A Dissertation
entitled
Microwave Assisted Synthesis of Alkaline Earth Phosphate Coating and its Applications
for Biomedical Implants
by
Yufu Ren
Submitted to the Graduate Faculty as partial fulfillment of the requirements for the
Doctor of Philosophy Degree in
Engineering
______Dr. Sarit B. Bhaduri, Committee Chair
______Dr. Vijay K. Goel, Committee Member
______Dr. Arunan Nadarajah, Committee Member
______Dr. Ahalapitiya Jayatissa, Committee Member
______Dr. Matthew Franchetti, Committee Member
______Dr. Amanda Bryant-Friedrich, Dean College of Graduate Studies
The University of Toledo
December 2017
Copyright 2017, Yufu Ren
This document is copyrighted material. Under copyright law, no parts of this document may be reproduced without the expressed permission of the author. An Abstract of
Microwave Assisted Synthesis of Alkaline Earth Phosphate Coating and its Applications for Biomedical Implants
by
Yufu Ren
Submitted to the Graduate Faculty as partial fulfillment of the requirements for the Doctor of Philosophy Degree in Engineering
The University of Toledo
December 2017
Bioimplant is a group of medical devices that aim to restore/replace the function of the defected/diseased tissue. As the potential candidate for orthopedic applications, the implant material needs to maintain the suitable mechanical properties and desirable surface chemistry to ensure long-term mechanical stability and foster the regeneration of host tissue at the defect site. Due to the advancement in the research of biomaterials, a wide range of materials including metals, ceramics, polymers, and composites now can be used in bone grafting procedure for different purposes. However, none of the currently available bioimplant materials have met all the requirements and expectations. For instance, most of the metallic implant materials with high strength present poor osseointegration properties.
Hence, there have been great efforts in developing the alkaline earth phosphate based bioactive coatings on implant materials to promote bone mineralization. This study covers diverse types of surface modifications of metallic implant material – magnesium alloy and polymeric implant material – polyetheretherketone (PEEK) with various alkaline earth phosphate coatings, nanostructuring, composite formation and surface pretreatments.
iii
In the first part of this thesis, a microwave assisted coating technique was developed to improve in vitro degradation behavior and biological properties of Mg alloys. The microwave irradiation dramatically accelerated coating deposition kinetics and notably shortened the coating process to minutes rather than hours/days consumed in the conventional biomimetic coating method. Moreover, the as-deposited calcium deficient hydroxyapatite (CDHA) and magnesium phosphate (MgP) layers presented outstanding corrosion resistance and bioactivity in physiological environment, which evidently enhanced the biological responses of Mg alloys.
Further on, organic-inorganic composite coatings were synthesized by combining the microwave assisted coating technique with spin coating process and the synergistic benefits of as-deposited composite coatings were explored. It was found the alkaline earth phosphate nanoparticles including FHA and AMP, either serving as the interlayer between the polymer film and Mg alloy substrate or incorporating in polymer matrix, were beneficial to enhance biomineralization capability of Mg alloys and subsequently facilitate the osseointegration process. Moreover, organic layer-biodegradable PLA film presented excellent interlocking with alkaline earth phosphate phases, which leads to the formation of favorably compact microstructures of composite coatings. Consequently, as-deposited
FHA/PLA and nAMP/PLA composite coatings were demonstrated to be efficacious in blocking the infiltration of the corrosive medium.
In the end, an approach incorporating surface activation pretreatment and amorphous alkaline earth phosphate coating deposition was developed to overcome the inherent biological inertness of polyetheretherketone (PEEK). The pretreatment involves the acid or alkali etching could generate the negative charge on the surface of PEEK, which
iv is fundamental to induce rapid deposition of amorphous magnesium phosphate (AMP) coating via the microwave assisted coating technique. The 3-D porous structure resulted by sulfonation in combination with bioactive AMP layer showed great ability to promote osteoblast attachment. More importantly, a more vigorous biomineralization process with greatly intensified bone-like apatite precipitation was observed on AMP coated PEEK samples in vitro, which reveals the tremendous potential of using amorphous alkaline earth phosphate coatings to improve the bioactivities of biomedical implants.
v
Acknowledgements
I would like to express my sincere gratitude to my advisor Dr. Sarit B. Bhaduri for his continuous support and motivation in past five years. Without his insightful guidance, this work would not have been completed. Also, I am very thankful for the valuable inputs from other members on my dissertation committee, Dr. Vijay Goel, Dr. Arunan Nadarajah,
Dr. Ahalapitiya Jayatissa and Dr. Matthew Franchetti.
Furthermore, I would like to acknowledge Dr. Huan Zhou for his great assistance in the beginning of my research. I would like to extend my thanks to Dr. Joseph Lawrence,
Dr. Lidia Rodriguez, Dr. Sam Imanieh, Dr. Boren Lin, Tammy Phares and John Jaegly for their tremendous help and worthy suggestions.
I would like to thank my fellow lab mates and friends at UT for the useful discussions, cooperation and encouraging during my graduate studies. Moreover, the financial support from the National Science Foundation (Grant no. 1312211) and College of Engineering are greatly appreciated.
My special thanks go to my parents and Yue, because of their persistent support throughout the entire journey of my graduate studies, from the day I applied to graduate school to the day of my graduation. Their love shaped me into who I am today.
vi
Table of Contents
Abstract ...... iii
Acknowledgements ...... v
Table of Contents ...... vi
List of Tables ...... xi
List of Figures ...... xii
List of Abbreviations ...... xvii
1 Introduction…...... 1
1.1 Overview...... 1
2 Microwave Processing of Biomaterials: A Review ...... 5
2.1 Abstract...... 5
2.2 Interaction of microwave and materials...... 6
2.3 Microwave assisted synthesis of CaPs and MgPs…...... 11
2.3.1 Hydroxyapatite and calcium deficient hydroxyapatite ...... 11
2.3.2 Ion doped hydroxyapatite ...... 24
2.3.3 Biphasic calcium phosphate ...... 33
2.3.4 Amorphous calcium phosphate ...... 34
2.3.5 Magnesium phosphate compound...... 39
2.4 Microwave assisted surface modification of biomaterials…...... 40
vii
3 Rapid Coating of AZ31 Magnesium Alloy with Calcium Deficient Hydroxyapatite
Using Microwave Energy ...... 43
3.1 Abstract...... 43
3.2 Introduction...... 44
3.3 Materials and methods...... 47
3.3.1 Material preparation ...... 47
3.3.2 Coating preparation ...... 47
3.3.3 Coating characterization ...... 48
3.3.4 Electrochemical test ...... 49
3.3.5 Immersion test ...... 50
3.3.6 Cytotoxicity test ...... 50
3.3.7 Statistical analysis ...... 51
3.4 Results…...... 51
3.4.1 Coating characterization ...... 51
3.4.2 Electrochemical behavior ...... 55
3.4.3 Degradation behavior ...... 56
3.4.4 Cytotoxicity...... 58
3.5 Discussion...... 60
3.6 Conclusion ...... 65
4 Microwave Assisted Magnesium Phosphate Coating on AZ31 Magnesium
Alloy……………………………………………………………………………...... 67
4.1 Abstract...... 67
4.2 Introduction...... 68
viii
4.3 Experimental...... 70
4.3.1 Material preparation ...... 70
4.3.2 Coating preparation ...... 70
4.3.3 Coating characterization ...... 71
4.3.4 Electrochemical test ...... 71
4.3.5 Immersion test ...... 72
4.3.6 Cytotoxicity test ...... 72
4.3.7 Statistical analysis ...... 73
4.4 Results and discussion…...... 74
4.4.1 Coating characterization ...... 74
4.4.2 In vitro degradation behavior ...... 79
4.4.2.1 Electrochemical test ...... 79
4.4.2.2 Immersion test ...... 81
4.4.3 Cytotoxicity...... 87
4.5 Conclusion ...... 88
5 Synthesis and In Vitro Evaluation of a Bilayer Coating Combining Poly (Lactic
Acid) and Fluorinated Hydroxyapatite on AZ31 Magnesium Alloy ...... 90
5.1 Abstract...... 90
5.2 Introduction...... 91
5.3 Experimental...... 94
5.3.1 Material preparation ...... 94
5.3.2 FHA deposition ...... 95
5.3.3 Spin coating ...... 95
ix
5.3.4 Coating characterization ...... 96
5.3.5 Electrochemical test ...... 97
5.3.6 In vitro degradation ...... 97
5.4 Results and discussion…...... 98
5.4.1 Coating characterization ...... 98
5.4.2 Electrochemical behavior ...... 102
5.4.3 In vitro degradation behavior ...... 104
5.5 Conclusion ...... 108
6 Nanostructured Amorphous Magnesium Phosphate / Poly (Lactic Acid) Composite
Coating for Enhanced Corrosion Resistance and Bioactivity of AZ31 Magnesium
Alloy……………………………………………………………………………...... 110
6.1 Abstract...... 110
6.2 Introduction...... 111
6.3 Experimental...... 113
6.3.1 Material preparation ...... 113
6.3.2 Coating preparation ...... 113
6.3.3 Coating characterization ...... 114
6.3.4 Electrochemical test ...... 115
6.3.5 Immersion test ...... 116
6.4 Results and discussion…...... 116
6.4.1 Characterization of nAMP powder ...... 116
6.4.2 Characterization of PLA and nAMP/PLA coatings ...... 118
6.4.3 In vitro degradation behavior ...... 118
x
6.4.3.1 Electrochemical test ...... 121
6.4.3.2 Immersion test ...... 127
6.5 Conclusion ...... 130
7 Microwave Assisted Coating of Bioactive Amorphous Magnesium Phosphate
(AMP) on Polyetheretherketone (PEEK)...... 131
7.1 Abstract...... 131
7.2 Introduction...... 132
7.3 Experimental...... 134
7.3.1 Material preparation ...... 134
7.3.2 Coating preparation ...... 135
7.3.3 Characterization and bioactivity evaluation ...... 135
7.3.4 Cell viability assay ...... 136
7.3.5 Statistical analysis ...... 137
7.4 Results and discussion…...... 138
7.4.1 Characterization ...... 138
7.4.2 Deposition mechanism ...... 142
7.4.3 In vitro properties ...... 143
7.4.4 Cytocompatibility and RT-PCR ...... 145
7.5 Conclusion ...... 147
8 Conclusion and Future Directions ...... 148
8.1 Conclusion...... 148
8.2 Future directions...... 149
References ...... 151
xi
List of Tables
2.1 Comparison of conventional thermal processing and microwave heating ...... 9
2.2 Methods, parameters and results of MW assisted synthesis of HA ...... 13
3.1 Compositions of coating solutions ...... 48
3.2 Ion concentrations of the SBF and human blood plasma ...... 49
3.3 Electrochemical parameters of CDHA coated AZ31 samples obtained from polarization curves ...... 56
4.1 Electrochemical parameters of newberyite and TMP coated AZ31 Mg alloy samples obtained from polarization curves...... 81
5.1 Electrochemical parameters of FHA/PLA coated AZ31 Mg alloy calculated from polarization curves ...... 104
6.1 Electrochemical parameters of nAMP/PLA coated AZ31 alloy calculated from polarization curves ...... 123
6.2 Fitting results of EIS spectra of bare AZ31, PLA and nAMP/PLA coated AZ31 Mg alloy samples ...... 126
xii
List of Figures
2-1 Schematic illustration of electromagnetic spectrum near MW frequencies ...... 7
2-2 SEM and TEM micrographs with diameter distribution of HAP nanowires prepared using microwave-assisted hydrothermal method ...... 20
2-3 SEM image showing HAp nanostructures and TEM image of a typical bundle of
HA. Inset: SAED pattern taken from an individual HAp nanorod ...... 22
2-4 SEM micrographs of (a, b) HA nanowhiskers synthesized using microwave assisted combustion process ...... 22
2-5 TEM micrographs of the samples with various Na3Cit content prepared by microwave-assisted hydrothermal method ...... 24
2-6 The TEM morphology of the HA with different addition of Ag ...... 26
2-7 Antibacterial activity of HA with different addition of Ag ...... 26
2-8 The TEM graphs of the Eu doped HA nanoparticles with corresponding luminescence of suspensions of Eu-HA when irradiated with UV light ...... 29
2-9 SEM images, particle size distribution and typical luminescent image of Eu doped
HA nanowhiskers synthesized using MASCS method ...... 30
2-10 In vivo PL imaging of the mice after subcutaneous injection with Eu3+/Gd3+–HAp nanorods ...... 30
2-11 Magnetization curves for pristine and Co2+ doped samples in an applied magnetic field of ±2 tesla ...... 32 xiii
2-12 Hexagonal unit cell of a hydroxyapatite crystal with possible carbonate substitutions ...... 33
2-13 SEM micrographs of TCP and biphasic HA-TCP samples prepared via microwave assisted combustion synthesis ...... 34
2-14 TEM images of ACP nanospheres and related clusters after transition…………..36
2-15 Schematic diagram of the ACP formation process via microwave assisted hydrothermal route and the functions of soybean lecithin as the template ...... 37
2-16 TEM micrographs, SAED patter and EDS spectrum of Eu-ACP mesoporous microspheres prepared the microwave-assisted solvothermal method ...... 38
2-17 TEM micrographs single AMP nanosphere and fused AMP nanospheres prepared by the microwave-assisted method ...... 40
2-18 Representive AFM images of apatite coating with various strontium concentration… ...... 41
3-1 Schematic diagram of set up of microwave assisted coating process...... 48
3-2 XRD patterns of bare AZ31, CS1 coated AZ31 and CS2 coated AZ31...... 52
3-3 FTIR spectra of CS1 coated AZ31 and CS2 coated AZ31 ...... 53
3-4 SEM images and respective EDS analysis of CDHA coated samples ...... 54
3-5 Cross-sectional morphology of CDHA coatings ...... 55
3-6 Potentiodynamic polarization curves of un-treated and CDHA coated AZ31 alloy samples tested in SBF ...... 56
3-7 Variance of weight loss of coated sample and non-coated sample with immersion time in t-SBF solution ...... 57
xiv
3-8 SEM images and respective EDS analysis of CDHA coated samples after 7 days’ immersion in SBF solution ...... 58
3-9 O.D. values of MC3T3-E1 cells seeded in extracts of coated samples and culture medium for 1, 3 and 5 days ...... 59
3-10 Optical micrographs of osteoblast cells (MC3T3-E1) after 3 days and 5 days incubation ...... 59
3-11 Schematic diagrams of microwave assisted formation mechanism of CDHA coatings on AZ31 alloy ...... 65
4-1 XRD patterns of bare AZ31 and MgP coated AZ31 with various parameters ...... 75
4-2 FTIR spectra of bare AZ31 and MgP coated AZ31 with various parameters ...... 76
4-3 SEM images and EDS spectra of as-deposited MgP coating prepared with different temperature and time...... 78
4-4 Cross-sectional SEM images of MgP coated samples ...... 79
4-5 Potentiodynamic polarization curves of the bare and treated AZ31 samples ...... 80
4-6 Variations of weight loss of samples and pH value of SBF solution as a function of immersion time ...... 83
4-7 SEM images of MgP coated samples after two weeks SBF immersion ...... 85
4-8 XRD patterns of bare AZ31 and MgP coated AZ31 after incubating in SBF solution for 15 days...... 87
4-9 Cell viability of MC3T3-E1 pre-osteoblasts cultured in medium extracts of different MgP coated AZ31 samples for 3 days ...... 88
5-1 Schematic diagram of the setup of spin coated and the PLA film deposition of spin coating...... 96
xv
5-2 XRD patterns of the AZ31 substrate, the FHA coated, and the PLA/FHA coated samples ...... 99
5-3 FTIR spectra of the PLA coating, the FHA coating, and the PLA/FHA composite coating...... 99
5-4 SEM images of as-deposited FHA coating, PLA/FHA coating and EDS spectrum of PLA/FHA coating ...... 101
5-5 Cross-sectional SEM micrograph of as-deposited PLA/FHA coating on Mg alloy
AZ31sample ...... 101
5-6 Potentiodynamic polarization curves of uncoated, FHA coated, PLA coated and
PLA/FHA coated AZ31 Mg alloy samples in SBF solution...... 103
5-7 Weight loss of surface treated samples and control sample and variations of pH values of SBF solution at various immersion periods in SBF ...... 107
5-8 Surface morphology and corresponding EDS spectra of FHA and PLA/FHA coated
AZ31 Mg alloy samples after 15 days immersion in SBF ...... 108
6-1 Photographs of PLA solution and nAMP/PLA suspension ...... 114
6-2 XRD pattern, TEM image and SAED pattern of nAMP powder ...... 117
6-3 XRD patterns of bare AZ31, PLA and PLA-AMP coated AZ31 magnesium alloy and Infrared spectra of nAMP, PLA and PLA-AMP coated AZ31 magnesium alloy ....119
6-4 Surface SEM images of Bare AZ31, PLA coated AZ31, nAMP/PLA coated AZ31 and respective EDS spectrum ...... 120
6-5 Cross-section SEM images of PLA coated AZ31 and nAMP/PLA coated AZ31 samples ...... 121
6-6 Polarization curves of bare AZ31, PLA and nAMP/PLA coated sample ...... 122
xvi
6-7 Nyqist plots of uncoated and coated samples in SBF solution and equivalent circuit models applied for fitting EIS spectra of bare AZ31, PLA and nAMP/PLA coated AZ31
Mg alloy samples ...... 125
6-8 Weight loss of coated samples and control sample and variations of pH values of
SBF solution at various immersion periods in SBF ...... 129
6-9 Surface morphology and corresponding EDS spectra of PLA and nAMP/PLA coated samples after 15 days immersion in SBF ...... 129
7-1 Surface morphology of NaOH, and sulfuric acid treated PEEK ...... 139
7-2 Water contact angle of bare PEEK and treated PEEK samples ...... 139
7-3 XRD patterns and FTIR spectra of bare and treated PEEK samples ...... 140
7-4 SEM images and EDS spectra of AMP coated PEEK samples ...... 141
7-5 Cross-sectional SEM images of AMP coated PEEK and TEM image associated with SAED of AMP coatings ...... 142
7-6 SEM images of bare and AMP coated PEEK samples after 1 week SBF immersion and EDS spectra of apatite formed on PEEK surface during SBF immersion ...... 144
7-7 Cell viability of MC3T3-E1 pre-osteoblasts cultured on the PEEK control, PEEK-
OH-AMP and PEEK-S-AMP for 3 days ...... 146
7-8 Osteogenic differentiation by measuring the mRNA expression level of alkaline phosphatase (ALP) and osteocalcin (OCN) after 7 days ...... 147
xvii
List of Abbreviations
α-MEM...... Alpha minimum essential medium ACP...... Amorphous calcium phosphate ALP...... Alkaline phosphatase AMP...... Amorphous magnesium phosphate ANOVA...... Analysis of variance
BCP...... Biphasic calcium phosphate
CaP...... Calcium phosphate CDHA...... Calcium deficient hydroxyapatite
EDS...... Energy dispersive X-ray spectroscopy EIS...... Electrochemical impedance spectroscopy
FBS ...... Fetal bovine serum FHA...... Fluorinated hyrxoyapatite FIB...... Focused ion beam FTIR...... Fourier transform infrared spectroscopy
HA...... Hydroxyapatite
MASCS...... Microwave assisted solution combustion synthesis MgP...... Magnesium phosphate MW...... Microwave
OCN...... Osteocalcin
PCR...... Polymerase chain reaction PDF...... Powder diffraction file PLA...... Poly lactic acid PEEK...... Polyether ether ketone
SAED...... Selected area electron diffraction SBF ...... Simulated body fluids SEM...... Scanning electron microscope
TCP...... Tricalcium phosphate xviii
TEM ...... Transmission electron microscope TMP...... Trimagnesium phosphate Tris...... Tris-hydroxymethylaminomethane
XRD...... X-Ray diffraction
xix
Chapter 1
Introduction
1.1 Overview
Natural bone is a complex and highly dynamic tissue, which undergo the biological remodeling comprising of mature bone resorption and new bone formation throughout the lifetime. Owing to this unique property, cortical bone can repair small defects by itself.
However, when complicated pathological fractures occur or the defects exceed the self- healable size, the bone grafting process is required to support the self-healing of defected sites. These kinds of defects are usually caused by the orthopedic trauma, infection and chronic diseases. According to the statistics, more than 0.5 million and 2 million of bone grafting procedures were performed yearly in the US and around the world respectively, and a steady annual growth is projected for the coming years [1]. Moreover, with the increase of life expectancy and aging of the population, the global demand and market size for novel orthopedic implants that can promote the new bone formation and shorten the healing process are dramatically expanding.
Orthopedic implant is a group of biomedical devices to be employed in human body for diverse applications, such as temporary fixation of fractures and permanent replacement
1
of defected or diseased bone tissues and joints. To date, orthopedic implants have been designed and produced using various biomaterials ranging from metals to polymers.
Metallic materials such as stainless steel, cobalt–chrome alloy and titanium alloys are most commonly used implant materials, due to their desirable mechanical properties, good corrosion resistance, and biocompatibility. The major drawback of these materials for orthopedic applications is their relatively higher elastic moduli than those of cortical bones.
As these implant materials are much stiffer than natural bone, which can result in stress shielding phenomenon leading to resorption of surrounding bone tissue and prosthetic loosening. Moreover, the implants manufactured using these materials generally serve as permanent prostheses or temporary fixation devices in human body, which can cause significant risks of infection in long-term application or secondary removal/revision operations. Thus, magnesium alloys have been developed as the next generation orthopedic implant materials, since the excellent match between of Mg alloys and natural bones in mechanical properties can greatly eliminate the stress shielding effects and their biodegradability can avoid further external interventions after the implantation. Polymers such as poly (L-lactic acid), polymethyl methacrylate (PMMA) and polyglycolide
(PGA) have been extensively used in biomedical applications, owing to their versatility and bioresorbability. However, most of the biopolymers are not suitable for bone fixation or replacement, as they lack mechanical stability and cannot provide demanding mechanical support to fractured tissue. One exception here is polyetheretherketone
(PEEK), a semi-crystalline thermoplastic with high chemical stability, good biocompatibility and a comparable elastic modulus (8.3 GPa) to that of cortical bone (17.7
GPa). Since the late 1990s, PEEK has been used as orthopedic implant biomaterial in spinal 2
surgery, mostly in the form of intervertebral cages. Further, its radiolucency, accountability for MRI imaging make it more appealing as the candidate for replacing metallic implant counterparts.
Despite recent advancement in the research of biomedical implant materials has offered a great deal of bulk implant material selections for all subspecialties in orthopedic surgery, the long-term success of implants remains to be challenging. Due to the aseptic loosening, infection, instability and other undesired activities, the high tendency of premature failure is observed in currently commercially available implants. For example,
10% of implants used for knee and hip replacements failed within first 10-20 years and necessitated the further revision procedure [2]. Thus, to optimize the clinical outcome, improve the stability and osseointegration of orthopedic implant in host tissue and minimize the adverse events, surface modifications, especially bioactive alkaline earth phosphates have been developed on prostheses. The motivation of applying alkaline earth phosphate based coatings such as CaP, MgP and their composites, is alkaline earth phosphate materials exhibit a great similarity to natural bone minerals and outstanding osteoconductive properties. As a result, a robust osseointegration with fast bony in-growth is expected to take place at the interface of bone-implant [3]. In addition, the expeditious osseointegration can remarkably enhance the stability of implants and mitigate the detrimental events associated with implants such as micromotion, microbial attachment, and colonization. In general, the advantages of alkaline earth phosphate coatings are readily evident: improving the corrosion resistance of metallic implant materials in physiological environment, facilitating earlier osseointegration and pain relief, promoting new bone formation, shortening healing process and delivering favorable patient outcome. 3
In this work, a microwave assisted coating technique was developed to perform the rapid surface modifications on biodegradable and non-degradable biomedical implant materials including AZ31 Mg alloy and PEEK. The microwave assisted coating technique combining the convenience of microwave chemistry and wet chemical precipitation, consumes much less time (minutes) than the time costed in conventional biomimetic coating process (hours/days). Moreover, various kinds of alkaline earth phosphate coatings with amorphous, crystalline, biphasic and composite structure were prepared and investigated to tailor the in vitro behaviors of implant materials.
4
Chapter 2
Microwave Processing of Biomaterials: A Review
2.1 Abstract
Biomaterials utilize a wide range of materials including ceramics, metals, polymers and biological substances to replace or restore the functions of various parts of human body.
As inspired by the numerous success of microwave chemistry, the use of microwave technology in biomaterials processing, especially in rapid synthesis of bioceramics has been steadily growing in last decade. This chapter initially outlines the interactions between microwave and materials and also draws out the principle of rate acceleration by microwave processing. Further, a meticulous attempt is made to summarize the development in microwave processing of biomaterials with the special attention paid to the microwave assisted synthesis of hydroxyapatite nanoparticles. Last, microwave assisted rapid surface modification of biomedical implants is discussed, which could help us explore novel applications of microwave technology in the field of biomaterials research.
5
2.2 Interaction of microwave and materials
In last 30 years, materials researchers have discovered and developed many advanced and functional materials which can be extensively employed in various sectors such as automobile, aerospace, nuclear and biotechnology. Thermal processing plays an important role in the microstructure, mechanical property and performance of developed materials and is widely considered as one of the most important steps in materials synthesis.
However, in conventional thermal treatment, heat energy can only be conveyed to the bulk material via convection, conduction and radiation, which can cause significant energy loss and time consumption in this process. Thus, along with the progress in the development of novel materials, researchers also recognized the importance and significance for the further exploration of the breakthrough technique which could deliver rapid, efficient, energy saving and environment-friendly thermal processing compared to the conventional heating techniques. As a consequence, a novel heating technique employing electromagnetic waves was developed to curb out the abovementioned limitations. Today, microwave heating technology has emerged as a popular and promising heat source with widespread applications in the research of advanced materials.
Microwaves (MWs) comprise of the electromagnetic waves with frequency ranging from 0.3 to 300 GHz and with the wavelength of 1 mm -1 m, which lie between infrared and radio frequency waves in electromagnetic spectrum (Fig. 2-1). To avoid the interference with the microwave radar and telecommunication devices, the frequency of
MWs for industrial and domestic uses is regulated between 2-8 GHz. The microwave ovens which are commonly used in homes and laboratories operate at 2.45 GHz. The microwave generators include magnetrons, klystrons, power grid tubes, traveling wave tubes and 6
gyrotrons. Among them, magnetrons are recognized to be more reliable, cost-effective and efficient. Microwaves are coherent, polarized and with the ability to be transmitted, absorbed, or reflected depending on the type of the material. Due to its rapid heating, shorter reaction time, higher yield, lower energy consumption, cleaner products and enhanced quality of synthesized materials, microwave assisted synthesis has received increasing attention and steadily become a promising synthetic route in green chemistry
[4]. Despite the many unique benefits of MW assisted synthesis have been widely recognized, application of MW technology in processing of biomaterials is still comparatively limited. However, the research interests in rapid preparation of inorganic, organic and composite biomaterials have been continuously thriving in the past decade. It is also anticipated the research activities in related fields will continue to boost in coming years.
Figure. 2-1 Schematic illustration of electromagnetic spectrum near microwave
frequencies.
7
In conventional heating or thermal processing, the material surfaces or walls of the reactor were firstly heated through radiation and convection. Then, heat was conducted from the hot surface and reactor to the cold core of materials and reactants respectively.
Thus, such heating pathway could cost longer time to achieve homogeneous heat distribution in interior of materials and result in unexpected thermal gradients inside processed materials. Conversely, the microwave energy could heat target materials at molecular level without heating the furnace and reactor. Table 2.1 summarized the differences between conventional thermal processing and microwave heating.
It is fundamental to understand the principles of interaction between microwave and materials, and the mechanism of microwave heating. To date, two major mechanisms had been proposed and widely accepted to explain the principle of microwave dielectric heating: 1) polar molecules such as water molecules in solvent are exposed to external electric field generated by microwave irradiation and try to reorient with rapid changing electric field. During this process, molecular interactions such as friction and collision can increase molecular kinetic energy and then convert the electromagnetic energy to volumetric heat throughout target materials; 2) the conduction mechanism can occur in metallic materials, semiconductors and ionic solutions. Hence, the free electrons or ions present in materials or solutions can start long range transportation in the orientation of electric field and induce large electric current under microwave irradiation. Then, the dielectric energy can be lost due to frictional, elastic, intermolecular forces and electrical resistance of materials. As a result of absorption losses, the volumetric heating is generated within materials and ionic solutions [5, 6].
8
Table 2.1 Comparison of conventional thermal processing and microwave heating [4, 7]
Conventional Processing Microwave Processing
Energy Energy is mainly transferred to The microwaves can directly
Transfer the materials through: interact with target materials at the
(i) Conduction molecular level to generate
(ii) Convection volumetric heat.
(iii) Radiation
Heat Thermal gradients. Conversion of electromagnetic
Transfer energy to thermal energy.
Heat External heating sources are Heat is produced internally and
Generation necessary and diffusion of heat throughout the whole volume of the
from the material surface is material.
responsible for heating up the
whole substance.
Heating rate Heating rate is usually low Heating rate is very high and
& Process which results in prolonged consequently shorten the processing time processing time. periods.
Direction of Heat flows are from the surface As the microwaves penetrate inside heat flow of the material to the core. Thus, the material, heat is generated
9
the surface maintains higher firstly at the core and flows outside
temperature than the core. towards the surface. Thus, the
internal site maintains a higher
temperature than external surface.
Influence of The heating profile is dependent The heating efficiency is largely material mainly on the material’s depending on the material’s properties capability to conduct heat i.e. capability to convert
thermal conductivity of the electromagnetic waves into thermal
substance influences most. energy i.e. permittivity,
di-electric loss, permeability and
magnetic loss are the most
influential factors.
Previous studies have shown that the microwave irradiation can remarkably increase the chemical reaction kinetics through promoting the nucleation and crystal growth by orders of magnitude. Compared with acceleration in crystal growth, the microwave irradiation appeared to be more effective in intensifying the nucleation of as-synthesized materials [8]. Unlike the hours/days consumed in chemical reactions heated by conventional conductive method, the microwave heating can shorten the reaction/chemical transformation process to minutes and even seconds to significantly reduce the time and energy consuming. In microwave chemistry, the key factor that affecting the chemical reaction rate is the ability of reaction medium/mixture to absorb the microwave energy and induce rapid in-core volumetric heating by direct coupling microwave at molecular level. 10
Thus, according to Arrhenius equation, the thermal microwave effects including rapid heating kinetic and high processing temperature in sealed reactor undoubtedly explained the reason for the dramatically enhanced reaction rate in microwave assisted synthesis process. Although most work has suggested that the extremely high reaction rate of microwave processing can be purely attributed to the thermal/kinetic phenomena of microwave energy, some specific microwave effects have also drawn growing attentions from researchers. Owing to the specials of microwave dielectric heating, the specific microwave effects include the superheating effect of solvent, selective heating of polar species in a less polar medium and formation of molecular radiators in homogenous solution [9]. The MW superheating effect can elevate the boiling temperature of liquid by
10-20◦C at atmospheric pressure, which is barely observed in conventional heating process.
Due to inhomogeneity of the applied electromagnetic filed and the multiphasic liquid system, selective heating is another specific effect reported in microwave processing. The selective heating can prompt the rapid temperature rise of strongly microwave absorbing reagents and formation of hot spots in reaction medium. The hot spots may maintain a steady-state in the system with slow heat transfer and consequently promote the rate of chemical reaction within the hot zone [9-11].
