deposition of biocompatible thin films using glow...
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DEPOSITION OF BIOCOMPATIBLE THIN FILMS
USING GLOW DISCHARGE
Since 1864
DOCTOR OF PHILOSOPHY
in
PHYSICS
by
SEHRISH SALEEM
2
2009-2015
79-GCU-PhD-PHY- 09
DEPARTMENT OF PHYSICS
GC UNIVERSITY LAHORE
DEPOSITION OF BIOCOMPATIBLE THIN FILMS
USING GLOW DISCHARGE
Submitted to GC University Lahore in partial
fulfilment of the requirements for the award of
degree of
DOCTOR OF PHILOSOPHY
in
3
PHYSICS
by
SEHRISH SALEEM
2009-2015
79-GCU-PhD-PHY- 09
DEPARTMENT OF PHYSICS
GC UNIVERSITY LAHORE
4
So Exalted be God, the True King! And do not hasten with
the Reading (Qur'an) before its revelation is accomplished to
thee, and say, `my Lord, increase me in knowledge.
(114) TA HA (20)
5
Dedications
To my Teachers, Parents, Sisters, Husband and Son
The reason of what I become today, thanks for your
great support and continuous care.
6
DECLARATION
I, Sehrish Saleem, Reg. No 79-GCU-PhD-PHY- 09 student of Prof. Dr. Riaz Ahmad in the subject of
Physics, hereby declare that the matter printed in the thesis titled “Deposition of biocompatible thin
films using glow discharge” is my own work and has not been printed, published and submitted as
research work, thesis or publication in any form in any University, Research Institution etc. in Pakistan
or abroad.
Signature of Deponent
Dated: Sehrish Saleem
7
RESEARCH COMPLETION CERTIFICATE
It is certified that the research work contained in this thesis titled “Deposition of biocompatible thin
films using glow discharge” has been carried out and completed by Ms. Sehrish Saleem, Reg No. 79-
GCU-PhD-PHY- 09 under my supervision.
Supervisor:
Dr. Riaz Ahmad
Professor
Chairperson
Department of physics
GC University Lahore
Co-supervisor:
Dr. Muhammad Nouman Sarwar
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Qureshi
Associate Professor
Department of physics
GC University Lahore
Date:
Dr. Riaz Ahmad
Professor
Chairperson
Department of physics
GC University Lahore
Controller of Examinations
GC University Lahore
ACKNOWLEDGEMENT
“One who not thanks to people not thanks to Allah”.
All praise to Allah Almighty, the most Merciful, Beneficent and Compassionate who gave me ability,
strength and knowledge to complete my research work successfully. I also offer my humblest gratitude
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to Holy Prophet (peace be upon him), who is a source of knowledge and symbol of guidance for
humanity.
This piece of work that is now encompassing your vision would never come to existence without the help,
motivation and supervision provided by Prof. Dr. Riaz Ahmad, Chairperson of department of physics.
He developed enthusiasm at the start, provided stimulation when i was nowhere, showed guideline when
I was confused, offered help I was stuck up, had strict supervision when I was faultering and reviewed
the thesis to the extent and it became presentable.
I am very thankful to all the teachers and the staff members for their benevolent behavior especially Dr.
Hassan A. Shah, Dr. M. Nouman Sarwar, Dr. Salamat Ali.
I am very grateful to Prof. Dr. Paul. K. Chu for providing me opportunities to work at Plasma
Laboratory, Department of Physics and Materials Science, City University of Hong Kong, Kowloon,
Hong Kong, China, where I did a part of my research. Prof. Dr. Paul has always been very nice and
generous to help me during my research work, during my stay in Pakistan as well as in Hong Kong.
I must acknowledge the financial support of Higher Education Commission of Pakistan, during my
PhD research within Pakistan and abroad.
I wish to thanks my entire lab. fellows Dr. Shoib Shah, Dr. Tousif Hussnain, Dr. Zeeshan Adil Umer,
Dr. M. Ali Hussnain, Mr. Saqib Jabbar, Mr. M. Sabtain Abbas, Mr. M. Farrukh Ethesham
Mubarik, Mr. Ali Shahzad and friends Dr. Uzma Ikhlaq, Ms Nida Khalid, Dr. Umm-i-kalsoom,
Dr. Laila Zafer, Ms. Alvina Butt, Ms. Tahira Shujah, Ms. Khizra Abbas, Ms. Rabia Ahson, , Ms.
Saira Liaqat, Ms. Mariam Khan, Dr. Nazneen Bangash, Ms. Sobia Sehr, and Ms. Pakiza Iqbal for
their moral support, encouragement, friendly behavior and for bearance me during my work and study.
I want to give my thanks to all staff members and students in the plasma lab in City University of Hong
Kong, especially, Dr. Ricky Fu, Dr. Guosong Wu, Dr. Xuming Zhang, Dr. W. H. Jin, Dr. Xu.
Ruizhen, Dr. Jamesh M. Ibrahim, Dr. Penghui Li, Mr. W. H. Wang, Dr. Patrick Neumann, Mr.
Ming Zhang, Mr. H. S. Cheng, Mr. K. H. Lai, Mr. Chenxi Wang, Mr. S. W. Wong, for their precious
discussion and helpand with whom I have worked together as a cooperating and progressive team.
I would like to appreciate the cooperative behavior of my Mother–in-Law and Mrs. Sajjad as well for
helping and giving me an easy environment during my research work. How can I forget to thank my
loving parents and sisters (I have no words to explain their love), who made it possible for me to
complete this work. It would not be possible to attain this goal without my parents love and prayers. Both
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have come to mean more to me in my life. I would not be the person I am today without both of their
influences upon me, their patience with me and their support and encouragement in all that I do.
Last but not least, a very special thanks to my husband Rana Muhammad Ayub and my cute, innocent
son (Muhammad Usman Ayub). His support, encouragement, quiet patience and unwavering love were
undeniably the basis upon which the last year of my life has been built. His tolerance of my occasional
rude and aggressive moods is a testament in itself of his unyielding devotion and love. I also thank to his
parents and Brother Rana Muhammad Sajjad.
To all these and all those I have left out of this page I say, “Thank you!”
Sehrish Saleem
TABLE OF CONTENTS
Declaration v
Research Completion Certificate vi
Acknowledgements vii
List of Figures xii
List of Tables xvi
Publications xix
Abstract xx
11
CHAPTER 1 INTRODUCTION 1
1.1 Biomaterials 1
1.2 Types of Biomaterials 5
1.3 Plasma based technology 9
1.4 Application of plasma based technology to biomaterials 10
1.5 High vacuum magnetron sputtering system 10
1.6 Plasma immersion ion implantation & Deposition 11
1.7 Layout of the Thesis 13
1.8 Reference 15
CHAPTER 2 LITERATURE SURVEY 18
2.1 Literature of Oxynitride Film 18
2.2 Literature Survey of ZrO2 22
2.3 Literature Survey of Ti-Al-O 25
2.4 Reference 28
CHAPTER 3 EXPERIMENTAL SETUP AND DIAGNOSTICS
TECHNIQUES 31
3.1 Introductions 31
3.2 High Power Plasma Magnetron Sputtering System 31
3.3 Plasma Immersion Ion Implantation and Deposition System 32
3.4 Pulsed DC Magnetron Sputtering System 35
3.5 Sample Preparation 36
3.6 Thin Film Characterization Techniques 37
3.6.1 X-ray Diffractometer 34
3.6.2 Scanning Electron Microscope 39
12
3.6.3 Energy Dispersive X-ray Spectroscopy (EDS or EDX) 41
3.6.4 Atomic Force Microscopy (AFM) 43
3.6.5 X-Ray Photoelectron Spectroscopy (XPS) 46
3.6.6 Nanoindentation 49
3.6.7 Universal Testing Machine 51
3.6.8 Wear test 52
3.6.9 Preparation of Simulated Body Fluid 55
3.6.10 Cell Culture 56
3.7 Reference 58
CHAPTER 4 RESULTS AND DISCUSSIONS 60
4.1 Synthesis of Titanium Oxynitride on NiTi 60
4.1.1 Introduction 60
4.1.2 Experimental Detail 61
4.1.3 Results and Discussion 62
4.2 Nano-structured Zirconium Oxide film using PIII & D
4.2.1 Introduction 73
4.2.2 Experimental Setup 74
4.2.3 Results and Discussion 75
4.3 Structural and mechanical properties biocompatibility of
Ti-Al-O 94
4.3.1 Introduction 94
4.3.2 Experimental Setup 95
4.3.3 Result and Discussion 96
4.4 References 112
13
CHAPTER 5 Conclusions and future work 119
5.1 Conclusions 119
5.2 Future Suggestions 120
BIBLIOGRAPHY 121
LIST OF FIGURES Figure: 1.1 Factors and their effect on biocompatibility 2
Figure: 1.2 Value of Elastic Modulus of Metallic Materials 4
Figure: 1.3 Applications of Biomaterials in different parts of body 8
Figure: 1.4 Photographs of cardiovascular components such as (a) rotary heart
pumps (b) artificial heart valves and (c) stents
9
Figure: 1.5 Working diagram of Ultra high vacuum PVD system 11
Figure: 1.6 Working diagram of PIII-D system 12
Figure: 3.1 Working diagram of ultra-high vacuum PVD system in
CityU Hong Kong
32
Figure: 3.2 Schematic diagram of plasma immersion ion implantation (PIII) 34
Figure: 3.3 Working diagram of Plasma immersion ion implantation and
Deposition System in CityU Hong Kong
35
Figure: 3.4 Schematic of chamber with probe ports 36
Figure: 3.5 Braggs Diffraction 38
Figure: 3.6 Schematic representation of XRD 39
Figure: 3.7 Emission of electrons and x-ray, when beam of electron fall on
specimen
40
Figure: 3.8 Systematic diagram of Scanning electron microscope 41
Figure: 3.9 Schematic representation of electron-beam sample interaction 42
Figure: 3.10 Schematic representation of SEM with both energy dispersive and
wavelength dispersive X-Ray spectroscopy systems
43
14
Figure: 3.11 Schematic diagram of Atomic Force Microscopy 44
Figure: 3.12 Graph of force versus distance 46
Figure: 3.13 Schematic representation of principle of XPS 47
Figure: 3.14 Basic structure of X-Ray Spectroscopy system 48
Figure: 3.15 Schematic diagram of nanoindentor 50
Figure: 3.16 Nanoindentor XP 51
Figure: 3.17 Universal Testing Machines 52
Figure: 3.18 Schematic representations of pin-on-disc wear test system 54
Figure: 3.19 The basic layout for cell culture hood 57
Figure: 4.1 XRD patterns of TiOxNy films deposited at different oxygen flow
rates
63
Figure: 4.2 XPS spectrum showing TiOxNy film growth 65
Figure: 4.3 High resolution XPS spectra of Ti2p obtained from TiOxNy films
prepared at different oxygen flow rates: (a) 15 sccm, (b) 10 sccm,
and (c) 5 sccm
66
Figure: 4.4 High-resolution XPS spectra of O 1s obtained from TiOxNy films
prepared at different oxygen flow rates: (a) 15 sccm, (b) 10 sccm,
and (c) 5 sccm
67
Figure: 4.5 High-resolution XPS spectra of N 1s obtained from TiOxNy
films prepared at different oxygen flow rates: (a) 15 sccm, (b) 10
sccm, and (c) 5 sccm
68
Figure: 4.6 AFM surface morphology: (a) untreated, (b)5 sccm, (c)10 sccm and
(d) 15 sccm oxygen flow rate
70
Figure: 4.7 Fluorescence microscopy images after culturing for 3 days: (a) 5
sccm, (b) 10 sccm, and (c) 15 sccm oxygen flow rate
72
Figure: 4.8 Survey scan spectrum of (a) untreated Ti6Al4V (b) ZrO2 deposited
Specimens
77
Figure: 4.9 Deconvoluted XPS spectrum of (a) Ti 2p (b) Zr 3d (c) O1s 78
15
Figure: 4.10 X-ray photoelectron spectroscopy (XPS) of (a) Ti 2p, (b) Al 2p, (c)
V 2p,(d) Zr 3d, (e) O1s as a function of sputtering time, specimens
treated at 15 kV (A- Series), at 20 kV (B-series), at 25 kV (Cseries)
79
Figure: 4.11 Graphical representation of Atomic concentration versus sputter time
(a) 15 kV (b) 20 kV (c) 25kV
80
Figure: 4.12 AFM micrographs of deposited film (a) untreated (b) 15 kV (c) 20
kV (d) 25 kV
81
Figure: 4.13 Variation in thickness at different voltages 82
Figure: 4.14 Friction Coefficient as a function of no of cycles. (A) at 2 N (B) at 85
7 N (a) Untreated, (b) 15 kV, (c) 20 kV, (d) 25 kV
Figure: 4.15 Variation of wear rates with applied voltages at (a) 2 N (b) 7 N load 86
Figure: 4.16 Depth profiles of wear track as a function of different voltages (A)
2 N, (B) 7 N, (a) untreated, (b) 15 kV, (c) 20kV, (d) 25 kV
87
Figure: 4.17 Variation in depth and area of grooves at different applied voltages,
(a) 2 N, (b) 7 N
88
Figure: 4.18 SEM micrographs of wear track of untreated and treated specimens
at 2 N load (a) untreated (b) 15 kV, (c) 20 kV, (d) 25 kV
90
Figure: 4.19 SEM micrographs of wear track of untreated and treated specimens
at 7 N load (a) untreated (b) 15 kV, (c) 20 kV, (d) 25 kV
91
Figure: 4.20 Nanohardness as a function of indentation depth (a) untreated (b)
15 kV, (c) 20 kV, (d) 25 kV
93
Figure: 4.21 Surface hardness as a function of applied voltage 93
Figure: 4.22 Systematic diagram of Plasma generating system 96
Figure: 4.23 XRD patterns of untreated and Al-O deposited specimens at
different plasma excitation powers
99
Figure: 4.24 Variation of average crystallite size as a function of different
excitation plasma excitation powers
100
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Figure: 4.25 Variation in residual stress as function different excitation plasma
excitation powers
100
Figure: 4.26 SEM micrographs of untreated and treated specimens at different
plasma excitation plasma excitation powers (a) untreated, (b) 100
watt, (c) 150 watt (d) 200 watt
101
Figure: 4.27 AFM micrographs of deposited specimens at different nitrogen
argon ratios (a) untreated, (b) 100 watt, (c) 150 watt (d) 200 watt
103
Figure: 4.28 Line analysis of deposited specimens at different plasma excitation
plasma excitation powers. (a) untreated (b) 100 watt, (c) 150 watt,
(d) 200 watt
104
Figure: 4.29 Comparison of stress-strain curve of untreated and treated
specimens at different plasma excitation plasma excitation powers,
106
(1) untreated, (b) 100 watt, (c) 150 watt, (d) 200 watt
Figure: 4.30 Variation in the YS and UTS at different plasma excitation plasma
excitation power
107
Figure: 4.31 Variation in percentage elongation at different plasma excitation
plasma excitation power
107
Figure: 4.32 SEM Fractrographs of specimens (a) untreated (b) 100 watt, (c)
150 watt, (d) 200 watt
108
Figure: 4.34 Variation of surface microhardness as a function of different
plasma excitation plasma excitation powers (100, 150, 200 watt)
109
Figure: 4.35 Fluorescence microscopy images after culturing for 3 days: (a) 100
watt, (b) 150 watt and (c) 200 watt
111
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List of Tables Table 1.1: Summary of Anticorrosion coating. 3
Table 1.2: Some mechanical properties of implanted devices. 5
Table 3.1: Ion concentrations in human blood plasma and SBF solution. 56
Table 4.1: Roughness values of untreated and deposited films. 70
Table 4.2: Elemental concentrations of the films. 71
Table.4.3: Roughness values of deposited film by AFM analysis. 82
Table.4.4: Min and Max of coefficient of friction. 84
Table.4.5: Roughness values of untreated and deposited films. 104
18
LIST OF SYMBOLS AND ABBREATIONS
Al Aluminium
AFM Atomic Force Microscopy
Ar Argon
CII Conventional ion Implantation
C Carbon
CSM Continuous Stiffness Measurement
CASP Centre for Advance Studies in Physics
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Cr2O3 Cromium oxide
DMEM Dulbecco’s Modified Eagle Medium
DLC Diamond-like carbon
DC Direct Current
DNA Deoxyribonucleic acid
DAPI 4’,6’-diaidino-2-phenylindole
EDS Energy-dispersive x-ray
e- Electron
EIS Electrochemical impedance spectroscopy
FBS Fetal bovine serum
H Hydrogen
HA Hydroxyapatite
HCIP Hydroxylion
HCl Hydrochloric acid
H2O water
Mg Magnesium
Mg(OH)2 Magnesium hydroxide
MgCl2 Magnesium cholide
MgO Magnesium Oxide
MgCo3 Magnesium carbonates
MAO Micro arc oxidation
N Nitrogen
Na2SO4 Sodium Sulfate
O Oxygen
OH Hydroxylion
PEO Plasma electrolytic anodization
PIII Plasma immersion ion implantation
PIII & D Plasma immersion ion implantation and deposition
PD Plasma deposition
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PVD Physical Vapour Deposition
PBS Phosphate-buffered saline
Rq Root mean square roughness
Rz Ten point mean height
Ra Average roughness
RF Radio frequency
SEM Scanning electron microsopy
SBF Simulated body fluids
Si Silicon
TiO2 Titanium oxide
Ti Titanium
UHV Ultra-high vacuum
UTM Universal testing machine
UTS Ultimate tensile strength
YS Yeild Strength
XRD X-ray diffraction
XPS X-ray photoelectron spectroscopy
Zr Zirconium
LIST OF PUBLICATIONS INCLUDING IN THIS THESIS
1- Sehrish Saleem, R. Ahmad, Uzma Ikhlaq, R. Ayub, Jin Wei-Hong, Xu Rui-Zhen, Li Peng-Hui,
Khizra Abbas, Paul K. Chu,
21
“Effects of N2/O2 flow rate on the surface properties and biocompatibility of nano-structured
TiOxNy thin films prepared by high vacuum magnetron sputtering”, Chin. Phys. B 24 (2015)
075202.
