A review on biomedical titanium alloys Recent progress and prospect

REVIEW
Titanium Alloys
www.aem-journal.com
A Review on Biomedical Titanium Alloys: Recent
Progress and Prospect
Lai-Chang Zhang* and Liang-Yu Chen*
stents, heart valves, cardiac simulator, and
replacement implants in knees, hips,
elbows, shoulders, ears, and dentistry.[1,5]
Among these parts, implants employed for
hip and knee joint replacements have a
great deal of requirements since osteoarthritis (inflammation in the bone joints),
osteoporosis (weakening of the bones), and
trauma usually lead to pain or loss in
function of the tissues.[2] These deteriorating diseases result in gradual degradation
in the mechanical properties of the bones
because of inordinate loading or loss of the
function for the self-healing process. It has
been reported that the population of elderly
people remarkably increases in the United
States and less developed countries after
2010 (from 4.9 million in 2002 to 39.7
million in 2010) and most of them over
the age of 40 may have such types of
degenerative diseases.[1,6] Therefore, using
biomedical artificial implants to recover the
function of diseased hard tissues by
surgical implantation is the best solution for patients. The data
on the total replacement surgery for joints in the United States
reveal that by the end of 2030, the total number of replacement
surgery will increase by 174% (0.57 million procedures) for hip
replacement and by 673% (3.48 million procedures) for knee
arthroplasties from the current rate.[7] The reason for the growth
in joint replacement is mainly ascribed to the increase in both
replacement surgeries and revision surgeries. Unfortunately, the
success rates of the revision surgeries are rather low as compared
to the first implantation. The revision surgery may cause
tremendous pain to the patients. The main reasons for the
implant failure can be ascribed to debris generation, metal ions
release and foreign body response in the human body, mismatch
of modulus between the bones and the implants or too low
strength to maintain the load.[3] As such, biomaterial implants
employed for load-bearing applications must have many reliable
properties, including appropriate mechanical properties, excellent corrosion resistance, high wear resistance, excellent
biocompatibility, osseointegration, and non-cytotoxicity to avoid
the revision surgery.[8]
One of the most important things for designing and
selecting the biomedical materials is their medical intentions.
Once a biomedical implant is placed, a considerable number of
reactions would take place at the interface of the host tissue and
the implant. These reactions determine the biocompatibility of
an implant, which also determines the success of the
implantation.[9] To be biocompatible, toxic elements are not
Compared with stainless steel and Co–Cr-based alloys, Ti and its alloys are
widely used as biomedical implants due to many fascinating properties, such
as superior mechanical properties, strong corrosion resistance, and excellent
biocompatibility. After briefly introducing several most commonly used
biomedical materials, this article reviews the recent development in Ti alloys
and their biomedical applications, especially the low-modulus β-type Ti alloys
and their design methods. This review also systemically investigates the
recently attractive progress in preparation of biomedical Ti alloys, including
additive manufacturing, porous powder metallurgy, and severe plastic
deformation, applied in the manufacturing and the influenced microstructures
and properties. Nevertheless, there are still some problems with the
long-term performance of Ti alloys, and therefore several surface modification
methods are reviewed to further improve their biological activity, wear
resistance, and corrosion resistance. Finally, the biocompatibility of Ti and its
alloys is concluded. Summarizing the findings from literature, future
prediction is also conducted.
1. Introduction
Biomedical materials, which are made of natural or artificial
materials, are commonly used as structures and implants of
human beings with the purpose of replacing their lost
or diseased biological structure and improving the quality of
life.[1–3] With growing aging population problems, the aged
people may suffer from diseases such as arthritis and joint pain,
and therefore there is a great increase in the demand for
replacing dysfunctional hard tissues by artificial instruments
made of biomedical materials.[4] Therefore, biomedical materials
have gained substantially increasing attention in recent decades.
According to the specific application, biomedical materials can
be used in various parts in the human body as intravascular
Prof. L. C. Zhang
School of Engineering
Edith Cowan University
270 Joondalup Drive, Joondalup, Perth, WA 6027, Australia
E-mail: [email protected]; [email protected]
Dr. L. Y. Chen
School of Science
Jiangsu University of Science and Technology
Mengxi Road 2#, Zhenjiang 212003, P. R. China [email protected]
The ORCID identification number(s) for the author(s) of this article
can be found under https://doi.org/10.1002/adem.201801215.
DOI: 10.1002/adem.201801215
Adv. Eng. Mater. 2019, 21, 1801215
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expected to be used in biomedical materials.[10] According to
previous studies, elements such as Ti, Nb, Mo, Ta, Zr, Au, W,
and Sn are highly biocompatible, while Al, V, Cr, Ni, etc. are
considered as hazardous elements for human body.[11,12] The
detailed information on pure metals and their biocompatibility
is shown in Figure 1. As a result, the safe and biocompatible
elements are more preferred to be considered when designing
biomedical materials.[10] Meanwhile, the mechanical properties
of biomedical materials determine the implants for specific
applications. As biomedical materials, their elastic modulus,
fatigue strength, tensile strength, and elongation are highly
important for load-bearing hard tissue replacements. For
instance, the strength of implant materials decides their loadbearing capability.[13] Besides fatigue and fracture failure, an
implant is required to have a comparable modulus to the
replaced bone.[14] Table 1 shows the tensile properties of human
cortical bone. Table 2 and 3 list the compressive properties of
human cortical bone and cancellous bone. Normally, if an
implant has a greater modulus than the replaced bone, stress
shielding effect, which prohibits the transfer of needed stress to
the adjacent bone, would lead to bone resorption around the
implant and the death of bone cells.[15] Therefore, an ideal
biomedical material is expected to have the combined
properties of low modulus (close to the bone) and high
strength to sustain a long-term service period and to eliminate
the revision surgery. Furthermore, the service period of the
implant is determined by its wear and corrosion resistance.[16]
The low wear and corrosion resistance of the implants in the
human body lead to the wear debris and/or releases the
Lai-Chang Zhang is a Professor of
Materials Engineering and the
Program Leader–Mechanical
Engineering in the School of
Engineering at Edith Cowan
University. After receiving his PhD in
Materials Science and Engineering
at the Institute of Metal Research,
Chinese Academy of Sciences, Prof.
Zhang held several positions at The
University of Western Australian, University of
Wollongong, IFW Dresden and Technische Universität
Darmstadt. His current research interest is focusing on
the metal additive manufacturing (e.g. selective laser
melting, electron beam melting), titanium alloys and
composites, and processing-microstructure-properties in
high-performance materials.
Liang-Yu Chen is an Associate
Professor in the School of Science at
Jiangsu University of Science and
Technology. He received his PhD
degree from Shanghai Jiaotong
University in 2016. His current
research interest focuses on the
biomedical titanium alloys, plasma
sprayed coatings and corrosion
science of metallic materials.
incompatible metal ions, which may cause allergic and toxic
reactions.[16] Therefore, the designed biomedical materials are
also expected to have both high corrosion and wear resistance,
aiming to maximize the longevity of the implants in the human
Table 1. Tensile properties of human cortical bones.
Figure 1. Pure metals and their biocompatibility. (Reproduced with
permission from ref. [11]. Copyright (1998), Elsevier).
Adv. Eng. Mater. 2019, 21, 1801215
Bone
Age
N
n
σUTS [MPa]
E [GPa]
emax [%]
ρ [g cm3]
Fibula
41.5
17
20
100
19.2
2.10
1.91
[59]
Fibula
71
17
16
80
15.2
1.19
1.73
[59]
Humerus
15–89
64
27
149
15.6
2.20
1.77
[60]
Humerus 15–89
64
16
151
16.1
1.90
1.72
[60]
Tibia
41.5
17
67
106
18.9
1.76
1.96
[59]
Tibia
71
17
34
84
16.2
1.56
1.83
[59]
Tibia
20–89
28
123
156
23.8
3.09
–
[61]
Femur
41.5
17
35
102
14.9
1.32
1.91
[59]
Femur
71
17
35
68
13.6
1.07
1.85
[59]
Femur
15–89
64
29
141
15.2
2.00
1.90
[60]
Femur 15–89
64
30
134
15.0
1.80
1.80
[60]
Femur
20–89
33
178
132
16.8
2.83
–
[61]
ref.
N is the number of bones tested and n is the number of samples obtained from the
tested bones; σUTS is ultimate tensile strength; E is elastic modulus; emax is ultimate
tensile strain, and ρ is density; means male; indicates female; 41.5 and 71
are the average age value from 36 to 75 years of age, 41.5 is a younger group, 71 is an
older group.
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Table 2. Compressive properties of human cortical bones.[12]
σUTS
Age
N
n
[MPa]
E
[GPa]
H
[MPa]
Tibia osteons (longitudinal)
57/61
2
2
–
22.5
614
[62]
Tibia interstitial
Lamellae (longitudinal)
57/61
2
2
–
25.8
736
[62]
Femur
20–89
19
95
194
17.6
–
[61]
Tibia
30–89
11
38
195
28.0
–
[61]
Tissue source
ref.
N is the number of bones tested and n is the number of samples obtained from
them; σUTS is ultimate tensile strength; E is elastic modulus; and H is hardness.
body. Lastly, high biocompatibility (such as osseointegration
and antibacterial) is the most significant consideration for
using implants in orthopaedic applications, which primarily
means the formation of a reliable bonding between the implant
and the adjacent bone without an intervening fibrous tissue and
infection.[17,18] A successful osseointegration is considered as
being no relative motion between the implant and the adjacent
bone during functional loading.[19] Hence, it is highly essential
that biomedical implants are expected to have an appropriate
surface to integrate well with the adjacent bone. As such,
surface topography, surface chemical compositions as well as
surface coating are vital factors for development of implants
with good biocompatibility.
Based on the above consideration, metals and alloys have
been used as load-bearing implants for a long period.[2,20–22]
Until now, 316L stainless steel, Co–Cr-based alloys, Ti, and its
alloys are most employed as metallic implant materials.[4,23–26]
The use of stainless steels as implants began in the 1920s.[27]
The stainless steels contain at least 10.5 wt% Cr to prevent the
formation of rust in unpolluted atmospheres.[28] Among
stainless steels, 316L stainless steel (18Cr–14Ni–2.5Mo, in wt
%, the same hereafter unless indicated) are most commonly
used for temporary implants and in some cases as a biomaterial
for permanent implants.[29] The “L” in 316L stainless steel
indicates low carbon, which can prevent the formation of
chromium carbides and enhance the resistance to pitting and
crevice corrosion.[30] However, stress corrosion cracking, which
cannot be prevented in 316L stainless steel, can be triggered by
the combined effect of tensile stress and a Cl-rich environment
(i.e., human body fluid), thereby resulting in an unexpected
sudden failure of the implant under the stresses much lower
than the strength of the alloy.[31] By comparison, Co–Cr-based
alloys have a better corrosion resistance in the Cl rich
environment than 316L stainless steel. The high Cr contents in
Co–Cr-based alloys results in the formation of a passive oxide
(Cr2O3) film spontaneously on their surfaces in the human body
environment.[32] Meanwhile, the mechanical properties of Co–
Cr-based alloys meet the demand criteria of implants.[32] These
alloys have often been used for sliding parts of artificial joints.[2]
Unfortunately, due to corrosion and wear, some unfriendly
elements such as Ni, Co, and Cr are found to be released from
the 316L stainless steel and Co–Cr-based alloys in the human
body.[33] It has been reported that the released Ni even caused
some skin related diseases such as dermatitis, and the presence
of Co showed carcinogenicity in numerous animal studies and
led to the neurological symptoms in some patients after some
years of implantation.[34,35] The released Cr would damage the
kidney, liver and blood cells via oxidation reactions.[36] As such,
316L stainless steels and Co–Cr-based alloys have potential
hazards when they are employed as biomedical implants.
Furthermore, both 316L stainless steels and Co–Cr-based alloys
exhibit much higher elastic moduli than bones in the human
body, which may lead to the stress shielding effect and
loosening of the implant after a long term of implantation.[13]
Therefore, these two alloys may not be the best choice for
orthopaedic implants. Ti has received extensive attention since
the 1950s due to its high specific strength (e.g. the ratio of
strength to density), excellent corrosion resistance, and
superior biocompatibility.[16,37] During the development of Ti,
most Ti products, such as commercially pure Ti (CP–Ti),[13,38,39]
Ti alloys,[40–44] and Ti-based composites,[45–54] have been
designed for the industries of aerospace, military, automobiles,
Table 3. Compressive properties of human cancellous bones.[12]
Age
N
n
σUTS [MPa]
E [GPa]
emax [%]
ρ [g cm3]
Lumbar vertebra (male)
14–89
60
32
4.60
0.06
6.70
0.20
[63]
Lumbar vertebra (female)
14–89
60
32
2.70
0.04
6.10
0.20
[63]
Tibia head (male)
14–89
60
32
3.90
0.03
8.30
0.22
[63]
Tibia head (female)
14–89
60
32
2.20
0.02
6.90
0.22
[63]
Tibia
16–83
31
395/183
8.80
0.64
2.20
0.45
[64]
Vertebra (vertical)
15–87
42
84
2.5
0.07
7.40
–
[64]
Vertebra (horizontal)
15–87
42
84
0.9
0.02
8.50
–
[64]
Proximal tibia
59–82
9
121
5.33
0.45
–
0.29
[12]
Femur
58–83
10
299
7.36
0.39
–
0.50
[12]
Lumbar spine
15–87
42
40
2.45
0.07
–
0.24
[12]
Lumbar spine
71–84
3
231
1.55
0.02
–
0.19
[12]
Tissue source
ref.
