Comparative Investigation on the Adhesion of Hydroxyapatit
description
Transcript of Comparative Investigation on the Adhesion of Hydroxyapatit
Author's Accepted Manuscript
Comparative Investigation on the Adhesion ofHydroxyapatite coating on Ti-6Al-4V Implant:A Review Paper
E. Mohseni, E. Zalnezhad, A.R. Bushroa
PII: S0143-7496(13)00168-1DOI: http://dx.doi.org/10.1016/j.ijadhadh.2013.09.030Reference: JAAD1419
To appear in: International Journal of Adhesion & Adhesives
Accepted date: 17 July 2013
Cite this article as: E. Mohseni, E. Zalnezhad, A.R. Bushroa, ComparativeInvestigation on the Adhesion of Hydroxyapatite coating on Ti-6Al-4VImplant: A Review Paper, International Journal of Adhesion & Adhesives, http://dx.doi.org/10.1016/j.ijadhadh.2013.09.030
This is a PDF file of an unedited manuscript that has been accepted forpublication. As a service to our customers we are providing this early version ofthe manuscript. The manuscript will undergo copyediting, typesetting, andreview of the resulting galley proof before it is published in its final citable form.Please note that during the production process errors may be discovered whichcould affect the content, and all legal disclaimers that apply to the journalpertain.
www.elsevier.com/locate/ijadhadh
1
Comparative Investigation on the Adhesion of Hydroxyapatite coating on Ti-
6Al-4V Implant: A Review Paper
E. Mohsenia, E. Zalnezhadb, A. R. Bushroac*
a,b,c Center of Advanced Manufacturing and Material Processing, Department of Engineering
Design and Manufacture, Faculty of Engineering, University of Malaya, Kuala Lumpur, 50603,
Malaysia
[email protected] [email protected] Corresponding author: [email protected]
Abstract
Hydroxyapatite (HA) has been used in clinical bone graft procedures for the past 25 years. Although a biocompatible material, its poor adhesion strength to substrate makes it unsuitable for major load-bearing devices. Investigations on various deposition techniques of HA coating on Ti-6Al-4V implants have been made over the years, in particular to improve its adhesion strength to the metal alloy and its long-term reliability. This review comprehensively analyzes nine techniques mostly used for deposition of HA onto Ti-6Al-4V alloys. The techniques reviewed are Plasma sprayed deposition, Hot Isostatic Pressing, Thermal Spray, Dip coating, Pulsed Laser deposition (PLD), Electrophoretic deposition (EPD), Sol-Gel, Ion Beam Assisted deposition (IBAD), and Sputtering. The advantages and disadvantages of each method over other techniques are discussed. The adhesion strength and the factors affecting the adhesion of HA coating on Ti-6Al-4V implants are also compared. Keywords: Adhesion; Hydroxyapatite; coating; Ti-6Al-4V implant
1. Introduction
Biological fixation is defined as the process where prosthetic components become firmly
bonded to the host bone by ongrowth or ingrowth without the use of bone cements [1-3]. In the
late 1960s, the concept of biological fixation of load-bearing implants using bioactive
2
hydroxyapatite (HA) coatings was proposed as an alternative to cemented fixation.
Hydroxyapatite (HA: Ca10(PO4)6(OH)2), a pure calcium phosphate phase, is a preferred
biomaterial for both dental and orthopedics use due to its favorable osteoconductive and
bioactive properties [4, 5]. HA has a similar chemical composition and crystal structure as the
apatite in the human skeletal system, and is therefore suitable for bone substitution and
reconstruction [6]. Furthermore, HA has shown significant success in implants due to its
favorable in vivo behavior [7, 8] and the presence of HA films prolongs the lifetime of
prostheses [9]. However, HA coatings are susceptible to fatigue failure, making it unsuitable for
load bearing implants [10, 11].
Nevertheless, there is a large demand for implants with excellent mechanical properties.
These implants should possess similar properties to the human bones, such as in the value of its
Young’s modulus, which result in less stress shielding effect [12] and extends its service life.
The implants can made into different shapes such as plates, rods, screws and pins [13].
Historically, titanium-based alloys are the most common material for this purpose since it is
known to be a tolerable metal in the human body [14].
Titanium (Ti) and its alloys are the most commonly used metallic materials for medical
implants in orthopedic and dental applications, due to their low density, high strength, non-
toxicity and excellent corrosion resistance [15]. However, there have been reports on
inflammatory reaction around these implants as a result from the creation of an avascular fibrous
tissue that encapsulated the implants [16, 17]. A coating of hydroxyapatite layer can be deposited
on the metal alloy to assist the osseointegration of these implants with surrounding tissues [16].
The bond strength between the coating layer and the metal substrate is a very critical
factor. Separation of the coating layer from the implant during service in the human body results
3
in adverse effects on the implants and the surrounding tissue caused by detached particles [18].
The main reason of using HA coating on metallic substrates is to keep the mechanical properties
of the metal such a load-bearing ability and, at the same time, to take advantage of the coating’s
chemical similarity and biocompatibility with the bone [19].
According to Blind et al., the HA coating allows rapid osteointegration as a result of
bone tissue bonding properties [20]. The first clinical results from HA coatings on titanium
dental implants were promising, showing excellent results, even with poor bone quality.
However, after a long period, mechanical failure would occur at the interface of HA and metallic
substrate [21]. The HA coating dissolves as a result of poor crystallized structure [22, 23],
decrease of adherence with the titanium surface and dramatic late implant failure [23, 24].
Moreover, HA itself has poor mechanical properties, with a bending strength of less than 100
MPa [25]. Thus, it can be concluded that the stability of the HA coating is the most critical
factor to ensure the success of this type of implant. Furthermore, the method used to deposit HA
powder onto the substrate could influence the coating characteristics such its adhesion strength
and reliability.
Several techniques have been used to create the HA coating on metallic implants, such as
plasma spraying process [26], thermal spraying [27], sputter coating [28], pulsed laser ablation
[29], dynamic mixing [30], dip coating [31], sol–gel [32], electrophoretic deposition [33],
biomimetic coating [34], ion-beam-assisted-deposition [35],and hot isostatic pressing [36].
Amongst the techniques listed, plasma spraying is the only process which is approved by the
Food and Drug Administration (FDA), USA for biomedical coatings due to its excellent coating
properties as compared to other coating processes [37]. However, plasma sprayed hydroxyapatite
coatings suffer from poor mechanical properties on tensile strength, wear resistance, hardness,
4
toughness and fatigue. Improvements in plasma spraying techniques over the years have
addressed many of these limitations. However, other coating methods are available which can be
used as an alternative to conventional techniques.
Limitations such as high porosity, poor uniformity in thickness, phase impurity, limited
crystallinity, and poor adhesion are common in HA coating. However, low coating adhesion
seems to be the major issue, limiting its extensive use for implants at a commercial scale [38-40].
Hence, improvement of bonding strength between the metallic substrate and ceramic coating is a
general requirement regardless of the techniques used.
This review focuses on adhesion strengths between HA coating and Ti-6Al-4V substrate,
fabricated using various techniques such as plasma sprayed deposition, hot isostatic pressing,
thermal spray, dip coating, pulsed laser deposition (PLD), electrophoretic deposition (EPD), sol-
gel, ion beam assisted deposition (IBAD). Parameters affecting the adhesion of coating and other
factors influencing the enhancement of bonding strength of coating surface and the substrate are
also discussed in detail.
2. Coating techniques
2.1 Plasma sprayed coating technique
Plasma spraying process involves melting of ceramics or metal powders using the heat of
ionized inert gas (plasma). The molten powders are then sprayed onto the surface to be coated,
forming the protective layer which provides a barrier against corrosion, wear or high
temperatures. The technique offers advantages such as low cost and rapid deposition rate [41,
42]. In addition, the risk of thermal degradation of the coating and substrate is much less than
5
other high-temperature processes since the gas in the plasma flame is chemically inert and the
target can be kept relatively cool [43]. However, plasma sprayed coatings suffers from poor
adhesion between the coatings and substrates [44], and the process may induce structural
changes in the microstructure of the coating material [45, 46].
2.1.1 Plasma sprayed hydroxyapatite (HA) coatings
Plasma spray was the first method used for the production of calcium phosphate coatings,
such as HA coating, due to its ease of application [26]. Plasma sprayed hydroxyapatite (HA)
coatings are biocompatible and able to bond directly to the bone [38], thus making plasma
spraying a favorable choice amongst the many techniques available for coating HA layers onto
metallic substrates [47]. Recent studies on plasma sprayed HA coatings (HACs) on titanium have
shown encouraging results in orthopedic implant applications. These studies reported that the
new bone could appose directly onto the HA coatings and very good adhesion between the
HACs and the new bone can be obtained [48-51]. The plasma sprayed HA coatings have also
assisted in overall quick bone recovery [52].
Nevertheless, the brittle nature of the HA coating makes it prone to crack and fracture,
non-uniformity in density of coating [53], wear of the coated layer, weak mechanical adhesion to
the substrate [44, 54], and alteration of structure [55].
Overall, plasma sprayed coating did not show significant improved long-life performance, better
mechanical integrity and reliability over uncoated implants [56, 57]. An alternative to plasma
spraying is the pulsed laser deposition (PLD) which enables the stoichiometric transfer of
sintered HA yields to form a thin and adherent bioactive coating on titanium substrate surface
[58].
2.1.2 Adhesion of plasma-sprayed hydroxyapatite (HA) coatings on Ti-6Al-4V
6
It is well understood that, the determination of the adhesion between the substrate and
coating is one of the main concerns when using plasma spraying techniques [59]. It’s quite
complicated that how coating adheres to a substrate and by today it is not completely understood.
Many theories describe the mechanism of adhesion, although, there is no single clear
interpretation for all adhesion behaviors [60]. Many factors seem to affect the adhesion: (1) Van
der Waals physical interaction forces mechanical anchorage; (2) mechanical anchorage; (3)
metallurgical processes and (4) chemical interaction [59].
Recent reports on alternative orthopedics implant fixation utilizing plasma sprayed HA
coatings (HACs) on Ti-6Al-4V have shown that the new bone was able to appose directly onto
the HA coatings, which resulted in a very good adhesion between the HACs and the new bone
[48-51]. From the viewpoint of materials science, characteristics of HACs are varied with the
spraying parameters such as phase composition, the microstructure, OH-ion content,
crystallinity, and the ration of calcium to phosphorus for the HACs. Among these parameters,
high bonding strength of HACs can be achieved by high spraying power due to a denser
microstructure caused by the greatest extent of coating melting.
Y.C. Yang, et al. experimented on six plasma sprayed HA on Ti- 6Al- 4V substrates by
varying the cooling conditions and the substrate temperatures [61]. The residual stresses and
bonding strengths were measured by XRD “ sin2φ “ technique and a standard adhesion test
(ASTM C-633). Results of the bonding strength evaluation shows that the HA coating with the
lowest residual stress exhibited a higher bonding strength (9.18±0.72 MPa).
The deposition stress and thermal stress are the two major sources of residual stresses in
plasma sprayed coating. Deposition stresses are produced during the cooling of sprayed particles
after solidification. Thermal stresses are generated from differential thermal contraction during
7
the post-fabrication cooling phase after coating [62, 63]. The residual stresses are present near
the interface of metal substrate and coating [64-66], due to the difference of thermal expansion
coefficients between both materials [62, 63]. These stresses may vary with substrate cooling
effects, parameters of spraying [62, 63], and coating thickness [67, 68]. Generally, it is believed
that the increased thickness of coating and the temperature of the specimen during plasma
spraying are the mains reasons for the rise in the residual stress.
In addition, high-powered, dense plasma sprayed HA coatings would have stronger
bonding strength than those sprayed using low power. The result is not solely due to the
difference in adhesive strength of HA coating. The value for bonding strength reflects the
combination of both cohesive (within the coating layers) and adhesive (coating to substrate)
strengths of a coating [61]. In a similar study, Tsui, Y. et al. claimed that the cohesive and
adhesive integrity of the coatings influence the long term performance of HA coated implants
considerably [69]. The adhesive strength is usually evaluated based on surface roughness,
coating properties, residual stress, and the mechanical interlocking between the coating and the
substrates, whereas the cohesive strength is determined by coating properties, such as
microstructure and crystallinity [61].
The bonding strength of HA coatings on metallic substrates can be evaluated using
several techniques such as the standard tensile adhesion test [69], interfacial indentation test [58],
tensile adhesion strength (TAS) [61], and indentation method [63]. However, there are
limitations on these techniques to accurately measure the adhesion strength, such as a probability
of penetration of glue into the coating layer , and a dependence of coating failure to the flaw
distribution at the edge of specimen [69]. However, Z. Mohammadi et al. have demonstrated that
the tensile adhesion strength test measured by the standard adhesion test ISO 13779-4, can be
8
used in conjunction with the interface indentation test to predict the effects of different
parameters on the adhesion properties of the HA coating by plasma spraying [70]. In general, the
HA coatings with the densest structure (i.e. lowest porosity, and predominantly amorphous
phase) have a higher tensile adhesion strength than those of lower density [61, 71]. The report by
Z. Mohammadi et al. [70] also showed that the tensile adhesion strength was in the range of ~25
MPa for HA coated on the Ti-6Al-4V.
2.2 Hot isostatic pressing technique
Hot isostatic pressing (HIP) is an enabling technology providing an efficient method for
the densification of ceramic powders which allows production of net-shape ceramics with
superior and consistent properties [72]. HIP is an alternative method of producing an HA coating
on a Ti substrate in which pressurized gas is used to exert the required load at the desired
temperature. This requires a gas-tight metal or glass encapsulation around the porous HA coated
implant [73]. In the HIP process, pressure and temperature are applied to the workpiece
simultaneously [74-77].
In hot isostatic pressing, high-pressure levels can be obtained since there is no
dependency on rigid tools with limited strength (such as graphite tools in uniaxial hot pressing)
to transmit the pressure to the body. Typical operating pressure ranges are 100-320 MPa (15-50
ksi), with temperatures exceeding above 2000 °C conducted in large industrial equipment [72].
The advantages of HIP are better temperature control as compared to uniaxial hot pressing, and a
9
resultant homogeneous material structure and properties. The reduced sintering temperature
enables control or even avoidance of grain growth and undesirable reactions. A very high
uniformity of properties as well as freedom from directionality can also, if desired, be obtained
[72]. Some researchers have used HIP treatments to densify plasma sprayed coatings, and results
have shown that HIP is useful in reducing the porosity and improving the physical and
mechanical properties of ceramic coatings [78].
Thus, the most important advantage of the hot isostatic pressing is the ability to control
the size and shape of the product to a very high precision without costly diamond machining
operations. Under ideal conditions no change of shape (just a change of scale) of the body
occurs. It has an inherent ability to produce parts with exceptionally accurate shape, virtually
with no dimensional or shape limitation [72].
2.2.1 Hot isostatic pressing of hydroxyapatite (HA) Coatings
Reports shows that, sort of problems such as porosity and crack appearance are
conducted with existing dc plasma sprayed Ha coating on Ti-6Al-4V [79]. In medical
applications some amount of porosity is needed for bony tissue to grow into the coating for
efficient fixation. In addition, the crack propagation needs to be healed for the composite coating
to have reasonable mechanical strength during usage. In this sense, HIP introduces its profound
advantages by improving the adhesion and physical properties of the plasma sprayed HA
coatings as a post- treatment [79].