2- Sehrish Saleem, R. Ahmad, R. Ayub, Uzma Ikhlaq, Weihong Jin, Paul K. Chu
“Investigation of nano-structured Zirconium oxide film on Ti6Al4V substrate to improve
tribological properties prepared by PIII-D”, Applied Surface Science 394 (2017) 586-597.
3- Sehrish Saleem, Riaz Ahmad, R. Ayub, Uzma Ikhlaq, Paul. K. Chu, “Study of surface roughness,
mechanical properties and biocompatibility of
Titanium/Alumina composite film”.
OTHER PUBLICATIONS
1- Uzma Ikhlaq, Akira Hirose, Riaz Ahmad, Amir Ikhlaq, Sehrish Saleem, Ramaswami
Sammynaiken, Chunzi Zhang, Jason Malley, The role of Ar flow rates on synthesis of
nanostructured zirconium nitride layer growth using plasma enhanced hot filament nitriding
(PEHFN) technique,Eur. Phys. J. Appl. Phys. (2015) 69: 10801
2- U. Ikhlaq, R. Ahmad, M. Shafiq, S. Saleem, M. S. Shah, T. Hussain, I. A. Khan,K. Abbas, and
M. S. Abbas, Nitriding molybdenum: Effects of duration and fill gas pressure when using 100-Hz
pulse DC discharge technique, Chin. Phys. B Vol. 23, No. 10 (2014) 105203
3- Umm-i-KALSOOM, R. AHMAD, Nisar ALI, I. A. KHAN, Sehrish SALEEM, Uzma IKHLAQ,
Nasarullah KHAN, Effect of Power and Nitrogen Content on the Deposition of CrN Films by
Using Pulsed DC Magnetron Sputtering Plasma, Plasma Science and Technology, Vol.15, No.7,
Jul. 2013
4- K. Abbas, R. Ahmad, I.A Khan, U. Ikhlaq, S. Saleem, R. Ahson, Role of Argon plasma on the
Structural and Morphological Properties of Silicon Nitride Films by Pulsed DC Glow Discharge,
2013 International Conference on Aerospace Science & Engineering (ICASE) 978-1-4799-0993-
3/13$31.00 © 2013 IEEE
5- U. Ikhlaq, R. Ahmad, S. Saleem, M.S. Shah, Umm-i-Kalsoom, N. Khan, and N. Khalid, Argon
gas concentration effects on nanostructured molybdenum nitride layer growth using 100 Hz
pulsed dc glow discharge, Eur. Phys. J. Appl. Phys. (2012) 59: 20801
22
6- M. S. Shah, U. Ikhlaq AND S. Saleem, Nitriding of titanium using capacitively coupled AC
plasma, PK ISSN 0022- 2941; CODEN JNSMAC, Vol. 52, (2012) PP 31-46
7- M. S. Shaw, R. Ahmad, U.Ikhlaq AND S. Saleem, Characterization of pulsed DC nitrogen plasma
using optical emission spectroscopy and Langmuir probe, PK ISSN 0022- 2941; CODEN
JNSMAC, Vol. 53, (2013) PP 01-12 .
Abstract The present research is motivated to make an ideal biomaterial which possessed high strength,
elastic modulus comparable to bone; good wear resistance and excellent biocompatibility. Since Ti and
its alloys are light weight, possessed excellent mechanical strength, high corrosion resistance and good
biocompatibility therefore they are frequently used in joint replacement, bone plates and screw, dental
root implant, vascular stents and spinal fixation devices. Commercial pure (CP) titanium is a good
candidate for biomedical application. However its low mechanical strength and surface hardness limits
its use in load bearing applications. On other hand Ti6Al4V and NiTi alloys are two most useable Ti
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alloys, but vanadium (V) and nickel (Ni) hampered their properties, as both are toxic and produce the
allergic reaction. Therefore it is imperative to improve their surface properties to make them an ideal
biomaterial. A plasma surface modification is an attractive method to improve the surface properties of
biomaterials; because it is not only economic and efficient, but we can also tailor only surface properties
without any change in bulk. In present research work we performed three experiments to improve the
surface properties of Ti and its alloy using the glow discharge.
In first experiment, titanium oxynitride films were deposited on NiTi samples by high vacuum
magnetron sputtering for various nitrogen and oxygen gas flow rates. The composition of deposited film
was characterized using X-ray diffraction (XRD) and x-ray photoelectron spectroscopy (XPS). The results
reveal the presence of TiN, and rutile and anatase phases of TiO2 in the titanium oxynitride thin films.
Energy dispersive spectroscopy (EDS) elemental mapping of samples after immersion in simulated body
fluids (SBF) shows that Ni is depleted from the surface and cell cultures corroborate the enhanced
biocompatibility in vitro.
In second experiment, zirconium oxide nanostructure thin film has been deposited on the surface
of Ti6Al4V alloy via plasma immersion ion implantation and deposition (PIII&D) technique at the
various voltages 15, 20 and 25 KV. The chemical composition and surface morphology of deposited film
is characterized by the X-ray photoelectron spectroscopy (XPS) and Atomic force microscope (AFM)
respectively. The XPS results confirm the formation of ZrO2 film. AFM results show the formation of
smooth film was formed with maximum roughness of 8.4 nm. The effects of the implantation voltages on
the wear characteristics are also investigated by pin-on-disk test. It is observed that wear resistance
improves with an increase in the applied voltage and is found to be maximum at 25 KV. Moreover the
nanohardness is improved in treated specimens and is almost doubled as compared to untreated specimen
at the maximum voltage. The variation in wear resistance and nanohardness is attributed to the formation
of hard nanostructure ZrO2 film on substrate surface.
In third experiment, Ti-Al-O composite film has been formed by using pulsed DC magnetron
sputtering system at various powers (100, 150, 200 watt). The effect of deposited film on mechanical
properties and biocompatibility of CP Ti has been studied. The composition of film has been examined
through X-ray diffractometer (XRD). Surface morphology of deposited film was observed using atomic
force microscopy (AFM) and scanning electron microscopy (SEM) techniques. It was found that surface
roughness of film increase with increasing plasma excitation power. To determine the strength of film,
tensile test was carried out using Universal testing machine. The hardness was also measured by Vickers
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microhardness tester. The results show that composite film improved the mechanical properties such as
YS, UTS and hardness of CP Ti without any reduction in percentage elongation. Moreover, the
biocompatibility of deposited film also performs by culturing the MC3T3-E1 cell for three days. Results
exhibit that composite film significantly improves the biocompatibility of titanium. Micrographs of cell
culture indicate that better cells growth/proliferation (elongated morphology) is observed on film prepared
at 150 watt.
Chapter 1 Introduction
1.1 Biomaterials
Materials which are used to support or improve the function of living parts of body
are known as biomaterials [1]. These materials are artificially implanted to replace as the
whole or to help the damaged part of body without any harmful effects to the tissues, organ
and other parts of body. These materials are extensively used in dental implant, artificial
knee, total hip joint replacement, surgeries, heart valves, stents, drug delivery devices,
contact lens etc [2-6]. In present time the number of diseases has been increased with an
increase in world population. It is estimated that 90% of world population suffer from joint
problems [7]. Also the number of orthopedic implant is also increasing day by day due to
huge traffic accidents. With an increase in the world population it is observed that in 2004,
the stent implanted in human body was more than two millions [8]. In whole world, millions
of people use the contact lens for vision purpose. To improve the quality of life the use of
biomaterials to replace the failed tissue or parts are increasing day by day. Therefore
numerous works has been carried out in medical field to enhance properties of existing
biomaterials and to develop the new class of biomaterials.
25
Although biomaterials successfully performing their functions in human body and
helping the human beings to recover their life, yet it is required to improve their
performance. Following are some important requirements for biomaterials
1) Biocompatibility
2) Corrosion
3) Mechanical Properties
Biocompatibility is the response of implanted material with the other tissue and organs
of the body [9]. It is required that an implanted material should be nontoxic. So that it does
not destroy or produce harmful effects on other parts of body. If implant material contains
some toxic elements which may be reactive, then the surface modification has been carried
out to avoid the release of these toxic elements from the surface of material to be used for
implant. An easy and effective way is to select such materials which have ability to form
the oxide, nitride or carbide layers [10]. It is observed that these layers are very effective
to enhance the biocompatibility of many implanted materials [10].
26
igure 1.1: Factors and their effect on biocompatibility
Corrosion is a process in which material is destroyed gradually by the chemical attack of
existing environment. Since it is desired that artificial implant has no effect on other parts
of human body, however, it is also vital that the body environment does not produce any
influence on the implant material. Since the environment of body is hot and salty, there
might be the chance that implanted materials degrade. Therefore it is required to prevent
the biomaterials from the corrosion, because it reduces the mechanical properties or
durability of materials. To avoid the corrosion effect the properties of materials are changed
either alloying other elements or changing the surface properties by coating of such
materials which are resistive to corrosion [11-13]. Some techniques which are used for
anticorrosion coating on biomaterials are given in the table 1.1.
Table 1.1: Summary of Anticorrosion Coating
Ref
Substrate
Method
Electrolyte
[11]
316-L stainless steel
Reactive magnetron
sputtering
pH 5.6 acetic acid and
sodium acetate buffer
solution
[12]
Plain carbon stee (Ck35)
magnetron
sputtering technique
1N sulfuric acid
solution
[13]
NiTi alloy
Plasma Immersion
Implantation
Deposition (PIII&D)
Ion
and
Hank’s solution
[14]
1Cr11Ni2W2MoV
Martensitic stainless
steel
(HCIP )
0.5 mol/L NaCl and
1mol/L H2SO4 diluted
aqueous
solution
[15]
Ti-6Al-4V
Plasma assisted
electron beam
technique
PVD
0.5N NaCl solution
[16]
NiTi coated Si
Dc magnetron sputtering
1 mol/L NaCl solution
27
[17]
316L stainless steel
Arc ion plating
Neutral Troyde’s
(SBF)
In many applications it is most important that biomaterials should possess good
mechanical properties. Some significant mechanical properties of biomaterials are yield
stress, ultimate tensile stress, ductility, hardness, durability, excellent wears resistance,
elastic modulus and toughness [6]. The load bearing ability of implant material can be
determined from its yield stress. The yield stress is found from the stress strain curve of
material, where a constant load is applied on material and its deformation is determined.
Since the bone is hard and brittle, so it is required that the yield stress of material should
be high. A material having low yield stress does not tolerate load resulting in the premature
failure [11]. Besides the yield stress, it is also necessary that the elastic modulus of
implanted material and bone should be comparable. An effect known as stress shielding is
occurred due to the mismatch of elastic modulus of implanted material and natural bone [1,
7]. Due to this effect, the mass and thickness of bone gradually reduced. So it is essential
to make the elastic modulus of implanted material comparable with that of bone. The elastic
modulus of bone ranges from 10-40 GPa. The most world widely used metallic materials as
implant are stainless steel, Ti and its alloy and Co base alloy. However the value of their
elastic modulus is much higher than that of bone. Therefore it is necessary to reduce elastic
modulus before using these materials as implant. This is done by either alloying such
elements having low value of elastic modulus or by surface modification of these materials
using different techniques.
28
Figure 1.2: Value of Elastic Modulus of Metallic Materials [18]
Another important factor of implant materials is its durability and fatigue resistance.
The implant material should be stable and have long life. For example a hip joint must
remain stable under load more than 10 years. Some mechanical properties of implant
devices are given below:
Table 1.2: Some Mechanical Properties of Implanted Device
Device
Properties
Heart valve
Flexible and tough
Hip prosthesis
Strong and rigid
Dialysis membrane
strong and flexible
29
Tendon material
strong and flexible
1.2 Types of Biomaterials
There exist various materials which can be used as biomaterials. However according to
the biomedical applications we can divide the biomaterials into three major categories
which are
1) Ceramics
2) Polymers
3) Metallic Materials
Ceramics are extensively used in different medical applications like artificial knee, bone
fixation and dental coating [1]. The unique properties of ceramic which make them
prominent to use as biomaterials are
1) Their good biocompatibility
2) Excellent compressive strength
3) High Toughness
4) Inert in Body
Ceramics are used direct as an implant materials or some time they coated on the surface
of other materials. The most widely used ceramics in medical applications are
1) Inert Bioceramic
2) Resorable Bioceramic
3) Bioactive ceramic
4) Porous Bioceramic
In most of medical applications polymers also play very vital role. Some important
properties of polymers which make them prominent from the other biomaterials are
1) Light weight
2) Flexibility
3) Good biocompatibility
30
4) Corrosion Resistance
Excellent mechanical properties and good electrical conductivity of metallic material make
them the most favorable for biomaterials, particularly for the orthopedic implants, where
the good mechanical strength is necessary [7]. Some important applications of metallic
materials are bone plate, Knee joints, total hip, vascular stents, dental implants [4, 7]. The
metallic biomaterials consist of metals and alloys. As metals have free electrons in valence
shell. It is observed that most of metals lose their electrons particularly in the solution.
Therefore metals have more probability to corrode rapidly as compared to other
biomaterials. Moreover their poor wear resistance limiting their use in medical field. To
avoid the corrosion and wear, the surface properties of materials are improved by coating
the different layers.
Metallic materials are generally divided into three main categories.
1) Stainless Steel
a) 302 stainless steel
b) 316 stainless steel
c) 316L stainless steel
2) Cobalt base alloys
a) Co-Cr-Mo
b) Co-Ni-Cr-Mo
c) Co-Cr-W-Ni alloy
3) Ti and its alloy
a) Pure Titanium ( Grade 1-4)
b) Ti-6Al-4V (6%Al, 4%V)
c) Ti-6Al-7Nb (6% Al, 7% Nb)
d) Ti-Ni alloy
In above given metallic materials Ti and its alloy are most frequently used metallic
materials for biomedical application due to their high biocompatibility, light weight,
excellent mechanical strength, and low elastic modulus value as compared to other metals
and alloys and good corrosion resistance [1]. Commercially pure Ti exists in four grades
31
and has hexagonal crystal structure. When Ti is heated above a certain temperature a phase
transformation from hexagonal to body center takes place. Ti6Al4V is the most suitable
alloy used for many implant devices due to its high mechanical strength and good
biocompatibility as compared to the pure titanium. However said alloy exhibits the poor
tribological properties [1]. Moreover vanadium is a toxic element and produces the allergic
reaction when used for long term implant. Another very useful Ti alloy in medical field is
NiTi alloy also known as shape memory alloy. The properties which make NiTi alloy
superior as compared to other Ti alloys are
1) Shape memory effect
2) Superelasticity
Shape memory effect refers to regain the original shape upon heating even after
deformation.
The property of material to return its parent shape upon the removal of stress is called
superelasticity. Some different applications of metallic materials as implant devices are
shown in the figure below.
32
Figure 1.3: Applications of Biomaterials in different parts of body [19]
33
Figure 1.4: Photographs of cardiovascular components such as (a) rotary heart
pumps, (b) artificial heart valves and (c) stents [20]
1.3 Plasma-based technology
Plasma in gaseous form comprises of charged particles and neutrals, often known
as fourth state of matter. When plasma is formed in reactive chemical environment many
unusual plasma surface interactions occur. Hence, to change the chemical, structural and
morphological properties of surface, plasma is most economical surface treatment
technique and attracts growing interests in materials engineering. Plasma based methods
like sputtering, cleaning, etching, ion implantation and deposition [21, 22] are broadly used
to alter surface of polymers, ceramics, metal, alloys and industrial components including
biomaterials [23-33].
34
1.4 Application of plasma based technology to biomaterials
Due to long term procedures to fulfil the requirements of biocompatibility it is
extremely difficult to develop new biomaterials. So as to the meet increasing needs of
biomaterials in medical field Plasma surface modification technique helps to modify the
surface biological and mechanical properties of conventional biomaterials. Hence,
materials that have promising bulk properties can have their surfaces redesigned to satisfy
to biomedical applications. Thin films formation on a substrate can shield the bulk material
by providing a barrier against physical and chemical environment that would corrode the
material. Hence, there are so many deposition techniques available for surface modification
of materials involving physical and chemical phenomenon. Surface modification through
plasma is a widespread technique to improve corrosion resistance, mechanical and
biocompatibility of biomaterials and medical devices [34]. Thin film coating depends on
many factors like process parameters and which deposition process is used. Many
researchers have presented a detailed review of thin-film deposition processes [35-38].
Different methods are used for depositing a thin film. In most of them vapor of some kind
are used for deposition.
In PVD techniques deposition of thin film is carried out through formation of vapors
by means of energetic bombardment of a source. The phenomenon involve in this process
are sputtering, evaporation ion plating etc.
1.5 High Vacuum Magnetron Sputtering System
In this research work Plasma Magnetron sputtering and PIII&D techniques are used
to improve the biocompatibility of Ti and its alloys. Sputtering depends on the ion and
target atoms, as well as the ion energy. Impinging ions will also help to emit the secondary
electrons which are accelerated away from the target by the applied negative potential.
These ionize more atoms of the sputtering gas, thereby sustaining the discharge. Usually a
negative potential is applied to the substrate. Therefore, the coating surface is also ion
bombarded during growth. This ion bombardment is applied in order to densify the growing
film by enhancing the surface mobility.
35
Permanent magnets below the target are used for applying magnetic field to increase the
ionization rate by emitted secondary electrons even further. The PVD system is shown in
Fig 1.5
Figure 1.5: Working diagram of ultra high vacuum PVD system
1.6 Plasma Immersion Ion Implantation and Deposition
PIII&D is an exclusive surface modification technique for different type of
materials including polymers, metals and ceramics. PIII&D system is comprised of plasma
source, vacuum chamber, power supply and target [39-41]. This device is easy to use and
comparatively low in price as a conventional beam line ion implanter because it does not
require additional focusing optics, masking, beam extraction optics and target manipulation
as in PIII. Also PIII&D system reduces the degradation rate without changes the bulk
properties and many surgical implant and tools with a complex design and shapes are
treated by this. The substrate material is directly placed in plasma and the process can easily
be controlled by adjusting the applied voltage, pulse width, plasma density and
configuration [23]. When negative voltage is supplied to target material in early stage then
36
the electrons are repelled on the time scale of the inverse electron plasma frequency to
create a positive space charge region. Ion matrix sheath can be described by the potential
profile. A quasi-static Child Law sheath develops followed by ion collection/extraction and
plasma sheath expansion. The plasma sheath evolution plays a vital role in PIII, since it
provides information about the implantation process. Process parameters can be predicted
by using it and results in the implantation profile, implantation dose and implantation
current. Schematic diagram of the PIII&D process installed at City University Hong Kong
is shown in Fig. 1.6.