N is the number of bones tested and n is the number of samples obtained from them; 395 samples for elastic modulus, 183 samples for ultimate strength and ultimate
strain; σUTS is ultimate tensile strength; E is elastic modulus; emax is ultimate tensile strain; and ρ is density.
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medicine, jewelry, and mobile phones. Most importantly, due to
the low density, high strength, high corrosion resistance,
complete inertness to body environment, superior biocompatibility, low elastic modulus, high capacity to join with bone and
other tissues, Ti and its alloys are regarded as the highly
desirable choice for orthopaedic implants.[1,54,55] Figure 2
shows the elastic moduli of some biomedical materials and
bone. Apparently, by comparing the data in Table 1–3, the elastic
moduli of Ti alloys are significantly closer to those of bones. The
density of Ti alloys (4.4–4.5 g cm3) is significantly less than
those of Co–Cr-based alloys (8.3–9.2 g cm3) and 316L stainless
steel (7.9 g cm3).[56] However, the strength of the Ti alloys can
be as high as 316L stainless steel.[13] Therefore, Ti alloys are
superior to other metallic implant biomaterials by comparing
their specific strength (i.e. the ratio between yield strength and
density). In the human body environment, a passive film
(mainly consisted of TiO2) can spontaneously form on the
surface of Ti alloy implant to protect the underlying substrate
from further corrosion.[57] Once the oxide film of the Ti alloy
implant is broken and exposes the substrate to the corrosive
environment, a new passive film would be rebuilt rapidly.[58]
Therefore, Ti alloys are more tolerant than 316L stainless steels
and Co–Cr-based alloys. As aforementioned, Ti is non-toxic for
the human body and has high biocompatibility.[36] Combined
these advantages, Ti and its alloys provide a better solution to
the problems of biomedical implants in the human body. In this
review, a brief introduction to biomedical Ti alloys is first
presented. Afterward, the development of biomedical Ti alloys
is reviewed in order to understand different types of Ti alloys
(i.e. α, near-α, (α þ β), β and shape-memory alloys) and their
corresponding microstructures and mechanical properties.
Third, new preparation methods of Ti and its alloys, such as
additive manufacturing, porous powder metallurgy, and severe
plastic deformation methods for biomedical applications, are
reviewed. Fourth, the surface modification of Ti alloys for
biomedical applications is introduced to further improve their
biological activity, wear resistance, and corrosion resistance.
Finally, the biocompatibility of Ti and Ti alloys is concluded.
2. Alloy Development and Mechanical
Behavior
Basically, pure Ti has a hexagonal close-packed (hcp) structure at
room temperature, called α phase.[65] A phase transformation
from hcp structure to body-centered cubic (bcc) structure would
take place for pure Ti at 883 C (allotropic temperature), namely
β phase.[66] Many previous investigations show that the phase
transformation temperature of pure Ti can be strongly
influenced by alloying elements.[67] In fact, alloying elements
are usually classified into α-stabilizers (such as Al, C, O) and
β-stabilizers (such as Mo, Ta, Nb). α-stabilizers increase the
temperature of β/α transus while β-stabilizers reduce the β/α
transus temperature.[68] Generally, β-stabilizers are transition
elements in the Periodic Table. The strength of added βstabilizing transition elements depends on additive amounts. By
tailoring the type and amount of alloying elements, β phase can
be retained at room temperature. As such, conventional Ti alloys
can be primarily classified into four groups of α-, near-α, (α þ β)-,
and β-type alloys. Meanwhile, Ti-based shape-memory alloys is
another group of Ti alloy with superelasticity and shape memory
effect.
2.1. α-Type and Near-α Ti Alloys
α-type Ti alloys, which only have α phase, consist of various grades
of CP–Ti and Ti alloys.[69] Generally, CP–Ti has four grades which
contain 0.20–0.50 wt% Fe and 0.18–0.40 wt% O.[70] Near-α Ti
alloys primarily consist of α phase and a minor amount of β phase
(less than 5 vol%), which are resulted from the addition of a very
small amount of β-stabilizers (1–2 wt%) compared with α-type Ti
alloys.[69] Both α-type Ti alloys and near-α Ti alloys have similar
properties, such as excellent corrosion resistance,
good weldability, and high creep resistance.
Hence they are suitable for high-temperature
applications.[71] However, their strength is significantly low at room temperature and usually
cannot be improved by heat treatment since hcp
structure is much stable.[71] As seen from Table 4,
four grades of CP–Ti have the yield strength
and ultimate tensile strength in the range of
170–480 MPa and 240–550 MPa, respectively.[28]
Due to low mechanical strength at room
temperature, α-type Ti alloys have not been used
as implants.[36] Nonetheless, CP–Ti is the first
generation of biomedical materials developed in
1950–1990 and is widely used biomaterial for
dental and medical applications since they can
promote the rapid osteointegration,[1] because the
passive film of TiO2 on the surface of CP–Ti
contains much OH ions which can react with the
mineral constituents on the surface of bones (e.g.,
Ca2þ and (PO4)3–).[56] Unfortunately, low meFigure 2. Elastic moduli of various biomedical alloys and bone. (Reproduced with
chanical strength and low fatigue strength
permission from ref. [1]. Copyright (1998), Elsevier).
prohibit α-type Ti alloys to be used in load-bearing
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Table 4. Mechanical properties of Ti and Ti alloys properties developed for orthopaedic implant application.[5,13,28,36,72–79]
Materials
σ0.2 [MPa]
σUTS [MPa]
E [GPa]
emax [%]
α-type
CP–Ti grade 1
170
240
105
24
CP–Ti grade 2
275
345
105
20
CP–Ti grade 3
380
445
105
18
CP–Ti grade 4
480
550
105
15
(α þ β)-type
Ti–6Al–4V (annealed)
825–869
895-930
110–114
6–10
Ti–6Al–4V ELI (mill annealed)
795–875
960-965
101–110
10–15
10
Ti–6Al–7Nb
795
860
105
Ti–5Al–2.5Fe
820
900
110
6
Ti–3Al–2.5V
585
690
100
15
β-type
Ti–13Nb–13Zr (aged)
836–908
973–1037
79–84
42–44
1000–1060
1060–1100
74–85
18-22
Ti–15Mo (annealed)
544
874
78
21
Ti–15Mo–5Zr–3Al (ST)
838
852
80
25
1000–1060
1060–1100
–
18-22
945–987
979–999
83
16-18
Ti–12Mo–6Zr–2Fe (annealed)
Ti–15Mo–5Zr–3Al (aged)
Ti–15Mo–2.8Nb–0.2Si–0.260 (annealed)
Ti–16Nb–10Hf
736
851
81
10
Ti–35.5Nb–7.3Zr–5.7Ta
793
827
55-66
20
Ti–29Nb–13Ta–4.6Zr (aged)
864
911
80
13.2
Ti–24Nb–4Zr–8Sn (Hot rolled)
700
830
46
15.0
Ti–24Nb–4Zr–8Sn (Hot forged)
570
755
55
13.0
Ti–9Mn
1023
1048
94
19.0
Ti–6Mn–4Mo
1090
1105
89
15.0
Ti–10Fe–10Ta–4Zr
960
1092
–
6.0
–
760
65
18.5
670–1150
835–1180
32
6.5–12.9
Ti–12Cr
Ti–36Nb–2Ta–3Zr–0.3O
Ti–24Nb–0.5O
665
810
54
22
Ti–24Nb–0.5N
665
665
43
13
Ti–23Nb–0.7Ta–2Zr
280
400
55
33
Ti–23Nb–0.7Ta–2Zr–1.2O
830
880
60
14
σ0.2, σUTS, E, emax are yield strength, ultimate tensile strength, elastic modulus, and ultimate strain, respectively. ST means solution treated.
applications.[56] Similarly, near-α Ti alloys have not been used for
biomedical applications until now. For non-load bearing but
corrosion resistant applications, CP–Ti is better than near-α Ti
alloys.[28,36]
2.2. (α þ β)-Type Ti Alloys
(α þ β)-Type Ti alloys has a higher content of β-stabilizers than
near-α Ti alloys. Correspondingly, (α þ β)-type Ti alloys possess
higher fraction of β phase (about 5–30 vol%).[1,80] (α þ β)-type Ti
alloys exhibit good fabricability, high strength at room
temperature and moderate strength at high temperature.[81] In
comparison to α-type Ti alloys, (α þ β)-type Ti alloys are heat
treatable,[82] and hence, their mechanical properties can be
optimized by heat treatment. The volume fractions and natures
Adv. Eng. Mater. 2019, 21, 1801215
of α and β phases may vary according to the alloy chemistry, heat
treatment temperature, and cooling rate.[83–86] Previous works
show that the strength of (α þ β)-type Ti alloys can be improved
by 30–50% via appropriate solution treatment and aging while
their elastic moduli maintain at a similar level.[36] Therefore, as
shown in Table 4, (α þ β)-type Ti alloys exhibit higher yield
strength and ultimate tensile strength than α-type Ti alloys.
Moreover, (α þ β)-type Ti alloys still have excellent corrosion
resistance and capability of fast osteointegration in the human
body environment by rapidly forming TiO2/OH film on the
surface of bones.[69]
Among (α þ β)-type Ti alloys, Ti–6Al–4V (Ti64) is the most
widely used biomedical alloy, accounting for 50% of the total Ti
production besides CP–Ti.[69] This alloy is commonly used in the
annealed condition and for orthopaedic and traumatology
implants.[82] Although Ti–6Al–4V is originally developed for
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aerospace purposes its attractive combined properties
of strength, fracture toughness accompany with excellent
corrosion resistance lead to the employment in biomedical
applications.[1,87] Ti–6Al–4V has been further improved by
decreasing the contents of interstitial impurities (such as O, N,
C, and H), which is, well known as Ti–6Al–4V ELI and has been
widely used for bone fixation plates and stem of artificial
hip joints.[88] As Ti–6Al–4V contains toxic V, Ti–6Al–7Nb, and
Ti–5Al–2.5Fe have been developed to replace Ti–6Al–4V.[88,89]
Ti–6Al–7Nb possesses better wear resistance than Ti–6Al–4V
and has been applied in medical devices including fracture
fixation plates, femoral hip stems, fasteners, wires, and
screws.[90] Ti–5Al–2.5Fe has been employed in the fabrication
of hip prostheses and hip prosthesis heads as it is metallurgically
similar to Ti–6Al–4V.[91] Although both Ti–6Al–7Nb and Ti–5Al–
2.5Fe can eliminate toxic V from Ti–6Al–4V, Al still exists in
these alloys, which may still lead to a certain disease (such as
Alzheimer).[12] Therefore, (α þ β)-type Ti alloys are not the
optimal choice for biomedical implants. CP–Ti, Ti–6Al–4V, Ti–
6Al–4V ELI, Ti–6Al–7Nb, and Ti–5Al–2.5Fe are the first
generation of biomedical materials that were developed during
1950–1990s.[1] The mechanical properties of these alloys, as well
as Ti–3Al–2.5V, are listed in Table 4. Apparently, the moduli of
(α þ β)-type Ti alloys are still significantly higher than those of
bones (Table 1–3), which may cause bone resorption and implant
loosening. Therefore, these reasons have motivated the
development of β-type Ti alloys which contain non-toxic βstabilizers (such as Mo, Ta, Zr) and have lower elastic modulus
and enhanced biocompatibility.[21,92]
2.3. β-Type Ti Alloys
Compared with (α þ β)-type Ti alloys, β-type Ti alloys contain
higher amounts of β-stabilizers (such as Mo, Ta, and Zr) and
lower amounts of α-stabilizers without formation of any
intermetallic phases.[71,93] β-Type Ti Alloys are heat treatable
and can be heated at solution treated condition followed by
aging at the temperatures between 450 and 650 C in order to
enhance their strength. Such a phenomenon is attributed to
the dispersion strengthening due to the partial transformation
of β to α phase.[11] In fact, the toughness, plasticity, heat
treatment capability, and age hardenability can be improved by
increasing the portion of β phase in the β-type Ti alloys,
resulting in decrease in elastic modulus.[71] Table 4 lists the
mechanical properties of some β-type Ti alloys. According to
Table 4, β-type Ti alloys have lower elastic moduli compared
with the other types of Ti alloys, which can facilitate to reduce
the stress shield effect. Meanwhile, β-type Ti alloys have
comparable strength and better biocompatibility and are
expected to have higher corrosion resistance in the human
body compared with (α þ β)-typesen Ti alloys, due to the abce
of micro-galvanic effect between different phases.[11,94,95]
Therefore, β-type Ti alloys hold the promise of accommodating
challenges in being used as biomedical materials. In the last
two decades, many new β-type Ti alloys[96] and gum metals (a
kind of β-type Ti alloys with specific alloy compositions appear
to be deformed via a dislocation-free mechanism)[97] with
comparable or optimized properties have been developed for
Adv. Eng. Mater. 2019, 21, 1801215
biomedical applications. The properties of such alloys are also
listed in Table 4.
