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Please cite thisnanotube arrayj.jmbbm.2013.01.

Research Paper

Fabrication and deformation behaviour of multilayer Al2O3/Ti/TiO2 nanotube arrays

S. Baradarana,n, W.J. Basirunb,d, E. Zalnezhada, M. Hamdic, Ahmed A.D. Sarhanc, Y. Aliasb

aDepartment of Engineering Design and Manufacture, Faculty of Engineering University of Malaya, 50603 Kuala Lumpur, MalaysiabDepartment of Chemistry, Faculty of Science, University of Malaya, Kuala Lumpur 50603, MalaysiacCenter of Advanced Manufacturing and Material Processing, University of Malaya, 50603 Kuala Lumpur, MalaysiadNanotechnology & Catalysis Research Centre (NanoCat), Institute of Postgraduate Studies, University Malaya, 50603 Kuala Lumpur,

Malaysia

a r t i c l e i n f o

Article history:

Received 24 September 2012

Received in revised form

24 January 2013

Accepted 26 January 2013

Keywords:

Magnetron sputtering

Adhesion

TiO2 nanotube

Anodizing

Nanoindentation

nt matter & 2013 Elsevie.1016/j.jmbbm.2013.01.02

hor. Tel.: þ601 7283 8175;saeid_baradaran@yahoo

article as: Baradarans. Journal of the M020

a b s t r a c t

In this study, titanium thin films were deposited on alumina substrates by radio frequency

(RF) magnetron sputtering. The mechanical properties of the Ti coatings were evaluated in

terms of adhesion strength at various RF powers, temperatures, and substrate bias

voltages. The coating conditions of 400 W of RF power, 250 1C, and a 75 V substrate bias

voltage produced the strongest coating adhesion, as obtained by the Taguchi optimisation

method. TiO2 nanotube arrays were grown as a second layer on the Ti substrates using

electrochemical anodisation at a constant potential of 20 V and anodisation times of

15 min, 45 min, and 75 min in a NH4F electrolyte solution (75 ethylene glycol: 25 water). The

anodised titanium was annealed at 450 1C and 650 1C in a N2 gas furnace to obtain different

phases of titania, anatase and rutile, respectively. The mechanical properties of the

anodised layer were investigated by nanoindentation. The results indicate that Young’s

modulus and hardness increased with annealing temperature to 650 1C.

& 2013 Elsevier Ltd. All rights reserved.

1. Introduction

Bioceramics have recently become one of the most important

biomaterials used for implants. They are the type of material

most compatible with the human body due to chemical

similarities that facilitate their direct bonding to bone.

According to their level of interaction with living tissue,

bioceramics can be divided into inert and active types.

Alumina (Al2O3) is a type of inert bioceramic utilised in

orthopaedic implantation because of its good mechanical

properties (high strength, high fracture toughness) and good

compatibility. In addition, it is employed in the fabrication of

r Ltd. All rights reserved.0

fax: þ603 7967 5330..com (S. Baradaran).

, S., et al., Fabricationechanical Behavior o

bone plates, screws, and femoral heads and widely applied in

total hip joint and knee replacement (Liu, 2007; Velmurugan

et al., 2010; Youn et al., 2011).

Over the last few decades, implant coating has found a

wide array of applications. The thin-film coating of implant

surfaces can be performed by various methods, including

plasma spraying, electrophoresis, dipping, electrochemical

deposition, pulsed laser deposition, ion beam dynamic mix-

ing, and ion beam deposition (Oh et al., 2005; Raja et al., 2005;

Kar et al., 2006; Best et al., 2008; Wang et al., 2008; Kodama

et al., 2009; Roy et al., 2010; Wang and Luo, 2011). Some of

these methods have severe limitations such as poor

and deformation behaviour of multilayer Al2O3/Ti/TiO2

f Biomedical Materials (2013), http://dx.doi.org/10.1016/

Table 1 – Factors and levels used in the experiment.

Exp. Factors

RF power

(W)

Temperature

(1C)

Substrate biases

voltage (V)

1 300 150 25

2 300 200 50

3 300 250 75

4 400 150 50

5 400 200 75

6 400 250 25

7 500 150 75

8 500 200 25

9 500 250 50

j o u r n a l o f t h e m e c h a n i c a l b e h a v i o r o f b i o m e d i c a l m a t e r i a l s ] ( ] ] ] ] ) ] ] ] – ] ] ]2

adhesion, micro-crack formation, phase changes at high

temperatures, non-uniformity, and improper microstructural

control, all of which make them inadequate for implant

systems (Raja et al., 2005; Kar et al., 2006). More recently,

physical vapour deposition (PVD) magnetron sputtering has

been suggested by many researchers as a versatile deposition

technique that offers many advantages including high

deposition rates, ease of sputtering any metal, alloy or

compound, the formation of high-purity films, extremely

high adhesion to films, and the ability to form dense coatings

(Swann, 1988; Ding et al., 1999; Kelly and Arnell, 2000; Nelea

et al., 2003; Shi et al., 2008; Toque et al., 2010).