Figure 1.6: Working diagram of PIII-D system
37
1.7 Layout of the thesis
The thesis is divided into four chapters. In first chapter, introduction of the
biomaterials and its long term limitations are discussed along with the basic materials
processing techniques and the use of the different plasma techniques as a promising source
for surface modifications of biomaterials. The second chapter contains review of literature
about biomaterials, thin films and plasma techniques related to the work present in this
thesis. In third chapter experimental setup and different surface diagnostic techniques are
described in detail.
Fourth chapter deals with the objectives of this research and includes:
The poor wear and corrosion resistance in some aggressive environment and release of
some toxic elements involve in titanium alloys like Nickel and vanadium in NiTi and
Ti6Al4V alloys respectively limits their use in medical applications. Following results are
discussed in this chapter:
(i) Formation of titanium oxynitride films as a barrier against nickel release in NiTi
shape memory alloy.
(ii) To enhance the wear and wettability of Ti6Al4V by the deposition of zirconium
oxide by using PIII&D technique.
(iii) Study the mechanical properties of Titanium by Deposition of Alumina using
Pulsed DC glow discharge.
PIII&D is an effective surface modification method to enhance the corrosion
resistance of titanium. Characterization techniques such as X-ray photoelectron
spectroscopy (XPS), Scanning electron microscope (SEM), X-ray Diffraction (XRD),
Atomic force microscope (AFM) and fluorescence microscope, the degradation behavior
of TiON films in SBF solution as well as in cell culture medium is assessed. The influence
of different oxygen nitrogen concentrations on formation of titanium oxynitride films and
deposition of zirconium oxide film at different voltages is also discussed. In addition,
38
structural, mechanical and biocompatibility of Ti-Al-O composite film coated on titanium
substrate has been investigated.
1.8 References
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Business Media: New York, NY, USA (2007).
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[2] H. S. Tran, M. M. Puc, C. W. Hewitt, D. B. Soll, S. W. Marra, V. A. Simonetti, J.
H. Cilley, A. J. DelRossi, J. Investigative Surgery 12 (1999) 133-140.
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[10] S. Izman, M. R. Abdul-Kadir, M. Anwar, E.M. Nazim, R. Rosliza, A. Shah and
M.A. Hassan, (2012), Surface modification techniques for biomedical grade of
titanium alloys:Oxidation, Carburization and Ion Implantation Processes. In A.K.
M, Nurul Amin (Ed.), Titanium Alloys-Towards achieving enhanced properties for
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142-144 (2001) 1078-1083.
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(1996) 207-280.
[25] X. Lu, S. S. K. Iyer, C. M. Hu, N. W. Cheung, J. Min, Z. Fan, P. K. Chu, Appl.
Phys. Lett 71 (1997) 2767-2769
[26] P. K. Chu, S. B. Felch, P. Kellerman, F. Sinclair, L. A. Larson, B. Mizuno, Solid
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Coat. Tech. 186 (2004) 239-244.
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Tech. 186 (2004) 295-298.
[29] P. Chen, S. S. Lau, P. K. Chu, K. Henttinen, T. Suni, I. Suni, N. D. Theodore, T.
L. Alford, J. W. Mayer, L. Shao, M. Nastasi, Appl. Phys. Lett. 87 (2005) 1119101–
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[31] Z. F. Di, P. K. Chu, M. Zhu, R. K. Y. Fu, S. H. Luo, L. Shao, M. Nastasi, P. Chen,
T. L. Alford, J.W. Mayer, M. Zhang, W. L. Liu, Z. T. Song, C. L. Lin, Appl. Phys.
Lett. 88 (2006) 142108-1 – 142108-3.
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[35] D. Brandon, W. D. Kaplan, Microstructural Characterization of Materials, John
Wiley & Sons (1999).
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Chapter 2
42
Literature Survey
The literature reviewed in this chapter contains information about modification of
biomaterial surfaces without altering bulk properties. Content also includes different
methods used for depositing thin films for surface modification of biomaterials.
2.1 TiON layer on NiTi alloy
Ti and its alloy are most attractive materials for medical applications due to their
high mechanical strength, durability, good biocompatibility and excellent corrosion
resistance. In addition their modulus of elasticity much lower as compared to the other
metallic material. Among Ti alloys NiTi and Ti6Al4V alloys are most frequently used in
different medical applications. Two unique characteristic which increase the demand of
NiTi alloy in implanted devices is shape memory effects (SME) and superelasticity.
However, Ni is toxic and release of Ni-ions from NiTi alloy produce the allergic reaction
in body. Hence it is desired to stop the release of Ni from the surface of the NiTi alloy. This
can be achieved by the surface modification of NiTi alloy using the various techniques.
Some methods are described here.
The biocompatibility of NiTi alloy after the deposition of TiO2 film was
investigated by Liu et al. [1]. XRD, AFM and corrosion test was conducted to determine
the effects of TiO2 film on biocompatibility of NiTi alloy. The thickness of film was about
205 nm calculated using the ellipsometry. The AFM results indicate that a uniform
nanostructure film was formed on the substrate surface. The roughness of film was about
3.9nm. The XRD results confirmed the formation of anatase and rutile phase of TiO2. In
addition TiO2 exist mainly in anatase phase in temperature range 400-500 oC, whereas with
an increase in the temperature (above 800 oC) anatase was converted into the rutile phase.
They observed that corrosion resistance of TiO2 deposited specimens was greater as
compared to the untreated NiTi alloy. The improvement of corrosion resistance was
attributed to the formation of TiO2 layer on the surface of NiTi alloy. The biocompatibility
of deposited film was characterized using the in vitro blood platelet adhesion test. Their
43
results showed that treated specimens have better biocompatibility as compared to the
virgin NiTi alloy.
Barocs et al. [2] studied the nitrogen ions effect on the corrosion behavior of NiTi
alloy to improve the biocompatibility of NiTi alloy. Nitrogen ions were implanted on
surface of NiTi alloy at the dose varying from 0.5x1017 to 8x1017 ion cm–2 at 150 keV. To
investigate the N-ions effects on NiTi substrate XPS, nanohardness and EIS test were
carried out. The XPS results confirmed the formation of titanium oxide, nitride and
oxynitride bond. The nanohardness value in N-ions implanted specimens was found to be
increased as compared to the virgin specimen. The corrosion resistance measured by EIS
test was improved in the specimen implanted at dose 2x1017 ion cm–2 as compared to the
other specimens. They observed that improvement in nanohardness and corrosion
resistance in implanted specimens was due to formation of titanium oxide and nitride layer
on substrate surface.
Poon et al. [3] examined the carbon ions effect on the corrosion and
biocompatibility of NiTi alloy. They coated that the carbon ions form a carbide thin layer
on the surface of substrate, which act as barrier and hindered the release of Ni-ions from
the NiTi surface. Moreover nanohardness of all implanted specimens is much higher as
compared to the un-implanted specimen. However with decrease in the depth the value of
nanohardness decreases. On the other hand the young modulus in carbon ions coated
specimens were found to be 3.5% less as compared to the un-coated specimen. It was
observed that carbon ions enhanced the corrosion resistance of NiTi alloy. The
improvement of the corrosion resistance in C-ions coated specimens may be due to the
decrease in the electrical conductivity. Cell culture results also showed that C-ions increase
the biocompatibility of NiTi alloy.
Wu et al. [4] studied surface, mechanical and biocompatibility of porous NiTi alloy
after oxygen implantation. To determine the Ni-release the specimens were placed in SBF
solution for 70 days at room temperature. They mentioned that the oxygen ion implantation
significantly minimize the release of Ni-ions from the NiTi surface.
44
Moreover, it was observed that more cells attached to the implanted specimens, which mean
that oxygen implantation reduces the cytotoxic effects of Ni. They mentioned due to
oxygen implantation a layer of stable TiO and NiO were formed on the surface of NiTi
alloy, which enhance the biocompatibility. The mechanical properties were examined by
performing the compression test. It was revealed that both the un-implanted and oxygen
implanted specimens had good superelasticity.
Habijan et al. [5] identified that magnetron sputtering method was a novel technique
to enhance the biocompatibility. They deposited the NiTi film on the glass substrate. The
mechanical properties and biocompatibility of film was calculated. The thickness of NiTi
film was about 6.5 µm. They reported that NiTi film formed by magnetron sputtering
technique not only enhanced the biocompatibility but also reduced the Ni release. Therefore
magnetron sputtering was a convenient technique for the production of NiTi stent.
The surface modification of NiTi alloy was conducted by the Liu at al. [6] using the
plasma ion implantation technique. Nitrogen ions were implanted on the surface of NiTi
alloy at the different voltages 20 kV, 30 kV and 40 kV. The surface topography of nitrogen
ion implanted and untreated specimens were characterized using the atomic force
microscope (AFM), whereas the biocompatibility was determined by cell culture test. The
amount of Ni-release after the implantation has also been calculated. The release of Ni-ions
minimize in the specimens implanted at the 30 kV and 40 kV. They observed that the
nitrogen ions implantation improved the biocompatibility and mechanical properties of
NiTi alloy. They concluded due to nitrogen ions implantation TiN film was formed on the
surface of NiTi alloy, which act as barrier and reduced Ni release and also enhanced the
cell proliferation.
In another study Yeung et al. [7] examined the nitrogen ions effect on the NiTi
alloy. The surface of N-implanted specimens was studied using the different techniques
and results compared with those of untreated, stainless steel and Ti6Al4V alloy. They found
7.7 GPa hardness of N-implanted NiTi specimens at the surface which was much higher
(7.7 GPa) as compared to the untreated specimen (5.2 GPa) and SS (6.7 GPa).
45
However hardness of virgin Ti alloy (9.2 GPa) was higher as compared to the Nimplanted
specimen. The surface topography results taken by AFM indicated that the roughness value
in N-implanted specimens was greater as compared to the other all specimen. They
observed that N-implantation favored the formation of nitride layer with decreased of Ni
release. Moreover corrosion resistance of N-implanted specimen was higher as compared
to the others.
Sun et al. [8] reported that titanium oxynitride coating improved the biomedical
application of NiTi alloy. The coating was performed using the PIII&D. The coated NiTi
specimens were characterized using the AFM, SEM, XRD and XPS. Cell culture, Wear
resistance and nanohardness of oxynitride coated film were also calculated and results
compared with uncoated NiTi specimen. They found nanosize particles generated in
PIII&D coated specimens. Moreover the oxynitride film hindered the release of Ni-ions
from substrate surface and had better biocompatibility as compared to untreated NiTi alloy.
The nanohardness and wear resistance NiTi alloy were found to be improved in all coated
specimens. They concluded that coating of oxynitride film using the PIII&D enhanced the
mechanical properties, wear resistance and biocompatibility of NiTi alloy.
The biocompatibility of NiTi alloy can also be enhanced by reducing the Ni-ions
release from the surface of NiTi alloy, because Ni was toxic and produce the allergic
reaction in the body. The release of Ni-ions from the surface of NiTi alloy after nitrogen
ion implantation was examined by Camargo et al. [9]. Nitrogen ions were implanted at
three different temperatures 250 oC, 290 oC and 560 oC. The release of Ni-ions in treated
and untreated NiTi alloy were determined after immersion the specimens in SBF solution
for 7 days. They found that nickel ion release was significantly decreased in the specimens
implanted at the highest temperature. They concluded at high temperature a dense layer of
nitrogen was formed on the surface of NiTi alloy, which acted as barrier and minimized the
Ni release.
The effect of TiN coating on the surface properties and biocompatibility of NiTi
alloy was discussed by Jin et al. [10]. The film was coated on NiTi substrate using the
filtered arching ion plate method. The surface morphology and composition of film were
46
determined by the SEM and XRD. The results confirmed the formation of TiN film on
substrate. The thickness of film measureed by the SEM was about 2.86 µm. The roughness
of film was found to be increased in all coated specimens as compared to the untreated
specimen. The biocompatibility of film was explored using the cell culture and cytotoxicity
test. The results indicated that TiN coated specimens displayed better biocompatibility as
compared to the uncoated specimen. The improvement in biocompatibility of all coated
specimens was attributed to the increase in surface roughness, because rough surface
provide more surface contents for cell attachment.
2.2 Nano-structure ZrO2 film on Ti6Al4V.
The effect of ZrO2 film on the bioactivity and cytocompatibility of silicon substrate
was investigated by Liu et al. [11] using the cathodic arc deposition. The surface
morphology of deposited film was characterized by Scanning Electron Microscope (SEM)
and Atomic Force Microscope (AFM), whereas the phase composition of film was
determined using the X-ray Diffraction (XRD) and Rutherford backscattering spectroscopy
(RBS). The XRD and RBS results confirm the formation of zirconium oxide film.
Moreover film was crystalline. The surface morphology of treated specimens taken by the
AFM showed film contains the nano-size particles and the roughness of film was less than
1 nm. The bioactivity of film was determined after dipping in the SBF solution. They
concluded that the deposition of zirconium oxide film on the surface of silicon using the
cathodic arc deposition significantly enhance the biocompatibility and cytocompatibility.
Dong et al. [12] investigated the wear behavior of untreated and treated Ti6Al4V
alloy using tribometer. They found that the wear behavior of untreated and thermally
treated oxide specimen were totally different. The wear resistance in all treated specimens
was found to be increased as compared to the untreated specimen. They concluded that the
increase in the wear resistance is due to the formation of rutile phase of TiO2 in treated
specimens.
Zirconia and alumina composite was prepared by Liang at al. [13] using the plasma
spray technique. Two samples with different ratio of Al2O3 were prepared and given the
47
name ZA15 and ZA30 respectively. The friction and wear resistance was evaluated with
the help of ball on disc technique. They found that the friction coefficient and wear rate of
specimen ZA30 was much lower as compared to the ZA15.
In another study Masmoudi et al. [14] studied the wear mechanism of Ti6Al4V
alloy. Pin on Disc technique was used to investigate the wear behavior in the three different
environments air, ringer solution and NaCl. All the specimens were analyzed at load of 1N
with sliding speed of 4mm s-1. They observed that coefficient of friction and wear rate of
air and non- passivated specimen was much higher as compared to the passivated specimen.
Mello et al. [15] examined the plasma immersion ion implantation (PIII) effect on
the tribological properties of Ti6Al4V alloy. The surface of Ti6Al4V alloy was modified
using the nitrogen ions. The composition of implanted specimens was characterized using
the X-Ray Diffraction and Auger Electron Spectroscopy. Nano indentation test was carried
out to determine the surface hardness. The friction coefficient and wear mechanism was
calculated using the pin on disc technique at the load value of 1 N and sliding speed 5 cm
s-1. They found that PIII significantly enhanced the surface layer from 100 nm to 1500 nm.
The surface hardness of nitrogen implanted specimen is remarkably increased. Moreover
the coefficient of friction in the plasma implanted specimens was found to be decrease. The
variation in the implanted specimen may be due to the formation of Ti2N and TiO2 layer
on the surface of substrate.
The tribological properties of NiTi alloy were examined by Wu et al. [16]. The
surface morphology of NiTi specimens were studied using the scanning electron
microscope (SEM), whereas chemical composition of surface was investigated by X-ray
diffraction (XRD) and X-ray photoelectron spectroscopy (XPS). The composition and
morphology of wear debris were explored using the SEM and EDS analysis. The phase
transformation temperature of NiTi alloy was also calculated. The wear behavior of porous
NiTi alloy was studied with help of ball on disc tester in air environment. All tests were
performed at room temperature with different load value 1N, 2N and 3N. They observed
that NiTi specimens containing the higher porosity displayed the better tribological
properties as compared to the low porosity specimens.
48
Yan et al. [17] also studied the wear mechanism of martensitic NiTi alloy. The wear
test was conducted using the ball on disc tester under the different load value (50 mN to
500mN) and varying number of cycles (10-1000 cycles). The results of friction coefficient
calculated under different load revealed that the degradation phenomena of NiTi alloy
consisted of three stages. In first stages (zero wear stage) the value of coefficient of friction
is very small. On the other hand value of coefficient of friction was much higher in stage-
2 and stage-3. The structure of wear debris showed that the crown like structure was formed
during the wear track in the stage-1 and stage-2. The variation in the coefficient of friction
in the stage-3 may be due to the formation of wear debris. Moreover the surface roughness
was calculated by atomic force microscope in tapping mode and found to be less than 15
nm.
Liu et al. [19] implanted the nitrogen ions using the plasma immersion ion
implantation technique to modify the tribological properties of NiTi shape memory alloy.
All the specimens were implanted with nitrogen ions at the different voltages 10, 20, 30
and 40 kV. The penetration depth of nitrogen ions at these voltages was also calculated
using the TRIM and found to be 26, 46, 66 and 84 nm respectively. The chemical
composition of N-implanted specimens was determined by the X-ray photoelectron
spectroscopy (XPS). The XPS results confirmed the formation of TiN layer in the
implanted specimens. The thickness of film was depending on the value of applied voltage.
The wear behavior of implanted specimens was determined using the pin on disc test at the
different load 1N, 2N and 4N. They examined that the wear resistance of NiTi alloy was
improved after the N-ion implantation. Particularly specimens implanted at higher voltage
(30 kVand 40 kV) contained the maximum wear resistance as compared to the virgin NiTi
alloy. On the other hand with an increase in the load value the wear resistance was found
to be decreased.