2.2.2 Adhesion of hot isostatic pressing of hydroxyapatite (HA) coatings on Ti-6Al-4V
K. Khor et al. [79] investigated the effect of post-sprayed HIP on plasma sprayed HA on
Ti-6Al-4V. Fig.1 shows the bond strengths of HA coated Ti-6Al-4V for the plasma sprayed
samples, and after HIP treatment at different temperatures with respect to the coating thickness.
10
In general, it was shown that the bonding strength generally improves after HIP. It is also shown
that the adhesion strength decreases with increasing coating thickness. The enhancement of the
adhesion strength in the 20 wt.% HA coating after HIP is apparent for coating below 160 μm.
However, the result of adhesion strengths for coatings thicker than 160 μm show that HIP may
have adverse effects on the coating strengths.
2.3 Thermal spray coating technique
Thermal spray technology is a group of coating processes that provide functional surfaces
to protect or improve the performance of a substrate or component. Many types and forms of
materials can be thermal sprayed to provide protection from corrosion, wear, and heat; to restore
and repair components; and for a variety of other applications [80]. Thermal spraying of
biomedical coating is a relatively new class of applications for thermal spray coating as
compared with other industrial applications, [81]. Thermal spray processes are grouped into three
major categories: flame spray, electrical arc spray, and plasma arc spray. These energies sources
are used to heat the coating material (in powder, wire, or rod form) to a molten and semi-molten
state. The resultant heated particles are accelerated and propelled towards a prepared surface by
either process gases or atomization jets. A schematic diagram of thermal spray coating is
illustrated in Fig. 2.
2.3.1 Thermal spray deposition of hydroxyapatite (HA) coatings
Thermal spraying of HAP on implant devices can be compared with plasma spray coating
technique, having the advantage of high deposition rate and low cost [82, 83]. Thermal spray
technique has the ability to produce HA layer with thickness from 30 to 200 µm depending on
11
the coating condition However films deposited by thermal spraying suffers from poor coating–
substrate adherence and non-uniform crystallinity which reduces the lifetime of implants [84,
85]. In addition, thermal spray requires high sintering temperature which may result in crack
propagation on the surface of the coating [86-90].
2.3.2 Adhesion of thermal spray deposition of hydroxyapatite (HA) coatings on Ti-6Al-4V
J. Hsiung et al. [91] have evaluated the applications and characterizations of biological
coating such as hydroxyapatite on titanium alloy, particularly Ti-6Al-4V, in artificial knee joint
by thermal spray coating technology. The process involves melting of HA powder and guiding
the molten mass via a jet stream of air to form a coating on the substrate, as shown in Fig. 3. The
thermal spray process conditions of the three coating materials are shown in Table 1,
highlighting the important parameters affecting the quality of the coating, such as inert gas
compositions, currents, voltage levels, powder feeding rates, and spraying distances.
The tensile test is commonly used to evaluate the bond strength in accordance to ASTM
C633 standard method [92]. A bonding strength of 33.2 Mpa was obtained by J. Hsiung et al.
[91] for the HA coating on Ti-6Al-4V by thermal spraying technique. In comparison, this result
is not satisfactory when compared to other coatings for the same application such as Al2O3,
ZrO2. In addition, results of microstructure analysis shows that the HA coatings suffers from
spalling, interface separation and high levels of porosity.
Several pre and post-treatments of HA coating were also investigated by J. Hsiung et al. [91].
Treatment conditions include high pressure cleaning, ultrasonic cleaning and cryogenic
treatments. [92]. Table 2 shows the result of the bond strength test using ASTM C633 [93],
12
indicating the bond strengths of samples cleaned with high pressure air are lower as compared
with those ultrasonically cleaned, and the bond strengths with cryogenic treatments are better
than those without cryogenic treatments. The result shows that the inclusion of ultrasonic
cleaning and cryogenic treatments can effectively improve the coating properties.
2.4 Dip coating technique
Dip coating involves the deposition of a wet liquid film by withdrawal of a substrate from
a liquid coating medium. The complete process of film formation involves several stages, as
shown in Fig 3. The process starts by immersion of the substrate in the solution of the coating
material. When the substrate is withdrawn from the coating fluid, a coherent liquid film is
entrained on the surface of the substrate. A thin layer of coating is formed upon evaporations of
solvents and any accompanying chemical reactions in the liquid film. Normally an additional
post-treatment such as curing or sintering is required to obtain the final coating. Dip coating
technique is similar to sol-gel coating technique, although the process is significantly faster in
which a complete transition can be achieved within a few seconds if volatile solvents are used
[94]. Dip coating is fairly popular in the industry and in laboratory applications due to its low
cost, simple processing steps and high coating quality.
2.4.1 Dip coating of hydroxyapatite (HA) coatings
HA can be homogenously coated onto metal substrates to obtain coating thickness in the
range of 0.05 to 0.5 mm. The surface uniformity of HA can be controlled well using this
technique, as can be seen in the Fig. 4. In addition, the processing time for dip coating can be
very short, even for substrate with complex shapes. The coating layer is deposited on the surface
of the substrate without decomposition or reaction with the metal substrate. However, this
13
technique requires high sintering post-treatments which may induce crack formations on the
surface of the substrate [95].
2.4.2 Adhesion of dip coating of hydroxyapatite (HA) coatings on Ti-6Al-4V
B. Mavis et al. [96] had developed several compositions of the liquid coating medium for
the dip coating of HA on Ti-6Al-4V substrates, using chemically precipitated hydroxyapatite
precursor powders. To evaluate the adhesion strength, two steel cylinders 5 mm in diameter were
attached to both sides (coated and uncoated after the coating layer was ground off) of the dipped
strips by a thin layer of glue. The adhesive strengths were determined by measuring the tensile
stress needed to separate the cylinders from the strips [97]. It is reported that, the HA coatings
obtained were highly porous, with bonding strengths of more than 30 MPa.
2.5 Pulsed Laser Deposited Coating Technique
Laser processing is a rapid and clean process which can be used for surface modification
and controlled micro-structuring of materials. In biomedical applications, laser has been used to
modify the surface texture of materials to improve its bio-functionality [98-102]. Pulsed laser
deposition (PLD) technique can be used to grow ceramic thin films. By using appropriate laser,
thin films such as semiconductor films [103], cuprate superconductor films [103, 104], and
ferroelectric films [105] can be deposited onto substrates. PLD process involves using high
power laser energy to vaporize the bulk coating material from a target. The vaporized material is
ejected from the target and condenses on the substrate. Repeated laser pulses will result in the
deposition of the thin film as a coating on the substrate [106].
14
The formation of thin film by PLS can be separated into the following three stages [103,
107, 108]:
1- Laser radiation interaction with the target.
2- Dynamic ablation of the materials.
3- Deposition of the ablation materials with the substrate, nucleation and growth of a thin
film on the substrate surface.
One of the main advantages of PLD technique is the ability to retain the stoichiometry of
the target in the deposition films [107]. This is due to the high ablation rate which causes all
compounds or elements to evaporate at the same time [106]. Conversely, limitations of PLD
include the splashing of the particulates deposition on the film. Some methods have been
developed to decrease splashing problem since it is a major issues of the PLD [109]. One method
is to apply a mechanical particle filter that includes a velocity selector acting as a high-velocity
pass filter to eliminate slow-moving particulate. The second method is using a smooth, high-
density target which can be obtained by polishing the target surface before each coating run. The
third method is by applying a lower deposition rate or low energy density. Furthermore, the
deposited films have only a small area of structural and thickness uniformity, due to the angular
distribution of the ablation plume. Several methods have been proposed to scale up the PLD
process for large area thin films, such as laser beam rasterizing across a rotating target. [106].
High quality hydroxyapatite thin films deposited by the PLD was first reported in 1992
[105, 110] and since then the process have been improved significantly to obtained well adhered
and highly crystalline HA thin films under certain conditions [29, 111-113].
2.5.1 Pulsed laser deposited hydroxyapatite (HA) coatings
15
Preparing hydroxyapatite thin films by pulsed laser deposition allows accurate control of
hydroxyapatite growth parameters at low deposition temperatures and the ability to produce
highly crystalline HA coatings[16, 112]. In-vitro evaluations shows that these HA coatings are
stable and osteoinductive [114, 115]. Nanostructured hydroxyapatite layer having unique
biological properties can be obtained by selection of suitable parameters for the deposition
process [16].
2.5.2 Adhesion of pulsed laser deposited hydroxyapatite (HA) coatings on Ti-6Al-4V
Adhesion strength of HA coating on metals depends on the microstructure of the
substrate, the surface chemistry and the PLD process parameters such as laser power density and
substrate temperature. [58, 116-118]. Various surface modification techniques have been used to
improve the metal-ceramic interface such as nitridation, surface oxidation and ion implantation
[119-123].
Blind, et al. reported that adhesion of pulsed laser deposited HA films on titanium alloy
is due to the existence of an oxide, specifically titanium dioxide, at the interface between the
substrate and the coating layer [20]. Another report suggests that there may be some effects of
epitaxy between the oxide and coating [124]. Fernandez- Pradas et al. [54] commented whether
the presence of a titanium oxide interface would favour adhesion of the HA coating to the Ti-
6Al-4V substrate is still a cause for debate. Some authors consider that such a layer favours
adhesion [125, 126]. Other studies have attributed the weak adhesion in the first calcium
phosphate coatings deposited by PLD at high temperatures to the formation of a titanium oxide
layer during the process of pressure stabilisation [105]. A study of the adhesion strength in
coatings deposited by ion bombardment on passivated and non-passivated substrates, suggest
that this oxide layer should be as thin as possible [127].
16
C. Koch et al. [16] investigated pulsed laser deposition of hydroxyapatite on Ti-6Al-4V
for medical and dental applications. A pull-off testing method was used to determine the coating-
to-substrate adhesion strength. Garcia-sanz et al. had also examined hydroxyapatite films
prepared using pulsed laser deposition using a pull-off test based upon a modified ASTM C-633
procedure [128]. The measured tensile strength of hydroxyapatite grown at 480 °C was 58 MPa
and failure was observed at the coating–substrate interface. Wang et al. obtained tensile bonding
strength values within the range of 30 MPa and 40 MPa for hydroxyapatite coatings grown on
Ti-6Al-4V in an argon–water atmosphere at 500–600 °C [129]. Zeng et al. determined the bond
strength values for hydroxyapatite films grown using 3rd harmonic YAG:Nd lasers (λ=355 nm),
and 4th harmonic YAG:Nd lasers (λ=266 nm) on Ti-6Al-4V substrates in an argon–water
atmosphere at 500–520 °C [116]. Films grown on unpolished titanium substrates had tensile
strength values of 30 MPa while films grown on polished titanium possess lower tensile
strength values of 20 MPa.
In a study to enhance the bonding strength of HA , Nelea et al. [110] utilized a TiN
interfacial layer between the Ti-6Al-4V substrate and HA coating. The study reported that the
adhesion was improved due to better bonding of HA to TiN, which is a ceramic, and then to the
surface of metallic substrate. Man et al. [40] and Cui et al. [130] described the utilization of a
pre-treatment process which included etching and laser surface nitriding on titanium to produce a
TiN dendritic scaffold network structure. This coralline-like structure provides additional surface
area for interlocking of the coating material.
Man et al. [119] reported the influence of pre-treatments on the adhesion of the HA
coating to the substrate. Five types of pre-treatments, shown in Fig. 5 were: (i) mirror finished
specimen, (ii) 60 grit grinded SiC paper (specimen 2), (iii) 320 grit grinded SiC paper (specimen
17
3), (iv) mirror finish with 1-μm diamond paste (specimen 1), and (v) 10 s etching with Knoll
solution after polishing (specimen 4). The surface roughness of the specimens were determined
using a profilometer (Taylor Hobson Surtronic 25) and the adhesion strengths between HA
coatings and the substrates were evaluated in accordance to ASTM C-633 [131]. The maximum
adhesion strength obtained was ~ 16 MPa for specimen 5 (nitrided+ etching).
Figure 5 shows the adhesion strength of deposited HA on different pre-treated specimens
and surface roughness. Generally, an increase in surface roughness increases the adhesion
strength. Based on these results, it can be concluded that significant enhancement in the adhesion
strength of pulsed laser deposited HA on Ti-6Al-4V can be obtained by laser surface nitriding
and subsequent etching [119].
A related study has concluded that a controlled surface microstructure can be obtained by
using few laser pulses without affecting the bulk mechanical property of titanium substrate [132].
Figure 6 plots the average surface roughness values, measured after laser treatment and after HA
coating versus their initial roughness.
Figure 7 compares the adhesion strengths of HA coating on substrates treated with 500–
18,000 laser pulses with those of untreated, polished titanium. The adhesion of HA to the
substrate is examined in accordance to ISO 20502:2005(E) [133] using a micro-scratch tester
(micro-combi tester; CSM Instrument Switzerland) equipped with a diamond Rockwell tip of
100 μm [132]. It was found that in all cases, the laser treated substrates would have higher
bonding strengths, which imply that the surface roughness directly influences the adhesion
strength. Varying the laser pulses would affect the surface morphology. Figure 6 shows that the
roughness increases with the increase in the number of laser pulses, which starts from ~ 0.4 μm
at 500 laser pulses/min up to ~ 1 μm at 12000 pulses/min. However, there is a significant
18
decrease in the roughness value for laser pulses in the range of 12000 to 18000 laser pulses/min.
Low rate of laser pulses ( 500, 1,000, and 3,000 pulse/min) would only etch the surface and may
not be able to control the surface roughness. The surface roughness is under control only after ~
3000 pulses/min. A surface with controlled structure/pattern is obtained using 18,000 pulses
[132].
The polished surface of specimen does not have much adhesion strength to the coating.
However, once the surface is treated with laser, the surface roughness increases which results in
increased adhesion (from 0 to 1000 pulses/min) due to the initial material removal from the
surface. However, at this stage, certain regions are unaffected by the laser and a control over the
adhesion at this stage is not predictable. Once the laser pulses reaches ~3000 pulses/min, the
surface attains a certain level of smoothness since the large, number of pulses would completely
remove the original top surface to uncover a fresh coating surface. Therefore, the morphology
and adhesion can be controlled by the number of laser pulses (3000 to 18000) [132]. Figure 7
shows the trend of adhesion strength versus number of laser pulse, showing that the adhesion
strength would gradually increase until 1000 pulse/min, then decreases in between 1000 to 3000
pulse/min, and increases again past 3000 pulse/min. The highest adhesion strength obtained was
10.87 N and 11.21 N at 2,000 and 18,000 laser pulses respectively, while untreated substrate
showed a lower adhesion strength value of 4.57 N [132].
HA coatings by PLD exhibit good biocompatible and mechanical properties making it
suitable for medical implants. PLD HA coatings, on titanium alloy such as Ti-6Al-4V, resulted
in higher adhesion between the coating and substrate and have only minor undesirable phase
under optimal conditions [54, 106].
2.6 Electrophoretic deposition coating technique
19
Electrophoretic deposition (EPD) is a process in which particles in a suspension is coated
onto an electrode under the effect of an electric field [134]. The colloidal particles suspended in a
liquid medium migrate under the influence of an electric field (electrophoresis) and are then
deposited onto an electrode. Electrophoretic deposition (EPD) is particularly advantageous for
ceramic film and coatings as well as laminar ceramic composites applications [134-137].