Throughout the last decades, pure titanium has been the

most widely applied biomaterial among metals due to its

fine biocompatibility, favourable mechanical properties, high

corrosion resistance, and adequately strong load-bearing

applications. Moreover, research on the nanostructured

morphology of titanium is becoming increasingly important

(Santos et al., 2008). The properties of nanostructured tita-

nium are mainly dependent on the material’s morphology

(Li et al., 2010). Nanostructured titanium layers can be

fabricated to have different surface areas and uniform size

distributions tailored for specific implant applications (Oh

et al., 2006; Zhu et al., 2007; Macak et al., 2008). Titanium

oxide (TiO2) nanotubes can be fabricated by various methods

such as hydrothermal deposition, electrodeposition, and

anodisation (Li et al., 2010; Zhang et al., 2010). Anodisation

is the most preferred process among these methods due to

the strong surface adhesion and mechanical integrity of the

resulting implants.

Moreover, to improve the strength of adhesion to sub-

strates, the coating parameter conditions should be opti-

mised (Camargo and Voorwald, 2007; Rajasekaran et al.,

2008). The conventional method is to use the ‘‘trial and error’’

approach, which is actually very time-consuming due to the

large number of experiments required. Therefore, a reliable

systematic approach for optimising coating parameters is

required (Ghani et al., 2004). The Taguchi optimisation

method is an efficient and effective approach through which

the response parameters affecting adhesion strength can be

optimised to identify the most significant response variables

with the minimum number of experiments (Farooq and Lee,

2002).

In addition, one of the most important attributes of

biomaterials used in medical and dental applications is their

mechanical properties. The consideration of such properties

becomes particularly important when these materials are

applied in bone implants, for which long-term in vivo struc-

tural stability is crucial. Typically, bone resorption occurs

when there is no compatibility between the mechanical

properties, such as the elastic moduli, of an implant material

and bone, leading to the loss of the implant and ultimately

failure (Huiskes et al., 1992). Therefore, characterising the

mechanical behaviour of implant materials is important. In

recent years, the mechanical characterisation of thin coat-

ings has been challenging because of the small length scales

involved. In this regard, the nanoindentation technique is

appropriate for testing biomaterials because of its low load

(mN) and small displacement (nm) resolution. Furthermore,

due to the small probe it employs, nanoindentation is an

Please cite this article as: Baradaran, S., et al., Fabricationnanotube arrays. Journal of the Mechanical Behavior oj.jmbbm.2013.01.020

effective technique for measuring the mechanical properties

of thin (10 mm) coated systems, including soft coatings on soft

substrates, hard coatings on soft substrates and multilayered

systems (Bahr et al., 2003; Deng et al., 2005a, 2005b; Ebenstein

and Pruitt, 2006).

To our knowledge, there is a lack of information regarding

the coating of metals on ceramics using customary methods,

which are mainly used to coat ceramics on metals and other

ceramics; thus, a comprehensive study in this line of research

is needed. In this work, alumina substrates were coated with

pure titanium metal by PVD magnetron sputtering. An

optimisation study on the conditions of magnetron sputter

coating (RF power, temperature and DC bias voltage) was

carried out using the Taguchi method to obtain the appro-

priate adhesion strength. TiO2 nanotubes were grown by

anodisation with ammonium fluoride (NH4F) salt dissolved

in a water and ethylene glycol (EG) solvent mixture at 20 V for

different time periods. FESEM and XRD were used to probe

the titanium nanotube coating structure. The adhesion

strength between the ceramic (alumina) substrates and Ti

coatings was evaluated using a micro scratch tester under

different conditions. The hardness and elastic modulus of

each coating were measured by performing a nanoindenta-

tion test on the nanotube layer at various annealing tem-

peratures and indentation depth.

2. Materials and experimental procedure

2.1. Experimental design

The most important stage in experimental design using the

Taguchi approach lies in the selection of the control para-

meters and identifying the orthogonal array (OA) (Ghani

et al., 2004). This experiment comprised three parameters

with three levels each; thus, the fractional factors design

used was a standard L9 (33) orthogonal array, which was

chosen due to its capability to verify the interactions among

parameters. The nine experiments performed and the details

regarding all combinations of the experimental conditions for

each control parameter are shown in Table 1.

and deformation behaviour of multilayer Al2O3/Ti/TiO2

f Biomedical Materials (2013), http://dx.doi.org/10.1016/

Table 3 – The (S/N) response values for adhesionstrength.