The surface modification of Ti6Al4V using the plasma electrolytic oxidation
(PEO) was reported by Peng et al. [20]. The surface morphology of film was determined
by the scanning electron microscope (SEM), whereas the elemental analysis was carried
out with the help of energy dispersive spectroscopy (EDX). The wear behavior of coated
specimens was investigated using the pin on disc technique in dry condition at the normal
49
load of 2N. They coated that as the result of PEO coating the coefficient of friction of
Ti6Al4V alloy reduced and improved the wear resistance.
More recently Nohava et al. [21] studied the effect of oxide, nitride and oxynitride
coating on the wear mechanism. The coating process was performed using the plasma vapor
deposition (PVD). The wear test was conducted using the pin on disc technique at different
temperatures 24 oC, 600 oC and 800 oC. They mentioned that nitride and oxynitride coating
showed the better wear resistance at room temperature as compared to the other
temperatures.
2.3 Alumina deposition effects on mechanical properties of Ti
Oxygen effects on lattice strain and mechanical properties of commercially pure Ti
and Ti6Al4V was investigated by Oh et al. [22]. In order to observe the oxygen effect on
mechanical properties, the tensile strength, percentage elongation and Vicker hardness was
calculated. It was observed that strength and hardness both increased with an increase in
the oxygen concentration. However, the percentage elongation of Ti significantly
decreases, whereas the percentage elongation of Ti6Al4V gradually decreases with a rise
in oxygen content. Their results showed that mechanical properties of Ti effected greatly
as compared to Ti6Al4V alloy with increase in oxygen contents. The variation in
mechanical properties of Ti and Ti6Al4V explained on the basis of difference in solid
solution hardening of oxygen between Ti and Ti6Al4V alloy. They mentioned that oxygen
contents have greater influence on the c/a ratio of Ti as compared to the Ti6Al4V, which
results in greater effects on strength and hardness of Ti.
Wei et al. [23] studied the effects of interstitial elements such as oxygen, nitrogen
and hydrogen on the mechanical properties of pure titanium. The mechanical properties
such as yield stress, ultimate tensile stress, percentage elongation, reduction in area,
hardness was investigated. They concluded that oxygen and nitrogen had greater effects on
the mechanical properties as compared to the hydrogen.
Yan et al. [25] also studied the oxygen effects on the ductility of Ti and its alloys.
50
Three types of Ti alloys α-Ti, α+β-Ti and β-Ti were fabricated. They observed that the
ductility of α-Ti is significantly effects due to the oxygen content. On the other hand the
ductility of β-Ti no noticeable decrease with an increase in the oxygen contents. The small
decrease in the β-Ti was due to the small suppression of ω-phase with an increase in oxygen
contents.
Xing et al. [26] studied the effect of alumina coating on microstructure and
mechanical properties. The alumina coating was prepared on the stainless steel substrate at
various temperatures by plasma spray method. The microstructure of coated film was
determined using the scanning electron microscope (SEM). They mentioned that
interlamellar interface bonding is increased with rise in the substrate temperature. To
determine the mechanical properties of alumina film they calculated the hardness and
elastic modulus. The results show that with rise in the substrate temperature both hardness
and elastic modulus also increase. The increase in hardness and elastic modulus was
attributed to increase in the interlamellar interface bonding.
Since the strength and elastic modulus are very important properties of materials
used as implant. The strength determined durability and life time of materials used as
implant. On other hand, if elastic modulus of implant does not match with bone, a effect
called stress shield occurred, which result in ultimate an early fracture. In order to obtain
the good combination of mechanical properties along with biocompatibility Vicente et al.
[27] investigated the effects of oxygen contents on structural, mechanical and
biocompatibility of TiZr alloy. Since oxygen is an interstitial element, so it formed the solid
solution, thus improving the mechanical properties of TiZr alloy. They observed that the
oxygen contents did not affects the microstructure of alloy. The microhardness found to be
increased with an increase in oxygen concentration. The oxygen contents also decreased
the elastic modulus as compared to Ti. Moreover cytocompatibility test showed no toxic
effect of material.
Santos et al. [28] determined the titania film effects on the mechanical properties of
Ti. The film was successfully prepared on substrate using the anodic oxidation method.
The mechanical properties of film were explored by studying the elastic modulus,
51
nanohardness and brittleness. The nanohardness results showed that hardness of all treated
specimens was greater than Ti. The elastic modulus in all treated specimens was found to
be small as compared to Ti. They concluded that anodic oxidation formed the hard titania
film, which ultimately reduced the elastic modulus of Ti.
2.4 References
[1] J. X. Liua, D. Z. Yanga, F. Shia, Y. J. Cai, Thin Solid Films 429 (2003) 225–230.
52
[2] R. Barcos, A. Conde, J. J. de Damborenea, J. A. Puertolas REVISTA DE
METALURGIA, 44 (2008) 326-334.
[3] R.W.Y. Poona, K.W.K. Yeungb, X.Y. Liua, P.K. Chua, C.Y. Chunga, W.W. Lub,
K.M.C. Cheungb, D. Chan, Biomaterials 26 (2005) 2265–2272.
[4] S. L. Wu, P. K. Chu, X. M. Liu, C. Y. Chung, J. P. Ho, C. L. Chu, S. C. Tjong , K.
W. Yeung, W. W. Lu, K. M. Cheung, K. D. Luk, J. Biomed Mater Res A. 79(1):
2006 139-46.
[5] T. Habijan, R. L. De Miranda, C. Zamponi, E. Quandt, C. Greulich, T. A.
Schildhauer, M. Koller, Mater. Sci. Eng. C 32 (2012) 2523–2528.
[6] X. M. Liu, S. L. Wu , Y. L. Chan, P. K. Chu , C. Y. Chung , C. L. Chu , K. W.
Yeung, W. W. Lu , K. M. Cheung , K. D. Luk, J. Biomed Mater Res A. 82 (2007)
469-78.
[7] K. W. Yeung, R. W. Poon, P. K. Chu, C. Y. Chung, X. Y. Liu, W. W. Lu, D. Chan,
S. C. Chan, K. D. Luk, K. M. Cheung J. Biomed Mater Res A. 82 (2007) 403-14.
[8] T. Sun, L. P. Wang, M. Wang, H. W. Tong, W. W. Lu, Mater. Sci.Eng. C 32 (2012)
1469–1479.
[9] E. N. D. Camargo, A. O. Lobo, M. M. D. Silva, M. Ueda, E. E. Garcia, L. Pichon,
H. Reuther, J. Otubo, JMEPEG 20 (2011) 798–801.
[10] S. Jin, Y. Zhanga, Q. Wanga, D. Zhanga, S. Zhang, Colloids and Surfaces B:
Biointerfaces 101 (2013) 343– 349.
[11] X. Liua, A. Huanga, C. Dingb, P. K. Chu, Biomaterials 27 (2006) 3904–3911.
[12] H. Dong, T. Bell, Wear 238 (2000) 131–137.
[13] B. Liang, G. Zhang, H. Liao, C. Coddet , C. Ding, Surface & Coatings Technology
203 (2009) 3235–3242.
53
[14] M. Masmoudi, M. Assoul, M. Wery, R. Abdelhedi, F. El Halouani, G. Monteil,
Applied Surface Science 253 (2006) 2237–2243.
[15] C.B. Melloa,b, M. Uedaa, M.M. Silvac, H. Reutherd, L. Pichone, C.M. Lepienski,
Wear 267 (2009) 867–873.
[16] S. Wu, X. Liu, G. Wu, K. W. K. Yeung, D. Zheng, C. Y. Chung, Z. S. Xu, P. K.
Chu, J.Biomed Mater Res Part A 2013 101A:2586–2601.
[17] L. Yan, Yong Liu n, Erjia Liu, Tribology International 66 (2013) 219–224.
[18] P.W. Shum, Y.F. Xu, Z.F. Zhou, W.L. Cheng, K.Y. Li, Wear 274– 275 (2012) 274–
280.
[19] X. Liu, S. Wu , Y.L. Chanb, P. K. Chua, C.Y. Chung, C.L. Chua, K.W.K. Yeung,
W.W. Lub, K.M.C. Cheungb, K.D.K. Luk, Materials Science and Engineering A
444 (2007) 192–197.
[20] B. Y. Peng, X. Nie, Y. Chen, International Journal of Aerospace Engineering
Volume 2014 (2014), Article ID 640364, 10 pages.
[21] J. Nohava, P. Dessarzin , P. Karvankova, M. Morstein, Tribology International
81(2015) 231–239.
[22] J. M. Oh, B.G. Lee, S. W. Cho, S. W. Lee, G. S. Choi, J. W. Lim Met. Mater. Int.17
(2011) 733-736.
[23] Y. Wei, H. K. D. H. Bhadeshia, T. Sourmail J. Mater. Sci. Technol. 21 (2005) 403-
407.
[24] M. Mittala, S. K. Natha, S. Prakasha Journal of Minerals & Materials
Characterization & Engineering, 10 (2011) 1041-1049.
[25] M. Yan, W. Xu, M. S. Dargusch, H. P. Tang , M. Brandt and M. Qian Powder
Metallurgy 57 (2014) 251-257.
[26] Y. Xing, C. Jiang, H. Chen, J. Hao, Acta Mechanica Solida Sinica 24 (2011) 461.
54
[27] F. B. Vicente, D. R. N. Correa, T. A. G. Donato, V. E. A. Chavez, M. A. R. Buzalaf,
C. R. Grandini, Materials 7 (2014) 542.
[28] J. E. Santos, N. K. Kuromoto, G. A. Soares, Materials Chemistry and Physics 102
(2007) 92.
Chapter 3
55
Experimental Setup and Characterization
Techniques
3.1 Introduction
This chapter contain brief introduction related to experimental systems used in this
research for depositing thin films, also characteristics of biomaterials selected for
deposition and as a substrate material. Additionally to study thin film morphology, different
characterization techniques such as XRD, SEM attached with EDS, AFM, XPS,
preparation of SBF solution, cell culture, wear test, wettability, nanoindentation and
universal testing machine are employed.
3.2 High Power Plasma Magnetron Sputtering System
Deposition of titanium oxynitride film was performed in the Ultra high vacuum PVD
system (magnetron system) in the Plasma Laboratory at City University of Hong Kong.
Vacuum achieved in working chamber was below 1.2x10-5 Torr and the substrate
temperature during deposition was) 400 oC. For substrate etching, Ar gas was fed into the
chamber in an Ar flow rate of 10 sccm for 20 min (by means of 30 W DC power.
Afterwards, deposition proceeded for 1 h in an Ar/N2+O2 environment. The samples were
rotated at 4 rpm and positioned at a distance of 23 mm from the target. The temperature
(400 oC), power (400 W), and deposition time (1 hour) were kept the same while the
nitrogen to oxygen flow rate ratio was changed (1:3, 3:1, 1:1) and the gas flow was constant
at 5 sccm.
56
Figure 3.1: Working diagram of ultra-high vacuum PVD system in CityU Hong
Kong
3.3 Plasma Immersion Ion Implantation and Deposition System
More than 120 laboratories around the world have PIII system due to its versatile
applications such as for strengthening the components of glass, metals, plastic, ceramics
and polymers. This equipment is also used for surface modifications of materials without
altering bulk properties to enhance the strength, produce wear and corrosion resistant, light
weight parts at much less time and cost. Subsequently, this processing system was
pioneered by Chu et. al. [1, 2] and Mizuno et. al. [3, 4] and commonly known as PIII
processing technique. Since then this technique is receiving popularity among researchers
for its numerous applications each year.
In PIII process, plasma of ion species to be implanted is form within the chamber
where target is directly immersed and high negative pulse (several kV) is applied. To
provide high potential the walls of the chamber are grounded. This negative potential which
is applied to target drifts the electrons away from the target and a positive ions sheath is
form around the target. With time, due to negative potential, the positive ions which are
57
present in the sheath are being dragged towards the target and get implanted on it from all
sides. Thus PIII, being a non-line-of sight technique, completely eliminates the complexes
and sophisticated mechanisms of CII and proves to be a comparatively simple and cost
competitive process technology.
Some of the advantages of PIII can be summarized as:
1. PIII equipment is relatively easy to maintain and operate.
2. Running cost is less.
3. Process time is independent of sample surface area and its size.
4. Samples of different size shape and weight can be treated.
5. Deposition along with implantation can be carried out at the same time, also etching
sputtering etc., and not just metals or semiconductor, even insulating samples can be
processed.
6. Number of samples can be treated simultaneously.
7. Surface of specimen can be implanted confirming uniform dose rate as well as good
conformity.
8. Implantation energy ranging 1 keV to 300 keV.
9. When insulating samples are treated charge build-up is alleviated by secondary electrons
indigenous to the plasma.
11. Low-temperature process.
12. Implantation of multiple species with multiple charges is possible in the same system
[5].
58
Figure 3.2: Schematic diagram of plasma immersion ion implantation (PIII)
Basic Mechanism of a PIII System
The schematic diagram of a PIII [6, 7] with its main building blocks is shown in
figure. 3.4. It comprises of working chamber, gases inlet, plasma generating setup, target
holder and a high voltage modulator. A pulse generator with varying pulse durations is used
to provide a negative bias (ranging from 1 to 300 kV) to the target. The magnitude of this
bias voltage is usually limited by the limitations of pulse modulator and the process related
constraints. Other basic accessories like, Langmuir probe to measure plasma density and
electron temperature, ionization gauge to measure neutral density during implantation, IR
pyrometer to monitor target temperature during implantation, etc. are also parts of a PIII
system. To implant the ions in a PIII system [8-10], the target to be treated is placed on an
isolated and properly insulated target stage in the vacuum chamber. Plasma of ions of
desired species, to be implanted on to the target is generated and a high voltage (negative),
with respect to the chamber walls is applied to the target.
59
Figure 3.3: Working diagram of PIII&D System in CityU Hong Kong
3.4 Pulsed DC Magnetron Sputtering System
Experimental work for deposition of aluminum oxide on titanium substrate has been
done in Plasma Technology Laboratory, Department of Physics GC University Lahore. The
chamber design for deposition is shown in figure 3.6. It consists of working chamber
(stainless steel) having cylindrical shape, height and diameter of this chamber are 32 cm.
Two electrodes arranged in parallel configuration 3 cm apart (adjustable if needed) and
60
surrounded with ceramic to avoid the additional discharge. One of these electrodes serve
as anode and other is cathode. Permanent magnets with alternating polarity are arranged
behind the cathode to form magnetron arrangements. Four ports (diameter of 10 cm each)
are used for multipurpose. One of these ports serves as view window to monitor the status
of plasma. Other ports are used to insert different probes such as Langmuir probe optical
fiber etc. Few small ports for gas inlet also attach to vacuum pump also to connect pressure
gauge to monitor the gas pressure. To create vacuum inside the chamber rotary and turbo
molecular pump is used to achieve pressures 10-2 to 10-3 mbar and < 10-6 mbar respectively.
Pulsed dc power is given to electrode as shown in figure. 3.4. It is applied to the bottom
electrode through the inductive load.
Figure 3.4: Schematic of chamber with probe ports
3.5 Sample Preparation
Sample Size, Polishing, grinding should be up to the mark before exposing these
samples to plasma. For this purpose first of all dimensions of samples should be appropriate
and according to requirement, because many thin film characterization techniques required
specific sample size such as tensile testing machine required long bar or strip or wire shape
samples, AFM, SEM, EDX need small size etc. Also make sure that the surface of samples
61
should be clean, smooth, scratch free and mirror like polish. For this the samples are
mechanically polished using SiC abrasive paper from (200 ~ 1200) on polishing machine.
The use of water during polishing is also necessary to avoid temperature effect which in
turns produces stress. Samples were then cleaned with ethanol, acetone sequentially for 10
min in an ultrasonic bath, and air dried.
3.6 Thin film Characterization Techniques
Different characterization techniques were performed to investigate the surface
properties of deposited films.
3.6.1 X-Ray Diffractometer
The small twisting of light as it passes around the boundary of an object is called
diffraction. Diffraction can be observed in all types of waves like water waves, sound
waves, EM waves etc.
XRD is the most helpful in the characterization of crystalline materials such as
metals, intermetallic, ceramic objects, minerals and inorganic composites. It is a quick and
nondestructive procedure which provides information on unit cell measurement. Crystal
structure, grain size, phase identification, micro strain, texture, disorder and imperfection
in the crystal, thin film composition, even complex structures such as DNA and proteins
can be analyzed by this useful technique.
Bragg’s Law
Bragg’s law is the basic working principal to identify the structures of crystals and
molecules using x-ray diffraction. This law narrates the spacing between the neighboring
crystal planes and the angle of diffraction. nλ = 2dsinθ (Bragg’s law)
Here θ= angle of scattering λ= wave length of X-ray
beam d= spacing between crystal planes and n
is an integer for constructive interference.
62
Figure 3.5: Braggs Diffraction.
Working of XRD
The main components of X-ray diffraction are
• X-ray source
• A machine for restricting wavelength series called “goniometer”
• Sample holder
• Radiation detector
• Signal processor and readout. Modes of Operation
The orientation of the observed crystal plane depends on scanning mode. For XRD, two
scanning modes can be used.
(i) θ-θ scan
(ii) θ-2θ scan,
63
Figure 3.6: Schematic representation of XRD[11]
3.6.2 Scanning Electron Microscopy
In this technique, a spotlight consisting of electrons having high energy to interacts
with the exterior of the sample to produce a variety of signals which provide the information
about surface topography and composition of the sample. Mostly a two dimensional image
is obtained after getting the data over a preferred area of the sample. By using this
technique, very small areas of width ranging from 1cm to 5 microns can be imaged. SEM
has resolution of 50 to 100 nm and magnification ranging from 20X to 30000X.
Basic Principal of SEM
A high energetic beam of electrons is made to fall on the desired area of the sample.
After decelerating the incident electrons in the sample, a variety of signals are obtained due
to the electron-sample interaction. SEM images are formed due to secondary electrons.
Back scattered electrons reveals the rapid phase discrimination. Crystal structures and
orientations of minerals can be determined by diffracted back scattered electrons. Due to
X-Rays generation, SEM is considered to be a non-destructive technique. Hence a sample
can be characterized several times without any volume loss.
