be kept in the pores or

on the fibres. To eliminate the difference, films with exactly the same

treatment were used for WCA testing. It was assumed that contact angle on

films can reflect the water wettability of fibrous matrices.

Water contact angle of PHMB coated samples were measured on the PCL film

with exactly the same treatment method. Results are shown in Fig.7.13.

Without any treatment, the PCL film was hydrophobic with a mean contact

angle of 78.0º. When one layer of PHMB was introduced on the surface,

contact angle decreased to 67.9º. Increasing the PHMB layers on the surface

lowered the contact angles. When the coating was 7 bilayers of PHMB, the

contact angle reached as low as 52.7º, which was ideal for tissue scaffold

applications [443].

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Fig.7.13. Water contact angles of different layers of PHMB coated films

7.2.2.3 Washing durability of PHMB coatings Fig.7.14 shows the Visible absorption of the PHMB coated fibrous samples

before and after sterilising process (washing with 70% ethanol and distilled

water for several times, detailed process was described in chapter 3). The

PHMB peak was still very high after sterilising treatment. The inset digital

photo shows the colours before and after sterilising treatment.

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Fig.7.14. Visible spectra of the PHMB-coated samples before and after sterilising treatment, Inserted images were digital photos of the corresponding samples.

7.2.2.4 Antibacterial properties The antibacterial results of the PHMB coated fibrous samples are provided in

Fig.7.15. The initial seeding number of bacteria was 7.575×106, while after 4hr

culturing and growth on pristine control sample, the number increased to

1.87×109, which was regarded as no bacterial killing property. For the fibrous

sample containing 1 layer of PHBM coating, although there was a certain

reduction of E.coli bacteria (The reduction rate was 71.88% and 80.13%

respectively), it was still not enough for the self-disinfection purpose. Some

bacteria were still found on the surface after 5 hours of incubation. When more

layers of PHMB were assembled on the surface, the retarding rate of bacterial

could reach 99.99%, as indicated in Fig.7.15. Especially for 7 layer coating,

most of plates showed no living E.Coli.

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Fig.7.15. Bacterial growth and antibacterial rate of fibrous matrices having different layers of PHMB treated samples after 4hrs contact

Although the 3, 5 and 7 bilayers of PHMB assemblies made it possible to kill

all bacteria in 5 hours, it was still necessary to investigate the long-term

antibacterial efficacy. At the first half hour, the bacteria on the control samples

were recovered from the liquids and adjusted to the environments. There was

no increase in logarithm value. In contrast, the bacteria on the PHMB treated

samples (7 bilayers) were killed immediately when they contacted with PHMB

coated surface, which was tested in 10 mins. There were no bacteria on the

PHMB-treated fibre surface. With time, bacteria on the control samples

propagated quickly, while no changes happened on the PHMB treated samples.

It proved that with PHMB on the surface, no bacteria grew.

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Fig.7.16. Bacterial growth as a function of time

7.2.2.5 Cell cytotoxicity The cytotoxicity study was done using fibroblast cells according the method

mentioned in Chapter 3. The extract media was used. The cell viability and the

cell growth morphology after 3 days culture are shown Fig.7.17 and Fig.7.18.

All the samples tested proved no toxicity to cells, with cell viability of more

than 95%, and cell morphology sustaining in a healthy polygon shape.

Fig.7.17. Cell viability results of the fibrous samples with different layers of PHMB (p>0.05)

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Fig.7.18. Cell growth morphology, a) & a’) on control samples, and on the PHMB assembled fibrous samples with b) & b’) 1 layer PHMB coating, c) & c’) 1 bilayer PHMB/PAA coating, d)

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& d’) 3 bilayers PHMB coating, e) & e’) 5 bilayers PHMB coating and f) & f’) 7 bilayers PHMB coating

7.2.2.6 Cell growth A simple 7 day culture of fibroblast was conducted for evaluating the cell

culture adaptability. MTS assay was used to identify the real viable cell

number after 7 day culture. The cell number data is given in Fig.7.19.

Although there were little differences in the mean cell number after 7 days

culture, the cell number showed no differences between the pristine samples

and the PHMB assembled samples. When taking into consideration of different

layers, 7 bilayers of PHMB coating had less cell number compared to samples

coated with fewer PHMB layers (Fig.7.19).