Coating parameters S/N response

Level 1 Level 2 Level 3

RF power 54.78 59.26 57.65

Temperature 55.98 57.15 58.56

DC bias voltage 55.64 56.74 59.32

j o u r n a l o f t h e m e c h a n i c a l b e h a v i o r o f b i o m e d i c a l m a t e r i a l s ] ( ] ] ] ] ) ] ] ] – ] ] ] 3

2.2. Preparation of alumina substrate

Alumina plates (AdValue Technology, USA) measuring

60 mm�10 mm�5 mm were cut to a length of 15 mm and

used as the titanium coating substrates. The ceramic sur-

faces were polished using SiC emery paper (120–1200 grit),

followed by wet polishing in a diamond slurry. After polish-

ing, the alumina substrates were sonicated for approximately

15 min in ethanol and acetone before being rinsed with

distilled water and dried at 100 1C for 30 min.

2.3. Preparation of Ti coating

Magnetron sputtering (SG control engineering Pte Ltd series)

was used to deposit titanium thin films on the alumina

substrates. A pure titanium target (99.995% purity) was used

in this investigation. The distance between the target and

substrate was fixed at 150 mm. The alumina substrates were

cleaned with acetone and ethanol in an ultrasonic bath prior

to deposition. The target was pre-sputtered in an argon

atmosphere to remove the oxide layer. The chamber was

evacuated to below 2�10�5 Torr before the argon gas used for

sputtering was introduced. The working pressure was kept at

5.2�10�3 Torr, the argon gas flow rate was fixed at 20 sccm

and the deposition time was 7 h.

2.3.1. Data analysis for high adhesion strengthFollowing the experimental runs, the data obtained was

analysed to optimise the deposition parameters and identify

which process parameters are statistically significant. Data

analysis was conducted using signal-to-noise (S/N) response

analysis. The methods used to calculate the S/N ratio are

classified into three main categories, depending on whether

small, large, or nominal values are required to obtain the best

results. In the case of the adhesion strength between the

coatings and substrates, larger values were required. The

following equation is used to calculate the S/N ratio (in dB)

under conditions in which large ratios yield the best results:

S=N¼�10 log1n

S1y2

� �ð1Þ

where y represents the observed data and n is the number of

observations. The S/N ratio functions as a performance

Table 2 – The measured titanium coating thickness, surface h

Exp Ti coating thickness (nm) Scratch force (mN)

First Seco

1 876 385 437

2 870 565 583

3 902 696 697

4 1102 637 656

5 1401 1130 1193

6 1011 1012 993

7 1307 933 951

8 1024 520 555

9 1119 870 848

Please cite this article as: Baradaran, S., et al., Fabricationnanotube arrays. Journal of the Mechanical Behavior oj.jmbbm.2013.01.020

measurement that can be used to develop processes insensi-

tive to noise factors. The degree of predictable performance

of a product or process in the presence of noise factors can be

defined from the corresponding S/N ratios. Table 2 shows the

S/N ratio calculated for adhesion strength, while Table 3

provides the S/N response values for the measured data,

which are plotted in Fig. 1. The desired ‘‘the higher the better’’

criteria implies that the highest S/N reflects the best

response, which results in the lowest effects due to noise

on the machine setup. This criterion was employed in this

study to determine the optimal coating parameters yielding

the highest adhesion strength. As observed in Fig. 1 and

according to the higher (S/N) response base, the RF power,

temperature, and DC bias voltage were significant in deter-

mining the best adhesion strength. An RF power of 400 W,

temperature of 250 1C, and DC bias voltage (75 V) appear to be

the best conditions for obtaining high adhesion strength. At

these levels, a confirmation test was carried out using the

best parameter combination to validate the findings. The

adhesion strength obtained from this confirmation test was

1238 mN, representing an improvement of 4.47% compared

with the highest adhesion strength value obtained from the

experiments shown in Table 2.

2.4. Preparation of TiO2 nanotubes

The Ti-coated samples with the highest adhesion strength

were anodised in an ammonium fluoride (NH4F, Sigma-

Aldrich CO., 0.5 wt%) electrolyte dissolved in a 75:25 ethylene

glycol (EG, J.T. Baker CO.) and water solvent mixture at room

temperature. A direct current (DC) power source (Model

E3641A, Agilent Technologies, Palo Alto, USA) was used to

ardness and calculated (S/N) ratio.

Calculated S/N ratio

nd Third Ave

391 404 52.13

601 583 55.31

703 699 56.89

647 647 56.22

1233 1185 61.47

1030 1012 60.10

977 954 59.59

552 542 54.68

860 859 58.68

and deformation behaviour of multilayer Al2O3/Ti/TiO2

f Biomedical Materials (2013), http://dx.doi.org/10.1016/

52

53

54

55

56

57

58

59

60

300 400 500 150 200 250 25 50 75

S/N

resp

onse

val

ues

Coating parameters level

RF power (W)

Temperature (°C)

DC bias voltage (V)

Fig. 1 – The S/N response graph for adhesion strength.

j o u r n a l o f t h e m e c h a n i c a l b e h a v i o r o f b i o m e d i c a l m a t e r i a l s ] ( ] ] ] ] ) ] ] ] – ] ] ]4

perform the anodisation. Each coated sample was connected

to the positive terminal (anode), and a graphite rod (D¼5 mm)

was connected to the negative terminal (cathode) of the

power source. The distance between the cathode and anode

was fixed at roughly 20 mm. The anodisation of the Ti surface

was carried out at a constant potential of 20 V for different

anodisation times, namely 15 min, 45 min, and 75 min. Dur-

ing anodisation, the surface colour changed from purple to

blue to light green and finally to light yellow. Thermal

annealing was performed in an argon gas furnace at 450 1C

and 650 1C for 2 h at heating and cooling rates of 5 1C/min.