2.3 Microwave assisted synthesis of calcium phosphates (CaPs) and magnesium phosphates (MgPs)
2.3.1 Hydroxyapatite and calcium deficient hydroxyapatite
Among the CaP compounds, hydroxyapatite (HA) has aroused most interest, as the bone substitutes, due to its similarity to inorganic constituent of cortical bones, superior
11
osteoinductivity and desirable mechanical strength [12, 13]. Despite HA can be extracted from natural source, the wet-chemical precipitation is still the most commonly employed
+ 2+ - 2– route to fabricate HA particles [13, 14]. However, other ions (Na , Mg , Cl , CO3 etc.) from starting chemicals could occupy the vacancies in HA lattice during the precipitation process, which results the formation of calcium deficient hydroxyapatite (CDHA) instead of stoichiometric HA (Ca/P - 1.67) [15-17]. The ionic additions in HA structure have shown many benefits in promoting osteoinductive and osteoconductive properties of HA, which suggests the great potential of using CDHA as the bone substitution material [15,
16, 18]. Table 2.2 listed the microstructures and Ca/P ratios of HA synthesized using microwave assisted technique with various starting materials and param
12
Table 2.2 Methods, parameters and results of MW assisted synthesis of HA
Materials MW Processing Additional HA Structure Ca/P Ref Treatment CaCl2 (10mM) Household MW Filtered and Not revealed Not [19] NaH2PO4 (6mM) Oven (1200W) dried revealed (NH4)2HPO4 5 min DI Water
CaCl2 (10 mM) Household MW Quenched in Tiny platelets (300 nm) Not [20] NaH2PO4 (6mM) Oven (750W) ice for 30 min loosely aggregated in revealed DI Water 5 min Filtered and spherulites (2–4 µm) pH=7.4 dried
CaCl2 (0.5 M) MW Oven Filtered, Tiny rods with a mean Not [21] (NH4)2HPO4 15 min washed and size of 220 nm revealed (0.3 M) dried Most HA particles are DI Water of single crystal pH=10
CaCl2 (0.2 M) Household MW Filtered, Agglomerated and 1.43- [22] Na2HPO4 (0.1 M) Oven (1000W) washed and irregular HA 1.56 with DI Water 5-20 min dried nanoparticles increased pH=7.4 Low crystallinity MW time CaCl2 MDS-6 MW Centrifuged, Nanowires with D of Not [23] Na2ATP Oven washed and about 30 nm and L up to revealed DI Water Sealed autoclave dried several microns 13
pH=5 150- 180 ◦C 30 and 60 min
CaCl2 MDS-6 MW Centrifuged, Nanorods or nanosheets Not [24] Pyridoxal-5’- Oven washed and Self-assemble to form revealed phosphate (PLP) Sealed autoclave dried 3D hierarchical DI Water 140- 180 ◦C nanostructures pH=4.5 5 min
CaCl2 MDS-6 MW Centrifuged, Porous hollow structure Not [25] Creatine-P-Na2·4 Oven washed and Microspheres are within revealed H2O Sealed autoclave dried the range 0.8–1.5 μm DI Water 120◦C and 10 min pH=10 10 × SBF Household MW Filtered, Bone-like structure at 1.44- [26] DI Water Oven washed and low MW power 1.61 pH=7.4 90W- 1200W dried Nanorods with 2-30 min decreased length at high MW power
CaCl2 (0.05 M) MW furnace Centrifuged Nano sized and needle 1.51 [27, 28] Na2HPO4 (0.03 and (300W) and freeze like structure (30-60 0.06 M) 2-120 min dried nm) DI Water Heating rate= 7.8 Crystallinity is pH=7.8 and 8.3 ◦C/min relatively low
14
Ca(NO3) 2·4H2O Household MW Centrifuged, Nanostrips Not [29] Na2HPO4 Oven (900W) washed and average W= 10 nm and revealed CTAB 5 min dried L= 55 nm pH=12
Ca(NO3) 2·4H2O Household MW Washed to Nanorods with D 1.5 [30, 31] Na3PO4·12H2O Oven (700W) remove varying from 60-80 nm Ca/P=1.67 30% - 100% unreacted ions and L about 400 nm Power Dried at 80 ◦C Particle size increases 0.5-1 min and 4 with longer MW time min Agglomerated spheres after 4 min MW
Ca(NO3) 2·4H2O Household MW Centrifuged, Nanorods with D Not [32-35] (0.1M) Oven (700W) washed and varying from 25-40 nm revealed Na2HPO4 (0.06 M) 2 min and 30 min dried and L varying from EDTA (0.1 M) 100-400 nm pH=9-13 Particle size increases with longer MW time
Ca(NO3)2·4H2O Household MW Centrifuged, Nanorods with D of 25 Not [36] (0.5 M) Oven (900W) washed and nm and L varying from revealed (NH4)2HPO4 (0.3 30 min dried 40-75 nm M) DI Water pH=10
15
Ca(NO3)2·4H2O Sonochemistry- Centrifuged, Mesoporous HA Not [37] (NH4)2HPO4 assisted MW washed and nanoparticles revealed DI Water MW reactor freeze-dried Rod-like morphology pH=10.5 (200W) with Heat treated at with length of 50–100 ◦ Ca/P=1.67 ultrasonic probe 550 C for 5 h nm and width of about 60◦C and 30 min 20 nm.
NaNO3 Household MW Filtered, Nano whiskers with D Not [38, 39] Ca(NO3) 2·4H2O Oven (600W) washed and of 100 nm and L about revealed KH2PO4 5 min dired 1 μm HNO3 Aspect ratio about 7-10 Urea DI Water Ca/P=1.67
Ca(NO3) 2·4H2O Household MW Filtered, Spherical particle with 1.67 [40] H3PO4 Oven (900W) washed and average D equals 40 nm NaOH 8 min dired pH=10
Ca(NO3) 2·4H2O Household MW Stirred to Needle like CDHA 1.5 [41] H3PO4 Oven (800W) form paste- particle with D varying NH4OH 15 min like from 7–16 nm and L Ca/P=1.5 precipitates varying from 16–39 nm Precipitates MWed Grinded
16
CaCl2 or Mars 5 MW Oven Centrifuged, HAp particles with size 1.29- [42] Ca(NO3) 2·4H2O or (800W) washed and varying between 29.5- 1.43 Ca(OH)2 15 min dried 45.4 nm and (NH4)2HPO4 Heat treated at crystallinity fraction ◦ DI water 950 C for 1 h varying between 0.53- pH=7.4 2.37
Ca(OH)2 Household MW Washed and Spherical particle with 1.67 [43] H3PO4 Oven (700W) dried the size ranging from 50 DI Water 300◦C and 25 min Heat treated at to 90 nm in diameter Ca/P=1.67 750 ◦C in air At pH=12, monolithic pH=6-12 HA with the particle of 50-70 nm
Ca(OH)2 High density MW Washed and Needle-like particle 1.57 [44] H3PO4 irradiation dried with aspect ratio Ca/P=1.67 90 s ranging from 2 to 5 and 1 MPa pressure average size below 6 nm
Eegshell derived Household MW Washed and Platelet particles: length Slightly [45] Ca(OH)2 (0.3M) Oven (700W) dried 33–50 nm and width 8– over 1.67 (NH4)2HPO4 (0.5M) 14 nm DI Water Ca/P=1.67
17
Ca(CH3COO)2 Microwave– Centrifuged, Hollow spheres Not [46, 47] Na2HPO4 hydrothermal washed and assembled with revealed NaH2PO4 synthesis system dried nanosheets to form a Ca/P=1 80-120◦C and 5- hierarchical flower-like pH=5 60 min structure D of spheres are in the range of 0.5-3 μm Thickness of nanosheet is around 50 nm
18
Over past two decades, microwave assisted process had delivered many interesting results in the rapid preparation of HA and CDHA. In 1991, lerner and his colleagues firstly reported the preparation of HA particles using calcium chloride as the calcium source, sodium dihydrogen phosphate and diammonium hydrogen phosphate as phosphorous source by MW irradiation in 5 min and subsequent MW irradiation of collected solids could significantly improve the crystallinity of precipitates [19]. A similar approach was further explored by Sarig et al. to produce plate-like HA nanoparticles using dilute calcium chloride and sodium phosphate solutions with Ca/P=1.67 and pH=7.4 [20]. It was found that the nHA precipitates loosely aggregated in spherulites with the dimensions around 2-
4 µm and MW energy could weaken the bonding between the cation and the hydration sphere which is fundamental for the formation of HA crystals in aqueous solutions. Needle- like and irregular calcium deficient hydroxyapatite (CDHA) nanoparticle (Ca/P=1.5) with low crystallinity can be prepared by microwave heating and using sodium hydrogen phosphate as the phosphorous source [22]. The authors also reported the application of citrate ions as the Ca-complexing agent to inhibit the crystal growth of HA crystals. Lately, phosphorous containing biomolecules have been studied as the phosphorous sources, which are helpful to avoid disordered crystal growth of HA. Qi et al. prepared HA nanowires with the diameter of 30 nm (Fig. 2-2) using adenosine 5'-triphosphate (ATP) as the phosphorous source via a one step and surfactant-free microwave-assisted hydrothermal method in a short period (30 min) [23]. Similarly, Zhao et al. employed the pyridoxal-5′-phosphate (PLP) as the organic phosphorous source in the experiment and prepared nanosheets and nanorods by microwave heating in 5 min [24]. It was also revealed that the as-prepared nanosheets self-assembled to form a 3D hierarchical nanostructure. 19
Creatine phosphate and Fructose 1,6-bisphosphate trisodium salt (FBP) had also been investigated as one of the potential organic phosphorous sources to produce HA particles
[25]. HA microspheres with the size ranging from 0.8–1.5 μm could be easily prepared using calcium chloride as the calcium source, creatine phosphate as phosphorous source by a simple microwave assisted hydrothermal process in 10 min. It is noteworthy that an environment with elevating pH value (pH=10) should be maintained for using creatine phosphate as the phosphorous source, compared to that of adenosine 5'-triphosphate (ATP) and pyridoxal-5′-phosphate (PLP).
Figure. 2-2 SEM micrographs (a), TEM micrographs (b) and diameter distribution (c) of
HAP nanowires prepared using an aqueous solution containing CaCl2·2H2O
and Na2ATP by the microwave-assisted hydrothermal method at 160 °C for
30 min [23]. Reproduced with the kind permission from Elsevier SA.
Another important aspect in the microwave assisted synthesis of HA particles is using calcium nitrate as the calcium source in aqueous solution to produce HA nanorods or nanostrips. Arami et al. prepared HA nanostrips directly by applying microwave irradiation to the precursor solution containing calcium nitrate, sodium hydrogen phosphate and CTAB. In this work, the cationic surfactant CTAB serving as the soft template assisted the epitaxial growth of HA crystals [29]. Moreover, multiple studies
20
reported the successful preparation of HA nanorods with the diameter varying from 20-80 nm and length varying 40-400 nm by microwave assisted synthesis methods [30-32, 34,
36, 37]. One of these studies carried out by Liu et al. explained crucial effects of OH− on the crystallite facets and changing complex stability of Ca–EDTA in altering the shapes of as-synthesized HA crystals (Fig. 2-3) [33]. Moreover, Liang et al. proposed the method combining template-free sonochemistry and MW irradiation for preparation of mesoporous
HA nanoparticles with high efficiency [37]. The HA nanoparticles were synthesized by sonochemistry assisted MW heating of aqueous precursors containing Ca(NO3)2·4H2O and
(NH4)2HPO4 in 30 min. The as-prepared HA particles showed rod-like structures with length of 50–100 nm and width of about 20 nm. The formation of mesoporous structure with a mean pore size of 2-3 nm could be attributed to the MW irradiation. Other than HA particles synthesized using aqueous solution, one step solid state reaction was also demonstrated to be a feasible route to prepare HA nanorods. Additionally, the prolonged
MW heating could result the increase of HA particles. To synthesize the HA with high aspect ratios and whisker-like structures, Jalota et al. reported a "combustion (autoignition) synthesis/Molten salt” hybrid method. This microwave assisted solution combustion synthesis (MASCS) process involved the microwave triggered auto ignition of Ca(NO3)2
(fluxing agent) and urea (fuel) in aqueous solution by using a household microwave oven at 600 watts for 5 min. Then the irradiated precursors were stirred for 1 hour to obtain nano-sized HA whiskers with a high aspect ratio of 10, as shown in Fig. 2-4 [38].
Interestingly, Elkady et al. reported the formation of spherical HA nanoparticles with a mean size of 68 nm by microwave heating of mixture that contained calcium nitrate and phosphoric acid for 8 min. In comparison to the HA nanorods and nanowhiskers prepared 21
in many other studies, the HA nanoparticle with a spherical structure reported in this work could be the result of high pH value (pH=10) of the mixture adjusted by sodium hydroxide solution.
Figure. 2-3 (a) Low magnification SEM image showing HAp nanostructures. (b) Higher
magnification SEM image of a typical bundle of the HAp nanostructures. (c)
TEM image of a typical bundle of HAp. Inset: SAED pattern taken from an
individual HAp nanorod (d) HRTEM image recorded from the tip of the
individual HAp nanorod [33]. Reproduced with the kind permission from IOP
Publishing Ltd.
a b
Figure. 2-4 SEM micrographs of (a, b) HA nanowhiskers synthesized using microwave
assisted combustion process [38]. Reproduced with the kind permission
from Cambridge University Press. 22
Unlike the extensively investigated calcium source such as calcium chloride and calcium nitrate, limited work has been done and reported by using calcium hydroxide and calcium acetate as the starting materials in the microwave assisted synthesis of HA.
However, some recent results explored the potentials of calcium hydroxide and calcium acetate serving as the calcium source in the preparation of HA with various compositions and structures. Needle-like HA nanoparticles with outstanding thermal stability can be prepared using MW irradiation on calcium precursors that consist of calcium hydroxide derived from eggshell wastes [45]. The as synthesized platelet particles showed a length ranging from 30-55 nm and width ranging from 8-14 nm. A MW-assisted solvothermal method was also developed to produce needle like CDHA (Ca/P=1.57) particle [44]. The
CaP precursors contain calcium hydroxide and phosphoric acid were placed in a sealed reactor and heated by high density MW irradiation to reach the processing temperature which is higher than boiling temperature of solvent. In addition, Wang et al. reported the preparation of HA hollow spheres using calcium acetate as the calcium source via the microwave assisted hydrothermal with copolymer template [46]. The concentration of copolymer template such as poly- (lactide)-block-poly (ethylene glycol) (PLA–PEG) was revealed to be a crucial aspect to the formation of hollow structures. Later, this research group also delivered a template-free method by replacing copolymer template with sodium citrate dehydrate to fabricate HA hollow spheres assembled by HA nanosheets (Fig. 2-5)
[47].
23
Figure. 2-5 TEM micrographs of the samples prepared by microwave-assisted
hydrothermal method for 30 min. (a) and (b) [Na3Cit] = 0 mM, 120 °C; (c)
and (d) [Na3Cit] = 2.5 mM, 140 °C; (e) and (f) [Na3Cit] = 5 mM, 80 °C [47].
Reproduced with the kind permission from Elsevier SA.
2.3.2 Microwave assisted synthesis of ion doped hydroxyapatite
The development of multifunctional biomaterials with enhanced biological activities and unique roles in bio-imaging, drug delivery and antibacterial activities has attracted considerable interests in recent years. Due to the flexibility of the apatitic structure, many cationic ions such Mg2+, Sr2+, Cu2+, Ag+ can be easily incorporated in hydroxyapatite crystals and replace Ca2+ in the lattice structure [48, 49]. Therefore, HA is widely recognized as the ideal host for a great number of ion substitutions and potential applications in fields of biodiagnosis, bone regeneration and treatment. As stated previously, microwave assisted synthesis is a favorable method to prepare HA particles with benefits including rapidness, versatility and energy saving. Thus, it is also highly 24
anticipated that microwave energy can facilitate the preparation of HA particles with diverse dopants.
Rameshbabu et al. prepared the silver (0-3 at%) substituted HA nanoparticles using microwave heating of aqueous solution containing calcium hydroxide, diammonium hydrogen phosphate and silver nitrate for 30 min [50]. The as-prepared AgHA particles showed a nano-sized needle like structure with width ranging from 15 to 20 nm and length around 60–70 nm (Fig. 2-6) The bactericidal activity and osteoblast spreading were more evident on AgHA particles with low Ag concentration (eg. 0.5 at%). Other researchers reported the Ag-HA prepared using similar approach with an Ag concentration as low as
0.3 wt% was optimal [51]. The low amount substitution of silver will not alter the structure and thermal stability of HA. Moreover, the surfactant agent (CTAB) is helpful to control morphologies and particle sizes of silver doped hydroxyapatite particles [52]. These studies also demonstrated that silver doped hydroxyapatite was highly active against common bacteria, such as staphylococcus aureus, bacillus subtilis, pseudomonas aeruginosa and escherichia coli, as shown in Fig. Later, the same research group further synthesized silver and zinc co-doped hydroxyapatite nanoparticles. The spherical Ag-Zn-
HA particles delivered better bioactivity and antibiotic properties, as compared to those of pristine HA particles. A structural analysis reported the doping Zn2+ in pure HA could transform partial of its hexagonal phase to monoclinic phase and the monoclinic phase was stabilized by the doped Zn2+ ions [53].
25
Figure. 2-6 The TEM morphology of the HA (a), 0.5 AgHA (b), 2 AgHA (c) and 5 AgHA
(d) samples in as synthesized condition [50].
Figure. 2-7 Antibacterial activity (a); E. coli, (b); P. aeruginosa, (c); S. aureus and (d); B.
Subtilis. (1) HA; (2) 0.3Ag-HA; (3) 1Ag-HA; (4) 5Ag-HA [51].
26
Magnesium ion (Mg2+) plays a key role in the bone metabolism by controlling the cellular activity of osteoblasts and stimulates osteoblasts proliferation. The low addition of
Mg2+ (0.06 mol%) in HA lattice results the formation of 2-D plate-like nanostructures rather than the common nanorods [54]. Magnesium-silver co-doping induced by microwave irradiation is beneficial to microhardness, photoluminescence, antimicrobial efficacy and in vitro bioactivity of hydroxyapatite [55]. Strontium (Sr) is a trace element in human body and has shown beneficial effects on many biological activities, such as promoting bone formation and inhibiting bone resorption [56]. Therefore, strontium substituted apatite has been increasingly synthesized and investigated for applications in osteoporotic bone treatment. Ravi et al. prepared Sr substituted calcium deficient phosphate with the (Ca + Sr)/P ratio of 1.61 by microwave processing [57]. The Sr-CDHA particles showed rod and acicular-like morphologies with the dimensions around 50 nm × 7 nm.
Moreover, excellent cell viability, good cellular localization with HPDLF cells and significant antimicrobial activity were observed on Sr-CDHA nanoparticles. Co- substitutions of Sr and Ce in pure HA has also been studied for synergistic benefits on bioactivity and antibacterial properties [58]. The XRD and HRTEM results suggested that the incorporation of Sr and Ce ions in HA lattice structure can decrease crystallinity and crystal size. While antibacterial activity and thermal stability of HA nanoparticles were significantly enhanced by the doped Sr2+ and Ce3+ [58]. Silicon (Si) is another essential trace element in natural bones and has been demonstrated to play an important role in bone regeneration through inducing the formation extracellular matrix in bone and cartilage [59].
To take advantage of favorable effects of Si, Si dope HA particles were synthesized by microwave heating of calcium phosphate precursors with addition of 27
Si(OCH2CH3)4(TEOS) as the silicon source for 30 min [60]. The incorporation of silicon in pure HA lattice occurred via silicate group replacing phosphate group. Substitution of
Si also resulted the decrease in particle size and altered the morphology of HA crystals to elongated ellipsoidal like structure.
Recently, hydroxyapatite serving as phosphate hosts that are doped with luminescent lanthanide cations has attracted considerable interests as the luminescent agent for bioimaging applications. This is due to its low toxicities, high thermal and chemical stabilities, good photostabilities and narrow emission/excitation peaks [61]. To date, a handful of work has investigated the possibility of using microwave energy to prepare lanthanide-doped hydroxyapatite nanoparticle and endow the material with fluorescent properties. Escudero et al. reported the formation of polycrystalline and spindle-like hydroxyapatite and fluorapatite based nanophosphors doped with europium by a microwave-assisted hydrothermal process at 180 °C [62]. The as-prepared Eu-HA nanoparticles exhibited a dimension of 191 × 40 nm and typical red luminescence (Fig 2-
8). However, the luminescent properties of fluorapatite were superior to those of hydroxyapatite. Andre et al. prepared Eu doped HA nanorods by using comparable microwave assisted hydrothermal method and monitored the structural evolution in HA lattice crystallization via the photoluminescence (PL) measurements from Eu3+ emission
[63]. The microwave assisted microemulsion process has also been studied to synthesize luminescent Eu3+ doped hydroxyapatite nanoparticles. The morphology and particle size of final products were found to be controlled by initial pH value of the precursor solution
[64]. Later, Wagner et al. proposed a simple microwave-assisted solution combustion synthesis (MASCS) to produce europium doped hydroxyapatite nanowiskers (Fig. 2-9) for 28
multifunctional bioimaging applications [65]. The Eu3+ doped HA nanowhiskers synthesized in that work presented several unique properties such as high crystallinity and aspect ratio, non-agglomeration, high chemical purity and low toxicity. These nanowhiskers were also found to be capable of emitting at red and far red wavelengths utilizing UV excitation. To enhance photoluminescence effects and further tailor biological response of hydroxyapatite, Chen et al. synthesized Eu3+/Gd3+ dual-doped HA nanorods using the microwave assisted hydrothermal method [66]. With the doping of Eu3+ and Gd3+, the sizes of nanorods were found to be smaller with a more centered distribution.
Furthermore, the excellent luminescent and magnetic properties of Eu3+/Gd3+ dual-doped
HA nanorods paved its way towards multiple applications in bioimaging and biodiagnostics fields (Fig. 2-10). Alshemary et al. obtained the mesoporous erbium-doped
(Er3+) hydroxyapatite nano-sized particles with a mean crystallite size of 25 nm using calcium nitrate, diammonium hydrogen phosphate and erbium(III) nitrate pentahydrate under microwave irradiation in a domestic microwave oven [67]. In addition,
Er3+ incorporated HA exhibited desirable ability to promote apatite formation in vitro and good photoluminescence properties including red and green emission.
Figure. 2-8 The TEM graphs of the Eu doped HA nanoparticles with corresponding
luminescence of suspensions of Eu-HA when irradiated with UV light [62]. 29
Figure. 2-9 SEM images, particle size distribution and typical luminescent image of Eu
doped HA nanowhiskers synthesized using MASCS method [65].
Figure. 2-10 In vivo PL imaging of the mice after subcutaneous injection with
Eu3+/Gd3+–HAp nanorods [66].
30
Besides above-mentioned dopants, many other ions such as Li+, Cu2+, Mn2+, Co2+, Ni2+,
Fe3+ had been incorporated in HA singly or multiply by microwave assisted methods with aspiration of promoting its biological activities and delivering specific functions. Badran et al. prepared Li doped HA nanoparticles (14 to 21 nm) by a sol-gel technique associated with microwave-hydrothermal treatment [68]. The Li additions were found to increase the crystallinity and Gamma attenuation coefficient of HA. The copper (Cu) doped mesoporous hydroxyapatite (HA) microspheres could be synthesized using microwave assisted hydrothermal procedure with creatine 26 phosphate as an organic phosphorus source [69]. The as-synthesized mesoporous microsphere with high specific surface area and drug delivery capability were assembled by nano-sized rod like or sheet like Cu-HA crystals. Magnetic ions including iron (Fe), cobalt (Co), nickel (Ni), etc., which can endow
HA particles with electromagnetic properties and expedite their applications in magnetically targeted drug delivery and magnetic resonance imaging (MRI) [70]. Thus, the rapid synthesis of Co2+ doped, Fe3+ doped and Co2+, Fe3+ co-doped HA nanoparticles by microwave irradiation treatment had been reported [53, 71-73]. The Co2+and Fe3+ dopants could reduce the crystallinity of HA and change its morphology to nanorods from spherical structure. In addition, the both Co2+ and Fe3+ dopant ions showed outstanding capability to convert the diamagnetic HA particles to be superparamagnetic, as shown in
Fig. 2-11. In comparison to Fe3+ doped HA, the Co2+ dope HA exhibited more sustained drug release, favorable electrical and magnetic behaviors. Most recently, studies involved using molybdenum (Mo), tungsten (W), boron (B), Tellurium (Te) etc. as doping elements to customize the structure, morphology, bioactivity, strength, dielectric and optical properties of HA to suit its various applications [74-77]. 31
Figure. 2-11 Magnetization curves for pristine and Co2+ doped samples in an applied
magnetic field of ±2 tesla (a) 0 Co and (b) 20 Co [71].
The ion substitutions in hydroxyapatite lattice could also take place between carbonate
2− 3− ions (CO3 ) and phosphate ions (PO4 ). Such ion substitutions had already been found and reported in carbonate apatite minerals and calcinated bones [78]. Kumar et al. prepared carbonate substituted hydroxyapatite (CHA) by heating calcium phosphate precursor modified with calcium carbonate in a domestic microwave oven [79]. The decrease in a-
2− axis of synthesized CHA could be attributed to the B-type substitution of CO3 , which
3− replaced the PO4 tetrahedral group in HA lattice. In contrast, the A-type carbonate substitution means the carbonate ions occupy the hydroxyl sites, as illustrated in Fig. 2-12.
And this kind of substitution could expand the a-axis of HA lattice [78]. Due to the lattice distortion caused by carbonate incorporation, the resorbability of substituted HA will consequently increase. Fluoride (F-) is another kind of ions that can be incorporated in HA structure to partially and completely replace hydroxyl groups (OH-). The fluorine substituted hydroxyapatite could be produced using microwave energy either by wet- chemical precipitation or by combustion synthesis [80, 81]. It is noteworthy that fluorapatite particles synthesized by microwave assisted solution combustion synthesis
32
(MASCS) showed unique nanotube structure with a “Y” like inner channel [81]. Moreover, due to high binding energy between F- and Ca2+, the low solubility and degradation are anticipated for fluorine substituted hydroxyapatite [82].
Figure. 2-12 Hexagonal unit cell of a hydroxyapatite crystal with possible carbonate
substitutions [83].
2.3.3 Microwave assisted synthesis of biphasic calcium phosphate
To regulate the resorption kinetics and achieve better biological properties of the scaffolds, the attentions have been drawn to the fabrication biphasic calcium phosphate system, such as HA - tricalcium phosphate (TCP) [84, 85]. The motivation is to tailor the ratio of the non-degradable phase (i.e. HA) and more soluble phase (i.e. TCP) to control the rate of biodegradation, promote bone remodeling around the implants [84, 86].
Following this concept, Manjubala et al. produced the biphasic CaP bioceramic with various HA/TCP ratios using a domestic microwave oven [87]. The enhanced bioresorbablility was detected with increasing content of TCP in biphasic powder, which possesses excellent osteoconductive and osteoinductive properties [87, 88]. This also demonstrated the feasibility of using microwave processing to synthesis biphasic CaPs for
33
bone substitute applications. Similarly, Farzadi et al altered the Ca/P ratio in the initial solution to control HA/β-TCP ratio of prepared biphasic products [89]. Compared to the conventional wet chemical method, microwave energy was found to improve the crystallinity and the amount of HA particles in BCP. Jalota et al. reported the preparation of HA-TCP biphasic nanowhiskers/nanoparticles by reducing the content of the fuel agent
(urea) in the initial microwave assisted combustion reaction [38]. Most of the as-prepared
TCP particles were rhombohedra like particles with a minor number of whiskers, while biphasic samples mainly consisted of whiskers, as seen in Fig. 2-13. Microwave assisted densification was also employed to fabricate the porous biphasic calcium phosphate ceramic with fine grain size, uniform microstructure, high porosity and enhanced cellular responses of osteoblast-like cells [90, 91].
Figure. 2-13 SEM micrographs of (a) TCP and (b) biphasic HA-TCP samples prepared
via microwave assisted combustion synthesis [38].
2.3.4 Microwave assisted synthesis of amorphous calcium phosphate (ACP)
Amorphous calcium phosphate (ACP) is an essential precursor in mineralization of bone tissues. As the rising interest in calcium phosphate compounds, ACP maintains high
34
solubility, excellent ability to initial biomineralization, and was widely recognized as the metastable precursor phase for reformation of apatite in biological environment [92, 93].
The most commonly used procedure to prepare ACP is called low temperature-wet chemical precipitation route, which involves the decomposition and nucleation of calcium and phosphate salts in aqueous and alcoholic medium [92, 94, 95]. Previous research efforts have revealed advantages of microwave assisted wet chemical synthesis, such as accelerated nucleation and crystal growth, yield improvement and emergence of new material phases [10]. Thus, it is expected the wet-chemical precipitation assisted by microwave irradiation could a novel, simple and rapid route to synthesize ACP particles for numerous applications in tissue engineering. Tang et al. firstly reported the formation of ACP nanoparticles in a weakly acidic medium after microwave irradiation in a domestic microwave oven at 80 °C [96]. The size of synthesized ACP spherical nanoparticles (10 nm) was smaller than that of HA nanorods which are prepared using the same manner.
Later, zhou et al prepared ACP nanospheres by microwave heating supersaturated biomimetic fluids (SBFs) with various ionic concentrations, as seen in Fig. 2-14 [97]. The pH value and ionic concentrations were found to play an import role in precipitation of
ACP nanospheres. Additionally, the highly active ACP could be stabilized by Mg2+ induced in precipitation process through inhibiting the growth of apatitic crystal. In other words, Mg2+ was beneficial for the deposition of spherical ACP nuclei. In vitro cell culture assay suggested the ACP nanospheres possessed comparable biocompatibility with HA.
35
Figure. 2-14 TEM images of (a) and (b) ACP nanospheres and (c) related clusters after
transition [97].
As inspired by successful application of organic phosphorous source in the microwave assisted synthesis of HA, organic phosphorous sources such as adenosine 5′-triphosphate disodium salt (ATP) and cytidine 5′-triphosphate disodium salt (CTP) had also be studied and adapted in the synthesis of nano-sized ACP particles. The ACP porous nanospheres with uniform size and high stability in phosphate buffer saline (PBS) solution were prepared using a facile microwave assisted hydrothermal method [98]. As shown in Fig. 2-
15, the organic phosphorous source ATP not only supplied phosphate ions in this process, but also acted as the stabilizer to inhibit the transformation of ACP to HA. Further, with the incorporation of soybean lecithin or block copolymer methoxyl poly(ethylene glycol)- block-poly(d,l-lactide) (mPEG-PLA) as the template in this process, the ACP porous microsphere with hollow features became more predominant [99, 100]. Besides ATP, another biocompatible biomolecule - cytidine 5′-triphosphate disodium salt (CTP) can also
3- serve as organic phosphorus source by hydrolysis to produce phosphate ions (PO4 ) in precursor and react with pre-existing calcium ions to form ACP nanoparticles under suitable pH environment and microwave irradiation [101].
36
Figure. 2-15 Schematic diagram of the ACP formation process via microwave assisted
hydrothermal route and the functions of soybean lecithin as the template [98,
99].
Lately, ACP has been investigated as the new multifunctional nano-system for the potential capabilities of simultaneous in vivo bioimaging and treatment. To enable ACP nanoparticles with high stability, antibacterial and photoluminescence properties, functional ions such as Zn2+, Mg2+ and Eu3+ had been incorporated in ACP structure using microwave assisted method. Qi et al. prepared amorphous calcium magnesium phosphate
(ACMP) with high stability in aqueous condition and excellent biological properties by microwave heating the precursor which contains CaCl2, MgCl2 and creatine phosphate
(CP) for 10 min [102]. The initial ratio of Ca/Mg was found to be important to adjust phases and morphologies of the final products. Moreover, the as-prepared ACMP hollow microspheres exhibited a high degradation rate and pH-sensitive behavior, which were desirable for their potential applications in drug delivery. Zhao et al. implemented the
37
similar approach to produce mesoporous ACP microspheres doped with zinc ions for antibacterial benefits [103]. It was demonstrated that the release of Zn2+ from Zn-ACP microspheres was pH dependable with more zinc ions release in lower pH environment.
Also, the antibacterial activities of Zn-ACP particles against bacteria S. aureus and E. coli were examined to be significant. The addition of Eu3+ in mesoporous ACP microsphere could result in decrease in pore size (17.8 to 10.1 nm with 10 mol% Eu3+ addition), while increase in specific surface area (209 to 315 m2 g -1 with 10 mol% addition) [104]. Fig. 2-
16 showed the typical morphology, amorphous nature and europium additions of the prepared Eu-ACP microspheres with 5 mol% Eu3+ doping. Ascribed to the florescent behavior of doped Eu3+, the Eu-ACP microspheres synthesized using microwave assisted solvothermal method presented competitive luminescent imaging abilities in vitro/in vivo
[104, 105].
Figure. 2-16 TEM micrographs (a,b), SAED pattern (c) and EDS spectrum (d) of 5 mol
% Eu-ACP mesoporous microspheres prepared the microwave-assisted
solvothermal method [104].
38
2.3.5 Microwave assisted synthesis of magnesium phosphate compounds
Biomaterials with magnesium containing have shown high biocompatibility, controlled biodegradability, promoted osteoblast cell attachment and proliferation [106-108]. Thus, research interests on magnesium based biomaterials including magnesium alloys and magnesium phosphate ceramic have been continuously emerging in last decade.
Microwave processing has shown a great success in synthesis of multifunctional calcium phosphate nanoparticles with various phase compositions, diverse microstructures and many specific properties. It is also expected that the microwave assisted method can be a feasible strategy to prepare its alternative - MgP nanoparticles and overcome the instability issue associated with magnesium phosphate compounds in aqueous medium. Amorphous magnesium phosphate (AMP) nanospheres (Fig. 2-17) could be synthesized from precursor
2+ 2− 3− containing Mg and HPO4 /PO4 with 5 min microwave irradiation [109]. The as- prepared AMP nanospheres exhibited high activity in physiological environment and could transform to mature MgP phase such as Bobierrite (Mg3PO4·8H2O) to sustain osteoblast cell proliferation. A simple microwave assisted hydrothermal method was developed for rapid synthesis of amorphous magnesium phosphate with flower-like hierarchical nanostructures [110]. The organic phosphorous salt fructose 1,6-bisphosphate trisodium
(FBP), treatment temperature and time were found to play important roles in the formation of flower-like AMP microspheres. In addition, the self-assembly behavior of AMP nanosheets in the synthesis process resulted the unique evolution of flower-like hierarchical nanostructures. Applying comparable microwave assisted hydrothermal method along with other organic phosphorous sources or inorganic source such as creatine phosphate (CP)
39
ATP or sodium dihydrogen phosphate respectively could produce magnesium phosphate hydrate nanosheets for various biomedical applications in drug delivery and protein absorption [111-113].
Figure. 2-17 TEM micrographs (a) single AMP nanosphere and (b) fused AMP
nanospheres prepared by the microwave-assisted method [109].
2.4 Microwave assisted surface modification of biomaterials
Surface characteristics have been demonstrated to play a key role in the long term success of medical implants [114]. By modifying the surface characteristics of implant materials, the biodegradation behavior can be tailored, the biocompatibility, bioactivity, osseointegration can be promoted, and the healing process can be subsequently shortened.