Fig.7.19. Mean cell number of fibroblasts on different samples after 7 days of cell culturing

Confocal microscopy images provided another proof of fibroblasts growth

state. As shown in Fig.7.20, in one week of cell culture, it was hard to observe

separate cells from these images. Cells already grew into multiple layers on the

fibre surface.

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Fig.7.20. Confocal microscopic images of fibroblasts after 7 days of cell culture, a) pristine fibre sample, and the fibrous sample with b) 1 PHMB layer, c) 1 PHMB/PAA bilayer, d) 3 PHMB/PAA bilayers, e) 5 PHMB/PAA bilayers, f) 7 PHMB/PAA bilayers. Scale bar = 100 μm

Fig.7.21 shows the SEM images of the fibrous sample after 7 days of cell

culture. For the untreated pristine samples, due to the hydrophobic surface, the

cell attachment on the fibre surface was not uniform. Some surface parts

contained much less cells compared to other parts. For the samples having 1

layer of PHMB coating, the cells covered the entire fibre surface. The fibrous

samples with 1 bilayer, 3 bilayers, and 5 bilayers, showed a similar tendency.

With increasing PHMB layers, cells seemed to not grow as usual. In addition,

cells grew not only on the fibre surface, but also between the fibres.

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Fig.7.21. SEM images of fibroblasts after 7 days of culturing on fibrous matrices, a) pristine fibre sample, and the fibrous sample with b) 1 PHMB layer, c) 1 PHMB/PAA bilayer, d) 3 PHMB/PAA bilayers, e) 5 PHMB/PAA bilayers, f) 7 PHMB/PAA bilayers. Scale bar = 100 μm

7.2.2.7 Cell proliferation Cell proliferation study was conducted on 7 bilayers of PHMB coated samples,

which had the best antibacterial properties and cell effectiveness.

The samples with 7 layers of PHMB assembly had similar number of cells

compared to the control sample in the first three days, and even after a longer

culture period. The cell growth speed on both samples was increased until 7

days, after which the cell number reduced (Fig.7.22).

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Fig.7.22. Fibroblast number on the control and the PHMB assembled (7 bilayers) samples.

SEM images in Fig.7.23 reveal the cell morphology of fibroblasts on the

fibrous samples after 3 and 7 days of cell culture. On day 3, very few cells

were observed on the matrices. After 7 days of culture, fibroblasts could form

multiple layers around the fibres of control samples. However for the PHMB

treated fibrous matrices, fibroblasts were found to be in bunches and existed

mainly on the boundary between fibres, rather than on the fibre surfaces.

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Fig.7.23. SEM images of fibroblasts after 3 and 7 days of culture, a) pristine fibre sample-3 days’ culture, b) pristine fibre sample-7 days’ culture, c) fibrous sample with 7 PHMB/PAA bilayers-3 days’ culture, and d) fibrous sample with 7 PHMB/PAA bilayers-3 days’ culture. Scale bar = 50 μm

7.2.3 Comparison

Electroless plating provided a quick method to assemble a thin silver

nanoparticle layer on the surface of PCL matrices, which could be achieved in

10 minutes. The content of silver on the surface was dependent on plating time

and coating layers. In contrast, the procedure for PHMB/PAA electrostatic

self-assembly on PCL fibrous matrices was much complicated and time-

consuming, but the content of PHMB could be controlled by changes of layers.

In the view of antibacterial properties, silver coating was more effective than

PHMB coating. Silver killed 99.99% of original seeded bacterial in 4 hours,

while for PHMB coating, only samples with larger number of layers (3 bilayers,

5 bilayers & 7 bilayers) had relative good bacterial killing ability.

Although samples treated with the two methods showed no difference in cell

toxicity study, the cell growth experiment indicated that silver treated samples

were only good for short time cell culture (1 day), but PHMB treated samples

could provide good environment for cell growth till 14 days.

7.3 Conclusion

Two antibacterial agents, silver and PHMB were applied to the PCL fibrous

matrix respectively by electroless plating and layer-by-layer electrostatic self-

assembling methods. Antibacterial experiment indicated that these treatments

were effective for killing bacteria. For cell culture, the samples coated with

silver nanoparticles were proved to be good for short time, e.g. 1 day, culture,

longer culture time would result in death of cells. The cells on the PHMB

assembled fibrous samples showed much better growth performance. The

layers of PHMB influenced cell growth mainly in terms of cell number rather

than cell morphology. Seven bilayers of PHMB coating gave a small cell

population after 7 days of culture, and the cell maintained not only on the

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surface but also on the fibre-junction area. Cell growth speed was higher in the

first two weeks, and it slowed down afterwards on PHMB treated samples as

well as on control samples. The PHMB may be a good antibacterial candidate

for self-disinfection tissue scaffolds.