64
Working of SEM
The main components of SEM are
• Electron gun (filament)
• Lenses
• Scan coils
• Sample stage
• Detectors
• Vacuum system
Figure 3.7: Emission of electrons and x-ray, when beam of electron fall on specimen
As instead of light, electrons are used to form SEM image. Therefore from electron gun, a
beam of electron is made to fall vertically through the column of the microscope. After
passing through the lenses, a focused and directed beam hits the sample. After hitting the
sample, secondary and back scattered electrons and X-Rays are produced. The emitted
secondary and back scattered electrons are collected and converted into the signals by the
detectors and finally these signals are sent to the viewing screen to form the image of the
sample.
65
Figure 3.8: Systematic diagram of Scanning electron microscope[12] 3.6.3
Energy Dispersive X-Ray Spectroscopy (EDS or EDX)
It is a chemical micro-analysis technique that can be coupled with some applications
including TEM, SEM and STEM. The EDS technique provides information about
elemental composition by detecting the characteristics x-rays. The EDS can be utilized for
sample area having diameter in nanometer.
The working principle of EDS is very much related with the SEM. The SEM’s
electron beam hits the sample surface, core electrons are emitted. Due to removal of core
electrons, core holes are created. The core holes are filled by the higher state electrons and
as a result difference in energy is emitted as x-rays. The emitted x-rays energy represents
the characteristics of the particular element of the sample.
66
Figure 3.9: Schematic representation of electron-beam sample interaction.
The EDS used a detector that measures the x-rays correspond to their energy. The
incident x-ray hits the detector and as a result a charge pulse proportional to x-ray energy
is created. The charge pulse is then converted into a voltage pulse by using a preamplifier
which is charge sensitive. The voltage pulses are then sent to analyzer. The multichannel
analyzer sort outs the pulses by voltage. The energy data is then sent to computer for further
data evaluation and display. A spectrum is obtained by using energy of x-ray versus counts.
The elemental composition of the sample can be determined by the evaluation of spectrum.
Sample Requirements – The sample must be suitable in a moderate vacuum environment
(pressures of 2 Torr or less). The ideal sample size in the SEM is in diameter is up to 200
mm (8 in.) (200 mm) whereas large sample size in diameter is up to approximately 300 mm
(12 in.), can be loaded with the limited stage movement. The maximum sample height is
of approximately 50 mm (2 in.).
Applications – The EDS is used for analysis of foreign material, evaluation of corrosion,
analysis of coating composition, phase identification etc.
67
Figure 3.10: Schematic representation of SEM with both energy dispersive and
wavelength dispersive X-Ray spectroscopy systems
3.6.4 Atomic Force Microscopy
It is a kind of high resolution scanning probe microscope which is used to measure
local properties like stiffness, hardness, friction and elasticity etc with a probe. A high
resolution image (upto the nanoscale) of any type of surface counting polymers, ceramics,
composites, glasses and even biological samples (DNA molecules and protein crystal
growth) etc can be obtained by this technique.
Basic principal of AFM
AFM works on the same principal as the profilometer does. A sharp tip (probe) at
the end of a cantilever scans the surface of the sample. This probe has length of a couple of
microns and diameter of 100Aº while the cantilever has length ranging from 100-200 μm.
68
Normally cantilever is made of silicon or silicon nitride. When the tip is come closer to the
surface of the specimen, cantilever deflects according to Hook’s law due to the forces
between the tip and the sample. Hence by measuring these forces (attractive or repulsive),
sample can be analyzed.
Working of AFM
The basic parts of AFM are
• Laser
• Mirror
• Photo detector
• Probe
• Cantilever
• Piezoelectric scanner
Figure 3.11: Schematic diagram of atomic force microscopy[13]
As the probe moves over the surface of the sample, the cantilever deflects due to
inter atomic forces between the tip and the sample surface. A laser beam is passed on
the cantilever which reflects due to the orientation of the cantilever. This reflected laser
69
beam is detected by a detector. Finally the output of the detector is transferred to a
computer for processing the data to provide a topographical image of the sample.
Modes of Operation
AFM operates in three scanning modes, depending on the applications.
• Contact mode,
• Non-contact mode
Tapping mode.
As the attractive or repulsive forces cannot be calculated directly, so Hook’s law is used to
measure these forces which is given as
F= - kz
Here
F =force to be measured
K= stiffness of the cantilever
Z= deflection of the cantilever
70
Figure 3.12: Graph of force versus distance
3.6.5 X-Ray Photoelectron Spectroscopy (XPS)
It is an analysis technique, describes the chemical information of the surfaces of the
solid materials. The surface area of conductors and insulators from few microns to
millimeters can easily be analyzed by XPS.
Methodology
For analysis purpose, sample is placed in the ultrahigh vacuum environment. The
sample surface is then irradiated with soft x-ray. The x-rays then eject core electrons from
the sample atoms, also known as Photo-electrons.
The core electrons are ejected and as a result, core holes are produced. The core holes are
filled by the outer electrons and in this transition (from higher to lower level), energy is
emitted. The emitted energy is responsible for auger electron emission or characteristics x-
rays
.
71
Figure 3.13: Schematic representation of principle of XPS
The energy of the photo-electrons of an atom corresponds to its binding energy. The
binding energy depends on composition of the atom. The speed or kinetic energy of the
electrons ejected from the sample represents the energy coming from incident energy minus
binding energy. The binding energy can be calculated if speed of emitted electrons is
known. A spectrum is obtained, representing the surface composition using energy of
photoelectron and auger electron. The energy corresponding to spectrum peak represents
the composition or elements of the sample volume. The relative amount of each element is
calculated by area under the spectrum peak. The chemical state of the element can be
determined by the peak precise position and shape.
XPS detects those electrons which are emitted near the sample surface. The
photoelectrons with more than 20 to 50 Å can’t escape from the surface with sufficient
energy, so difficult to detect by the detector.
72
Figure 3.14: Basic structure of x-ray Spectroscopy system [14]
Survey Scan – In the survey scan, the peaks energy represents the elemental compositions
of analyzed surface. Except helium and hydrogen, all elements are detected.
Quantitation – The concentrations of sample elements can be determined by taking are
under the specific peak of element. For quantitative analysis of sample surface, area under
the peak is multiplied with the sensitivity factor.
Depth Profile - The composition of elements are determined (in terms of depth) into the
specimen. For this purpose, the AES analysis is alternated with ion sputtering and it
removes material from the surface of specimen.
Sample Preparation- The sample for XPS analysis requires a specific preparation. The
sample size in any lateral direction should not exceed 25 mm (1 in.). The sample height is
in the range of 12 mm (½ in.). The sample must be ultra-high vacuum (>10-9 Torr)
environmental friendly.
73
Applications- The XPS can be used for the analysis of contaminated thin film, quantity of
elemental composition, for chemical state identification etc.
3.6.6 Nanoindentation
Nanoindentation technique is mostly used to find out the mechanical properties of
solids like hardness and young modulus stiffness toughness etc. with very high resolution.
Basically in these method values of force and displacement is measured during the elastic-
plastic contact of a hard reference body (Indenter) with the sample. Hence it is almost
similar to micro-hardness-testing technique, but it offers some significant advantages over
micro-hardness:
very high depth and spatial resolution
no optical surveying of the indents necessary
measurement of hardness and Young’s modulus simultaneously
generation and measuring of exactly definable scratches The applications
of nanoindentation include:
Characterization of thin films coating, surfaces and microstructures
hardness
Young’s modulus
fracture toughness
scratch hardness
Adherence
determination of storage- and loss modulus
time dependence of scratches
determination of creep behavior
single grain analysis
74
Measurement method in detail
To calculate the important parameters like hardness (H) and young modulus (E) of
the thin films it is necessary to calculate elastic stiffness of the contact.i.e S. Where S is
measure from the slope of graph between load-unload and displacement. But from this
method one can find out the values of S, hardness and young modulus only at maximum
penetration depth only. Whereas in CSM method we can measure values during loading
and not just at the point of initial unload. In this method, continuous data obtained as a
function of penetration depth. This method is very helpful for thin films on substrates.
Figure 3.15: Schematic diagram of nanoindentor [15]
75
Figure 3.16: Nanoindentor XP [16]
3.6.7 Universal Testing Machine
The mechanical behavior of untreated and treated specimens is investigated by the
instrument known Universal Test Machine 50 KN AG-1 Shimadzu Autograph (UTM),
installed at CASP, GC University Lahore. We can conduct the different types of test like
bending, compression and tensile using the said machine. The machine is equipped with a
computer and results of various test performed is taken by the Trapezium software.
The machine has ability to adjust the different parameters as cross- head speed,
diameter and gauge length of specimens before initiating the test. First of all, the diameter
of specimen is measured using the vernier calliper. After that, the specimen is placed
between the two jaws of machine. The value of diameter, gauge length and cross head speed
is adjusted in the software. When the test was started stress-strain curve is plotted on the
software of machine. We can also find out the YS, ductility, UTS, fracture stress using the
data obtained from machine. Moreover, machine has also ability to perform the test from
room temperature to low temperature up to -200 oC and high temperature up to 300 oC.
76
Figure 3.17: Universal Testing Machine
3.6.8 Wear Test
Wear test is performed to investigate the wear mechanism and to predict the wear
performance. This evaluation provides sufficient information about material either it is
suitable for specific wear application or not. For surface engineering analysis, the test is
performed to check the potential of using a specific surface engineering technology so as
to reduce the wear for a particular application and to assess the effect of processing
parameters on the wear performance in order to obtain optimized surface treatment
conditions.
77
Working Principle
The pin on disc wear tester mainly consists of a pin and a flat disc specimen. The
pin is loaded and is in the perpendicular direction of the rotating disc specimen. This
combination produces a circular wear path. The pure sliding conditions are used to evaluate
the material properties such as wear and friction etc. One serves as specimen whereas other
as counterface. Different geometries of pin are used. The most common way is to use pin
as counterface and the geometry of pin is ball-like. The ball is composed of commercially
available materials such as tungsten carbide, bearing steel and alumina (Al2O3). The name
of the machine is based on usage of ball, so called as ball-ondisc tester.
Procedure
The specimen should be cleaned properly prior the test. Then measured the
specimen dimensions (nearest to 2.5 μm) or weigh the specimen (nearest 0.0001 g).
Introduced the pin in the pin-holder and adjusted it in a direction perpendicular to the disc
surface for necessary contact condition. Then proper mass was added to the lever system
so that to achieved the desired force which pressed the pin against the disc. The motor speed
and revolution counter were adjusted according to desire speed and number of revolution.
Then test was performed while specimen was in contact under load. After performing the
test, the specimen was removed and re-measured the specimen dimension or re-weighed
the specimen.
Wear Measurement/Quantification
The wear measurement is performed to investigate the quantity of material loss after
a wear test. In reality, the mass loss is determined for a specific time period. The removal
of material can be expressed as mass (weight) loss, linear dimensional change, volume loss,
wear type, the size and geometry of the specimen and availability of measurement facility.
Mass loss
The mass loss can be calculated by weighing the sample before and after the wear
test. The change in mass or weight represents the mass loss by the wear test. The weighing
machine is a precision balance and mostly suitable for sample surface which is
78
unsymmetrical and irregular in shape after wear test. The unit used for mass loss is gram
(g) or milligram (mg).
Figure 3.18: Schematic representations of pin-on-disc wear test system [17]
Volume Loss
The volume loss is calculated for a uniform material with known density. The
volume loss is measured in mm3 or μm3 for pin and disc separately. For volume loss, the
initial spherical radius of the pin is R and initially the disc is flat. The condition is that one
of the two members shows significant wears. The assumption is that disc does not have
significant wear.
79
Linear dimension loss
The linear dimension change gives a better understanding in many engineering
situations. The certain dimension e.g. length, diameter or thickness is more important to
the standard function of the engineering system. For linear dimension loss calculation, a
surface profilometer such as a stylus type, microscope or micrometers are used; the unit to
express the linear dimension loss is μm or mm.
Wear rate
Wear rates can be estimated by the wear volume loss, mass loss or change in linear
dimension under unit sliding distance or unit applied force. Different ways are used to
express the wear rate.
Wear resistance
The wear resistance is described as the resistance offered by the material to the
wear. There is no specific unit for wear resistance. It can be estimated by the inverse of
volume loss or mass loss. The relative wear resistance can be calculated as the ratio of wear
loss of a reference material over that the wear loss of the investigated material under same
testing parameters.
Sample Preparation and Test Specimens
A variety of materials can be used for wear test. The material should have specific
dimension and it should withstand with stress imposed during the wear test. The materials after
wear test can be described by dimensions, material type, surface finish, composition, form,
processing treatments, microstructure, and indentation hardness.
3.6.9 Preparation of simulated body fluid (SBF)
To study the apatite growth and integration of bioactive material in vitro
environment Kokubo and his colleagues prepared the solution named as SBF [18]. SBF is
a solution which contains inorganic ion concentrations almost similar to those of human
extracellular fluid, this solution used to reproduce apatite formation on biomaterials in
vitro, it can be used as apatite coating on various materials under appropriate conditions
80
also for evaluation of bioactivity of artificial material in vitro. The comparison between
human blood plasma and simulated body fluid is listed in table 3.1
Table 3.1: Ion concentrations in human blood plasma and SBF solution
Concentration (mmol/dm3)
Ion Human blood Plasma SBF
Na+ 142 142
K+ 5 5
Mg2+ 1.5 1.5
Ca2+ 2.5 2.5
Cl- 103 147.8
HCO3- 27 4.2
HPO42- 1 1
SO42- 0.5 0.5
The pH of this solution is adjusted to 7.25 at temperature 36.5 oC by using 50 mM
(=mmol/dm3) of tris (hydroxymetyl) aminomethane and approximately 45 mM of HCl.
And if the ability of specimen for apatite formation is not so good than pH of SBF is
sometimes adjusted to pH 7.40.
3.6.10 Cell Culture Observation
Cell culture means that first removal of cells from parent body i.e. plant or an animal
and then their growth on artificial substrate under appropriate conditions. These conditions
are different for different cells, but generally the environment for culturing the cells are
comprises of appropriate vessel which contains substrate or medium (amino acids,
vitamins, minerals, carbohydrates), gases like oxygen and carbonate, hormones, growth
factors. Other parameters such as pressure, temperature and pH are also regulates.
The major application of cell culturing is that it provides information about biochemistry
and physiology of cells, the effects of harmful compounds and drugs on the cells. Cell
culture also used in manufacturing of biological compounds like therapeutic proteins and
vaccines. The most important advantage of using cell culture is reproducibility and
81
consistency of results that can be achieved from using batch of clonal cells. Furthermore
the environment for cell culture laboratories is free from pathogenic microorganisms (i.e.
such as virus, parasites, bacteria and fungi), also some basic equipment that is necessary
for cell culture laboratory like Cell culture hood, incubator, water bath etc, for culturing the
cells following procedure is needed:
First of all sterilize the samples than soak the samples within 70% alcohol and washed with
PBS. Than MC3T3-E1 pre-osteoblasts cells were seeds on substrate in a highglucose
DMEM supplemented with 10% FBS in 75 cm2 tissue culture flakes and were incubated in
a humidified atmosphere of 5% CO2 at 37° C. 50,000 cells in a volume 1 mL were seeded
on each sample on 12-well plates for cell morphology observation and proliferation assay.
After culturing for 24 h, the cells were rinsed with PBS and fixed with 4%
paraformaldehyde in a humidified atmosphere of 5% CO2 at 37 °C for 20 min.
After fixing, the cells were stained with FITC-Phalloidin and 4′, 6-diamidino-
2phenylindole (DAPI) sequentially and then observe the cell morphology by fluorescence
microscopy.
Figure 3.19: The basic layout for cell culture hood
82
3.7 References
[1] P. K. Chu, S. Qin, C. Chan, N. W. Cheung, and L. A. Larson, Mater. Sci. Eng. R
17 (1986) 207-280.
[2] N. W. Cheung, Mater. Chem. Phys. 57 (1998) 1-16.
[3] B. Mizuno, I. Nakayama, N. Aoi, and M. Kubota, 19 th Conference on solid state
devices and materials, Tokyo, Japan (1987) 319-322.
[4] B. Mizuno, I. Nakayama, N. Aoi, M. Kubota, and T. Komeda, Appl. Phys. Lett. 53
(1988) 2059-2061.
[5] Dushyant Gupta, International Journal of Advancements in Technology, ISSN
0976-4860.
[6] J. R. Conrad, U.S. Patent 4764394, Wisconsin Alumni Research Foundation,
Madison, WI, 1988.
[7] J. R. Conrad, J. L. Radtke, R. A. Dodd, F. J. Worzala, and N. C. Tran, J. Appl.
Phys. 62 (1987) 4591-4596.
[8] A. Anders, Surf. Coat. Tech., 93 (1997) 157-167.
[9] D. J. Rej, in D.A. Glocker, S. I. Shah (Eds.), Handbook of thin film process
technology, IOP, Bristol, (1996), E2.3:1- E2.3:25.
[10] J. V. Mantese, I. G. Brown, N. W. Cheung, and G. A. Collins, MRS Bull., 21 (1996)
52-56.
[11] http://lipidlibrary.aocs.org/Biochemistry/content.cfm?ItemNumber=40299.
[12] https://www.purdue.edu/ehps/rem/rs/sem.htm.
[13] http://artsci.ucla.edu/BlueMorph/researchAFM.html.
[14] https://en.wikipedia.org/wiki/X-ray_photoelectron_spectroscopy.
[15] http://www.saint-gobain-recherche.fr/svi/Nanoindentation/NoteNano.htm.
[16] http://composites.usc.edu/facilities/mechanical-testing/nano-indenter-agilent-
mtsxp.htm.
83
[17] http://emrtk.unimiskolc.hu/projektek/adveng/home/kurzus/korsz_anyagtech/1_ko
nzultacio_ elemei/wear_testing_measurement.htm.
[18] Ohtsuki C., How to prepare Simulted Body Fluid, http://mswebs.naist.jp/
LABs/tanihara/ohtsuki/SBF/index.html.