Recent advances in microwave processing of biomaterials have led to the rapid surface modification of bio-implant materials accelerated by microwave energy. The microwave assisted coating technique was firstly developed on non-degradable bioimplants such as Ti alloy to improve their biological activities and biomineralization capabilities. By taking the advantage of conventional biomimetic coating process and microwave chemistry, a uniform, ultrathin and crack-free apatite coating could be deposited on Ti6Al4V alloy 40
within a few minutes [115]. It was revealed that the doping Mg2+, Sr2+, changing initial
Ca/P ratio, alkali/heat treatment can significantly modify the morphology and properties of as-deposited apatite coatings. For example, incorporating Mg2+ in the apatite coating can smoothen the coating surface and notably enhance the cellular response [115]. The decrease of surface roughness and nucleus size was observed due to the presence of strontium in the coating solution, as shown in Fig. 2-18. In addition, with certain amount of strontium substituted in apatite coating structure could promote the proliferation and differentiation of osteoblasts [116]. Later, alkalization associated with heat treatment of Ti substrate at 600°C was developed with benefit in facilitating the generation of apatite coating in the microwave assisted deposition process [117]. Most recently, a hybridization method involving laser micropatterning, microwave heating and in situ synthesis was performed to produce titania/hydroxyapatite/tricalcium phosphate (TiO2/HA/TCP) composite coating on Ti6Al4V alloy substrate for enhanced osteoconductive properties and apatite inducing capability [118].
Figure. 2-17 Representive AFM images of apatite coating with strontium concentration
of (a) 0, (b) 2%, (c) 4% and (d) 6% in a 5 μm × 5 μm area [116]. 41
Magnesium alloys are widely recognized as revolutionary implant materials, due to their uniqueness of biodegradation and desirable mechanical properties. To successfully utilize commercial Mg alloys in clinical application, coating Mg alloys with protective and biocompatible materials is necessary. Therefore, the rapid coating of Mg alloys with alkaline earth phosphate bioceramic using microwave energy has been developed. The microwave irradiation can dramatically promote the nucleation rate of bioceramic in aqueous condition, create rapid covering of Mg substrate with nascent nuclei and subsequently retard the initial degradation of Mg alloys. For instance, the calcium deficient hydroxyapatite coating could be deposited on AZ31 magnesium alloy in less than 10 minutes via a microwave assisted coating process [119]. The time consumed in this process was much less than that commonly costed in biomimetic treatment. Especially, the as- prepared HA/CDHA coating showed a functional structure with a dense inner layer blocking corrosive medium and a loose outer layer enhancing apatite precipitation [119,
120]. Moreover, a recent study reported the preparation of magnesium phosphate coating on Mg alloys by microwave heating the sealed vials containing MgP precursors and Mg alloy substrates with various temperatures and periods [121]. It was found that the increasing treatment temperature could partially transform the newberyite phase to trimagnesium phosphate phase with enhanced chemical stability and corrosion resistance.
Besides the metallic implant materials, microwave assisted coating technique has also been applied on PEEK, chitosan scaffolds and carbon/carbon composite to produce amorphous and crystalline calcium phosphate coatings for improved bioactivities and mechanical properties [122-124].
42
Chapter 3
Rapid Coating of AZ31 Magnesium Alloy with Calcium Deficient Hydroxyapatite Using Microwave Energy
3.1 Abstract
Due to their unique biodegradability, magnesium alloys have been recognized as suitable metallic implant materials for degradable bone implants and bioresorbable cardiovascular stents. However, the extremely high degradation rate of magnesium alloys in physiological environment has restricted its practical application. This paper reports the use of a novel microwave assisted coating technology to improve the in vitro corrosion resistance and biocompatibility of Mg alloy AZ31. Results indicate that a dense Calcium
Deficient Hydroxyapatite (CDHA) layer was uniformly coated on AZ31 substrate in less than 10 minutes. Weight loss measurement and SEM were used to evaluate corrosion behaviors in vitro of coated samples and of non-coated samples. It was seen that CDHA coatings remarkably reduced the mass loss of AZ31 alloy after 7 days immersion in SBF.
In addition, the prompt precipitation of bone-like apatite layer on the sample surface during
43
immersion demonstrated good bioactivity of the CDHA coatings. Proliferation of osteoblast cells was promoted in 5 days incubation, which indicated that the CDHA coatings could improve the cytocompatibility of the AZ31 alloy. All the results suggest that CDHA coatings, serving as a protective layer, can enhance the corrosion resistance and biological response of magnesium alloys. Furthermore, this microwave assisted coating technology could be a promising method for rapid surface modification of biomedical materials.
3.2 Introduction
Recently, magnesium and its biodegradable alloys have been extensively studied as revolutionary implant materials due to their unique properties such as high strength/weight ratio, comparable mechanical properties with natural bone and excellent biocompatibility[125-129]. Furthermore, the Mg2+ ion has been demonstrated to stimulate initial osteogenesis and promote new bone formation [130]. However, Mg and its alloys degrade much too quickly in physiological environment due to the high concentration of chloride ions [125, 131, 132]. Moreover, hydrogen gas evolved as a corrosion by-product may cause problems against the surrounding tissues, thus hindering the attachment of osteoblasts to the implant surface [133].
To successfully employ these alloys, some kind of protective coating is desirable. To date, there have been many technologies developed to produce such coatings on magnesium and its alloys [134-136]. The most common methods are electroplating, conversion coating, anodizing, organic coating and thermal spraying [137-141]. Although
44
electroless nickel plating and chromate conversion coatings have shown promising corrosion resistance, the toxicity in the nature of these treatment processes restricts their application in the human body. Oxide films produced by anodizing coatings have shown adverse results in fatigue behavior of substrate and poor biocompatibility [142, 143].
Furthermore, pinholes and blisters formed in organic polymer coating processes and inherent porosity produced by thermal spraying facilitates galvanic corrosion between the porous coating and magnesium substrate [134]. Besides the disadvantages addressed above, poor biocompatibility of the deposited coating is another shortcoming in these techniques [134]. Therefore, recent publications have focused on bioactive ceramic coatings such as calcium phosphate coatings, which presented good corrosion resistance as well as outstanding biocompatibility and bioactivity [129, 132, 144-146].
Calcium phosphates (CaP), especially Hydroxyapatite/Calcium Deficient
Hydroxyapatite (HA/CDHA, Ca10(PO4)6(OH)2) have been broadly used as biomedical materials for their chemical similarity to bone mineral. In addition, CaP exhibits non- toxicity, high biocompatibility and superior ability to induce physicochemical bonds between the implant and cortical bone i.e. osseointegration [147-150]. To date, several forms of calcium phosphate coating have been developed on magnesium and its alloys by various techniques, such as chemical precipitation[143, 151-156], electrodeposition [157-
159], sol-gel [160] and hydrothermal method [161, 162]. For example, Yang et al. [153] produced hydroxyapatite (HA) coatings on AZ31 alloy by immersing the AZ31 substrate in supersaturated calcification solutions and applying heat treatments afterwards. Song et al. [159] produced HA coatings on AZ91 alloy by electrodeposition with subsequent treatment in NaOH solution. Ramin et al. [160] applied nanostructured HA coatings on 45
AZ91 alloy by a sol-gel derived dip coating technique followed by prolonged aging. Liu et al. [143] produced HA/DCPD (Dicalcium Phosphate Dihydrate) coatings on micro-arc oxidized (MAO) Mg substrates by combining MAO and chemical precipitation. Among these techniques, chemical precipitation is the most popular method to deposit hydroxyapatite coatings on Mg alloys with complex shapes. Nevertheless, the disadvantages of this method are also evident. After a series of pre-coating treatment like heat treatment and passivation in alkaline solution, immersion in supersaturated calcification solutions takes at least 24 hours [152]. Depending on Ca/P ratios, temperature and solution pH, phase compositions in coating showed huge variations [154]. Besides that, non-uniform coatings were formed due to the accumulated hydrogen bubbles on the substrate surface and Mg2+ ions released from substrates during biomimetic coating process, resulting in inhibition of CaP crystallization [150]. For these reasons, it is important to develop novel coating techniques with shorter immersion time scales and without compromising the bioactivity, biocompatibility, corrosion resistance and mechanical properties of Mg alloys.
This paper represents our attempt in coating Mg-alloys with a calcium phosphate phase with the objective to slow down the degradation rate while simultaneously making the surface more biocompatible. Over many years, our group has developed coating technology including biomimetic coating to coat various substrates to render them more bioactive [115, 124, 146, 163-167]. Our contributions to the field include the development of special ‘Simulated Body Fluid’, referred to a t-SBF. This t-SBF was mainly used in coating Ti-alloy with Calcium Deficient Carbonated Hydroxyapatite (CDHA). We have developed compositions to accelerate the process [163]. Also, we developed drug 46
containing CDHA coatings [146]. More recently we have used a microwave assisted coating process to accelerate the coating deposition kinetics [115]. With a microwave- assisted process, the coating deposition can take place within a very short time and was successfully used on polymer and Ti substrates [115, 124]. After microwave assisted surface modification, implants showed elevated cytocompatibility and bioactivity [115,
124].
AZ31 alloy, an important class of Mg alloy, was selected as the substrate in this work.
AZ31 is not only a low-cost Mg alloy with outstanding mechanical properties and corrosion resistance, but also a classic template used for coating studies in many previous papers
[139, 152, 156]. Thus, an easy comparison can be made. This study aims to fabricate hydroxyapatite coating (HA/CDHA) on AZ31 alloy through microwave irradiation of CaP solution with various Ca/P ratios. The morphology and compositions of the coating layer were characterized before and after t-SBF immersion, and the corrosion behavior and cytocompatibility were evaluated.
3.3 Materials and methods
3.3.1 Material preparation
Each AZ31 (3% Al, 1% Zn) disk with Ø 25 × 2 mm was cut equally into four quarters.
Sample surfaces were sequentially polished with finer grit-sized SiC paper up to #1200 grit to achieve homogeneous roughness, then ultrasonically rinsed in ethanol and distilled water for 5 minutes respectively. After the cleaning process, samples were dried at 60 °C for 30 minutes and cooled to room temperature.
3.3.2 Coating preparation 47
Two sets of CaP coating solution were prepared by following compositions as shown in Table 3.1 [156]. Two quartered samples were placed in one 200 ml beaker, filled with
100 ml coating solution. The top of the beaker was covered by an alumina fiberboard
(Zircar Ceramic, Florida, NY). Afterwards, the entire set up was moved into a 1200 W microwave oven (Panasonic) and irradiated at maximum power for 4 minutes. The above steps were repeated once more to ensure the uniformity of the coatings. The Fig. 3-1 illustrated the setup of microwave assisted coating process. Finally, the coated samples were dried in a 100 °C oven.
Table 3.1 Compositions of coating solutions.
H O Ca(NO ) 4H O NaH PO NaHCO Ca/P 2 3 2 2 2 4 3 CS1 200ml 0.6612g 0.2016g 0.0672g 1.67 CS2 200ml 0.6612g 0.1344g 0.0672g 2.50
Figure. 3-1. Schematic diagram of set up of microwave assisted coating process.
3.3.3 Coating characterization
The phase composition of as-coated samples was identified by X-ray diffraction (XRD,
Ultima III, Rigaku)) with monochromated Cu Kα radiation (44KV, 40mA) over a 2θ range
48
of 20–80°. Surface morphology and elemental composition were analyzed using scanning electron microscopy (SEM, S4800, Hitachi) equipped with an energy dispersive X-ray spectroscopy (EDS, Oxford INCA). Elemental analysis was performed at 15 KV with a working distance of 15mm. Functional groups present in the coatings were characterized by a Fourier Transform Infrared Spectroscopy (FTIR, UMA-600 Microscope, Varian
Excalibur Series) using an ATR diamond crystal for 256 scans in the range between 4000-
700 cm-1 with a resolution of 1cm-1.
3.3.4 Electrochemical test
To evaluate the corrosion behavior of both coated and uncoated samples in simulated body fluid (t-SBF), potentiodynamic polarization tests were conducted using Gamry
Reference 600 potentiostat. t-SBF was prepared following the previously reported recipe which better mimics human blood plasma. Ion concentrations are listed in table 3.2 [164].
Meanwhile, a conventional three-electrode cell system comprising the sample with an exposed area of 1.2 cm2 as working electrode, Ag/AgCl electrode as reference electrode, and platinum wire as counter electrode was employed in this study. The potentiodynamic polarization curves were measured at a scan rate of 1mV/s after a steady open circuit potential (OCP) was reached.
Table 3.2 Ion concentrations of the SBF and human blood plasma.
Ion Concentration (mM) Human blood plasma Simulated body fluid (pH 7.4)
Na+ 142.0 142.0 K+ 5.0 5.0 Mg2+ 1.5 1.5
49
Ca2+ 2.5 2.5 Cl- 103.0 125.0 - HCO3 27.0 27.0 2- HPO4 1.0 1.0 2- SO4 0.5 0.5
3.3.5 Immersion test
Degradation behavior of uncoated and coated samples was evaluated by immersion tests in simulated body fluid (t-SBF) according to ASTM-G31-72. Each sample was kept in 60 ml t-SBF in a thermostatic water bath at 37 ± 0.5 °C for 7 days. In addition, t-SBF solution was refreshed every other day. After soaking for a given amount of time, samples were rinsed with flowing distilled water to remove corrosion products and then dried in air.
Weight loss was calculated using following formula at 4 time points (1, 3, 5 and 7 days) to evaluate the degradation rate.
푚0−푚1 Weight loss= 퐴
Where 푚0 is the weight of sample before immersion test (mg), 푚1 is the weight of sample after certain time of immersion in SBF (mg), and 퐴 is the exposed area of sample (cm2).
Post-immersion samples were also characterized by scanning electron microscopy (SEM) coupled with an energy dispersive spectroscopy (EDS).
3.3.6 Cytotoxicity test
Cytotoxicity tests were carried out by indirect assay according to ISO 10993-5:1999.
MC3T3-E1 (CRL-2593™, ATCC, Manassas, VA, USA) preosteoblast cells were initially cultured in alpha minimum essential medium (α-MEM, Thermo Scientific HyClone),
50
augmented with 10% Fetal Bovine Serum (FBS, Thermo Scientific HyClone) at 37 °C in a humidified atmosphere of 5% CO2. The culture medium was replenished every other day until the cells reached a confluence of 90%. Extracts were prepared after 24h incubation at
2 37 °C and 5% CO2 with a surface area to extract medium volume ration of 1.5 cm /ml. α-
MEM culture medium was employed as the negative control. For the cytotoxicity test,
MC3T3-E1 cells were seeded in a 96-well cell culture plate (Flacon™ BD Biosciences, USA) at a density of 5000 cells/well for 24h to allow attachment. After that, the culture medium was replaced by 100μl extracts or 100μl α-MEM. Cellular viability (lactate dehydrogenase,
LDH) was measured after 1 day, 3 days and 5 days using a CytoTox 96® Non-Radioactive
Cytotoxicity Assay kit (Promega, USA).
3.3.7 Statistical analysis
All test results represented means ± SD performed at least in triplicate. One-way analysis of variance (one way-ANOVA) with turkey test was conducted to determine the statistical difference between groups and p < 0.05 was considered significant.
3.4 Results
3.4.1 Coating characterization
X-ray diffraction patterns of HA coated and uncoated substrates are shown in Fig. 3-2.
After the microwave assisted coating treatment, new peaks corresponding to HA (PDF No.
09-432) confirm formation of crystallized HA coating on the AZ31 substrate. In addition, the relative intensity of the diffraction peaks of Mg reduced significantly due to this surface modification. Compared with standard peaks of HA (PDF No. 09-432), the diffraction peak
51
originated from the (002)HA plane shows higher relative intensity, which means that [002] is the preferred orientation for HA to grow along. The differences in intensity between CS1 and CS2 coatings indicate higher crystallinity was achieved with the CS1 coating. Fig.3-3 shows the FT-IR spectra of both the CS1 coating and the CS2 coating on the AZ31 substrate. Hydroxyl (OH) group stretching is observed at 3556 cm-1 and 3550 cm-1 in the spectra of CS1 and CS2 coated AZ31, respectively. However, these weak peaks are mixed
-1 3- with broad bands of absorbed water (3000-3700 cm ). The phosphate (PO4 ) band
-1 3- (asymmetric stretching ν3 mode) is centered at 1030 cm . Moreover, the phosphate (PO4
) band consists of several visible peaks and shoulders indicating the formation of crystalline
HA [168]. Due to asymmetric stretching ν3 and out of plane bending ν2 vibration mode,
2- -1 carbonate brand (CO3 ) is present at 1473, 1446 and 856 cm . The presence of carbonate
2- group in IR spectra is either surface absorption or incorporation of CO3 in HA structure.
2- - 3- It has been demonstrated that CO3 can substitute for both OH and PO4 in the apatite
2- - - 2- lattice. Besides CO3 substituting for OH , OH vacancy left by substitution between CO3
3- and PO4 may contribute to the low intensity of the hydroxyl (OH) band in IR spectra
[169].
Figure. 3-2. XRD patterns of bare AZ31, CS1 coated AZ31 and CS2 coated AZ31.
52
Figure. 3-3. FTIR spectra of CS1 coated AZ31 and CS2 coated AZ31.
Fig. 3-4 presents the surface morphology and elemental composition of the as- deposited coatings on AZ31 specimens. Neither the substrate nor cracks were visible in
SEM micrographs, demonstrating that the coatings had outstanding uniformity. CS1 and
CS2 coatings exhibited different surface features. As shown in Fig. 3-4 (a) and (c), the
CS1-treated sample is densely covered with irregular rod-like crystals. The CS2 coating
(Fig. 3-4 b and 3d) showed a tabular structure comprised of flake-like crystals measuring approximately 2-4 microns in length. It may also be noticed that plenty of acicular spheres were present on both coating surfaces. These spherical precipitations are mostly considered as nascent crystal nuclei [157]. The rods/flakes were formed on the Mg surface, while these microspheres may be formed in solution and bonded to the coating layer. Therefore, as seen in SEM, the CS1 coating has many more spheres on the surface than the CS2 coating, indicating either better bioactivity or the ability to absorb ions from solution to form a new layer on surface. The EDS spectra revealed Ca, P and O were main elements present in both coatings. Additionally, the Ca/P ratios were calculated to be 1.30 for the CS1 coating and 1.46 for the CS2 coating. The lower Ca/P ratio indicated the formation of Calcium
Deficient Hydroxapatite (CDHA), as opposed to Stoichiometric Hydroxyapatite (Ca/P
1.67). This can be result of Mg2+, Na+ ions substituting for Ca2+ and the incorporation of 53
2- CO3 in the HA structure. Cross-sectional images of coatings are shown in Fig. 3-5. The average coating thickness of both coatings is around 5μm. The CS1 coating (Fig. 3-5a) showed inner rod-like CDHA crystals were generally perpendicular to the substrate surface. Conversely, outer CDHA crystals were randomly oriented. As shown in Fig. 3-
5b, flake-like CDHA crystals piled up to form the CS2 coating on the AZ31 substrate. In addition, the absence of gaps between the coating and substrate implies that the coating was well integrated with the substrate.
Figure. 3-4. SEM images and respective EDS analysis of (a,c) CS1 coated and (b, d) CS2
coated samples.
54
Figure. 3-5. Crosssectional morphology of (a) CS1 coating and (b) CS2 coating
3.4.2 Electrochemical behavior
Representative polarization curves of both bare and as-coated AZ31 magnesium alloys are shown in Fig. 3-6 and corresponding potentiodynamic parameters are summarized in
Table 3.3. Compared to bare AZ31 sample, the corrosion potential (Ecorr) of CS2 coated sample shifted to more positive potential with an increase over 100 mV. The corrosion
-5 2 current density (icorr) of CS2 coated sample was suppressed to about 7.9 ×10 A/cm , which is about one order of magnitude lower than that of un-treated AZ31 magnesium alloy. This result indicates CS2 CDHA coating obviously enhanced corrosion resistance of AZ31 magnesium alloys in t-SBF. In comparison with CS2 coated sample, the corrosion potential
(Ecorr) of CS1 coated sample was increased by about 15 mV, while current density (icorr) was reduced by about 50 μA/cm2. Above results demonstrate CS1 coating is more effective in blocking the infiltration of aggressive ions to AZ31 substrate, therefore resulting in better corrosion protection.
55
Figure. 3-6. Potentiodynamic polarization curves of un-treated, CS1 coated and CS2
coated AZ31 alloy samples tested in SBF.
Table 3.3 Electrochemical parameters of samples obtained from polarization curves.
Sample Corrosion potential, Ecorr Corrosion current density, 2 (VAg/AgCl) icorr (A/cm )
AZ31 -1.507 6.72×10-4
AZ31-CS1 -1.372 2.41×10-5
AZ31-CS2 -1.388 7.93×10-5
3.4.3 Degradation behavior
Fig. 3-7 illustrates the corrosion rate of bare and the CDHA coated AZ31 samples in t-
SBF after an immersion time of 1 to 7 days. The values are given in terms of mass loss in every square centimeter. The weight loss of the un-coated AZ31 sample increased sharply in first 2 days, and then changed to be moderate in following days. CS1 coated and CS2 coated samples also presented similar trends in weight change during immersion test.
However, in comparison with un-coated sample, the corrosion rate of CDHA coated
56
samples decreased significantly at each time point. After 7 days of immersion, the mass loss of the un-coated sample was 9 times that of the CS1 coated sample and 3 times that of the CS2 coated sample. It could be speculated that CDHA coating provided effective protection for the AZ31 alloy against rapid corrosion in SBF.
Figure. 3-7. Variance of weight loss of coated sample and non-coated sample with
immersion time in t-SBF solution
After soaking in t-SBF for 7 days, post-immersion samples were characterized by SEM and EDS, as shown in Fig. 3-8. It can clearly be seen that a dense layer consisting of spherical particles precipitated on the CS1 coating, from Fig. 3-8a. EDS spectra reveals that the newly deposited layer was rich in Ca, P and O. This demonstrates that the CS1 coating was covered with a dense apatite layer during the SBF incubation. It needs to be pointed out that this dense apatite layer can increase the coating thickness and enhance corrosion resistance of AZ31 in vitro and in vivo. Similar apatite precipitation was also observed on the CS2 coated sample (Fig. 3-8b) indicating good bioactivity of both types of coatings. Compared with the CS2 post-immersion sample, there are no obvious cracks
57
present on the CS1 post-immersion sample. This corroborated the results of weight loss discussed above as the CS1 coating performed better for corrosion protection in vitro.
Figure. 3-8. SEM images and respective EDS analysis of (a) CS1 coated and (b) CS2
coated samples after 7 days’ immersion in SBF solution
3.4.4 Cytotoxicity
Fig. 3-9 shows the cell proliferation data after incubating in the culture medium and various extracts for 1, 3 and 5 days. The results are presented as optical density readings
(OD490), which were assumed to be proportional to the number of cells. No statistical differences were observed between coating extracts and culture medium after 24 hours incubation. At day 3 and day 5, all 3 groups showed statistically significant cellular proliferation (p < 0.05). It is clear that cells cultured with coating extracts showed the same growth trend as cells cultured with α-MEM media. This means the CS1 coated sample and the CS2 coated sample both have good cytocompatibility. Although no statistical differences were found (p > 0.05), CS1 coating apparently exhibited higher cell
58
proliferation than CS2 coating and control at all time intervals. Fig. 3-10 shows optical microscopy morphologies of osteoblast cells cultured in α-MEM media and different coating extracts. It can be seen that number of cells increased with incubation time, and cell morphologies in extracts of coated samples were normal and healthy, similar to that of cells incubated in culture media.
Figure. 3-9 O.D. values of MC3T3-E1 cells seeded in extracts of coated samples and
culture medium for 1, 3 and 5 days.
Figure. 3-10 Optical micrographs of osteoblast cells (MC3T3-E1) after 3 days and 5 days
incubation: (a, d) the negative control, (b, e) extracts of AZ31-CS1 coated,
and (c, f) extracts of AZ31-CS2 coated. Scale bar is 100 μm. 59
3.5. Discussion
Both XRD and FTIR spectra confirmed the formation of a HA coating on the AZ31 substrate through novel microwave assisted coating treatment with various initial Ca/P ratios. Moreover, no obvious differences were observed from the spectra of CS1 coating and CS2 coating, which implies the consistency of this microwave treatment.
SEM images of CS1 and CS2 coatings clearly suggested substrates were coated with a dense CDHA layer. The formation mechanism of HA coating on Mg alloy substrates has been widely studied [156, 161, 170]. It is known that higher pH environment and relatively high temperature can facilitate the nucleation and precipitation of HA [161, 170, 171]. The growth mechanism is illustrated in Fig. 3-11 and summarized as following: corrosion reaction occurs immediately after immersing AZ31 substrate into a supersaturated coating solution. As a result, hydroxyl ions (OH-) accumulate on the surface leading to the formation of Mg(OH)2 as a passive layer on the sample surface. Simultaneously, on account of the sudden increase of pH on substrate surface, heterogeneous nucleation of HA is promoted significantly [161, 170]. In addition, microwave assisted coating treatment brought out several unique phenomena, beneficial for precipitation of HA crystals. For
2+ 3- instance, the concentration of Ca , PO4 and local pH value increased expeditiously due to dielectric heating by the absorption of microwave energy in water. Nucleation of HA was accelerated because of the rapid rise in local pH value. The abundant nascent nuclei scattered on the as-coated sample surface revealed the high nucleation rate achieved in this treatment. Furthermore, plentiful calcium and phosphate ions gathered around the sample surface can advance further crystallization and growth of HA. In a chemical precipitation coating of magnesium alloy such as biomimetic coating, hydrogen bubbles that accumulate 60
on the surface of the magnesium alloy during the coating process can result in the formation of cracks, pores and some other undesirable defects in the coating structure. However, due to the violent movement of liquids caused by microwave irradiation, hydrogen bubbles are not able to attach to the sample surface during this coating process. Additionally, it is unlikely that, as in the case for conventional apatite, the coating initially deposits around small cracks and then expand to whole surface area; numerous HA nuclei first spread over the sample surface through the microwave assisted coating technology [115, 172]. Thus, continuous anodic dissolution of Mg can be retarded, and the resulting evolution of hydrogen gas is suppressed during the early stage of this microwave assisted coating process. Yanovska et al. [154] reported that HA coating deposited under a certain magnetic field could improve the crystallinity of HA with a preferred orientation to the c axis.
Magnetic field distribution in the microwave oven may also play an important role to the deposition of crystallized HA coating on the AZ31 alloy. These advantages enable this novel microwave assisted coating technology to homogeneously produce a crack-free and dense HA coating on AZ31 alloy in minutes. Moreover, this technique allows us to reduce immersion time to a few minutes rather than the many hours reported in available HA coating methods [151, 152, 156, 160].
Quantitative EDS results showed that the Calcium to Phosphate ratio of the CS1 coating and the CS2 coating was below that of Stoichiometric HA. This is because both Mg2+ and
+ 2+ 2- Na can substitute for Ca in the HA lattice, while CO3 can simultaneously substitute for
3- - PO4 or OH [169]. Theoretical calculations suggested precipitation of CDHA
2+ + 2- incorporated with Mg , Na , CO3 is more favorable than Stoichiometric HA [171]. These substitutions might also attribute to the enhanced diffusion of ions between the HA and 61
solution by microwave irradiation. Though the incorporation of ionic substitutions into HA can distort the HA lattice and reduce the crystallinity of HA, these minor species not only have the advantage of mimicking natural bone tissue but also have a helpful impact on the biological response of bone cells such as inducing growth of bone-like apatite, promoting new bone remodeling and stimulating osteoblast proliferation [173].
Upon immersion in a physiological solution like SBF, anodic dissolution of Mg takes place instantly, as shown in reaction (1). Following this reaction, Mg(OH)2 precipitates on the AZ31 sample surface with the simultaneous release of H2 gas.
Mg→Mg2+ +2e-,
- - 2H2O+2e →H2+2(OH) , (1)
2+ - Mg + 2(OH) → Mg(OH)2 .
Due to its low solubility in water, Mg(OH)2 can act as a protective layer against further anodic dissolution of magnesium [174]. However, the presence of chloride ions (Cl-) can transform Mg(OH)2 into more soluble MgCl2, following reaction (2).
2+ − Mg + 2Cl → MgCl2 (2)
The aggressive attack by corrosive mediums like chloride ions can cause rapid dissolution of Mg(OH)2 and that dissolution is faster than the formation of the new protective layer.
These corrosive species then preferentially aggregate and attack the exposed Mg alloy substrate, which results in pitting corrosion and rapid degradation of Mg alloy [175]. The more severely AZ31 corrodes, the more weight loss is expected. Therefore, the coating’s protective effect was evaluated by comparing weight loss of the coated sample and non- coated sample in t-SBF for 7 days. The significantly (p < 0.05) less weight loss and low corrosion current achieved by both CS1 coated and CS2 coated samples demonstrates the 62
as-deposited CDHA coating effectively protected the AZ31 substrate from fast degradation in t-SBF. It is noteworthy that the CS1 coated sample showed greater corrosion resistance than the CS2 coated sample. In addition, the CS1 coated sample maintained good integrity after 7 days immersion, whereas micro cracking was observed in the CS2 coated post- immersion sample. This may be explained by the following: 1) the pileup structure in the
CS2 coating is not dense enough, and 2) the adhesion strength of the coating is not as strong as that of the CS1 coating. From the cross-section micrograph of the CS2 coating, it is seen that part of the coating was peeled off during the polishing process. This implies that the
CS2 coating was not effectively integrated with the AZ31 substrate. In contrast, rod like
CDHA crystals were closely packed, especially at the region near the substrate surface as presented in the CS1 coated sample. This compact structure in the CS1 coating definitely provided more protection against corrosive media. CDHA coating not only improves corrosion resistance but also promotes biomineralization in physiological environment and enhances the bonding between implants and fractured bones [145]. After 7 days immersion in t-SBF, the coated samples were covered by numerous precipitates of bone-like apatite, offering the excellent bioactivity of CDHA coating. Because of relative weakness in corrosion resistance, precipitates of bone-like apatite on the CS2 coated sample were not as uniform as the apatite layer on the CS1 coated sample. As mentioned before, microwave assisted coating treatment also induced plenty of nascent nuclei on top of the coating. Ca2+
3- and PO4 ions were continuously supplied to the coating surface in SBF solution. Thus, bone-like apatite continuously grew from these nuclei with increasing immersion time in physiological solution. Moreover, especially in CS1 coating, the rod-like structure with large surface area, can amplify the ion diffusion with body fluid, which is beneficial for 63
nucleation of apatite. Whereas the close-packed inner layer will block the severe corrosive attacks caused by aggressive ions such as Cl-. In other words, the outer layer comprising scattered rods had much larger surface area than crowded inner layer in CS1 coating.
Therefore, this kind of coating structure can effectively improve the bioactivities and corrosion resistance of magnesium alloy substrate.
Preliminary cytocompatibility assessment is an important assay to evaluate biological properties of biomaterials prior to in vivo evaluation. Introducing culture medium onto the
Mg alloy surface can result in strong color change of the medium, which indicates huge pH increase of the medium. Significantly low cell density was observed by culturing HeLa cells in medium with a high pH value [176]. In this work, coated samples were incubated in α-MEM culture medium for 24 hours before performing the cytotoxicity assay. There was no unusual color change in the medium after 24 hours incubation. This means the
CDHA coated samples present good corrosion resistance at an early stage of culturing.
Cytotoxicity assay after 5 days demonstrated that both CS1 and CS2 coatings promoted the proliferation of preosteoblast cells. This is because the ionic dissolutions from CDHA coatings played a role as a growth factor in the differentiation of osteoblasts and expedited new bone formation [173, 177].
64
Figure. 3-11 Schematic diagrams of microwave assisted formation mechanism of CDHA
coatings on AZ31 alloy.
3.6 Conclusions
Biodegradable CDHA coating was fabricated onto AZ31 substrates using a novel microwave assisted coating technique in minutes. Surface characterizations by XRD and
FTIR confirmed the presence of CDHA coating on the sample surface. CDHA coatings produced in aqueous solutions with various Ca/P ratios showed different microstructures: the CS1 coating was composed of rod-like CDHA crystals that were mostly perpendicular to the AZ31 substrate; the CS2 coating maintained a pileup structure with flake-like CDHA crystals. Results of in vitro immersion assessment revealed that both CS1 and CS2 coated samples showed significantly improved corrosion resistance and excellent bioactivities.
After 7 days immersion in SBF, a dense layer of apatite precipitated on the CS1 coated sample surface, which is helpful for protecting integrity of the AZ31 substrate and 65
supporting new bone growth. Cell culture assays demonstrated that the CDHA coated
AZ31 alloy possessed good cytocompatibility and promoted cellular proliferation after 5 days incubation. The present study suggests that microwave assisted coating treatment can rapidly induce the deposition of a uniform protective CDHA layer on the AZ31 surface.
Furthermore, CDHA coated samples present enhanced ability in corrosion resistance and biological response to meet the requirements for application as orthopedic implants.
66
Chapter 4
Microwave Assisted Magnesium Phosphate Coating on AZ31 Magnesium Alloy
4.1 Abstract
Due to the combination of many unique properties, magnesium alloys have been widely recognized as a suitable metallic material for fabricating degradable biomedical implants.
However, extremely high degradation kinetics of magnesium alloys in physiological environment have hindered their clinical applications. This paper reports for the first time the use of a novel microwave assisted coating process to deposit magnesium phosphate coatings on Mg alloy AZ31 and improve its the in vitro corrosion resistance. Newberyite and trimagnesium phosphate hydrate (TMP) layers with distinct features were fabricated at various processing times and temperatures. Subsequently, the corrosion resistance, degradation behavior, bioactivity and cytocompatibility of the magnesium phosphate coated AZ31 samples were investigated. The potentiodynamic polarization tests reveal that the corrosion current density of AZ31 magnesium alloy in SBF is significantly suppressed by the deposited magnesium phosphate coatings. Additionally, it is seen that magnesium
67
phosphate coatings remarkably reduced the mass loss of AZ31 alloy after immersion in
SBF for two weeks and promoted precipitation of apatite particles. The high viability of preosteoblast cells cultured with extracts of coated samples, indicates that the magnesium phosphate coatings can improve the cytocompatibility of AZ31 alloy. These attractive results suggest that magnesium phosphate coatings, serving as the protective and bioactive layer can enhance the corrosion resistance and biological response of magnesium alloys.