Furthermore, the method used low-cost equipment, easy to set-up, and is able to coat complex
shapes and patterns. A high degree of control on the coating results can be achieved by
regulating the deposition conditions and the ceramic powder size and shape [138]. EPD is a
cheaper method than chemical vapor deposition, sol gel deposition, and sputtering for producing
films of a wide range of thickness, from less than 1 mm to more than 100 µm thick [139].
However, limitations of the technique includes low adhesion strength, and cracking on the coated
surface due to post-process shrinkage.
EPD has shown its potential use in biomedical applications in recent years [140-142].
The interest in electrophoresis for biomedical applications [143-147] stems from a variety of
reasons such as the possibility of stoichiometric deposition, high purity material to a degree not
easily achievable by other processing techniques and the possibility of forming coatings and
bodies of complex shape [140]. Considering all advantages and disadvantages of this technique,
electrophoretic deposition is one of the favorable coating techniques which can be utilized for
hydroxyapatite coating.
2.6.1 Electrophoretic deposition hydroxyapatite (HA) coatings
20
There is a growing interest in processing of HA powders using EPD technique, owing to
its uniformity and good sinterability of the deposits, possibility of impregnation of porous
substrates, and composite consolidation [142, 148]. However, reports on the use of EPD for
depositing HA on titanium substrate are thus far, relatively limited. Nie et al. [149] and Soares
et al. [150] have used EPD to deposit HA on Ti-6Al-4V substrates and have obtained uniform
thin coating with good mechanical strength. Stoch et al. [146] have also coated HA on titanium
implants with intermediate layer of silica. EPD process of HA is a colloidal process where HA
powders are deposited directly from a stable colloid suspension by using a DC electric field [25].
Electrophoretic deposition of HA can be processed at room temperature or lower, which
avoids problems related to formation of amorphous phases. The nature of the bond is more
metallurgical rather than mechanical, thus HA coatings using EPD are expected to have
improved adhesion strength as compared to thermal sprayed techniques. However, a major
drawback is the presence of porosities which may later on leads to corrosion and delamination of
the titanium caused by penetration of body fluids into the substrate. Post-treatment high
temperature sintering can be utilized to minimize the porosity by increasing the coating density.
Unfortunately, cracks in the coating can form during high temperature sintering due to the
difference in the thermal expansion coefficients and large reduction of the pore volume between
the titanium and HA [151].
For nanostructured materials, the mismatch in thermal expansion coefficient is not a
significant problem [152]. In nano-ceramics, the thermal expansion coefficient is fairly matched
with the metal alloy because the large quantity of atoms located at the grain boundary improves
mobility [152-154]. However, the success of electrophoretic deposited HA has been limited to
conventional materials in the range of micron-sized grains [134, 140, 154]. Limitations on the
21
mechanical properties of the micron size HA are poor fracture toughness, adhesion, and
compressive strengths. There is a need for the HA coating and the substrate to have sufficient
interfacial bond strength since the coating would endure high interfacial stresses during in vivo
service.
2.6.2 Adhesion of electrophoretic deposited hydroxyapatite (HA) coatings on Ti-6Al-4V
Zhang, et al. [151] have developed a unique room temperature EPD process to deposit
nanostructured HA coating having adhesion strength of 50-60 MPa, which is 2-3 times better
than thermal-sprayed HA coating. The interfacial bond strength was measured in accordance to
ASTM Standard F 1501-95 using a tensile tester [151]. The corrosion resistance of this
nanostructured HA is 50 to 100 times higher than conventional HA coating. Fig. 8 shows the
corrosion resistance results for both EPD coatings and thermal sprayed coatings, where the
corrosion current of n-HA coating is 50-100 times smaller than the thermal sprayed coating in
simulated human body fluid at room temperature.
High quality HA nano-coating can be produced using EPD technique. The adhesion
stress obtained was 60 MPa, measured using a direct-pull-tests, which exceeds the 50 MPa
requirements of the food and drug administration (FDA) [155]. A 2 months in vitro testing also
showed that the bonding strength of the EPD n-HA coating on the titanium alloy was able to be
maintained in the range of 50-60 MPa, which is significantly better than plasma sprayed HA
coatings [151].
Ma, et al. [139] reported that HA particles were successfully deposited onto a titanium
substrate via a single electrophoretic deposition. Good adhesion between the coating and
substrate was verified by scanning electron miscopy examination and shear strength tests,
22
following methods outlined by Wei et al. [148] and ASTM standard F1044-87. The shear stress
of the HA coating after sintering at 1000 ˚C was 3.34 MPa, indicating a good adhesion of the
coating has been obtained. Figures 9 and 10 show SEM micrographs of the cross-section and the
surface deposit of the 1000 ˚C sintered HA coating, respectively. It can be seen that a layer of
HA coating as thick as 400 µm has adhered well into titanium substrate and no delamination or
crack was observed at both the interface and the surface. The deposition was found to be uniform
with the coating thickness maintained consistently along the surface of the sample. No
observable crack, which is one of the common problems of EPD, was detected. It is believed that
the good deposition result is due to the stable and dispersed HA suspension used for the
deposition [148].
Studies on EPD coating of HA on titanium alloys show that particle size is an important
factor for the process as the mobility of the charged particles is proportional to the size of the
particles [156]. Ferrari et al. [157] have also reported that the charges, hence the conductivity of
the suspension, play an essential role and has an optimum value for the process. Nevertheless,
the colloidal stability of the suspension could also be a main factor to obtain good coating
uniformity and bonding strength in the EPD process [142].
Like many similar techniques for coatings involving ceramics, EPD coating of HA
requires a densification stage involving the sintering of the coated implants. This requirement
poses a dilemma, especially since high sintering temperature is sometime necessary. Low
sintering temperatures results in weak bond with low-density coatings whereas high sintering
temperatures can lead to the degradation of the HA and the metal substrate (oxidation and
impaired mechanical properties) as a result of the metal substrate catalyzing decomposition of
the HA to anhydrous calcium phosphates [158, 159].
23
A high sintering temperature may also lead to phase transformation and grain growth of
the metal substrate, causing significant decrease in mechanical properties. It has been
demonstrated that the mechanical properties of these titanium alloys degrade significantly when
heated above 1050˚C [138]. Therefore, it is recommended to keep the densification temperatures
below 1000˚C to minimize degradation of the HA and the metal substrate.
The sintering phase for EPD implants improves densification and the bonding of the
coating. However, HA may decompose in the process [160]. An interlayer can be used in
between the HA and the metal substrate to moderate the problem of HA decomposition. Nie et
al. deposited a dense layer of titanium dioxide (TiO₂) as the inner layer between HA top layer
and titanium alloy substrate to achieve a very good combination of mechanical integrity,
chemical stability and bioactivity [149].
Kumar and Wang [161] investigated the coating of TiO₂ powders on Ti-6Al-4V
substrates as the first layer, followed by the HA- TiO₂ composite layers of different weight
ratios, coated onto the TiO₂ layer. Wei et al. [138] studied on the adhesion strength of HA
coating in which HA powders are used as both inner and outer layer. Hence, no change occurred
in the structure of coating layers. Sintering was also applied after the deposition of every single
layer. In the HA coating on TiO₂ deposited substrate, the decomposition of HA is decreased; and
generally adhesion of coating, which is tested according to ASTM F1044-99, was enhanced with
the reduction of voltage value used for TiO₂ coating [160]. Table 3 shows the result of adhesion
strengths of HA coated samples with and without TiO2 inner layer deposited using different
voltages.
2.7 Sol-Gel derived coating technique
24
The Sol-gel method is one of the simplest technique to manufacture thin films which can
produce almost any single or multicomponent oxide coating on glass or metals [162, 163]. Sol-
gel derived coating can be used for optical, electronic, magnetic or coating with chemical
functions [164]. Sol-gel derived ceramic films are widely used as a protective layer against
corrosion and oxidation of stainless steel [165], Ag [166], and Al [167] substrates. The sol-gel
process involves the formation of solid materials, mainly inorganic non-metallic materials from
solution. This can be a solution of monomeric, oligomeric, polymeric or colloidal precursors
[168].
The sol-gel process, shown in Fig. 11 [169], consist of: (i) producing a homogeneous
solution of purified precursors in an organic solvent which can be mixed with the reagent used in
the next step or water; (ii) shaping the solution to the ‘sol’ form by using treatment with a
suitable reagent, e.g. water for oxide ceramics; (iii) changing the sol to a ‘gel’ by
polycondensation; (iv) converting the gel to the finally preferred shape like thin film, fiber, and
(v) finally converting( sintering) the shaped gel to the desired ceramic material at temperatures
(~500˚C) much lower than those required in the conventional procedure of the melting the oxides
together [168, 170-172].
Olding et al. [172], reports that sol-gel techniques has considerable advantages such as :
1. Ability to produce thin bond-coating to provide excellent adhesion between the metallic
substrate and the top coat.
2. Corrosion resistance performance due to ability to form thick coating.
3. Ability to shape materials even complex geometries in the gel state.
4. Production of high purity samples.
5. Low temperature sintering, usually in the range of 200 to 600 °C [173].
25
6. A simple, economic and effective method to produce high quality coatings.
However, the sol-gel technique has disadvantages such as high permeability, low wear-
resistance, and difficult porosity control, which has limited its utilization in the industry. For
crack-free coating, the maximum thickness of the coating is only 0.5 μm [172]. Furthermore,
trapped organics during the thermal process would result in coating failure. Recent advancement
in high substrate sensitive sol-gel also suffers from thermal expansion mismatch. Nevertheless,
there is a wide room for improvement in the technique and further investigation should be done
to improve this highly potential method for biomaterial coating.
2.7.1 Sol-Gel derived hydroxyapatite (HA) coatings
The sol-gel is a low temperature process, thus does not suffer from the implications of
structural instability of hydroxyapatite at elevated temperatures [174-177]. A major processing
stage involves solution chemistry, whereby a sol is produced from suitable alkoxides or salts to
yield a hydroxyapatite composition upon heating [178].
Gross, et al. [178] described that the production of sol-gel hydroxyapatite coatings on
titanium substrates using alkoxide precursors requires more control on firing temperature and the
aging time. X-ray diffraction of the coatings heated to various temperatures, as illustrated in Fig.
12, indicated that the titanium substrate would start to oxidize at temperatures starting at 800˚C.
Thus for sol-gel hydroxyapatite coating, it is suggested that the processing temperature should be
around 800˚C to reduce possible phase transformation in the metallic substrate as well as the
occurrence of oxidation. [178]. Nanograined hydroxyapatite coating with an average grain size
of 50 nm was achieved using this technique. Figure 13 shows a scanning electron micrograph of
26
a coating fired to 800˚C for 10 min. Densification of the coating can then be obtained with a
longer duration of firing at 800˚C.
Fabrication of sol-gel deposited HA on implants HA[173, 179, 180] requires extremely
stringent processing parameters, particularly for the thermal processing phase such as the
duration and calcining temperature, chemical compositions of the precursor , types of substrate,
and number of HA- coated layers. Major issues include the crystalline phases, adhesion strength
and biocompatibility of the resulted coatings.
2.7.2 Adhesion of Sol-Gel deposited hydroxyapatite (HA) coatings on Ti-6Al-4V
Tests have shown that pure HA suffers relatively high dissolution rate in simulated body
fluid that would affects its long-term stability. High dissolution may lead to disintegration of the
coatings and hinder the fixation of implant to the host tissue [181, 182]. To address this issue,
Zhang, et al. [183] incorporated fluorine ion, which exists in human bone and enamel, into HA
crystal structures. Mixing of fluorine into HA, or fluoridation, decreases the solubility of HA
while still maintaining its biocompatibility [184].
Zhang, et al. [183] have successfully deposited dense, crack-free fluoridated
hydroxyapatite (FHA, Ca10(PO4)6 (OH)2−xFx) coatings ( 1.5 μm) through sol–gel dip coating on
Ti-6Al-4V substrates. Scratch testing has shown an increase of over 35% in the adhesion
strengths of the coating to Ti-alloy. The increase in adhesion is more prominent for high
annealing temperatures. This increase is most likely due to the formation of chemical bonding at
the interface and the incorporation of fluorine in HA which provided relief of thermal mismatch.
Figure 14 illustrates the coefficient of friction in terms of relative voltage as a function of
normal load while scratching (a) pure HA coating; (b) fluoridate HA (FHA6) coating on Ti-6Al-
27
4V. At the beginning of the scratch and because of the “soft” nature of the coating, coefficient of
friction increases as load increases. The fluctuation in the diagram, before point 1, is caused by
the surface roughness. After point 1, the indenter would start to advance into the coating,
resulting in a sharp increase in friction coefficient. The indenter would completely peel off the
coating and scratches the substrate as the load increases to point 2, or 370 mN for pure HA
(shown in curve a), which results in a sudden increase in friction at about 470 mN for FHA6.
Comparison of curves (a) and (b) in Fig. 14 shows that curve “b” appears to have less
fluctuation before the indenter completely digs in and the adhesion of coating and substrate is
better since there is a slower gradient rise after the indenter digs in. A sharp increase of friction
would indicate a brittle peeling-off of the coating from the substrate surface. Since curve “b”
lacks the sharp change in friction, it is thus a more ductile interface and subsequently have better
coating-substrate bonding than those of curve “a” (pure HA) [185].
Figure 15 shows the “upper critical load”,Lc, of all FHA coatings as a function of firing
temperature and fluorine. Both firing temperatures and fluorine content seems to have a
significant effect on the adhesion strength of the coating. Increasing firing temperature or
fluorine concentration results in a dramatic raise of the critical load. For coatings with the same
amount of fluorine content, higher adhesion is due to higher annealing temperatures. Similarly, at
the same firing temperature, adhesion strength increases with fluorine content.
Zhang, et al. [186], in similar studies [183], reported that FHA is a potential
replacement for pure HA coating on metallic implants due to FHA’s significant biocompatibility
and resistance to biodegradation [184, 187]. Ding, et al. [188] identified two critical aspects as
the main contributors for long-term stability of the ceramic-coated implants: high adhesion
strength of substrate to coating and low solubility of the coating. Incorporation of fluoride ions
28
into HA lattice structure results in reduction of HA solubility. However, reports on adhesion
improvements, especially on adhesion studies after in vitro dissolution test have yet to be studied
extensively. In vitro dissolution tests can be used to investigate the influence of dissolution
behavior on the adhesion. Zhang, et al. [186], evaluated the adhesion of FHA coated on Ti-6Al-
4V using sol-gel technique before and after dissolution tests. The dissolution tests were
conducted by soaking FHA coatings in a Tris-buffered physiological saline solution (TPS)
(0.9%NaCl, pH7.4) at a fixed temperature of 37 °C for a duration of 3 weeks (Fig. 16). It worth
to mention that the “P “value in the Fig. 16 is one-way ANOVA test was conducted to assess the
statistical significance of the adhesion and toughness results.
Figure 16 shows the nominal adhesion strength between the coating and the Ti-6Al-4V
substrate. “Adhesion failure” and “cohesion failure” cannot be recognized by “nominal”.
