Articulo. Development of Ti–Nb–Zr Alloys With High Elastic Admissible Strain

Page 1

7/21/2019 Articulo. Development of Ti–Nb–Zr Alloys With High Elastic Admissible Strain

http://slidepdf.com/reader/full/articulo-development-of-tinbzr-alloys-with-high-elastic-admissible-strain 1/12

Development of Ti–Nb–Zr alloys with high elastic admissible strain

for temporary orthopedic devices

Sertan Ozan a,b, Jixing Lin c,d, Yuncang Li e, Rasim Ipek b, Cuie Wen a,e,⇑

a Faculty of Engineering and Industrial Sciences, Swinburne University of Technology, Hawthorn, Victoria 3122, Australiab Department of Mechanical Engineering, Ege University, 35100 Bornova, Izmir, Turkeyc Advanced Material Research and Development Center, Zhejiang Industry & Trade Vocational College, Wenzhou, Zhejiang 325003, Chinad Department of Materials Science and Engineering, Jilin University, Changchun, Jilin 130025, Chinae School of Aerospace, Mechanical and Manufacturing Engineering, RMIT University, Melbourne, Victoria 3083, Australia

a r t i c l e i n f o

 Article history:

Received 13 September 2014

Received in revised form 19 March 2015

Accepted 20 March 2015

Available online 25 March 2015

Keywords:

Elastic admissible strain

Cytocompatibility

TNZ (Ti–Nb–Zr) alloys

Mechanical properties

Young’s modulus

a b s t r a c t

A new series of beta Ti–Nb–Zr (TNZ) alloys with considerable plastic deformation ability during

compression test, high elastic admissible strain, and excellent cytocompatibility have been developed

for removable bone tissue implant applications. TNZ alloys with nominal compositions of Ti–34Nb–

25Zr, Ti–30Nb–32Zr, Ti–28Nb–35.4Zr and Ti–24.8Nb–40.7Zr (wt.% hereafter) were fabricated using the

cold-crucible levitation technique, and the effects of alloying element content on their microstructures,

mechanical properties (tensile strength, yield strength, compressive yield strength, Young’s modulus,

elastic energy, toughness, and micro-hardness), and cytocompatibilities were investigated and compared.

Microstructural examinations revealed that the TNZ alloys consisted of  b  phase. The alloy samples dis-

played excellent ductility with no cracking, or fracturing during compression tests. Their tensile strength,

Young’s modulus, elongation at rupture, and elastic admissible strain were measured in the ranges of 

704–839 MPa, 62–65 GPa, 9.9–14.8% and 1.08–1.31%, respectively. The tensile strength, Young’s modulus

and elongation at rupture of the Ti–34Nb–25Zr alloy were measured as 839 ± 31.8 MPa, 62 ± 3.6 GPa, and

14.8 ± 1.6%, respectively; this alloy exhibited the elastic admissible strain of approximately 1.31%.

Cytocompatibility tests indicated that the cell viability ratios (CVR) of the alloys are greater than those

of the control group; thus the TNZ alloys possess excellent cytocompatibility.

  2015 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.

1. Introduction

According to a report by the United Nations in 2013   [1],   the

population of the age group 60 years and older is increasing

rapidly, and in the more developed regions of the world, the ratio

of elderly people is expected to increase by 45% as of 2050.

Orthopedic biomaterials are implanted into human bodies in order

to heal bone tissue diseases developed due to aging, various con-genital defects in bone tissue, injuries to bone tissue and joints

from traffic and sports accidents  [2,3].

With the increasing utilization of orthopedic implant materials,

if biologically and mechanically compatible materials are not used

for implants, it is inevitable that the number of revision surgeries

will increase   [4]. Genotoxic, cytotoxic, carcinogenic, mutagenic,

allergenic, neurological effects are taken into account while

evaluating the biological compatibility of an implant material  [5].

The formation of stress shielding may occur in cases where an

implant material with a higher Young’s modulus than that of bone

tissue is used   [6]. Resorption in bone tissue can occur when it

carries lower amounts of biological load due to a rigid implant

material carrying most of the loads  [7]. In such cases, the implant

material may loosen [6]. The fact that titanium (Ti) alloys have a

lower Young’s modulus in comparison with 316L stainless steeland Co–Cr alloys is an important factor in their increasing use as

implant materials [3].

Since Young’s modulus of the (a + b) Ti–6Al–4V alloy (at

110 GPa) [8] is about 4 times greater than that of cortical bone tis-

sue (max. 27 GPa) [9],  b  Ti alloys with a relatively lower Young’s

modulus in comparison with a + b  Ti alloys have been developed

for orthopedic implant applications  [10–16]. As it is well known,

among the Ti alloys, Ti–6Al–4V is still the most used one for ortho-

pedic applications, even though it contains vanadium, which is

defined as toxic   [17], as well as having aluminum, for which

neurological side-effects [18], and genotoxic effects [19] have been

http://dx.doi.org/10.1016/j.actbio.2015.03.023

1742-7061/  2015 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.

⇑ Corresponding author at: School of Aerospace, Mechanical and Manufacturing

Engineering, RMIT University, Melbourne, Victoria 3083, Australia. Tel.: +61 3 9925

7290.

E-mail address:  [email protected] (C. Wen).

Acta Biomaterialia 20 (2015) 176–187

Contents lists available at  ScienceDirect

Acta Biomaterialia

j o u r n a l h o m e p a g e :   w w w . e l s e v i e r . c o m / l o c a t e / a c t a b i o m a t

Page 2

7/21/2019 Articulo. Development of Ti–Nb–Zr Alloys With High Elastic Admissible Strain

http://slidepdf.com/reader/full/articulo-development-of-tinbzr-alloys-with-high-elastic-admissible-strain 2/12

reported. Another important consideration is that the cold-forming

capabilities of  b  Ti alloys (e.g. Ti–15Mo–5Zr–3Al) are much better

in comparison witha + b Ti alloys (e.g. Ti–6Al–4V) [20]. The super-

ior cold-forming ability of  b  Ti alloys decreases the manufacturing

cost, thus creating an advantage for them in becoming commercial

products  [4].

The role of Ca–P formations on implant’s surface is to accelerate

the assimilation of implant material and bone-tissue   [21]. For

instance, on the surface of bio-inert Zirconia, calcium phosphate

formation enabling a chemical bond with the surrounding bone-

tissue does not occur, it only ensures morphological fixation while

staying in the body  [22]. Ca–P formations enable a tight bond

between the implant and surrounding bone-tissue   [23]. On the

other hand, it is crucial to prevent the formation of fibrous tissue

between the bone-tissue and implants [24]. A direct contact with-

out fibrosis tissue between implant material and bone is defined

as osseointegration. [24]. Implants may be isolated due to fibrous

tissue formed between the bone tissue and implant  [25]. It is

beyond doubt that fibrous tissue formation between the implant

and surrounding tissue is not desired, it is a prerequisite for both

temporary and permanent implant materials. Unlike osseointegra-

tion, chemical bonding between implant and bone tissue is related

to calcium phosphate forming ability of implants  [26]. The fast

formation of fixation betweenthe implant materialand bone is pro-

vided by calcium phosphate precipitates on the surface of the

implant material, which thereby increases bone conduction   [27].