2.4.1. TiO2 growth mechanismThe growth of TiO2 nanotube arrays is considered to occur in two

steps: (i) chemical dissolution, followed by (ii) electrochemical

oxidation (Chang et al., 2011). The structure and geometric size

of Ti nanotubes are controlled by the chemical dissolution step.

The electrochemical oxidation stage can be adjusted by three

critical factors: (i) anodisation voltage, (ii) anodisation time, and

(iii) acid concentration. It is notable that the formation of TiO2

nanotubes depends mainly on the ammonium fluoride concen-

tration in the electrolyte (Crawford et al., 2007). In this study, the

titanium coating anodes were dissolved by water electrolysis.

Titanium ions migrated through the oxide layer by electric-field-

assisted processes and accumulated at the entrances of nano-

pores in the coatings. F�, O2� and OH� ions from the electrolyte

migrated through the oxide layer to the Ti metal/oxide interface

and react with the Ti metal. Pitting corrosion due to F� led to the

initial formation of bowl-like pores on the substrate surfaces. At

the same time, Ti4þ ions moved toward the electrolyte and

Ti(OH)4 precipitates formed via the instantaneous hydrolysis

reaction at the entrance of the nanopores. The chemical reac-

tions governing these processes are as follows:

(i)

Plenaj.jm

Ti-Ti4þþ4e�

(ii)

Ti4þþ4H2O-Ti(OH)4þ4Hþ

(iii)

Ti(OH)4-TiO2þ2H2O

Reaction (iv) is the total reaction that results in the TiO2

product:

(iv)

Tiþ2H2O-TiO2þ2H2

The reaction between F� and TiO2 forms the TiF62�

complex, which can dissolve into the electrolyte. The

oxidation rate is limited by the barrier layer’s thickness at

ase cite this article as: Baradaran, S., et al., Fabrication annotube arrays. Journal of the Mechanical Behavior of Bi

bbm.2013.01.020

the Ti metal/oxide interface and is kept active by the

chemical dissolution process at the oxide/electrolyte

interface. The corresponding reaction is as follows:

(v)

TiO2þ4Hþþ6F�-TiF62�þ2H2O

2.5. Characterisation

The coating morphology was observed using a high-

resolution FEI Quanta 200 F field emission scanning electron

microscope (FESEM). The statistical analysis of each tube

dimension (i.e., inner diameter, outer diameter, tube length)

was performed using the SPSS statistical software package

version 19 (SPSS Institute, Chicago, IL), with the assistance of

image analysis software before annealing. The source of

mentioned variations in Fig. 5 (f–h) is within the limit for

each sample. The measurements were directly extracted

from the FESEM images. Three samples were measured for

each processing condition. The adhesion strength was mea-

sured quantitatively using a Micro Materials NanoTest (Wrex-

ham, U.K.) with a diamond indenter; the sliding velocity was

5 mm/s, and the load was gradually increased to 9.2 mN/s.

Three scratches were performed for each sample. An Al2O3/

Ti/TiO2 cross-section was prepared by a relatively destructive

method using a diamond cutting tool machine. Sample X-ray

diffraction (XRD) patterns (as prepared and annealed sam-

ples) were recorded on an Empyrean X-ray diffractometer

with Cu Ka radiation (l¼1.54178 A) operating at 45 kV and

30 mA with a step size of 0.0261, scanning rate of 0.11/s and 2yrange from 101 to 801. The coatings’ mechanical properties

were investigated by nanoindentation using a commercial

nanoindenter (Micro materials Ltd. Wrexham, U.K.) with a

continuous stiffness measurement capability. The nanoin-

dentation experiments were performed using a Berkovich

diamond tip with a radius of 20 nm and constant depths of

550 nm and 800 nm. Calibration for load and hardness was

performed on a reference sample of fused silica. The inden-

tation velocity was 3–6 nm/s; once the maximum prescribed

depth was reached, loading was stopped and the load was

held constant for 10 s. The source of variations gathered from

five test sites is within the limit for each sample. Indentation

experiments were conducted for three samples per proces-

sing condition.