Most of these β-type Ti alloys have a higher density than other
types of Ti alloys due to the higher content of higher density
alloying elements (such as Mo, Ta, Nb, and Zr).[97] These alloying
elements are rare, thereby leading to high cost of raw materials.
The high melting points of these alloying elements also cause
the segregation and difficulty in preparation of these alloys.[98]
Consequently, recent developments have focused on developing
new low-cost β-type Ti alloys which consist of low-cost alloying
elements such as Fe, Mn, Sn, and Cr, and possess favorable
mechanical properties for biomedical applications.[73,99] It
should be noted that attempts for designing more low-cost βtype Ti alloys are never ceased, and therefore the future
biomedical implant applications may be achieved by these
low-cost β-type Ti alloys which also show favorable properties.[3]
Usually, β-type Ti alloys are designed by considering the
stability of β phase of the titanium alloys. The stability of β phase
in Ti alloys can be predicted based on the molybdenum
equivalency (Moeq) and DV-Xα cluster method.[66] The equivalent
Mo content used in multi-component Ti alloys can be expressed
by the following equation[66]:
½Ta ½Nb ½W ½V
þ
þ
þ
þ 1:25½Cr
5
3:6
2:5 1:5
þ 1:25½Ni þ 1:7½Mn þ 1:7½Co þ 2:5½Fe
Moeq wt% ¼ ½Mo þ
ð1Þ
where [x] indicates the content of the element [x] in wt%.
In the DV-Xα cluster method, two parameters, i.e. bond order
(Bo) and metal d-orbital energy level (Md), are defined. The
former is the measurement of the bond strength between Ti and
an alloying element. The latter is related to the electronegativity
and the metallic radius of the alloying elements.[11,100] For a
given alloy, the average values of bond order (Bo) and d-orbital
energy level (Md) are given as follows[101,102]:
Bo ¼
X
X i ðBoÞi
X
Md ¼
X i ðMdÞi
ð2Þ
ð3Þ
where X i is the atomic fraction of element i in the alloy, (Bo)i and
(Md)i are the values of bond order and metal d-orbital energy
level for element i, respectively.[101,102] In this method, phase
stability map can be constructed using Bo and Md, as shown in
Figure 3.[11] The location of designed alloy is varied depending on
its chemical compositions.
Some new β-type alloys, such as in the Ti–Fe–Ta, Ti–Fe–Nb, Ti–
Nb–Fe, Ti–Zr–Fe–Cr, and Ti–Nb–Fe–Cr alloy systems, have been
developed according to the aforementioned methods.[65–67,103–107]
Such low-cost β-type alloys exhibit comparable mechanical
properties to the conventional β-type alloys (Table 5). In these
developed new β-type alloys, the stability of β phase can be
predicted by the value of Moeq. It is observed that the value of
Moeq decreases with increasing the fraction of β-stabilizers (e.g.,
Nb and Ta) and the resultant strength decreases. However, the
strength of the Ti–27Nb–7Fe–xCr does not continuously
decrease after the addition of Cr (β-stabilizer) is greater than
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Figure 3. Phase stability map based on Bo and Md parameters. Modulus (in GPa) of each Ti alloy is shown in the bracket. (Reproduced with permission
from ref. [11]. Copyright (1998), Elsevier).
4% but it maintains at a stable level in spite of the increase in
Moeq.[106] As shown in Figure 4a–d, the fraction of β phase in the
Ti–11Nb–xFe alloys designed by the current group increases
with increasing the content of Fe (β-stabilizer).[65] Correspondingly, the value of Moeq increases from 25 to 35. Impressively,
both the strength and elongation of the Ti–11Nb–xFe increase
simultaneously (Figure 4e). Such an enhancement of strength is
ascribed to the solid solution strengthening of Nb and Fe in β
phase. As such, Ti–11Nb–3.5Fe containing more β phase, which
exhibits large ductility and high strength, has an enhanced yield
Table 5. β stability indicators and mechanical properties of some designed Ti alloys.
β stability indicators
Bo
Md
Moeq
H [GPa]
E [GPa]
σ0.2 [MPa]
emax [%]
ref.
Ti–8Fe
2.7803
2.3444
20.0
5.75
127
1562
17.1
[67]
Ti–8Fe–5Ta
2.7849
2.3417
21.0
5.00
119
1062
37.3
[67]
Ti–8Fe–8Ta
2.7878
2.3399
21.6
4.38
107
1030
38.1
[67]
Ti–9Fe–9Ta
2.7875
2.3253
24.3
3.75
99
1010
39.5
[67]
Ti–10Fe–10Ta
2.7872
2.3105
27.0
3.38
92
1000
40.0
[67]
Ti–7Fe
2.7810
2.3570
17.5
5.92
130
1847
8.0
[66]
Ti–1Nb–7Fe
2.7830
2.3560
17.8
5.58
124
1785
10.0
[66]
Ti–4Nb–7Fe
2.7871
2.3550
18.6
5.13
115
1539
24.3
[66]
Ti–6Nb–7Fe
2.7911
2.3530
19.2
4.18
97
1144
32.1
[66]
Ti–9Nb–7Fe
2.7960
2.3520
20.0
3.72
89
1010
40.5
[66]
Ti–11Nb–7Fe
2.8000
2.3500
20.6
3.57
86
985
41.5
[66]
Ti–27Nb–7Fe
2.8303
2.3400
25.0
6.25
125
775
40.2
[106]
Ti–27Nb–7Fe–2Cr
2.8301
2.3190
27.5
3.80
108
1100
30.0
[106]
Ti–27Nb–7Fe–4Cr
2.8300
2.2976
30.0
3.75
88
950
34.1
[106]
Ti–27Nb–7Fe–6Cr
2.8298
2.2768
32.5
3.25
81
925
36.2
[106]
Ti–27Nb–7Fe–8Cr
2.8296
2.2555
35.0
3.50
77
930
36.7
[106]
Alloy nominal Compositions [wt%]
Bo is average bond order; Md is average d-orbital energy level Moeq is molybdenum equivalency; H is hardness; E is elastic modulus; σ0.2 is yield strength; and emax is ultimate
compressive strain.
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Figure 4. SEM micrographs of Ti–11Nb–xFe alloys: a) x ¼ 0.5, b) x ¼ 3.5, c) x ¼ 6, and d) x ¼ 9; e) compressive stress–strain curves of the Ti–11Nb–xFe
alloys at room temperature and f) elastic energy versus modulus plot comparing the Ti–11Nb–xFe alloys and some other Ti alloys. (Reproduced with
permission from ref. [65]. Copyright (2016), Elsevier).
strength and plastic deformation as compared to Ti–11Nb–
0.5Fe. Figure 4f shows the elastic energy versus modulus
comparing the Ti–11Nb–xFe alloys with some other Ti alloys.
Apparently, Ti–11Nb–6Fe and Ti–11Nb–9Fe have higher elastic
energies than CP–Ti, Ti–6Al–4V, Ti–15Mo, etc, indicating a
better bearing capacity in heavy loading compared with other Ti
alloys. Such results demonstrate that the phase constituents of
the designed β-type alloys can be tailored by adjusting the alloy
compositions, whereby the resultant properties can also be
optimized. Furthermore, Ti–Zr–Fe–Cr and Ti–Nb–Fe–Cr series
of β-type alloys with C15 laves phase precipitates exhibit high
strength (yield strength: 785 to 1285 MPa) and large plasticity
(plastic strain: 8.0 to 41.5%).[105,106] It was found that the C15
laves phase precipitates possess larger plasticity than the other
laves phase precipitates (such as C14) in the Ti alloys. Although
initiation and propagation of cracks were preferred in the C15
laves phase precipitate regions, C15 laves phase precipitates
could dissipate the deformation energy to some extent in the
Adv. Eng. Mater. 2019, 21, 1801215
form of heat during crack propagation. Therefore, an appropriate amount of C15 laves phase precipitates could improve the
strength of the alloys while maintaining large plasticity. The
high strength and large plasticity may ensure the load-bearing
capability of developed new alloys in the human body.
It has been reported that the deformation of solution-treated
β-type Ti alloys is governed by stress-induced martensite (SIM),
mechanical twinning, and dislocation slip.[102] As proposed by
Morinaga,[108] the DV-Xα cluster method also could be used to
reveal the relationship between the deformation behavior and β
phase stability for β-type Ti alloys. Very recently, Wang et al.[109]
pointed out that there was still problematic in the use of the
Bo Md diagram. For instance, it was found that the Ti–10V–
3Fe–3Al (wt%) lays within the twinning region in Bo Md
diagram; however, in reality, the stress-induced α” martensite
was observed in the microstructure after room temperature
deformation.[102] Therefore, Wang et al.[87] proposed a semiempirical approach to predict the deformation behavior of β-type
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Figure 5. The e=a-Δr diagram in which the Twinning/SIM and Slip regions are distinguished according to the reviews of published works. The use of
empty and full bullets represents the Slip- and Twinning/SIM-dominated deformation mechanisms, respectively. SIM means stress-induced martensite.
(Reproduced with permission from ref. [109]. Copyright (2018), Elsevier).
Ti alloys by using the average electron-to-atom ratio (e=a) and
atomic radius difference (Δr) as a supplement to the Bo Md
diagram, and thereby an e=a Δr map was suggested, as shown
in Figure 5. The main features of the e=a Δr map indicate that:
1) the slip region has a higher value of e=a than the
twinning/SIM region, which mainly takes place when e=a < 4.2
and Δr > –2.5) the critical values for e=a of the twinning/SIM
region have a peak when Δr is close to 0, and then decline as the
Δr increases. Therefore, in addition to the β phase stability, it is
proved that e=a and Δr also influence the deformation behavior
of β-type Ti alloys by controlling the difficulty of lattice shear.
However, the current alloy design methods mentioned in this
section are primitive at best. In order to develop more new
titanium alloys to fully meet the demand of the biomedical
applications, new methods with better performance prediction
are required.
2.4. Ti-Based Shape Memory Alloys
Ti-based shape-memory alloys, such as TiNi, TiNiAg, TiZr, and
TiNbSn, are also frequently used as biomedical materials due
to their unusual mechanical properties.[110–113] The nearequiatomic TiNi alloy is initially found to have superelasticity
and shape-memory effects involving high strength, high recover
strain, and a relatively low elastic modulus, which are based on a
reversible martensitic transformation.[114] Such advantages
enable TiNi alloy to have multifunctional applications. Even
so, biomedical and aerial applications remain the main object for
TiNi and other Ti-based shape-memory alloys. For biomedical
purposes, TiNi alloy exhibits comparable biocompatibility to
stainless steel and other Ti implants.[115] TiNi alloy also shows
slightly better corrosion resistance than Co-Cr-based alloys in
artificial saliva.[116] Thierry et al.[115] compared the thrombogenicity of TiNi vs. stainless steel stents in an ex vivo shunt porcine
Adv. Eng. Mater. 2019, 21, 1801215
model and the results showed that TiNi stents represent only a
few of thrombus mainly located at the strut intersections while
stainless steel stents clearly exhibit more thrombus. Many other
studies also have been conducted to investigate the biological
performance of TiNi alloy and have a consensus that TiNi alloy
exhibits biocompatible behavior both in vitro and in vivo.[117–119]
Meanwhile, the high strength and relatively low stiffness enable
TiNi alloy to be many bone implants, including maxillofacial
implant, stents, cervical and joint replacements, bone tissue
engineering, and spine fracture fixation.[110,120] Especially, the
unique shape-recovery behavior facilitates the mechanical
stability of TiNi alloy implant within the host tissue. The
recoverable strains in superelastic TiNi implant are achievable
during a patient’s walking and running activities,[121] which
facilitates the growth of bone cells near and/or in TiNi (porous)
implants. It is well known that implant strain can stimulate cell
growth by in vivo experiments using a guinea-pig model.[122]
Compared with TiNi alloy, the other metallic implants always
tend to be stiffer and therefore exhibit lower strains. Therefore,
TiNi alloy and other Ti-based shape-memory alloys implants
have the advantages for osseointegration.