Methods such as alkaline treatment [28], micro-arc oxidation [29]

and sol–gel   [30]   are used in order to enhance the adherence of 

the implant material to the bone tissue or, in other words, to

increase chemical bonding ability by providing calcium phosphate

formation on the surface. Permanent implants (e.g. artificial joints

and dental implants) should have an assimilation ability with bone

tissue when assessed on the basis of clinic needs [31]. It is expected

that such an implant should offer excellent bioactivity   [32]. It is

well known that artificial joints take on the natural joint function

by replacing the severely damaged joint. It is crucial to introduce

artificial joints having the ability of assimilation with bone tissuewhich ensures patients to carry out daily activities properly for

the rest of their lives [33]. When viewed under this aspect, the bone

conductivity of permanent implant materials should be high  [31].

However, the reverse is required in removable implant applica-

tions [34]; that is, calcium phosphate precipitates are undesirable

on the surface of removable implants so as to minimize the bone

conductance of the implant material, and to prevent assimilation

between the bone tissue and the implant material  [27]. Contrary

to permanent implants, the bonding between implant and bone

should be weak enough to prevent refracture of the bone during

removal surgery [35]. Trauma implants, e.g. bone nails, lose their

function after a healing period and they are removed after fracture

union [36]. Unlike artificial joints, trauma implants (e.g. bone nails)

do not lead to any restriction of movement function when removedfrom the body at an assigned time after healing period of the bone

tissue, i.e., it is not necessary to remain in the body permanently as

artificial joints. Because of the tendency of calcium phosphate pre-

cipitation on the surface of some Ti alloys, assimilation occurs

between the bone tissue and the implant material which makes

them unadaptable as a choice of temporary implants [31]. It is cru-

cial to avoid assimilation of these devices with bone tissue for an

easy removal surgery [37]. In cases where there is no assimilation

between bone and implant material, the occurrence of new bone

fractures during surgery to remove the implant material is mini-

mized, and so the removal surgery is easier  [34]. Implant removal

becomes necessary with patient demands and complaints   [38].

Especially in the cases of implants used for athletes engaged in

contact sports   [39]   and for children  [40], the implants may beremoved after a healing period   [34].   This necessitates the design

of materials with a low Young’s modulus and high mechanical

strength, as well as low bone conductance for removable implant

applications   [31,34,41]. Problems that arise during implant

removal surgery due to the assimilation of the bone and the

implant material can be prevented by inhibition of calcium phos-

phate precipitations on the surfaces of the implant materials [34].

It is possible to prevent bone atrophy during the time that the

implant stays in the body by using implant materials with a low

Young’s modulus   [42]. By minimizing or eliminating the stress

shielding effect with the use of implant materials that have a low

Young’s modulus, problems relating to long recovery times of 

patients following removal operations [39] can be resolved.

The objective of this study was to develop Ti alloys with low

rigidity, high mechanical strength, and reasonable elongation at

rupture for removable bone tissue implant applications. To this

end, studies showing that zirconium (Zr) prevents calcium phos-

phate formation on the surface of a material [27,43,44] and studies

describing the mechanical properties and phase stability changes

of Ti–Nb–Zr alloys containing high amounts of Zr  [45,46]   were

taken into account, and TNZ alloys with high Zr element (between

25 wt.% and 40.7 wt.%) were manufactured. Niobium (Nb), another

of the alloy elements used in this study, is a  b  isomorphous alloy

element [47]. Whereas the effect of Zr on binary Ti alloys is evalu-

ated as either neutral or weak  b  stabilization, Zr functions as an

effective   b   stabilizer in multi-element Ti alloys that contain Nb

or Ta  [48]. In this study, the effects of Zr and Nb content on the

microstructures, mechanical properties, and cytocompatibility of 

TNZ alloys with different electronic parameters were investigated

and compared.

2. Materials and methods

 2.1. Alloy design

Using a trial-and-error method for the design of Ti alloys results

in time-consuming and uneconomical manufacturing processes[49].  The approach developed by Morinaga et al.   [50],   known as

the d-electron alloy design method, exploits the relationship

between the phase stability and elastic properties of Ti alloys by

using the electronic parameters of average bond order (Bo) and aver-

agemetal d-orbital energylevel (M d) [48]. In theliterature,thereare

reports of many Ti alloys that have been designed using the d-elec-

tron alloy design method [11,51–53]. It was reported by Abdel et al.

[48]   that Young’s modulus decreases along the   b/b + x   phase

boundary with an increase in   Bo   and   M d   values, as seen in the

Bo M d   diagram (Fig. 1) (adapted and redrawn from Ref.   [48],

Copyright (2006), with permission from Elsevier).   b/b + x   phase

boundary presented with a dotted curve in Fig. 1, means that, the

phase stability of titanium alloys depends on the amount of other

alloying elements (e.g. Hf, Zr), as explained in Ref.  [45].Bo is the measure of the covalent bond strength between Ti and

the alloying element, whereas the   M d  of the alloying transition-

metal is a value that relates the metallic radius of the elements

and their electronegativities [48]; these two parameters are values

that have been theoretically determined using the orbital method

in body-centered cubic titanium using the DV-Xa molecular orbital

method [48,50]. The average Bo  and M d  values are calculated using

the formula given by [54]:

Bo ¼ R X iðBoÞi   ð1Þ

M d ¼ R X iðM dÞi   ð2Þ

where X i  is the atomic ratio of the given element, (Bo)i   is the bondorder value of the given element   [54], and (M d)i   is the metal

S. Ozan et al./ Acta Biomaterialia 20 (2015) 176–187    177

Page 3

7/21/2019 Articulo. Development of Ti–Nb–Zr Alloys With High Elastic Admissible Strain

http://slidepdf.com/reader/full/articulo-development-of-tinbzr-alloys-with-high-elastic-admissible-strain 3/12

d-orbital energy level of the alloying transition-metal value of the

given element [54].

In this study, we combined the d-electron alloy design method

[50] with the molybdenum equivalence (Moeq) [55] and electron to

atom ratio ((e/a)ratio)   [20]   approaches in the design of the TNZ

alloys. The Moeq   values were designed to be larger than 6 and

the (e/a)ratio of the TNZ alloys were in the range of 4.18–4.25. The

electronic parameters of the alloys including the  Bo M d  values,

Moeq, and (e/a)ratio are summarized in Fig. 1 (adapted and redrawn

from Ref.   [48], Copyright (2006), with permission from Elsevier).

The locations of the designed TNZ alloys are shown on the

Bo M d  graph, above and as close as possible to the  b/b + x  phase

boundary.

 2.2. Material preparation

Since Ti is very active chemically at high temperatures, difficul-

ties may arise in obtaining a chemically homogeneous structure in

the production of Ti alloys in traditional arc or induction melting

furnaces, so the cold-crucible levitation melting method is used

for the production of Ti alloys via casting [51]. This method ensures

that the alloy is melted without any contact between the melt and

the melting crucible, and is also not affected by the melting point

difference between Ti and the alloying element [51]. The nominal

compositions of the TNZ alloys were Ti–34Nb–25Zr, Ti–30Nb–

32Zr, Ti–28Nb–35.4Zr and Ti–24.8Nb–40.7Zr; these alloys are

hereafter denoted TNZ-1, TNZ-2, TNZ-3, and TNZ-4, respectively.

As seen in  Fig. 1  (adapted and redrawn from Ref.   [48], Copyright(2006), with permission from Elsevier), with an increase in the

concentration of Zr or a decrease in the concentrations of Nb and

Ti, the position of the alloy moves from TNZ-1 to TNZ-4 along

the b/b + x boundary. Simultaneously, the Bo M d values increase,

while the Moeq value and the (e/a)ratio decrease. In the diagram, the

positions of the alloys are situated just above the boundary with as

close proximity as possible. Ingots of the alloys were re-melted 5

times in order to guarantee chemical homogeneity. Commercially

pure titanium (CP-Ti) Grade-2, furnished as cold worked, was used

as a control group for tensile test, as well.