Chapter 4
4.1 Synthesis of Titanium Oxynitride on NiTi 4.1.1-Introduction
Shape memory effects and superelasticity are the two main properties of NiTi shape
memory alloys (SMA) which make them important materials for different applications [1–
84
3]. Due to their high corrosion resistance, good biocompatibility, and excellent mechanical
properties, NiTi alloys have been widely used in medical fields, such as orthodontic arc
wires, cardiovascular stents, and orthopedic devices [4–8]. However presence of Ni, limits
NiTi use in medical filed. It is observed that Ni ions are toxic and produced the allergic
reaction. It may also be carcinogenic in case of prolonged exposure. Therefore, to increase
the biocompatibility of NiTi alloy, Ni ions release should be minimized from NiTi alloy.
For this purpose, surface modification of NiTi alloys are carried out by different techniques
such as sol-gel method, plasma immersion ion implantation, plasma spray, and plasma
deposition [9–12]. The key advantage of plasma surface modification is its ability to
improve only selective surface properties, whereas the bulk properties remain the same
[11]. Several research groups have made great effort to enhance the bioactivity of NiTi
alloy. Liu et al. [9] mentioned that the formation of TiO2 layer on NiTi substrate, improved
the corrosion and blood compatibility. Sun et al. [10] concluded that (Ti, O, N)/Ti
composite coating was best for medical applications because it improves the
biocompatibility, wear resistance, and mechanical properties of NiTi alloy. More recently,
Hossary et al. [13] found that plasma nitriding formed the TiN layer, which not only
enhanced the corrosion resistance, but also effectively prevented the Ni release from the
NiTi surface.
From the literature, it is clear that biocompatibility of NiTi alloys are mostly studied
after formation of oxide or nitride layer on NiTi surface. Recently, metallic oxynitrides
have attracted much scientific interest owing to their diverse properties [14– 16]. In current
research work, we investigate and compare effects of the gas flow rate on the properties of
TiON films deposited by reactive magnetron sputtering. The surface chemistry and
morphology are evaluated and discussed together with the biological results.
4.1.2. Experimental details
Ni (50.8 at. %) Ti alloys (SE508) samples purchased from USA, were used as
substrate in this experiment. The samples grounded with 200, 400, 600, 800, 1000, and
1200 grit SiC paper, cleaned with ethanol, acetone, and distilled water sequentially for 10
min in an ultrasonic bath, and air dried. A round-shaped Ti (99.99%) purity with 0.25 inch
thickness and 2 inch diameter are used as target for sputtering. Deposition was performed
85
in the ultra-high vacuum PVD system (magnetron system) in the Plasma Laboratory at City
University of Hong Kong. The base vacuum was below 1.2 x 10 -5 Torr and the substrate
temperature during deposition was 400 oC. High purity argon was fed into the chamber to
generate the plasma generation and the substrate was first etched in an Ar flow rate of 10
sccm for 20 min at a DC power of 30 W. Afterwards, deposition proceeded for 1 h in an
Ar/N2+O2 environment. The samples were rotated at 4 rpm and positioned at a distance of
23 mm from the target. Oxygen and nitrogen were used as reactive gases to form titanium
oxynitrides along with argon as sputtering gas. The temperature (400 oC), power (400 W),
and deposition time (1 hour) were kept the same, while the nitrogen to oxygen flow rate
ratio was changed (1:3, 3:1, 1:1) and the gas flow was constant at 5 sccm. XRD was
employed to examine the change in the preferred orientation on the X’Pert Pro
diffractometer using Cu Kα (1.541 A°) radiation at 40 kV and 30 mA. The chemical
composition of the outermost layer of the films was determined by x-ray photoelectron
spectroscopy (XPS) (PHI Model 5802) performed at a base pressure of 10 -8 Torr using
MgKα x-ray (1253.6 eV). A Gaussian peak fitting model was adopted to deconvolute the
spectra using XPSPEAK41 software. Atomic force microscopy (AFM, Park Scientific
Instruments) was used to determine the surface topography and roughness of the sample
surface before and after deposition. Scanning electron microscopy (SEM, JEOL JSM-820)
and energy-dispersive x-ray spectrometry (EDS) were utilized to examine and analyze the
surface morphology and chemical composition of the specimens.
The biological experiments were conducted with MC3T3-E1 pre-osteoblasts. The cells
were cultured in a DMEM supplemented with 10% FBS in 75 cm2 tissue culture flakes and
were incubated in a humidified atmosphere of 5% CO2 at 37 oC. 50000 cells in a volume 1
mL were seeded on each sample on 12-well plates for cell morphology observation and
proliferation assay. After culturing for 24 h, the cells were rinsed with PBS and fixed with
4% paraformaldehyde in a humidified atmosphere of 5% CO2 at 37 oC for 20 min. After
fixing, the cells were stained with FITC-Phalloidin and 4 6diamidino-2-phenylindole ,׳
(DAPI) sequentially and then were observed by fluorescence microscopy.
86
4.1.3 Results and discussion
XRD Analysis
The XRD patterns are depicted in Fig 4.1. It is clear that untreated NiTi sample
consists of three peaks at 2θ positions of 42.56o, 61.80 o, and 77.9 o corresponding to (110),
(200), and (211) phases for cubic structure, respectively. In the deposited samples, TiO,
TiN, Ti2N, and TiO2 phases are observed.
The existence of these new phases in deposited samples confirms the formation of titanium
oxynitride film on the NiTi substrate. Since TiO or TiN has the same crystallographic
structure (cubic) with very close lattice parameters as calculated for (200) and (111) phases,
(i.e., αTiN = 0.427 nm and αTiO = 0.429 nm), XRD is not the suitable technique to
distinguish these two structures [17]. The occurrence of TiO2 at angels 36.7 o and 54.4 o
having planes (211) and (202), respectively, indicates the anatase phase, whereas planes
corresponding to (103) and (213) at angels 36.7 o and 61.6 o, respectively, are rutile.
Presence of these phases is useful to improve the bioactivity as reported by Shu et al. [18]
and is therefore favorable for apatite growth. Liu et al. [9] also observed that TiO2 layer
formed on NiTi substrate consists of anatase and rutile phases, enhancing the
biocompatibility of NiTi alloy. Moreover, with an increase in oxygen concentration, the
intensity of R-TiO2 (202) also increases. This increase in intensity may be attributed to
more oxygen incorporation at higher concentration of oxygen because the probability of
ionized oxygen is larger as compared to that of nitrogen. The evolution of the average
crystallite size of (200) can be estimated using Scherer’s equation [19].
where K is a constant that depends on cell geometry (shape factor) and its value varies
between 0.9–1, λ is the x-ray wavelength and is equal to 1.54 A°, the FWHM is the full
width at half maxima of particular diffraction peak, and θ is Bragg’s angle of the diffraction
peak.
87
Figure 4.1: XRD patterns of TiOxNy films deposited at different oxygen flow rates
The results are 112, 21, and 74 nm for oxygen flow rates of 5, 10, and 15 sccm, respectively.
The minimum value of crystallite (i.e., 21 nm) is observed at the same flow rate of nitrogen
and oxygen. This variation trend of average crystallite size with increasing oxygen
concentration may be due to variation in oxygen incorporation.
XPS Analysis
The XPS survey scan spectra are acquired over a binding energy range from 0 to
1400 eV with an energy step of ∆E = 0.8 eV (Fig. 4.2). The peaks are assigned to Ti, N, O,
and C (surface contamination). The peaks within range from 1230 eV to 976 eV correspond
to the Auger peaks of Ti LMM, OKLL and CKLL [20]. The main component of the C 1s
peak at 284.93 eV corresponds to the C–C bonding. This carbon is detected due to the
surface contamination from the atmosphere. No nickel peak is observed due to the titanium
oxynitride surface layer which forms a barrier against Ni out-diffusion from the substrate
to minimize allergic and toxic effects [21]. As the presence of oxygen in TiN coatings can
promote the formation of bone-like materials in vivo [22]. In fact, a titanium oxynitride
film on the surface is favorable to calcium phosphate nucleation due to the presence of
mixed-valence states of the surface Ti atoms and localization of negative charges on the
88
oxidized surface. The high-resolution spectra from 450 eV to 475 eV for Ti 2p, 392 eV to
408 eV for N 1s, and 525 eV to 542 eV for O 1s for samples deposited at different p (O2)/p
(N2) flow rates are presented in Figs. (4.3–4.5). The core level spectra of the films deposited
at different flow rates can be convoluted to identify various compounds/phases that have
been formed on the surface of composite thin films such as nitrides, oxides, and oxynitrides
in the film as given in Figs. (4.3–4.5). As shown in Fig. 4.3, the deconvoluted peaks of Ti
2p into Ti 2p3/2 and Ti 2p1/2 found at 459.15 eV and 458.9 eV, respectively, were identified
as Ti4+ in TiO2 [23]. As shown in Fig. 4.4, the O 1s spectrum shows a shoulder peak on the
high binding energy side and it can be deconvoluted into two peaks at 533.30 eV and 532.23
eV which are assigned to TiOxNy with the hydroxyl (O–H) and (C–O–C) groups [20, 24].
The N-1s spectrum in Fig. 4. 5 shows Ti–O at binding energies of 396.2 eV, Ti–N–O at
397.4 eV [17, 25], whereas the peaks at 400.3 and 402.20 eV are due to N–O bonds [26,
27]. Hence, the TiOxNy films contain a mixture of Ti–N–O, Ti–N, and Ti–O chemical states
and the chemical composition can be calculated using the integrated area of the Ti-2p, O-
1s, and N-1s peaks [28].
89
Figure 4.2: XPS spectrum showing TiOxNy film growth
Figure 4.3: High resolution XPS spectra of Ti2p obtained from TiOxNy films
prepared at different oxygen flow rates: (a) 15 sccm, (b) 10 sccm, and (c) 5 sccm.
90
Figure 4.4: High resolution XPS spectra of O1s obtained from TiOxNy films
prepared at different oxygen flow rates: (a) 15 sccm, (b) 10 sccm, and (c) 5 sccm.
91
Figure 4.5: High resolution XPS spectra of N1s obtained from TiOxNy films
prepared at different oxygen flow rates: (a) 15 sccm, (b) 10 sccm, and (c) 5 sccm.
92
AFM Results
The surface morphology before and after titanium oxynitride deposition on NiTi
alloy was studied by AFM and given in Fig. 4.6. Figure 4.6(a) depicts that the surface of
untreated NiTi alloy is quite smooth and almost free of particles. However, the surface of
treated specimen consists of uniformly-distributed, large dome-shaped islands (spiky
cones) which become smooth and then again enhance with increasing oxygen flow rates
(Figs. 4.6(b)–4.6(d)). The occurrence of large dome-shaped islands confirms the formation
of titanium oxynitride film on substrate surface. The surface roughness parameters, namely,
Rq, Ra and Rz , before and after deposition are listed in Table 1.1. The 3D micrographs
indicate that the roughness of treated specimens is higher as compared to the untreated
specimen. Moreover, roughness increases with increase in oxygen concentration. It is
quoted in the literature that the rough surface is more favorable, to enhance the
biocompatibility of material, as compared to smooth surface [29]. Wirth et al. [30] specified
that a rough surface was more beneficial to stimulate the cell proliferation. In addition, it
does not affect the osteoblasts morphology. The increase in biocompatibility with
roughness may be due to rough surface providing more surface area for initial cell
attachment. The increase in available surface area can enhance the bioactivity, and thus the
rate of bone formation also improves. The increase in surface roughness of the NiTi alloy
after titanium oxynitride formation is attributed to the growth mechanism which is
influenced by both the sputtering and deposition effects [31].
93
Figure 4.6: AFM Surface morphology: (a) Untreated, (b) 5 sccm, (c) 10
sccm, and (d) 15 sccm oxygen flow rate.
Table 4.1: Roughness values of untreated and deposited films.
Parameters O2/N2
(sccm)
Rq (nm) Ra (nm) Rz (nm)
Untreated 7.98
6.46 3.96
5:15 8.11
6.44 9.65
10:10 8.95 7.21
17.9
15:5 19.2 15.5
17.8
94
EDX Analysis
EDS results after immersing in simulated body fluid (SBF) with a pH of 7.42 at 37o
± 0.5 oC for 3 days are shown in Table 4.2. The composition of SBF is: 7.996 g/L of
NaCl, 0.35 g/L of NaHCO3, 0.224 g/L of KCl, 0.228 g/L of K2HPO4.3H2O, 0.305 g/L of
MgCl2.6H2O, 0.278 g/L of CaCl2, 0.071 g/L of Na2SO4, as well as 6.057 g/L
(CH2OH)3CNH2 [32]. The results show the distribution of Ti, Ni, Ca, P, N and O,
confirming the presence of the deposited film. The concentration of nitrogen and oxygen
is significant as compared to that of other elements. Although the amounts of Ca and P are
insignificant, a large amount of surface nitrogen and oxygen acts as barrier against nickel
release. Sufficient blocking of Ni release from the NiTi-SMA is imperative to biomedical
implants because even a small amount of Ni (e.g., 0.5 ppm) has been reported to induce
allergic reactions [33] and strong adverse reactions of osteoblasts have been observed in
vitro [34].
Table 4.2: Elemental concentrations of the films.
Relative concentrations (at %)
Oxygen flow rate N O Ca P Ti Ni
Untreated 51.40 48.60
5 sccm 89.2 8.1 0 2.7 0 0
10 sccm 85 13 0 2 0 0
15 sccm 61.29 27.91 2.07 3.06 3.66 2.01
Cell Culture Study
In order to study the biological acceptability, mouse MC3T3-E1 pre-osteoblasts are
cultured. Figure 4.7 displays the fluorescent microscopy images of the cells after 3 days.
The cells can clearly attach and proliferate on the samples and no obvious cytotoxic effects
can be observed from the deposited samples. The MC3T3-E1 cells on the samples prepared
at 5 and 15 sccm oxygen flow rates have higher activity than those treated at 10 sccm. The
rectangular and elongated morphologies of the cells in Figs. 4.7(a) and 7(c) reveal the
tendency to occupy a larger area compared to the film deposited at the same flow rate of
oxygen and nitrogen in Fig. 4.7(b) which shows a triangular morphology. Titanium
95
oxynitride has formed a dense layer on the NiTi surface, which acts as barrier and stops the
Ni release from the substrate. The cell reaction is governed by the coating instead of the
substrate, thus providing the opportunity for accelerated tissue repair and regeneration [35].
Figure 4.7: Fluorescence microscopy images after culturing for 3 days: (a) 5 sccm,
(b) 10 sccm, and (c) 15 sccm oxygen flow rate.
96
4.2 Nano-structured Zirconium Oxide film using PIII&D
4.2.1 Introduction
Plasma immersion ion implantation and deposition (PIII&D) is becoming the
versatile technique to modify the surface properties of materials for the last decade [3638].
The unique advantage of PIII&D on other plasma surface modification techniques is that it
can perform simultaneously implantation, deposition and etching [36]. In addition PIII&D
is most efficient and economic technique used to enhance the surface properties of
materials. Recently, PIII&D is widely used in medical field to improve the bio-medical
applications of polymers, metals and its alloys [36]. Ti6Al4V alloy is most suitable material
for biomedical implants as compared to the pure Ti and its other alloys due to its high
mechanical strength, modulus of elasticity comparable to bone, good biocompatibility and
excellent corrosion resistance [39-40]. However, Ti6Al4V alloy exhibits the poor
tribological properties, which limits its use in load-bearing applications [41-43]. Liang et
al [44] reported that degradation of artificial implants like knee, elbow and hip joints
occurred after 10–15 years of use. The main causes behind this degradation are wear
failures. Moreover, it is observed that wear particles promote the corrosion process which
ultimate result in genetic damage [45]. Therefore, it is imperative to make these implant
materials wear and corrosion resistant. The wear resistance of implant material can be
enhanced by reducing the COF or increasing the surface roughness and hardness [46]. This
can be achieved by coatings of different materials on the surface of Ti and its alloy.
Dong and Bell [47] & Guleyuz et al [48] oxidized Ti6Al4V which resulted in the
formation of TiO2 layer on substrate surface improving the wear resistance. Tan et al. [49]
mentioned that the fretting wear of NiTi alloy significantly enhanced after the oxygen ion
implantation. In another study Wan et al. [50] found that the Ti6Al4V specimens coated
with TiO/TiN exhibited the good wear resistance as compared to the
TiN coated and uncoated specimens. More recently Obadele et al. [51] coated Ti6Al4V
with ZrO2 in Ti matrix which decreased the coefficient of friction and enhanced the wear
resistance.
97
Although many researchers have made considerable efforts to enhance the wear
behavior of Ti6Al4V alloy but to the best of our knowledge, no one has prepared the nano
structure zirconium oxide film at various voltages by PIII&D technique. Therefore the aim
of present study to form the nano structured ZrO2 film and its effects on tribological
properties of Ti6Al4V alloy.
4.2.2 Experimental details
Ti6Al4V samples are cut in dimensions of 10x10x2 mm3 and then are mechanically
grinded with SiC paper grit (200-1200) and ultrasonically rinsed in acetone, ethanol and
distilled water (10 min) each and then dried in air before its exposure to plasma. PIII&D is
carried out in the PIII housed in plasma laboratory, City University of Hong Kong. The
details of PIII equipment and implanter are mentioned in reference [52] and in chapter 3.
The vacuum in chamber was achieved by mechanical and turbo molecular pump. Argon
gas was used for cleaning the specimens by means of sputtering. To deposit Zirconium
oxide thin film, a pure zirconium cathode was used to produce zirconium ions. High purity
argon (Ar) and oxygen (O2) were introduced into the chamber to generate the plasma at 30
sccm and 20 sccm respectively. Specimens were treated at three different applied voltages
15 kV, 20 kV and 30 kV for 120 min treatment time.
The chemical compositions of the outermost layer of the films were determined by
XPS (PHI Model 5802) performed at a base pressure of 10-8 Torr using MgKα X-ray
(1253.6 eV). A Gaussian peak fitting model was adopted to deconvolute the spectra using
XPSPEAK41 software. The beam of Argon ion was used to sputter off deposited film upto
50 nm depths for the elemental analysis. The dry sliding wear tests were performed using
a pin-on disk machine (TEER Coating Ltd., Model POD-2). A standard WC ball with a
diameter of 5mm was used. The tests were conducted at a constant sliding speed of 300
mms−1 under loads of 2 and 7N. Roughness of untreated and treated specimens was
measured using the AFM. SEM (JEOL JSM-820) and EDS were utilized to examine and
analyze the surface morphology and chemical composition of the specimens. The
mechanical properties such as Young’s modulus (E) and hardness (H) were evaluated by
nano-indentation (MTS Nano Indenter XP, USA).