4.2 Introduction
Magnesium based biodegradable alloys are widely recognized as implant materials due to their combinations of properties such as low Young’s modulus, high strength/weight ratio, and excellent biodegradability and biocompatibility [125-127, 178, 179]. However, bare magnesium alloys have a fast kinetic of degradation in physiological environment, which would subsequently trigger the early mechanical failure of magnesium implants, formation of gas cavities and many other detrimental events [132, 180].
To date, various surface modification techniques have been developed to tailor the corrosion rates of magnesium alloys, extend the mechanical integrity of magnesium based implants in vivo and promote their biological responses [132, 144]. Among the protective coatings, calcium phosphate (CaP), especially hydroxyapatite (HA, Ca10(PO4)6(OH)2) is one of the most favorable coating compound, since it does not only retard the rapid degradation of magnesium alloys, but also promotes the direct bonding between gap of the implants and host tissues [129, 144].
This paper combines our research expertise on magnesium based biomaterials and microwave assisted coating technology. Magnesium, being in the same group (alkaline 68
earth) of calcium in the periodic table, presents a great deal of chemical similarities to calcium. In addition, magnesium is the main substitute of calcium in body minerals [181].
The presence of magnesium phosphate (MgP) in physiological and pathological mineralized tissues also demonstrates the outstanding biocompatibility of the magnesium phosphate compounds [108, 182]. Various compositions of magnesium phosphates are gaining great momentum in the development of newer bioceramic for orthopedic applications, on account of their better mechanical properties, fast degradation kinetics, and comparable biocompatibility to calcium phosphate compounds [109, 181, 183-187].
In contrast to calcium phosphate coatings on magnesium alloys which have been extensively studied, the corrosion resistance, bioactivity and biocompatibility of magnesium phosphate coatings have not been systematically explored. Additionally, the deposition method of magnesium phosphate coating has mostly been limited to phosphate conversion coating. For instance, single phase newberyite (MgHPO4•3H2O) coatings can be produced by immersing magnesium alloy substrate in phosphate solutions for 3 days
[188, 189]. Zhao et al. [190] proposed a two-step method to deposit two distinct layers of newberyite and struvite on AZ31 magnesium alloy for enhanced corrosion resistance.
Fouladi et al. studied the effect of phosphating temperature and time on the microstructure and corrosion resistance of magnesium phosphate coatings The results demonstrate that dense newberyite coatings remarkably enhance the corrosion resistance of stainless steel substrates [191]. Morks developed a novel magnesium phosphate (bobierrite) coating on low carbon steels for improved anti-corrosion properties [192]. It’s also revealed that the magnesium phosphate coating on carbon steel maintains a better corrosion resistance than zinc phosphate coatings [193]. 69
The second aspect of this work is the application of a microwave assisted coating process, which is a low-cost technique with the benefits including high efficiency, energy saving and relatively low coating formation temperature. Our previous efforts have shown that microwave energy can significantly promote the deposition kinetics of calcium phosphate coating, thereby shorten the coating process to minutes instead of hours [115,
119]. Moreover, due to the nature of wet-chemical precipitation, the microwave assisted coating technique is applicable to various biomedical implant materials even with complex structures.
Therefore, the aim of this study is to use a microwave assisted method to coat AZ31 magnesium alloy with magnesium phosphate within a short time scale. With the incorporation of the magnesium phosphate coating, the corrosion resistance and bioactivity of magnesium alloys will be significantly improved.
4.3 Experimental
4.3.1 Material preparation
Commercially available AZ31 magnesium alloy (3% Aluminium, 1% Zinc) blocks
(1cm 1cm 2mm) were ground with abrasive SiC paper up to #1200 grit to ensure the uniform surface roughness, then ultrasonically cleaned in ethanol and DI (deionized) water for 10 min respectively. Afterwards, samples were dried at 60 °C for 30 min and cooled down in air.
4.3.2 Coating preparation
The coating bath was prepared by dissolving 2.564g Mg(NO3)2·6H2O, 0.7198g
NaH2PO4 in 100ml DI water. The pH value of the coating bath was adjusted to 5.6 by 1M 70
NaOH solution. 50ml coating solution with AZ31 substrates were poured into a reaction vessel with a tight lid. Next, the vessel was placed in a microwave system (MARS 230/60,
CEM) and heated by microwave for 10 min and 30 min at 80°C, 120°C and 160°C respectively. Finally, the AZ31 substrates were rinsed with DI water and dried overnight in a 60°C oven.
4.3.3 Coating characterization
The phase compositions of as-deposited coatings were identified by X-ray diffractometer (XRD, Ultima III, Rigaku) using monochromated Cu Kα radiation (44KV,
40mA). The diffraction patterns were measured over a 2θ range of 10–80° with a step size of 0.05°/step. The surface, cross-sectional morphologies and chemical compositions were analyzed using scanning electron microscopy (SEM, S4800, Hitachi) coupled with an energy dispersive X-ray spectroscopy (EDS, Oxford INCA). The elemental analyses were determined with accelerating voltage of 20kV and the working distance of 15mm. Fourier
Transform Infrared Spectroscopy (FTIR, UMA-600 Microscope, Varian Excalibur Series, resolution 1cm-1) in ATR geometry was employed to investigate functional groups present in deposited coatings. The spectra were collected by 256 scans in a wavenumber range of
4000-700 cm-1.
4.3.4 Electrochemical test
Potentiodynamic polarization tests were conducted in order to evaluate the corrosion resistance of the magnesium phosphate coated AZ31 samples in simulated body fluid (t-
SBF), which was prepared according the recipe reported by Tas et al. [164]. All the electrochemical measurements were carried out using Gamry Reference 600 potentiostat system and Echem Analyst software for data acquisition and analysis respectively. In 71
addition, a conventional three-electrode cell system comprising the modified sample as working electrode, Ag/AgCl as reference electrode, and graphite rod as counter electrode was employed in this study. Before the potentiodynamic polarization curves were measured with a scan rate of 1mV/s, the open circuit potential (OCP) was set to stabilize for 30 min.
4.3.5 Immersion test
The in vitro degradation behaviors of MgP coated samples were evaluated by an immersion test according to ASTM-G31-72. Each sample was kept in 60 ml t-SBF in a thermostatic water bath at 37 ± 0.5 °C for 15 days. Additionally, t-SBF liquid was replenished every other day. The pH value of SBF solution was recorded during the immersion test. After a certain time of soaking, samples were immersed in chromic acid
(200 g/L CrO3 and 10 g/L AgNO3) for 5 min and rinsed with DI water flow to remove the newly formed corrosion products. The weight losses of samples were measured at each time point to arrive at the degradation rate. The degradation rate, presenting the relative percentage of mass loss was derived from following relation:
푊1−푊0 푊퐿 = − ( ) ×100 (1) 푊0
where WL, W0, and W1 are the percentage of weight loss, and the weights before and after immersion for different time intervals, respectively. Post-immersion specimens were also characterized by SEM and EDS to evaluate their abilities to induce biomineralization.
4.3.6 Cytotoxicity test
The surface cytotoxicity properties of bare and MgP coated AZ31samples were tested by indirect assay according to ISO 10993-5:1999. Osteoblast-like cells MC3T3-E1 (CRL-
72
2593™, ATCC, Manassas, VA, USA) were initially cultured in alpha minimum essential medium (α-MEM, Thermo Scientific HyClone) supplemented using 10% Fetal Bovine
Serum (FBS, Thermo Scientific HyClone) with a humidified atmosphere of 5% CO2 at
37°C. The cells were sub-cultured every two days to reach a confluence of 90%. Extracts were prepared by incubating samples in α-MEM for 24h at 37 °C and 5% CO2 with the ratio of surface area to extract medium volume set to be 1.5 cm2/ml. The complete α-MEM culture medium without specimen immersed acted as a control group. For cytotoxicity test,
MC3T3-E1 cells were seeded in 96-wells cell culture plate (Flacon™ BD Biosciences, USA) at a density of 5000 /well and incubated for 24h to allow attachment. Then the culture medium was then replaced by 100μl sample extracts. After further incubation for 3 days, the cells were treated with 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide
(MTT, Sigma-Aldrich, St. Louis, MO, USA) for 4 h. The formazan precipitate was dissolved in dimethyl sulfoxide (DMSO) and measured using a microplate reader at a wavelength of 570 nm to assess live cells.
4.3.7 Statistical analysis
All test results were expressed as means ± SD with at least in triplicate. The statistical differences between groups were determined by one-way analysis of variance (one way-
ANOVA) with turkey test and ρ< 0.05 is considered to be significant.
73
4.4 Results and discussion
4.4.1 Coating characterization
The crystal phases and compositions of the magnesium phosphate coatings prepared via the microwave assisted process were characterized using XRD and FTIR. Fig. 4-1 presents X-ray diffraction patterns of bare AZ31 substrate and the substrates treated with different temperatures and coating periods. The diffraction patterns of all the treated samples exhibit sharp diffraction peaks, indicating the relative high crystallinity of the as- deposited magnesium phosphate coatings. Previous studies have demonstrated the capability of microwave assisted process to synthesize amorphous magnesium phosphate
(AMP) nano-particles [109, 187]. Conversely, the formation of highly crystallized magnesium phosphate precipitates can be attributed to the slightly acidic condition in current treatment process. Boistelle reported the precipitation of newberyite
(MgHPO4·3H2O) can only occur when the surrounding pH is lower than 5.8 [194].
Regarding each sample, the diffraction peaks of the coating prepared at 80 °C and 120 °C correspond to the newberyite (MgHPO4·3H2O). This suggests the coating is a single-phase layer of newberyite. In contrast to coatings prepared at a lower temperature, coatings produced at 160°C showed biphasic structure consisting with both newberyite (minor phase) and trimagnesium phosphate hydrate (Mg3(PO4)2 ·4H2O) phases. It can be speculated that most of the newberyite phase transformed to trimagnesium phosphate hydrate (TMP) when temperature increased to 160°C. The similar temperature induced phase transitions were also reported in the calcination of amorphous magnesium phosphate
(AMP) and the synthesis of biphasic calcium phosphate (BCP) [184, 195, 196]. Moreover, the incongruent dissolution of newberyite can lead to the precipitation of trimagnesium 74
2+ 3- phosphate [197]. At a proper pH and relevant concentrations of Mg and PO4 in aqueous environment, newberyite can partially transform to bobierrite [181, 198].
All six kinds of coatings present very similar IR spectra, as shown in Fig. 4-2. The
3 – -1 absorption peaks attributed to phosphate band (PO4 ) were centered at 1161cm , 1054
-1 -1 3 – -1 cm and 1014 cm . The splitting of phosphate (PO4 ) band in the range of 1000 cm to
1100 cm-1 also implies the crystalline nature of the as-prepared magnesium phosphate coatings [169, 195]. Moreover, the broad absorption peaks at 3225 cm-1 and 1641 cm-1 can be assigned to v1 symmetric stretch and v2 bend of structural water respectively. Overall, the FTIR results are in good agreement with XRD analysis.
Figure. 4-1 XRD patterns of bare AZ31 and MgP coated AZ31 with various
parameters.
75
Figure. 4-2 FTIR spectra of bare AZ31 and MgP coated AZ31 with various
parameters.
The surface features of as-prepared coatings were examined by SEM and shown in Fig.
4-3. The coatings deposited at 80°C and 120°C present comparable morphologies, both coatings comprised of micron sized newberyite precipitates with polyhedral and tabular structures. As shown in Fig. 4-3(a) and (c), a dense layer with uniform-sized newberyite crystals is fabricated on AZ31 sample surfaces in 10 min. This demonstrates the high efficiency of the microwave assisted coating process in accelerating deposition kinetics of magnesium phosphate. The newberyite film produced by long-term phosphating treatment also shows identical microstructure, as compared to newberyite coatings prepared in this work [188, 191]. It is noticed that most of the tabular newberyite precipitates transformed to polyhedral structure following the increase of treatment period and temperature.
Moreover, along with the increase of the processing duration, the size of the deposited newberyite crystals slightly decreased. The diameter of the polyhedral particles prepared
76
at 120°C and 30 min are close to 5m, while particles prepared at 120°C and 10 min maintain a diameter above 10m. The fast dissolution of newberyite in the coating solution could be the reason of shrinkage of the newberyite crystals synthesized with prolonged treatment. In contrast, no significant change in size of the newberyite particles was observed when the temperature elevated from 80 °C to 120°C. It’s revealed that temperature in range of 80-120°C presents minor effects in altering the crystal size of newberyite precipitates. Additionally, the levels of pH and supersaturation in aqueous solution could play an important role in the morphology evolution and growth of newberyite crystals [194, 199]. At 160°C however, the AZ31 substrate are uniformly covered with the precipitates of flower-like trimagnesium phosphate hydrate (TMP). The flower-like structure is assembled by stacking sheets along the [010] [200, 201]. Similarly, the flower-like trimagnesium phosphate hydrate precipitates present smaller size with prolonged coating process. EDS results confirmed the presence of Mg, P and O, which are major elemental constituents of magnesium phosphate. Therefore, the EDS results demonstrate the successfully formation of MgP coatings on AZ31 substrate.
77
Figure. 4-3 SEM images and EDS spectra of as-deposited MgP coating prepared by
(a) 80°C-10min, (b) 80°C-30min, (c) 120°C-10min, (d) 120°C-30min, (e)
160°C-10min, (f) 160°C-10min, (g) 80°C-30min, (h) 120°C-30min and
(i) 160°C-30min.
The cross-section of magnesium phosphate coated AZ31 samples were prepared by focus ion beam and examined using scanning electron microscopy (SEM, FEI Quanta 3D
FEG). Fig. 4-4 shows the cross-sectional images of magnesium phosphate coated samples with various parameters. All of the coatings are closely integrated with AZ31 magnesium substrate implying good adhesion of the magnesium phosphate layers is achieved. It can be seen that coatings prepared at 80°C and 120°C present the identical thickness of 7-9m and 5-7m respectively. The decrease of coating thickness can be related to the finer
78
newberyite crystal size at high temperature. In contrast, the thicknesses of magnesium phosphate coatings produced at 160°C are over 17m. Additionally, less micro cracks are observed on coatings prepared at 160°C which indicates these layers are more compact than the others.
Figure. 4-4 Cross-sectional SEM images of samples prepared by (a) 80°C-10min, (b)
80°C-30min, (c) 120°C-10min, (d) 120°C-30min, (e) 160°C-10min and
(f) 160°C-30min (Some of the microcracks were identified by arrows).
4.4.2 In vitro degradation behavior
4.4.2.1 Electrochemical test
To evaluate the corrosion resistance of bare and magnesium phosphate coated AZ31 samples in physiological environment, a potentiodynamic test was employed. Fig. 4-5 shows the polarization curves of bare and magnesium phosphate coated AZ31 magnesium alloy samples in SBF and the extrapolated electrochemical parameters are listed in Table
4.1. As compared to the untreated AZ31 sample, the corrosion potential (Ecorr) of AZ31 substrates modified by newberyite precipitates shifted by 100-220 mV toward the more positive potentials. However, a dramatic increase in the value of corrosion potential (Ecorr) of trimagnesium phosphate hydrate (TMP) coated AZ31 is observed. This indicates the
79
magnesium phosphate coatings, especially newberyite-trimagnesium phosphate hydrate
(TMP) biphasic coating remarkably suppressed the thermodynamic reaction activity of
AZ31 Mg alloy in SBF. As compared to the bare AZ31 sample, all of the magnesium phosphate coated AZ31 samples exhibit significant reduction in the corrosion current density (icorr). It is well known that the corrosion current density is directly proportional to the corrosion activities. In analogy with other newberyite coated samples, newberyite coating deposited at 80°C for 10 min tend to corrode more easily, which can be ascribed to the rougher surface morphology resulting in an ease of the infiltration of aggressive ions.
Among the coated samples, the AZ31 substrates treated at 160°C show minimum corrosion current density (1.108 μA/cm2) which is almost three orders of magnitude lower than that of non-coated AZ31 substrate. The excellent electrochemical stability of the thick trimagnesium phosphate hydrate (TMP) layer and its compact structure are the main factors promoting its anti-corrosion performance. Taken together, all of these results demonstrate the magnesium phosphate coatings perform well toward corrosion protection.
Figure. 4-5 Potentiodynamic polarization curves of the bare and treated AZ31
samples. 80
Table 4.1 Electrochemical parameters of samples obtained from polarization curves.
2 Sample Ecorr/V Icorr(μA/cm ) Bare AZ31 -1.507 859.0 80°C-10min -1.274 142.7 80°C-30min -1.400 19.94 120°C-10min -1.427 2.143 120°C-30min -1.307 3.745 160°C-10min -0.4673 1.108 160°C-30min -0.4095 1.260
4.4.2.2 Immersion test
The weight losses of the coated and un-coated AZ31 magnesium alloys with various immersion times are shown in Fig. 4-6 (a). In a physiological environment, the corrosive species such as chloride ions (Cl-) can aggregate and attack the Mg alloy substrate, which results in pitting corrosion and rapid degradation of Mg alloy [175]. The more severely the magnesium alloy corrodes, the more weight loss is expected. Therefore, the extent of magnesium phosphate coatings’ protective effects was evaluated by comparing weight losses of the coated and non-coated sample in SBF for 15 days. It can be seen the weight loss increases with the prolongation of immersion period. Furthermore, the weight loss due to anodic dissolution of magnesium alloy in SBF was observed to slow down after the first two days’ immersion. This is because the simultaneously formed Mg(OH)2 and phosphate containing protective layer suppressed the initial burst of corrosion reaction [174, 189]. As compared with the bare AZ31 sample, either the newberyite coated or the newberyite -
81
TMP biphasic coated AZ31 samples present much less mass loss at each time point indicating the significant improvement of corrosion resistance. After two weeks of incubation in SBF, the uncoated AZ31 sample showed a highest weight loss percentage
(9.90.8), whereas the sample treated at 160°C for 10 min presents most minimal weight loss percentage (2.00.2). These results further demonstrate the supreme corrosion resistance achieved by samples coated at 160°C.
As shown in Fig. 4-6 (b), the pH values of SBF solution rapidly increase to 8.3 during the initial immersion of AZ31 magnesium substrate. This is due to the alkalization behavior of AZ31 magnesium alloy in SBF. Conversely, magnesium phosphate coated sample shows slight increase of pH value at the early stage of the immersion and definitely lower pH value. In the second week of immersion test, the pH values of the SBF solution gradually become stable. The stabilization and decrease of pH can be the result of the
- consumption of hydroxyl ions (OH ) in SBF solution due to the precipitation of Mg(OH)2.
During the entire immersion test, the pH of magnesium phosphate coated samples shows less tendency towards elevation and also varies in a more desirable range compared with the bare AZ31 substrate. Especially, the pH value of SBF after incubation with biphasic magnesium phosphate coated sample is about 7.5 after 15 days of incubation, which means sample coated at 160°C presents a more favorable degradation behavior compared to that of uncoated sample. These observations are in good accordance with the results of electrochemical test.
82
Figure. 4-6 Variations of weight loss (a) of samples and pH value (b) of SBF solution
as a function of immersion time.
Figures 4-7 (a)-(f) illustrate the surface morphology and representative EDS spectra of magnesium phosphate coatings prepared with different parameters after immersion in SBF solution. Meanwhile, the phase evolution of magnesium phosphate coatings is examined by XRD and depicted in Fig. 4-8. Due to the fast dissolution of newberyite crystals, several micro-cracks originated on the newberyite coating produced at 80°C, which evidently implies the relatively weak corrosion resistance of the newberyite coating deposited at low 83
temperature. As seen in Figures 4-7 (a) and (b), significant amounts of cavities were formed inside the polyhedral newberyite crystals, which result in hollow structure of newberyite crystals after two weeks of immersion in SBF. Thus, it is anticipated that the extent of dissolution is more at internal resorption sites of the newberyite crystals, compared with those of external surfaces. Klammert et. al also reported the more significant degradation from the interior of the newberyite cement, thereby remarkably enlarging its surface area
[202]. It is important to note that the formation of spherical apatite particles on the coating surface demonstrating bioactivity of the as-deposited newberyite layer. The abundant traces of Ca and P, as dispayed in EDS spectrum (Fig. 4-7g) also confirms the precipitation of calcium phosphate minerals during the immersion in SBF. Moreover, the XRD pattern
(Fig. 4-8) of post-immersion newberyite coating prepared at 80°C presents a single broad diffraction peak at 2θ=30° that can be due to the presence of amorphous magnesium phosphate [184]. Thus, it can be speculated that with the continuous dissolution of the newberyite crystals in SBF, the as-deposited crystalline newberyite coating gradually transforms to the amorphous structure. Interestingly, a totally different phase evolution of the newberyite coating prepared at 120°C is identified. As seen in Fig. 4-8, the newberyite phase completely transforms to trimagnesium phosphate hydrate (TMP) after soaking in
SBF for 15 days. Accordingly, significant changes regarding the morphologies of deposited newberyite coating are expected. As seen in Fig. 4-7 (c) and (d), the micron-sized polyhedral newberyite crystals completely disappeared after the immersion test, whereas the hierarchical flower-like trimagnesium phosphate hydrate (TMP) crystals appeared to be uniformly distributed on AZ31 magnesium alloy surface.
84
Figure. 4-7 SEM images of samples (a) 80°C-10min, (b) 80°C-30min, (c) 120°C-10min,
(d) 120°C-30min, (e) 160°C-10min and (f) 160°C-30min after 2 weeks SBF
immersion.
The morphology evolution is in good accordance with the changes in phase composition. Previous studies have demonstrated the in vitro transformation of amorphous magnesium phosphate and newberyite to bobierrite (Mg3(PO4)2 ·8H2O) [109, 181]. Due to the superior ability of divalent magnesium ion to coordinate water molecules, the trimagnesium phosphate can bind with various molecules of crystallization water and form at least four kinds of crystalline trimagensium phosphate phases [108]. Therefore, the newberyite may not only transform to bobierrite (Mg3(PO4)2 ·8H2O) in physiological environment, but may also transform to the trimagnesium phosphate hydrate
(Mg3(PO4)2·4H2O) with less levels of structural water, as explored in this work. It may be
85
possible that the rapid degradation and phase evolution of newberyite coating deposited at
120°C is due to its small crystal size. On account of its lower solubility in SBF, the trimagnesium phosphate hydrate (TMP) possesses greater stability than newberyite [108].
Thus, it’s anticipated that the TMP coating can provide more effective long-term corrosion resistance. Even after 15 days of immersion, the trimagnesium phosphate hydrate (TMP) layer with the flower-like TMP crystals still maintains identical morphology, which can be clearly distinguished from the newly-precipitated apatite globules. Moreover, the corresponding XRD pattern only shows the diffraction peaks belonging to trimagnesium phosphate hydrate (TMP) layer and magnesium substrate. This indicates complete degradation of newberyite crystals incorporated in original coating structures and the amorphous nature of the newly-formed apatite. However, the quantitative EDS analysis reveals a low concentration of Ca on the coating surface, which may imply the apatite is highly substituted with Mg2+. It is noteworthy that diverse benefits of magnesium substitution in calcium phosphate compounds have been investigated. These include stabilizing amorphous calcium phosphate (ACP) and stimulating the proliferation, differentiation and mineralization of osteoblast cells [16,40,41]. After incubating in SBF for two weeks, the absence of micro-cracks on the magnesium phosphate coatings prepared at 120°C and 160°C presents their protective signatures in physiological condition.
Additionally, the tailored degradation rate is beneficial for magnesium alloy to stimulate mineralization and osseointegration in vivo.
86
Figure. 4-8 XRD patterns of bare AZ31 and MgP coated AZ31 after incubating in
SBF solution for 15 days.
4.4.3 Cytotoxicity
Preliminary cytotoxicity assessment is an important assay to evaluate biological response of biomedical implant materials prior to in vivo evaluation. Fig. 4-9 shows the viability of MC3T3-E1 cells after incubating in the culture medium and various extracts for 3 days. Compared to the control group, the samples treated at 80°C-10min show statistically low (p < 0.05) cell viability, which may be attributed to the excessive hydroxyl
(OH-) ions released in culture medium. It is revealed that the rapid alkaline pH shift in the vicinity of sample surface can cause the adverse effects for cellular proliferation and adhesion [203]. However, the MC3T3-E1 cells cultured with extracts of other five magnesium phosphate coated AZ31 samples all exhibit more than 90% viability, suggesting the non-toxicity and outstanding cytocompatibility of the as-deposited newberyite and newberyite - trimagnesium phosphate hydrate (TMP) biphasic coatings. 87
Moreover, as indicated by the results of electrochemical and immersion tests, the longer processing times or higher treatment temperatures can enhance the corrosion resistance of magnesium phosphate coatings, subsequently promote their cellular responses and improve the cytocompatibility of AZ31 magnesium alloy.
Figure. 4-9 Cell viability of MC3T3-E1 pre-osteoblasts cultured in medium extracts
of different MgP coated AZ31 samples for 3 days. Statistically significant
differences at *p < 0.05 vs. control.
4.5 Conclusion
This work reports the first attempt to deposit magnesium phosphate coatings including the single phase newberyite layer and newberyite - trimagnesium phosphate hydrate (TMP) biphasic layer on AZ31 magnesium alloy via a microwave assisted coating process for enhanced in vitro corrosion resistance and cytocompatibility. The results indicate that: 1) treatment temperatures play an import role in the phase composition, morphology and degradation behaviors of magnesium phosphate coatings. 2) The as-prepared newberyite layers and trimagnesium phosphate hydrate (TMP) layers present distinct morphology: as
88
the newberyite layer is composed of tabular and polyhedral precipitates; trimagnesium phosphate hydrate (TMP) layer maintains a pile-up structure with flower-like TMP crystals. 3) Electrochemical and immersion tests indicate that the corrosion resistance and bioactivity of AZ31 magnesium alloy is significantly improved by the magnesium phosphate coatings. 4) Due to the high electrochemical stability, thicker and denser structures of trimagnesium phosphate hydrate (TMP) layer, the AZ31 samples treated at
160°C possess the most promising in vitro degradation properties. 5) The spherical/globular apatite precipitates formed on coated surfaces after 2 weeks of immersion in SBF further demonstrate some ability of the magnesium phosphate coatings to induce biomineralization. 6) The cytotoxicity assessment proves the outstanding biocompatibility of the newberyite and trimagnesium phosphate hydrate (TMP) coatings.
To conclude, this study points out the great potential of magnesium phosphates, especially the trimagnesium phosphate hydrate (TMP) layer serving as the protective and bioactive coating on magnesium alloys for orthopedic applications.
89
Chapter 5
Synthesis and In Vitro Evaluation of a Bilayer Coating Combining Poly (Lactic Acid) and Fluorinated Hydroxyapatite on AZ31 Magnesium Alloy
5.1 Abstract
Magnesium alloys have received considerable attentions as next generation metallic implant materials, owing to their distinctive degradability, good biocompatibility and favorable mechanical properties. However, the premature failure issue caused by uncontrolled degradation of Mg alloys in physiological environment has restricted its clinical application. Thus, the FHA and PLA/FHA coatings were developed on AZ31 Mg alloy surface by incorporating the microwave assisted coating process with spin coating technique. The FHA coating showed a nano porous structure, while PLA/FHA coating was found to be much denser and defect-free. More importantly, the with the protection of as- deposited coating, the galvanic corrosion reactions of AZ31 Mg alloy in SBF were significantly retarded. In particular, the corrosion current density of PLA/FHA coated sample was about two orders of magnitude lower than that of untreated AZ31 sample. In vitro degradation test further demonstrated the outstanding corrosion resistance of
PLA/FHA coating and superior biomineralization properties of FHA coating. As a result, 90
it can be speculated that the FHA coating and PLA/FHA hybrid coating are promising candidates as the protective layer on Mg alloys for biomedical applications.
5.2 Introduction
Metallic materials such as titanium alloy and stainless steel are the most widely used implant materials in orthopedic operations for fixation of bone fracture and bone replacement, as they have superior mechanical properties for load carrying applications and low toxicity to the host tissue [204, 205]. However, the mismatch of elastic modulus between Ti alloys, stainless steel and natural bones can cause stress shielding effects which need be corrected by the revision surgery [204, 206]. Also, currently available orthopedic implants are made with non-degradable materials, i.e. Ti alloys and stainless steel, which require permanent affiliation with damaged tissue or removal procedures by a follow up surgery after healing. Thus, considerable efforts have been taken to develop novel implant materials with biodegradable properties in vivo. Among the potential candidates, magnesium and magnesium based alloys have gained increasing interests, owing to their desirable degradability, high strength/weight ratio excellent mechanical and biological compatibility [125, 127, 129]. Furthermore, the elastic modulus of Mg alloys is smaller than those of conventional metallic implant materials and more comparable with that of cortical bone, which could help eliminate the stress shielding issues associated with conventional metallic implants. Mg is an essential mineral component in human body, which ensures that the by-products of Mg dissolution are biologically safe. Additionally,
Mg2+ ion has been found to stimulate initial osteogenesis and promote the remodeling of fractured bones [130]. However, extremely high degradation rate Mg and its alloys in physiological environment, which can result in corrosion induced mechanical failure and 91
localized alkalization, is the foremost obstacle towards their clinical application [125, 132].
Further, H2 gas evolution associated with degradation process of Mg alloys could cause severe consequences against the surrounding tissues, which consequently undermining the osseointegration properties and obstructing the attachment of osteoblasts to the implant surface [133]. Thus, the degradation rate of Mg alloys should be regulated accordingly to meet the time scale for the completion of bone regeneration and exclude the early loss of mechanical stability of Mg alloys during the healing process [125, 129]. To improve the corrosion resistance of Mg alloys and successfully employ them in human body, two major strategies had been proposed: one is tailoring the microstructure of Mg alloys by alloying and mechanical deformation, the other is surface modification by depositing ceramic, polymer or composite coatings on Mg alloys surfaces [132]. In comparison to altering the microstructure and composition of Mg alloys, surface modification can be implemented to various Mg alloy substrates and has shown a more significant effect on corrosion resistance. Therefore, surface modification is recognized as a more favorable technique to optimize the degradation behavior and biological response of Mg alloys.
Coating deposition is one of most promising surface modification techniques to control the degradation of Mg alloys, as the deposited coatings can serve as the barrier layer to block the infiltration of attacking ions in the medium. Since calcium phosphate compounds
(CaP), especially hydroxyapatite (HA/CDHA, Ca10(PO4)6(OH)2) with great chemical similarity to bone mineral have shown well recognized advantages of non-toxicity, high biocompatibility and superior ability to directly bond with cortical bone i.e. osseointegration [147-150], CaP based bioceramics have emerged as the most attractive coating materials on biomedical implants. As a result, diverse calcium phosphate coatings 92
have been developed on Mg alloys via different techniques, such as chemical precipitation[143, 151-156], electrodeposition [157-159], sol-gel [160] and hydrothermal method [161, 162]. In particular, wet chemical precipitation is of special interest, as it can be conveniently employed and applicable to Mg alloy substrates with complex structures.