Without fluoridation (sample F0), the adhesion strength is about 19 MPa. Fluoridated samples
(F1 and F2) show significant increase in adhesion strength to about 26–27 MPa. Zhang, et al.
[186] concluded that, the strength range starts from about 19 MPa for pure hydroxyapatite (x=0)
up to about 26 MPa for x=1. However, after 21 days of soaking the coating in Tris-buffered
physiological saline solution, the adhesion strength increases to about 30 MPa for pure HA and
to over 40 MPa for FHA.
Comparing the sol–gel and thermal spraying methods for the same FHA coatings on Ti-
6Al-4V, Gu, et al. [189] described that after soaking, the adhesion strengths of thermal sprayed
specimens tends to decline, with reductions up to 75%. For example the adhesion strength had
decreased from 27 MPa before soaking down to 19 MPa after soaking in synthetic body fluid
(SBF) for 2 weeks. The reduction in adhesion strength of thermal spray deposited HA coatings
is probably due to the presence of cracks in the coating [83].
29
Cheng, et al. [190] used a pull-out method and scanning scratch technique to evaluate the
bonding strength of FHA coatings on Ti-6Al-4V. Figure 17 shows the result of measurements
by pull-out strength, showing the strength is about 11 MPa for pure HA coating (FHA0), with
considering of F content, the strength intensifies up to about 22 MPa, and then decreases to
around 17–18 MPa. Coating peeling-off value is about 390mN for pure HA. In contrast, the
coating peeling- off increases with increasing F content, 447 mN for FHA1, 450 mN for FHA2,
449 mN for FHA3 and 478mN for FHA4. The result of the study confirms that the presence of F
in FHA coatings has improved the adhesion strength [190].
2.8 Ion beam assisted deposition technique
Surface modification techniques based on the bombardment method have been used since
the mid-1970s, and many have been developed and are now widely used for surface engineering
of materials such as ceramics, bioceramics, and metals. Examples of such methods are ion beam
deposition, ion beam mixing and ion beam assisted deposition (IBAD) [191-195].
IBAD is a vacuum deposition process based on the combination of ion beam
bombardment and physical vapor deposition. The major characteristic of IBAD is the
bombardment with a specific energy ion beam during coating deposition. Many parameters can
affect the composition, mechanical properties, chemical properties, and structural properties of
the deposited coating in the IBAD process. The most important processing parameters in IBAD
are evaporation rate or sputtering rate, coating materials, ion species, ion beam current density
and ion energy [196].
IBAD has the ability to prepare bio-coatings with considerably higher adhesive strength
as compared to traditional coating methods. The high adhesive strength is the result of
interaction between the substrate and coating atoms, assisted by ion bombardment. This results
30
in an atomic intermixed zone in the substrate-coating interface [196]. IBAD process is highly
reliable, reproducible and is conducted at low substrate temperature, without unfavorably
affecting the bulk substrate characteristics. Furthermore, the process has superior control over
coating microstructure and chemical composition [197].
2.8.1 Ion beam assisted deposition of hydroxyapatite (HA) coatings
As it mentioned earlier, there are several methods to make HA coating on Ti-6Al-4V,
among which plasma spraying is the most frequently used [198, 199]. However, long-term
clinical follow-up has demonstrated that there are significant deficiencies in the plasma-sprayed
HA coatings. The limited cohesive strength of the coatings and the limited strength of the
coating-metal substrate interface are the main problem with plasma-sprayed coting technique.
Moreover, heat treatments in plasma-sprayed HA coatings results in cracks in the coating layer
because of thermal expansion mismatch between the metal substrate and coated layer. This leads
to a severe decreasing in bond strength [200-203]. In order to produce more permanent bone-
bonding calcium phosphate coatings, ion beam assisted deposition (IBAD) is introduced as an
alternative technique for plasma spraying technique. Previous studies shows that implants coated
with HA by the IBAD method demonstrate a very good adhesion to the substrate [204].
2.8.2 Adhesion of ion beam assisted deposition of hydroxyapatite (HA) coatings on Ti-6Al-4V
In the IBAD process, a wide atomic intermixed zone between the coatinsg material and
the substrate can be created, assisted by the bombardment with energetic ions during deposition.
This creates a strong adhesion of the coating to the substrate [205, 206]. Ohtsuka et al. first used
50 keV Ca+ implantation into Ti, followed by Ca+ IBAD to deposit HA coating on Ti substrate
and has obtained higher adhesive strength than conventional methods [204]. It has been
demonstrated that Ca+ implantation alone into Ti was unable to provide the bioactive surface.
31
Cui, et al. [207] proposed using Ar+ IBAD to form highly adhesive hydroxyapatite
coatings on titanium alloy. The coatings prepare by IBAD was compared to those formed by ion
beam sputtering deposition (IBSD) of calcium phosphate coatings. Scratch test is used to
investigate the adhesive strength of the IBSD and IBAD coatings on the substrates. Figure 18
shows the typical Fz- Fy curves of scratch test results for the specimens prepared by IBSD and
IBAD. Markers “A” and “B” indicate the points of the first occurrence of coating detachment
from the substrate. Fz and Fy , as the normal and tangential forces respectively, are affecting the
diamond indenter during the test. A load speed of 2000 gf /min was chosen for the tests. The
results have shown that the critical loads were 660 gf for IBSD and 1050 gf for IBAD samples.
Generally, it was seen that the adhesive strength of the coatings prepared by IBAD technique is
almost twice that of the IBSD coatings.
It has been shown that the adhesion strengths of coatings prepared by IBSD and plasma
sprayed technique are generally similar [127]. Thus, it can be deduced from the comparative
results between IBSD and IBAD that the adhesive strength of IBAD coatings would be
reasonably higher than that of plasma sprayed depositions. The main benefit of IBAD is the
improved adhesion strength due to the wide atomic intermixed zone at the interface of the
coating and substrate [204, 206]. Thus, the issue of low adhesion strength, which exists in
plasma sprayed coatings can be significantly eliminated by using the IBAD technique [207].
Choi, et al. [35] have used an Ar ion beam in the coating of HA on Ti-6Al-4V deposited
by IBAD technique. Figure 19 illustrates the bonding strength as a function of the ion beam
current, before and after the heat treatment. Increasing the current would increase the ion
bombardment and broadens the atomic intermixed zone during the deposition. This results in the
increase of adhesion strength between the substrate and coating layer [207].
32
Several studies have shown that heat treatments would decrease the bond strength [200-
203]. Figure 20 shows the SEM micrographs of the coating layer before and after heat
treatments. The morphologies were found to be relatively similar regardless of the current level.
Before heat treatment, the layer was rather featureless, as shown in Fig. 20 (A). The lines at the
interface are Wallner lines frequently observed when hard coating layers are detached from a
metal substrate [208]. However, after heat treatment, the layer became severely cracked, as
shown in Fig. 20 (B). This is probably due to the thermal expansion mismatch between the
coating and the substrate [209]. These cracks are the main reason for the reduction in bond
strengths. The micrograph also reveals that the metal surface was slightly oxidized, presumably
by OH in the coating layer [209]. Overall significant improvement in the bond strength is
resulted by Choi, et al. [35] using an Ar ion beam while deposition.
Hamdi and Ide-Ektessabi [197] have proposed the deposition of hydroxyapatite layer
using a combination of technique of IBAD and simultaneous vapor deposition (SVD), namely
ion-beam-assisted simultaneous vapor deposition (IBASVD). Figure 21 illustrates the result of
coating detachments for two sets of IBASVD samples as function of different annealing
temperatures. Both types of samples resulted in similar curve patterns with the minimum
detachment forces recorded at 700 ˚C annealing temperature and the maximum adhesion strength
at 1200 ˚C. In all cases the adhesion strength for the 260 µA/cm2 sample was higher than the 180
µA/cm2 sample. In general, the recorded data for both samples are extremely higher than the
maximum adhesion strength obtainable by the SVD samples, which was less than 100 mN [210].
It is suggested that the increase in adhesion strength was the result of the formation of a mixed
layer between the substrate and the HA film , consisting of a gradient fill of Ca, P and the
element of the substrate [207, 211]. Hamdi and Ide-Ektessabi [197] described that the energetic
33
ions assisted the reactions between the migrated atoms and the substrate atoms to generate an
intermixed layer, which have specific properties different from the deposited films and the
substrate. It was also understood that high current density of ion beam resulted in a wider atomic
intermixed zone, which consequently improved the overall adhesion strength.
2.9 Sputter coating technique
Sputter deposition is a physical vapor deposition (PVD) method of depositing thin films
by sputtering. This involves ejecting material from a source, known as a "target", onto a
"substrate" such as a silicon wafer. It was reported that initial sputtering using multi-component
ceramic targets such as superconducting oxides, HA and other CaP materials would produce
coatings whose chemistries were different upon deposition than the bulk target [212, 213].
Sputtering utilizes a gas plasma (argon, neon, krypton or xenon) to remove material from a
negatively charged target which is then deposited as a thin film coating onto the substrate.
Studies have shown successful deposition of thin HA layers on titanium substrates using RF
magnetron sputtering [214].
2.9.1 Sputter coating of hydroxyapatite (HA) coatings
Sputtering techniques have been used to deposit homogeneous thin films coatings of high
adhesion strength with thicknesses ranging from 0.5 to 3 μm. However, sputter coated HA films
on metals were found to be of low crystallinity [214-216]. The low crystallinity increases the rate
of dissolution of the coating in the living body. Post-treatment thermal process can be used to
crystallize the film, hence reducing the possibility of dissolution. However, conventional thermal
treatment in the electric furnace increases the likely formation of cracks and may degrade the HA
films.
2.9.2 Adhesion of sputter coating of hydroxyapatite (HA) coatings on Ti-6Al-4V
34
K. Ozeki et al. [217] compared the thermal treatments of the HA coated on titanium alloy
substrate prepared by sputter coating with those prepared by plasma spraying technique. The
substrates were sandblasted using Al2O3 (125-180 µm) abrasive before coating. The specimens
were post-treated with a hydrothermal process for 24 hours. The film thickness obtained for
sputter coating was 1.2 µm while the thickness for plasma spraying was 60-100 µm.
Figure 22 shows the shear strength results of the sputter coating, the plasma sprayed
coatings and the non-coated columns over a period of time. The sputter coating showed the
highest bonding strength overall with recorded strengths of 3.3 ± 0.2, 5.7 ± 0.5, and 8.6 ± 1.6
MPa after two, four, and 12 weeks, respectively. The plasma sprayed coatings resulted in
strength values of 1.9 ± 0.25, 4.0 ± 0.3, and 6.6 ± 0.7 MPa, respectively, for the same period of
time. The strength values of the non-coated columns were 0.4 ± 0.3 and 1.1 ± 0.3 MPa after four
and 12 weeks, respectively. The strength of the sputter coating exceeded that of the plasma
sprayed coating by more than 70, 40, and 30% after a period of two, four and twelve weeks,
respectively. K. De Groot et al. reported that coating thicknesses above 100 μm were associated
with fatigue failure under tensile loading [218]. According to S. Hasegawa et al., thin plasma
sprayed coatings are bound more strongly than thick coatings [219].
S.J. Ding et al. [220] investigated on a series of thin (<10 μm), single layered HA/Ti
coatings deposited on Ti-6Al-4V substrate using an RF magnetron-assisted sputtering system.
For the experiments, six HA/Ti targets with different compositions (95HA/5Ti, 90HA/10Ti,
85HA/15Ti, 75HA/25Ti, 50HA/50Ti, and 25HA/75Ti) were prepared. Generally it was found
that the coating with higher Ti contents resulted higher adhesion strengths. The highest adhesion
strength (of the 25HA/ 75Ti coating), evaluated using a Sebastian adhesion test system
(Sebastian Five, Quad Group, Spokane, WA) [127] was even higher than 80 MPa, which
35
exceeded the maximum value achievable using the bonding resin in the pull-out test. Table 4,
reports the adhesion strength and their corresponding failure point for different compositions and
Fig 23 shows the adhesion strength for each composition.
The high adhesion strength of sputtered monolithic HA coating is higher than most
plasma sprayed HA coatings [221, 222], and is believed to be attributed to the sputter cleaning
and ion bombarding processes. The sputter cleaning process would remove contaminants and
adsorbed gas molecules from the surface of the substrate to produce a clean, highly active
surface [223]. The ion bombarding process during sputtering would enhance atomic diffusion
and mixing near the interface region [207, 224]. Mechanical interlocking effect may have
contributed to the higher average adhesion strength of coating sputtered on the rougher surface
(Ra = 0.7 mm) as compared to the lower value obtained for the smoother surface (Ra = 0.06
mm). However this effect was not as significant for sputtering with Ti-containing targets.
Results from S.J. Ding et al. [220] have shown that all coatings had adhesion strengths
between 60 and 80 MPa. Furthermore if the sputtering uses a target comprising of more than 15
vol % Ti, the resulting coating adhesion strength and hardness were significantly higher than
those of monolithic HA coating.
3. Discussion
There have been numerous studies on coatings of hydroxyapatite (HA) onto Ti-6Al-4V
because of its significant utilization in orthopedic prostheses and implants. Table 5 summarizes
the previous discussion on the various techniques for coating of HA on Ti-6Al-4V, with
comparison on their advantages and disadvantages.
Plasma spraying is the most frequently investigated method to coat HA onto Ti-6Al-4V
specimen, [198, 199]. Plasma spray is the first method used for HA coating, owing to its ease of
36
application [26]. Moreover, the determination of the adhesion between the coating and the
substrate has been always a main concern when using plasma spraying technique [59]. High
spraying power results in high adhesion strength of HACs due to significant melting of the
coating material which forms dense microstructure. However, the high-temperature process can
lead to phase transformation and grain growth of the metal substrate which may cause significant
decrease in the mechanical properties of the metal.
Results of the study [61] has established the relationship between residual stress and
bonding strength especially for plasma sprayed hydroxyapatite coatings. This stress in the
coating is influenced by the spraying parameter, coating thickness [67, 68], and substrate cooling
effect (i.e. temperature of substrate) [62, 63]. Generally, the residual stresses increase with the
increase in the thickness of coating and the temperature of the specimen during plasma spraying.
Moreover, high-power sprayed HA coatings generally possess higher adhesion strength than
those sprayed with lower power. In some cases, the adhesion of the plasma sprayed HA can be
significantly improved by a subsequent hot isostatic pressing operation.
The adhesion strength is a reflection of the combination of cohesive (within the coating
layers themselves) and adhesive (coating to substrate) strengths of a coating [61]. The cohesive
strength is obtained by coating properties, such as the microstructure and crystallinity, but the
adhesive strength is mostly influenced by coating properties, such as surface roughness, residual
stress, and the mechanical interlocking between substrate and HACs [61].
Overall, it was found that plasma sprayed coating has not improved the service-life
performance of uncoated implants. In addition, there are issues with poor reliability and
mechanical integrity [56, 57]. The pulsed laser deposition (PLD) is a better alternative than the
plasma spray technique because the PLD transfers sintered HA stoichiometrically to deposit a
37
thin adherent coating onto titanium substrate surface [58]. The substrate temperature is lower in
PLD as compared to plasma spray and different calcium phosphate compositions can be
deposited by changing the parameters of deposition [112, 114, 225]. In addition, undesirable
phases of HA coatings by PLD are reduced under optimal conditions and generally have better
coating to substrate adhesion [54, 106].