98
5.2.3 Results and Discussion
X-ray photoelectron spectroscopy (XPS)
Untreated and zirconium oxide deposited specimens at three different voltages 15,
20 and 25 kV are analyzed. Survey scan spectra (energy step E = 0.8 eV and binding
energy range from 0 to 1200 eV) of untreated and treated specimens are shown in Fig.4.8.
The peaks in untreated spectrum assigned as Ti 2p, Al (2p, 2s) and V 2p. Presence of O 1s
in untreated spectrum is also observed indicating surface contamination or native oxides
on the surface of substrate. The peaks at 976 eV correspond to the Auger peaks of OKLL
[53]. Spectrum of deposited specimens clearly shows the presence of significant peaks of
zirconium (Zr 3s, Zr 3p, Zr 3d and Zr 4p) along with oxygen (O 1s). For identification and
confirmation the presence of different compounds/phases, highresolution spectra for Ti 2p
over binding energy ranging from 550 eV to 570 eV, for Zr 3d (175 eV to 190 eV) and for
O1s (525 to 540 eV) are plotted and deconvolution of these is done by using Casa XPS
Peak41 software and results are shown in Fig. 4.9(a-c).
In Fig 4.9(a), four peaks assigned to titanium oxide at binding energies 455.1, 457.6, 460.1
and 461.8 eV, whereas two peaks show bonding between Ti-Zr-O at binding energy (BE)
values of 464.3 and 467.2 eV respectively. Similarly the deconvolution spectrum of Zr 3d
(Fig. 4.9b) also shows the presence of Zr-O bonding at BE values 186.7, 185.73 and 182.9
eV for Zr-O formation and Zr-Ti-O bonding at BE values of 184.15 eV. Furthermore, the
existence of Zr 3d5/2 and Zr 3d1/2 also observed at low BE values of 179.13 eV and high
BE value 181.7 eV respectively in Fig. 4.9 (b). Deconvoluted spectrum of O1s (Fig. 4.9c)
only show the presence of bonding of metal oxides which are Zr-O at BE of 531.2 eV, Ti-
O at BE of 532.6 eV and Al-O at BE of 530.4 eV. All deconvoluted peaks are matched by
standard data of NIST web site [54].
Fig. 4.10 shows the variation of XPS spectra of deposited films as a function of sputter time
for different voltages (15 kV, 20 kV and 25 kV). Sputter rate for all such spectrum plotted
in Fig. 4.10 are 6.955 nm/min upto 50 nm depth. It is quite clear in Ti2p spectra that less
amount of titanium is observed at surface of deposited specimens but with the gradual
increase in depth, amount of titanium improved. Also significant chemical shifting towards
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higher binding energy is observed in titanium oxidation states confirming the existence of
titanium dioxide (Ti4+) on the surface of specimen whereas pure metal (Ti) is observed at
lower binding energy. Similar trend has been observed for depth profile of Al 2p and V 2p
spectra that with an increase in the depth, vanadium and aluminum peaks also improved.
However in case of Zr 3d and O 1s the amount of zirconium and oxygen decreases with
increasing depth. The spectrum of treated specimen at 20 kV as shown in Fig. 4.10 (B-
series) exhibit the similar results as in (A-series), but variation in spectra of specimen
treated at 25 kV (C-series) exhibit different results shown in Fig 4.10. The spectra indicates
that presence of Ti 2p, Al 2p and V2p peaks are negligible at surface, with an increase in
depth these peaks are remains negligible. On the other side, the spectra corresponding to
Zr 3d and O1s show strong peaks even at deeper part i.e zirconium and oxygen peaks exist
at 50 nm depth which confirm the formation of relatively thick zirconium oxide film. It is
concluded from results that maximum deposition is achieved maximum voltage (25 kV).
Additionally shifting at different depths of the samples shows the formation of various
compounds in oxygen 1s spectrum.
Figures 4.11 (a-c) depict the variation in atomic concentration as a function of sputter time.
Atomic concentration of different elements with 4 keV Ar sputter time are shown in Fig.
4.11. Results indicate that atomic concentrations crossponding to Ti 2p, Al 2p and V 2p
increase whereas in case of Zr 3d and O 1s atomic concentrations decrease as shown in
Figs.4.11(a&b). By comparison of Figs. 4.11(a) and 4.11(b) no noticeable difference is
observed for sample treated at 15 and 20 kV, whereas the specimen treated at 25 kV (Fig.
4.11(c)) shows that Zr 3d and O 1s are almost stable and no significance of V 2p and Al 2p
are observed with depth, indicating more diffusion of Zr and O as compared to sample
treated at lower voltages. These results also supporting the results mentioned in Fig. 4.10.
100
101
Figure 4.8: Survey scan spectrum of (a) untreated Ti6Al4V (b) ZrO2 deposited
specimen
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Figure 4.9: Deconvoluted XPS spectrum of (a) Ti 2p (b) Zr 3d (c) O1s
Figure 4.10: Detail X-ray photoelectron spectroscopy (XPS) spectra of (a) Ti 2p, (b)
Al 2p, (c) V 2p, (d) Zr 3d, (e) O1s, Each individual graph consist of eight different
spectrum after successive Ar sputtering, (A-Series) specimens treated at 15 kV, (B-
series) at 20 kV, (C-series) at 25 kV
103
Figure 4.11: Graphical representation of Atomic concentration versus sputter time
(a) 15 kV (b) 20 kV (c) 25 kV
104
Atomic Force Microscopy
Atomic force microscopy (AFM) is employed to investigate the effect of different
applied voltages on deposition of zirconium oxide. Surface features (morphology)
untreated and deposited specimens are shown in Fig.4.12. The roughness values, Ra, Rq and
Rz are mentioned in Table 2.1. 3-D micrograph of untreated specimen shows some
scratches on the surface which may be produced during polishing as shown in Fig.4.12 (a).
Deposition at 15 kV and 20 kV as shown in Figs. 4.12(b& c) show almost similar
morphology showing uniformly distribution of nano structured layer but some hillocks are
also observed due to which roughness of the film increases as compared to untreated
sample. Although roughness is of the order of few nano meters but the trend indicates that
roughness increases with applied voltage. Surface morphology of specimen at 25 kV is
quite different showing waves like patterns/structures over all area. Maximum roughness
is achieved at 25 kV which is about 8.45 nm. The increase in surface roughness is related
to the growth mechanism which is influenced by both the sputtering and deposition effects
[55]. Actually, with the increase in applied voltage sputter and deposition rate increase
resulting in the formation of thick deposited layer with high roughness [56].
Figure 4.12: AFM micrographs of deposited film (a) untreated (b) 15 kV (c) 20 kV
(d) 25 kV
Table 4.3: Roughness values of deposited film by AFM analysis
105
Rq nm Ra nm Rz nm
Untreated 1.58 1.26 1.15
15 kV 2.51 1.94 0.535
20 kV 7.64 6.41 7.20
25 kV 8.45 7.02 1.55
Thickness
The average thickness of deposited films was measured at three different areas using the
ellipsometry (incident angle 70o). The influence of bias voltages on thickness of deposited
films is given in Fig. 6. The average thickness is found to be approximately 60, 83 and 108
nm for bias voltages 15, 20 and 25 kV respectively. Thickness increases linearly with the
bias voltages which are in agreement with the results reported by Wu et al. [57].
Figure 4.13: Variation in thickness at different voltages.
Wear Testing
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The instrument used to study the wear resistance is a conventional pin-on-disk
system. During wear tests specimens are rotated at a linear velocity of 300 mm/min for
2000 number of cycles. A tungsten carbide (WC) ball of 5 mm diameter is mounted at the
end of pin. All tests are conducted in air and without lubrication at normal loads of 2 N and
7 N. The wear rate K (mm3/N-m) of the deposited film is calculated by the given
relationship [58].
K
Where V (mm3) is the volume of the removed particles (debris), N (N) the normal applied
load, v (m/s) is the sliding velocity and t (s) is the test duration. The resultant wear track
left on the specimens is analyzed using a surface profilometer in order to accurately
determine its depth. The variations of coefficient of friction (COF) of untreated and treated
specimens as a function of sliding cycles (2000) against tungsten carbide (WC) balls are
shown in Fig. 4.14. Initial COF of untreated specimen is ~ 0.42 at load of 2 N and reaches
its maximum value ~ 0.8, whereas at 7 N load COF varies from ~ 0.3 to 0.5. Minimum and
maximum values of COF against the 2 N and 7 N loads are mentioned in Table 4.4. It can
be clearly seen that at 2 N load COF decreased in all treated specimens as compared to
untreated specimen. The significant decrease in COF is observed at 25 kV at 2 N load,
indicating that good wear resistant ZrO2 films can be achieved at 25 kV along with other
parameters. Referring to Fig. 4.14 (b), for all deposited specimens, friction coefficient
initially drops as compared to untreated specimen but becomes stable and attains the
comparable values as untreated specimen.
The variation in wear rate at the various applied voltages at 2 N and 7 N loads is given in
Fig. 4.15 (a-b). It is clear from the figure that the wear rate continuously decreases with
increase in the applied voltage for both loads in all treated specimens as compared to the
untreated specimen. The decrease in the wear rate in the treated specimens is due to
decrease in the COF. However, the prominent variation in the both wear rate and COF at
the 25 kV may be due to the formation of hard top layer on the substrate surface which is
also supported by the XPS depth profile results, where ZrO2 remains present on the surface
as compared to the other treated specimens. Wang et al [59] also prepared the zirconia film
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on zirconium alloy at various frequencies and found that a thick and hard film is formed at
maximum frequency 4000 Hz, which improved the wear behavior of substrate. In another
study Zdravecka et al. [60] also mentioned that the formation of hard film is responsible to
decrease in the COF and ultimate results in improvement of wear resistance. To achieve
better understanding, the wear tracks obtained using a surface texture tester is displayed in
Fig 4.16. In the wear mechanism, a material is removed from the surface of solid. The
material removed can be measured by means of depth profile.
The variation in depth and area of grooves of the untreated and treated samples are shown
in Fig. 4.17. The depths of these grooves are of the order of few micrometers. Graphical
representation of these grooves is shown in Figs. 4.17(a & b). Depth and area of these
grooves represent that untreated specimen exhibits deep and wide wear groove, whereas
the implanted specimens display narrow and relatively compact grooves. The improvement
and compactness of these grooves is caused by surface hardening due to ion implantation
[61]. From above discussion we conclude that 25 kV is the best voltage for anti-wear
zirconium oxide coating.
Table. 4.4: Min. and Max. of coefficient of friction.
Min-Max at 2 N Min-Max at 7 N
Untreated 0.42-0.8 0.3-0.51
15 k eV 0.32-0.75 0.14-0.51
20 k eV 0.31-0.77 0.14-0.51
25 k eV 0.13-0.49 0.13-0.49
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Figure 4.14: Friction Coefficient as a function of no of cycles. (A) at 2 N (B) at 7 N
(a) untreated, (b) 15 kV, (c) 20 kV, (d) 25 kV
109
Figure 4.15: Variation of wear rates with applied voltages at (a) 2 N (b) 7 N load.
110
111
Figure 4.16: Depth profiles of wear track as a function of different voltages (A) 2 N,
(B) 7 N, (a) untreated, (b) 15 kV, (c) 20kV, (d) 25 kV
112
Figure 4.17: Variation in depth and area of grooves at different applied
voltages, (a) 2 N (b) 7 N
SEM microstructure of worn surfaces of untreated and treated specimens after the wear
track at loads 2 N and 7 N are shown in Fig. 4.18- 4.19 respectively under 500 x
magnifications. Fig. 4.18(a) shows SEM micrograph of untreated spacemen, exhibiting
rough surface that shows adhesive wear behavior because many adhesive craters and
grooves are observed after wear ploughing. On the other hand the surface of treated
specimens is almost same except the film deposited at 25 kV which indicates wearing
relatively less surface when the counterpart slides as shown in Fig. 4.18 (d). The main
reason of less wear at 25 kV is due to its low COF and presence of more hard ZrO2 layer
on the top surface. Moreover, the less wear resulting in less debris from the surface, which
is desirable for bio-implant.
Again in micrographs at the normal load of 7 N as shown in Fig. 4.19, untreated specimen
showed spallation, whereas micrograph of treated specimens at 15 kV and 20 kV are almost
similar and show relatively better morphology than untreated but asperities on the surface
are deformed and worn out which may be due to rubbing which might be due their poor
adhesion of film to substrate. Wear track of specimen deposited at maximum voltage
exhibit better morphology as compared to specimens treated at lower voltages.
Nanohardness Measurements
The mechanical behavior is analyzed by nanoindentation (Nano Instruments XP,
MTS). The indentation experiments are conducted in displacement control to a depth range
upto 200 nm on each specimen to measure the film properties as shown in Fig.4.20. From
graphical representation, it is observed that with the increase in depth, hardness decreases
near the film substrate interface and become stable at other places. It is reported by many
researchers [62-64] that hardness increases with the decrease in depth. This is because of
thickness effect, as we know that thickness of our implanted specimens is about 100 nm.
Therefore, beyond the 100nm depth hardness results are due to the deposited layer as well
as the bulk substrate [65].
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114
Figure 4.18. SEM micrographs of wear track of untreated and treated
specimens at 2 N load (a) untreated (b) 15 kV, (c) 20 kV, (d) 25 kV
Figure 4.19: SEM micrographs of wear track of untreated and treated specimens at 7
N load (a) untreated (b) 15 kV, (c) 20 kV, (d) 25 kV
Fig.4.21 shows the hardness at the surface of thin films at different voltages. Each point in
Figure represents the average of three indentations. From the graph it is clear that
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improvement in hardness is observed for specimen implanted using PIII implanter. The
values of hardness are higher at higher applied voltage. The surface hardness varies from
5.5 GPa to 10.5 GPa, which is about 2 times compared to untreated specimen. This increase
in hardness might be due to higher deposition rate with increasing voltage which led to
thicker deposited layer [66].
The improvement in wear mechanism with increase in the hardness can also be explained
on the basic of archard equation [67].
Where
V = total volume loss during the sliding
K = wear constant, F = applied normal load, S = sliding distance, and H = hardness
The above equation also shows the inverse relation between volume loss and hardness.
With an increase in the surface hardness the amount of weight loss also reduces. Therefore,
in the present study the gradually decrease in the wear rate is also due to increase in the
surface hardness in all treated specimens. Since the hardness value of film formed at 25 kV
is almost twice as compared to the untreated specimen so a significant decrease in wear
rate at this voltage is observed.
116
Figure 4.20: Nanohardness as a function of indentation depth (a) untreated (b) 15 kV,
(c) 20 kV, (d) 25 kV
Figure 4.21: surface hardness as a function of applied voltage
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4.3 Mechanical properties and biocompatibility of Ti-Al-O
composite film
4.3.1-Introduction
In last few decades, significant struggles have been made to develop the
biomaterials possessed good combination of mechanical properties and biocompatibility
[68-71]. When bone is replaced by the biomaterial, it is imperative that implant material
should possess the similar mechanical properties as that of bone, otherwise their might be
chance of premature failure [72-73]. Since the metallic biomaterials have excellent
strength, good electrical and thermal conductivity as compared to the ceramic and
polymers, therefore they are most frequently used in medical applications particularly in
hip joint, knee joint, bone plate and dental implant [73].
The metallic materials which are most frequently used in biomedical field are
stainless steel, cobalt and titanium based alloys [74-76]. However among these materials
titanium and its alloy are most suitable materials for biomedical applications due to their
excellent mechanical strength, elastic modulus comparable to bone, good biocompatibility
and high corrosion resistance [74, 76]. Although pure titanium is good candidate for
implant, but its low tensile strength and surface hardness limits its use in medical
application. Therefore it is required to improve the surface properties of Ti before using as
implant material. Many surface modification techniques has been develop in order to
improve hardness, surface roughness and bioactivity of Ti implant [77-78]. However,
among these techniques plasma surface modification is most attractive and cost effect
technique. The unique benefit of plasma surface modification is its ability to improve the
selective surface property whereas, the bulk remains unchanged.
Ceramic coating such as Al2O3, TiO2 and HA is widely used to improve the surface
properties of implant materials [79-80]. Besides the ceramics coating recently metal-
ceramic composite coating is getting the great interest in medical field due to its excellent
corrosion resistance, high hardness and wear resistance [73]. Titanium/alumina composite
has an additional advantage that both has similar coefficient of thermal expansion, which
118
make them more suitable particularly at high temperature as compared to other
titanium/SiC composite [81-82].
In the present study titanium/aluminum oxide composite film has been prepared on Ti
using the magnetron sputtering method at different plasma excitation powers, and their
effect on mechanical properties and biocompatibility has been investigated.
4.3.2 Experimental Setup
The commercially pure (99%) Ti material is used as substrate in this present
research. The dimensions of samples are 10x10x3 mm3. Before expose to plasma samples
are mechanically grinded and polished using Stuers-Rotor-3 machine, at SiC paper grit
range (200-1200), and then ultrasonically rinsed in ethanol, acetone and distilled water (10
min) each and then dried in air before expose to plasma. Plasma deposition of aluminum-
oxide was carried out in the plasma laboratory, Government College, Lahore. Plasma
device mentioned in Fig.4.21 consists of a cylindrical shaped stainless steel vacuum
chamber having 32 cm height and diameter. Two circular electrodes (diameter and
thickness are 14 cm and 2 cm respectively) each are adjustable in a parallel configuration.
Inter electrode distance between electrodes are adjusted 3 cm. Vacuum in chamber is
achieved down to a 1x10-2 mbar by using a rotary pump. The total gas pressure of oxygen-
argon plasma is kept constant throughout deposition process. Capsule type gauge is used
to measure the vacuum inside the chamber. Pulsed DC plasma excitation power is applied
to the bottom electrode which serve as anode whereas, top electrode is grounded which
serve as cathode and samples are mounted on it whereas upper electrode serve as cathode
and grounded.