However, various cations and anions present in the reaction solution could be induced into the lattice of HA, which may alter the structure, stability, solubility and biological properties of deposited HA coatings. For example, with the partial and complete
- - substitution of OH by F , the fluorine doped hydroxyapatite apatite (Ca10(PO4)6 (OH)2−xFx, where 0 Recently, organic-inorganic composite coatings fabricated by combining bioactive ceramic and biocompatible polymer have drawn considerable attentions. The as-prepared organic-inorganic coatings with less intrinsic defects have shown remarkable protection against the corrosion reactions and excellent durability [211]. So far, two distinct structures of organic-inorganic composite coatings have been proposed and explored: 1) single layer 93 structure of inorganic bioactive nanoparticle uniformly distributed in biopolymer matrix; 2) bilayer structure of biopolymer film covering the bioceramic layer for pore-sealing function. Herein, we aim at studying the corrosion resistance and in vitro properties of AZ31 Mg alloy modified with the PLA/FHA bilayer coating. A nano-sized fluorinated hydroxyapatite coating was firstly deposited on AZ31 magnesium alloy as the bottom layer via a facile microwave assisted coating technique. It is well known that microwave irradiation can dramatically enhance rate of chemical reaction and transformation due to the rapid heating characteristic of microwave processing. Consequently, the coating deposition can take place within minutes/seconds in the microwave accelerated coating process [115, 124]. Then, the biocompatible and biodegradable PLA film was sealed on top of the as-deposited FHA coating for further anti-corrosion enhancement. PLA is an FDA approved biodegradable polymer and has shown corrosion resistance property to a certain extent [141]. Therefore, it is expected PLA film would suppress the degradation behavior of magnesium alloy and simultaneously meet the clinical requirements. 5.3 Experimental 5.3.1 Material preparation Commercially available AZ31 magnesium alloy (3% Aluminium, 1% Zinc) sheets were firstly machined to dimensions of 1cm 1cm 2mm. Afterwards, AZ31 Mg alloy blocks were ground with abrasive SiC paper up to #1200 grit to ensure the homogeneous surface roughness, then ultrasonically degreased in ethanol and rinsed inDI (deionized) 94 water for 10 min respectively. After the cleaning process, samples were dried at 60 °C for 30 min and cooled down in air. 5.3.2 FHA deposition The coating solution for fluorinated hydroxyapatite (FHA) deposition was prepared by dissolving 1.9836g Mg(NO3)2·6H2O, 0.6048g NaH2PO4 and 0.0705g NaF in 200ml DI water. Two AZ31 substrates were placed in one 200 ml beaker, filled with 100 ml coating solution. The top of the beaker was covered by an alumina fiberboard (Zircar Ceramic, Florida, NY). Afterwards, the entire setup was placed in a 1200 W microwave oven (Panasonic, Japan) and heated at maximum power for 5 minutes. The above steps were repeated once more to ensure the uniformity of the coatings. Finally, the coated samples were dried in an oven at 60°C. 5.3.3 Spin coating For the deposition of PLA film, PLA powders were weighed and dissolved into chloroform obtain 4% (w/v) solution with good transparency. The PLA film was prepared using one-step spin coating process. The PLA solution was uniformly spread on the bare AZ31 Mg substrates and FHA treated AZ31 blocks by using a micropipette and spin coated for 30 s at a rotating speed of 3000 rpm. Subsequently, the coated surface was immediately dried in air at room temperature. The abovementioned coating procedure was repeated for 3 times to ensure a uniform and thicker PLA film. The Fig. 5-1 represents the schematic diagram of the spin coating process. 95 Figure. 5-1 Schematic diagram of the setup of spin coated and the PLA film deposition of spin coating. 5.3.4 Coating characterization The phase structure of as-coated samples and bare AZ31 substrate was identified by X- ray diffraction (XRD, Ultima III, Rigaku)) with monochromated Cu Kα radiation (44KV, 40mA) over a 2θ range of 10–60°. Surface morphology and elemental composition were examined using scanning electron microscopy (SEM, S4800, Hitachi) equipped with an energy dispersive X-ray spectroscopy (EDS, Oxford INCA). In addition, an accelerated voltage of 20 KV with a working distance of 15mm was employed for EDS analysis. Functional groups present in the FHA coating and PLA film were characterized by a Fourier Transform Infrared Spectroscopy (FTIR, UMA-600 Microscope, Varian Excalibur Series) using an ATR diamond crystal for 256 scans in the range between 4000-700 cm-1 with a resolution of 1cm-1. For cross-section imaging, the coated samples were milled by focused ion beam to produce a clean cross-section. 96 5.3.5 Electrochemical test To evaluate the corrosion behavior of both coated and uncoated samples in simulated body fluid (t-SBF), potentiodynamic polarization tests were conducted using Gamry Reference 600 potentiostat. t-SBF was prepared by dissolving 6.547 g NaCl, 2.268 g NaHCO3, 0.373 g KCl, 0.178 g Na2HPO4 · 2H2O, 0.305 g MgCl2 · 6H2O, 0.368 g CaCl2 · 2H2O, 0.071 g Na2SO4 and 6.057g Tris-buffer, and 40ml of 1 M HCl respectively, in 1 L ultra pure water [164].. All the reagents were from Fisher sci, and dissolved in solution one by one. For the potentiodynamic test, a conventional three-electrode arrangement was employed. It is composed of the sample as working electrode, Ag/AgCl electrode as reference electrode, and graphite rod as counter electrode. The potentiodynamic polarization curves were measured at a scan rate of 1mV/s after a steady open circuit potential (OCP) was achieved. 5.3.6 In vitro degradation Degradation behaviors of uncoated and coated samples were evaluated by an immersion test in simulated body fluid (t-SBF) according to ASTM-G31-72. Each sample was kept in 60 ml t-SBF in a thermostatic water bath at 37 ± 0.5 °C for 15 days. In addition, t-SBF solution was replenished every other day. The pH value of SBF solution was monitored using pH meter every two days during the immersion test. After a certain period of immersing, samples were rinsed with distilled water flow to remove the formed corrosion byproducts. The weight losses of samples were recorded at each time point to determine the degradation rate. The degradation rate, in terms of the relative percentage of mass loss was derived from following equation: 97 푊1−푊0 푊퐿 = − ( ) ×100 (1) 푊0 where WL, W0, and W1 are the percentage of weight loss, and the weights before and after immersion for different time intervals, respectively. After immersing in SBF for two weeks, bare and coated samples were also characterized by scanning electron microscopy (SEM) to evaluate extent of corrosion and biomineralization properties. 5.4 Results and discussion 5.4.1 Coating characterization X-ray diffraction patterns of FHA, PLA/FHA coated and bare AZ31 substrates are shown in Fig. 5-2. All the samples presented significant diffraction peaks corresponding to Mg substrate, which implies the as-deposited coatings were very thin. After the FHA coating treatment, diffraction peaks in the 2-theta range of 25-30° and 50-55° were distinguished from characteristic peaks of Mg. These peaks which can be assigned to FHA phase, indicate the formation of single phase and poorly crystallized FHA layer on the AZ31 Mg substrate. The intensity of FHA diffraction peaks on PLA/FHA coated sample was further weakened due to amorphous PLA film reduced the X-ray penetration. Although The replacement of OH– by F– in HA lattice leads to contraction in the a-axis dimension, FHA presents quite similar cyrstal structure with HA and belongs to the same space group of HA and (space group: P63/m; parameters: a=b=9.462 Å and c=6.849 Å, α=β=90°, γ=120°). Fig. 5-3 shows the FT-IR spectra of PLA, FHA and PLA/FHA composite coatings 3- on the AZ31 substrate. The absorption peak attributed to phosphate (PO4 ) band -1 (asymmetric stretching ν3 mode) is centered at 1050 cm [168]. Moreover, the broad 98 -1 -1 absorption peaks at 3300 cm and 1652 cm can be assigned to v1 symmetric stretch and v2 bend of structural water respectively. It is worth mentioning many extra absorption peaks observed in PLA and PLA/FHA do not belong to the FHA. These characteristic peaks are the result of symmetric and asymmetric stretching of ether group (C – O – C) and carbonyl (C = O) in PLA structure. Overall, the FTIR results are in line with XRD analysis. Figure. 5-2 XRD patterns of the AZ31 substrate, the FHA coated, and the PLA/FHA coated samples. Figure. 5-3 FTIR spectra of the PLA coating, the FHA coating, and the PLA/FHA composite coating. 99 Fig. 5-4 presents the surface features along with elemental composition of FHA and PLA/FHA coated specimens. Neither the substrate nor cracks were visible in SEM micrographs, indicating the good uniformity of the deposited coatings. The as-deposited FHA coating exhibited a relatively porous and nested structure with irregularly arranged FHA nanospheres. As seen in high magnification image of FHA coating (Fig. 5-4b), the sizes of irregular pores formed on FHA coating ranged from 10 nm to 200nm. After the sealing by PLA film, the surface was altered to be featureless and highly smooth, as seen in Fig. 5-4c. The EDS results (Fig. 5-4d) of PLA/FHA coating revealed the presence of Ca, P and F in the composite coating, which is indicative of the formation of FHA coating as an interlayer between PLA film and Mg substrate. In addition, the high amount of C and O concentration on the coating surface is the configuration of organic polymer PLA. The cross-section profile of PLA/FHA coated sample showed the compact coating with the thickness around 1.6 µm comprising two distinct layers, as displayed in Fig. 5-5. The inner layer is a nano-FHA layer with a relatively thin thickness of 400 nm, while the outer layer is a PLA polymeric film with a relatively thick thickness of 1.2 µm. More importantly, the PLA film effectively sealed the pores present in FHA layer, which could potentially block the pathway of the infiltration of attacking ions and improve the corrosion resistance of Mg alloys. Besides, rough surface of as-deposited FHA layer was found to increase thickness of PLA film. Previous studies demonstrated that the coating thickness plays an important role in the anti-corrosion behavior of Mg alloys, where the thicker can provide more effective and long-term corrosion resistance[212]. The organic PLA film and inorganic FHA layer were closely bonded together, indicating the high interlocking strength of the bilayer structure. Moreover, no distinct gaps between the deposited FHA coating and AZ31 100 substrate are located, which is evident for the good adherence between FHA layer and AZ31 substrate. Figure. 5-4 SEM images of as-deposited (a) FHA coating, (b) FHA coating with high magnification, (c) PLA/FHA coating and EDS spectrum of (d) PLA/FHA coating Figure. 5-5 Cross-sectional SEM micrograph of as-deposited PLA/FHA coating on Mg alloy AZ31. 101 5.4.2 Electrochemical behavior The electrochemical Potentiodynamic polarization that can intensify the corrosion reaction and accelerate the corrosion process, is an efficient tool to evaluate the corrosion resistance of materials in a short period. Fig. 5-6 shows the potentiodynamic polarization curves of the bare and coated AZ31 samples, while corresponding potentiodynamic parameters calculated based on these curves are listed in Table 5.1. Compared to that of uncoated AZ31 sample, the polarization curves of FHA, PLA and PLA/FHA coated samples shifted towards higher corrosion potentials and lower corrosion current densities, which are evidences for favorable anti-corrosion behavior. Particularly, the corrosion potentials (Ecorr) of FHA and PLA/FHA coated samples moved to more positive potential with an increase slightly less than 90 mV and 180 mV respectively. The more positive corrosion potential indicates higher thermodynamic stability state of the sample. Obviously, the enhancement in thermodynamic stability of the FHA coated sample can be attributed to the ceramic nature of deposited FHA layer and its high chemical stability in physiological conditions. Despite the corrosion potential of PLA coated sample shifted to lower value, potential does not necessarily represent the corrosion resistance of PLA coatings. As corrosion potential majorly specifies the thermodynamic properties of materials, while the corrosion current density directly indicates the corrosion resistance properties. As seen in Table 5.1, the corrosion current density (icorr) of PLA and FHA coated sample was reduced to about 95.58 and 104.37 μA/cm2 respectively, which are almost one order of magnitude lower than that of un-treated AZ31 magnesium alloy. These results demonstrated that both PLA and FHA coatings are capable to prevent the ionic diffusion to Mg substrate and improve the corrosion resistance to some extent. Moreover, 102 by stacking PLA film and FHA layer on AZ31 magnesium alloy surface, a significant decrease in corrosion current density was achieved. The corrosion current density (icorr) of PLA/FHA coated sample was nearly two orders of magnitude lower than that of the bare AZ31 sample, which strongly suggests the superior anti-corrosion performance of PLA/FHA coating. The thick and compact structure of the deposited bilayer coating, PLA seals the pores of FHA layer and minimizes the intrinsic defects are widely expected to be reasons that notably enhance the corrosion resistance of PLA/FHA coated sample [213- 215]. Figure. 5-6 Potentiodynamic polarization curves of uncoated, FHA coated, PLA coated and PLA/FHA coated AZ31 Mg alloy samples in SBF solution. 103 Table 5.1 Electrochemical parameters of samples calculated from polarization curves Sample Corrosion potential, Ecorr Corrosion current density, icorr 2 (VAg/AgCl) (μA/cm ) Bare AZ31 -1.510 796.31 AZ31-FHA -1.424 104.37 AZ31-PLA -1.697 95.58 AZ31-FHA-PLA -1.336 9.91 5.4.3 In vitro Degradation behavior Upon immersing in SBF solution, the anodic dissolution reaction occurred on the surface of Mg alloys, according to following reaction: Mg→Mg2+ +2e- (2) Meanwhile, due to occurrence of galvanic corrosion, the cathodic reaction took place simultaneously, which is also accopanied with hydrogen evolution. The equation (3) represents the cathodic reation occured in the electrolyte. - - 2H2O+2e →H2+2(OH) (3) Later, the incorporation of Mg2+ and (OH)- could induce the insoluble precipitates Mg(OH)2 and result in the passive layer (Equation. 4) on Mg surface which could hinder the further anodic dissolution of Mg [174]. 2+ - Mg + 2(OH) → Mg(OH)2 (4) - However, with the presence of chloride ions (Cl ) in electrolyte, insoluble Mg(OH)2 was transformed into highly soluble MgCl2, following reaction (5). 104 2+ − Mg + 2Cl → MgCl2 (5) Due to the aggressive attack by corrosive mediums, the Mg(OH)2 passive layer was easily dissolved and the more active surface was exposed, which led to severe pitting corrosion and rapid degradation of Mg alloy [175]. Based on the degradation mechanism discussed above, the weight loss of Mg alloy and the pH elevation of the SBF solution are evidently two key parameters indicating the degradation properties of Mg alloy in vitro. Here, the mass loss and pH variations were continuously recorded during the immersion test to evaluate the in vitro degradation behavior of the coated AZ31 Mg alloy samples. As seen Fig. 5-7 (a), the bare AZ31 samples showed an intense degradation in the first day of immersion test, which corresponds to results of rapid degradation of Mg at initial stage [216]. Moreover, the PLA and FHA coated samples exhibited similar profiles of weight loss, but with less weight loss at each time point. In contrast, the PLA/FHA coated samples presented least weight loss, which is less than one fifth of the weight bare AZ31 lost after 15 days incubation in SBF. This fact further demonstrated that the PLA/FHA coating notably improved the corrosion resistance and suppressed rapid degradation of AZ31 Mg alloy in physiological environment. The surrounding pH value is important for many cellular responses and unsuitable pH conditions such as rapid alkaline pH shift can cause several adverse effects on cell proliferation and adhesion [203]. Thus, the changes of pH value in SBF solution were closely monitored in immersion test. As seen in Fig. 5-7 (b), the pH of solution incubating the bare AZ31 sample lifted to 8.4 at day 1, due to the severe corrosion reaction took place at the early stage. Later, the pH value steadily decreased, owing to the presence of passive layer and precipitation of apatite. Conversely, the FHA, PLA and PLA/FHA coated 105 specimens showed a mild increment in pH value at the beginning. Especially, the pH value of solution incubating PLA/FHA coated sample stabilized around 7.6, which is close to value in blood plasma and favors the potential application of PLA/FHA coated Mg alloys in human body. Other than the excellent corrosion resistance brought by the deposited coatings, the presence of F- is also expected to decrease the pH value [217]. Overall, the in vitro degradation behavior of the samples was in good agreement with the results of the electrochemical test. The surface morphologies and elemental compositions of FHA and PLA/FHA coated samples after immersing in SBF for 15 days are shown in Fig. 5-8. As the result of severe corrosion reaction, sizable and deep cracks were generated on the surface of FHA coated sample. The loose and porous structure of the fluorine doped hydroxyapatite coating could be the reason that results insufficient corrosion protection. Conversely, no significant change in morphology of PLA/FHA coated. Except few micro cracks and corroded pits, the PLA/FHA coated sample mostly remained intact after the immersion test. This evinces that the outer PLA film adequately made the AZ31 Mg alloy substrate less prone to the galvanic corrosions. Accompanying initiation of cracks on the surface of FHA coated AZ31 sample, numerous bone-like apatite globules were uniformly deposited on the surface of FHA coated AZ31 sample. The corresponding EDS spectrum also revealed high concentration of Ca and P on the FHA surface after immersion in SBF, which is indicative of the formation of a dense apatite layer. In contrast, only a few clusters consisted of Mg(OH)2 and apatite precipitates on the surface of PLA/FHA coating after incubating in SBF for 15 days. The outstanding bioactivity of FHA coating that remarks promising ossointegration properties in vivo, can be ascribed to two aspects: 1) the porous structure 106 of as-deposited FHA coating significantly enlarged the surface area of FHA layer, amplified the release of hydroxy ions and ionic diffusion in physiological liquid. Thus, the nucleation and crystal growth of apatite was promoted accordingly. 2) because of the smaller ion radii and stronger electronegativity, the presence of F- in the coating structure can result in a more negatively charged surface, and induce more formation of Ca-rich ACP precipitates. As the Ca-rich ACP precipitates are considered to be the nucleation site of bone like apatite, FHA coating is anticipated to enhance the nucleation rate in SBF solution [218]. Figure. 5-7 (a) Weight loss of surface treated samples and control sample and (b) variations of pH values of SBF solution at various immersion periods in SBF. 107 Figure. 5-8 Surface morphology and corresponding EDS spectra of (a) FHA and (b) PLA/FHA coated AZ31 Mg alloy samples after 15 days immersion in SBF. 5.5 Conclusion In this work, the FHA coating and PLA/FHA bilayer coating were produced on AZ31 Mg alloy by combining a facile microwave assisted coating technique and spin coating. The PLA/FHA hybrid coating consisted of a PLA polymeric layer and a nano-FHA ceramic layer. The inner FHA layer exhibited a porous structure withthe thickness of 400 nm, while the outer PLA layer with the thickness of 1.2 µm presented a compact structure and effectively reduced the defects in FHA layer. Electrochemical tests and in vitro test were performed to investigate the corrosion resistance and degradation behavior of the FHA and PLA/FHA coated samples. As expected, the PLA/FHA coating dramtically suppressed the corrosion current density and degradation rate of AZ31 magnesium in the physiological 108 environment. Moreover, the FHA coating showed great ability to induce bone-like apatite precipitation and great potential for bone regeneration applications. 109 Chapter 6 Nanostructured Amorphous Magnesium Phosphate / Poly (Lactic Acid) Composite Coating for Enhanced Corrosion Resistance and Bioactivity of Biodegradable AZ31 Magnesium Alloy 6.1 Abstract In view of the combination many interesting properties, magnesium alloys have attracted considerable interest as suitable metallic biomaterials for bioresorbable orthopedic implants. Nevertheless, their fast degradation in physiological environments poses challenges for their practical applications. Here, we report that spin coating of composites of nano amorphous magnesium phosphate (nAMP) and poly (lactic acid) (PLA) on AZ31 magnesium alloy. The idea is to use the nAMP/PLA composite film while tailoring the degradation and enhancing the bioactivity of magnesium alloys. SEM examinations show that as-deposited nAMP/PLA film is smooth, crack-free and the nAMP particles are well distributed in PLA matrix. The electrochemical test including potentiodynamic polarization and EIS associated with immersion test results reveal that the corrosion activities of nAMP/PLA coated AZ31 magnesium alloy in SBF are markedly suppressed. Furthermore, it is seen that massive bone-like apatite precipitates formed on 110 surface of nAMP/PLA coated sample, which is indicative of the superior biomineralization capability achieved by nAMP/PLA nanocomposite coating. Thus, the nAMP/PLA composite coating has great potential to be employed as the protective and bioactive coating on biodegradable magnesium alloys for orthopedic applications. 6.2 Introduction Over the last decade, magnesium alloys have attracted tremendous interest as biodegradable orthopedic implant materials, due to their low density, suitable elastic moduli and degradability in physiological environment [125, 126, 178]. The controlled degradation of magnesium implants in vivo could eliminate the second surgical intervention to avoid potential risks associated with such procedures. The high degradation kinetics of magnesium alloys, however is a double-edged sword. The rapid degradation of magnesium alloys in physiological condition usually leads to the locally elevated pH value, formation of hydrogen gas cavities and premature mechanical failure of implants [130, 178, 180]. To retard the gas evolution and enhance the corrosion resistance of magnesium alloys, various forms of surface modification have been proposed. The protective coatings could be, but not limited to chemical conversion coatings [219], calcium phosphate based bioceramic coatings [119, 156] and polymeric coatings [220, 221]. Of special interest are biodegradable polymeric coatings which can not only serve as the barrier layer, but can also be applied as the drug delivery vehicles [132]. A case in point is poly-lactic acid (PLA) which is already cleared by the FDA for clinical applications [222]. The degradation of PLA through bulk erosion mechanism is relatively slow, and the degradation product (lactic acid) is metabolically digestible [223]. To improve degradation resistance of 111 magnesium alloy, PLA coatings have been deposited on magnesium and its alloys via various methods, such as spin coating, dip coating and spray [224-226]. In general, the PLA film prepared by spin coating is more uniform with fewer defects, thereby providing better corrosion resistance. However, the degradation of PLA generates the acidic product that can lower the local pH value, subsequently hinder the biomineralization and osseointegration [227]. Therefore, incorporating the nano-sized bioactive ceramic in PLA matrix is a feasible solution to overcome its drawback, while regulating the degradation of magnesium alloy simultaneously. Magnesium phosphate compounds are the alternative to the well-known calcium phosphates and have not been extensively explored. Amorphous magnesium phosphate (AMP) is a sub-set of MgO-P2O5 group, which has been attracting increasing interests due to their outstanding biocompatibility, biodegradability and bioactivity [183, 184]. Similar to its counterpart amorphous calcium phosphate (ACP), AMP is also considered to be a promising candidate to be used as a bone substitute, filler in biocomposite and coatings which will not only support sustained release of magnesium and phosphate ions, but also can significantly promote the biological activities of AMP based biomaterials[92, 185, 228]. In spite of these widely-recognized benefits, the research activities regarding the synthesis and evaluation of AMP are still at a nascent scale [109, 187]. Thus, to solve the low yield of previous synthesis methods, Babaie et al developed the ethanol induced precipitation method to produce copious amount of AMP nano-particles [186]. Recently, employing AMP in MgP cement have received considerable attention, as the AMPs based cement could overcome the issue of relatively low bioactivity and mechanical properties associated with conventional CaP cements [181, 183, 229]. However, the Mg-P 112 compositions and processing parameters still need to be further regulated to achieve clinically relevant properties. To best of our knowledge, this is the first attempt to use nAMP/PLA composites as protective coatings on AZ31 magnesium alloys. The hypothesis is the nAMP particles dispersed in PLA matrix would remarkably improve the bioactivity and corrosion resistance, compared to pristine PLA film. Herein, the nAMP/PLA film was fabricated via spin coating and the vitro degradation properties of nAMP/PLA coated AZ31 sample were investigated. 6.3 Experimental 6.3.1 Material preparation Commercially available AZ31 magnesium alloy plates were cut into small substrates with dimensions of 1cm 1cm 2mm and ground up to 1200 grit SiC abrasive paper to ensure the homogeneous roughness. Subsequently, all of the polished samples were ultrasonically cleaned in ethanol for 10 minutes and dried in air. 6.3.2 Coating preparation PLA powders was weighed and dissolved into chloroform obtain 4% (w/v) solution. The nAMP powder were synthesized by following the procedure reported in our previous publication [186]. Magnesium nitrate [Mg(NO3)2.6H2O] solution was added to a solution containing diammonium hydrogen phosphate [(NH4)2HPO4] at the certain molar ratio of Mg:P:2:1. Afterwards, the precipitates were washed, centrifuged and oven dried. nAMP/PLA composite suspension was prepared by adding nAMP powder (20 wt.% of PLA) to the PLA solution. Finally, the mixture was magnetically stirred continuously for more than 4 h and followed by ultrasonic dispersion for 10 min. The Figure. 6-1 shows the 113 photographs of pristine PLA solution and nAMP/PLA suspension. The PLA solution is transparent, whereas the nAMP/PLA suspension presents a milky color. Moreover, no sedimentation is observed at bottom of the container, which indicates a good dispersion of nAMP particles in PLA solution. The coating process was conducted using a spin coater. The PLA solution and composite suspension were applied on the magnesium substrate by using a micropipette and spin coated for 30 s at a rotating speed of 3000 rpm. Subsequently, the coated surface was immediately dried in air at room temperature. The above-mentioned coating procedure was repeated for 3 times to ensure a uniform and thicker coating. Figure. 6-1 Photographs of PLA solution and nAMP/PLA suspension. 6.3.3 Characterization Phase composition of as-prepared nAMP powder and as-coated samples were identified by X-ray diffraction (XRD, Ultima III, Rigaku)) with monochromated Cu Kα radiation (44KV, 40mA) over a 2θ range of 10–45° and 10–60° respectively. The transmission electron microscopy (TEM, HD-2300, Hitachi, USA) was employed to investigate the morphologies and particle sizes of the nAMP particles. Surface features and 114 elemental compositions of the as-deposited coatings were analyzed using scanning electron microscopy (SEM, S4800, Hitachi) equipped with an energy dispersive X-ray spectroscopy (EDS, Oxford INCA). Elemental analyses were performed at 20KV with the working distance of 15mm. Functional groups present in coatings were characterized by a Fourier Transform Infrared Spectroscopy (FTIR, UMA-600 Microscope, Varian Excalibur Series) using an ATR diamond crystal for 256 scans at the range between 4000-700 cm-1 with a resolution of 1cm-1. 6.3.4 Electrochemical test To evaluate the corrosion behavior of both coated and uncoated samples in simulated body fluid (t-SBF), potentiodynamic polarization tests and Electrochemical impedance spectroscopy (EIS) were carried out using Gamry Reference 600 potentiostat. t-SBF was prepared following the previously reported recipe which better mimics human blood plasma [164]. A conventional three-electrode cell system comprising of the prepared sample as working electrode, a Ag/AgCl electrode as reference electrode, and a graphite rod as counter electrode was employed in this study. The potentiodynamic polarization curves were measured at a scan rate of 1mV/s after the open circuit potential (OCP) was stabilized for 30 min. Electrochemical impedance spectroscopy (EIS) was conducted at the open circuit potential 10 mV sinusoidal amplitude over a frequency range of 100 kHz to 10 mHz. All the electrochemical tests were performed at room temperature. Gamry Echem Analyst software was used for impedance fitting and data analysis. 115 6.3.5 Immersion test Degradation behaviors of uncoated and coated samples were evaluated by an immersion test in simulated body fluid (t-SBF) according to ASTM-G31-72. Each sample was kept in 60 ml t-SBF in a thermostatic water bath at 37 ± 0.5 °C for 15 days. In addition, t-SBF solution was refreshed every two days. The pH value of SBF solution was monitored during the immersion test. After a certain time of soaking, samples were rinsed with distilled water flow to remove the formed corrosion products. The weight losses of samples were measured at each time point to determine the degradation rate. The degradation rate, presenting the relative percentage of mass loss was derived from following relation: 푊1−푊0 푊퐿 = − ( ) ×100 (1) 푊0 where WL, W0, and W1 are the percentage of weight loss, and the weights before and after immersion for different time intervals, respectively. Post-immersion samples were also characterized by scanning electron microscopy (SEM) to evaluate their abilities to stimulate biomineralization in vitro. 6.4 Results and discussions 6.4.1 Characterization of nAMP powder The diffraction pattern of synthesized powder is shown in Fig. 6-2 (a). According to the figure, there is only a broad bump centered at 2θ=30° that can be indexed as amorphous magnesium phosphate. As shown in TEM image (Fig. 6-2b), the sample consists of nano- sized spheres with diameters ranging from 40 to 70 nm. The selected area electron 116 diffraction (SAED) pattern (inset of Fig. 6-2b) features diffused halos, which demonstrates the amorphous nature of the as-synthesized powder. The amorphous magnesium phosphate is recognized as the initial phase precipitated in concentrated Mg-P solution [228]. However, the methods of synthesizing AMP have been limited to aqueous/ethanol precipitation and microwave irradiation, due to the instability of MgP in aqueous medium [109, 186, 187, 195, 229]. Especially, the fabrication of nano-sized AMP powder remains challenging, which yields highly tailored reaction conditions. Our previous efforts have shown that the microwave accelerated reaction and freeze drying are two practicable routes to synthesize nano-sized AMP powder [109, 186]. Figure. 6-2 (a) XRD pattern of as-synthesized nAMP powder and (b) TEM image and SAED pattern of as-synthesized nAMP powder. 117 6.4.2 Characterization of PLA and nAMP/PLA coatings Fig. 6-3 (a) shows the XRD patterns of non-coated and coated AZ31 samples. Except for the diffraction peaks belonging to magnesium, no other characteristic peaks are observed. This means PLA film and nAMP/PLA composite film present a non-crystalline structure. Furthermore, the amorphous structure of nAMP/PLA composite implies no chemical reactions occurred between nAMP particles and PLA solution, as opposed to the phase evolution of AMP in PVA solution [183]. The infrared spectra of the AMP powder and coated samples are illustrated in Fig. 6-3 (b). The FTIR spectra of nAMP and nAMP/PLA coated samples exhibit absorption peaks at around 1006 and 1049 cm-1, which 3- can be assigned to the 3 bending of phosphate (PO4 ) ions [169, 220]. This also confirms the presence of nAMP particles in PLA matrix film. In addition, the broad bump located at 3255 cm-1 and the absorption band at 1650 are ascribed to the structural water in nAMP and the stretching of OH-. More importantly, no functional groups related to chloroform solvent is identified in the composite film, as chloroform is potentially harmful to human health [230]. This also implies the complete evaporation of chloroform solvent was achieved during the spin coating process. 