3. Results and discussion

3.1. Adhesion strength

The most essential properties of implant tissue materials are

their stability and durability because of the long-term pre-

sence required in the living body (Tomsia et al., 2005; Zhang

et al., 2007). Regarding applications, it is important to char-

acterise and thereby improve the adhesion strength between

coating layers and implant materials. The adhesion of a

coating to a substrate and the integrity of the substrate/

coating interface are always connected to the performance

and reliability of any coated device. In this study, a total of

nine experiments were carried out to achieve high adhesion

strength and consistency. Table 2 presents a summary of the

titanium coating thickness and adhesion test results at

d deformation behaviour of multilayer Al2O3/Ti/TiO2

omedical Materials (2013), http://dx.doi.org/10.1016/

j o u r n a l o f t h e m e c h a n i c a l b e h a v i o r o f b i o m e d i c a l m a t e r i a l s ] ( ] ] ] ] ) ] ] ] – ] ] ] 5

different parameters. As observed, Coating’s thickness is

increased with increasing RF power and substrate bias vol-

tage from 300 W to 400 W and 25 V to 75 V, respectively

(Fig. 1). When the RF power increased up to 400 W, the

sputtering rate increased because the ionised and sputtered

particles became more energetic, reducing the distance

between the energised atoms and subsequently making the

surface more adhesive. On the other hand when it increases

more to 500 W, collisions of sputtered particles with chamber

particles (argon gas and ions) increased. Hence, a drastic fall

in the sputtering rate for high power sputtering occurs.

Furthermore, with increasing substrate bias voltage, the

coating became more compact and denser which this event

is more effective in improving the adhesion of the coatings.

This improvement can be attributed to the additional energy

available for film growth. In this case, high-energy atoms

with greater mobility settled into lower-energy surface sites,

leading to higher coating adhesion and density (Gangopadhyay

et al., 2010; Singh et al., 2012). Finally, the combination of an RF

power of 400 W, temperature of 250 1C and substrate bias voltage

of 75 V provided suitable conditions for the strong adhesion of Ti

to alumina.

Fig. 2(a) and (b) illustrates the scratch length and failure

point as well as a graph of depth and load versus distance

during the scratch force test. It can be observed from the

graphs that coating (a) has the best adhesion strength. This

coated specimen had a scratch length of 1090.36 mm and a

failure point and load of approximately 995.33 mm and

1238 mN, respectively. Moreover, the failure point and scratch

length of the specimen exhibiting the lowest adhesion

strength were 388.27 mm and 1102.63 mN, respectively, while

Fig. 2 – Scratch force (adhesion) test results for a coated samp

critical load: (a) 400 W RF power, 250 1C and 75 V bias voltage; (

Please cite this article as: Baradaran, S., et al., Fabricationnanotube arrays. Journal of the Mechanical Behavior oj.jmbbm.2013.01.020

the graph of load versus distance shows a critical load of

437 mN.

Moreover, the critical loads were defined based on the

scratch profile of the load- displacement graph. The scratch

direction was from left to right for all of the scratch images

obtained in this study. Three critical loads Lc1, Lc2, and Lc3

were defined for the failure of the coatings. Lc1 is the first

critical load, corresponding to cohesive failure characterised

by trackside cracking. Lc2 is the second critical load, corre-

sponding to a mixture of adhesive and cohesive failure

characterised by delamination on the trackside and mild

tensile cracking. Lc3 is the load at which the coatings are

delaminated along the scratch path. Fig. 3(a and b) indicates

that several types of failure mode were observed during the

scratching of the titanium coatings on the alumina sub-

strates. The as-deposited coating began to fail when it

showed some cracking on the track side; the load at which

this was observed was recorded as Lc1. This was followed by

mild tensile cracking along the scratch path. As the load was

gradually increased, delamination occurred on the trackside,

and the load was recorded as Lc2. Prior to this mode of failure,

no coating delamination was observed along the scratch

path, but some cracking of the coating’s trackside was

evident. As the load continued to increase, the trackside

cracking and delamination became more severe and the

spacing between tensile cracks became smaller. Finally, the

coating was delaminated from the substrate along the

scratch path; the corresponding load was recorded as Lc3.

As this continued, the spallation on the trackside replaced

the trackside delamination and the scratch width became

le and force and depth versus depth graphs indicating the

b) 300 W RF power, 150 1C and 25 V bias voltage.

and deformation behaviour of multilayer Al2O3/Ti/TiO2

f Biomedical Materials (2013), http://dx.doi.org/10.1016/

Fig. 3 – Failure modes observed during scratch testing of as-deposited titanium coating at different stages.

Fig. 4 – Cross-section of sample (a) alumina, (b) titanium and

(c) titanium dioxide.

j o u r n a l o f t h e m e c h a n i c a l b e h a v i o r o f b i o m e d i c a l m a t e r i a l s ] ( ] ] ] ] ) ] ] ] – ] ] ]6

constant. This phenomenon is illustrated in Fig. 3b (Beake

and Lau, 2005; Toque et al., 2010; Singh et al., 2012).

3.2. Microstructure characterisation (Ti coating and TiO2

nanotube arrays)

Fig. 4(a–c) illustrates the cross-section of a sample prepared

by RF magnetron sputtering (PVD) and the anodised surface

on which TiO2 nanotube arrays were grown as a second layer.