 2.3. Microstructural characterization

Disk-shaped samples with a diameter of 8 mm and thickness of 2 mm were cut from the ingots for phase analysis, microstructure

examination, and micro-hardness measurements using electrical

discharge machining (EDM). The samples used for microstructure

examination, phase analysis, and micro-hardness measurements

were consecutively ground with SiC grinding papers of up to

4000 grit. Following the grinding process, the samples were

polished using a colloidal silica and H2O2   mixture suspension,

and afterward the samples were cleaned for 5 min in an ultrasonic

bath using ethanol. Phase analyses of the samples were carried out

using a Bruker D8 Advance X-ray diffractometer (XRD) with a Cu

Ka radiation source at room temperature. Microhardness measure-

ments were made on the polished surfaces of the samples in accor-

dance with procedure listed in ASTM standard E384-94   [56]   by

using microhardness tester (Buehler Micromet 2100). Nine differ-

ent measurements were made, and the minimum and maximum

microhardness values were not taken into account when calculat-

ing the average micro-hardness values. The samples for

microstructure characterization were etched with Kroll solution.

The microstructure of TNZ alloys was examined by optical micro-

scopy (Leica-MEF4M).

 2.4. Mechanical property testing 

Tensile samples with a gauge section of 8 mm 2 mm 1 mm

were cut using EDM. Both surfaces of the samples were ground

with SiC grinding papers of up to 2400 grit. Tensile tests were

carried out at 0.5 mm/min cross-head speed at room temperature

using an Instron 5567 testing system with advanced video exten-

someter. The gauge length of 8 mm was marked prior to the test

so that the video extensometer could detect it. Using the stress–

strain curves obtained from the tensile tests, tensile strength, yield

strength, Young’s modulus, elastic energy, toughness and elonga-

tion at rupture of the TNZ alloys were measured.

The slope of the linear part of the stress–strain curves was deter-

mined as Young’s modulus, and the 0.2% offset yield method was

used to determine the yield strength of the TNZ alloys. Tensile tests

were repeated on 5 samples for each alloy, and the average values

obtained from the 5 measurements were used in the calculationof the mechanical properties. The fracture surfaces of the tensile

samples were examined using field-emission scanning electron

microscopy (SEM, ZEISS SUPRA 40 VP). Cylindrical samples 5 mm

in diameterand 8 mmin lengthweremachined using EDM for com-

pression tests. The tests were carried out at a cross-head speed of 

0.5 mm/min, using an MTS testing system at room temperature in

accordance with procedure listed in ASTM standard E9-09   [57].

The compressive yield strength of the TNZ alloys was calculated

from the averages of 3 samples for each alloy. As it is seen in

Fig. 5   (indicated by blue arrow), the compressive stress–strain

graph of TNZ alloys display ‘‘pre-deformation phenomenon’’   [58]

which makes it impossible to use the 0.2% offset approach for

calculating the compressive yield strength of the alloys.

Consequently, ‘‘two lines method’’ [58] was used to be able to cal-culate the compressive yield strength of the alloys.

 2.5. MTS assay

Disk samples 8 mm in diameter and 2 mm in thickness for cell

culture were cut from the ingots using EDM. All samples for the

cytocompatibility assessment were sterilized in a muffled furnace

at 180 C for 3 h. The cytocompatibility of all alloys was evaluated

using a direct cell method according to the International Standard

ISO10993-5  [59]. The disk shaped TNZ and CP-Ti specimens were

put into the wells of the cell culture plate. At a density of 5 103

cells per well (Barwon Biomedical Research, Geelong Hospital,

Victoria, Australia) were seeded onto the surface of the samples.

In order to determine the in vitro proliferation of cells, MTSassay was used   [60,61]. The cell morphology was observed by

Fig. 1.  Location of TNZ alloys on  B o M d   map and electronic parameters of TNZ

alloys (adapted and redrawn fromRef. [48], Copyright (2006), withpermission from

Elsevier).

178   S. Ozan et al. / Acta Biomaterialia 20 (2015) 176–187 

Page 4

7/21/2019 Articulo. Development of Ti–Nb–Zr Alloys With High Elastic Admissible Strain

http://slidepdf.com/reader/full/articulo-development-of-tinbzr-alloys-with-high-elastic-admissible-strain 4/12

means of SEM after cell culturing. Prior to SEM observation of 

SaOS2 cells attached to the surface of samples, a series of processes

were carried out. Samples, on which cells were seeded, were fixed

in glutaraldehyde (3.9%) at room temperature for 1 h. Following

this process, dehydration of cells was performed by washing con-

secutively in ethanol solutions with a concentration of 60%, 70%,

80%, 90%, and 100%. Chemical drying of samples was performed

by using hexamethyldisilazane. Finally, by coating the samples

with gold, the samples were made ready for SEM observation. In

this study, all cell culture experiments were conducted in triplicate

and results were stated as means ± standard deviations.

 2.6. Statistical analysis

The mechanical property values (tensile and yield strength,

Young’s modulus, elongation at rupture, elastic energy, toughness,

compressive yield strength, micro-hardness) and cell adhesion

density values were analyzed by using one-way ANOVA

(q < 0.05) followed by Tukey HSD post hoc test. q  < 0.05 was taken

into consideration statistically significant.

3. Results and discussion

 3.1. Microstructure of the TNZ alloys

It was determined following microstructural examinations that

the TNZ alloys display primary grain boundaries together with sec-

ondary grain structure, as shown in   Fig. 2. In  Fig. 2(c), dendritic

structure was clearly observable.

The XRD patterns of the TNZ alloys are given in Fig. 3. As shown,

only  b  diffraction peaks were detected in each TNZ specimen. The

diffraction peaks of b Ti shifted toward lower 2h degree values with

increases in the amount of Zr or decreases in the amounts of Nb

and Ti in the alloys. Similar results were reported by Ning et al.

[62]; however, in that study [62], the Nb level was kept constant

while increasing the Zr and decreasing the Ti, and it was suggested

that, as a result of the increase in Zr, the atomic radius of which is

larger in comparison with Ti, a greater number of Zr atoms

replaced Ti atoms, resulting in a shift of the peaks toward lower

2h degree values [62].

 3.2. Micro-hardness of the TNZ alloys

Fig. 4  presents the microhardness of the alloys in comparison

with that of other Ti alloys as published in the literature, respec-

tively. As seen in Fig. 4, it is worth noting that the average micro-hardness levels of TNZ-1 (241.3 HV), TNZ-2 (244.4 HV), TNZ-3

(244.0 HV) and TNZ-4 (235.6 HV) are higher than those of CP-Ti

(134.0 HV)   [63], Ti–29Nb–13Ta–4.6Zr (solution treated condition

– 180 HV)   [64]   and Ti–13Nb–13Zr (solution treated condition –

235 HV)   [65], and lower than that of Ti–6Al–4V (as-cast – 294

HV) [66]. The micro-hardness value of the TNZ-1,2,3 alloys was sig-

nificantly (q < 0.05) higher than that of the TNZ-4 alloy. When

ranked the average micro-hardness value of TNZ alloys among

themselves, the alloys are ranked TNZ-2 > TNZ-3 > TNZ-1 > TNZ-4.