The superelasticity and shape-memory effects are very sensitive
to chemical compositions of TiNi alloys since different compositions may lead to different phase constituents in TiNi alloys,
which finally results in distinctive transformation hysteresis
behavior and properties. Fu et al.[123] prepared TiNi films with
different Ti/Ni ratios by co-sputtering of a Ti50Ni50 (at%) target and
a separated Ti target at the temperature of 723 K. The TiNi film
possesses a quite low residual stress at room temperature and a
two-step phase transformation among martensite, R phase (i.e.,
the rod-like eutectic phase) and austenite in the case of the Ti
content of 51.3 at%. When the Ti content decreases to 47.3 at% or
increases to 53.0 at%, residual stress becomes significantly high
since the intrinsic stress is great and the relaxation of stress
triggered by the R phase transformation is limited in such contents
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TiNi alloys.[123] Therefore, based on this mechanism, many
researchers alloyed some other elements such as Nb, Cu, Fe, Zr,
and Ag in Ti alloy to manipulate the transformation hysteresis
behavior of TiNi alloy,[111,112,124–127] therefore the transformation
temperatures, mechanical properties, and fatigue properties
in recent decades. Hsieh et al.[125] studied the martensitic
transformation in Ti50.5-xNi49.5Zrx and Ti51.5-xNi48.5Zrx alloys
(x ¼ 0–25 at%) and found that the transformation peak temperature can rise to 50–450 C depending on the different additions of
Zr. Addition of Cu or Fe can slightly increase the range
of transformation hysteresis of TiNi alloy.[126,127] By contrast,
TiNiNb alloy has a wider transformation hysteresis and therefore
better shape memory effect.[128] Meanwhile, the addition of Nb
results in formation of more R phase in TiNiNb alloy, which would
lead to more superelastic recovery and elastic recovery. The
addition of Ag in the TiNi alloy has an antibacterial function
and the resultant TiNiAg alloy can also obtain a maximum
recovery strain of 6.4% with a total pre-strain of 7%.[111] However,
Ni is a basic component in most Ti-based shape-memory alloys,
hence the release of Ni ions from the corrosion of implant during
service is inevitable. As mentioned above, excessive release of Ni
ions may cause an allergic response, leading to severe health
problems. Therefore, real applications of Ti-based shape-memory
alloys should consider the problem of Ni ions release.[129] To
decrease the release of Ni ions from the surface of Ti-based shapememory alloys, surface treatments are commonly used, which is
introduced in Section 4.
3. New Preparation Methods of Ti Alloys for
Biomedical Applications
In the development of Ti alloys, many preparation methods,
such as solidification/casting, forging, powder metallurgy, and
sintering, have been employed. As these conventional methods
have been introduced in detail in many reviews, herein, only
some new preparation methods of Ti alloys for biomedical
applications are reviewed in this article.
3.1. Selective Laser Melting
Biomedical implants often have complex shapes and their sizes
are varied according to patients. As a result, the production of
biomedical implants using conventional technologies always
accompanies with significant time, material, and energy costs.
In recent decades, selective laser melting (SLM), as one of
additive manufacturing (AM) techniques, has been widely
employed to solve the conventional manufacturing problems of
Ti-based materials.[130–134] SLM technology is a layer-wise
process in accordance with a pre-designed 3D model under a
protective atmosphere. Ti alloy parts are produced layer by layer
by selectively melting, which consolidates the powder at a high
rate by using a computer-controlled laser.[130,135] SLM provides a
wide range of advantages compared with the conventional
manufacturing methods, such as lower production time, higher
material utilization, almost no geometric constrictions, and
further post-processing.[136] Due to the fast heating/cooling rate
in SLM processing, SLM-produced alloys have distinctive
microstructures in comparison with the conventional counterparts.[83,137,138] SLM-produced (α þ β)-type Ti–6Al–4V consists of
dominant fine acicular α’ martensite and some prior β grains,
while the conventional Ti–6Al–4V alloy (Grade 5) reveals typical
equiaxed α þ β microstructure.[136] This distinctive microstructure leads to the comparable or even better mechanical
properties of SLM-produced Ti–6Al–4V compared with the
conventional counterparts, as summarized in Table 6. SLM
investigations with respect to (α þ β)-type Ti alloys have also been
employed on Ti–6Al–7Nb, which has a favorable set of
mechanical properties, corrosion resistance, and biocompatibility.[139] Similar to Ti–6Al–4V, SLM-produced Ti–6Al–7Nb also
has primarily α’ martensitic plates.[140] As there is no powder
feedstock β-type Ti alloys available for SLM, little SLM
investigation has been conducted on β-type Ti alloys. The
current research group has conducted some SLM investigations
on β-type Ti–24Nb–4Zr–8Sn.[5,141–143] As shown in Table 6, the
tensile properties of Ti–24Nb–4Zr–8Sn samples fabricated by
SLM, hot rolling and hot forging are closely comparable.
Table 6. Mechanical properties of different types of titanium alloys processed by selective laser melting and conventional methods.
H [HV]
E [GPa]
σ0.2 [MPa]
σUTS [MPa]
emax [%]
ref.
Selective laser melting
261 13
106 3
555
757
19.5
[144]
Selective laser melting
–
–
500
650
17
[145]
Sheet forming
–
–
280
345
20
[130]
Full annealed
–
–
432
561
14.7
[39]
409
109
1110
1267
7.28
[146]
[1,13]
Processing method
CP–Ti
Ti–6Al–4V
Selective laser melting
Casting/superplastic forming
346
110
847
976
5.1
220 6
53 1
563 38
665 18
13.8 4.1
[5]
Ti–24Nb–4Zr–8Sn
Selective laser melting
Hot rolling
–
46
700
830
15.0
[147]
Hot forging
–
55
570
755
13.0
[148]
H is hardness; E is elastic modulus; σ0.2 is yield strength; σUTS is ultimate tensile strength; and emax is fracture strain.
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However, because of excessively high oxygen concentration in
the powder used, the SLM-produced Ti–24Nb–4Zr–8Sn does not
show the advantageous superelastic behavior as the counterparts
produced by conventional technologies.[5] Based on the above
results, it demonstrates that the SLM-produced Ti alloys can
meet the mechanical demand for biomedical applications.
Considering the application of biomedical implants in the
human body, the current research group have investigated the
corrosion behavior of SLM-produced Ti–6Al–4V in 3.5 wt% NaCl
solution compared with the commercial Grade 5 alloy.[83]
Electrochemical measurements showed that the passive current
density of SLM-produced Ti–6Al–4V alloy (0.841 0.0275 mA
cm2) is two times higher than that of commercial Grade 5 alloy
(0.390 0.0125 mA cm2) and the fitted film resistance of
commercial Grade 5 alloy (94 7 kV cm2) is also higher than
that of SLM-produced Ti–6Al–4V (40 6 kV cm2). These results
indicate that SLM-produced Ti–6Al–4V has unfavorable corrosion resistance than the commercial Grade 5 alloy, attributing to
the relatively large amount of acicular α’ martensite and less
amount of β-Ti phase in the microstructure of SLM-produced
Ti–6Al–4V.[83] Although heat treatment was carried out to
eliminate the acicular α’ martensite in SLM-produced Ti–6Al–
4V, the corrosion resistance of heat-treated SLM-produced Ti–
6Al–4V was still not improved.[84] By contrast, SLM-produced Ti–
TiB biocomposite showed better corrosion resistance than the
conventional counterparts in simulated body fluid.[53] All the
differences in corrosion behavior between SLM-produced Ti
alloys and conventional ones can be ascribed to their different
microstructures which result from the processing characteristics. However, there is still lack of investigation on the
corrosion behavior of SLM-produced Ti alloys (and other
materials, such as Al alloys and stainless steel[87,137,149–152]),
therefore more studies are needed to deeply understand the
corrosion behavior of SLM-produced alloys. Although SLM is
widely used to investigate the processing and mechanical
properties of titanium alloys (especially Ti–6Al–4V alloy), the
presence of α’ phase in the microstructure of the SLM-produced
Ti–6Al–4V has resulted in undesirable service properties of the
alloy. The current researches have attempted to avoid the
presence of α’ phase in SLM-produced titanium alloys.
3.2. Electron Beam Melting
Electron beam melting (EBM), as another popular metal additive
manufacturing technique, is considered to be an innovative
industrial production technology to produce biomedical Ti
alloys.[86,153,154] Similar to SLM, EBM technology also adopts a
layer-wise processing method using a high-energy electron beam
as a heat source to fabricate alloy parts on a powder bed based on
a computer aided designed (CAD) model. EBM can rapidly
prototype metal components with complex shapes using
minimal processing steps.[153] In comparison to SLM technology, EBM process is conducted in a vacuum environment and
can preheat the powder used, thereby producing alloy parts with
higher densities.[155] From an industrial point of view, EBM
becomes a great alternative method for manufacturing biomedical Ti alloys. During EBM process, the electron beam moves fast
and melts the powder to form a molten pool in a vacuum
Adv. Eng. Mater. 2019, 21, 1801215
environment, following with solidification of the melt. Therefore, the built layer underneath is influenced by the solidification
of the building layer. Due to the low cooling rate and heat
accumulation during EBM process, the overall temperature of a
built part is maintained at a high level.[153] Such processing
characteristics lead to a distinctive microstructure in the EBMproduced alloys compared with SLM-produced counterparts as
well as conventionally produced ones.[86,156]
To date, EBM has been widely used in the production of
Ti–6Al–4V. It has been reported that both EBM-produced
Ti–6Al–4V and wrought Ti–6Al–4V reveal a basket-weave
microstructure with columnar prior β grains delineated by α
grain boundary, regardless of the fair difference in microstructural fineness.[86] The prior β grains in an EBM-produced
Ti–6Al–4V have a strong h001i texture component along the
build direction.[157] Table 7 lists the mechanical properties of
some EBM-produced Ti alloys. As seen from Table 7, ultimate
tensile strength (UTS) and elongation of EBM-produced Ti–6Al–
4V have a small fluctuation. This is because the microstructure
of the EBM-produced alloy is influenced by its cooling rate which
is associated with the sample size in the processing environment.[158] The EBM-produced Ti–6Al–4V has a slightly higher
corrosion potential (–0.28 0.03 V vs. SCE) and a slightly lower
corrosion current (0.007 0.001 mA cm2) than those of the
wrought Ti–6Al–4V (Ecorr: 0.29 0.02 V vs. SCE; Icorr:
0.012 0.001 mA cm2), indicating its slightly better corrosion
resistance in phosphate buffered saline.[86] The reason can be
attributed to the higher fraction of β phase and significantly
refined lamellar α/β phases in the EBM-produced Ti–6Al–4V
compared with the wrought Ti–6Al–4V. In comparison to the
SLM-produced Ti–6Al–4V, EBM-produced Ti–6Al–4V exhibits a
better corrosion behavior. Such a distinction results from the
different phase constituents between the EBM-produced Ti–
6Al–4V and the SLM-produced Ti–6Al–4V.[83,136] It is speculated
that the corrosion behavior would be significantly similar in
monolithic phase Ti alloys irrespective of the manufacturing
methods adopted.
3.3. Preparation Methods of Porous Ti Alloys
Although the current Ti alloys can solve most of the existed
problems of biomedical applications, their elastic moduli are still
higher than those of human bones. In order to minimize the
stress shielding effect, new Ti alloys with lower moduli are
expected. Besides developing the new low-modulus β-type Ti
alloys as mentioned above, porous Ti alloy structures also have
lower moduli than their bulky counterparts. Therefore, many
researchers dedicated to developing porous Ti alloys in recent
years. Porous Ti alloys which possess sufficient macro/micropores for bone cell ingrowth and vascularization have adjusted
elastic moduli in a considerably wide range.[113,171,172] Although
porous metals with higher porosities may improve the
regeneration of bone cells, the porous structures have an
opposite influence on the mechanical properties. Therefore, it is
significant to balance the porosity of porous metal parts for bone
ingrowth and mechanical performance. As such, preparation
methods of porous Ti alloys with sufficient strength have been
investigated extensively.[128,171,173,174] The previous works show
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Table 7. The mechanical properties of EBM-manufactured titanium materials.
Materials
σUTS [MPa]
σ0.2 [MPa]
emax [%]
H [GPa]
E [GPa]
ref.
Ti–6Al–4V
950–990
910–940
14–16
–
120
[159]
Ti–6Al–4V
910 10
830 5
–
3.21 0.02
118 5
[160]
Ti–6Al–4V
1100–1400
–
12–25
4.3
–
[161]
Ti–6Al–4V
1150-1200
1100–1150
16–25
3.6–3.9
–
[162]
Ti–6Al–4V
775
735
2.3
3.69
93
[163]
Ti–6Al–4V
994–1029
883–938
11.6–13.6
–
–
[164]
Ti–6Al–4V
930
865
12–17
–
–
[165]
Ti–6Al–4V
944.5–964.5
823.4–851.8
13.2–16.3
3.19–3.27
–
[166]
Ti–6Al–4V
–
832–1049
–
–
–
[157]
Ti–6Al–4V
990–1180
900–1100
18–23
–
–
[167]
336 26
253 13
0.27 0.1
–
166 2
[168]
γ–TiAl
–
1400
–
4.1
–
[169]
Ti–24Nb–4Zr–8Sn
–
–
–
2.5
–
[170]
Ti–48Al–2Nb–0.7Cr–0.3Si
σUTS is ultimate tensile strength; σ0.2 is yield strength; emax is maximum plastic strain; H is hardness and E is elastic modulus.
that pore size, shape, orientation and distribution of fabricated
porous Ti alloys can be controlled by different preparation
methods, thereby resulting in the distinction in mechanical
properties.[171,175,176] Some popular methods for porous Ti alloys
are introduced below.