Clearly, the thickness of the titanium coating is approxi-

mately 1.4 mm, where the first and second layers measure

approximately 620 nm and 745 nm in thickness, respectively

(Fig. 4(b and c)). A matte surface was observed on the bottom

of the TiO2 coating, as shown in Fig. 4b. The concave dimples

formed at the interface caused an increase in mechanical

interlocking and thereby enhanced the adhesion strength at

the interface (Crawford et al., 2007).

Fig. 5(a–c) shows the samples’ morphology at various

anodisation times. As shown in Fig. 4a, during the first stage

of anodisation, pore nucleation occurs on the Ti surface due

to the pitting corrosion induced by fluoride ions at an

anodisation time of 15 min. In Fig. 5a, it is evident that

nanopore formation occurred on top of the nanotubes, which

were non-homogenous and showed poor organisation on the

surface with no remnant TiO2 layer during the first 15 min. At

an anodisation time of 45 min (Fig. 5b), the porous structure

is converted to the tubular structure of the TiO2 nanotubes.

Fig. 5(c–e) shows the top, bottom view and cross-section of

the fractured TiO2 coating at 75 min of anodisation. The

bottom of the TiO2 nanotube coating shows a series of evenly

spaced ‘‘bumps’’ (Fig. 5d) that represent the pore tips of each

individual nanotube. It is clear that the nanotubes are well

defined with an average inner diameter, outer diameter and

Please cite this article as: Baradaran, S., et al., Fabricationnanotube arrays. Journal of the Mechanical Behavior oj.jmbbm.2013.01.020

tube length of approximately 58.3711.2 nm, 64.878.1 nm,

and 620.4734.7 nm, respectively (Fig.5(f–h)). The cross-

section of the TiO2 nanotubes shows a bamboo-shaped

structure with good density, which could result in enhanced

chemical activity and stronger interactions. It is noteworthy

that the coating density is an important factor contributing to

good mechanical properties, as will be discussed in the next

section (Zhang et al., 2007).

3.3. Crystal properties

Fig. 6 shows the diffuse diffraction pattern obtained by thin-

film X-ray diffraction analysis, where the oxide layer of

nanotube arrays formed by electrochemical anodisation is

crystallised at various annealing temperatures. As is known,

the performance of titanium dioxide depends highly on the

phase of the material. In recent years, the anatase and rutile

phases have attracted great attention, especially in medical

and electronics applications. The anatase phase is much

more efficient for the nucleation and growth of osteal cells

due to the superior bioactive properties. The crystal lattice of

the anatase phase complements the hydroxyapatite phase

well, but the rutile phase is more suitable in providing

enhanced mechanical properties (Chang et al., 2011). There-

fore, the anodised samples were annealed at different tem-

peratures at a heating and cooling rate of 5 1C/min to obtain

rutile-phase TiO2 nanostructures. Fig. 6(a–d) shows the XRD

patterns of alumina (a), alumina coated with titanium (b) and

the oxide layer of nanotube arrays prepared on a titanium

surface annealed at 450 1C and 650 1C for 2 h in a furnace with

N2 gas (c and d). The peaks for alumina, as the main

substrate, are clearly detected in Fig. 6(b–d). The temperature

of the phase transition from anatase to rutile is approxi-

mately 550 1C–600 1C, as previously reported (Lee and Park,

2003; Fan et al., 2007; Sadeghzadeh Attar et al., 2008; Sun

et al., 2011). Without taking the substrate peaks (Fig. 6(a and

b)) into account, all of the coatings’ XRD peaks can be

ascribed to the highly crystalline TiO2 anatase phase

(Fig. 6c) and rutile phase (Fig. 6d); no other phases were

detected.

3.4. Nanomechanical characteristics

In commercial nanoindenters, the load and displacement are

typically recorded continuously during the indentation pro-

cess, resulting in a load-displacement curve describing the

interaction between the tip and sample during the indenta-

tion process, as shown in (Fig. 7(a and b)). In recent years,

numerous studies on material properties have focused on the

measurement of Young’s modulus (E) and indentation hard-

ness (H) from indentation data (VanLandingham, 2003; Oliver

and Pharr, 2004; Ebenstein and Pruitt, 2006). Hardness and

and deformation behaviour of multilayer Al2O3/Ti/TiO2

f Biomedical Materials (2013), http://dx.doi.org/10.1016/

p

Fig. 5 – FESEM images of nanotube array samples: (a) 15 min of anodisation; (b) 45 min of anodisation; (c) top of surface after

75 min of anodisation; (d) bottom of nanotube; (e) cross-sectional view; and (f) outer dimensions of nanotube.

j o u r n a l o f t h e m e c h a n i c a l b e h a v i o r o f b i o m e d i c a l m a t e r i a l s ] ( ] ] ] ] ) ] ] ] – ] ] ] 7