 3.3. Compressive properties of the TNZ alloys

The compressive stress–strain graphs of the TNZ alloys are

shown in Fig. 5. It can be seen that under compression, the alloysexhibited considerable plastic deformation ability. As seen in

Fig. 5, the room temperature compression stress–strain curve of 

the TNZ alloys is consisted of ‘‘elastic stage (I), a plastic yield plateau

stage (II), a parabola stage (III)’’ [67], consecutively. TNZ-2 displayed

the highest compressive yield strength among all the TNZ alloys,

and TNZ-3 displayed a compressive yield strength very close to

that of TNZ-2. Representative images of the alloy samples taken

before and after compression are inserted in  Fig. 5, showing that

neither cracking nor fracturing was observed. Compression tests

were stopped after reaching a strain value of 65%.

The average compressive yield strength of the TNZ alloys is

given in   Fig. 6. There are no significant (q > 0.05) differences

among the compressive yield strength of the alloys. However,

when a ranking in terms of average compressive yield strength is

made, the alloys are ranked TNZ-2 > TNZ-3 > TNZ-1 > TNZ-4. The

small differences among the compressive yield strengths of the

Fig. 2.  Optical micrographs of as-cast TNZ alloys: (a) TNZ-1; (b) TNZ-2; (c) TNZ-3; (d) TNZ-4.

S. Ozan et al./ Acta Biomaterialia 20 (2015) 176–187    179

Page 5

7/21/2019 Articulo. Development of Ti–Nb–Zr Alloys With High Elastic Admissible Strain

http://slidepdf.com/reader/full/articulo-development-of-tinbzr-alloys-with-high-elastic-admissible-strain 5/12

alloys may have been caused by different second-phase particles

[68], grain size, or solid-solution strengthening effects   [69].

Similarly, it can be observed that the microhardness levels of the

alloys present the same ranking order (Fig. 4).

It is worth noting that the compressive yield strength of the TNZ

alloys is about 4 times greater than the compressive yield strength

of bone tissue (Human Haversian: 180 MPa)   [70]. However, in

recent years, metal foam materials have been manufactured that

compromise on the mechanical strength of the implant material

in order to ensure that the Young’s modulus value of the material

is similar to that of bone tissue [61,71]. Even though the design of 

foam material takes bone ingrowth into consideration [72], the use

of such materials, especially for orthopedic implants carrying high

loads, is still uncommon due to the fact that metal foams possess

Fig. 3.  XRD patterns of as-cast TNZ alloys.

Fig. 4.  Microhardness of TNZ alloys in comparison with those of other Ti alloys published in the literature.

180   S. Ozan et al. / Acta Biomaterialia 20 (2015) 176–187 

Page 6

7/21/2019 Articulo. Development of Ti–Nb–Zr Alloys With High Elastic Admissible Strain

http://slidepdf.com/reader/full/articulo-development-of-tinbzr-alloys-with-high-elastic-admissible-strain 6/12

lower mechanical strength, i.e. fatigue strength, tensile strength

and compression strength, in comparison with their solid counter-parts [73]. The fact remains that, metal foams also possess lower

corrosion resistance [74].

 3.4. Tensile properties of the TNZ alloys

Fig. 7 shows the stress–strain graphs of the TNZ alloys obtained

via tensile tests. The tensile properties (tensile strength, yield

strength, and elongation at rupture) of the alloys are given in

Fig. 8. TNZ-1 exhibited the highest average tensile strength and

yield strength, and largest elongation. The ranking of the average

tensile strength, yield strength, and elongation of the alloys from

the highest to lowest is TNZ-1 > TNZ-2 > TNZ-3 > TNZ-4. In particu-

lar, the elongation at rupture of the alloys was in the range of 9.9–14.8%. When the TNZ alloys were evaluated statistically among

themselves; it was determined that the tensile and yield strength,

elongation at rupture, toughness, elastic energy of TNZ-1 were sig-

nificantly (q < 0.05) higher than TNZ-3 and TNZ-4 alloy. The fact

remains that, TNZ-2 had significantly (q < 0.05) higher tensile

and yield strength, elongation at rupture, toughness and elastic

energy than that of the TNZ-4, as well. However, there were no sig-

nificant differences (q > 0.05) in terms of Young’s modulus among

TNZ alloys. The tensile strength, yield strength, Young’s modulusand elongation at rupture of the CP-Ti, used as a control in this

study, were measured as 758 ± 14.1MPa, 607 ± 10.3 MPa,

109 ± 1.1 GPa, and 26 ± 0.2%, respectively. By statistically compar-

ing Young’s modulus of CP-Ti and TNZ alloys, it was determined

that CP-Ti had significantly (q < 0.05) higher Young’s modulus in

comparison with TNZ alloys.

It is noticeable that the elongation values of these as-cast TNZ

alloys are comparable to or slightly higher than those of the

thermomechanically processed biomedical Ti alloys currently in

use, such as Ti–12Mo–6Zr–2Fe   [75], Ti–6Al–7Nb   [76], Ti–15Mo

[77], Ti–13Nb–13Zr [78] and as-cast Ti–6Al–4V [79] with elonga-

tions of 12%, 10%, 12%, 15% and 8%, respectively. It should be noted

that as-cast Ti alloys generally present relatively lower elongation

values compared to those of their thermomechanically processedcounterparts [8]. However, the as-cast TNZ alloys studied in this

research exhibited fairly large elongation values that are compara-

ble to those of the annealed Ti alloys noted above.

In the case of the elastic deformation of material, the ability to

release absorbed energy on unloading is known as resilience, and

this property is measured by the resilience modulus or elastic

energy   [80]. In calculating the elastic energy of the TNZ alloys,

the yield strength was taken as the elastic limit   [80]. Toughness

is a measure indicating the ability of a material to absorb energy

during tensile deformation till fracture [80]. The toughness values

obtained by calculating the area under the tensile curve along with

the elastic energy values denoting the absorbed energy when the

alloys deformed elastically, and Young’s modulus values of 

the alloys are given in Fig. 9. The elastic energy and toughness of the alloys were calculated using a Matlab-R2012b. The ranking

Fig. 5.  Compressive stress–strain curve of as-cast TNZ alloys.

Fig. 6.  Compressive yield strength of as-cast TNZ alloys.

S. Ozan et al./ Acta Biomaterialia 20 (2015) 176–187    181

Page 7

7/21/2019 Articulo. Development of Ti–Nb–Zr Alloys With High Elastic Admissible Strain

http://slidepdf.com/reader/full/articulo-development-of-tinbzr-alloys-with-high-elastic-admissible-strain 7/12

of the toughness and elastic energy of the alloys from the highest

to lowest is TNZ-1 > TNZ-2 > TNZ-3 > TNZ-4. As seen in Fig. 9, the

average Young’s modulus of the alloys was in the range of 

62–65 GPa.

Ti alloys that are used as implant material are required to pos-

sess high mechanical strength and a low Young’s modulus, so thatthey can safeguard the host bone tissue without causing stress

shielding, which could lead to bone resorption and cause implant

loosening   [81]. The mechanically compatible performance of an

orthopedic implant material can be evaluated by the elastic admis-

sible strain [82], the ratio of yield strength to Young’s modulus of 

the material given by [4]:

d ¼  r=E    ð3Þ

Fig. 7.  Tensile stress–strain curves of as-cast TNZ alloys.

Fig. 8.  Tensile mechanical properties (tensile strength, yield strength and elongation at rupture) of TNZ alloys.

182   S. Ozan et al. / Acta Biomaterialia 20 (2015) 176–187 

Page 8

7/21/2019 Articulo. Development of Ti–Nb–Zr Alloys With High Elastic Admissible Strain

http://slidepdf.com/reader/full/articulo-development-of-tinbzr-alloys-with-high-elastic-admissible-strain 8/12

where d is the elastic admissible strain, r is the yield strength and E 

is Young’s modulus of the material.