3.3.1. Selective Laser Melting and Electron Beam Melting
Both selective laser melting and electron beam melting
technologies are capable of producing Ti alloy parts with
complex shapes.[141] By pre-designing 3D models and optimizing parameters, porous Ti alloys scaffolds with almost full
density can be produced.[27,130,177,178] Due to the distinction in
structures and scanning areas compared with the solid Ti alloy
parts, porous Ti alloy parts have a smaller melt pool because of
the smaller input energy levels needed for obtaining high
density. The current research group compared the microstructure and properties of rhombic dodecahedron Ti–24Nb–4Zr–
8Sn scaffolds with a porosity of 75% prepared by EBM and SLM
technology, respectively. The spot size of the laser (40 mm) is
much smaller than that of electron beam (200 mm), which
results in a smaller melt pool of SLM than that of EBM. As a
result, the SLM as-produced surface with 50 mm layer thickness
is smoother than that of EBM with 70 mm layer thickness, as
shown in Figure 6. Correspondingly, the scaffolds with the
same structure exhibit somewhat difference in mechanical
properties.[155] The SLM-produced Ti–24Nb–4Zr–8Sn parts
exhibit slightly lower elastic modulus (0.95 0.05 GPa) and
higher compressive strength (50 MPa) than the EBMproduced Ti–24Nb–4Zr–8Sn (elastic modulus: 1.34 0.04
GPa; compressive strength: 45 MPa). This difference can be
diminished by heat treatment. Furthermore, Liu et al.[133]
investigated the mechanical behavior of SLM-produced porous
Ti alloys with 75% porosity using cubic, topology optimized, or
rhombic dodecahedron structures and the results revealed that
these three structures had different local stress concentrations
Adv. Eng. Mater. 2019, 21, 1801215
and stress distributions during deformation, leading to a
distinction in energy absorption, as shown in Figure 7. Such a
result demonstrates that the mechanical properties of porous Ti
alloys can be controlled by tailoring their architecture. Table 8
summarizes the compressive mechanical properties and phase
constituents of EBM- or SLM-produced porous Ti alloys.
Superelastic properties are found in the EBM-produced Ti–
24Nb–4Zr–8Sn porous parts with a porosity between 67.9% and
91.2%. In comparison to the Ti–6Al–4V porous parts, the
Ti–24Nb–4Zr–8Sn porous parts exhibit a greater fatigue
strength due to the superelastic properties and the larger
plastic zone near the fatigue crack tip. For the same fatigue
strength, the modulus of the Ti–6Al–4V porous parts is two
times larger than that of the Ti–6Al–4V porous parts.[177] For
these reasons, Ti–24Nb–4Zr–8Sn porous parts are more
suitable for biomedical implants.
3.3.2. Spark Plasma Sintering
Conventional sintering of Ti and its alloys requires a high
temperature (1200–1400 C) in a high vacuum environment for a
long time (24 to 48 h), which results in harmful changes in the
microstructure and mechanical properties of the sintered
products.[188] Recently, spark plasma sintering (SPS), which
adopts the application of an external current to assist powder
consolidation, have attracted much attentions.[189–192] This
technique sinters the powder under strong electric field and
stress field and low sintering temperature. It also has the
advantages of fast heating/cooling rate and the combination of
sintering and heat treatment.[193] Therefore, SPS technology can
efficiently save energy to sinter the powder used. Current studies
on the preparation of porous Ti alloys by SPS mainly focus on the
conditions of low temperature and low pressure that can easily
sinter Ti and its alloys in powders,[194,195] because high current
discharge produces ionization in the plasma can melt the local
oxide film on the particles surface and allow the formation of
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Figure 6. a) SEM micrographs for the surface morphologies of EBM- and SLM-produced rhombic dodecahedron structural Ti–24Nb–4Zr–8Sn scaffolds
and b) melting schematic diagram. (Reproduced with permission from ref. [155]. Copyright (2016), Elsevier).
junctions between particles. The porous CP–Ti prepared at
700 C for 8 min under 50 MPa by combining SPS with NaCl
dissolution method (NaCl can be easily removed by dissolution
in water) showed a yield strength from 27 to 94 MPa and elastic
modulus from 6.2 to 36.1 GPa and represented the highly
interconnected structure and uniform distribution of pores.[196]
The typical microstructures of porous CP–Ti with the porosity of
55% but various pore sizes are represented in Figure 8.[196]
Porous Ti–5Mn was also prepared by using SPS technique
associated with ammonium hydrogen carbonate and by blowing
agent TiH2 under a low-pressure condition at different temperatures for 5 min.[197] The porosity of prepared porous Ti–5Mn
reduced from 56% to 21% with the increase in sintering
temperature from 950 to 1100 C, while their elastic moduli
increased from 35.0 to 51.8 GPa.[197] Moreover, SPS can produce
field effect between powder particles and effectively prevent the
growth of the grains during sintering and therefore it becomes a
nanostructured/ultrafine-grained material forming technology.[198] Hence, SPS provides a novel method to fabricate
nanostructured/ultrafine-grained Ti and its alloys for biomedical
applications.
3.3.3. Space Holder Method
Space holder method is one of the powder metallurgy
techniques using space holder to sinter the metal powder,
Adv. Eng. Mater. 2019, 21, 1801215
which can control the pores including volume fraction, size,
shape, and distribution depending on the amount and
characteristics of space holder used.[199] Space holders should
be strictly chosen since the left space holder may be
detrimental to the human body. In general, NH4HCO3, NaCl,
starch, Mg powder, polymethyl methacrylate, and urea are
used as space holders, which are bio-inert or biocompatible.[199] Such space holders can be evaporated at a relatively
low temperature or can be removed by water.[12] Therefore,
once the sintering and the spacer holder removal are finished,
the porous metal materials are produced. There are many
successful examples of the production of porous CP–Ti and Ti
alloys. Venezuela et al.[118] used urea as spacer holder to
fabricate porous CP–Ti, which was heat treated at 180 C for
2 h in air to eliminate the spacer particles and then sintered at
1200 C for 1 h. The resultant microstructure showed
interconnected pores and small isolated pores and had an
average pore size of 480 mm and total porosity of 36%. Porous
Ti–6Al–4V was ever produced by Esen et al.[200] using 30–70
vol% Mg powder (660 mm in average diameter) as spacer
holder. The porosities of final products ranged from 43% to
64% with an average pore size from 485 to 572 mm, resulting
in the elastic modulus and yield strength of final products
ranging from 1.4 to 14.7 GPa and from 28 to 150 MPa,
respectively. β-type Ti alloys can also be synthesized by this
method. Wang et al.[171] prepared porous Ti–10Nb–10Zr (wt%)
using NH4HCO3 as spacer holder in the size of 500–800 mm.
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Figure 7. Stress–strain curves for multiple cyclic load-unloading curves and finite element method results for three structures. (Reproduced with
permission from refs. [133,155]. Copyright (2018), Elsevier).
After adding 20–60 (wt%) NH4HCO3, the porosities of
prepared Ti–10Nb–10Zr ranging from 42 to 74% and the
porous Ti–10Nb–10Zr alloy with 59% porosity exhibited the
highest elastic modulus and yield strength of 5.6 GPa and
137 MPa, respectively. Porous Ti–24Nb–4Zr–8Sn was fabricated using 30 vol% with polymethyl methacrylate and the
products showed the porosities in the range of 35–43% with
an average pore size of 244–372 μm.[201] The compressive
strength and elastic modulus of prepared porous Ti–24Nb–
4Zr–8Sn were in the range of 21–260 MPa and 1.5–6.6 GPa,
respectively. All these results match the mechanical properties
of bones exhibited in Table 1–3. Therefore, space holder
method is a useful method for the production porous Ti and Ti
alloys since this method can produce porous parts with
controllable pore size, shape, volume fraction, and distribution using an appropriate space holder material which can be
removed at a relatively low temperature without excessive
contamination. Although this method to produce porous Ti
parts is convenient and has a lot of advantages, the accuracy of
pore size is not expected. Furthermore, the pore shape in the
produced Ti parts is restrained by the shape of the space
holder, which also influences the structure of the produced Ti
parts. Therefore, unlike the porous structure produced by
additive manufacturing technologies, the porous Ti alloys
produced by the space holder method have limited porous
structure choice.
Adv. Eng. Mater. 2019, 21, 1801215
3.3.4. Loose Sintering of Powder
Loose sintering of powder is the simplest method for producing
porous Ti and Ti alloys based on the partial sintering of powders,
and it is a cost-effective and most widely used powder metallurgical
technique.[202] Unless necessary, spacer holders are not needed
during sintering. Therefore, the possible contamination is
minimized. For this method, the porosity increases with
increasing the initial powder size or with decreasing the sintering
pressure, while the sintering temperature has little effect on the
densification of the produced scaffold.[12] Correspondingly, the
elastic modulus and yield strength of the scaffold linearly reduce
with increasing their porosity. Oh et al.[202] used this method to
prepare porous CP–Ti in a vacuum of 1 103 Pa and the result
showed that the porosities are varied according to the applied
pressure and particle size. The porous CP–Ti with 30 vol% porosity
exhibits an elastic modulus of 25 GPa and compressive yield
strength of 61 MPa, which is comparable with cortical bone
(Table 1). However, due to partial sintering of the powder used, the
porosity of produced scaffold is also associated with the degree of
particle interconnectivity and particle size. The pore size and shape
are determined by the powder size and shape. Therefore, the
porosity of the produced scaffold is limited to 50% and its pores are
not in spherical shape.[203] Therefore, the Ti and its alloy scaffolds
prepared by this method are brittle with low toughness and prone
to crack at a relatively low stress. Other powder metallurgical
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Table 8. Compressive mechanical properties of additive manufactured Ti alloys.[133,153]
Materials
Unit shape
Method
Porosity [%]
E [GPa]
σmax [MPa]
ref.
Ti–6Al–4V
Dodecahedron
SLM
75
1.40
42.8
[179]
Ti–6Al–4V
Rhombic dodecahedron
EBM
75
2
40
[180]
Ti–6Al–4V
Diamond structure
EBM
80
Parallel: 1.6 0.3
Parallel: 29.3 0.8
[20]
Perpendicular: 0.9 0.1
Perpendicular: 21.0 0.7
Ti–6Al–4V
Diamond crystal lattice
EBM
60–87
Parallel: 0.4–6.5
Parallel: 16.3–118.8
[181]
Ti–6Al–4V
Rhombic dodecahedron structure
EBM
62–86
Perpendicular: 0.89–6.34
Perpendicular: 12.4–112.8
[182]
Ti–6Al–4V
Rhombic dodecahedron structure
EBM
62.7–83.8
–
Perpendicular: 20–90
[183]
Ti–6Al–4V
Rhombic dodecahedron structure
EBM
58–88
Perpendicular: 0.5–6.5
Perpendicular: 10–100
[180]
Ti–6Al–4V
Honeycomb-like structure
EBM
66.3 2.1
Parallel: 2.5 0.5
Parallel: 116 10
[184]
Ti–6Al–4V
Hatched structure
EBM
60
Parallel: 12.9 0.9
Parallel: 148.4 3.5
[185]
Perpendicular: 3.9 2.1
Perpendicular: 127.1 29.2
Ti–6Al–4V
Cubic structure
EBM
49.75–70.32
Perpendicular: 0.57–2.92
Perpendicular: 7.28–163.02
Ti–6Al–4V
Cubic structure
EBM
58–88
Perpendicular: 0.5–15
Ti–6Al–4V
Lattice structures
EBM
88.8–94.7
Parallel: 0.025–0.079
Parallel: 2.21–6.01
Perpendicular: 0.05–0.21
Perpendicular: 2.28–8.78
Ti–6Al–4V
G7 (Magics software) structure
EBM
58–88
Perpendicular: 0.5–5 GPa
Perpendicular: 8–80
[180]
Ti–6Al–4V
Solid cellular foam structure
EBM
55–82
Perpendicular: 0.75–9.97
–
[187]
Ti–6Al–4V
Hollow cellular foam structure
EBM
55–89
Perpendicular: 0.48–9.77
–
[187]
Ti–6Al–4V
Stochastic foam structure
EBM
90.08–91.65
Perpendicular: 0.19–0.49
Perpendicular: 3.8–4.5
[182]
G7 (Magics software) structure
EBM
70
Perpendicular: 0.7 0.1
Perpendicular: 35 2
[38]
Ti–24Nb–4Zr–8Sn
Perpendicular:
20-200
[159]
[180]
[186]
Ti–24Nb–4Zr–8Sn
Rhombic dodecahedron structure
EBM
75
Parallel: 1.34 0.04
Parallel: 46
[155]
Ti–24Nb–4Zr–8Sn (annealed)
Rhombic dodecahedron structure
EBM
67.9–91.2
Parallel: 0.475–1.8
Parallel: 11.6–50
[177]
EBM is electron beam melting; SLM is selective laser melting; E is elastic modulus and σmax is ultimate compressive strength.
techniques such as microwave sintering,[204] combustion synthesis,[205] metal injection molding,[206] capsule-free hotisostatic
pressing,[110] solid-state isothermal foaming,[199,207] freeze (reverse
freeze) casting,[208] slip casting,[209] and slurry foaming[207] are also
reported to produce porous Ti alloys for potential biomedical
applications.