Young’s modulus are parameters that are independent of one

another. The experimental hardness values obtained in this

study were described earlier in Table 4; these results include

the unloading data, from which Young’s modulus for the TiO2

nanophase can be calculated. The equations used to calculate

Young’s modulus are as follows:

Please cite this article as: Baradaran, S., et al., Fabricationnanotube arrays. Journal of the Mechanical Behavior oj.jmbbm.2013.01.020

1�W2s

ES¼

1ER�

1�W2D

EDð2Þ

ER ¼1

2h

ffiffiffiffiffiffiffiffiffiffip

24:5

rdLDh

� �ð3Þ

and deformation behaviour of multilayer Al2O3/Ti/TiO2

f Biomedical Materials (2013), http://dx.doi.org/10.1016/

Fig. 6 – XRD profiles for (a) alumina, (b) alumina coated with

titanium and (c) alumina/titanium/TiO2 nanotube arrays

sintered at 650 1C for 2 h (T: titanium, A: anatase, R: rutile).

Fig. 7 – Schematic of (a) a typical load versus indenter

displacement data and (b) the indentation process.

Pmax¼maximum load applied; hmax¼penetration depth;

hc¼contact depth (the height of contact between the tip and

the sample); hf¼final depth of contact impression after

unloading; S¼initial unloading stiffness.

j o u r n a l o f t h e m e c h a n i c a l b e h a v i o r o f b i o m e d i c a l m a t e r i a l s ] ( ] ] ] ] ) ] ] ] – ] ] ]8

ES ¼Kffiffiffipp

2ffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi24:56hp �

dpmax

dhmaxð4Þ

In these equations Es and Ws are the Young’s modulus and

Poisson’s ratio of the samples, respectively, and E0 and W0 are

the indenter’s (composed of diamond) corresponding proper-

ties. The composite modulus (ER) can be obtained from the

unloading slope (dL/dh) and requires determining the plastic

depth when unloading begins (hp) (Eq. 3). The plastic depth

differs from the total depth due to elastic contraction; it can

be obtained by extrapolating the unloading line from its

initial load and depth when the load is zero (Fig. 8). A value

of W¼0.28 was chosen because this is Poisson’s ratio for

commercial TiO2 (Mayo et al., 1990). dPmax/dhmax is the

contact stiffness Eq. (4) determined from the maximum load

and maximum depth at the initial unloading slope. K is the

shape constant of an indenter and is 1.034 for a Berkovich tip

Please cite this article as: Baradaran, S., et al., Fabricationnanotube arrays. Journal of the Mechanical Behavior oj.jmbbm.2013.01.020

(Chen and Duh, 2001). In this section, we describe the

nanomechanical characterisation of the titanium nanotube

structures on alumina coated with a titanium layer measur-

ing 1.4 mm in thickness. Fig. 8(a and b) shows the representa-

tive load versus displacement curves for the annealed

samples at 450 1C and 650 1C, respectively. The load forces

were applied at the different indentation depths of 550 nm

and 800 nm. In general, material properties such as strength

and hardness can change as a result of heat treatment at

different annealing temperatures. Table 4 shows the Young’s

modulus and hardness obtained from the continuous stiff-

ness measurement system. It is important to note that

Young’s modulus and hardness can be determined instanta-

neously as a function of depth. Some researchers refer to

them as the apparent hardness and Young’s modulus

because of the effect of the substrate on the coatings. More-

over, the measurement of Young’s modulus and hardness is

reliable when the indenter depth is less than 10% of the film

thickness: otherwise the substrate effect should be taken into

account. (Crawford et al., 2007; Soares et al., 2008; Toque

et al., 2010). Therefore, in this study, we fixed the depth at

550 nm and 800 nm to show the effect of substrate on the

obtained values when the displacement of the surface go

beyond the limit of 10% of the film thickness. According to

the results, the films’ Young’s modulus and hardness

increased with the increase of penetration indenter tip into

the surface with depth of 800 nm. The cause of this phenom-

enon is that during the indentation process, the densification

of the nanotube layer and the wear of the dense surface were

the two main deformation processes that occurred under the

and deformation behaviour of multilayer Al2O3/Ti/TiO2

f Biomedical Materials (2013), http://dx.doi.org/10.1016/

Table 4 – Young’s modulus characteristics and hardness of TiO2 nanotubes in various annealed temperature and depth.