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C H A P T E R E I G H T

Summary and Outlook

8.1 Conclusions In this thesis, a new scaffold fabrication method combining fibre melt

processing with particle leaching technique has been developed, which has

great potential for tissue engineering. The scaffolds prepared not only have the

basic characteristics of fibrous scaffolds, but also provide stable mechanical

properties with controlled and interconnected pores which make them an

excellent candidate for tissue engineering scaffolds.

Based on this novel 3D scaffold, surface techniques were used to improve

surface roughness, surface bioactivity and surface antibacterial properties. The

surface modified 3D scaffolds revealed enhanced cell performances. The main

conclusions from this PhD study are highlighted below.

1) 3D inter-bonded fibrous scaffolds were fabricated by combining melt-

bonding of nonwoven with a particle leaching process. The melt procedure

ensures good mechanical properties and stability of the 3D scaffolds, and the

use of short fibres facilitates to form 3D structure with isotopic pores. The

particle leaching technique allows controlling the pore size and porosity. The

resulting 3D structure was demonstrated to be an excellent candidate for in

vitro cell culture.

2) Layer-by-layer electrostatic coating was used to uniformly generate a thin

layer of silica nanoparticles on the PCL fibrous matrices. This coating

technique allows good control of the nanoparticle layers on the fibre surfaces.

This coating not only changes the surface roughness and water wettablity, but

also provides a profound benefit for cell attachment, proliferation and

differentiation.

3) Fibrous scaffold has also been modified with apatite to enhance the

osteoblasts bioactivity. PCL fibres could be activated by vacuum plasma

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treatment and subsequently immersing into SBF solution. The coating resulted

in enhanced proliferation rates for bone cells and alkaline phostatase formation.

4) Antibacterial agents have also been introduced to the surface of 3D fibrous

scaffold by electroless plating of Ag or layer-by-layer coating of PHMB. The

antibacterial coating showed non toxicity but only a short term cell

biocompatibility for Ag, while maintaining normal cell function till 14 days for

PHMB.

This research has shown that 3D inter-bonded fibrous scaffolds with functional

surface have great potential for biomedical applications.

8.2 Suggestion for Future Work This research opens up a new avenue of using 3D fibrous matrices with

stabilised porous structure and nano-structured surface for tissue scaffolding

applications. However, many questions remain to be answered. For example,

the materials selection, the large-scale and uniform dynamic seeding of cells,

the gene expression of silica nanoparticles treated surface, the possibility of

integrating nano-tomography, bioactive surface and antibacterial properties

into one step coating, effect of growth factor and most importantly, the in vivo

applicability and the long-term effect to tissue and organ after implantation.

Tackling the following technical hurdles may form the focus of further studies.

1) The main material used in this study, polycaprolactone (PCL), is a

thermoplastic polymer. There are many other kinds of polymers, such as

natural and nondegradable polymers, that need to be examined. The relevant

cell culture experiments should also be conducted.

2) Dynamic seeding gives better efficiency and cell distribution than static

seeding method for cell culture. This seeding method was not used due to the

lack of facility. It should be performed in future studies.

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3) The effect of silica nanoparticles on the gene expression is still not clear,

although they are nontoxic to cells. What’s more, the in vivo experiment on

animal is necessary before clinical trials. The apatite coated samples face

similar challenges that may be tackled.

4) For antibacterial surface treatment, more experiments are needed to examine

the generality of the antibacterial activity. That is to say, when it imparts the

sterilising property, it may simultaneously induce biological responses to

human mammalian cells, resulting in the failure of the whole system. This

aspect requires further study.

5) The possibility of growth factor induction into the 3D inter-bonded fibrous

scaffolds, the influence of different growth factors on scaffold performance

both in vitro and in vivo, and the delivery mechanism of the growth factor are

also critical issues to be examined.

6) In the long term, in vivo implantation experiment is necessary before any

clinical trials on human body can be contemplated.

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