Although the porous Ti alloys have significantly great
potential applications in biomedicine, there are still a few
limitations from a clinical point of view.[210] For example, the
pores in the porous Ti alloys may not be interconnected,
therefore the ingrowth of the bone cells and osseointegration
might be hindered.[210] Meanwhile, a pre-shaped implant may
not be used in the situation of an abnormal anatomical structure.
Hence, the production of the porous Ti alloys should be better
assessed and suited for the patient. To date, the research for
some methods mentioned above to produce porous Ti alloys is
still at a very early stage. Few works have reported the fatigue and
corrosion properties of the porous samples.[155,211,212] Besides
strength and elastic modulus, these two properties are also vital
for the porous Ti alloys to be used as biomedical materials, which
needs further and deeper research in the future.
3.4. Severe Plastic Deformation
Compared with coarse-grained materials, ultrafine-grained
materials (in submicron or nanometer) often have good physical
Adv. Eng. Mater. 2019, 21, 1801215
and chemical properties, such as higher hardness, higher
strength, longer fatigue life, and better low-temperature superelasticity.[25,213–222] Some ultrafine-grained materials even exhibit good thermostability, high corrosion resistance, high wear
resistance, and excellent biocompatibility.[25,214,223–230] Ultrafinegrained materials, especially nano-grained materials, have fewer
atoms in each grain compared with the coarse-grained materials
and hence possess more surface atoms, leading to high surface
energy.[227] Therefore, the bone cells are prone to attach
themselves to the surface of such materials, resulting in high
osseointegration. Khang et al.[231] compared the influence of
nanometer and submicron surface features on vascular and
bone cell adhesion on CP–Ti using flat smooth surfaces and
found that the adhesion of bone cells was very high on these
surfaces. For instance, the cell density on the surface of
nanocrystalline CP–Ti (2000 cell/cm3) was much higher than
that of conventional CP–Ti (1400 cell/cm3).[1] Similarly, nanocrystalline Ti–6Al–4V also had a higher cell density (1600 cell/
cm3) than conventional Ti–6Al–4V (950 cell/cm3).[1] Such
phenomenon was also found in other metal and ceramic
nanomaterials for biomedical applications, such as Co–Cr-based
alloys, hydroxyapatite, and TiO2.[232,233] As such, the ultrafinegrained materials are expected to be the next generation of
biomedical materials. So far, many different severe plastic
deformation (SPD) methods have been proposed, including
equal channel angular extrusion (ECAP)[220,221,228,234,235] and
accumulative roll bonding (ARB),[236–239] high-pressure torsion
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Figure 8. Typical microstructures of porous CP–Ti with the porosity of 55% but various pore sizes: a) 125 mm, b) 250 mm, c) 400 mm, and d) 800 mm.
(Reproduced with permission from ref. [196]. Copyright (2018), Wiley).
(HPT),[240] multi-directional forging,[241] twist extrusion,[242]
cyclic-extrusion–compression,[243] friction stir processing
(FSP).[244]
Of these various methods, ECAP, ARB, and FSP have been
applied to Ti alloys for biomedical applications. ECAP produces
ultrafine-grained materials by introducing high strains and
simple shear using a common extrusion machine.[221] An
et al.[235] prepared nanostructured CP–Ti with 200–300 nm in
size by ECAP and the in vivo (the femurs of New Zealand rabbits)
study showed that the viability, proliferation, and adhesion of
cells cultured on the ultrafine-grained CP–Ti are superior to that
of conventional CP–Ti with the average grain size of 30 mm. Nie
et al.[234] also produced nanostructured CP–Ti with the average
grain size of 250 nm and examined the material in vitro and in
vivo. The nanostructured CP–Ti showed good bioaffinity and
17% increment in protein adsorption and in vivo experiment in
the tibia of Beagles dogs presented no fibrous tissue encapsulation in the gap between the autogenous bone and implant made
of nanostructured CP–Ti.[234] ARB is an intense plastic straining
process, which is conducted by together rolling overlapped
strip.[238] Milner et al.[236] successfully prepared CP–Ti with an
average grain size of 100 nm by ARB at 450 C. Cojocar et al.[237]
conducted cold ARB on α-type Ti–10Zr–5Nb–5Ta alloy and
obtained the nanostructured/ultra-fine grained structure. β-type
Ti alloy of Ti–25Nb–3Zr–3Mo–2Sn was also successfully
prepared by ARB and showed a significant increase in
Adv. Eng. Mater. 2019, 21, 1801215
strength.[238] Kent et al.[239] demonstrated that the grains in
nanostructured Ti–25Nb–3Zr–3Mo–2Sn is stable during the
prolonged annealing treatment at 400 C and have a limited
growth at 500 C. It should be pointed out that ECAP and ARB
technique significantly improve the strength of Ti alloys while
maintaining a low elastic modulus. Unlike the ECAP and ARB,
FSP is an attractive solid-state surface modification technology,
which produces a surface layer with ultrafine-grained microstructure. All these surface severe plastic deformation methods
can be implemented at room temperature to obtain adherent
surface layer. These methods have attracted much attention due
to their convenience and reliability for a great potential for
biomedical applications. Unfortunately, the limitations of severe
plastic deformation are also apparent. Because a large external
force is needed during deformation, specific equipment is
always needed. Therefore, these technologies are generally
associated with high energy cost and equipment cost. Furthermore, the size is limited for the sample fabricated by some of
these technologies, such as ECAP, ARB, and HPT, which
restricts the wider applications of the severe plastic deformation
methods.
Based on the statements above, the mechanical properties of
Ti alloys prepared by the methods mentioned in this review can
be simply summarized in Figure 9. The mechanical properties of
cortical bones and cancellous bones are represented by a light
gray ellipse and a red star. The Ti alloys are located in the
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Figure 9. Ultimate strength and modulus of bones and Ti alloys. CON
means conventional Ti alloys. Ultimate strength includes ultimate tensile
strength and ultimate compressive strength. The mechanical property of
cancellous bones is represented by a red star (at very left bottom in the
figure). NEW indicates the Ti alloys developed by the DV–Xα cluster method
or molybdenum equivalency method, SLM is selective laser melted Ti alloys,
EBM is electron beam melted Ti alloys, SPD is the Ti alloys prepared by
severe plastic deformation, P is the porous Ti alloys produced by powder
metallurgy, selective laser melting or electron beam melting.
different positions in this map according to their compositions,
microstructures and preparing methods. High strength is a basic
demand for load-bearing applications of the implants in the
human body. Apparently, all bulky Ti alloys have higher strength
than the bones irrespective of preparing methods. However, the
elastic moduli of bulky Ti alloys are still much higher than the
bones. The Ti alloy implants with an appropriate modulus are
one of the most important prerequisites of the successful
implantation for long-term service. The ultrafine-grained Ti
alloys prepared by severe plastic deformation have higher
strength and cell bioaffinity than the conventional counterparts,
their moduli are still comparable to the conventional counterparts. Therefore, after solving the questions of strength and
biocompatibility, further reducing the moduli of Ti alloys is a
critical consideration. As seen in Figure 9, only porous Ti alloys
have comparable mechanical properties to the bones. Unlike
bulky metallic materials, the mechanical properties of porous
metallic materials can be adjusted by their porosities as well as
the chemical compositions and microstructures. The porous Ti
alloys also have the advantages of bone cell ingrowth and
vascularization. Therefore, compared to the development of new
low-modulus titanium alloys, porous Ti alloys with refined
microstructure may be an important future development for
biomedical materials.
4. Surface Modification of Ti Alloys for
Biomedical Applications
Although Ti and its alloys have been widely used for hard tissue
replacement and the cardiovascular applications, there are still
Adv. Eng. Mater. 2019, 21, 1801215
some problems for their long-term performance.[245] First, Ti
alloys are bio-inert. Although they have excellent biocompatibility, the compositions of Ti-based alloys are significantly different
from the surrounding bones. Therefore, no fiber capsules are
formed around the biomedical implants, which leads to the
mechanical osseointegration between the Ti alloy implants and
bones but not strong chemical bone bonding.[245] Second, due to
the low resistance to plastic shear, work hardening properties
and low adhesion of the surface oxide film, the sub-surface of Ti
alloys cannot be protected perfectly. Hence, bare Ti alloys are not
capable of resisting to the adhesive and abrasive wear triggered
by relative movement between implants and bones, resulting in
loosening of implants and final failure.[246] Third, a stable
passive film can form on the surface of Ti alloys, thereby having
high corrosion resistance in Ti alloys. However, owing to the
complexity of the human body environment, the surface passive
film may be peeled and dissolved by the combined effect of
external force and body fluid.[247] As a result, the releasing
substance or wear debris may be toxic or cause inflammation
and thrombus in the human body.[247] Such problems are often
restricted by the surface properties of the implants. Hence,
various surface modification methods have been employed to
improve the biofunction, wear, and corrosion resistance of the Ti
alloy implants.
4.1. Surface Modifications for Biofunction
So far, there are several common methods for preparing
the bioceramic bioactive coatings, such as plasma spraying,[248,249]
electrochemical deposition,[250] sputtering,[251,252] dip coating,[253]
sol-gel,[254] bionic growth method,[255] sintering,[256] micro-arc
oxidation.[257] The prepared bioceramic coatings include hydroxyapatite,[249,258] fluorapatite,[254,259,260] β-tricalcium phosphate,[251]
β-wollastonite,[248] bioactive glasses,[261] and so on. Among these
bioceramic coatings, the first three ones have been extensively
investigated. Hydroxyapatite (Ca10(PO4)6(OH)2) is the principal
constituent of hard tissues (e.g., bone and tooth) in human body
and have excellent biocompatibility and bioactivity, which
encourages the growth of new bone tissues on the hydroxyapatite-coated implants and inhibits the release of metal ions
from entering the surrounding bone tissues.[249,258] Many
research demonstrated that Ca10(PO4)6(OH)2 coated Ti and Ti
alloy implants prepared by various methods have an earlier
osteoid formation than the uncoated Ti and Ti alloy
implants.[249,252,254,258,262] However, the main drawbacks of the
Ca10(PO4)6(OH)2 coating are the low bond strength and high
residual stress.[263] Therefore, the Ca10(PO4)6(OH)2 coating is
prone to peel off from the implants during long-term service in
the human body. As such, some researchers introduced transition
layer (e.g., ZrO2 and TiO2) between substrate and Ca10(PO4)6(OH)2
coating, resulting in an improvement in bonding strength and the
resultant stability.[264,265] Furthermore, fluorapatite (Ca5(PO4)3F)
is a stable mineral in the natural world and actually it can be
obtained from replacing OH by F in Ca10(PO4)6(OH)2.[260] Since
the solubility of Ca5(PO4)3F in body fluid is lower than that
of Ca10(PO4)6(OH)2, Ca5(PO4)3F is more stable than
Ca10(PO4)6(OH)2 in the human body.[259] Many studies demonstrate that the concentration of F- has an important influence on
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the performance of the bioceramic coatings. Tredwin et al.[266]
compared the bond strength and interaction of Ca10(PO4)6(OH)2,
and
fluoridated
hydroxyapatite
(Ca10(PO4)6F2-xOHx)
Ca5(PO4)3F with Ti–6Al–4V produced via the sol-gel method
and found that the bond strength increased with increasing the
increasing F substitution. Wang et al.[267] also prepared
Ca10(PO4)6F2-xOHx coatings with various concentrations of F
by the sol-gel dip-coating method and pointed out that the
Ca10(PO4)6F2-xOHx coatings of Ca10(PO4)6F0.8–1.1(OH)1.2–0.9 have
a stronger stimulating effect on cell proliferation and differentiation activity. The solubility of β-tricalcium phosphate (β-Ca3(PO4)2)
is much higher than those of hydroxyapatite and fluorapatite,
thereby easily degrading in the human body.[268] The degradation
of β-Ca3(PO4)2 may participate in the formation of new bone,[269]
thereby accelerating the growth of bone tissue. Additionally,
research with respect to use wollastonite (Ca3(Si3O9)) and diopside
(MgCaSi2O6) coatings to improve the bioactivity of Ti and Ti alloy
implants was also reported.[247,248]
For TiNi series shape memory alloys, surface treatments such
as surface coating,[270,271] heat treatment,[272] and ion implantation[273] are commonly used in order to reduce the release of Ni
ions. TiN and TiO2 coatings are deposited on TiNi alloy by
physical vapor deposition and the release of Ni ions, respectively,
decreases by the factors of 10 and 14 for TiN and TiO2 coatings
compared with uncoated counterparts.[270] Crystalline hydroxyapatite coating on TiNi alloy prepared by chemical treatment can
facilitate to reduce the Ni ion release.[271] This method also
enhances the bioactivity of TiNi implants. Heat treatment in air
promotes the formation of TiO2 film on the surface of TiNi alloy,
thereby reducing Ni ion release.[272] Wu et al.[272] have found an
optimal temperature of 450 C for the surface treatment of
TiNi alloy in air. Lowering the heat treatment temperature
(300–450 C) can also reduce Ni ion release of TiNi alloy
compared with the untreated counterpart, while the output is
less than that treated at 450 C. By contrast, annealing at a higher
temperature (550–800 C) has a negative effect.[272] Oxygen
plasma immersion ion implantation produces a barrier layer on
TiNi alloy to reduce the Ni ions leaching with almost no
geometry limitation; therefore this method can be adopted to
porous sample in a complicated shape.[273] The treated TiNi alloy
shows only one third of Ni ions release compared with the
untreated counterpart in stimulated blood fluid for 28 days.[273]
Antibacterial property is an important aspect for implants.