Annealed temperature Max. depth (nm) Max. load (mN) Hardness (GPa) Reduced modulus (GPa) Young’s modulus (GPa)

S(4 5 0) 550 25.473.6 1.970.2 45.776.5 42.876.3

S(4 5 0) 800 38.175.8 3.970.3 88.6710.6 86.3711.2

S(6 5 0) 550 42.275.1 2.870.8 117.5711.8 117.5713.2

S(6 5 0) 800 5774.7 4.570.3 165714.1 162.279.1

Fig. 8 – Characteristic load vs. nanoindentation depth for samples at (a) 450 1C and (b) 650 1C in different indentation depth

550 nm and 800 nm.

j o u r n a l o f t h e m e c h a n i c a l b e h a v i o r o f b i o m e d i c a l m a t e r i a l s ] ( ] ] ] ] ) ] ] ] – ] ] ] 9

indenter tip. This means that the areas under the indenter tip

were densified and the other sides of the indenter were

subjected to shear stresses that induced both densification

and wear. Therefore, as the indenter penetrated the coating

surface, the modulus increased due to the increasing con-

tribution of the substrate to the overall strength of the

layered material. In addition, the densification of the area

surrounding the indentation provided minor contributions to

the material strength and led to wear between the indenter

surface and dense coating surface. As the TiO2 coating under

the indenter became increasingly dense, the modulus

increased. Therefore, it was difficult to extract a precise value

of the Young’s modulus of the porous coating itself (Crawford

et al., 2007). Moreover, the indentation results also increased

with the increase of annealing temperature due to the

phase transformation (anatase to rutile). The residual stress

(hysteresis loop) during loading and unloading indicates that

the nanotube surface undergoes elastic energy dissipation,

generating higher stiffness, an elastic response, frictional

energy, and compressive plastic deformation. Moreover, the

hysteresis loop area increased with the annealing temperature.

This area represents the energy dissipated due to material

plasticity. In general, the surface energy of the anatase phase is

lower than that of the rutile phase. The phase transition

becomes increasingly apparent with increasing annealing

temperature (Chang et al., 2011). In addition, it is notable that

the anatase phase is metastable with respect to rutile under all

conditions of temperature and pressure and reverts sponta-

neously to rutile. Anatase has a tetragonal crystal structure

where the Ti–O octahedra share four corners. Rutile has a

crystal structure similar to that of anatase with the exception

that the octahedra share four edges instead of four corners.

This leads to the formation of chains, which are subsequently

arranged in a four-fold symmetry. In comparison, the rutile

structure is more densely packed than anatase. Furthermore,

the stable form of titania is rutile, to which the other forms

Please cite this article as: Baradaran, S., et al., Fabricationnanotube arrays. Journal of the Mechanical Behavior oj.jmbbm.2013.01.020

of titania transform at sufficiently high temperatures

(Navrotsky and Kleppa, 1967).

The elastic modulus data calculated at an annealing

temperature of 450 1C are in close agreement with the

Young’s modulus values reported by Crawford et al. (2007,

2009), Shokuhfar et al. (2009) who also used nanoindentation.

Their reported values were in the range of 36–43 GPa and

23–44 GPa, while those reported for an annealing temperature

of 650 1C are quite different from reported values of approxi-

mately 105–129 GPa (de Souza et al., 2011b). Additionally, the

hardness values at calculated at different annealing tem-

peratures are similar to those reported in previous studies

(de Souza et al., 2011a; de Souza et al., 2011b). The values

reported by Souza et al. were approximately 2–3 GPa.

4. Conclusion

We have reported the microstructural characterisation and

deformation behaviour of TiO2 nanotube arrays on alumina

substrates. The following conclusions can be drawn:

(1)

anf B

Ti was successfully deposited as a first layer on alumina

substrates by reactive RF magnetron sputtering (PVD),

which the strong surface adhesion obtained at an RF

power of 400 W, temperature of 250 1C and substrate bias

voltage of 75 V using the Taguchi optimisation method.

For the film exhibiting the strongest adhesion, the scratch

length was 1090.36 mm, the failure point was approxi-

mately 995.33 mm and the load was 1238 mN. The failure

point and scratch length in the film exhibiting the weak-

est adhesion were 388.27 mm and 1102.63 mN, respec-

tively. The graph of load versus distance shows that the

critical load was 437 mN.

(2)

TiO2 nanotubes were synthesised as a second layer by

electrochemical oxidation using a mixture of NH4F,

d deformation behaviour of multilayer Al2O3/Ti/TiO2

iomedical Materials (2013), http://dx.doi.org/10.1016/

j o u r n a l o f t h e m e c h a n i c a l b e h a v i o r o f b i o m e d i c a l m a t e r i a l s ] ( ] ] ] ] ) ] ] ] – ] ] ]10

Plnj.j

ethylene glycol and water as an electrolyte. The FESEM

results indicate that the best anodisation time, i.e., that

allowing for the most suitable nanotube formation, was

75 min.

(3)

The phase transition temperature for the transformation

of the anatase phase to the rutile phase ranged from

450 1C to 650 1C.

(4)

The results of mechanical indentation tests indicate that

Young’s modulus and hardness increased with annealing

temperature.

Acknowledgment

The authors wish to thank Pardis Moslemzadeh Tehrani for

valuable discussions. This work has been supported by the

University of Malaya, grant No: (HIR F0004-21001), FRGS grant

number FP039 2010B and RG 181 12SUS.

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and deformation behaviour of multilayer Al2O3/Ti/TiO2

f Biomedical Materials (2013), http://dx.doi.org/10.1016/