The mechanical properties of the TNZ alloys investigated in this

study, and of Ti–16Nb–10Zr   [83] and some of the commercial Tialloys  [8,81,84]   published in the literature are listed in   Table 1.

As can be seen, the elastic admissible strain values of the TNZ

alloys are in the range of 1.08–1.31%, and the highest elastic admis-

sible strain value for commercial Ti alloys stated in Table 1 is 0.99%

for annealed Ti–35.3Nb–5.1Ta–7.1Zr [8]. It can be concluded that

the TNZ alloys would offer promising performance in practical

applications in terms of mechanical compatibility. The Young’s

modulus values of TNZ-1, TNZ-2, TNZ-3, and TNZ-4 are 62.0 ± 3.6,

65.0 ± 4.2, 64.0 ± 4.5 and 63.0 ± 4.2 GPa, respectively. Young’s

modulus values of the TNZ alloys are lower than those of pure

CP-Ti (104 GPa)   [8]  and mill-annealed Ti–6Al–4V (110 GPa)  [84],

and close to those of Ti–35.3Nb–5.1Ta–7.1Zr (55 GPa)   [8]   and

Ti–29Nb–13Ta–4.6Zr (60 GPa) [81].

The fracture surfaces of the TNZ alloys after tensile tests areshown in Fig. 10. It can be seen that the alloys exhibited a ductile

fracture surface with numerous dimples, exemplified by the white

arrows in Fig. 10(a–d). It is noticeable that the number of dimples

for TNZ-1 is particularly large and the dimples observed on the

fracture surface are deeper, in comparison with TNZ-2, TNZ-3,and TNZ-4. In addition, cleavages (indicated by the black arrows

in Fig. 10) were observed accompanying the dimples on the frac-

ture surfaces of the alloys. In particular, cleavage was the predomi-

nant fracture mechanism in TNZ-4. It is well known that the

fracture surfaces of ductile materials are rougher, whereas those

of brittle materials are smooth. Intergranular fracturing, which is

one of the indicators of brittle fracturing, was not observed on

the fracture surfaces of the alloys.

 3.5. Cytocompatibility of the TNZ alloys

The healthy adhesion of osteoblast cells to the implant surface,

and their growth and proliferation, are important indicators of thecytocompatibility of the implant materials   [61]. To this end, the

Fig. 9.  Tensile mechanical properties (Young’s modulus, elastic energy and toughness) of TNZ alloys.

 Table 1

Comparison of mechanical properties of TNZ alloys with those of some commercial titanium alloys.

Ti alloys Yield strength

(MPa)

Tensile strength

(MPa)

Young’s

modulus (GPa)

Elongation (%) Elastic admissible

Strain (%)

Alloy type References

Pure Ti (Grade 4) 485 550 104.1 15 0.47   a   [8]

Ti–16Nb–10Zr (annealed) 485 520 70 22 0.69   b + a’’   [83]

Ti–15Mo (annealed) 544 874 78 21 0.70   b   [84]Ti-6-4 ELI (mill annealed) 875 965 110 10–15 0.80   a + b   [84]

Tiadyne 1610 (aged) 736 851 81 I0 0.91   b   [84]

Ti–35.3Nb–5.1Ta–7.1Zr (annealed) 547 597 55 19 0.99   b   [8]

Ti–29Nb–13Ta–4.6Zr (annealed) – 549 60 41.6 –   b   [81]

CP-Ti (cold worked) 607 ± 10.3 758 ± 14.1 109 ± 1.1 26 ± 0.2 0.56   a   [This study]

Ti–34Nb–25Zr (TNZ-1) 810 ± 48.0 839 ± 31.8 62 ± 3.6 14.8 ± 1.6 1.31   b   [This study]

Ti–32Zr–30Nb (TNZ-2) 782 ± 18.8 794 ± 10.1 65 ± 4.2 13 ± 1.8 1.20   b   [This study]

Ti–35.4Zr–28Nb (TNZ-3) 729 ± 26.3 755 ± 28.3 64 ± 4.5 11.3 ± 1.5 1.14   b   [This study]

Ti–40.7Zr–24.8Nb (TNZ-4) 682 ± 52.9 704 ± 49.5 63 ± 4.2 9.9 ± 1.1 1.08   b   [This study]

S. Ozan et al./ Acta Biomaterialia 20 (2015) 176–187    183

Page 9

7/21/2019 Articulo. Development of Ti–Nb–Zr Alloys With High Elastic Admissible Strain

http://slidepdf.com/reader/full/articulo-development-of-tinbzr-alloys-with-high-elastic-admissible-strain 9/12

in vitro cytocompatibility of the TNZ alloys was assessed using

SaOS2 cells.

Following the cell culture of the TNZ alloys for 24 h, the surfaces

of alloys were examined in terms of the morphology of osteoblastcells using SEM. The adherence of osteoblast cells to the surfaces of 

the samples is shown in Fig. 11. It can be seen that the osteoblast

cells started to proliferate following 24 h of cell culture. The osteo-

blast cells on the alloy surfaces display a polygonal morphology,

and a large number of filopodia have formed, providing anchoring

of the osteoblast cells on the materials’ surfaces (Fig. 11). The for-

mation of filopodia on the surfaces after cell culture is an indica-

tion that the initial cell adhesion and the growth of the cells

were healthy and strong  [85]. Thus it can be concluded that the

TNZ alloys are favorable for cell growth and proliferation.

SEM images of the SaOS2 osteoblast cells on the surfaces of the

TNZ alloys after cell culture for 7 days are given in Fig. 12. It can be

seen that the SaOS2 cells display a healthy growth on the surfaces

of the alloys. Numerous osteoblast cells firmly adhered to andspread over the alloy surfaces. There was no difference in terms

of the morphology of the osteoblast cells on the four different alloy

surfaces (Fig. 12). Unhealthy cells are distinguished from healthy

cells on the basis of their reduced cell morphology caused by

shrinkage of the cells   [86]. When this morphological property

specific to unhealthy cells was considered, it was observed that

very healthy cells proliferated on the surfaces of the alloys.

Fig. 13 shows a comparison of cell adhesion density for the TNZ

alloys and CP-Ti after 7 days of cell culture. After assessment of the

images of the osteoblast cells on the alloy surfaces, along with the

cell adhesion densities given in  Fig. 13, it can be concluded that

osteoblast cells adhered well to the surfaces, and spread and grewin a healthy way. As can be seen from Fig. 13, the average cell adhe-

sion density of the TNZ alloys is greater than that of CP-Ti. Ranking

of the TNZ alloys and CP-Ti in terms of average cell adhesion den-

sity is TNZ-4 > TNZ-3 > TNZ-2 > TNZ-1 > CP-Ti. No significant dif-

ferences (q > 0.05) were confirmed in cell adhesion densities

between CP-Ti and TNZ-1, TNZ-2, TNZ-3 alloys. However, the cell

adhesion density of TNZ-4 was significantly (q < 0.05) higher than

CP-Ti. According to the obtained results, an increase in Zr in the

TNZ alloys and decreases in Nb and Ti result in an increase in cell

adhesion density. The highest cell adhesion density among the

alloys is that of TNZ-4, which also has the highest Zr (40.7%) and

the lowest Nb (24.8%) and Ti (35.5%). The three elements of Ti,

Nb, and Zr are all biocompatible elements, as previously reported

[87–89].Cells were seeded in the wells having only the medium within,

as well. As a negative control, they were incubated for the same

days. The cell viability ratios (CVR) of CP-Ti and the TNZ alloys

were calculated by dividing the live SaOS2 cell number inside the

experimental wells by the live SaOS2 cell number inside the con-

trol; the control group has a CVR of 1 and is accepted as being

cytocompatible as a reference in determining the cell viability

ratios of samples [85]. Fig. 14 presents the CVR values that enabled

assessment of the in vitro cytotoxicity of the TNZ alloys and CP-Ti.