Bacterial infection on the surface of implant results in the
bacterial attachment which may lead the implant to be resistant
to the host’s immune system. To avoid this situation,
antibacterial implants of Ti alloys are developed based on
the idea of anti-adhesive or anti-infective.[274] The anti-adhesive
method has a shortcoming for intraosseous implants since
the anti-adhesive surface of implant prevents the surface from
pollution, thereby influencing osseointegration.[275] The antiinfective method enables the implant to kill bacteria on its
periphery based on local drug delivery.[275] Therefore, antiinfective method or combined anti-infective method and antiadhesive strategies are more widely used. Mei et al.[276] used
anodization and Ag plasma immersion ion implantation to
produce TiO2 nanotubes containing Ag particle on CP–Ti for
dental application and both in vitro and in vivo experiment
demonstrated sustained antibacterial properties. Combining the
Adv. Eng. Mater. 2019, 21, 1801215
osteoconductive coating (such as calcium phosphate) with
bactericidal agents (such as anti-microbial Ag particle) can
enhance osseointegration and prevent infection of orthopaedic
implants simultaneously.[277]
Besides the in-organic coatings, biofunctionalized surface
modification with biopolymer, biomolecules and protein are
usually conducted with the aim of osseointegration and
antibacterial function. Physiological bone is a composite
material that contains an organic matrix phase with inorganic
mineral phase. Collagen I is composed of almost 90% of the
organic matrix phase which is known to be easily adhesive to
microbial.[277] Therefore, collagen-mimetic peptide coatings are
synthesized on implants to afford osteoinductive cues while also
reduce bacterial adhesion.[278] The engineered collagen mimetic
protein has specific extracellular matrix proteins that can rapidly
delivery integrin on osteogenic cells. Meanwhile, this mimic also
shows a reduction in adhesion of some types of bacterial
compared with human collagen I (100–1000 times) and
uncoated Ti alloy implant (10–100 times). Other collagen
coatings, which release absorbable antibiotics and enhance
osteoblast functions, have also been reported.[279,280] Incorporating growth factors, such as bone morphogenetic protein, into
biomolecular or biopolymer coatings are also a widely used
method to improve the adhesion of osteogenic cells.[281] It is
reported that bone morphogenetic protein has been grafted to an
anti-infective chitosan layer to produce a coating with both
osseointegration and antibacterial functions.[282] Bone morphogenetic protein in the coatings can significantly enhance both Ca
deposition and alkaline phosphatase activity compared with
coatings without bone morphogenetic protein.[282] A newer
method for osseointegration is the recruitment of osteogenic
cells by releasing signaling factors, which can induce osteogenic
cells to find the osteoinductive coating. Cytokine stromal cellderived factor 1α is one of the most well investigated recruitment
cues, which can improve the efficacy of bone morphogenetic
protein.[283] Moreover, cytokine stromal cell-derived factor 1α
may reduce inflection since it is a chemoattractant for
neutrophils.[284] Biopolymer chitosan coatings have been
commonly investigated in both antibacterial application and
bone tissue engineering since it can combine with osteogenic
agents to produce hybrid materials for biomedical applications.
It has been reported that the chitosan coating combined with
cell-adhesive arginine-glycine-aspartic acid is capable of reducing S. epidermidis and S. aureus (two types of bacterial) adhesion
by 85% and 67%, respectively, and increases the bone cells
growth.[285] Chitosan has also been used as a drug-carrier coating
combined with other antimicrobial agents, such as adsorbed
vancomycin, to enhance its antibacterial capability.[286] As stated
above, whatever in-organic or organic coatings used, an ideal
orthopaedic implant coating is desired to be multifunctional,
which both increases the ability of osteointegration and
antibacterial. Therefore, combining different technologies to
produce such ideal coatings is expected.
4.2. Surface Modifications for Improving Wear Resistance
In general, high wear resistance can prevent the loosening of
biomedical implants. Although Ti alloy has high specific
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strength and high corrosion resistance, their wear resistance is
not desirable.[1] Surface modification techniques including
physical deposition methods (such as ion implantation,[287]
laser cladding,[288,289] physical vapor deposition,[290] and plasma
spray coating[291]), thermochemical surface treatments (such as
nitriding,[292] carburization,[293] oxygen diffusion hardening,[294,295] and boriding[296]), and surface severe plastic
deformation (such as friction stir processing[244,297–299] and
surface mechanical attrition treatment[300,301]) have been
employed to improve the surface properties of Ti alloys.
TiN[290] and diamond-like carbon (DLC)[302] coatings are
commonly used in both physical deposition methods and
thermochemical surface treatments. For instance, TiN film and
its ramifications (e.g., TiON and TiAlN) have high hardness,
good wear resistance, and chemical inertness, which can be
prepared by both physical deposition methods and thermochemical surface treatments.[287,290,293] Subramanian et al.[290]
compared TiN, TiON, and TiAlN as surface coatings on CP–Ti
prepared by direct current reactive magnetron sputtering
method and indicated all coatings show good cytocompatibility
with L-929 mouse fibroblast cell lines. Choudhury et al.[303]
investigated the durability of functional DLC coated Ti–6Al–4V
for application in artificial hip joints and the results showed a
significant wear reduction to the DLC coated Ti–6Al–4V
compared with that of non-coated counterparts. Meanwhile, it
is reported that DLC films-coated Ti alloys, as well as other
materials, do not cause any cytotoxicity and inflammation during
long-term implantation.[304,305] Oxygen diffusion hardening
(ODH)[294] is another technology which has been investigated
since it can drastically improve the abrasive wear of Ti alloys. The
hardening effect of this technology results from the adherent
surface modification by oxygen diffusion, which does not
delaminate or spall like the deposited coatings.[306]
Although physical deposition methods and thermochemical
surface treatments can improve the wear resistance of Ti alloy
implants to some extent, some problems still exist. The physical
deposition methods are prone to delaminate between the coating
and substrate under the condition of repeated loading. The
thermochemical surface treatments are usually conducted at
high temperatures, leading to a twist of the substrate. Therefore,
surface severe plastic deformation methods are well developed in
recent years.[244,297,300,307,308] Surface mechanical attrition
treatment is a newly emerging method to achieve a nanostructured surface layer on metallic materials by peening with a large
number of high-frequency impacts over a short time.[301] A
gradient nanostructured surface layer of CP–Ti can be formed
after surface mechanical attrition treatment.[309] The grain size is
extremely small in the top-surface of CP–Ti and gradually
increases with increasing depth. Zhu et al.[310] successfully
prepared a nanostructured surface with the thickness of 50 um
on CP–Ti by shot blasting, demonstrating the feasibility of this
method. Subsequently, Huang et al.[311] examined the nanomechanical properties of nanostructured CP–Ti treated by
surface mechanical attrition treatment and found that surface
nanocrystallization leads to a significant increase in hardness
and scratch resistance. The surface nanostructured layer also
reveals a greatly better friction property than that of the CP–Ti
matrix. Wen et al.[312] conducted surface mechanical attrition
treatment on Ti–6Al–4V, resulting in a significant increase in
Adv. Eng. Mater. 2019, 21, 1801215
strength with acceptable ductility of the treated Ti–6Al–4V
compared with untreated counterpart, which indicates a
potential biomedical application. The in vivo and in vitro
experiments also elucidated the Ti alloy implants with severe
plastic deformed surface have good compatibility with
cells.[313,314] Friction stir processing is another attractive solidstate surface modification technology and focuses on the
localizing thermomechanical processing, which uses a professional fiction stir welding machine with a friction stir processing
tool of tungsten steel to severely deform the surface of the alloy
by fiction stirring at a relatively high rotation speed.[244,315,316]
Wang et al.[244] conducted fiction stir processing on the Ti–35Nb–
2Ta–3Zr and promoted the formation of α” martensite and ω
phase on its surface microstructure, resulting in surface
hardening. As it is known that, some beneficial elements or
compounds (e.g., Zn, Ag, and TiO2) are not contained in the
commonly used Ti alloy. Fortunately, using friction stirring
processing (FSP) can prepare surface composite layer. As shown
in Figure 10a–c, grooves in Ti–6Al–4V filled with TiO2
nanoparticles are processed before FSP. The amount of TiO2
nanoparticles can be controlled by the depth of the grooves.
During processing, not only the microstructure is refined but
also the TiO2 nanoparticles are uniformly distributed in the
surface layer of Ti–6Al–4V. After FSP, the TiO2 nano-powder is
mixed within the surface layer with refined microstructure.[316,317] In vitro experiments (Figure 10d, e) show that cell
viability and proliferation are substantially improved by FSP at
an appropriate parameter.[316,317] Zn participates in the synthesis
of various enzymes in the human body, maintaining cell
physiological function and promoting bone growth and
maintain cell physiological function.[318] Nanostructured Ag
particles have wide applications in biomedical fields due to its
antimicrobial properties.[319] Lv et al.[307] and Xie et al.[308]
prepared functional Ti–6Al–4V /Zn and Ti6–Al–4V/Ag composite layers by this method. Above all, the Ti alloys prepared by
these three methods are expected to have the surface roughness
in the nanometer range. Because our bone cells consist of
inorganic minerals with a grain size of 20–80 nm long and 2–
3 nm in diameter, the specific reaction can take place between
the surrounding tissue and the implant made of ultrafinegrained materials, leading to an enhancement of bioaffinity.[1]
Therefore, ultrafine-grained Ti and Ti alloys produced by severe
plastic deformation exhibit better biocompatibility. Combining
with the intrinsic properties of Ti and Ti alloys, they are expected
to have a great potential for biomedical applications.
4.3. Surface Modifications for Improving Corrosion
Resistance
Further improving the corrosion resistance of Ti alloys can
enhance their stability during the long-term service in the
human body. There are a lot of methods to improve the corrosion
resistance of Ti alloys, such as passivation,[320,321] laser surface
modification,[288,322,323] sol-gel,[254] ion implantation,[287] physical vapor deposition,[290] electroplating,[324] and micro-arc
oxidation.[325] It should be noted that some methods are the
same as the ones used for improving the bioactivity and/or wear
resistance of Ti alloys. As surface modifications for Ti alloys
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Figure 10. Friction stir processing system and the Ti–6Al–4V plates: a) the surface of Ti–6Al–4V plate with the grooves, b) grooves with different depths
for filling TiO2 nanoparticles, and c) friction stir processing for preparing composite layer on the surface of Ti–6Al–4V, d) the cell viability of Ti–6Al–4V
with and without friction stir processing in the cytotoxicity test, and e) cell proliferation results. Black star means p < 0.05 versus Ti–6Al–4V. Red star
represents p < 0.05 versus other two friction stir processing groups. (Reproduced with permission from ref. [267]. Copyright (2017), Elsevier).
should consider their properties of bioactivity, wear resistance,
and corrosion resistance simultaneously due to the complex
environment in the human body, improving one of these three
properties cannot be conducted at the cost of the other two
properties. Passivation and laser surface modification are usually
used for enhancing the corrosion resistance of Ti alloys as well as
other surface properties. Bloyce et al.[326] improved both
corrosion and wear resistance of CP–Ti, Ti–0.05Pd, Ti–6Al–4V
using palladium-treated thermal oxidation (one of the passivation methods) in air. Güleryüz et al.[327] investigated the
optimum oxidation conditions of Ti–6Al–4V to further evaluate
its corrosion-wear performance and found that the treated Ti–
6Al–4V exhibit excellent corrosion resistance after oxidation at
600 C for 60 h and 25 times higher wear resistance compared
with the untreated counterpart. Kumar et al.[321] compared two
passivation methods of thermal oxidation and anodizing
(another passivation method) and indicated the corrosion
performance and wear resistance of thermally oxidized CP–Ti,
are better than that of the anodized one. Laser surface
modification can be performed by various treatments in a wide
range because of the high coherence and directionality of
laser.[322] Laser surface remelting of CP–Ti and Ti alloys can
induce a layer of fine martensite and a film of Ti oxides (such as
Adv. Eng. Mater. 2019, 21, 1801215
TiO, TiO2, and Ti2O3), leading to an increase in both corrosion
resistance and wear resistance.[322,328] due to the rapid
solidification and the formation of oxides. Laser surface alloying
technique can generate an alloyed layer of 0.5–1.0 mm thickness
by the reactions between the pre-placed powders/gases and the
melting substrate.[322] For example, the TiN layer aforementioned can also be prepared by this method in a dilute nitrogen
environment, which has excellent high-temperature and corrosion resistance, and a high degree of hardness.[329] Ti–6Al–4V
with an alloyed surface layer containing TiB and TiN can be
produced by alloying BN powder.[330] The prepared surface layer
has a micro-hardness of about 1700 HV associated with good
wear and corrosion resistance.[330] Laser cladding is also a
commonly used technique that uses a defocused laser beam to
melt cladding materials onto the surface of Ti alloys.[322]
Cladding materials are melted to form a metallurgical bond with
the underlying Ti alloys. As TiN layer is demonstrated to have
better corrosion resistance than Ti alloy,[331] Balla et al.[332]
prepared TiN reinforced Ti–6Al–4V composite coating by laser
cladding and the results showed that the coatings containing
40 wt% TiN exhibit the highest wear resistance and good
biocompatibility in vitro. Other materials are also used as
cladding materials to improve the corrosion resistance of Ti
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alloys.[333] However, considering the biomedical applications of
Ti alloys, the selection of cladding materials must be prudent.