118 Figure. 6-3 (a) XRD patterns of bare AZ31, PLA and PLA-AMP coated AZ31 magnesium alloy and (b) Infrared spectra of nAMP, PLA and PLA-AMP coated AZ31 magnesium alloy. The surface morphology of the bare AZ31 and surface treated samples is shown in Fig. 6-4. As compared to the bare AZ31 sample, both PLA coated and nAMP/PLA samples present a much smoother surface without visible pores. Due to the incorporation of nAMP particles in PLA, tiny dots are observed in the high magnification SEM image of the nAMP/PLA coated sample (Fig. 6-4d). This suggests that the nAMP particles stayed on the top surface of PLA matrix rather than setting down on the bottom. Even though with some level of agglomeration, the nano AMP particles preserved nano structures and exhibited a homogeneous distribution in the nAMP/PLA composite film, as marked by the green dash. The corresponding energy dispersive spectrum is displayed as the inset of Fig. 6-4d. The detected elements such as Mg and P are indicative of the presence of AMP particles. Additionally, carbon and oxygen are two major elements of PLA. The cross- section profiles of the coated samples were obtained using focused ion beam (FIB) and 119 examined by scanning electron microscope (SEM). As seen in Fig. 6-5 (a), the PLA coating shows a typical and uniform polymeric film with the thickness around 550 nm. Compared to the pristine PLA film, the nAMP/PLA composite film is slightly thicker and presenting a thickness in the range of 600 nm to 750 nm. The slightly thicker composite film can be attributed to the relatively higher viscosity of nAMP/PLA suspension, as the same spin- coating speed was applied to two kinds of films. The dispersed nAMP particles are also identified by the green oval in cross-section image of nAMP/PLA coated samples. Moreover, no distinct gaps between the deposited films and AZ31 substrate are located, which is evident for the filling of the nanocomposite suspension in abrasive scratches, the good adherence and physical interlocking of pristine PLA and nAMP/PLA coatings [231]. Figure. 6-4 Surface SEM images of (a) Bare AZ31, (b) PLA coated AZ31, (c) nAMP/PLA coated AZ31 and (d) nAMP/PLA coated AZ31 with higher magnification and respective EDS spectrum. 120 Figure. 6-5 Cross-section SEM images of (a) PLA coated AZ31 and (b) nAMP/PLA coated AZ31. 6.4.3 In vitro degradation behavior 6.4.3.1 Electrochemical test The potentiodynamic test was performed to study the corrosion resistance of un-coated and coated AZ31 samples. The potentiodynamic curves of all the samples are illustrated in Fig. 6-6 and the extrapolated electrochemical parameters are summarized in Table 6.1. In comparison to the corrosion potential (Ecorr) of bare AZ31 (Ecorr = -1.501V), the Ecorr of PLA and nAMP/PLA coated AZ31 sample decreased 197 mV and 122 mV respectively. The thicker coating of nAMP/PLA composite film could be a reason for less negative Ecorr, as compared to the PLA coated sample [220]. In general, the corrosion potential (Ecorr) only indicates the reaction rate between anodic and cathodic reactions, is not proportionally correlated to the corrosion resistance. Despite the corrosion potential of bare AZ31 sample is moderately more positive, it still shows the lowest corrosion resistance with highest 2 corrosion current density (icorr) of 796.31 μA/cm . Conversely, the PLA coated sample and 121 nAMP coated sample exhibit the corrosion current density of 95.58 μA/cm2 and 13.52 μA/cm2 respectively, which is at least one order of magnitude lower than that of uncoated AZ31 sample. Thus, the corrosion resistance of various samples can be represented in following orders from highest to lowest resistance: nAMP/PLA coated AZ31 > PLA coated AZ31 > uncoated AZ31. Especially, the low corrosion rate of nAMP/PLA coated AZ31 sample signifies the great protective effects achieved by the synergy of PLA coating and nAMP particles. Since the solubility of AMP in SBF is minor, the positive effect of incorporating nAMP in PLA matrix could include, occupying the potential defective sites in PLA matrix, reducing active corrosion area and hindering the direct infiltration of aggressive ions in physiological environment [232, 233]. Figure. 6-6 Polarization curves of bare AZ31, PLA and nAMP/PLA coated sample. 122 Table 6.1 Electrochemical parameters of samples obtained from polarization curves. Sample Corrosion potential, Ecorr Corrosion current density, 2 (VAg/AgCl) icorr (μA/cm ) Bare AZ31 -1.510 796.31 AZ31-PLA -1.697 95.58 AZ31-PLA-AMP -1.622 13.52 To gain a better understanding of the corrosion mechanism of PLA coating and nAMP/PLA composite coating in SBF, the electrochemical impedance spectroscopy (EIS) analysis was conducted. The EIS Nyquist plots of various samples in SBF solution is depicted in Fig. 6-7(a). Due to the small dimensions of the Nyquist curve for bare AZ31, an expanded plot was included in Fig.6-7(a) as an inset figure. The bare and PLA coated AZ31 sample exhibit two capacitive loops, which can be attributed to charge transfer reaction and mass transportation during the corrosion process [234]. Moreover, the capacitive loop at mid-frequency observed on PLA coated AZ31 sample implies the electrochemical activities also initiated at the interface between PLA coating and AZ31 substrate [235]. In comparison, the nAMP/PLA coated sample presents a single and markedly enlarged capacitive loop. In general, the large diameter of capacitive loop in Nyquist plot represents low corrosion rate [236]. This finding is in good accordance with potentiodynamic results and demonstrates nAMP/PLA composite coating can provide more efficient corrosion protection, compared the pristine PLA coating. To obtain the best fit, the equivalent circuits (Fig. 6-7 b and c) were employed for modeling the experimental 123 data [224]. As shown the Fig. 6-7, the equivalent circuit is composed of solution resistance (Rs), pore resistance (Rp), barrier resistance (Rb), charge transfer resistance (Rct). In addition, constant phase element was used to substitute the pure capacitance for compensating non-ideal capacitive behavior. The constant phase element (CPE) is described as 1 푍퐶푃퐸 = 훼 (2) 푌0(푗휔) where Y0 is the magnitude of the capacitance, ω is angular frequency with ω=2πf, and 훼 is the exponential term of CPE representing the deviation from the ideal capacitance, which ranges from 0 to 1. In this case, three constant phase elements are applied, which correspond to the capacitance of PLA film (CPEPLA), oxide layer (CPEOX) and electrical double layer (CPEDL). The values of fitted data are shown in Table 6.2. It can be seen the exponential parameters of PLA and nAMP/PLA coating are close to the unity. This indicates coated samples preserve small deviations from the ideal capacitance and relatively homogeneous current distribution on the film, which implies good uniformity of as-prepared polymer and composite coatings [236, 237]. Furthermore, the CPEPLA value of nAMP/PLA coated sample is smaller than that of PLA coated sample signifying the less active corrosion area retained on the surface of nAMP/PLA composite coated sample [238]. In comparison to charge transfer resistance of bare AZ31, those of PLA and nAMP/PLA coated samples increase by more than 6-fold and 110-fold to 1051.97 Ω.cm2 and 17022.6 Ω.cm2 respectively. It’s clearly evident that the significant enhancement of corrosion resistance of coated AZ31 magnesium alloy can be ascribed to the protection of PLA and nAMP/PLA films. Alabbasi et. al reported the degradation resistance of AZ31 124 alloy resistance can increase linearly with the growth in thickness of the PLA film [224]. Thus, the significant improvement in polarization resistance (Rp) of the nAMP/PLA coated sample can be partially attributed to the slightly thicker nanocomposite coating, compared to pristine PLA film. More importantly, the addition of AMP particles can relieve the intrinsic pores/defects in PLA film and impede the penetration of corrosive media, which is the playing a major role in enhancing the anti-corrosion performances of magnesium alloys. This demonstrates the great potential of nAMP/PLA composite films serving as the protective layer of magnesium alloy to tailor its degradation rate in physiological environment. Figure. 6-7 (a) Nyqist plots of uncoated and coated samples in SBF solution and equivalent circuit models applied for fitting EIS spectra of (b) bare AZ31and (c) PLA and nAMP/PLA coated samples. 125 Table 6.2 Fitting results of EIS spectra of bare AZ31, PLA and nAMP/PLA coated AZ31 samples 2 Resistance/(Ω.cm ) CPEPLA CPEOX CPEDL Sample Rs Rp Rb Rct Y01 훼1 Y02 훼2 Y03 훼3 Bare AZ31 37.03 - 21.89 144.45 - - 5.46019E-05 1 0.002337473 0.553 AZ31-PLA 121.11 161.9 1166.2 1051.97 8.81087E-06 0.766 0.00013592 0.742 0.055597295 0.819 AZ31-PLA-AMP 189.46 331.88 9239.58 17022.6 5.70043E-07 0.809 3.88007E-05 0.801 0.042768959 1 126 6.4.3.2 Immersion test It is well known that immersing magnesium alloys in aqueous solution can result in series of anodic dissolution reactions of magnesium alloys. Along with the degradation of magnesium alloy in physiological condition, the hydrogen evolution and alkalization of solution will simultaneously occur. Thus, to evaluate the in vitro degradation behavior, the uncoated and coated samples were immersed in SBF for two weeks to monitor the variations of mass loss and pH values of SBF solution. Fig. 6-8 (a) shows the weight loss of bare and coated AZ31 samples as a function of immersion time. It can be seen that the bare AZ31 presents a rapid degradation at the initial stage of immersion in SBF. Then the degradation rate gradually becomes moderate. In contrast, the PLA coated sample and nAMP/PLA coated sample exhibit a much slower degradation in SBF solution. After 15 days incubation in SBF, the mass loss of nAMP/PLA coated sample is one fifth of the weight that bare AZ31 lost. Fig. 6-8 (b) shows the variation of pH value of SBF solution containing the bare and coated AZ31 samples during the 15 days immersion test. Once the AZ31 samples are immersed in SBF, the pH value of SBF solution rises due to severe corrosion reactions of the magnesium alloy substrate. It can be seen the SBF solution incubating bare AZ31 shows a pH value of 8.4 and 8.2 after 1 day and 15 days accordingly. However, SBF solution maintains a relatively lower pH value with the incubation of coated samples. The pH values of SBF solution containing PLA coated and nAMP/PLA coated sample stabilize at 7.9 and 7.8 accordingly. Therefore, it can be speculated that the galvanic reactions of magnesium substrate in SBF are significantly retarded by the nAMP/PLA composite coatings. 127 Fig. 6-9 shows the surface features of PLA coated and nAMP/PLA coated samples after 15 days immersion in SBF. As shown in the figure, no visible micro cracks are identified on both of the samples, which indicates long term corrosion resistance of the as-prepared polymeric and composite films. It is also noticed the sphere like apatite particles scattered on the PLA coating during the immersion in SBF, as similar features of apatite formation on biodegradable polymer and AMP particles have been reported before [195, 225, 239, 240]. In contrast to the fragmentary distribution of apatite precipitates on PLA coating, it is interesting to note that the nAMP/PLA coating is completely covered by apatite globules within 7 days immersion in SBF. EDS spectra also indicate higher amount of Ca and P based compounds accumulated on nAMP/PLA coated surface, compared to that of pristine coated surface. Thus, with the addition of nAMP particles, the bioactivity of PLA in SBF is remarkably promoted. The enhanced bioactivity of nAMP/PLA coated AZ31 sample can possibly be ascribed to following phenomena : 1) The degradation of AMP particle and constant release of Mg2+ can facilitate the nucleation of amorphous calcium phosphate (ACP), which is known as the active site for apatite deposition [109]; 2) The incorporation of AMP in PLA matrix can improve the hydrophilicity of PLA film and subsequently promote the apatite deposition [241]; 3) The Mg2+ may inhibit the crystallization and agglomeration of apatite precipitates, which results a uniform coverage of apatite layer on PLA matrix [228, 241]. Overall, the findings evidence the extraordinary ability of nAMP/PLA composite coating to induce rapid biomineralization, which is crucial for the implant osseointegration and fixation. 128 Figure. 6-8 (a) Weight loss of coated samples and control sample and (b) variations of pH values of SBF solution at various immersion periods in SBF. Figure. 6-9 Surface morphology and corresponding EDS spectra of (a) PLA and (b) nAMP/PLA coated samples after 15 days immersion in SBF. 129 6.5 Conclusion The homogenous and crack-free PLA film and nAMP/PLA nanocomposite film were successfully spin coated onto biodegradable AZ31 magnesium alloy. The PLA film showed an average coating thickness of 500 nm. With the well dispersed AMP nano- particles in PLA matrix, the thickness of nAMP/PLA composite film slightly increased to 700 nm. Both PLA and nAMP/PLA coated AZ31 samples underwent significantly low degradation in SBF, according to the results of electrochemical measurements by potentiodynamic and EIS, and immersion test. Especially, with the addition of nano-sized AMP particles (20 wt.% of PLA) in PLA film, the in vitro corrosion resistance and bioactivity of AZ31 samples were further improved. Therefore, the nAMP/PLA nanocomposite film is anticipated to be a promising coating material for magnesium alloys to tailor their degradation behavior and promote the biological responses. 130 Chapter 7 Microwave Assisted Coating of Bioactive Amorphous Magnesium Phosphate (AMP) on Polyetheretherketone (PEEK) 7.1 Abstract Polyetheretherketone (PEEK) with great thermal and chemical stability, desirable mechanical properties and promising biocompatibility is being widely used as spine, orthopedic and dental implant materials. However, the bioinert surface of PEEK can hinder direct osseointegration between the host tissue and PEEK based implants. The important signatures of this paper are as follows. First, we report for the formation of osseointegrable amorphous magnesium phosphate (AMP) coating on PEEK surface using microwave energy. Second, coatings consist of nano-sized AMP particles with a stacked thickness of 800nm. Third, coatings enhance bioactivity in-vitro and induce significantly high amount of bone-like apatite coating, when soaked in simulated body fluid (SBF). Fourth, the as- deposited AMP coatings present no cytotoxicity effects and are beneficial for cell adhesion at early stage. Finally, the high levels of expression of osteocalcin (OCN) in cells cultured on AMP coated PEEK samples indicate that AMP coatings can promote new bone formation and hence osseointegration. 131 7.2 Introduction Since late 1990s, Polyetheretherketone ((-C6H4-O-C6H4-O-C6H4-CO-)n; PEEK) had emerged as a promising implant biomaterial for trauma, orthopedic and spine operations, due to its outstanding chemical and wear resistance, thermoplastic property and non- toxicity. In addition, mechanical properties of PEEK, such as elastic modulus, tensile strength, stiffness and fatigue are comparable to those of cortical bones [242-245]. On account of its high thermal, chemical, radiation and mechanical stability, PEEK can be effectively sterilized and shaped to desirable structure [242]. Moreover, the radiolucency of PEEK can facilitate the clear assessment of status of bony fusion [242, 243]. The surfaces of PEEK are biologically inert, thus presenting limited ability to directly bond with the surrounding tissues. The bioinert nature of PEEK significantly hinders its clinical applications in situations where osseointegration is critical. One strategy to address this drawback is to incorporate the hydroxyapatite (HA) as a filler in PEEK matrix or deposit HA layer on PEEK surface to promote the osseointegration of PEEK [246-254]. Apatitic phases is a group of bioactive ceramics, have been extensively investigated as the coating materials on metallic implants, because of its great similarity to bone minerals and excellent osteoconductive properties [144]. Alternatively, various techniques have been proposed and explored to deposit HA or other forms of calcium phosphate coatings on PEEK surface, such as thermal plasma spray [246], cold spray [249], spin coating [252], chemical deposition [255], sputtering [256], ion beam assisted deposition [257], biomimetic [258] and microwave assisted coating [124]. It’s noteworthy that the PEEK surface can also be modified by chemical and physical methods, such as sulfonation and plasma treatment, to enhance apatite deposition in physiological condition [259, 260]. 132 Recently, our group developed a novel microwave assisted coating process, which has been demonstrated to be a promising method to perform rapid surface modifications of various bioimplants [115, 119, 124]. It is known that thermal plasma spray involves high temperature, which can damage polymer substrate [256]. In comparison, the physical deposition techniques like spray and sputtering are only applicable to 2-dimensional structure. Although chemical deposition does not have the aforementioned limitations, the process, especially biomimetic technique is fairly time-consuming and susceptible to bacterial contamination. The microwave assisted coating process can deposit uniform bioactive coating on the polymer substrate with complex shape at relatively low temperatures, and within a short time-scale. Of late, compounds in the MgO-P2O5 binary system are getting a great deal of attention in potential applications in orthopedics and dentistry due to their optimal combination of biocompatibility and biodegradability. A fundamental reason for their success is that magnesium is the fourth most abundant element within the human body. In spite of that, research activities on magnesium phosphates are at still at a relatively smaller scale, as compared to their counterparts, calcium phosphates. Among magnesium phosphates, two crystalline compounds, such as struvite (NH4MgPO4· 6H2O) and newberyite (MgHPO4· 3H2O) are front runners in orthopedic applications. Their main applications have been in developing non-exothermic, self-setting, orthopedic cement compositions with high compressive strength, biocompatibility and biodegradability [181, 183, 261-263]. As opposed to crystalline magnesium phosphates, our group has directed concentrated effort in exploring amorphous magnesium phosphate (AMP) [109, 184-186]. We reported that the amorphous phosphate nano-particles elicit high levels of cellular response and gene 133 expression [185]. However, the variety of studies regarding the synthesis and further application of magnesium phosphate is still limited. To best of our knowledge, this is the first attempt to use amorphous magnesium phosphate (AMP) layer as the bioactive interface between polymer implant and host tissue. The objectives of this study are 1) synthesis of AMP coating on PEEK substrate via microwave assisted coating process, 2) investigate the in-vitro behavior such as bioactivity and cytocompatibility of the AMP coated PEEK. 7.3 Experimental 7.3.1 Material preparation Medical grade PEEK disks (Orchid implant solutions, Michigan) with a dimension of Ø10 x 2 mm3 were employed in this study. All the samples were ground up to #1200 SiC abrasive paper to ensure the smooth surfaces and ultrasonically cleaned in acetone for 10 min to remove the attached debris. Then the PEEK disks were divided into two groups. One group was incubated in 10 M NaOH for 48 h. The other batch of PEEK samples was immersed in concentrated sulfuric acid (95-98 wt.%, Fihser Sci) for 10 min to produce a uniform porous structure and rinsed with deionized (DI) water subsequently. After the pre- treatment, the first batch of PEEK samples was denoted as PEEK-OH and the second batch was denoted as PEEK-S. 7.3.2 Coating preparation The coating bath was prepared by dissolving 2.033g MgCl2 and 1.1998g NaH2PO4 in 200 ml DI water. 1M NaOH solution was used to adjust the pH value of coating solution to 6.8. The pretreated PEEK disks were placed in a 200 ml pyrex beaker, which was filled 134 with 100 ml coating solution. Then the beaker with PEEK samples and coating bath were irradiated in a microwave oven (Panasonic) at 1200 W power for 5 min. To ensure the uniformity of deposited layer, the above coating procedure was replicated. The coated PEEK samples were subsequently taken out from coating bath and rinsed with DI water followed by air drying. 7.3.3 Characterization and bioactivity evaluation The contact angle measurements were carried out using a contact angle meter (Model CAM-MTCRO, Tantec) to evaluate the wettability of samples. The surface morphologies and elemental compositions of the pretreated and AMP coated specimens were analyzed using scanning electron microscopy (SEM, S4800, Hitachi) coupled with an energy dispersive X-ray spectroscopy (EDS, Oxford INCA). Elemental analyses were performed at an accelerating voltage of 20KV with the working distance of 15mm. Phase compositions of the samples were identified by X-ray diffraction (XRD, Ultima III, Rigaku)) with monochromated Cu Kα radiation (44KV, 40mA) over a 2θ range of 10– 45°. Functional groups present in as-deposited coatings were characterized by a Fourier Transform Infrared Spectroscopy (FTIR, UMA-600 Microscope, Varian Excalibur Series) using an ATR diamond crystal for 256 scans at the range between 4000-700 cm-1 with a resolution of 1cm-1. The transmission electron microscopy (TEM, HD-2300, Hitachi, USA) was employed to investigate the microstructure of as-deposited AMP coatings. To evaluate the in vitro bioactivity, both coated and uncoated samples were immersed in simulated body fluid (t-SBF) at 37°C for 7 days. The ion concentrations of t-SBF that better mimic human blood plasma [164]. In addition, the supersaturation conditions were 135 maintained by replenishing the t-SBF solution every other day. After immersion, the samples were rinsed with deionized water and dried in air for further characterization of SEM and EDS. 7.3.4 Cell viability assay MC3T3-E1 (CRL-2593 ™ , ATCC, Manassas, VA, USA) preosteoblast cells were employed to study the effects of AMP coatings on preosteoblast proliferation and differentiation. The preosteoblast cells were initially cultured in alpha minimum essential medium (α-MEM, Thermo Scientific HyClone), augmented with 10% Fetal Bovine Serum (FBS, Thermo Scientific HyClone) at 37 °C in a humidified atmosphere of 5% CO2. The culture medium was replenished every other day until the cell reached a confluence of 90%. α-MEM culture medium acted as a control medium. For cytotoxicity test, MC3T3-E1 cells were seeded to the sterilized samples in a 24-wells cell culture plate (Flacon ™ BD Biosciences, USA) at a density of 10000 /well for 24h to allow attachment. After further incubation for 3 days, the cells were treated with 3-(4,5-dimethylthiazol-2-yl)-2,5- diphenyltetrazolium bromide (MTT, Sigma-Aldrich, St. Louis, MO, USA) for 4 h. The formazan precipitate was dissolved in dimethyl sulfoxide (DMSO) and measured using a microplate reader at a wavelength of 570 nm to assess live cells. To evaluate the expression of osteogenesis-related genes, some cells were subject to RNA isolation after 7 days incubation by the TRIzol regent (Invitrogen, Carlsbad, CA, USA), and the total RNA was reversely transcribed to complementary DNA (cDNA) using M-MLV reverse transcriptase (Promega, Madison, WI, USA). The expressions of the alkaline phosphatase (ALP) and osteocalcin (OPN) were quantified using real-time 136 polymerase chain reaction (PCR) with SsoFast EvaGreen Supermix (Bio-Rad, Hercules, CA, USA) and a 2-step amplification program (30 s at 95 °C and 60 s at 62 °C) performed on an Eppendorf Realplex thermal cycler. The relative mRNA expression level of each gene was determined by the CT (cycle threshold) values with normalization to the house- keeping gene glyceraldehyde-3- phosphate dehydrogenase (GAPDH). The forward and reverse primers for targeted genes are listed as follows: osteocalcin (OCN; forward 5′- GCAATAAGGTAGTGAACAGACTCC-3′ and reverse 5′- CTTTGTAGGCGGTCTTCAAGC-3′) and alkaline phosphatase (ALP; forward 5′- ATCTTTGGTCTGGCTCCCATG-3′ and reverse 5′-TTTCCCGTTCACCGTCCAC-3′). 7.3.5 Statistical analysis All test results are represented by means ± SD with at least in triplicate. One-way analysis of variance (one way-ANOVA) with turkey test was employed to determine the statistical difference between groups and ρ< 0.05 is considered to be significant. 7.4. Results and discussion 7.4.1 Characterization of un-coated and AMP coated PEEK samples In the present study, we use two different surface modification techniques to make PEEK’s surface prone to functional group attachment. 1) Alkali treatment, wherein OH- groups are deposited on the polymer surface, 2) Acid treatment, which makes the substrate surface porous, both of which activate PEEK’s surface with negative charge and enhance fixation of functional groups [124, 259, 264-266]. Figure 7-1(a-c) shows the surface features of NaOH and sulfuric acid treated samples. It can be clearly seen in Figure 7-1(a), 137 that 48 h NaOH treatment did not alter the surface morphology of PEEK. Even, the scratches due to the grinding process are visible. Whereas, 10 minutes of sulfuric acid treatment and subsequent immersion in water formed numerous pits of varying sizes on the PEEK surface (Figure 7-1b). The size of the generated pits ranged from 100 nm to 2 µm, which indicates that sulfuric acid treatment was able to modify the PEEK substrate and resulted in nanostructured surface on PEEK. A higher magnification image (Figure 7- 1c) of the nanostructured acid treated PEEK surface revealed that they consist of a stretched, three-dimensional, highly porous structure. Outward diffusion of excess sulfuric acid from the polymer surface during water immersion and alterations in the original chemical structure due to chemical introduction of SO3H group are the reasons for the formation of nanostructured pores. The SO3H group also causes a prolonged swelling of polymer chain which results in the formation of a three-dimensional network [259, 266]. Figure. 7-1 Surface morphology of (a) NaOH, (b) and (c) sulfuric acid treated PEEK samples. 138 Figure. 7-2 Water contact angle of samples, * represents the p 0.05 compared to bare PEEK. The surface hydrophilicity of the bare and treated PEEK samples was measured using contact angle meter (Model CAM-MTCRO, Tantec). Figure 7-2 illustrates the water contact angel measurements of various samples. Compared to untreated PEEK, the water contact angle of NaOH etched sample decreased from 74° to 43°. Conversely, the water contact angel of sulfuric acid treated PEEK sample increased to 97°. In spite of the acid treatment which introduced hydrophilic SO3H groups, the surface of PEEK-S exhibited significant hydrophobicity. This is because of the pronounced effect of the altered surface morphology of PEEK, which then consisted of nanostructured three-dimensional porous surface. After the coating treatment, the PEEK-OH and PEEK-S samples showed a much lower water contact angle of 25° and 12° accordingly, indicating that the as-deposited coating on the polymer surface is remarkably hydrophilic in nature. The X-ray diffraction patterns of the specimens are shown in Figure 7-3a. Only the diffraction peaks corresponding to PEEK are identified. This demonstrates that the as- 139 deposited magnesium phosphate coating presents an amorphous structure. To further characterize the control and coated PEEK samples, FTIR analysis was conducted. As shown in Figure 7-3b, the AMP coated PEEK-OH and PEEK-S samples exhibit a -1 3- significant absorption peak at 1095 cm , which can be assigned to phosphate group (PO4 ) present in amorphous magnesium phosphate phase. In addition, the broad absorption band located at around 3250 cm-1 originates from the structural water of AMP coatings. The presence of phosphate and absorbed water band confirms the formation of amorphous magnesium phosphate layer on PEEK surface. Figure. 7-3 XRD patterns (a) and FTIR spectra (b) of bare and treated PEEK samples. The SEM images of the coated PEEK samples are presented in Figure 7-4(a-b). The micrographs clearly indicate that the surfaces of PEEK-OH and PEEK-S samples are covered with micro-sized amorphous magnesium phosphate (AMP) flakes. Although coating surface of PEEK-S-AMP sample is rougher, both the AMP coated PEEK-OH and PEEK-S samples present comparable morphology. The representative EDS result reveals 140 that the coating layer is composed of Mg and P, which further demonstrates the homogeneous deposition of AMP particles on PEEK surface. Figure. 7-4 SEM images and EDS spectra of AMP coated PEEK samples (a) PEEK- OH-AMP and (b) PEEK-S-AMP. The cross-section of PEEK-OH-AMP and PEEK-S-AMP samples were prepared using focused ion beam and characterized using scanning electron microscopy (SEM). As shown in Figure 7-5(a-b), the thickness of AMP coating is around 800 nm. In addition, AMP coating on both PEEK-OH and PEEK-S samples exhibit a porous structure. It is noteworthy that the diffusion pathways of sulfuric acid are observed in cross-sectional image of PEEK-S-AMP sample. Moreover, the pores present in the top layer of PEEK-S sample are narrowed due to the accumulation of the AMP particles. Figure 7-5c shows the TEM image and corresponding selected area electron diffraction (SAED) pattern. 141 According to the TEM image, it can be seen that the AMP flakes are comprised of numerous nano sized AMP spheres. The SAED pattern presents several diffused rings, which is indicative of amorphous nature of as-deposited AMP coating. Figure. 7-5 Cross-sectional SEM images of AMP coated PEEK samples (a) PEEK- OH-AMP, (b) PEEK-S-AMP and (c) TEM image associated with SAED of AMP coatings. 7.4.2 Deposition mechanism With the continuous release of Na+ and H+ ions from the alkaline and acid etched PEEK - - surfaces in coating bath, the negatively charged groups consisted of OH or SO3 rapidly accumulated on the pretreated PEEK surfaces, which result in an overall negative charge on PEEK surface. Consequently, the negatively charged PEEK surface will selectively combine with the plentiful of Mg2+ ions in the coating liquid. Therefore, abundant nucleation sites of AMP precipitates spread over the surface of PEEK and generate overall positive charge. Finally, Mg2+ ions incorporate with negatively charged phosphate ions to form AMP particles and produce a uniform AMP layer on the PEEK surface. Additionally, the whole coating process is significantly promoted by the microwave energy, which not only elevates the AMP nucleation rate by orders of magnitude, but also ensures a 142 homogeneous deposition of AMP particles [10, 119, 124]. On account of abovementioned benefits associated with microwave assisted coating technique, the bioactive AMP coatings can be deposited onto PEEK surface in minutes, rather than hours consumed in the conventional biomimetic process. 7.4.3 In vitro properties Figures 7-6(a-c) shows the SEM images of SBF immersed bare and AMP coated PEEK samples. As seen in Figure 7-6a, very few particles of spherical appearance appear on the surface of bare PEEK upon immersing it in SBF for 7 days. On the contrary, Figures 7- 6(b-c) confirms the formation of a dense coating on the surface of SBF immersed PEEK- OH-AMP and PEEK-S-AMP samples. A careful analysis of the micrographs reveals that the spherical particles deposited on bare PEEK and the dense coating formed on AMP coated samples consist of petal-like morphology. The EDS results (Figure 7-6d) of the dense coating confirm the presence of high amounts of Ca and P, which can be identified as the major compositions of apatite precipitates [119, 166]. However, as-deposited AMP coating mostly disappeared after SBF immersion, which indicates the high degradation rate of AMP particles. In the present immersion study, we increased the rigor of the experimental conditions by not increasing the strength of t-SBF to 1.5 which is the norm. Moreover, the soaking period was only for 7 days, instead of the conventional 21 days. To our pleasant surprise, the AMP coated surface was completely covered with apatite within 7 days with similar features to reports presented by Kokubo and Takadama as well as in several studies from our group [164, 267] . Compared to the scarce spherical apatite particles on bare PEEK, it needs to infer the high bioactivity of AMP coatings on PEEK substrates and its great potential to promote biomineralization in vitro. A final comment is 143 about the high degradation rate of AMP. The degradation of bioactive AMP and the formation of bone-like apatite is a simultaneous mechanism. So it can be concluded that the simultaneous mechanism of rapid bone growth will compromise the degradation and help in osseointegration within a very short time (e.g. 7 days). Figure. 