The CVR values of the TNZ alloys were all greater than those of the

control group; hence it can be deduced that the alloys possess

exceptional cytocompatibility. In addition, the cell proliferation

ratios were similar in the experimental wells of the CP-Ti and

TNZ disks, so their cell viability ratio values are also very close(Fig. 14).

Fig. 10.  Fracture surfaces of as-cast TNZ alloys (a) TNZ-1; (b) TNZ-2; (c) TNZ-3; (d) TNZ-4.

Fig. 11.   Morphology of SaOS2 cells attached on the surface of TNZ-2 alloy after cell

culture for 24 h.

184   S. Ozan et al. / Acta Biomaterialia 20 (2015) 176–187 

Page 10

7/21/2019 Articulo. Development of Ti–Nb–Zr Alloys With High Elastic Admissible Strain

http://slidepdf.com/reader/full/articulo-development-of-tinbzr-alloys-with-high-elastic-admissible-strain 10/12

4. Conclusions

The microstructures, mechanical properties and cytocompati-

bility of the as-cast Ti–Nb–Zr alloys have been examined. The con-

clusions are as follows:

1. Microstructural examinations revealed that the TNZ alloys dis-

play primary grain boundaries together with secondary grain

structure.

2. The average compressive yield strength and microhardness

values of the TNZ alloys exhibited the same ranking: TNZ-

2 > TNZ-3 > TNZ-1 > TNZ-4. The alloys exhibited considerable

ductility under compression and no fracturing or cracking was

observed on the samples after the compression tests. The com-

pressive yield strength of thealloys rangedfrom713 to 751 MPa.

3. Tensile test results showed that the TNZ alloys have elongation

at rupture values in the range of 9.9–14.8%, which is highly

desirable for biomedical applications. The tensile strength of 

the TNZ alloys was in the range of 704–839 MPa and Young’s

modulus was in the range of 62–65 GPa.

4. The TNZ alloys exhibited elastic admissible strain values in the

range of 1.08–1.31%. Ti–34Nb–25Zr exhibited an elastic admis-

sible strain of 1.31%, the highest value compared to TNZ alloys

studied in this project.

Fig. 12.  SEM images of SaOS2 cells attached on to surfaces of TNZ alloys after cell culture for 7 days: (a) TNZ-1; (b) TNZ-2; (c) TNZ-3; (d) TNZ-4.

Fig. 13.  SaOS2 cell adhesion density of as-cast TNZ alloys and CP-Ti after cell culture for 7 days.

S. Ozan et al./ Acta Biomaterialia 20 (2015) 176–187    185

Page 11

7/21/2019 Articulo. Development of Ti–Nb–Zr Alloys With High Elastic Admissible Strain

http://slidepdf.com/reader/full/articulo-development-of-tinbzr-alloys-with-high-elastic-admissible-strain 11/12

5. The tensile strength, yield strength, elongation, elastic energy,

and toughness of the TNZ alloys all decreased when the Zr con-

tent increased and the amounts of Ti and Nb decreased. In other

words, the tensile mechanical properties decreased with a

decrease in the (e/a)ratio and Moeq  values.

6. The TNZ alloys exhibited excellent cytocompatibility on SaOS2

cells.

7. The TNZ alloys designed in this study are suitable candidates for

removable orthopedic implant applications due to their high Zr

content, high elastic admissible strain value, superior plastic

deformation ability during compression test, and excellent

cytocompatibility.

 Acknowledgments

SO would like to acknowledge the Council of Higher Education

(CoHE) of Turkey for PhD research scholarship. CW acknowledges

the financial support from the Australian Research Council (ARC)

through the ARC Discovery Project DP110101974.

 Appendix A. Figures with essential color discrimination

Certain figures in this article, particularly Figs. 5 and 7–9 are dif-

ficult to interpret in black and white. The full color images can be

found in the on-line version, at http://dx.doi.org/10.1016/j.actbio.

2015.03.023.

References

[1] Division UNDoEaSAP. World population prospects: the 2012 revision, Key

findings and advance tables. New York; 2013. p. 1-50.

[2] Navarro M, Michiardi A, Castano O, Planell JA. Biomaterials in orthopaedics. J R 

Soc Interface/R Soc 2008;5:1137–58.

[3]  Geetha M, Singh AK, Asokamani R, Gogia AK. Ti based biomaterials, the

ultimate choice for orthopaedic implants – a review. Prog Mater Sci

2009;54:397–425.

[4] Abdel-Hady Gepreel M, Niinomi M. Biocompatibility of Ti-alloys for long-termimplantation. J Mech Behav Biomed Mater 2013;20:407–15.

[5] Biesiekierski A, Wang J, Gepreel MA, Wen C. A new look at biomedical Ti-based

shape memory alloys. Acta Biomater 2012;8:1661–9.

[6]  Huiskes R, Weinans H, Vanrietbergen B. The relationship between stress

shielding and bone-resorption around total hip stems and the effects of flexible

materials. Clin Orthop Relat Res 1992:124–34.

[7] Black J. Biological performance of materials, fundamentals of biocompatibility.

4th ed. Boca Raton, FL: CRC Taylor & Francis; 2006.

[8]  Niinomi M. Mechanical propertiesof biomedical titanium alloys. Mater Sci Eng

1998;243:231–6.

[9] Rho JY, Tsui TY, Pharr GM. Elastic properties of human cortical and trabecular

lamellar bone measured by nanoindentation. Biomaterials 1997;18:1325–30.

[10]  Hao YL, Li SJ, Sun SY, Zheng CY, Yang R. Elastic deformation behaviour of Ti–

24Nb–4Zr–7.9Sn for biomedical applications. Acta Biomater 2007;3:277–86.

[11]  Kuroda D, Niinomi M, Morinaga M, Kato Y, Yashiro T. Design and mechanical

properties of new beta type titanium alloys for implant materials. Mater Sci

Eng 1998;243:244–9.

[12]  Li Q, Niinomi M, Nakai M, Cui Z, Zhu S, Yang X. Effect of Zr on super-elasticity

and mechanical properties of Ti–24 at% Nb–(0, 2, 4) at% Zr alloy subjected to

aging treatment. Mater Sci Eng, A 2012;536:197–206.

[13]   Matsumoto H, Watanabe S, Hanada S. Beta TiNbSn alloys with low Young’s

modulus and high strength. Mater Trans 2005;46:1070–8.

[14]  Miyazaki S, Kim HY, Hosoda H. Development and characterization of Ni-free

Ti-base shape memory and superelastic alloys. Mater Sci Eng, A 2006;438–

440:18–24.

[15]   Ozaki T, Matsumoto H, Watanabe S, Hanada S. Beta Ti alloys with low Young’s

modulus. Mater Trans 2004;45:2776–9.

[16]   Yang G, Zhang T. Phase transformation and mechanical properties of the

Ti50Zr30Nb10Ta10 alloy with low modulus and biocompatible. J Alloy Compd2005;392:291–4.

[17]   Okazaki Y. A new Ti-15Zr-4Nb-4Ta alloy for medical applications. Curr Opin

Solid State Mater Sci 2001;5:45–53.

[18]   Flaten TP. Aluminium as a risk factor in Alzheimer’s disease, with emphasis on

drinking water. Brain Res Bull 2001;55:187–96.