TiO2 and TiN layers can be used as an interfacial layer between coating and the metal
substrate as reported in studies related to the adhesion of crystalline PLD HA thin films on Ti-
6Al-4V substrates [20]. Some authors consider that this interfacial layer favours adhesion due to
better bonding of HA to TiN which is then,directly bonded to the substrate [125, 126]. These
layers can be created using pre-treatment processes, such as laser surface nitriding and etching
on titanium, which have been reported to improve the bonding strength of the coating. Thus,
laser surface nitriding and subsequent etching of the substrate is an effective pre-treatment
method for improving the adhesion strength of HA coated onto Ti-6Al-4V by PLD [119].
EPD is a technique which is gaining attention due to its ability to economically produce
films of a wide range of thicknesses as compared to conventional methods such as thermal
spraying, sol gel deposition, and sputtering [139]. Moreover, EPD of HA has ability to be
processed at room temperature, reducing the possibility of formation of the amorphous phase in
HA. The good uniformity and bonding strength results is mostly due to the colloidal stability of
the suspension [142]. The EPD technique can also produce nanostructured HA coating having
bond strength 2-3 times better than thermal sprayed HA coating.
Similar to PLD, studies have shown that an intermediate layer, such as silica or TiO2,
improves the adhesion strength of coating fabricated using EPD [146]. Dense titanium dioxide
(TiO2) films possess a very good combination of bioactivity, chemical stability and mechanical
38
integrity [149]. A TiO2 inner layer would also reduce the decomposition of HA and increases the
and overall adhesion strength of coating [160].
The sol-gel technique is a simple technique which can create single or multicomponent
oxide coating on glass or metals [162, 163]. However, there is a coating thickness limit of 0.5 μm
[172]. Fluoridation of HA can enhance the coating’s resistance to biodegradation while still
maintaining good biocompatibility [184, 187]. An increase in fluoridation ratio would increase
the adhesion strength by about 40%. The strength range for FHA is about 26 MPa which is
higher than the value of the bonding strength of 19 MPa for pure hydroxyapatite. The fracture
toughness increases about 200 to 300% and the scratch test results in adhesion improvement of
35 % for fluoridated HA coatings as compared to pure hydroxyapatite coating [183, 186, 190].
The enhancement in adhesion strength is believed to be caused by the formation chemical
bonding at the interface and the relief of thermal mismatch resulting from the incorporation of
fluorine (F) into the HA structure.
Dip coating can be generally compared with sol-gel coating technique. The technique is
simple, economical and is able to generate high coating quality. Dip coating process is rapid,
where the complete transition can be completed within a few seconds or less if volatile solvents
are used.
IBAD technique can deposit highly adhesive HA coating on Ti-6Al-4V due to atomic
interactions between the substrate and coating materials, assisted by ion bombardment [196]. The
main advantage of IBAD compare to other methods, such as IBSD or plasma spraying, is that
there is a wide atomic intermixed zone at the coating-substrate interface which significantly
improves the adhesive strength of the coating. Heat treatment of IBAD coated samples reduces
39
the adhesion strength, due formation of cracks in the layer and the thermal expansion mismatch
between the coated layer and the metal substrate [200-203].
Figure 24 shows adhesion strength values of HA coatings on Ti-6Al-4V coated using
various techniques. The sputtering technique has the highest adhesion of coating to the substrate
compares to other methods which can be attributed to the sputter cleaning and ion bombardment
processes.
4. Conclusion
Adhesion strength of HA on Ti-6Al-4V substrate has been reviewed in detail. Nine
common techniques of deposition such as plasma sprayed deposition, hot isostatic pressing,
thermal spray, dip coating, pulsed laser deposition (PLD), electrophoretic deposition (EPD), sol-
gel, ion beam assisted deposition (IBAD), and sputtering were evaluated and discussion were
made on the coating parameters affecting the adhesion strength of the coating. Advantages and
disadvantages of each method were discussed and a quantitative comparison was made on the
different techniques of HA coating on Ti-6Al-4V substrate. Based on this review, the best
adhesion of HA coating to substrate is obtained by sputtering deposition technique while the
worse bonding strength was obtained by PLD at 1000 laser pulses. Using an interfacial layer
(such as TiO2 or TiN) as the initial coating layer on the substrate followed by HA coating layer
can enhance the bonding strength. Pretreatments such as nitriding, followed by etching, can
enhance the adhesion strength in PLD. Moreover, post-treatments also have similar effects on
other techniques such as IBAD and thermal spray.
Acknowledgement
40
The authors would like to acknowledge the University of Malaya for providing the necessary
facilities and resources for this research. This research was fully founded by the Ministry of
Higher Education, Malaysia with the high impact research grant number of
um.c/625/1/HIR/MOHE/ENG/27.
References:
[1] W.L. Jaffe, D.F. Scott, Current Concepts Review-Total Hip Arthroplasty with Hydroxyapatite-Coated Prostheses*, The Journal of Bone & Joint Surgery, 78 (1996) 1918-34. [2] K.A. Thomas, Hydroxyapatite coatings, Orthopedics, 17 (1994) 267. [3] P. Ducheyne, J.M. Cuckler, Bioactive ceramic prosthetic coatings, Clin. Orthop. Relat. Res, 276 (1992) 102-14. [4] R.S. Corpe, D.E. Steflik, R.Y. Whitehead, M.D. Wilson, T.R. Young, C. Jaramillo, Correlative experimental animal and human clinical retrieval evaluations of hydroxyapatite (HA)-coated and non-coated implants in orthopaedics and dentistry, Critical reviews in biomedical engineering, 28 (2000) 395-8. [5] R. Sakkers, R. Dalmeyer, R. Brand, P. Rozing, C. Van Blitterswijk, Assessment of bioactivity for orthopedic coatings in a gap�healing model, Journal of Biomedical Materials Research, 36 (1998) 265-73. [6] D. Tadic, F. Peters, M. Epple, Continuous synthesis of amorphous carbonated apatites, Biomaterials, 23 (2002) 2553-9. [7] R.E. Holmes, R. Bucholz, V. Mooney, Porous hydroxyapatite as a bone-graft substitute in metaphyseal defects. A histometric study, The Journal of bone and joint surgery. American volume, 68 (1986) 904. [8] R.W. Bucholz, A. Carlton, R. Holmes, Interporous hydroxyapatite as a bone graft substitute in tibial plateau fractures, Clinical orthopaedics and related research, (1989) 53. [9] J. Wilson, L. Hench, D. Greenspan, Bioceramics, vol. 8, Pergamon-Elsevier, 1994. [10] M. Thomas, R. Doremus, M. Jarcho, R. Salsbury, Dense hydroxylapatite: fatigue and fracture strength after various treatments, from diametral tests, Journal of Materials Science, 15 (1980) 891-4. [11] G. With, H. Dijk, N. Hattu, K. Prijs, Preparation, microstructure and mechanical properties of dense polycrystalline hydroxy apatite, Journal of Materials Science, 16 (1981) 1592-8. [12] M.-F. Hsieh, L.-H. Perng, T.-S. Chin, Hydroxyapatite coating on Ti6Al4V alloy using a sol–gel derived precursor, Materials chemistry and physics, 74 (2002) 245-50. [13] K. Kawagoe, M. Saito, T. Shibuya, T. Nakashima, K. Hino, H. Yoshikawa, Augmentation of cancellous screw fixation with hydroxyapatite composite resin (CAP) in vivo, Journal of Biomedical Materials Research, 53 (2000) 678-84. [14] H.A. Luckey, F. Kubli Jr, Titanium alloys in surgical implants, ASTM International, 1983.
41
[15] D.M. Brunette, Titanium in medicine: material science, surface science, engineering, biological responses, and medical applications, Springer Verlag, 2001. [16] C. Koch, S. Johnson, D. Kumar, M. Jelinek, D. Chrisey, A. Doraiswamy, et al., Pulsed laser deposition of hydroxyapatite thin films, Materials Science and Engineering: C, 27 (2007) 484-94. [17] V. Nelea, C. Morosanu, M. Iliescu, I. Mihailescu, Microstructure and mechanical properties of hydroxyapatite thin films grown by RF magnetron sputtering, Surface and Coatings Technology, 173 (2003) 315-22. [18] S. Wang, W.R. Lacefield, J.E. Lemons, Interfacial shear strength and histology of plasma sprayed and sintered hydroxyapatite implants< i> in vivo</i>, Biomaterials, 17 (1996) 1965-70. [19] P. Silva, J. Santos, F. Monteiro, J. Knowles, Adhesion and microstructural characterization of plasma-sprayed hydroxyapatite/glass ceramic coatings onto Ti-6A1-4V substrates, Surface and Coatings Technology, 102 (1998) 191-6. [20] O. Blind, L.H. Klein, B. Dailey, L. Jordan, Characterization of hydroxyapatite films obtained by pulsed-laser deposition on Ti and Ti-6AL-4v substrates, Dental materials, 21 (2005) 1017-24. [21] R. Poser, F. Magee, J. Kay, A. Hedley, In-vivo Characterization of a Hydroxylapatite Coating, Transactions of the 16th annual meeting of the society for biomaterials, 1990, p. 170. [22] W. Tong, Z. Yang, X. Zhang, A. Yang, J. Feng, Y. Cao, et al., Studies on diffusion maximum in x�ray diffraction patterns of plasma�sprayed hydroxyapatite coatings, Journal of Biomedical Materials Research, 40 (1998) 407-13. [23] F. Garcia, J. Arias, B. Mayor, J. Pou, I. Rehman, J. Knowles, et al., Effect of heat treatment on pulsed laser deposited amorphous calcium phosphate coatings, Journal of Biomedical Materials Research, 43 (1998) 69-76. [24] C. Watson, D. Tinsley, A. Ogden, J. Russell, S. Mulay, E. Davison, Implants: A 3 to 4 year study of single tooth hydroxylapatite coated endosseous dental implants, British dental journal, 187 (1999) 90-4. [25] C. Wang, J. Ma, W. Cheng, R. Zhang, Thick hydroxyapatite coatings by electrophoretic deposition, Materials Letters, 57 (2002) 99-105. [26] K. De Groot, R. Geesink, C. Klein, P. Serekian, Plasma sprayed coatings of hydroxylapatite, Journal of Biomedical Materials Research, 21 (2004) 1375-81. [27] K.A. Gross, C.C. Berndt, Thermal processing of hydroxyapatite for coating production, Journal of Biomedical Materials Research, 39 (1998) 580-7. [28] S.-J. Ding, Properties and immersion behavior of magnetron-sputtered multi-layered hydroxyapatite/titanium composite coatings, Biomaterials, 24 (2003) 4233-8. [29] L. Cleries, E. Martınez, J. Fernandez-Pradas, G. Sardin, J. Esteve, J. Morenza, Mechanical properties of calcium phosphate coatings deposited by laser ablation, Biomaterials, 21 (2000) 967-71. [30] M. Yoshinari, Y. Ohtsuka, T. Dérand, Thin hydroxyapatite coating produced by the ion beam dynamic mixing method, Biomaterials, 15 (1994) 529-35. [31] S. Kaciulis, G. Mattogno, A. Napoli, E. Bemporad, F. Ferrari, A. Montenero, et al., Surface analysis of biocompatible coatings on titanium, Journal of electron spectroscopy and related phenomena, 95 (1998) 61-9. [32] P. Li, K.d. Groot, T. Kokubo, Bioactive Ca 10 (PO 4) 6 (OH) 2− TiO 2 composite coating prepared by sol-gel process, Journal of Sol-Gel Science and Technology, 7 (1996) 27-34.
42
[33] Y. Han, T. Fu, J. Lu, K. Xu, Characterization and stability of hydroxyapatite coatings prepared by an electrodeposition and alkaline�treatment process, Journal of Biomedical Materials Research, 54 (2000) 96-101. [34] P. Habibovic, F. Barrere, C.A. Blitterswijk, K. Groot, P. Layrolle, Biomimetic hydroxyapatite coating on metal implants, Journal of the American Ceramic Society, 85 (2002) 517-22. [35] J.-M. Choi, H.-E. Kim, I.-S. Lee, Ion-beam-assisted deposition (IBAD) of hydroxyapatite coating layer on Ti-based metal substrate, Biomaterials, 21 (2000) 469-73. [36] H. Wie, H. Herø, T. Solheim, Hot isostatic pressing-processed hydroxyapatite-coated titanium implants: light microscopic and scanning electron microscopy investigations, The International journal of oral & maxillofacial implants, 13 (1998) 837. [37] M. Mittal, S. Nath, S. Prakash, Improvement in mechanical properties of plasma sprayed hydroxyapatite coatings by Al2O3 reinforcement, Materials Science and Engineering: C, (2013). [38] L. Sun, C.C. Berndt, K.A. Gross, A. Kucuk, Material fundamentals and clinical performance of plasma�sprayed hydroxyapatite coatings: A review, Journal of Biomedical Materials Research, 58 (2001) 570-92. [39] R.R. Wang, G.E. Welsch, O. Monteiro, Silicon nitride coating on titanium to enable titanium–ceramic bonding, Journal of Biomedical Materials Research, 46 (1999) 262-70. [40] H. Man, N. Zhao, Z. Cui, Surface morphology of a laser surface nitrided and etched Ti–6Al–4V alloy, Surface and Coatings Technology, 192 (2005) 341-6. [41] H. Herman, Plasma spray deposition processes, MRS Bulletin, 13 (1988) 60-7. [42] J.L. Ong, D. Chan, Hydroxyapatite and their use as coatings in dental implants: a review, Critical reviews in biomedical engineering, 28 (2000) 667. [43] H. Herman, Plasma-sprayed coatings, Scientific American;(USA), 259 (1988). [44] M. Filiaggi, N. Coombs, R. Pilliar, Characterization of the interface in the plasma-sprayed HA coating/Ti-6Al-4V implant system, Journal of Biomedical Materials Research, 25 (1991) 1211. [45] V. Palka, E. Poštrková, H. Koerten, Some characteristics of hydroxylapatite powder particles after plasma spraying, Biomaterials, 19 (1998) 1763-72. [46] L. Ellies, D. Nelson, J. Featherstone, Crystallographic changes in calcium phosphates during plasma-spraying, Biomaterials, 13 (1992) 313-6. [47] J. Lemons, Hydroxyapatite coatings, Clinical orthopaedics and related research, (1988) 220. [48] D. Buser, R. Schenk, S. Steinemann, J. Fiorellini, C. Fox, H. Stich, Influence of surface characteristics on bone integration of titanium implants. A histomorphometric study in miniature pigs, Journal of Biomedical Materials Research, 25 (2004) 889-902. [49] J. Jansen, J. Van De Waerden, J. Wolke, K. De Groot, Histologic evaluation of the osseous adaptation to titanium and hydroxyapatite�coated titanium implants, Journal of Biomedical Materials Research, 25 (2004) 973-89. [50] K. Søballe, E.S. Hansen, H. Brockstedt-Rasmussen, C.M. Pedersen, C. Bünger, Hydroxyapatite coating enhances fixation of porous coated implants: a comparison in dogs between press fit and noninterference fit, Acta Orthopaedica, 61 (1990) 299-306. [51] S.D. Cook, K.A. Thomas, J.F. Kay, M. Jarcho, Hydroxyapatite-coated porous titanium for use as an orthopedic biologic attachment system, Clin Orthop, 230 (1988) 303-12.