Samples are treated for different plasma excitation powers (100, 150 and 200 Watt) by
keeping pressure 2 mbar, gas concentration (10%O2+90%Ar) and 4 hours process time
constant. The treated samples are then analyzed by X-ray diffraction (XRD) with Cu Kα
(λ = 1.54060Ǻ) radiation to investigate phases, average crystallite size and stresses induced
in the deposited film. The film morphology is studied by using HITACHI S3400N scanning
electron microscopy (SEM). The surface roughness is determined by
119
Atomic Force Microscope (AFM) (Veeco CP-II). To obtain the surface hardness, indenter
with a maximum force of 100 gf is used by Wilson Wolpert 402MVD Vickers Hardness
Tester.
To investigate the biological behavior MC3T3-E1 pre-osteoblasts were cultured in a high-
glucose Dulbecco’s modified eagle medium (DMEM) added with 10% fetal bovine serum
(FBS) in 75 cm2 tissue culture flakes and were incubated in a humidified atmosphere of
5% CO2 at 37o C. 50000 cells in a volume 1 mL were seeded on each sample on 12-well
plates for cell morphology observation and proliferation assay. After culturing for 24 h, the
cells were rinsed with PBS and fixed with 4% paraformaldehyde in a humidified
atmosphere of 5% CO2 at37o C for 20 min. After fixing, the cells were stained with FITC-
Phalloidin and 40, 6-diamidino-2-phenylindole (DAPI) sequentially and then were
observed by fluorescence microscopy.
Figure 4.22: Systematic diagram of Plasma generating system
4.3.3 Results and Discussion
XRD Analysis
The xrd patterns of untreated and Al-O deposited samples of Ti at different plasma
excitation powers (100, 150, 200 watt) are shown in Fig. 4.22 by keeping the total gas
120
pressure, concentration of gasses and treatment time constant at 2 mbar, (90%Ar + 10%O2)
and 4 hour respectively. The pattern of untreated Ti sample indicates the signatures of
diffraction peaks corresponding to (100), (002), (101), (102), (110), (103), (200), (201)
plane reflections of titanium for hexagonal closed packed structure [83]. For treated
specimens we didn’t observed any new phase but significant shifting, variation in peaks
broadening and increase in intensities are the evidence for presence of Al-O into the
titanium matrix and form Ti-Al-O composite film. Different phases of TiO2 are observed
at 2θ positions of 34.9o, 38.25o,40.2o, 52.9o, 76.16o, 77.22o are assign to (101), (004), (141),
(215), (301) planes respectively (matched by ICSD card no 01-082-
1123,01-075-1537), also AlTiO2 (421) and (511) at 2θ position of 62.8o and 70.7o (ICSD
card no 00-050-0834 ) are observed.
From xrd pattern it is clear that peak intensity crossponding to TiO2 (004) plane
(preferred orientation) increases with the increase in applied plasma excitation power upto
150 watt. This increase in peak intensity is attributed to fact that at low plasma excitation
power (100 watt), the number of energetic aluminum atoms sputtered from the target
material is lower, which results in formation of thin film with low roughness value.
Whereas, with increasing plasma excitation power up to 150 watt formation of thick film
is observed with higher crystallinity and roughness. As in magnetron sputtering with
increasing plasma excitation power, sputter rate also increases which in turns increases the
deposition rate. Many researchers [84-86] reported that increase in crystallinity with the
increase in plasma excitation power. We attribute this to the increase in surface mobility
with increasing plasma excitation power, which is required to form a highly crystalline
film. In this respect, it is inferred that the high plasma excitation power in the magnetron
sputtering process energizes Ar ions and provides translational kinetic energy to the
adatoms. The surface diffusion of these adatoms is then enhanced with the momentum
transfer to the growing surface, which finally leads to a crystalline film [87].
Further increase in plasma excitation power up to 200 Watt a slight decrease in peak
intensity is observed which may explain as: at 200 watt plasma excitation power the
sputtered atoms have sufficient energy and therefore would re-sputter the already
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developed film on the substrate. This may results in the decrease in film thickness, intensity
and roughness. These results are good agreement with AFM study (in section
3.6). The reason for the reduction of peak intensity is might be due to the recrystallization
phenomenon during resolidification [88].
The average crystallite size is calculated by using scherrer’s formula [89] for TiO2
(004) plane.
Where D is the average crystallite size, λ is the wavelength of x-rays (0.9-1), FWHM is the
full width at half maximum (in radians), Ɵ and is the diffraction angle (in radians).
It is mention from graph that average crystallite sizes are found to be 89 nm, 73 nm,
and 29 nm for the plasma excitation powers of 100, 150 and 200 watt respectively in
oxygen-argon mixed plasma. It is observed that average crystallite size decreases with an
increase in plasma excitation power. This reduction in average crystallite size is attributed
to the diffusion of oxygen atoms causing the peak broadening [90] due to variation of d-
spacing. The development of residual stresses and possible defective structure of oxide
layers [91] may be the other reasons.
The residual strain in the deposited film is measured by using following relation.
Where do and d are the observed and standard plane spacing respectively taken from
ICSD (inorganic crystal structure database), used to calculate the nature of residual stresses
(compressive or tensile) present in the lattice by multiplying the strain value to the young
modulus of crossponding plane .
𝑹𝒆𝒔𝒊𝒅𝒖𝒂𝒍 𝑺𝒕𝒓𝒆𝒔𝒔 = 𝜺 𝑿 𝒌
Where k is the young modulus of the material [92].
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From graph (Fig. 4.24), residual stresses present in the film are compressive in
nature, which initially increase at plasma excitation power 150 watt but with further
increase in plasma excitation power i.e 200 watt compressive stresses are going to relax,
hence reveals an anomalous behavior. This presence of compressive stresses can be
attributed to the refinement in average crystallite size and better packing of film [93].
123
Figure 4.23: XRD patterns of untreated and Al-O deposited specimens at different
plasma excitation powers
Figure 4.24: Variation of average average crystallite size as a function of different
excitation plasma excitation powers
124
Figure 4.25: Variation in residual stress as function different excitation plasma
excitation powers
SEM Morphology
The SEM micrograph of untraded and plasma treated specimen at different plasma
excitation powers is given in the Fig. 4.25 (a-d). Fig. 4.25 (a) shows the surface of untreated
specimen which is almost plain and free of particles. However the surface of specimen
treated at 100 watt consists of fine rounded grains which are not uniformly distributed all
over the surface. Moreover, the whole surface consists of bright and dark grey color
particles. The grain size of white particles is larger comparatively the dark grey particles.
Fig. 4.25 (c) indicates the specimens treated at 150 watt, the micrograph shows the
formation of grains (rounded shape), which are uniformly distributed over the entire
surface. Moreover, relatively compact film is formed as comparatively less empty spaces
are observed at this power. The micrograph of specimen deposited at 200 watt again shows
similar grains but again surface is not properly covered. The surface morphology strongly
depends upon plasma excitation power, which affects the mechanism of diffusion of atoms
and nucleation during the film growth. As mentioned in xrd results that maximum
deposition in achieved at 150 watt plasma excitation power.
Further increase in plasma excitation power caused resputtering of deposited layer.
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Figure 4.26: SEM micrographs of untreated and treated specimens at different plasma
excitation powers (a) untreated, (b) 100 watt, (c) 150 watt (d) 200 watt.
AFM Study
The surface topography of TiAlO composite film deposited at different plasma
excitation powers were studied using the AFM and given in the Fig. 4.26. The surface of
untreated specimen is almost free of particles with roughness value 60 nm. On the other
hand surface of treated specimen at 100 watt consists of large hillocks and irregular grains,
which confirm formation of film on the substrate surface. However with an increase in
plasma excitation power size of hillocks also increased and irregular grains converted into
a regular pattern. The Ra and Rq roughness of film was also determined and given in Table.
4.5. The Rms value increased with an increase in plasma excitation power and then
decrease at the maximum power. The line analysis of treated and untreated specimens was
also given in the Fig. 4.27 which represents the variation in the grain size.
The surface roughness strictly depends on plasma excitation power. The
improvement in surface roughness may be due to bombardment of high energy particle,
which can lead to substantial surface heating and local fusion [94]. It is also observed that
samples treated at higher plasma excitation power are greatly damaged by the plasma which
may be due to high dose of oxygen causing surface swelling of the substrate which also
increases surface roughness [95, 96]. However the slight decrease in the roughness at 200
watt may be due to the formation of smooth film on substrate surface, which also promote
the uniform grain pattern. The increase in the surface roughness is an important factor
which enhances the biocompatibility of film. It is mentioned by many researchers that a
rough surface is more beneficial to enhance the biocompatibility of implant material as
compared to the uniform surface [97]. As with increase in the surface roughness, larger
number of sites is available for cell attachment, resulting an increase in surface area of
implant. Moreover an implant material having higher surface roughness has greater
biomechanical interaction with bone. Surface roughness also enhances the surface energy.
With an increase in the surface energy the more protein are absorbed by matrix, which
ultimate enhanced the cell proliferation [98].
126
Figure 4.27: AFM micrographs of deposited specimens at different plasma
excitation powers (a) untreated, (b) 100 watt, (c) 150 watt (d) 200 watt
127
Figure 4.28: Line analysis of deposited specimens at different plasma excitation
powers. (a) untreated (b) 100 watt, (c) 150 watt, (d) 200 watt
Table 4.5: Roughness values of untreated and deposited films
128
Plasma excitation
plasma excitation
power (watt)
Root mean square
roughness Rms
(nm)
Average roughness
Ra (nm)
Untreated
58 50
100
101.51 80.7
150
142.2 110.05
200
112 89.98
Tensile Results
The comparison of stress-strain (S-S) curves of untreated and treated specimens at
the different plasma excitation powers are given in the Fig. 4.28. It is clear from the figure
that the yield stresses (YS), ultimate tensile stress (UTS) and percentage elongation of
treated specimens is improved after the deposition of composite film. The variation in YS
and UTS with increase in plasma excitation power is given in the Fig. 4.29. The increase
in YS and UTS in treated specimens is attributed to presence of interstitial element oxygen.
Since YS and UTS of Ti alloy depends upon the interstitial element such as oxygen [99].
The radius of oxygen is smaller as compared to the titanium, so oxygen atoms diffuse into
titanium matrix and produces the lattice strain, which impede dislocation motion, resulting
in strengthening of titanium matrix. Bin et al. [100] also reported that oxygen formed the
solid solution in the pure Ti matrix which increases the tensile strength of Ti. In another
study Oh et al. [101] investigated oxygen effect on mechanical properties and lattice strain
of Ti and Ti6Al4V alloy. They found that due to oxygen solid solution hardening c/a ratio
increases which increase the strength of Ti and Ti6Al4V alloy. However, the oxygen effect
on mechanical properties of Ti is more significant as compared to Ti6Al4V. The results of
S-S curves also agree with the hardness results. The change in the percentage elongation of
Ti with increase in plasma excitation is given in the Fig. 4.30. The results show that the
percentage elongation also increases with an increase in the plasma excitation power.
129
Figure 4.29: Comparison of stress-strain curve of untreated and treated specimens at
different plasma excitation powers, (1) untreated (b) 100 watt (c) 150 watt (d) 200 watt
130
Figure 4.30: Variation in the YS and UTS at different plasma excitation
plasma
Figure 4.31: Variation in percentage elongation at different plasma excitation powers
131
Fracture Analysis
The fracture surface of untreated and treated specimens after the deformation have been
given in the Fig. 4.31 (a-d). It is clear from the figure that fracture surface of treated
specimens consist of composite fracture that is plate like and voids as compared to the
untreated specimen. The plate like fracture is dominant in area near to surface which has
become hard due to formation of composite film, whereas bulk of material consist of voids
which indicate the ductile fracture. However, number of voids slightly increases and their
size also becomes large in the all treated specimens. The fractrographs are well correlated
with stress strain curves which also indicate that with increase in the plasma excitation
power the surface of specimens become hard without decrease in percentage elongation.
Figure 4.32: SEM Fractrographs of specimens (a) untreated (b) 100 watt (c) 150
watt (d) 200 watt
132
Hardness Results
A micro-hardness tester (Wilson Wolpert 402 MVD. China made) is used for surface
hardness (HV) measurements. The load 10 g is applied for a dwell time of 10 sec for all
indentations. Hardness values are obtained from the average of five measurements on each
sample.
Variation of surface hardness at different plasma excitation powers (100, 150, 200 watt) is
shown in Fig. 4.32. Microhardness increases with increasing plasma excitation power as
compared to untreated specimen. This increased in surface hardness is due to the formation
of TiO2 and Al-Ti-O composite films and increased oxygen content in substrate matrix
[102]. Maximum hardness (319 HV) is achieved at maximum plasma excitation power of
200 W, which mainly attributed to the crystallite refinement and incorporation of oxygen.
This is supported by the Hall-Petch, which states that harder films are achieved when films
having smaller average crystallite size, hardness are inversely related to the square root of
grain diameter and development of compressive residual stresses. It is generally observed
that compressive residual stresses and formation of modified layers account for high
hardness [103-106].
Figure 4.33: Variation of surface microhardness as a function of different
plasma excitation powers (100, 150, 200 watt)
133
Cell Culture Study
To investigate the biological acceptability, mouse MC3T3-E1 pre-osteoblasts are
cultivated on deposited specimens. Fig. 4.33 shows the images of the cells after 3 days
taken by fluorescent microscopy. It can be seen from micrographs that cell culture on
treated specimen at plasma excitation power 100 watt is not properly spread. However,
when we increase the plasma excitation power up to 200 watt, more cells are attached and
proliferate and show no obvious cytotoxic effects and have higher growth activity. The
rectangular and elongated morphologies of the cells in Figs. 4.33(b-c) indicates that cells
growth occupy a larger area compared to the film deposited at 100 watt as shown in Fig.
4.33(a). The improvement in the biocompatibility with the increase in plasma excitation
power may be due to increase in surface roughness. Since with the increase in roughness
the available surface area also increases which ultimately improve the cell attachment with
surface [97].
134
Figure 4.34: Fluorescence microscopy images after culturing for 3 days (a) 100 watt
(b) 150 watt and (c) 200 watt
135
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Chapter 5 Conclusions and Future work 5.1 CONCLUSIONS
This chapter consists of the conclusions obtained from the present research work.
To make a good combination of mechanical properties along with the biocompatibility,
the surface of Ti and its alloys are successfully modified by deposition of titanium
142
oxynitride, nanostructure zirconium oxide (ZrO2) and Ti-Al-O composite films using the
plasma methods. The conclusion of present investigation is given below.
In the first experiment, titanium oxynitride films are deposited on NiTi by DC reactive
magnetron sputtering using different flow rate ratios of nitrogen to oxygen and the effects
of the flow rates on the surface and biological properties are investigated. The analysis of
XRD and XPS confirm the presence of titanium oxynitride film on the NiTi substrate. The
surface morphology results indicate that after titanium oxynitride deposition the roughness
of NiTi surface increases, which promotes the cell adhesion proliferation. EDS show
insignificant amount of calcium and phosphorus, but the oxynitride film acts as a barrier
against out-diffusion of nickel. Cell cultures indicate that the NiTi with surface titanium
oxynitride fosters proliferation and shows no obvious cytotoxicity. The improvement in
biocompatibility is attributed to the presence of rutile and anatase phase of TiO2.
In second experiment, nanostructure zirconium oxide (ZrO2) film has successfully
implanted on Ti-6Al-4V substrate by using PIII&D system at various voltages 15, 20 and
25 kV. AFM results confirm the formation of nanostructured film. The maximum value of
average roughness is observed at 25 kV i.e 7.02 nm. XPS depth profile results confirm
formation of hard zirconium oxide film at 25 kV because sharp peaks crossponding to Zr3d
are observed even after the Ar sputtering of surface for 8 min. Friction coefficient and wear
rate in all treated specimens was found to be decrease with increasing voltages, on other
hand hardness values increase with the increase in voltages. Maximum hardness value is
about 10.54 GPa obtained at maximum voltage (25 kV), which is almost two times than
that of untreated specimen. The improvement in the wear resistance and hardness is
attributed to the presence of nanostructure ZrO2 film on the surface. Moreover we
concluded that to improve tribological properties of T6Al4V alloy 25 kV is the best voltage,
as a hard layer is formed at this voltage as compared to other voltages, which is also
supported by XPS results.
In third experiment, Ti-Al-O composite films are effectively deposited on Titanium
substrate by using pulsed DC magnetron sputtering system. The X-ray diffraction technique
confirms the formation of composite film. Average crystallite size and residual stresses of
143
Ti2O (004) plane depend on plasma excitation power. Micrographs of scanning electron
microscopy (SEM) shows the formation of granular grains on the surface that varies with
the plasma excitation power. Atomic force microscopic analysis reveals that variation in
film roughness is directly related with process parameters. Results obtained from Vicker's
hardness testing machine shows that hardness of composite film increases with an increase
in the plasma excitation power and the maximum value of hardness at 200 watt is 316 HV.
Tensile test shows that YS and UTS of Ti are increased as result of formation of composite
film without any decrease in the percentage elongation. Fluorescence microscopy images
after culturing cells for 3 days revealed that wide spread coverage over all area on substrate
and good cell adhesion is observed at 150 watt plasma excitation power also no obvious
cytotoxicity is observed. The improvement in mechanical properties and biocompatibility
in all treated specimens is attributed to the formation of Ti-Al-O composite film.
5.2 FUTURE SUGGESTIONS
In this current study the described work has been concerned with the formation of
biocompatible thin films using different plasma deposition techniques on titanium and its
alloys. Overall the aim of this research to achieve required goals was successful, our
experimental results confirm that mechanical and biocompatibility of substrate material
enhanced by means of surface modifications of biomaterials but still it is far away from
biomedical field. In the future, some characterization techniques like electrochemical
impedance spectroscopy, wettability of surface, and more in vivo and in vitro studies must
be carried out.
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