[19]  Lima PD, Vasconcellos MC, Montenegro RC, Bahia MO, Costa ET, Antunes LM,

et al. Genotoxic effects of aluminum, iron and manganese in human cells and

experimental systems: a review of the literature. Hum Exp Toxicol

2011;30:1435–44.

[20]   Boyer R, Welsch G, Collings EW, editors. Materials properties handbook:

titanium alloys. Materials Park, OH, USA: ASM International; 1994.

[21]   Anselme K. Biomaterials and interface with bone. Osteoporosis Int

2011;22:2037–42.

[22]   Liu XY, Huang AP, Ding CX, Chu PK. Bioactivity and cytocompatibility of 

zirconia (ZrO2) films fabricated by cathodic arc deposition. Biomaterials

2006;27:3904–11.

[23]   Yuan H, de Groot K. Calcium phosphate biomaterials: an overview. In: Reis RL,

Weiner S, editors. Learning from nature how to design new implantable

biomaterialsis: from biomineralization fundamentals to biomimetic materialsand processing routes. Netherlands: Springer; 2005. p. 37–57.

Fig. 14.  Cell viability ratio of as-cast TNZ alloys and CP-Ti.

186   S. Ozan et al. / Acta Biomaterialia 20 (2015) 176–187 

Page 12

7/21/2019 Articulo. Development of Ti–Nb–Zr Alloys With High Elastic Admissible Strain

http://slidepdf.com/reader/full/articulo-development-of-tinbzr-alloys-with-high-elastic-admissible-strain 12/12

[24]   Carlsson L, Röstlund T, Albrektsson B, Albrektsson T, Brånemark P-I.

Osseointegration of titanium implants. Acta Orthopaedica 1986;57:285–9.

[25]  Liu XY, Chu PK, Ding CX. Surface modification of titanium, titanium alloys, and

related materials for biomedical applications. Mater Sci Eng R Rep

2004;47:49–121.

[26]   Skripitz R, Aspenberg P. Tensile bond between bone and titanium – a

reappraisal of osseointegration. Acta Orthop Scand 1998;69:315–9.

[27]  Kobayashi E, Ando M, Tsutsumi Y, Doi H, Yoneyama T, Kobayashi M, et al.

Inhibition effect of zirconium coating on calcium phosphate precipitation of 

titanium to avoid assimilation with bone. Mater Trans 2007;48:301–6.

[28]  Kim HM, Miyaji F, Kokubo T, Nakamura T. Effect of heat treatment on apatite-forming ability of Ti metal induced by alkali treatment. J Mater Sci Mater Med

1997;8:341–7.

[29]  Tsutsumi Y, Niinomi M, Nakai M, Tsutsumi H, Doi H, Nomura N, et al. Micro-

arc oxidation treatment to improve the hard-tissue compatibility of Ti–29Nb–

13Ta–4.6Zr alloy. Appl Surf Sci 2012;262:34–8.

[30]  Li YC, Xiong JY, Wong CS, Hodgson PD, Wen CE. Bioactivating the surfaces of 

titanium by Sol–Gel process. Mater Sci Forum 2009;614:67–71.

[31]  Hanawa T. Research and development of metals for medical devices based on

clinical needs. Sci Technol Adv Mater 2012;13:064102.

[32]   Simka W, Krzakala A, Maselbas M, Dercz G, Szade J, Winiarski A, et al.

Formation of bioactive coatings on Ti–13Nb–13Zr alloy for hard tissue

implants. RSC Adv 2013;3:11195–204.

[33]   Fujibayashi S, Nakamura T. Current status of bioactive coatings in Japan.

Fifteen years of clinical experience with hydroxyapatite coatings in joint

arthroplasty. Paris: Springer; 2004. p. 115–24.

[34]  Zhao X, Niinomi M, Nakai M, Ishimoto T, Nakano T. Development of high Zr-

containing Ti-based alloys with low Young’s modulus for use in removable

implants. Mater Sci Eng, C 2011;31:1436–44.

[35] Niinomi M. Low modulus titanium alloys for inhibiting bone atrophy.

Biomater Sci Eng 2011. Prof. Rosario Pignatello (Ed.), ISBN: 978-953-307-

609-6, InTech, 2011. Available from: <http://www.intechopen.com/books/

biomaterials-science-and-engineering/low-modulus-titanium-alloys-

forinhibiting-bone-atrophy >.

[36]  Sidky A, Buckley RE. Hardware removal after tibial fracture has healed. Can J

Surg 2008;51:263–8.

[37] HanawaT, TsutsumiY. Calciumphosphate formationon titanium andzirconium

and its application to medical devices. Bioceramics Dev Appl 2010;1:1–4.

[38]  Vos DI, Verhofstad MHJ. Indications for implant removal after fracture healing:

a review of the literature. Eur J Trauma Emerg Surg 2013;39:327–37 .

[39]   Evans NA, Evans RON. Playing withmetal: fracture implants and contact sport.

Br J Sports Med 1997;31:319–21.

[40]   Schmittenbecher PP. Implant removal in children. Eur J Trauma Emerg Surg

2013;39:345–52.

[41]   Zhao X, Niinomi M, Nakai M, Miyamoto G, Furuhara T. Microstructures and

mechanical properties of metastable Ti–30Zr–(Cr, Mo) alloys with changeable

Young’s modulus for spinal fixation applications. Acta Biomater

2011;7:3230–6.[42]   Niinomi M, Nakai M. Titanium-based biomaterials for preventing stress

shielding between implant devices and bone. Int J Biomater 2011;2011:

836587.

[43]  Hanawa T, Okuno O, Hamanaka H. Compositional change in surface of Ti–Zr

alloys in artificial bioliquid. J Jpn Inst Met 1992;56:1168–73.

[44]   Hanawa T, Hiromoto S, Asami K, Okuno O, Asaoka K. Surface oxide films on

titanium alloys regenerated in Hanks’ solution. Mater Trans 2002;43:3000–4.

[45]   Abdel-Hady M, Fuwa H, Hinoshita K, Kimura H, Shinzato Y, Morinaga M. Phase

stability change with Zr content in beta-type Ti–Nb alloys. Scripta Mater

2007;57:1000–3.

[46]  You L, Song X. First principles; study of low Young’s modulus Ti–Nb–Zr alloy

system. Mater Lett 2012;80:165–7.

[47]   Henry S, Dragolich K, DiMatteo N. Fatigue data book: light structural

alloys. Materials Park, OH: ASM International; 1995.

[48]  Abdel-Hady M, Hinoshita K, Morinaga M. General approach to phase stability

and elastic properties of  b-type Ti-alloys using electronic parameters. Scripta

Mater 2006;55:477–80.

[49]  Laheurte P, Prima F, Eberhardt A, Gloriant T, Wary M, Patoor E. Mechanical

properties of low modulus beta titanium alloys designed from the electronicapproach. J Mech Behav Biomed Mater 2010;3:565–73.

[50] Morinaga M, Yukawa N, Maya T, Sone K, Adachi H. Theoretical design

of titanium alloys. In: Sixthworld conference on titanium, vol. 1. France; 1988.

p. 1601–6.

[51]   Matsugi K, EndoT, Choi Y-B, SasakiG. Alloydesignof Ti alloys usingubiquitous

alloying elements and characteristics of their levitation-melted alloys. Mater

Trans 2010;51:740–8.

[52]  Zhang DC, Lin JG, Jiang WJ, Ma M, Peng ZG. Shape memory and superelastic

behavior of Ti–7.5Nb–4Mo–1Sn alloy. Mater Des 2011;32:4614–7.