43
[52] A. Moroni, V. Caja, C. Sabato, E. Egger, F. Gottsauner-Wolf, E. Chao, Bone ingrowth analysis and interface evaluation of hydroxyapatite coated versus uncoated titanium porous bone implants, Journal of Materials Science: Materials in Medicine, 5 (1994) 411-6. [53] W. Lacefield, P. Ducheyne, J. Lemons, Bioceramics: Material Characteristics Versus In Vivo Behavior, The New York Academy of Science, New York, (1988) 72. [54] J. Fernández-Pradas, M. Garcıa-Cuenca, L. Cleries, G. Sardin, J. Morenza, Influence of the interface layer on the adhesion of pulsed laser deposited hydroxyapatite coatings on titanium alloy, Applied surface science, 195 (2002) 31-7. [55] P. Ducheyne, J. Cuckler, S. Radin, E. Nazar, Plasma sprayed calcium phosphate lining on porous metal coatings for bone ingrowth, Handbook of Bioactive Ceramics (Ninth Edition), Vol. IICRC Press, Boca Raton, FL, USA, (1990). [56] K. Khor, Y. Gu, C. Quek, P. Cheang, Plasma spraying of functionally graded hydroxyapatite/Ti–6Al–4V coatings, Surface and Coatings Technology, 168 (2003) 195-201. [57] J.G. Morales, R.R. Clemente, B. Armas, C. Combescure, R. Berjoan, J. Cubo, et al., Controlled nucleation and growth of thin hydroxyapatite layers on titanium implants by using induction heating technique, Langmuir, 20 (2004) 5174-8. [58] W. Lo, D. Grant, M. Ball, B. Welsh, S. Howdle, E. Antonov, et al., Physical, chemical, and biological characterization of pulsed laser deposited and plasma sputtered hydroxyapatite thin films on titanium alloy, Journal of Biomedical Materials Research, 50 (2000) 536-45. [59] D. Matejka, B. Benko, Plasma spraying of metallic and ceramic materials, John Wiley and Sons, Baffins Lane, Chichester, West Sussex, PO 19 1 UD, UK, 1989. 280, (1989). [60] K. Mittal, Adhesion measurement of thin films, thick films and bulk coatings, Astm International, 1978. [61] Y.-C. Yang, E. Chang, Influence of residual stress on bonding strength and fracture of plasma-sprayed hydroxyapatite coatings on Ti–6Al–4V substrate, Biomaterials, 22 (2001) 1827-36. [62] P. Scardi, M. Leoni, L. Bertamini, Residual stresses in plasma sprayed partially stabilised zirconia TBCs: influence of the deposition temperature, Thin solid films, 278 (1996) 96-103. [63] S. Takeuchi, M. Ito, K. Takeda, Modelling of residual stress in plasma-sprayed coatings: Effect of substrate temperature, Surface and Coatings Technology, 43 (1990) 426-35. [64] S. Brown, I. Turner, H. Reiter, Residual stress measurement in thermal sprayed hydroxyapatite coatings, Journal of Materials Science: Materials in Medicine, 5 (1994) 756-9. [65] A. Noutomi, M. Kodama, Y. Inoue, T. Ono, M. Kawano, N. Tani, Residual stress measurement on plasma sprayed coatings, Welding International, 3 (1989) 947-53. [66] B. Eigenmann, B. Scholtes, E. Macherauch, Determination of residual stresses in ceramics and ceramic-metal composites by X-ray diffraction methods, Materials Science and Engineering: A, 118 (1989) 1-17. [67] A. Evans, G. Crumley, R. Demaray, On the mechanical behavior of brittle coatings and layers, Oxidation of Metals, 20 (1983) 193-216. [68] J. Mencik, Mechanics of components with treated or coated surfaces(Book), Dordrecht, Netherlands: Kluwer Academic Publishers, 1996., (1996). [69] Y. Tsui, C. Doyle, T. Clyne, Plasma sprayed hydroxyapatite coatings on titanium substrates Part 1: Mechanical properties and residual stress levels, Biomaterials, 19 (1998) 2015-30. [70] Z. Mohammadi, A. Ziaei-Moayyed, A. Mesgar, Adhesive and cohesive properties by indentation method of plasma-sprayed hydroxyapatite coatings, Applied surface science, 253 (2007) 4960-5.
44
[71] S. Kweh, K. Khor, P. Cheang, An in vitro investigation of plasma sprayed hydroxyapatite (HA) coatings produced with flame-spheroidized feedstock, Biomaterials, 23 (2002) 775-85. [72] H.T. Larker, R. Larker, Hot isostatic pressing, Materials Science and Technology, (1991). [73] H. Herø, H. Wie, R.B. Jørgensen, I. Ruyter, Hydroxyapatite coatings on Ti produced by hot isostatic pressing, Journal of biomedical materials research, 28 (1994) 343-8. [74] M. Nakamura, Y. Kaieda, Microstructure and mechanical properties of sintered TiAl, Powder metallurgy, 31 (1988) 201-9. [75] L.M. Sheppard, Fabrication of ceramics; The challenge continues, American Ceramic Society Bulletin;(USA), 68 (1989). [76] D.W. Shin, K.-K. Orr, H. Schubert, Microstructure�Mechanical Property Relationships in Hot Isostatically Pressed Alumina and Zirconia�Toughened Alumina, Journal of the American Ceramic Society, 73 (1990) 1181-8. [77] F.F. Lange, Powder processing science and technology for increased reliability, Journal of the American Ceramic Society, 72 (1989) 3-15. [78] Y. Fu, A. Batchelor, Hot isostatic pressing of hydroxyapatite coating for improved fretting wear resistance, Journal of materials science letters, 17 (1998) 1695-6. [79] K. Khor, C. Yip, P. Cheang, Post-spray hot isostatic pressing of plasma sprayed Ti� 6Al� 4V/hydroxyapatite composite coatings, Journal of materials processing technology, 71 (1997) 280-7. [80] R.B. Heimann, Thermal spraying of biomaterials, Surface and Coatings Technology, 201 (2006) 2012-9. [81] H. Liang, B. Shi, A. Fairchild, T. Cale, Applications of plasma coatings in artificial joints: an overview, Vacuum, 73 (2004) 317-26. [82] Y. Yang, K.-H. Kim, J.L. Ong, A review on calcium phosphate coatings produced using a sputtering process—an alternative to plasma spraying, Biomaterials, 26 (2005) 327-37. [83] P. Cheang, K. Khor, Addressing processing problems associated with plasma spraying of hydroxyapatite coatings, Biomaterials, 17 (1996) 537-44. [84] C. Hanyaloglu, B. Aksakal, J. Bolton, Production and indentation analysis of WC/Fe–Mn as an alternative to cobalt-bonded hardmetals, Materials characterization, 47 (2001) 315-22. [85] M. Hamdi, S. Hakamata, A. Ektessabi, Coating of hydroxyapatite thin film by simultaneous vapor deposition, Thin Solid Films, 377 (2000) 484-9. [86] Z. Zyman, J. Weng, X. Liu, X. Zhang, Z. Ma, Amorphous phase and morphological structure of hydroxyapatite plasma coatings, Biomaterials, 14 (1993) 225-8. [87] T. Shunyan, J. Heng, D. Chuanxian, Effect of vapor�flame treatment on plasma sprayed hydroxyapatite coatings, Journal of biomedical materials research, 52 (2000) 572-5. [88] J. Weng, X. Liu, X. Zhang, Z. Ma, X. Ji, Z. Zyman, Further studies on the plasma-sprayed amorphous phase in hydroxyapatite coatings and its deamorphization, Biomaterials, 14 (1993) 578-82. [89] Z. Zyman, J. Weng, X. Liu, X. Li, X. Zhang, Phase and structural changes in hydroxyapatite coatings under heat treatment, Biomaterials, 15 (1994) 151-5. [90] C.M. Roome, C.D. Adam, Crystallite orientation and anisotropic strains in thermally sprayed hydroxyapatite coatings, Biomaterials, 16 (1995) 691-6. [91] J. Hsiung, J. Tzeng, K. Kung, H. Chen, A Study of Thermal Spray Coating on Artificial Knee Joints, Life Science Journal, 10 (2013). [92] J. Hsiung, H. Kung, H. Chen, K.-Y. Chang, Applications of Thermal Spray Coating in Artificial Knee Joints, Life Science Journal, 9 (2012).
45
[93] A. Standard, G65," Standard Test Method for Measuring Abrasion Using the Dry Sand/Rubber Wheel Apparatus, Annual Book of ASTM Standards, 3 (2008) 245-56. [94] J. Puetz, M. Aegerter, Dip coating technique, Sol-Gel Technologies for Glass Producers and Users, Springer, 2004, pp. 37-48. [95] T. Li, J. Lee, T. Kobayashi, H. Aoki, Hydroxyapatite coating by dipping method, and bone bonding strength, Journal of Materials Science: Materials in Medicine, 7 (1996) 355-7. [96] B. Mavis, A.C. Taş, Dip Coating of Calcium Hydroxyapatite on Ti�6Al�4V Substrates, Journal of the American Ceramic Society, 83 (2000) 989-91. [97] W. Weng, J.L. Baptista, Preparation and Characterization of Hydroxyapatite Coatings on Ti6Al4V Alloy by a Sol�Gel Method, Journal of the American Ceramic Society, 82 (1999) 27-32. [98] R. Lahoz, J.P. Espinós, G.F. de la Fuente, A.R. González-Elipe, “in situ” XPS studies of laser induced surface cleaning and nitridation of Ti, Surface and Coatings Technology, 202 (2008) 1486-92. [99] R. Delgado�Ruíz, J. Calvo�Guirado, P. Moreno, J. Guardia, G. Gomez�Moreno, J. Mate�Sánchez, et al., Femtosecond laser microstructuring of zirconia dental implants, Journal of Biomedical Materials Research Part B: Applied Biomaterials, 96 (2011) 91-100. [100] I. Zavestovskaya, Laser nanostructuring of materials surfaces, Quantum Electronics, 40 (2010) 942-54. [101] A. Kurella, N.B. Dahotre, Review paper: surface modification for bioimplants: the role of laser surface engineering, Journal of biomaterials applications, 20 (2005) 5-50. [102] A. Gaggl, G. Schultes, W. Müller, H. Kärcher, Scanning electron microscopical analysis of laser-treated titanium implant surfaces—a comparative study, Biomaterials, 21 (2000) 1067-73. [103] S.M. Kaczmarek, Pulsed laser deposition-today and tomorrow, Proc. SPIE, 1997, pp. 129-34. [104] H. Karl, B. Stritzker, Reflection high-energy electron diffraction oscillations modulated by laser-pulse deposited YBa_ {2} Cu_ {3} O_ {7-x}, Physical review letters, 69 (1992) 2939-42. [105] C.M. Cotell, D.B. Chrisey, K.S. Grabowski, J.A. Sprague, C.R. Gossett, Pulsed laser deposition of hydroxylapatite thin films on Ti�6Al�4V, Journal of Applied Biomaterials, 3 (2004) 87-93. [106] Q. Bao, C. Chen, D. Wang, Q. Ji, T. Lei, Pulsed laser deposition and its current research status in preparing hydroxyapatite thin films, Applied surface science, 252 (2005) 1538-44. [107] I.W. Boyd, Thin film growth by pulsed laser deposition, Ceramics international, 22 (1996) 429-34. [108] R. Kelly, A. Miotello, Mechanisms of pulsed laser sputtering, Pulsed laser deposition of thin films, (1994) 55-85. [109] J.T. Cheung, History and fundamentals of pulsed laser deposition, Pulsed laser deposition of thin films, (1994) 14-5. [110] V. Nelea, C. Ristoscu, C. Chiritescu, C. Ghica, I. Mihailescu, H. Pelletier, et al., Pulsed laser deposition of hydroxyapatite thin films on Ti-5Al-2.5 Fe substrates with and without buffer layers, Applied surface science, 168 (2000) 127-31. [111] J. Fernández-Pradas, L. Cleries, E. Martinez, G. Sardin, J. Esteve, J. Morenza, Calcium phosphate coatings deposited by laser ablation at 355 nm under different substrate temperatures and water vapour pressures, Applied Physics A: Materials Science & Processing, 71 (2000) 37-42.
46
[112] J. Fernández-Pradas, G. Sardin, L. Clèries, P. Serra, C. Ferrater, J. Morenza, Deposition of hydroxyapatite thin films by excimer laser ablation, Thin solid films, 317 (1998) 393-6. [113] B. Tucker, C. Cottell, R. Auyeungt, M. Spector, G. Nancollas, Pre-conditioning and dual constant composition dissolution kinetics of pulsed laser deposited hydroxyapatite thin films on silicon substrates, Biomaterials, 17 (1996) 631-7. [114] J.M. Fernandez-Pradas, L. Cleries, G. Sardin, J. Morenza, Hydroxyapatite coatings grown by pulsed laser deposition with a beam of 355 nm wavelength, Journal of materials research, 14 (1999) 4715-9. [115] L. Cleries, J. Fernández–Pradas, J. Morenza, Bone growth on and resorption of calcium phosphate coatings obtained by pulsed laser deposition, Journal of Biomedical Materials Research, 49 (1999) 43-52. [116] H. Zeng, W.R. Lacefield, XPS, EDX and FTIR analysis of pulsed laser deposited calcium phosphate bioceramic coatings: the effects of various process parameters, Biomaterials, 21 (2000) 23-30. [117] M. Ball, S. Downes, C. Scotchford, E. Antonov, V. Bagratashvili, V. Popov, et al., Osteoblast growth on titanium foils coated with hydroxyapatite by pulsed laser ablation, Biomaterials, 22 (2001) 337-47. [118] J.V. Rau, A. Generosi, S. Laureti, V.S. Komlev, D. Ferro, S.N. Cesaro, et al., Physicochemical investigation of pulsed laser deposited carbonated hydroxyapatite films on titanium, ACS Applied Materials & Interfaces, 1 (2009) 1813-20. [119] H. Man, K. Chiu, F. Cheng, K. Wong, Adhesion study of pulsed laser deposited hydroxyapatite coating on laser surface nitrided titanium, Thin solid films, 517 (2009) 5496-501. [120] G. Dinda, J. Shin, J. Mazumder, Pulsed laser deposition of hydroxyapatite thin films on Ti–6Al–4V: effect of heat treatment on structure and properties, Acta Biomaterialia, 5 (2009) 1821-30. [121] P. Rajesh, C. Muraleedharan, M. Komath, H. Varma, Pulsed laser deposition of hydroxyapatite on titanium substrate with titania interlayer, Journal of Materials Science: Materials in Medicine, 22 (2011) 497-505. [122] H. Kim, R.P. Camata, S. Chowdhury, Y.K. Vohra, In vitro dissolution and mechanical behavior of< i> c</i>-axis preferentially oriented hydroxyapatite thin films fabricated by pulsed laser deposition, Acta Biomaterialia, 6 (2010) 3234-41. [123] H. Pelletier, V. Nelea, P. Mille, D. Muller, Nano-scratch study of pulsed laser-deposited hydroxyapatite thin films implanted at high energy with N+ and AR+ ions, Journal of Materials Science, 39 (2004) 4185-92. [124] C. Cotell, Pulsed laser deposition and processing of biocompatible hydroxylapatite thin films, Applied surface science, 69 (1993) 140-8. [125] E. Park, R.A. Condrate, Graded coating of hydroxyapatite and titanium by atmospheric plasma spraying, Materials Letters, 40 (1999) 228-34. [126] F.J. Kummer, W.L. Jaffe, Stability of a cyclically loaded hydroxyapatite coating: effect of substrate material, surface preparation, and testing environment, Journal of Applied Biomaterials, 3 (1992) 211-5. [127] J. Ong, L. Lucas, W. Lacefield, E. Rigney, Structure, solubility and bond strength of thin calcium phosphate coatings produced by ion beam sputter deposition, Biomaterials, 13 (1992) 249-54.