7-6 SEM images of PEEK samples after 1 week SBF immersion (a) Bare PEEK, (b)PEEK-OH-AMP, (c)PEEK-S-AMP and (d) EDS spectra of apatite formed on PEEK-S-AMP surface during SBF immersion. 7.4.4 Cytocompatibility and RT-PCR As shown in Fig. 7-7, the cell viability measured by MTT assay presents significant increase (p 0.05) for PEEK-S-AMP sample after 3 days. It’s revealed that the AMP coating incorporating with 3D porous structure of the PEEK substrate is favorable for pre- osteoblast cells attachment. The comparable cell numbers on the bare PEEK and PEEK- 144 OH-AMP specimens can be the result of saturation on the disk surface, as reported by Lee et al [249]. However, the PEEK-S-AMP sample maintains a much larger surface area compared to that of bare PEEK sample. This could further enhance the attachment of cells. Additionally, the osteogenic differentiation properties of bare and coated PEEK samples were assessed by quantitative real-time RT-PCR of ALP and OCN mRNA expression and the results are illustrated in Fig. 7-8(a-b). ALP is generally regarded as the early marker for osteoblast differentiation [268]. The expression of ALP at early time points can be recognized as the rapid differentiation of osteoblast cells. The ALP activities for the coated PEEK samples are approximate to that of the control and bare PEEK after 7 days. This suggests that the AMP coating does not induce deleterious effects on the ALP expressions and suppress ALP activities at the early stage of cell proliferation. The similar low expression level of ALP regarding the AMP particles had been reported before [228]. In addition, the crystallized magnesium phosphate compounds such as newberyite and cattiite also present significantly lower level expression of ALP activities after 7 days culture, compared to that of calcium phosphates [108]. The fast release of magnesium and consumption of calcium in the biomineralization process could lower the β- glycerophosphate (BGP) induced ALP activity at some level [228, 269]. Interestingly, compared to untreated PEEK sample, the expression levels of OCN are greatly promoted for PEEK-OH-AMP and PEEK-S-AMP sample. OCN is a late-stage marker of osteoblast differentiation, which implies the AMP coating would enhance the extracellular matrix (ECM) mineralization [270] and new bone formation. The relatively high level of OCN expression rather than ALP expression could be attributed to the enhanced and accelerated biomineralization of AMP coated PEEK in vitro. Moreover, it’s noteworthy magnesium 145 phosphate compounds, especially amorphous magnesium phosphate possesses comparable bone regenerative capacity to that of calcium phosphates, which suggests the AMP coating could be promising candidate to serve as the osteoconductive layer on implant substrates [108, 185]. Figure. 7-7 Cell viability of MC3T3-E1 pre-osteoblasts cultured on the PEEK control, PEEK-OH-AMP and PEEK-S-AMP for 3 days. * represents the p 0.05 compared to bare PEEK. 146 Figure. 7-8 Osteogenic differentiation by measuring the mRNA expression level of alkaline phosphatase (ALP) and osteocalcin (OCN) after 7 days. 7.5 Conclusion In this study, we successfully deposited a bioactive coating comprising of amorphous magnesium phosphate (AMP) via a microwave assisted coating technique in 10 minutes. The SEM, EDS, FTIR, XRD and TEM results confirmed the homogeneous coverage and amorphous nature of as-deposited AMP layers. The in-vitro assessments demonstrated that the AMP coatings significantly promoted the biomineralization activities. In addition, the AMP coatings incorporating with 3D structured PEEK substrate presented superior ability to improve the attachment of preosteoclast cells. The high level of bone formation related gene expression of OCN revealed that AMP coating is beneficial to the remodeling of bones. Thus, it is anticipated that PEEK coated surfaces with AMP would be quite suitable for clinical applications as orthopedic and dental implants. 147 Chapter 8 Conclusion and Future Directions 8.1 Conclusion The major goal of this thesis is to develop the microwave assisted rapid surface modification techniques and produce promising biocompatible coatings that could effectively improve the degradation behavior and biological properties of biomedical implant materials. To have a good understanding of microwave chemistry, progress in microwave processing of biomaterials and explore potential applications of microwave assisted synthesis methods in biomaterials field, the chapter 2 outlined interactions between microwave and materials, thoroughly summarized the benefits and advances in microwave assisted synthesis of bioceramics, and insightfully pointed out the possibility of using microwave assisted process to conduct the rapid surface modification of bioimplant materials. Thus, various types of protective and bioactive coatings were developed on AZ31Mg alloy via a microwave assisted coating technique. For example, chapter 3 and chapter 4 reported the preparation of calcium phosphate (CaP) based calcium deficient hydroxyapatite coatings, magnesium phosphate based newberyite and newberyite- trimagnesium phosphate coatings on AZ31 Mg alloy substrate. It was found that both initial 148 Ca/P ratio and treatment temperature play key roles in the fabrication of CaP/MgP coatings and can critically alter the structure and phase composition of as-deposited coatings. More importantly, as-prepared CDHA and MgP coatings significantly improved the corrosion resistance and in vitro properties of AZ31 Mg alloy, and tailored it more suitable for future biomedical applications. By taking the advantage of synergistic effects of organic and inorganic biomaterials, the composite coatings involve the alkaline earth phosphate ceramic and biocompatible polymer PLA were designed, synthesized and reported in chapter 5 and chapter 6. The hybrid PLA/FHA and nanocomposite nAMP/PLA coating were prepared by incorporating the microwave assisted deposition process with spin coating. The PLA polymer with compact microstructure can functionally prevent the penetration of corrosive medium present in SBF. In addition, the FHA and nAMP serve as bioactive agents can simultaneously promote the biomineralization capability of Mg alloys in vitro. In chapter 7, the feasibility of employing microwave assisted coating technique in surface modification of polymeric implant materials instead of metallic implant materials was investigated. Herein, the highly bioactive amorphous magnesium phosphate coating was deposited on the surface of biopolymer PEEK using microwave energy. The results revealed that the AMP coating associated with proper pretreatment notably promoted the biological and cellular responses of PEEK. 8.2 Future directions Current work has been focusing on the in vitro evaluation of biodegradable Mg alloys with diverse surface modifications. In vitro evaluation is an essential step toward in vivo 149 trials and further practical applications. However, in vivo assessment is quite different to in vitro evaluations, as in vivo system presents much higher tolerance regarding the degradation activities of Mg alloys. Moreover, the degradation rate can be remarkably affected by the composition of electrolyte and the interaction between Mg alloys and specific electrolyte. Therefore, it is fundamental to understand the degradation mechanism of surface modified Mg alloy samples in vivo and corresponding tissue remodeling process before we can safely employ Mg alloy based implants in human body. Despite the significance of biological aspects and surface chemistry of implants has been addressed, the mechanical properties of surface modified bioimplants also play a crucial role in their long-term successes. To avoid the early loosening and premature failure, the wear behaviors of coated implants for load carrying applications need to be properly regulated. This requires the as-deposited coatings to be strongly adhered to the substrates and protect bulk implants from excessive wear such as abrasion and fatigue. Hence, future work should aim to investigate the adhesive strength and wear resistance of alkaline earth phosphate coated biomedical implants. Magnesium phosphate is a good alternative to the extensively studied calcium phosphate, and presented outstanding biological properties. However, the research and application of MgP based materials in biomedical field are still in a nascent phase. As reported in chapter 4 and chapter 7, MgP is a promising candidate for the protective and bioactive coating to control the degradation behavior of Mg alloys and improve biological properties of bioimplant materials. Thus, it is attractive to explore further applications of MgP such as coating for polymer scaffolds, filler in biocomposites and drug delivery systems. 150 References [1] P.V. Giannoudis, H. Dinopoulos, E. Tsiridis, Bone substitutes: an update, Injury 36(3) (2005) S20-S27. [2] S. Kurtz, F. Mowat, K. Ong, N. Chan, E. Lau, M. Halpern, Prevalence of primary and revision total hip and knee arthroplasty in the United States from 1990 through 2002, JBJS 87(7) (2005) 1487-1497. [3] S.B. Goodman, Z. Yao, M. Keeney, F. Yang, The future of biologic coatings for orthopaedic implants, Biomaterials 34(13) (2013) 3174-3183. [4] E. Thostenson, T.-W. Chou, Microwave processing: fundamentals and applications, Composites Part A: Applied Science and Manufacturing 30(9) (1999) 1055-1071. [5] D.E. Clark, W.H. Sutton, Microwave processing of materials, Annual Review of Materials Science 26(1) (1996) 299-331. [6] E. Grant, B.J. Halstead, Dielectric parameters relevant to microwave dielectric heating, Chemical society reviews 27(3) (1998) 213-224. [7] M. Bhattacharya, T. Basak, A review on the susceptor assisted microwave processing of materials, Energy 97 (2016) 306-338. 151 [8] S.H. Jhung, T. Jin, Y.K. Hwang, J.S. Chang, Microwave effect in the fast synthesis of microporous materials: which stage between nucleation and crystal growth is accelerated by microwave irradiation?, Chemistry-A European Journal 13(16) (2007) 4410-4417. [9] C.O. Kappe, B. Pieber, D. Dallinger, Microwave effects in organic synthesis: myth or reality?, Angewandte Chemie International Edition 52(4) (2013) 1088-1094. [10] Y.-J. Zhu, F. Chen, Microwave-Assisted Preparation of Inorganic Nanostructures in Liquid Phase, Chemical Reviews 114(12) (2014) 6462-6555. [11] I. Bilecka, M. Niederberger, Microwave chemistry for inorganic nanomaterials synthesis, Nanoscale 2(8) (2010) 1358-1374. [12] W. Suchanek, M. Yoshimura, Processing and properties of hydroxyapatite-based biomaterials for use as hard tissue replacement implants, Journal of Materials Research 13(1) (1998) 94-117. [13] S.V. Dorozhkin, Bioceramics of calcium orthophosphates, Biomaterials 31(7) (2010) 1465-1485. [14] S.V. Dorozhkin, Calcium orthophosphates in nature, biology and medicine, Materials 2(2) (2009) 399-498. [15] S.V. Dorozhkin, M. Epple, Biological and medical significance of calcium phosphates, Angewandte Chemie International Edition 41(17) (2002) 3130-3146. [16] F.A. Tylavsky, L.A. Spence, L. Harkness, The importance of calcium, potassium, and acid-base homeostasis in bone health and osteoporosis prevention, The Journal of nutrition 138(1) (2008) 164S-165S. 152 [17] R. Takeda, T. Nakamura, Effects of high magnesium intake on bone mineral status and lipid metabolism in rats, Journal of nutritional science and vitaminology 54(1) (2008) 66-75. [18] A. Siddharthan, S. Seshadri, T.S. Kumar, Rapid synthesis of calcium deficient hydroxyapatite nanoparticles by microwave irradiation, Trends Biomater Artif Organs 18(2) (2005) 110-113. [19] E. Lerner, S. Sarig, R. Azoury, Enhanced maturation of hydroxyapatite from aqueous solutions using microwave irradiation, Journal of Materials Science: Materials in Medicine 2(3) (1991) 138-141. [20] S. Sarig, F. Kahana, Rapid formation of nanocrystalline apatite, Journal of crystal growth 237 (2002) 55-59. [21] R. Murugan, S. Ramakrishna, Aqueous mediated synthesis of bioresorbable nanocrystalline hydroxyapatite, Journal of Crystal Growth 274(1) (2005) 209-213. [22] C. Silva, M. Graça, M. Valente, J. Góes, A. Sombra, Microwave preparation, structure and electrical properties of calcium–sodium–phosphate biosystem, Journal of non- crystalline solids 352(32) (2006) 3512-3517. [23] C. Qi, Q.-L. Tang, Y.-J. Zhu, X.-Y. Zhao, F. Chen, Microwave-assisted hydrothermal rapid synthesis of hydroxyapatite nanowires using adenosine 5'-triphosphate disodium salt as phosphorus source, Materials Letters 85 (2012) 71-73. [24] X.Y. Zhao, Y.J. Zhu, C. Qi, F. Chen, B.Q. Lu, J. Zhao, J. Wu, Hierarchical Hollow Hydroxyapatite Microspheres: Microwave‐Assisted Rapid Synthesis by Using Pyridoxal‐ 5′‐Phosphate as a Phosphorus Source and Application in Drug Delivery, Chemistry–An Asian Journal 8(6) (2013) 1313-1320. 153 [25] C. Qi, Y.J. Zhu, B.Q. Lu, X.Y. Zhao, J. Zhao, F. Chen, J. Wu, Hydroxyapatite Hierarchically Nanostructured Porous Hollow Microspheres: Rapid, Sustainable Microwave‐Hydrothermal Synthesis by Using Creatine Phosphate as an Organic Phosphorus Source and Application in Drug Delivery and Protein Adsorption, Chemistry- A European Journal 19(17) (2013) 5332-5341. [26] T.T. Demirtaş, G. Kaynak, M. Gümüşderelioğlu, Bone-like hydroxyapatite precipitated from 10× SBF-like solution by microwave irradiation, Materials Science and Engineering: C 49 (2015) 713-719. [27] A. López‐Macipe, J. Gómez‐Morales, R. Rodríguez‐Clemente, Nanosized hydroxyapatite precipitation from homogeneous calcium/citrate/phosphate solutions using microwave and conventional heating, Advanced Materials 10(1) (1998) 49-53. [28] J. Torrent‐Burgués, J. Gómez‐Morales, A. López‐Macipe, R. Rodríguez‐Clemente, Continuous precipitation of hydroxyapatite from Ca/citrate/phosphate solutions using microwave heating, Crystal Research and Technology 34(5‐6) (1999) 757-762. [29] H. Arami, M. Mohajerani, M. Mazloumi, R. Khalifehzadeh, A. Lak, S. Sadrnezhaad, Rapid formation of hydroxyapatite nanostrips via microwave irradiation, Journal of Alloys and Compounds 469(1) (2009) 391-394. [30] J.M. Cao, J. Feng, S.G. Deng, X. Chang, J. Wang, J.S. Liu, P. Lu, H.X. Lu, M.B. Zheng, F. Zhang, Microwave-assisted solid-state synthesis of hydroxyapatite nanorods at room temperature, Journal of materials science 40(23) (2005) 6311-6313. [31] Z. Zyman, A. Goncharenko, D. Rokhmistrov, M. Epple, Nanocrystalline calcium‐ deficient hydroxyapatite prepared by a microwave‐assisted solvent‐free reaction, Materialwissenschaft und Werkstofftechnik 42(2) (2011) 154-157. 154 [32] J. Liu, K. Li, H. Wang, M. Zhu, H. Yan, Rapid formation of hydroxyapatite nanostructures by microwave irradiation, Chemical physics letters 396(4) (2004) 429-432. [33] J. Liu, K. Li, H. Wang, M. Zhu, H. Xu, H. Yan, Self-assembly of hydroxyapatite nanostructures by microwave irradiation, Nanotechnology 16(1) (2004) 82. [34] A. Lak, M. Mazloumi, M.S. Mohajerani, S. Zanganeh, M.R. Shayegh, A. Kajbafvala, H. Arami, S.K. Sadrnezhaad, Rapid Formation of Mono‐Dispersed Hydroxyapatite Nanorods with Narrow‐Size Distribution via Microwave Irradiation, Journal of the American Ceramic Society 91(11) (2008) 3580-3584. [35] G.S. Kumar, A. Thamizhavel, E. Girija, Microwave conversion of eggshells into flower-like hydroxyapatite nanostructure for biomedical applications, Materials Letters 76 (2012) 198-200. [36] R. Vani, S.B. Raja, T. Sridevi, K. Savithri, S.N. Devaraj, E. Girija, A. Thamizhavel, S.N. Kalkura, Surfactant free rapid synthesis of hydroxyapatite nanorods by a microwave irradiation method for the treatment of bone infection, Nanotechnology 22(28) (2011) 285701. [37] T. Liang, J. Qian, Y. Yuan, C. Liu, Synthesis of mesoporous hydroxyapatite nanoparticles using a template-free sonochemistry-assisted microwave method, Journal of Materials Science 48(15) (2013) 5334-5341. [38] S. Jalota, A.C. Tas, S.B. Bhaduri, Microwave-assisted synthesis of calcium phosphate nanowhiskers, Journal of materials research 19(6) (2004) 1876-1881. [39] D.E. Wagner, J. Lawrence, S.B. Bhaduri, Microwave-assisted solution combustion synthesis of high aspect ratio calcium phosphate nanoparticles, Journal of Materials Research 28(22) (2013) 3119-3129. 155 [40] M. Elkady, M. Mahmoud, H. Abd-El-Rahman, Kinetic approach for cadmium sorption using microwave synthesized nano-hydroxyapatite, Journal of Non-Crystalline Solids 357(3) (2011) 1118-1129. [41] A. Siddharthan, S. Seshadri, T.S. Kumar, Microwave accelerated synthesis of nanosized calcium deficient hydroxyapatite, Journal of Materials Science: Materials in Medicine 15(12) (2004) 1279-1284. [42] S. Türk, İ. Altınsoy, G. ÇelebiEfe, M. Ipek, M. Özacar, C. Bindal, Microwave–assisted biomimetic synthesis of hydroxyapatite using different sources of calcium, Materials Science and Engineering: C 76 (2017) 528-535. [43] B.-T. Lee, M.-H. Youn, R.K. Paul, K.-H. Lee, H.-Y. Song, In situ synthesis of spherical BCP nanopowders by microwave assisted process, Materials Chemistry and Physics 104(2) (2007) 249-253. [44] D. Smolen, T. Chudoba, I. Malka, A. Kedzierska, W. Lojkowski, W. Swieszkowski, K.J. Kurzydlowski, M. Kolodziejczyk-Mierzynska, M. Lewandowska-Szumiel, Highly biocompatible, nanocrystalline hydroxyapatite synthesized in a solvothermal process driven by high energy density microwave radiation, International journal of nanomedicine 8 (2013) 653. [45] D.S.R. Krishna, A. Siddharthan, S. Seshadri, T.S. Kumar, A novel route for synthesis of nanocrystalline hydroxyapatite from eggshell waste, Journal of Materials Science: Materials in Medicine 18(9) (2007) 1735-1743. [46] K.W. Wang, Y.J. Zhu, X.Y. Chen, W.Y. Zhai, Q. Wang, F. Chen, J. Chang, Y.R. Duan, Flower‐Like Hierarchically Nanostructured Hydroxyapatite Hollow Spheres: Facile 156 Preparation and Application in Anticancer Drug Cellular Delivery, Chemistry–An Asian Journal 5(12) (2010) 2477-2482. [47] K.-W. Wang, Y.-J. Zhu, F. Chen, G.-F. Cheng, Y.-H. Huang, Microwave-assisted synthesis of hydroxyapatite hollow microspheres in aqueous solution, Materials Letters 65(15) (2011) 2361-2363. [48] T.J. Webster, E.A. Massa-Schlueter, J.L. Smith, E.B. Slamovich, Osteoblast response to hydroxyapatite doped with divalent and trivalent cations, Biomaterials 25(11) (2004) 2111-2121. [49] Y. Matsumura, S. Sugiyama, H. Hayashi, J.B. Moffat, Lead-calcium hydroxyapatite: cation effects in the oxidative coupling of methane, Journal of Solid State Chemistry 114(1) (1995) 138-145. [50] N. Rameshbabu, T. Sampath Kumar, T. Prabhakar, V. Sastry, K. Murty, K. Prasad Rao, Antibacterial nanosized silver substituted hydroxyapatite: synthesis and characterization, Journal of Biomedical Materials Research Part A 80(3) (2007) 581-591. [51] N. Iqbal, M.R.A. Kadir, N.A.N.N. Malek, N.H. Mahmood, M.R. Murali, T. Kamarul, Rapid microwave assisted synthesis and characterization of nanosized silver-doped hydroxyapatite with antibacterial properties, Materials Letters 89 (2012) 118-122. [52] N. Iqbal, M.R.A. Kadir, N.A.N.N. Malek, N.H.B. Mahmood, M.R. Murali, T. Kamarul, Characterization and antibacterial properties of stable silver substituted hydroxyapatite nanoparticles synthesized through surfactant assisted microwave process, Materials Research Bulletin 48(9) (2013) 3172-3177. 157 [53] M. Robles-Águila, J. Reyes-Avendaño, M. Mendoza, Structural analysis of metal- doped (Mn, Fe, Co, Ni, Cu, Zn) calcium hydroxyapatite synthetized by a sol-gel microwave-assisted method, Ceramics International 43(15) (2017) 12705-12709. [54] V.K. Mishra, B.N. Bhattacharjee, O. Parkash, D. Kumar, S.B. Rai, Mg-doped hydroxyapatite nanoplates for biomedical applications: a surfactant assisted microwave synthesis and spectroscopic investigations, Journal of Alloys and Compounds 614 (2014) 283-288. [55] K.T. Arul, J.R. Ramya, G. Bhalerao, S.N. Kalkura, Physicochemical characterization of the superhydrophilic, magnesium and silver ions co-incorporated nanocrystalline hydroxyapatite, synthesized by microwave processing, Ceramics International 40(9) (2014) 13771-13779. [56] J. Buehler, P. Chappuis, J. Saffar, Y. Tsouderos, A. Vignery, Strontium ranelate inhibits bone resorption while maintaining bone formation in alveolar bone in monkeys (Macaca fascicularis), Bone 29(2) (2001) 176-179. [57] N.D. Ravi, R. Balu, T. Sampath Kumar, Strontium‐Substituted Calcium Deficient Hydroxyapatite Nanoparticles: Synthesis, Characterization, and Antibacterial Properties, Journal of the American Ceramic Society 95(9) (2012) 2700-2708. [58] Z. Li, Z. Liu, M. Yin, X. Yang, Q. Yuan, J. Ren, X. Qu, Aptamer-capped multifunctional mesoporous strontium hydroxyapatite nanovehicle for cancer-cell- responsive drug delivery and imaging, Biomacromolecules 13(12) (2012) 4257-4263. [59] E.M. Carlisle, In vivo requirement for silicon in articular cartilage and connective tissue formation in the chick, The journal of nutrition 106(4) (1976) 478-484. 158 [60] S.K. Padmanabhan, E.U. Haq, A. Licciulli, Rapid synthesis and characterization of silicon substituted nano hydroxyapatite using microwave irradiation, Current Applied Physics 14(1) (2014) 87-92. [61] G. Wang, Q. Peng, Y. Li, Lanthanide-doped nanocrystals: synthesis, optical-magnetic properties, and applications, Accounts of chemical research 44(5) (2011) 322-332. [62] A. Escudero, M.E. Calvo, S. Rivera-Fernández, J.s.M. de la Fuente, M. Ocaña, Microwave-assisted synthesis of biocompatible europium-doped calcium hydroxyapatite and fluoroapatite luminescent nanospindles functionalized with poly (acrylic acid), Langmuir 29(6) (2013) 1985-1994. [63] R. André, E. Paris, M. Gurgel, I. Rosa, C. Paiva-Santos, M.S. Li, J.A. Varela, E. Longo, Structural evolution of Eu-doped hydroxyapatite nanorods monitored by photoluminescence emission, Journal of Alloys and Compounds 531 (2012) 50-54. [64] C. Yang, P. Yang, W. Wang, S. Gai, J. Wang, M. Zhang, J. Lin, Synthesis and characterization of Eu-doped hydroxyapatite through a microwave assisted microemulsion process, Solid State Sciences 11(11) (2009) 1923-1928. [65] D.E. Wagner, K.M. Eisenmann, A.L. Nestor-Kalinoski, S.B. Bhaduri, A microwave- assisted solution combustion synthesis to produce europium-doped calcium phosphate nanowhiskers for bioimaging applications, Acta biomaterialia 9(9) (2013) 8422-8432. [66] F. Chen, P. Huang, Y.-J. Zhu, J. Wu, C.-L. Zhang, D.-X. Cui, The photoluminescence, drug delivery and imaging properties of multifunctional Eu 3+/Gd 3+ dual-doped hydroxyapatite nanorods, Biomaterials 32(34) (2011) 9031-9039. 159 [67] A.Z. Alshemary, M. Akram, Y.-F. Goh, M.R.A. Kadir, A. Abdolahi, R. Hussain, Structural characterization, optical properties and in vitro bioactivity of mesoporous erbium-doped hydroxyapatite, Journal of Alloys and Compounds 645 (2015) 478-486. [68] H. Badran, I. Yahia, M.S. Hamdy, N. Awwad, Lithium-doped hydroxyapatite nano- composites: Synthesis, characterization, gamma attenuation coefficient and dielectric properties, Radiation Physics and Chemistry 130 (2017) 85-91. [69] W. Yu, T.-W. Sun, Z. Ding, C. Qi, H. Zhao, F. Chen, Z. Shi, Y.-J. Zhu, D. Chen, Y. He, Copper-doped mesoporous hydroxyapatite microspheres synthesized by a microwave- hydrothermal method using creatine phosphate as an organic phosphorus source: application in drug delivery and enhanced bone regeneration, Journal of Materials Chemistry B 5(5) (2017) 1039-1052. [70] A. Ashokan, D. Menon, S. Nair, M. Koyakutty, A molecular receptor targeted, hydroxyapatite nanocrystal based multi-modal contrast agent, Biomaterials 31(9) (2010) 2606-2616. [71] V.S. Chandra, K. Elayaraja, K.T. Arul, S. Ferraris, S. Spriano, M. Ferraris, K. Asokan, S.N. Kalkura, Synthesis of magnetic hydroxyapatite by hydrothermal–microwave technique: Dielectric, protein adsorption, blood compatibility and drug release studies, Ceramics International 41(10) (2015) 13153-13163. [72] V. Sarath Chandra, G. Baskar, R. Suganthi, K. Elayaraja, M. Ahymah Joshy, W. Sofi Beaula, R. Mythili, G. Venkatraman, S. Narayana Kalkura, Blood compatibility of iron- doped nanosize hydroxyapatite and its drug release, ACS applied materials & interfaces 4(3) (2012) 1200-1210. 160 [73] N. Petchsang, W. Pon-On, J. Hodak, I. Tang, Magnetic properties of Co-ferrite-doped hydroxyapatite nanoparticles having a core/shell structure, Journal of Magnetism and Magnetic Materials 321(13) (2009) 1990-1995. [74] E.Ö. Tunçay, T.T. Demirtaş, M. Gümüşderelioğlu, Microwave-induced production of boron-doped HAp (B-HAp) and B-HAp coated composite scaffolds, Journal of Trace Elements in Medicine and Biology 40 (2017) 72-81. [75] M. Abutalib, I. Yahia, Novel and facile microwave-assisted synthesis of Mo-doped hydroxyapatite nanorods: Characterization, gamma absorption coefficient, and bioactivity, Materials Science and Engineering: C 78 (2017) 1093-1100. [76] M. Shkir, M. Kilany, I. Yahia, Facile microwave-assisted synthesis of tungsten-doped hydroxyapatite nanorods: A systematic structural, morphological, dielectric, radiation and microbial activity studies, Ceramics International 43(17) (2017) 14923-14931. [77] I. Yahia, M. Shkir, S. AlFaify, V. Ganesh, H. Zahran, M. Kilany, Facile microwave- assisted synthesis of Te-doped hydroxyapatite nanorods and nanosheets and their characterizations for bone cement applications, Materials Science and Engineering: C 72 (2017) 472-480. [78] R. Zapanta-LeGeros, Effect of carbonate on the lattice parameters of apatite, Nature 206(4982) (1965) 403-404. [79] T.S. Kumar, I. Manjubala, J. Gunasekaran, Synthesis of carbonated calcium phosphate ceramics using microwave irradiation, Biomaterials 21(16) (2000) 1623-1629. [80] N. Rameshbabu, T.S. Kumar, K.P. Rao, Synthesis of nanocrystalline fluorinated hydroxyapatite by microwave processing and its in vitro dissolution study, Bulletin of Materials Science 29(6) (2006) 611-615. 161 [81] M. Nabiyouni, H. Zhou, T.J. Luchini, S.B. Bhaduri, Formation of nanostructured fluorapatite via microwave assisted solution combustion synthesis, Materials Science and Engineering: C 37 (2014) 363-368. [82] V. Orlovskii, V. Komlev, S. Barinov, Hydroxyapatite and hydroxyapatite-based ceramics, Inorganic Materials 38(10) (2002) 973-984. [83] H. Madupalli, B. Pavan, M.M. Tecklenburg, Carbonate substitution in the mineral component of bone: Discriminating the structural changes, simultaneously imposed by carbonate in A and B sites of apatite, Journal of Solid State Chemistry 255 (2017) 27-35. [84] R. LeGeros, S. Lin, R. Rohanizadeh, D. Mijares, J. LeGeros, Biphasic calcium phosphate bioceramics: preparation, properties and applications, Journal of materials science: Materials in Medicine 14(3) (2003) 201-209. [85] G. Daculsi, Biphasic calcium phosphate concept applied to artificial bone, implant coating and injectable bone substitute, Biomaterials 19(16) (1998) 1473-1478. [86] C.P. Klein, J. de Blieck-Hogemrst, J. Wolket, K. De Groot, Studies of the solubility of different calcium phosphate ceramic particles in vitro, Biomaterials 11(7) (1990) 509- 512. [87] I. Manjubala, M. Sivakumar, In-situ synthesis of biphasic calcium phosphate ceramics using microwave irradiation, Materials chemistry and physics 71(3) (2001) 272-278. [88] I. Manjubala, T. Sastry, R.S. Kumar, Bone in-growth induced by biphasic calcium phosphate ceramic in femoral defect of dogs, Journal of biomaterials applications 19(4) (2005) 341-360. 162 [89] A. Farzadi, M. Solati-Hashjin, F. Bakhshi, A. Aminian, Synthesis and characterization of hydroxyapatite/β-tricalcium phosphate nanocomposites using microwave irradiation, Ceramics International 37(1) (2011) 65-71. [90] B. Li, X. Chen, B. Guo, X. Wang, H. Fan, X. Zhang, Fabrication and cellular biocompatibility of porous carbonated biphasic calcium phosphate ceramics with a nanostructure, Acta biomaterialia 5(1) (2009) 134-143. [91] X. Wang, H. Fan, Y. Xiao, X. Zhang, Fabrication and characterization of porous hydroxyapatite/β-tricalcium phosphate ceramics by microwave sintering, Materials Letters 60(4) (2006) 455-458. [92] C. Combes, C. Rey, Amorphous calcium phosphates: synthesis, properties and uses in biomaterials, Acta Biomaterialia 6(9) (2010) 3362-3378. [93] Y. Li, W. Weng, In vitro synthesis and characterization of amorphous calcium phosphates with various Ca/P atomic ratios, Journal of Materials Science: Materials in Medicine 18(12) (2007) 2303-2308. [94] Y. Li, W. Weng, K. Cheng, P. Du, G. Shen, J. Wang, G. Han, Preparation of amorphous calcium phosphate in the presence of poly (ethylene glycol), Journal of Materials Science Letters 22(14) (2003) 1015-1016. [95] A. Rodrigues, A. Lebugle, Influence of ethanol in the precipitation medium on the composition, structure and reactivity of tricalcium phosphate, Colloids and Surfaces A: Physicochemical and Engineering Aspects 145(1) (1998) 191-204. [96] Q.-L. Tang, K.-W. Wang, Y.-J. Zhu, F. Chen, Single-step rapid microwave-assisted synthesis of polyacrylamide–calcium phosphate nanocomposites in aqueous solution, Materials Letters 63(15) (2009) 1332-1334. 163 [97] H. Zhou, S. Bhaduri, Novel microwave synthesis of amorphous calcium phosphate nanospheres, Journal of Biomedical Materials Research Part B-Applied Biomaterials 100B(4) (2012) 1142-1150. [98] C. Qi, Y.J. Zhu, X.Y. Zhao, B.Q. Lu, Q.L. Tang, J. Zhao, F. Chen, Highly Stable Amorphous Calcium Phosphate Porous Nanospheres: Microwave‐Assisted Rapid Synthesis Using ATP as Phosphorus Source and Stabilizer, and Their Application in Anticancer Drug Delivery, Chemistry-A European Journal 19(3) (2013) 981-987. [99] G.-J. Ding, Y.-J. Zhu, C. Qi, B.-Q. Lu, F. Chen, J. Wu, Porous hollow microspheres of amorphous calcium phosphate: soybean lecithin templated microwave-assisted hydrothermal synthesis and application in drug delivery, Journal of Materials Chemistry B 3(9) (2015) 1823-1830. [100] G.-J. Ding, Y.-J. Zhu, C. Qi, B.-Q. Lu, J. Wu, F. Chen, Porous microspheres of amorphous calcium phosphate: block copolymer templated microwave-assisted hydrothermal synthesis and application in drug delivery, Journal of colloid and interface science 443 (2015) 72-79. [101] J. Zhao, Y.-J. Zhu, G.-F. Cheng, Y.-J. Ruan, T.-W. Sun, F. Chen, J. Wu, X.-Y. Zhao, G.-J. Ding, Microwave-assisted hydrothermal rapid synthesis of amorphous calcium phosphate nanoparticles and hydroxyapatite microspheres using cytidine 5′-triphosphate disodium salt as a phosphate source, Materials Letters 124 (2014) 208-211. [102] C. Qi, Y.-J. Zhu, F. Chen, J. Wu, Porous microspheres of magnesium whitlockite and amorphous calcium magnesium phosphate: microwave-assisted rapid synthesis using creatine phosphate, and application in drug delivery, Journal of Materials Chemistry B 3(39) (2015) 7775-7786. 164 [103] J. Zhao, Y.-J. Zhu, J.-Q. Zheng, F. Chen, J. Wu, Microwave-assisted hydrothermal preparation using adenosine 5′-triphosphate disodium salt as a phosphate source and characterization of zinc-doped amorphous calcium phosphate mesoporous microspheres, Microporous and Mesoporous Materials 180 (2013) 79-85. [104] F. Chen, P. Huang, C. Qi, B.-Q. Lu, X.-Y. Zhao, C. Li, J. Wu, D.-X. Cui, Y.-J. Zhu, Multifunctional biodegradable mesoporous microspheres of Eu 3+-doped amorphous calcium phosphate: microwave-assisted preparation, pH-sensitive drug release, and bioimaging application, Journal of Materials Chemistry B 2(41) (2014) 7132-7140. [105] F. Chen, Y.-J. Zhu, K.-H. Zhang, J. Wu, K.-W. Wang, Q.-L. Tang, X.-M. Mo, Europium-doped amorphous calcium phosphate porous nanospheres: preparation and application as luminescent drug carriers, Nanoscale research letters 6(1) (2011) 67. [106] J. Wei, J. Jia, F. Wu, S. Wei, H. Zhou, H. Zhang, J.-W. Shin, C. Liu, Hierarchically microporous/macroporous scaffold of magnesium–calcium phosphate for bone tissue regeneration, Biomaterials 31(6) (2010) 1260-1269. [107] J. Lu, J. Wei, Y. Yan, H. Li, J. Jia, S. Wei, H. Guo, T. Xiao, C. Liu, Preparation and preliminary cytocompatibility of magnesium doped apatite cement with degradability for bone regeneration, Journal of Materials Science: Materials in Medicine 22(3) (2011) 607- 615. [108] F. Tamimi, D.L. Nihouannen, D.C. Bassett, S. Ibasco, U. Gbureck, J. Knowles, A. Wright, A. Flynn, S.V. Komarova, J.E. Barralet, Biocompatibility of magnesium phosphate minerals and their stability under physiological conditions, Acta Biomaterialia 7(6) (2011) 2678-2685. 165 [109] H. Zhou, T.J.F. Luchini, S.B. Bhaduri, Microwave assisted synthesis of amorphous magnesium phosphate nanospheres, Journal of Materials Science: Materials in Medicine 23(12) (2012) 2831-2837. [110] C. Qi, Y.-J. Zhu, B.-Q. Lu, J. Wu, F. Chen, Amorphous magnesium phosphate flower-like hierarchical nanostructures: microwave-assisted rapid synthesis using fructose 1, 6-bisphosphate trisodium salt as an organic phosphorus source and application in protein adsorption, RSC Advances 5(19) (2015) 14906-14915. [111] C. Qi, Y.-J. Zhu, B.-Q. Lu, G.-J. Ding, T.