[53]   Saito T, Furuta T, Hwang JH, Kuramoto S, Nishino K, Suzuki N, et al.

Multifunctional alloys obtained via a dislocation-free plastic deformation

mechanism. Science 2003;300:464–7.

[54] MorinagaM, Yukawa H. Alloydesign with theaid of molecular orbital method.

Bull Mater Sci 1997;20:805–15.

[55]  Bania PJ. Beta-titanium alloys and their role in the titanium industry. JOM J

Miner Met Mater Soc 1994;46:16–9.

[56] ASTM E384-11e1. Standard test method for Knoop and Vickers hardness of 

materials. WestConshohocken, PA: ASTMInternational; 2011.<www.astm.org>.

[57] ASTM E9-09. Standard test methods of compression testing of metallic

materials at room temperature. West Conshohocken, PA: ASTM

International; 2009. <www.astm.org>.

[58]   Alshehri O, Yavuz M, Tsui T. Manifestation of external size reductioneffects on

the yield point of nanocrystalline rhodium using nanopillars approach. Acta

Mater 2013;61:40–50.

[59]   Biological evaluation of medical devices. Part 5: Test for in vitro

cytotoxicity. Arlington, VA: International Organization for Standardization,

ANSI/AAMI; 1999.

[60]  Li Y, Xiong J, Hodgson PD, Wen Ce. Effects of structural property and surface

modification of Ti6Ta4Sn scaffolds on the response of SaOS2 cells for bonetissue engineering. J Alloy Compd 2010;494:323–9.

[61]  Li Y, Xiong J, Wong CS, Hodgson PD, Wen Ce. Ti6Ta4Sn alloy and subsequent

scaffolding for bone tissue engineering. Tissue Eng Part A 2009;15:3151–9.

[62]   Ning C, Ding D, Dai K, Zhai W, Chen L. The effect of Zr content on the

microstructure, mechanical properties and cell attachment of Ti–35Nb–xZr

alloys. Biomed Mater 2010;5:045006.

[63]   Bertrand E, Gloriant T, Gordin DM, Vasilescu E, Drob P, Vasilescu C, et al.

Synthesis and characterisation of a new superelastic Ti–25Ta–25Nb

biomedical alloy. J Mech Behav Biomed Mater 2010;3:559–64.

[64]   Ushida K, Tsuge K, Akahori T, Hattori T, Niinomi M, Ishikura K, et al.

Mechanical strength and biocompatibility of meta-stable beta type Ti–5Fe–

3Nb–3Zr for biomedical applications. J Jpn Inst Met 2012;76:397–401.

[65]   Geetha M, Singh AK, Muraleedharan K, Gogia AK, Asokamani R. Effect of 

thermomechanical processing on microstructure of a Ti–13Nb–13Zr alloy. J

Alloy Compd 2001;329:264–71.

[66]  Ho WF, Ju CP, Lin JHC. Structure and properties of cast binary Ti–Mo alloys.

Biomaterials 1999;20:2115–22.

[67]   Yang Y, Wu SQ, Li GP, Li YL, Lu YF, Yang K, et al. Evolution of deformation

mechanisms of Ti–22.4Nb–0.73Ta–2Zr–1.34O alloy during straining. Acta

Mater 2010;58:2778–87.

[68]   Li C, Zhan Y, Jiang W. Beta-type Ti–Mo–Si ternary alloys designed for

biomedical applications. Mater Des 2012;34:479–82.

[69]  Nie L, Zhan Y, Liu H, Tang C. Novel  b-type Zr–Mo–Ti alloys for biological hard

tissue replacements. Mater Des 2014;53:8–12.

[70]   Currey JD. Bones: structure and mechanics. Princeton: Princeton University

Press; 2002.

[71]  Wang X, Li Y, Xiong J, Hodgson PD, Wen Ce. Porous TiNbZr alloy scaffolds for

biomedical applications. Acta Biomater 2009;5:3616–24.

[72]   Hutmacher DW. Scaffolds in tissue engineering bone and cartilage.

Biomaterials 2000;21:2529–43.

[73] Ryan G, PanditA, Apatsidis DP. Fabricationmethods ofporous metals foruse in

orthopaedic applications. Biomaterials 2006;27:2651–70.

[74]  Dabrowski B, Swieszkowski W, Godlinski D, Kurzydlowski KJ. Highly porous

titanium scaffolds for orthopaedic applications. J Biomed Mater Res B Appl

Biomater 2010;95:53–61.

[75] ASTMF1813-13. Standardspecification for wrought titanium-12molybdenum-6

zirconium-2 iron alloy for surgical implant (UNS R58120). West Conshohocken,PA: ASTM International; 2013. <www.astm.org>.

[76] ASTM F1295-05. Standard specification for wrought titanium-6 aluminum-7

niobium alloy for surgical implant applications (UNS R56700). West

Conshohocken, PA: ASTM International; 2005. <www.astm.org>.

[77] ASTM F2066-08. Standard specification for wrought titanium-15 molybdenum

alloy for surgical implant applications (UNS R58150). West Conshohocken,

PA: ASTM International; 2008. <www.astm.org>.

[78] ASTM F1713-08. Standard specification for wrought titanium-13niobium-

13zirconium alloy for surgical implant applications (UNS R58130). West

Conshohocken, PA: ASTM International; 2013. <www.astm.org>.

[79] ASTM F1108-04. Standard specification for titanium-6aluminum-4vanadium

alloy castings for surgical implants (UNS R56406). West Conshohocken,

PA: ASTM International; 2004. <www.astm.org>.

[80]  Callister WD. Materials science and engineering: an introduction. 7th ed. New

York: John Wiley and Sons; 2007.

[81]  Akahori T, Niinomi M, Koyanagi Y, Kasuga T, Toda H, Fukui H, et al. Mechanical

properties of biocompatible beta-type titanium alloy coated with calcium

phosphate invert glass-ceramic layer. Mater Trans 2005;46:1564–9.

[82]  Song Y, Xu DS, Yang R, Li D, Wu WT, Guo ZX. Theoretical study of the effects of alloying elements on the strength and modulus of beta-type bio-titanium

alloys. Mater Sci Eng 1999;260:269–74.

[83]   Zhang J, Sun F, Hao Y, Gozdecki N, Lebrun E, Vermaut P, et al. Influence of 

equiatomic Zr/Nb substitution on superelastic behavior of Ti–Nb–Zr alloy.

Mater Sci Eng, A 2013;563:78–85.

[84]   Wang K. The use of titaniumfor medical applications in theUSA. Mater SciEng

1996;213:134–7.

[85]  Li Y, Wen C, Mushahary D, Sravanthi R, Harishankar N, Pande G, et al. Mg–Zr–

Sr alloys as biodegradable implant materials. Acta Biomater 2012;8:3177–88.

[86]  Li Y, WongC, XiongJ, Hodgson P, Wen C. Cytotoxicity of titanium and titanium

alloying elements. J Dent Res 2010;89:493–7.

[87]   Eisenbarth E, Velten D, Muller M, Thull R, Breme J. Biocompatibility of beta-

stabilizing elements of titanium alloys. Biomaterials 2004;25:5705–13.

[88]  Geurtsen W. Biocompatibility of dental casting alloys. Crit Rev Oral Biol Med

2002;13:71–84.

[89]   Yamamoto A, Honma R, Sumita M. Cytotoxicity evaluation of 43 metal salts

using murine fibroblasts and osteoblastic cells. J Biomed Mater Res

1998;39:331–40.

S. Ozan et al./ Acta Biomaterialia 20 (2015) 176–187    187