47
[128] F. Garcia-Sanz, M. Mayor, J. Arias, J. Pou, B. Leon, M. Perez-Amor, Hydroxyapatite coatings: a comparative study between plasma-spray and pulsed laser deposition techniques, Journal of Materials Science: Materials in Medicine, 8 (1997) 861-5. [129] C. Wang, J. Lin, C. Ju, H. Ong, R. Chang, Structural characterization of pulsed laser-deposited hydroxyapatite film on titanium substrate, Biomaterials, 18 (1997) 1331-8. [130] M. Chen, X. Yang, Y. Liu, S. Zhu, Z. Cui, H. Man, Study on the formation of an apatite layer on NiTi shape memory alloy using a chemical treatment method, Surface and Coatings Technology, 173 (2003) 229-34. [131] A. International, A.S.f. Testing, Materials, Annual book of ASTM Standards, American Society for Testing & Materials, 2004. [132] P. Rajesh, C. Muraleedharan, M. Komath, H. Varma, Laser surface modification of titanium substrate for pulsed laser deposition of highly adherent hydroxyapatite, Journal of Materials Science: Materials in Medicine, 22 (2011) 1671-9. [133] I.O.f. Standardization, I.E. Commission, Information technology: security techniques: code of practice for information security management, ISO/IEC, 2005. [134] P. Sarkar, P.S. Nicholson, Electrophoretic deposition (EPD): mechanisms, kinetics, and application to ceramics, Journal of the American Ceramic Society, 79 (1996) 1987-2002. [135] H. Maiti, S. Datta, R. Basu, High�Tc Superconductor Coating on Metal Substrates by an Electrophoretic Technique, Journal of the American Ceramic Society, 72 (2005) 1733-5. [136] P. Sarkar, X. Huang, P.S. Nicholson, Zirconia/Alumina functionally gradiented composites by electrophoretic deposition techniques, Journal of the American Ceramic Society, 76 (2005) 1055-6. [137] R. Fischer, E. Fischer, G. De Portu, E. Roncari, Preparation of ceramic micro-laminate by electrophoresis in aqueous system, Journal of materials science letters, 14 (1995) 25-7. [138] M. Wei, A. Ruys, B. Milthorpe, C. Sorrell, J. Evans, Electrophoretic deposition of hydroxyapatite coatings on metal substrates: A nanoparticulate dual-coating approach, Journal of Sol-Gel Science and Technology, 21 (2001) 39-48. [139] J. Ma, C. Liang, L. Kong, C. Wang, Colloidal characterization and electrophoretic deposition of hydroxyapatite on titanium substrate, Journal of Materials Science: Materials in Medicine, 14 (2003) 797-801. [140] P. Ducheyne, S. Radin, M. Heughebaert, J. Heughebaert, Calcium phosphate ceramic coatings on porous titanium: effect of structure and composition on electrophoretic deposition, vacuum sintering and< i> in vitro</i> dissolution, Biomaterials, 11 (1990) 244-54. [141] K. Yamashita, M. Nagai, T. Umegaki, Fabrication of green films of single-and multi-component ceramic composites by electrophoretic deposition technique, Journal of Materials Science, 32 (1997) 6661-4. [142] I. Zhitomirsky, L. Gal-Or, Electrophoretic deposition of hydroxyapatite, Journal of Materials Science: Materials in Medicine, 8 (1997) 213-9. [143] P. Ducheyne, W. Van Raemdonck, J. Heughebaert, M. Heughebaert, Structural analysis of hydroxyapatite coatings on titanium, Biomaterials, 7 (1986) 97-103. [144] C. Kim, P. Ducheyne, Compositional variations in the surface and interface of calcium phosphate ceramic coatings on Ti and Ti-6Al-4V due to sintering and immersion, Biomaterials, 12 (1991) 461-9. [145] K. Yamashita, K. Kitagaki, T. Umegaki, Thermal instability and proton conductivity of ceramic hydroxyapatite at high temperatures, Journal of the American Ceramic Society, 78 (1995) 1191-7.
48
[146] A. Stoch, A. Brożek, G. Kmita, J. Stoch, W. Jastrzebski, A. Rakowska, Electrophoretic coating of hydroxyapatite on titanium implants, Journal of Molecular Structure, 596 (2001) 191-200. [147] R.K. Singh, F. Qian, V. Nagabushnam, R. Damodaran, B. Moudgil, Excimer laser deposition of hydroxyapatite thin films, Biomaterials, 15 (1994) 522-8. [148] M. Wei, A. Ruys, M. Swain, S. Kim, B. Milthorpe, C. Sorrell, Interfacial bond strength of electrophoretically deposited hydroxyapatite coatings on metals, Journal of Materials Science: Materials in Medicine, 10 (1999) 401-9. [149] X. Nie, A. Leyland, A. Matthews, Deposition of layered bioceramic hydroxyapatite/TiO< sub> 2</sub> coatings on titanium alloys using a hybrid technique of micro-arc oxidation and electrophoresis, Surface and Coatings Technology, 125 (2000) 407-14. [150] G. Soares, L.Á. de Sena, A.M. Rossi, M. Pinto, C.A. Muller, G.D. de Almeida Soares, Effect of electrophoretic apatite coating on osseointegration of titanium dental implants, Key Engineering Materials, 254 (2004) 729-32. [151] Z. Zhang, M.F. Dunn, T. Xiao, A.P. Tomsia, E. Saiz, Nanostructured hydroxyapatite coatings for improved adhesion and corrosion resistance for medical implants, MRS Proceedings, Cambridge Univ Press, 2001. [152] H. Gleiter, Materials with ultrafine microstructures: retrospective and perspectives, Nanostructured materials, 1 (1992) 1-19. [153] R. Siegel, J. Eastman, Synthesis, characterization, and properties of nanophase ceramics, MRS Proceedings, Cambridge Univ Press, 1988. [154] K. Yamashita, E. Yonehara, X. Ding, M. Nagai, T. Umegaki, M. Matsuda, Electrophoretic coating of multilayered apatite composite on alumina ceramics, Journal of Biomedical Materials Research, 43 (1998) 46-53. [155] G. Nahler, Food and Drug Administration (FDA), Dictionary of Pharmaceutical Medicine, (2009) 76-. [156] K. Yamashita, M. Matsuda, Y. Inda, T. Umegaki, M. Ito, T. Okura, Dielectric depression and dispersion in electrophoretically fabricated BaTiO3 Ceramic films, Journal of the American Ceramic Society, 80 (1997) 1907-9. [157] B. Ferrari, A. Sanchez-Herencia, R. Moreno, Aqueous electrophoretic deposition of AL< sub> 2</sub> O< sub> 3</sub>/ZrO< sub> 2</sub> layered ceramics, Materials Letters, 35 (1998) 370-4. [158] A. Ruys, M. Wei, C. Sorrell, M. Dickson, A. Brandwood, B. Milthorpe, Sintering effects on the strength of hydroxyapatite, Biomaterials, 16 (1995) 409-15. [159] M. We, A. Ruys, M. Swain, B. Milthorpe, C. Sorrell, Hydroxyapatite-coated metals: interfacial reactions during sintering, Journal of Materials Science: Materials in Medicine, 16 (2005) 101-6. [160] O. Albayrak, O. El-Atwani, S. Altintas, Hydroxyapatite coating on titanium substrate by electrophoretic deposition method: effects of titanium dioxide inner layer on adhesion strength and hydroxyapatite decomposition, Surface and Coatings Technology, 202 (2008) 2482-7. [161] R. Roop Kumar, M. Wang, Functionally graded bioactive coatings of hydroxyapatite/titanium oxide composite system, Materials Letters, 55 (2002) 133-7. [162] D. Uhlmann, T. Suratwala, K. Davidson, J. Boulton, G. Teowee, Sol—gel derived coatings on glass, Journal of non-crystalline solids, 218 (1997) 113-22. [163] M. Guglielmi, Sol-gel coatings on metals, Journal of Sol-Gel Science and Technology, 8 (1997) 443-9.
49
[164] L. Hu, T. Yoko, H. Kozuka, S. Sakka, Effects of solvent on properties of sol—gel-derived TiO< sub> 2</sub> coating films, Thin solid films, 219 (1992) 18-23. [165] K.i. Miyazawa, K. Suzuki, M.Y. Wey, Microstructure and Oxidation�Resistant Property of Sol�Gel�Derived ZrO2�Y2O3 Films Prepared on Austenitic Stainless Steel Substrates, Journal of the American Ceramic Society, 78 (1995) 347-55. [166] D.G. Young, M.A. Duran, Guide to the identification and geographic distribution of Lutzomyia sand flies in Mexico, the West Indies, Central and South America (Diptera: Psychodidae), DTIC Document, 1994. [167] P. Innocenzi, M. Guglielmi, M. Gobbin, P. Colombo, Coating of metals by the sol-gel dip-coating method, Journal of the European Ceramic Society, 10 (1992) 431-6. [168] C.J. Brinker, G.W. Scherer, Sol-gel science: the physics and chemistry of sol-gel processing, Academic Pr, 1990. [169] R.C. Mehrotra, Chemistry of alkoxide precursors, Journal of non-crystalline solids, 121 (1990) 1-6. [170] C.J. Brinker, C.S. Ashley, R.A. Cairncross, K.S. Chen, A.J. Hurd, S.T. Reed, et al., Sol—gel derived ceramic films—fundamentals and applications, Metallurgical and Ceramic Protective Coatings, Springer, 1996, pp. 112-51. [171] T. Troczynski, Q. Yang, Process for making chemically bonded sol-gel ceramics, Google Patents, 2001. [172] T. Olding, M. Sayer, D. Barrow, Ceramic sol–gel composite coatings for electrical insulation, Thin solid films, 398 (2001) 581-6. [173] H. Varma, S. Kalkura, R. Sivakumar, Polymeric precursor route for the preparation of calcium phosphate compounds, Ceramics international, 24 (1998) 467-70. [174] L.L. Hench, Bioceramics: from concept to clinic, Journal of the American Ceramic Society, 74 (1991) 1487-510. [175] C.S. Chai, K.A. Gross, B. Ben-Nissan, Critical ageing of hydroxyapatite sol–gel solutions, Biomaterials, 19 (1998) 2291-6. [176] A. Deptuła, W. Łada, T. Olczak, A. Borello, C. Alvani, A. Di Bartolomeo, Preparation of spherical powders of hydroxyapatite by sol-gel process, Journal of non-crystalline solids, 147 (1992) 537-41. [177] W. Van Raemdonck, P. Ducheyne, P. De Meester, Calcium phosphate ceramics, Metal and ceramic biomaterials, 2 (1984) 143-66. [178] K. Gross, C. Chai, G. Kannangara, B. Ben-Nissan, L. Hanley, Thin hydroxyapatite coatings via sol–gel synthesis, Journal of Materials Science: Materials in Medicine, 9 (1998) 839-43. [179] P. Layrolle, A. Ito, T. Tateishi, Sol�Gel Synthesis of Amorphous Calcium Phosphate and Sintering into Microporous Hydroxyapatite Bioceramics, Journal of the American Ceramic Society, 81 (2005) 1421-8. [180] M.-F. Hsieh, L.-H. Perng, T.-S. Chin, H.-G. Perng, Phase purity of sol–gel-derived hydroxyapatite ceramic, Biomaterials, 22 (2001) 2601-7. [181] L. Gineste, M. Gineste, X. Ranz, A. Ellefterion, A. Guilhem, N. Rouquet, et al., Degradation of hydroxylapatite, fluorapatite, and fluorhydroxyapatite coatings of dental implants in dogs, Journal of Biomedical Materials Research, 48 (1999) 224-34. [182] S. Overgaard, M. Lind, K. Josephsen, A.B. Maunsbach, C. Bünger, K. Søballe, Resorption of hydroxyapatite and fluorapatite ceramic coatings on weight�bearing implants: A quantitative and morphological study in dogs, Journal of Biomedical Materials Research, 39 (1998) 141-52.
50
[183] S. Zhang, Z. Xianting, W. Yongsheng, C. Kui, W. Wenjian, Adhesion strength of sol–gel derived fluoridated hydroxyapatite coatings, Surface and Coatings Technology, 200 (2006) 6350-4. [184] S. Barinov, S. Tumanov, I. Fadeeva, V.Y. Bibikov, Environment effect on the strength of hydroxy-and fluorohydroxyapatite ceramics, Inorganic materials, 39 (2003) 877-80. [185] S. ZHANG, X.L. BUI, Adhesion improvement of magnetron-sputtered amorphous carbon coating on cemented carbide, Adhesion Aspects of Thin Films, Volume 2~ autofilled~, 37. [186] S. Zhang, Y. Wang, X. Zeng, K. Khor, W. Weng, D. Sun, Evaluation of adhesion strength and toughness of fluoridated hydroxyapatite coatings, Thin solid films, 516 (2008) 5162-7. [187] H.-W. Kim, Y.-M. Kong, C.-J. Bae, Y.-J. Noh, H.-E. Kim, Sol–gel derived fluor-hydroxyapatite biocoatings on zirconia substrate, Biomaterials, 25 (2004) 2919-26. [188] S. Ding, C.-P. Ju, J.C. Lin, Morphology and immersion behavior of plasma-sprayed hydroxyapatite/bioactive glass coatings, Journal of Materials Science: Materials in Medicine, 11 (2000) 183-90. [189] Y. Gu, K. Khor, P. Cheang, In vitro studies of plasma-sprayed hydroxyapatite/Ti-6Al-4V composite coatings in simulated body fluid (SBF), Biomaterials, 24 (2003) 1603-11. [190] K. Cheng, C. Ren, W. Weng, P. Du, G. Shen, G. Han, et al., Bonding strength of fluoridated hydroxyapatite coatings: A comparative study on pull-out and scratch analysis, Thin solid films, 517 (2009) 5361-4. [191] Q. Zhang, Z. Long, C. Ren, B. Guo, D. Xu, T. Ma, Multi-beam mixing implantation system and its applications, Surface and Coatings Technology, 103 (1998) 195-9. [192] A. Ektessabi, Ion beam processing of bio-ceramics, Nuclear Instruments and Methods in Physics Research Section B: Beam Interactions with Materials and Atoms, 99 (1995) 610-3. [193] R. Bambauer, P. Mestres, R. Schiel, J. Schneidewind, R. Goudjinou, R. Latza, et al., Surface treated large bore catheters with silver based coatings versus untreated catheters for extracorporeal detoxification methods, ASAIO journal, 44 (1998) 303-8. [194] A. Ektessabi, Surface modification of biomedical implants using ion-beam-assisted sputter deposition, Nuclear Instruments and Methods in Physics Research Section B: Beam Interactions with Materials and Atoms, 127 (1997) 1008-14. [195] P. Sioshansi, E.J. Tobin, Surface treatment of biomaterials by ion beam processes, Surface and Coatings Technology, 83 (1996) 175-82. [196] F. Cui, Z. Luo, Biomaterials modification by ion-beam processing, Surface and Coatings Technology, 112 (1999) 278-85. [197] M. Hamdi, A. Ide-Ektessabi, Preparation of hydroxyapatite layer by ion beam assisted simultaneous vapor deposition, Surface and Coatings Technology, 163 (2003) 362-7. [198] R.G. Geesink, M.T. Manley, Hydroxylapatite coatings in orthopaedic surgery, Raven Press, 1993. [199] H. Aoki, Science and medical applications of hydroxyapatite, Ishiyaku Euroamerica, 1991. [200] H. Ji, P. Marquis, Effect of heat treatment on the microstructure of plasma-sprayed hydroxyapatite coating, Biomaterials, 14 (1993) 64-8. [201] J. Chen, J. Wolke, K. De Groot, Microstructure and crystallinity in hydroxyapatite coatings, Biomaterials, 15 (1994) 396-9. [202] F. Brossa, A. Cigada, R. Chiesa, L. Paracchini, C. Consonni, Post-deposition treatment effects on hydroxyapatite vacuum plasma spray coatings, Journal of Materials Science: Materials in Medicine, 5 (1994) 855-7.