-W. Sun, F. Chen, J. Wu, Microwave- assisted rapid synthesis of magnesium phosphate hydrate nanosheets and their application in drug delivery and protein adsorption, Journal of Materials Chemistry B 2(48) (2014) 8576-8586. [112] C. Qi, Y.-J. Zhu, C.-T. Wu, T.-W. Sun, F. Chen, J. Wu, Magnesium phosphate pentahydrate nanosheets: Microwave-hydrothermal rapid synthesis using creatine phosphate as an organic phosphorus source and application in protein adsorption, Journal of colloid and interface science 462 (2016) 297-306. [113] C. Qi, Y.-J. Zhu, G.-F. Cheng, Y.-J. Ruan, G.-J. Ding, T.-W. Sun, F. Chen, J. Wu, Microwave-assisted rapid synthesis of magnesium phosphate hierarchical structures using adenosine 5′-triphosphate disodium salt as a phosphorus source, Materials Letters 140 (2015) 79-82. [114] T. Albrektsson, P.-I. Brånemark, H.-A. Hansson, J. Lindström, Osseointegrated titanium implants: requirements for ensuring a long-lasting, direct bone-to-implant anchorage in man, Acta Orthopaedica Scandinavica 52(2) (1981) 155-170. 166 [115] H. Zhou, M. Nabiyouni, S.B. Bhaduri, Microwave assisted apatite coating deposition on Ti6Al4V implants, Materials Science & Engineering C-Materials for Biological Applications 33(7) (2013) 4435-4443. [116] H. Zhou, S. Kong, Y. Pan, Z. Zhang, L. Deng, Microwave-assisted fabrication of strontium doped apatite coating on Ti6Al4V, Materials Science and Engineering: C 56 (2015) 174-180. [117] F. Zhang, S. Cai, G. Xu, F. Wang, N. Yu, R. Ling, X. Wu, Preparation and bioactivity of apatite coating on Ti6Al4V alloy by microwave assisted aqueous chemical method, Ceramics International 42(16) (2016) 18466-18473. [118] M.-T. Choy, K.-W. Yeung, L. Chen, C.-Y. Tang, G.C.-P. Tsui, W.-C. Law, In situ synthesis of osteoconductive biphasic ceramic coatings on Ti6Al4V substrate by laser- microwave hybridization, Surface and Coatings Technology 330 (2017) 92-101. [119] Y.F. Ren, H. Zhou, M. Nabiyouni, S.B. Bhaduri, Rapid coating of AZ31 magnesium alloy with calcium deficient hydroxyapatite using microwave energy, Materials Science & Engineering C-Materials for Biological Applications 49 (2015) 364-372. [120] S. Shen, S. Cai, M. Zhang, G. Xu, Y. Li, R. Ling, X. Wu, Microwave assisted deposition of hydroxyapatite coating on a magnesium alloy with enhanced corrosion resistance, Materials Letters 159 (2015) 146-149. [121] Y. Ren, E. Babaie, B. Lin, S.B. Bhaduri, Microwave-assisted magnesium phosphate coating on the AZ31 magnesium alloy, Biomedical Materials 12(4) (2017) 045026. [122] G.K. Bayrak, T.T. Demirtaş, M. Gümüşderelioğlu, Microwave-induced biomimetic approach for hydroxyapatite coatings of chitosan scaffolds, Carbohydrate polymers 157 (2017) 803-813. 167 [123] Y. Su, K. Li, L. Zhang, S. Farhan, S. Liu, G. Shi, Ca-P bioactive coating prepared by combining microwave-hydrothermal and supersonic atmospheric plasma spraying methods, Materials Science and Engineering: C 72 (2017) 371-377. [124] H. Zhou, V.K. Goel, S.B. Bhaduri, A fast route to modify biopolymer surface: A study on polyetheretherketone (PEEK), Materials Letters 125 (2014) 96-98. [125] M.P. Staiger, A.M. Pietak, J. Huadmai, G. Dias, Magnesium and its alloys as orthopedic biomaterials: A review, Biomaterials 27(9) (2006) 1728-1734. [126] F. Witte, N. Hort, C. Vogt, S. Cohen, K.U. Kainer, R. Willumeit, F. Feyerabend, Degradable biomaterials based on magnesium corrosion, Current Opinion in Solid State & Materials Science 12(5-6) (2008) 63-72. [127] F. Witte, The history of biodegradable magnesium implants: A review, Acta Biomaterialia 6(5) (2010) 1680-1692. [128] Y.F. Zheng, X.N. Gu, Research Activities of Biomedical Magnesium Alloys in China, Jom 63(4) (2011) 105-108. [129] S. Shadanbaz, G.J. Dias, Calcium phosphate coatings on magnesium alloys for biomedical applications: A review, Acta Biomaterialia 8(1) (2012) 20-30. [130] F. Witte, V. Kaese, H. Haferkamp, E. Switzer, A. Meyer-Lindenberg, C.J. Wirth, H. Windhagen, In vivo corrosion of four magnesium alloys and the associated bone response, Biomaterials 26(17) (2005) 3557-3563. [131] G.L. Song, Control of biodegradation of biocompatable magnesium alloys, Corrosion Science 49(4) (2007) 1696-1701. [132] H. Hornberger, S. Virtanen, A.R. Boccaccini, Biomedical coatings on magnesium alloys - A review, Acta Biomaterialia 8(7) (2012) 2442-2455. 168 [133] T. Kraus, S.F. Fischerauer, A.C. Hanzi, P.J. Uggowitzer, J.F. Loffler, A.M. Weinberg, Magnesium alloys for temporary implants in osteosynthesis: In vivo studies of their degradation and interaction with bone, Acta Biomaterialia 8(3) (2012) 1230-1238. [134] R.C. Zeng, W. Dietzel, F. Witte, N. Hort, C. Blawert, Progress and challenge for magnesium alloys as biomaterials, Advanced Engineering Materials 10(8) (2008) B3-B14. [135] J.E. Gray, B. Luan, Protective coatings on magnesium and its alloys — a critical review, Journal of Alloys and Compounds 336(1–2) (2002) 88-113. [136] X.B. Chen, N. Birbilis, T.B. Abbott, Review of Corrosion-Resistant Conversion Coatings for Magnesium and Its Alloys, Corrosion 67(3) (2011). [137] H. Huo, Y. Li, F. Wang, Corrosion of AZ91D magnesium alloy with a chemical conversion coating and electroless nickel layer, Corrosion Science 46(6) (2004) 1467- 1477. [138] C.D. Gu, J.S. Lian, J.G. He, Z.H. Jiang, Q. Jiang, High corrosion-resistance nanocrystalline Ni coating on AZ91D magnesium alloy, Surface & Coatings Technology 200(18-19) (2006) 5413-5418. [139] L.-H. Chiu, C.-C. Chen, C.-F. Yang, Improvement of corrosion properties in an aluminum-sprayed AZ31 magnesium alloy by a post-hot pressing and anodizing treatment, Surface and Coatings Technology 191(2-3) (2005) 181-187. [140] M. Campo, M. Carboneras, M.D. López, B. Torres, P. Rodrigo, E. Otero, J. Rams, Corrosion resistance of thermally sprayed Al and Al/SiC coatings on Mg, Surface and Coatings Technology 203(20-21) (2009) 3224-3230. 169 [141] L. Xu, A. Yamamoto, Characteristics and cytocompatibility of biodegradable polymer film on magnesium by spin coating, Colloids Surf B Biointerfaces 93 (2012) 67- 74. [142] S.A. Khana, Y. Miyashita, Y. Mutoh, T. Koike, Effect of anodized layer thickness on fatigue behavior of magnesium alloy, Materials Science and Engineering a-Structural Materials Properties Microstructure and Processing 474(1-2) (2008) 261-269. [143] G.Y. Liu, J. Hu, Z.K. Ding, C. Wang, Bioactive calcium phosphate coating formed on micro-arc oxidized magnesium by chemical deposition, Applied Surface Science 257(6) (2011) 2051-2057. [144] S.V. Dorozhkin, Calcium orthophosphate coatings on magnesium and its biodegradable alloys, Acta Biomaterialia 10(7) (2014) 2919-2934. [145] J.L. Ong, D.C.N. Chan, Hydroxyapatite and their use as coatings in dental implants: A review, Critical Reviews in Biomedical Engineering 28(5-6) (2000) 667A-707A. [146] H. Zhou, J.G. Lawrence, A.H. Touny, S.B. Bhaduri, Biomimetic coating of bisphosphonate incorporated CDHA on Ti6Al4V, Journal of Materials Science-Materials in Medicine 23(2) (2012) 365-374. [147] S. Jalota, S.B. Bhaduri, A.C. Tas, In vitro testing of calcium phosphate (HA, TCP, and biphasic HA-TCP) whiskers, Journal of Biomedical Materials Research Part A 78A(3) (2006) 481-490. [148] S.V. Dorozhkin, Calcium Orthophosphate-Based Bioceramics, Materials 6(9) (2013) 3840-3942. [149] S. Weiner, H.D. Wagner, THE MATERIAL BONE: Structure-Mechanical Function Relations, Annual Review of Materials Science 28(1) (1998) 271-298. 170 [150] H. Zhou, S. Bhaduri, Novel microwave synthesis of amorphous calcium phosphate nanospheres, J Biomed Mater Res B Appl Biomater 100(4) (2012) 1142-50. [151] J. Waterman, A. Pietak, N. Birbilis, T. Woodfield, G. Dias, M.P. Staiger, Corrosion resistance of biomimetic calcium phosphate coatings on magnesium due to varying pretreatment time, Materials Science and Engineering B-Advanced Functional Solid-State Materials 176(20) (2011) 1756-1760. [152] W. Cui, E. Beniash, E. Gawalt, Z. Xu, C. Sfeir, Biomimetic coating of magnesium alloy for enhanced corrosion resistance and calcium phosphate deposition, Acta Biomaterialia 9(10) (2013) 8650-8659. [153] J.X. Yang, F.Z. Cui, Q.S. Yin, Y. Zhang, T. Zhang, X.M. Wang, Characterization and Degradation Study of Calcium Phosphate Coating on Magnesium Alloy Bone Implant In Vitro, Ieee Transactions on Plasma Science 37(7) (2009) 1161-1168. [154] A. Yanovska, V. Kuznetsov, A. Stanislavov, S. Danilchenko, L. Sukhodub, Calcium–phosphate coatings obtained biomimetically on magnesium substrates under low magnetic field, Applied Surface Science 258(22) (2012) 8577-8584. [155] D.A. Cortes, H.Y. Lopez, D. Mantovani, Spontaneous and biomimetic apatite formation on pure magnesium, Thermec 2006, Pts 1-5 539-543 (2007) 589-594. [156] F.-z. Cui, J.-x. Yang, Y.-p. Jiao, Q.-s. Yin, Y. Zhang, I.-S. Lee, Calcium phosphate coating on magnesium alloy for modification of degradation behavior, Frontiers of Materials Science in China 2(2) (2008) 143-148. [157] H.X. Wang, S.K. Guan, X. Wang, C.X. Ren, L.G. Wang, In vitro degradation and mechanical integrity of Mg-Zn-Ca alloy coated with Ca-deficient hydroxyapatite by the pulse electrodeposition process, Acta Biomaterialia 6(5) (2010) 1743-1748. 171 [158] R.G. Guan, I. Johnson, T. Cui, T. Zhao, Z.Y. Zhao, X. Li, H. Liu, Electrodeposition of hydroxyapatite coating on Mg-4.0Zn-1.0Ca-0.6Zr alloy and in vitro evaluation of degradation, hemolysis, and cytotoxicity, J Biomed Mater Res A 100(4) (2012) 999-1015. [159] Y.W. Song, D.Y. Shan, E.H. Han, Electrodeposition of hydroxyapatite coating on AZ91D magnesium alloy for biomaterial application, Materials Letters 62(17-18) (2008) 3276-3279. [160] R. Rojaee, M. Fathi, K. Raeissi, Controlling the degradation rate of AZ91 magnesium alloy via sol-gel derived nanostructured hydroxyapatite coating, Materials Science & Engineering C-Materials for Biological Applications 33(7) (2013) 3817-3825. [161] M. Tomozawa, S. Hiromoto, Growth mechanism of hydroxyapatite-coatings formed on pure magnesium and corrosion behavior of the coated magnesium, Applied Surface Science 257(19) (2011) 8253-8257. [162] M. Tomozawa, S. Hiromoto, Y. Harada, Microstructure of hydroxyapatite-coated magnesium prepared in aqueous solution, Surface & Coatings Technology 204(20) (2010) 3243-3247. [163] A.C. Tas, S.B. Bhaduri, Rapid coating of Ti6Al4V at room temperature with a calcium phosphate solution similar to 10× simulated body fluid, Journal of Materials Research 19(09) (2004) 2742-2749. [164] S. Jalota, S.B. Bhaduri, A.C. Tas, Effect of carbonate content and buffer type on calcium phosphate formation in SBF solutions, Journal of Materials Science-Materials in Medicine 17(8) (2006) 697-707. 172 [165] S. Jalota, S.B. Bhaduri, A.C. Tas, Osteoblast proliferation on neat and apatite-like calcium phosphate-coated titanium foam scaffolds, Materials Science & Engineering C- Biomimetic and Supramolecular Systems 27(3) (2007) 432-440. [166] S. Jalota, S. Bhaduri, S.B. Bhaduri, A.C. Tas, A protocol to develop crack-free biomimetic coatings on Ti6Al4V substrates, Journal of Materials Research 22(6) (2007) 1593-1600. [167] S. Jalota, S.B. Bhaduri, A.C. Tas, Using a synthetic body-fluid (SBF) solution of 27 mM HCO3- to make bone substitutes more osteointegrative, Materials Science & Engineering C-Biomimetic and Supramolecular Systems 28(1) (2008) 129-140. [168] S.J. Gadaleta, E.P. Paschalis, F. Betts, R. Mendelsohn, A.L. Boskey, Fourier transform infrared spectroscopy of the solution-mediated conversion of amorphous calcium phosphate to hydroxyapatite: New correlations between X-ray diffraction and infrared data, Calcified Tissue International 58(1) (1996) 9-16. [169] F.Z. Ren, Y.H. Ding, Y. Leng, Infrared spectroscopic characterization of carbonated apatite: A combined experimental and computational study, Journal of Biomedical Materials Research Part A 102(2) (2014) 496-505. [170] H.X. Wang, S.J. Zhu, L.G. Wang, Y.S. Feng, X. Ma, S.K. Guan, Formation mechanism of Ca-deficient hydroxyapatite coating on Mg-Zn-Ca alloy for orthopaedic implant, Applied Surface Science 307 (2014) 92-100. [171] X. Lu, Y. Leng, Theoretical analysis of calcium phosphate precipitation in simulated body fluid, Biomaterials 26(10) (2005) 1097-1108. 173 [172] J.E. Gray-Munro, M. Strong, The mechanism of deposition of calcium phosphate coatings from solution onto magnesium alloy AZ31, J Biomed Mater Res A 90(2) (2009) 339-50. [173] D. Laurencin, N. Almora-Barrios, N.H. de Leeuw, C. Gervais, C. Bonhomme, F. Mauri, W. Chrzanowski, J.C. Knowles, R.J. Newport, A. Wong, Z.H. Gan, M.E. Smith, Magnesium incorporation into hydroxyapatite, Biomaterials 32(7) (2011) 1826-1837. [174] A. Yamamoto, S. Hiromoto, Effect of inorganic salts, amino acids and proteins on the degradation of pure magnesium in vitro, Materials Science & Engineering C- Biomimetic and Supramolecular Systems 29(5) (2009) 1559-1568. [175] Y.W. Song, D.Y. Shan, R.S. Chen, F. Zhang, E.H. Han, Biodegradable behaviors of AZ31 magnesium alloy in simulated body fluid, Materials Science & Engineering C- Biomimetic and Supramolecular Systems 29(3) (2009) 1039-1045. [176] S. Keim, J.G. Brunner, B. Fabry, S. Virtanen, Control of magnesium corrosion and biocompatibility with biomimetic coatings, Journal of Biomedical Materials Research Part B-Applied Biomaterials 96B(1) (2011) 84-90. [177] A. Abdal-hay, T. Amna, J.K. Lim, Biocorrosion and osteoconductivity of PCL/nHAp composite porous film-based coating of magnesium alloy, Solid State Sciences 18 (2013) 131-140. [178] R. Zeng, W. Dietzel, F. Witte, N. Hort, C. Blawert, Progress and Challenge for Magnesium Alloys as Biomaterials, Advanced Engineering Materials 10(8) (2008) B3- B14. 174 [179] N.T. Kirkland, N. Birbilis, M.P. Staiger, Assessing the corrosion of biodegradable magnesium implants: A critical review of current methodologies and their limitations, Acta Biomaterialia 8(3) (2012) 925-936. [180] F. Witte, J. Fischer, J. Nellesen, C. Vogt, J. Vogt, T. Donath, F. Beckmann, In vivo corrosion and corrosion protection of magnesium alloy LAE442, Acta Biomaterialia 6(5) (2010) 1792-1799. [181] H. Zhou, A.K. Agarwal, V.K. Goel, S.B. Bhaduri, Microwave assisted preparation of magnesium phosphate cement (MPC) for orthopedic applications: A novel solution to the exothermicity problem, Materials Science & Engineering C-Materials for Biological Applications 33(7) (2013) 4288-4294. [182] K.H. Bichler, E. Eipper, K. Naber, V. Braun, R. Zimmermann, S. Lahme, Urinary infection stones, International Journal of Antimicrobial Agents 19(6) (2002) 488-498. [183] E. Babaie, B. Lin, V.K. Goel, S.B. Bhaduri, Evaluation of amorphous magnesium phosphate (AMP) based non-exothermic orthopedic cements, Biomedical Materials 11(5) (2016) 055010. [184] E. Babaie, Y. Ren, S.B. Bhaduri, Microwave sintering of fine grained MgP and Mg substitutes with amorphous tricalcium phosphate: Structural, and mechanical characterization, Journal of Materials Research 31(08) (2016) 995-1003. [185] M. Nabiyouni, Y. Ren, S.B. Bhaduri, Magnesium substitution in the structure of orthopedic nanoparticles: A comparison between amorphous magnesium phosphates, calcium magnesium phosphates, and hydroxyapatites, Materials Science and Engineering: C 52 (2015) 11-17. 175 [186] E. Babaie, H. Zhou, B. Lin, S.B. Bhaduri, Influence of ethanol content in the precipitation medium on the composition, structure and reactivity of magnesium–calcium phosphate, Materials Science and Engineering: C 53 (2015) 204-211. [187] C. Qi, Y.-J. Zhu, B.-Q. Lu, J. Wu, F. Chen, Amorphous magnesium phosphate flower-like hierarchical nanostructures: microwave-assisted rapid synthesis using fructose 1,6-bisphosphate trisodium salt as an organic phosphorus source and application in protein adsorption, RSC Adv. 5(19) (2015) 14906-14915. [188] T. Ishizaki, I. Shigematsu, N. Saito, Anticorrosive magnesium phosphate coating on AZ31 magnesium alloy, Surface and Coatings Technology 203(16) (2009) 2288-2291. [189] M.I. Rahim, A. Tavares, F. Evertz, M. Kieke, J.-M. Seitz, R. Eifler, A. Weizbauer, E. Willbold, H. Jürgen Maier, B. Glasmacher, P. Behrens, H. Hauser, P.P. Mueller, Phosphate conversion coating reduces the degradation rate and suppresses side effects of metallic magnesium implants in an animal model, Journal of Biomedical Materials Research Part B: Applied Biomaterials (2016). [190] H. Zhao, S. Cai, Z. Ding, M. Zhang, Y. Li, G. Xu, A simple method for the preparation of magnesium phosphate conversion coatings on a AZ31 magnesium alloy with improved corrosion resistance, RSC Adv. 5(31) (2015) 24586-24590. [191] M. Fouladi, A. Amadeh, Effect of phosphating time and temperature on microstructure and corrosion behavior of magnesium phosphate coating, Electrochimica Acta 106 (2013) 1-12. [192] M.F. Morks, Magnesium phosphate treatment for steel, Materials Letters 58(26) (2004) 3316-3319. 176 [193] M. Fouladi, A. Amadeh, Comparative study between novel magnesium phosphate and traditional zinc phosphate coatings, Materials Letters 98 (2013) 1-4. [194] R. Boistelle, F. Abbona, Morphology, habit and growth pf newberyite crystals (MgHPO4·3 H2O), Journal of Crystal Growth 54(2) (1981) 275-295. [195] N. Ostrowski, B. Lee, D. Hong, P.N. Enick, A. Roy, P.N. Kumta, Synthesis, Osteoblast, and Osteoclast Viability of Amorphous and Crystalline Tri-Magnesium Phosphate, ACS Biomaterials Science & Engineering 1(1) (2015) 52-63. [196] R.Z. LeGeros, S. Lin, R. Rohanizadeh, D. Mijares, J.P. LeGeros, Journal of Materials Science: Materials in Medicine 14(3) (2003) 201-209. [197] A.W. Taylor, A.W. Frazier, E.L. Gurney, Solubility products of magnesium ammonium and magnesium potassium phosphates, Transactions of the Faraday Society 59 (1963) 1580. [198] P.W. Brown, J. Gulick, J. Dumm, The System MgOP2O5H2O at 25oC, Journal of the American Ceramic Society 76(6) (1993) 1558-1562. [199] F. Abbona, H.E. Lundager Madsen, R. Boistelle, Crystallization of two magnesium phosphates, struvite and newberyite: Effect of pH and concentration, Journal of Crystal Growth 57(1) (1982) 6-14. [200] K.O. Kongshaug, H. Fjellvåg, K.P. Lillerud, The synthesis and crystal structure of a hydrated magnesium phosphate Mg3(PO4)2·4H2O, Solid State Sciences 3(3) (2001) 353- 359. [201] J.-x. Yang, Y.-p. Jiao, Q.-s. Yin, Y. Zhang, T. Zhang, Calcium phosphate coating on magnesium alloy by biomimetic method: Investigation of morphology, composition and formation process, Frontiers of Materials Science in China 2(2) (2008) 149-155. 177 [202] U. Klammert, A. Ignatius, U. Wolfram, T. Reuther, U. Gbureck, In vivo degradation of low temperature calcium and magnesium phosphate ceramics in a heterotopic model, Acta Biomaterialia 7(9) (2011) 3469-3475. [203] F. Seuss, S. Seuss, M.C. Turhan, B. Fabry, S. Virtanen, Corrosion of Mg alloy AZ91D in the presence of living cells, Journal of Biomedical Materials Research Part B: Applied Biomaterials 99B(2) (2011) 276-281. [204] M. Niinomi, M. Nakai, J. Hieda, Development of new metallic alloys for biomedical applications, Acta Biomaterialia 8(11) (2012) 3888-3903. [205] M. Niinomi, Mechanical biocompatibilities of titanium alloys for biomedical applications, Journal of the mechanical behavior of biomedical materials 1(1) (2008) 30- 42. [206] J. Nagels, M. Stokdijk, P.M. Rozing, Stress shielding and bone resorption in shoulder arthroplasty, Journal of shoulder and elbow surgery 12(1) (2003) 35-39. [207] W. Dhert, C. Klein, J. Jansen, E. Van der Velde, R. Vriesde, P. Rozing, K. De Groot, A histological and histomorphometrical investigation of fluorapatite, magnesiumwhitlockite, and hydroxylapatite plasma‐sprayed coatings in goats, Journal of Biomedical Materials Research Part A 27(1) (1993) 127-138. [208] J. Li, Y. Song, S. Zhang, C. Zhao, F. Zhang, X. Zhang, L. Cao, Q. Fan, T. Tang, In vitro responses of human bone marrow stromal cells to a fluoridated hydroxyapatite coated biodegradable Mg–Zn alloy, Biomaterials 31(22) (2010) 5782-5788. [209] J. Wang, Y. Chao, Q. Wan, Z. Zhu, H. Yu, Fluoridated hydroxyapatite coatings on titanium obtained by electrochemical deposition, Acta Biomaterialia 5(5) (2009) 1798- 1807. 178 [210] N. Guha‐Chowdhury, A. Clark, C. Sissons, Inhibition of purified enolases from oral bacteria by fluoride, Molecular Oral Microbiology 12(2) (1997) 91-97. [211] M. Zheludkevich, R. Serra, M. Montemor, I.M. Salvado, M. Ferreira, Corrosion protective properties of nanostructured sol–gel hybrid coatings to AA2024-T3, Surface and Coatings Technology 200(9) (2006) 3084-3094. [212] X. Gu, W. Zheng, Y. Cheng, Y. Zheng, A study on alkaline heat treated Mg–Ca alloy for the control of the biocorrosion rate, Acta Biomaterialia 5(7) (2009) 2790-2799. [213] X. Li, Z. Weng, W. Yuan, X. Luo, H.M. Wong, X. Liu, S. Wu, K. Yeung, Y. Zheng, P.K. Chu, Corrosion resistance of dicalcium phosphate dihydrate/poly (lactic-co-glycolic acid) hybrid coating on AZ31 magnesium alloy, Corrosion Science 102 (2016) 209-221. [214] H. Bakhsheshi-Rad, E. Hamzah, M. Abdul-Kadir, S.N. Saud, M. Kasiri-Asgarani, R. Ebrahimi-Kahrizsangi, The Mechanical Properties and Corrosion Behavior of Double- Layered Nano Hydroxyapatite-Polymer Coating on Mg-Ca Alloy, Journal of Materials Engineering and Performance 24(10) (2015) 4010-4021. [215] R.-C. Zeng, W.-C. Qi, Y.-W. Song, Q.-K. He, H.-Z. Cui, E.-H. Han, In vitro degradation of MAO/PLA coating on Mg-1.21 Li-1.12 Ca-1.0 Y alloy, Frontiers of Materials Science 8(4) (2014) 343-353. [216] E. Meng, S. Guan, H. Wang, L. Wang, S. Zhu, J. Hu, C. Ren, J. Gao, Y. Feng, Effect of electrodeposition modes on surface characteristics and corrosion properties of fluorine- doped hydroxyapatite coatings on Mg–Zn–Ca alloy, Applied Surface Science 257(11) (2011) 4811-4816. [217] H. Bakhsheshi-Rad, E. Hamzah, M. Daroonparvar, R. Ebrahimi-Kahrizsangi, M. Medraj, In-vitro corrosion inhibition mechanism of fluorine-doped hydroxyapatite and 179 brushite coated Mg–Ca alloys for biomedical applications, Ceramics International 40(6) (2014) 7971-7982. [218] M. Fathi, E.M. Zahrani, Mechanical alloying synthesis and bioactivity evaluation of nanocrystalline fluoridated hydroxyapatite, Journal of Crystal Growth 311(5) (2009) 1392- 1403. [219] T. Ishizaki, M. Okido, Y. Masuda, N. Saito, M. Sakamoto, Corrosion Resistant Performances of Alkanoic and Phosphonic Acids Derived Self-Assembled Monolayers on Magnesium Alloy AZ31 by Vapor-Phase Method, Langmuir 27(10) (2011) 6009-6017. [220] N.J. Ostrowski, B. Lee, A. Roy, M. Ramanathan, P.N. Kumta, Biodegradable poly(lactide-co-glycolide) coatings on magnesium alloys for orthopedic applications, Journal of Materials Science: Materials in Medicine 24(1) (2012) 85-96. [221] W. Zhang, Y. Chen, M. Chen, S. Zhao, J. Mao, A. Qu, W. Li, Y. Zhao, N. Huang, G. Wan, Strengthened corrosion control of poly (lactic acid) (PLA) and poly (ε- caprolactone) (PCL) polymer-coated magnesium by imbedded hydrophobic stearic acid (SA) thin layer, Corrosion Science 112 (2016) 327-337. [222] B. O’Brien, W. Carroll, The evolution of cardiovascular stent materials and surfaces in response to clinical drivers: A review, Acta Biomaterialia 5(4) (2009) 945-958. [223] L.S. Nair, C.T. Laurencin, Biodegradable polymers as biomaterials, Progress in Polymer Science 32(8-9) (2007) 762-798. [224] A. Alabbasi, S. Liyanaarachchi, M.B. Kannan, Polylactic acid coating on a biodegradable magnesium alloy: An in vitro degradation study by electrochemical impedance spectroscopy, Thin Solid Films 520(23) (2012) 6841-6844. 180 [225] A. Abdal-hay, N.A.M. Barakat, J.K. Lim, Hydroxyapatite-doped poly(lactic acid) porous film coating for enhanced bioactivity and corrosion behavior of AZ31 Mg alloy for orthopedic applications, Ceramics International 39(1) (2013) 183-195. [226] Y. Chen, Y. Song, S. Zhang, J. Li, C. Zhao, X. Zhang, Interaction between a high purity magnesium surface and PCL and PLA coatings during dynamic degradation, Biomedical Materials 6(2) (2011) 025005. [227] H. Zhou, J.G. Lawrence, S.B. Bhaduri, Fabrication aspects of PLA-CaP/PLGA-CaP composites for orthopedic applications: A review, Acta Biomaterialia 8(6) (2012) 1999- 2016. [228] H. Zhou, M. Nabiyouni, B. Lin, S.B. Bhaduri, Fabrication of novel poly(lactic acid)/amorphous magnesium phosphate bionanocomposite fibers for tissue engineering applications via electrospinning, Materials Science and Engineering: C 33(4) (2013) 2302- 2310. [229] N. Ostrowski, V. Sharma, A. Roy, P.N. Kumta, Systematic Assessment of Synthesized Tri-magnesium Phosphate Powders (Amorphous, Semi-crystalline and Crystalline) and Cements for Ceramic Bone Cement Applications, Journal of Materials Science & Technology 31(5) (2015) 437-444. [230] C.P. Weisel, W.K. Jo, Ingestion, inhalation, and dermal exposures to chloroform and trichloroethene from tap water, Environ Health Perspect 104(1) (1996) 48-51. [231] I. Johnson, S.M. Wang, C. Silken, H. Liu, A systemic study on key parameters affecting nanocomposite coatings on magnesium substrates, Acta Biomaterialia 36 (2016) 332-349. 181 [232] I. Johnson, K. Akari, H. Liu, Nanostructured hydroxyapatite/poly(lactic-co-glycolic acid) composite coating for controlling magnesium degradation in simulated body fluid, Nanotechnology 24(37) (2013) 375103. [233] A. Abdal-hay, A. Hasan, K. Yu, M.-H. Lee, A.S. Hamdy, K.A. Khalil, Biocorrosion behavior of biodegradable nanocomposite fibers coated layer-by-layer on AM50 magnesium implant, Materials Science and Engineering: C 58 (2016) 1232-1241. [234] F. Zucchi, V. Grassi, A. Frignani, C. Monticelli, G. Trabanelli, Electrochemical behaviour of a magnesium alloy containing rare earth elements, Journal of Applied Electrochemistry 36(2) (2005) 195-204. [235] W. Araujo, Undoped polyaniline anticorrosive properties, Electrochimica Acta 46(9) (2001) 1307-1312. [236] T. Ishizaki, J. Hieda, N. Saito, N. Saito, O. Takai, Corrosion resistance and chemical stability of super-hydrophobic film deposited on magnesium alloy AZ31 by microwave plasma-enhanced chemical vapor deposition, Electrochimica Acta 55(23) (2010) 7094- 7101. [237] B. Li, Y. Han, K. Qi, Formation Mechanism, Degradation Behavior, and Cytocompatibility of a Nanorod-Shaped HA and Pore-Sealed MgO Bilayer Coating on Magnesium, ACS Applied Materials & Interfaces 6(20) (2014) 18258-18274. [238] J. Jayaraj, S. Amruth Raj, A. Srinivasan, S. Ananthakumar, U.T.S. Pillai, N.G.K. Dhaipule, U.K. Mudali, Composite magnesium phosphate coatings for improved corrosion resistance of magnesium AZ31 alloy, Corrosion Science 113 (2016) 104-115. 182 [239] A. Abdal-hay, N.A.M. Barakat, J.K. Lim, Influence of electrospinning and dip- coating techniques on the degradation and cytocompatibility of Mg-based alloy, Colloids and Surfaces A: Physicochemical and Engineering Aspects 420 (2013) 37-45. [240] R. Zhang, P.X. Ma, Biomimetic Polymer/Apatite Composite Scaffolds for Mineralized Tissue Engineering, Macromolecular Bioscience 4(2) (2004) 100-111. [241] N.L. Morozowich, J.L. Nichol, H.R. Allcock, Investigation of Apatite Mineralization on Antioxidant Polyphosphazenes for Bone Tissue Engineering, Chemistry of Materials 24(17) (2012) 3500-3509. [242] S.M. Kurtz, J.N. Devine, PEEK biomaterials in trauma, orthopedic, and spinal implants, Biomaterials 28(32) (2007) 4845-4869. [243] J.M. Toth, M. Wang, B.T. Estes, J.L. Scifert, H.B. Seim, A.S. Turner, Polyetheretherketone as a biomaterial for spinal applications, Biomaterials 27(3) (2006) 324-334. [244] D.P. Jones, D.C. Leach, D.R. Moore, Mechanical properties of poly(ether-ether- ketone) for engineering applications, Polymer 26(9) (1985) 1385-1393. [245] R.K. Ponnappan, H. Serhan, B. Zarda, R. Patel, T. Albert, A.R. Vaccaro, Biomechanical evaluation and comparison of polyetheretherketone rod system to traditional titanium rod fixation, The Spine Journal 9(3) (2009) 263-267. [246] S.W. Ha, J. Mayer, B. Koch, E. Wintermantel, Plasma-sprayed hydroxylapatite coating on carbon fibre reinforced thermoplastic composite materials, Journal of Materials Science: Materials in Medicine 5(6-7) (1994) 481-484. 183 [247] K.H. Tan, C.K. Chua, K.F. Leong, C.M. Cheah, P. Cheang, M.S. Abu Bakar, S.W. Cha, Scaffold development using selective laser sintering of polyetheretherketone– hydroxyapatite biocomposite blends, Biomaterials 24(18) (2003) 3115-3123. [248] G.M. Wu, W.D. Hsiao, S.F. Kung, Investigation of hydroxyapatite coated polyether ether ketone composites by gas plasma sprays, Surface and Coatings Technology 203(17- 18) (2009) 2755-2758. [249] J.H. Lee, H.L. Jang, K.M. Lee, H.-R. Baek, K. Jin, K.S. Hong, J.H. Noh, H.-K. Lee, In vitro and in vivo evaluation of the bioactivity of hydroxyapatite-coated polyetheretherketone biocomposites created by cold spray technology, Acta Biomaterialia 9(4) (2013) 6177-6187. [250] B.-D. Hahn, D.-S. Park, J.-J. Choi, J. Ryu, W.-H. Yoon, J.-H. Choi, J.-W. Kim, C.- W. Ahn, H.-E. Kim, B.-H. Yoon, I.-K. Jung, Osteoconductive hydroxyapatite coated PEEK for spinal fusion surgery, Applied Surface Science 283 (2013) 6-11. [251] M.R. Abdullah, A. Goharian, M.R. Abdul Kadir, M.U. Wahit, Biomechanical and bioactivity concepts of polyetheretherketone composites for use in orthopedic implants-a review, Journal of Biomedical Materials Research Part A 103(11) (2015) 3689-3702. [252] S. Barkarmo, A. Wennerberg, M. Hoffman, P. Kjellin, K. Breding, P. Handa, V. Stenport, Nano-hydroxyapatite-coated PEEK implants: A pilot study in rabbit bone, Journal of Biomedical Materials Research Part A 101A(2) (2013) 465-471. [253] G.L. Converse, W. Yue, R.K. Roeder, Processing and tensile properties of hydroxyapatite-whisker-reinforced polyetheretherketone, Biomaterials 28(6) (2007) 927- 935. 184 [254] G.L. Converse, T.L. Conrad, C.H. Merrill, R.K. Roeder, Hydroxyapatite whisker- reinforced polyetherketoneketone bone ingrowth scaffolds, Acta Biomaterialia 6(3) (2010) 856-863. [255] D. Almasi, S. Izman, M. Assadian, M. Ghanbari, M.R. Abdul Kadir, Crystalline ha coating on peek via chemical deposition, Applied Surface Science 314 (2014) 1034-1040. [256] A. Rabiei, S. Sandukas, Processing and evaluation of bioactive coatings on polymeric implants, Journal of Biomedical Materials Research Part A 101(9) (2013) 2621-2629. [257] J.W. Durham, A. Rabiei, Deposition, heat treatment and characterization of two layer bioactive coatings on cylindrical PEEK, Surface and Coatings Technology 301 (2016) 106- 113. [258] M. Pino, N. Stingelin, K.E. Tanner, Nucleation and growth of apatite on NaOH- treated PEEK, HDPE and UHMWPE for artificial cornea materials, Acta Biomaterialia 4(6) (2008) 1827-1836. [259] Y. Zhao, H.M. Wong, W. Wang, P. Li, Z. Xu, E.Y.W. Chong, C.H. Yan, K.W.K. Yeung, P.K. Chu, Cytocompatibility, osseointegration, and bioactivity of three- dimensional porous and nanostructured network on polyetheretherketone, Biomaterials 34(37) (2013) 9264-9277. [260] S.W. Ha, M. Kirch, F. Birchler, K.L. Eckert, J. Mayer, E. Wintermantel, C. Sittig, I. Pfund-Klingenfuss, M. Textor, N.D. Spencer, M. Guecheva, H. Vonmont, Journal of Materials Science Materials in Medicine 8(11) (1997) 683-690. [261] G. Mestres, M.-P. Ginebra, Novel magnesium phosphate cements with high early strength and antibacterial properties, Acta Biomaterialia 7(4) (2011) 1853-1861. 185 [262] C. Großardt, A. Ewald, L.M. Grover, J.E. Barralet, U. Gbureck, Passive and ActiveIn VitroResorption of Calcium and Magnesium Phosphate Cements by Osteoclastic Cells, Tissue Engineering Part A 16(12) (2010) 3687-3695. [263] U. Klammert, E. Vorndran, T. Reuther, F.A. Müller, K. Zorn, U. Gbureck, Low temperature fabrication of magnesium phosphate cement scaffolds by 3D powder printing, Journal of Materials Science: Materials in Medicine 21(11) (2010) 2947-2953. [264] X. Liu, P. Chu, C. Ding, Surface modification of titanium, titanium alloys, and related materials for biomedical applications, Materials Science and Engineering: R: Reports 47(3-4) (2004) 49-121. [265] H.M. Kim, T. Himeno, M. Kawashita, J.H. Lee, T. Kokubo, T. Nakamura, Surface potential change in bioactive titanium metal during the process of apatite formation in simulated body fluid, J Biomed Mater Res A 67(4) (2003) 1305-9. [266] L. Ouyang, Y. Zhao, G. Jin, T. Lu, J. Li, Y. Qiao, C. Ning, X. Zhang, P.K. Chu, X. Liu, Influence of sulfur content on bone formation and antibacterial ability of sulfonated PEEK, Biomaterials 83 (2016) 115-126. [267] T. Kokubo, H. Takadama, How useful is SBF in predicting in vivo bone bioactivity?, Biomaterials 27(15) (2006) 2907-2915. [268] S. Gronthos, A.C.W. Zannettino, S.E. Graves, S. Ohta, S.J. Hay, P.J. Simmons, Differential Cell Surface Expression of the STRO-1 and Alkaline Phosphatase Antigens on Discrete Developmental Stages in Primary Cultures of Human Bone Cells, Journal of Bone and Mineral Research 14(1) (1999) 47-56. [269] F. Kircelli, M.E. Peter, E. Sevinc Ok, F.G. Celenk, M. Yilmaz, S. Steppan, G. Asci, E. Ok, J. Passlick-Deetjen, Magnesium reduces calcification in bovine vascular smooth 186 muscle cells in a dose-dependent manner, Nephrology Dialysis Transplantation 27(2) (2011) 514-521. [270] V. Viereck, H. Siggelkow, S. Tauber, D. Raddatz, N. Schutze, M. Hüfner, Differential regulation of Cbfa1/Runx2 and osteocalcin gene expression by vitamin-D3, dexamethasone, and local growth factors in primary human osteoblasts, Journal of Cellular Biochemistry 86(2) (2002) 348-356. 187