51
[203] J. Chen, W. Tong, Y. Cao, J. Feng, X. Zhang, Effect of atmosphere on phase transformation in plasma�sprayed hydroxyapatite coatings during heat treatment, Journal of Biomedical Materials Research, 34 (1997) 15-20. [204] Y. Ohtsuka, M. Matsuura, N. Chida, M. Yoshinari, T. Sumii, T. Dérand, Formation of hydroxyapatite coating on pure titanium substrates by ion beam dynamic mixing, Surface and Coatings Technology, 65 (1994) 224-30. [205] R. Kant, B. Sartwell, Ion beam modification of TiN films during vapor deposition, Materials Science and Engineering, 90 (1987) 357-65. [206] F. Smidt, Use of ion beam assisted deposition to modify the microstructure and properties of thin films, International Materials Reviews, 35 (1990) 61-128. [207] F. Cui, Z. Luo, Q. Feng, Highly adhesive hydroxyapatite coatings on titanium alloy formed by ion beam assisted deposition, Journal of Materials Science: Materials in Medicine, 8 (1997) 403-5. [208] J.-M. Choi, Y.-M. Kong, K. Sona, H.-E. Kim, I. Lee, Formation and characterization of hydroxyapatite coating layer on Ti-based metal implant by electron-beam deposition, JOURNAL OF MATERIALS RESEARCH-PITTSBURGH-, 14 (1999) 2980-5. [209] J.W. Choi, Y.M. Kong, H.E. Kim, I.S. Lee, Reinforcement of hydroxyapatite bioceramic by addition of Ni3Al and Al2O3, Journal of the American Ceramic Society, 81 (1998) 1743-8. [210] M. Hamdi, A. Ektessabi, Influence of annealing temperature on simultaneous vapor deposited calcium phosphate thin films, Journal of Vacuum Science & Technology A: Vacuum, Surfaces, and Films, 19 (2001) 1566-70. [211] J. Liu, Z. Luo, F. Cui, X. Duan, L.M. Peng, High�resolution transmission electron microscopy investigations of a highly adhesive hydroxyapatite coating/titanium interface fabricated by ion�beam�assisted deposition, Journal of Biomedical Materials Research, 52 (2000) 115-8. [212] J.L. Ong, L.A. Harris, L.C. Lucas, W.R. Lacefield, D. Rigney, X�ray Photoelectron Spectroscopy Characterization of Ion�Beam Sputter�Deposited Calcium Phosphate Coatings, Journal of the American Ceramic Society, 74 (1991) 2301-4. [213] W. Lacefield, E. Rigney, L. Lucas, J. Ong, J. Gantenberg, Sputter deposition of Ca-P coatings onto metallic implants, (1990). [214] K. Ozeki, T. Yuhta, H. Aoki, I. Nishimura, Y. Fukui, Crystal chemistry of hydroxyapatite deposited on titanium by sputtering technique, Bio-Medical Materials and Engineering, 10 (2000) 221-7. [215] K. Ozeki, T. Yuhta, Y. Fukui, H. Aoki, Phase composition of sputtered films from a hydroxyapatite target, Surface and Coatings Technology, 160 (2002) 54-61. [216] K. Ozeki, T. Yuhta, H. Aoki, I. Nishimura, Y. Fukui, Push-out strength of hydroxyapatite coated by sputtering technique in bone, Bio-medical materials and engineering, 11 (2001) 63-8. [217] K. Ozeki, Y. Fukui, H. Aoki, Hydroxyapatite Coated Dental Implants by Sputtering, Biocybernetics and Biomedical Engineering, 26 (2006) 95-101. [218] K. De Groot, J. Wolke, J. Jansen, Calcium phosphate coatings for medical implants, Proceedings of the Institution of Mechanical Engineers, Part H: Journal of Engineering in Medicine, 212 (1998) 137-47. [219] S. Hasegawa, K. Furuya, Y. Ukegawa, M. Fushimi, H. Aoki, M. Akao, et al., Osteocompatibility of hydroxyapatite-coated titanium by thermal decomposition method, Proceedings of the First International Symposium on Apatite, Mishima, Japan, 1991, pp. 259-66.
52
[220] S.J. Ding, C.P. Ju, J.H.C. Lin, Characterization of hydroxyapatite and titanium coatings sputtered on Ti�6Al�4V substrate, Journal of biomedical materials research, 44 (1999) 266-79. [221] F. Brossa, A. Cigada, R. Chiesa, L. Paracchini, C. Consonni, Adhesion properties of plasma sprayed hydroxylapatite coatings for orthopaedic prostheses, Bio-medical materials and engineering, 3 (1993) 127-36. [222] J.C. Lin, M. Liu, C. Ju, Structure and properties of hydroxyapatite-bioactive glass composites plasma sprayed on Ti6Al4V, Journal of Materials Science: Materials in Medicine, 5 (1994) 279-83. [223] W. Lacefield, Hydroxyapatite coatings, Annals of the New York Academy of Sciences, 523 (1988) 72-80. [224] K. Koski, J. Hölsä, J. Ernoult, A. Rouzaud, The connection between sputter cleaning and adhesion of thin solid films, Surface and Coatings Technology, 80 (1996) 195-9. [225] L. Cleries, J. Fernandez-Pradas, J. Morenza, Behavior in simulated body fluid of calcium phosphate coatings obtained by laser ablation, Biomaterials, 21 (2000) 1861-5.
Table 1 Thermal spray condition of HA powders [92].
Parameters Argon
(l/min)
Helium
(l/min)
Current Voltage Powder
Rate
(g/min)
Spray
Distance
(mm)
Surface
Speed
(m/min)
Travers
Speed
(mm)
Cooling
Setting 41 60 700 52 30 115 75 8 yes
Table 2 Bond strength test results with different pretreatment and cryogenic treatment [92].
Coating Bonding Strength ( MPa)
Without Cryogenic
Treatment
With Cryogenic Treatment
Ultrasonic High
Pressure Air
Ultrasonic High
Pressure Air
HA 26.56 18.91 36.65 29.30
53
Table 3 Adhesion strengths of HA coated samples with and without TiO2 inner layer
deposited using different voltages [160].
Samples
(substrate + inner layer + Ha)
Shear Strength
(MPa)
Ti-6Al-4V + ----- + HA 13.8 ( s=1.8)
Ti-6Al-4V + TiO2 (50 V)+ HA 11.9 ( s=1.8)
Ti-6Al-4V + TiO2 (20 V)+ HA 13.1 ( s=1.8)
Ti-6Al-4V + TiO2 (10 V)+ HA 21.0 ( s=1.8)
Note. S: standard deviation.
Table 4 Adhesion Strength and Failure Mode of Coatings [220].
Coating Code
Adhesion Strength (MPa) Failure Mode (Ra= 0.06 µm) (Ra= 0.06 µm) (Ra= 0.7 µm)
HA 59.9 ±12.4 (41) 71.8±14.7 (25) R/C, C/S
95HA/5Ti 59.5 ± 6.5 (20) 60.7±5.8 (23) R/C, C/S
90HA/10Ti 58.4 ± 6.2 (18) 54.5±6.1 (12) R/C, C/S
85HA/15Ti 64.8 ± 6.2 (17) 69.5 ± 10.3 (19) R/C, C/S
75HA/25Ti 64.0 ± 6.9 (32) 65.3 ± 6.5 (50) R/C, C/S
50HA/50Ti 75.1 ± 5.5 (22) 72.9 ± 5.4 (28) R/C
54
25HA/75Ti 81.9 ± 5.2 (19) 79.8 ± 6.3 (17) R/C
Ti 79.5 ± 9.1 (15) 85.1 ± 5.1 (34) R/C
Table 5: Different techniques to deposit HA coating.
Technique Thickness Advantages Disadvantages
Plasma Spraying < 20 µm rapid deposition ; sufficiently low cost; fast bone healing, less risk for coating degradation
Poor adhesion, alternation of HA structure due to coating process; non- uniformity in coating density; extreme high temperature up to 1200 ºc, phase transformation and grain grow of substance due to high temperature procedure; increase in residual stress; unable to produce complete crystalline HA coating
Thermal Spraying
30- 200 µm High deposition rates; low cost;
Line of sight technique; high temperatures induce decomposition; rapid cooling produces amorphous coatings; lack of uniformity; crack appearance; low porosity; coating spalling and interface separation between the coating and the substrate
Sputter Coating 0.5- 3 µm Uniform coating thickness on flat substrates; dense coating; homogenous coating; high adhesion
Line of sight technique; expensive time consuming; produces amorphous coatings; low crystallite which accelerates the dissolution of the film in the body
Pulsed Laser Deposition
0.05- 5 mm Coating with crystalline and amorphous; coating with dense and porous; ability to produce wide range of multilayer coating from different materials; ability to produce high crystalline HA coating; ability to restore complex stoichiometry; high degree of control on deposition parameters
Line of sight technique; splashing or particle deposition; need surface pretreatment; lack of uniformity
Dip Coating < 1 µm Inexpensive; coatings applied quickly; can coat complex substrates; high surface uniformity; good speed of coating;
Requires high sintering temperatures; thermal expansion mismatch; crack appearance
Sol-gel 0.1- 2.0 µm Can coat complex shapes; Low processing temperatures; relatively cheap as coatings are very thin; simple deposition method; high purity; high corrosion resistant; fairly good adhesion
Some processes require controlled atmosphere processing; expensive raw materials; not suitable for industrial scale; high permeability; low wear resistance; hard to control the porosity;
55
Electrophoretic Deposition
0.1- 2.0 mm Uniform coating thickness; rapid deposition rates; can coat complex substrates; simple setup, low cost, high degree of control on coating morphology and thickness, good mechanical strength; high adhesion for n-HA
Difficult to produce crack-free coatings; requires high sintering temperatures; HA decomposition during sintering stage
Hot Isostatic Pressing
0.2- 2.0 mm Produces dense coatings; produce net-shape ceramics; good temperature control; homogeneous structure; high uniformity; high precision; no dimensional or shape limitation
Cannot coat complex substrates; high temperature required; thermal expansion mismatch; elastic property differences; expensive; removal/interaction of encapsulation material
Ion Beam Assisted Deposition
<0.03 µm Low temperature process; high reproducibility and reliability; high adhesion; wide atomic intermix zone are coating-to-substrate interface
Crack appearance on the coated surface
Figure 1 Tensile bond strength result of plasma sprayed Ti-6Al-4V/ 20 wt.%
hydroxyapatite coating (as sprayed and HIPed) [79].
56
Figure 2 A schematic diagram of thermal spray coating [82].
Figure 3 Fundamental stages of dip coating (the finer arrows indicate
the flow of air) [94].
57
Figure 4 SEM micrographs from cross-sectional view of HA coatings (via SOL 2) on Ti-
6Al-4Vsubstrates after heating at 840°C [96].
58
Figure 5 Comparison of adhesion strength for HA on substrates with different pre-treatments
[119].
Figure 6 Average surface roughness of titanium substrates treated with different laser pulses and
HA coating compared with control sample [132].
59
Figure 7 Failure values obtained by scratch test (Lc1, Lc2 and Lc3) for the HA coatings on
different irradiated and non-irradiated titanium substrate [132].
Figure 8 Electro-polarization corrosion curves for both EPD n-HA coating and HA thermal
sprayed coating [151].
60
Figure 9 Cross section SEM micrograph of the EPD deposited under the identified optimum
suspension condition [140].
Figure 10 SEM micrograph of the uncrack deposit surface [140].
100µm
61
Figure 11 Steps in the sol-gel process for ceramic materials [169].
Figure 12 X-ray diffraction of sol-gel coatings preferred to 500˚C on titanium substrates and
then fired at various temperatures [179].
62
Figure 13 A scanning electron micrograph of a coating fired to 800˚C for 10 min, the field of
view is 250 nm × by 250 nm [179].
Figure 14 Coefficient of friction in terms of relative voltage as a function of normal load while
scratching (a) pure HA coating; (b) fluoridate HA (FHA6) coating on Ti-6Al-4V [174].
63
Figure 15 Adhesion strength of pure HA and fluoridated HA coatings on Ti-6Al-4V substrates
as indicated by upper critical load in scratch test. Firing temperatures are indicated [174].
Figure 16 Pull-out adhesion strength of FHA coating before and after soaking in TPS solutions.
* indicates a significant increase of adhesion strength with respect to F0 (as prepared coatings);
64
** indicate a significant increase of adhesion strength with respect to F0 (after soaking in TPS
for 21 days) [186].
Figure 17 Pull-out strength of coatings with different F content [190].
65
Figure 18 Fz-Fy curve of scratch test from specimen prepared by (a) IBSD and (b) IBAD [207].
Figure 19 Layer-metal substrate bond strengths, before and after heat treatment, as a function of
ion beam current [207].
66
Figure 20 SEM micrographs of the coating layer (A) before and (B) after the heat treatment
[208].
67
Figure 21 Adhesion strength of the IBASVD samples at different elevated temperatures [197].
68
Figure 22 Bone bonding strengths of sputtered films [217].
69
Figure 14 Adhesion strength of coatings [220].
70
Figure 24 Quantative comparison of different coating techniques.
Plas
ma
Spra
y
PLD
(100
0 Pu
lsed
Las
er)
PLD
(Nitr
ided
/ Etc
hing
) Pre
trea
tmen
t EPD
(n-H
A)
EPD
(With
out T
iO2
Inte
rfac
ial L
ayer
)
EPD
(TiO
2 In
terf
acia
l Lay
er)
Sol-G
el So
l-Gel
(FH
A) IB
AD
IBA
D (H
eat T
reat
men
t)
The
rmal
Spr
ay
The
rmal
Spr
ay (w
ith tr
eatm
ent)
Dip
Coa
ting
Hot
Isos
tatic
Pre
ssin
g (a
fter
Pla
sma
Spra
y)
Sput
teri
ng