The passive film (or layers) with high wear resistance prepared
by various methods are also expected to have good corrosion
resistance.[333–335] Therefore, it can be concluded that suitable
surface modification technologies employed to enhance the
corrosion resistance of Ti alloys should not be compromising
their other advantageous properties.
coatings or surface treatment technologies.[341] Therefore,
multifunctional coatings prepared by combining different
technologies are likely to receive more attentions. As stated
above, β-type Ti alloys with multifunctional coatings are expected
to be a significant development for biomedical Ti alloys in order
to achieve better biomedical applications.
6. Conclusions
5. Biocompatibility of Ti and Ti Alloys
To meet the requirements of mechanical properties, corrosion
resistance and wear resistance are merely the prerequisites for Ti
and Ti alloys to be implant materials. Ideal implants can serve in
the human body for a long-term without revision surgery. After
being implanted in the human body, implants induce a lot of
reactions in the biological environment with body fluid, proteins
and cells, determining the ultimate success of implantation.
Therefore, biocompatibility tests are mandatory biomedical Ti
and Ti alloy implants. As well known, Ti and Ti alloys are inert
materials in the human body environment, which can
spontaneously form a stable passive film of TiO2 on the surface
of the implants. Even if the passive film is damaged, it is rebuilt
rapidly. This is an important reason that Ti and Ti alloys possess
good biocompatibility. However, although Ti and Ti alloys have
these advantages, the concerns on the biocompatibility of Ti and
Ti alloys have been widely discussed and studied since the
1970s.[336] For example, due to their inert characteristics, fibrous
tissue capsules are prone to form on the surfaces of implants,
while foreign body response may take place. For some Ti alloys,
several elements are toxic in elemental states, such as V in Ti–
6Al–4V alloy and Ni in TiNi series shape-memory alloys, hence
the release of these elements are harmful. Furthermore, some
properties like osseointegration and antibacterial function are
desired. Considering that specific interactions take place
between the surfaces of implants and the surrounded tissues,
the surface compositions and surface topography of implants are
therefore expected to be vital in contacting bone tissue. Table 9
summarizes some typical results for the biocompatibility of Ti
and Ti alloys. It is found that the biocompatibility of Ti and Ti
alloys can be mainly improved in two aspects. One is to reduce
the toxic elements used. Cell viability and growth rate for β-type
Ti alloys with nontoxic elements are much higher than those for
Ti–6Al–4V.[337,338] Hence, β-type Ti alloys are developed rapidly
in recent years. Up to date, the investigation of β-type Ti alloys is
still at a very early stage. Further studies about β-type Ti alloys are
urgently required. Another is the surface modification which can
improve the capability of osseointegration and antibacterial
function for Ti and Ti alloys. Although a large number of studies
show a lot of methods to enhance the osseointegration and/or
antibacterial function including organic and inorganic coatings,
specific coating often has a single function. For instance,
Collagen I coated and uncoated CP–Ti have no significant
difference in the growth rate of osteoblast-like cells.[339] By
contrast, human mesenchymal cells grow faster on the Collagen
I coated CP–Ti than those on the uncoated counterpart.[340]
These findings indicate the coatings have a distinctive response
to different types of cells. Similar results are also found in other
Adv. Eng. Mater. 2019, 21, 1801215
This article briefly reviews a number of recent developments of
Ti alloys for biomedical applications, including their processing,
microstructures, and properties. First, biomedical materials are
succinctly introduced in the view of the requirements of artificial
implants. Afterward, a detailed comparison among 316L
stainless steel, Co–Cr-based alloys, and Ti alloys is carried out
in according to their properties for using as implant materials.
Due to the excellent combined properties of low elastic modulus,
high specific strength, excellent corrosion resistance, complete
inertness to body environment and superior biocompatibility, Ti
and its alloys are a better choice for manufacturing biomedical
implants.
As reviewed, α-type and near-α Ti alloys were developed as the
first generation of biomedical materials. However, low strength
at room temperature and poor wear resistance limit them to be
used as implant materials. (α þ β)-type Ti alloys exhibit better
mechanical properties and have been used for orthopaedic
implants. Unfortunately, the moduli of (α þ β)-type Ti alloys are
significantly higher than those of bones and contains harmful
elements (e.g. Al and V). Therefore, lower elastic modulus β-type
Ti alloys with non-toxic β-stabilizers (e.g. Nb, Zr, Ta, and Sn) and
enhanced biocompatibility were developed in the last two
decades for biomedical applications. Ti-based shape memory
alloys are also frequently used as biomedical materials due to
their superelasticity and shape-memory effects.
Up to date, some new technologies are emerged to produce Ti
alloys for biomedical applications. Additive manufacturing (AM)
techniques, especially selective laser melting and electron beam
melting, have been frequently applied to produces Ti alloy parts
with the combined advantages of low production time, high
material utilization, almost no geometric constrictions and
further post-processing. Besides the AM techniques, porous
powder sintering technologies including spark plasma sintering,
space holder method and loose sintering of powder have been
widely employed, which produce Ti alloy scaffolds with
comparable mechanical properties to the bones in the human
body. Severe plastic deformation methods have also been applied
to Ti alloys for biomedical applications due to better combined
properties of strength, corrosion resistance and bioaffinity to
bone cells of ultrafine-grained materials. However, low biological
activity to surrounding tissue, poor wear resistance, and
inadequate corrosion resistance are still the questions for Ti
alloys during long-term service. Various surface modification
methods are conducted on Ti alloys to solve these problems. For
improving biological activity, the biofunctionalized surface
modification including inorganic and organic coatings are
commonly used methods. Wear and corrosion resistance of Ti
alloys can be improved simultaneously through various ways
such as physical deposition methods, thermochemical surface
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Table 9. Summary of the biocompatibility for Ti and Ti alloys.
Materials
Processing
CP–Ti
Titanium-based
dental or
prosthetic
implants
Evaluation
Comment
ref.
Lymphocyte migration inhibition
All blood samples from patients showed no immune
reactions after 3–12 months implantation
[342]
Memory lymphocyte
immune-stimulation assay
Patients had various symptoms after implantation; 54
patients (total 56) were negative in patch testing
[343]
Ti–6Al–4V
Wrought
Proliferation assay
No proliferative response
[344]
Ti–6Al–4V
Hot rolled
Dissolution test in Hank’s solution
Measurable dissolution was found in a solution
containing ethylenediamine tetra-acetate or Na-citrate
[345]
Ti–15Zr–4Nb–
4Ta
Aged
Rat tibia implantation for
6–48 weeks
Relative growth of cells for Ti–15Zr–4Nb–4Ta was much
higher than that for Ti–6Al–4V
[337]
Ti–29Nb–13Ta–
4.6Zr
Wrought
Cell viability test
Cell viability for Ti–29Nb–13Ta–4.6Zr was much better
than that for Ti–6Al–4V
[338]
Nanocrystalline
CP–Ti
Equal channel angular extrusion
Bonelike apatite immersion;
osteoblast cell lines test
Good bioaffinity and 17% increment in protein
adsorption
[234]
Nanocrystalline
CP–Ti
Equal channel angular extrusion
Beagles dog tibia implantation
No fibrous tissue encapsulation in the gap between the
autogenous bone and implant
[234]
Osteoblast adhesion test
Enhanced osteoblast adhesion
[232]
Nanocrystalline
Ti–6Al–4V
CP–Ti
Laser engineered net shaping
Human fetal osteoblasts
Cells well spread on the porous sample
[346]
CP–Ti
Fused deposition modeling
L929 mouse fibroblast
Excellent bone cell attachment and proliferation
[347]
Ti–6Al–4V
Electron beam melting
Human fetal osteoblasts
Decreased cell proliferation in greatly rough surface
[348]
Ti–6Al–4V
Electron beam melting
Human adipose-derived adult stem
cells
Increased cell proliferation in porous samples
[349]
Ti–6Al–4V
Electron beam melting
Rabbit femur and tibia
Displaying similar biological short-term behavior with
conventional counterpart
[350]
Ti–6Al–4V
scaffold
Selective laser melting
Human periosteum-derived cell
culture
Enhanced cell seeding while maintaining nutrient
transport
[351]
Ti–6Al–4V
Metal injection molding
MC-3T3-E1 pre-osteoblasts culture
Cells are proliferated, adhesive and differentiated.
[352]
CP–Ti
Acid processed in HF-HNO3
Osteoblast-like cells (MG63)
culture
No change in cell proliferation rate.
[353]
CP–Ti
Coarse grit-blasted and acid-etched in
HCl-H2SO4
Osteoblast-like cells (MG63)
culture
Lower enzyme activity in isolated.
[353]
CP–Ti
Radio frequency plasma sprayed
Sr-hydroxyapatite
Osteogenic cell response test
Osteoid formation at 7th week and 12th week for on Srhydroxyapatite coated implants and hydroxyapatite after
implantation coated implants
[262]
CP–Ti
Coated with α- or β-tricalcium phosphate
Rabbit cranial bones
α-tricalcium phosphate coating degraded faster; both
[354]
coatings were absorbed by osteoclast-like cells
Ti–6Al–4V
Coated with fluoridated hydroxyapatite
Osteoblast-like cell number test
Similar presence of cells on the surface with different
roughness; fluoridated hydroxyapatite supported
osteoblast adhesion and growth
[355]
Ti–6Al–4V
Coated with fluoridated hydroxyapatite
Osteoblastic cell response test
Similar cell morphologies and good cell viability in all
tested samples. Coatings with 0.8–1.1 mol/L Fconcentrations had a stronger stimulating effect on cell
proliferation and differentiation activities
[267]
Porous CP–Ti
cylinder
Chemical and heat treatment
Rabbit medial femoral condyle
Improved bioactivity without compromising mechanical
properties
[356]
Porous CP–Ti
cylinder
Hydroxyapatite coating by sol-gel method
Murine macrophage
RAW 264.7 test
Better osteogenic effect than uncoated sample
[357]
Porous CP–Ti
scaffold
Alkali-acid and alkali-heat coatings
Canine femur
Improved the osseointegration
[358]
Porous CP–Ti
Microarc oxidation and electrochemical
Bone marrow stem cell culture
Enhanced cell proliferation; improved levels of osteorix,
[359]
(Continued)
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Table 9. (Continued)
Materials
scaffold
Processing
Evaluation
reduction in alkaline solution
Comment
ref.
alkaline phosphates, osteopontin and osteocalcin
Microarc oxidation and electrochemical
reduction in alkaline solution
Canine femur
Enhanced bone cell regeneration and increased
bone-implant contact ratio
[359]
Porous Ti–6Al–
4V scaffold
Polydopamine-assisted
hydroxyapatite coating
MC3T3-E1 cells proliferation test
Increased cell attachment and proliferation; higher
expression of alkaline phosphatase, osteopontin,
osteocalcin and collagen
[360]
Porous Ti–6Al–
4V scaffold
Polydopamine-assisted
hydroxyapatite coating
Rabbit femoral condylar
Improved osteointegration and bone regeneration
[360]
Ti–6Al–4V
Friction stir processing surface with TiO2
composite layer
Cell viability and proliferation tests
Improved cell viability and proliferation
[316]
Ti–6Al–4V
Friction stir processing surface with Ag
composite layer
Antimicrobial tests
Improved antimicrobial properties
[308]
Tantalum plasma immersion ion implantation
Cytotoxicity and cell proliferation
test
Ni release was inhibited up to 60%; muscle cells and
osteoblasts were proliferated
[361]
Cell incubation
Cells shows 1.33 to 2-fold increase; After 28-day
implantation, all pores were completely filled with cells
[362]
Porous CP–Ti
scaffold
TiNi
Porous TiNi
CP–Ti
Coated with collagen I
Mesenchymal cell culture
Faster growth rate than uncoated sample
[340]
CP–Ti
Coated with collagen I
Osteoblast-like cell culture
No significantly differences with uncoated sample
[339]
CP–Ti
Covalently linked with short biomimetic
peptides
Sprague-Dawley rat osteoblast cell
culture
Increased osteoblast cell adhesion
[363]
CP–Ti
Immobilized with synthetic peptide
Dog mandible
Higher growth rate of osteoblast-like MC3T3 cells
[364]
Coated with collagen I
Mesenchymal cell culture
Increased cell adhesion, proliferation and differentiation
regarding with both the amount and stability of collagen I
[365]
Coated with biodegradable poly (D, L-lactic acid)
integrated with gentamicin and teicoplanin
Antibacterial test (Staphylococcus
epidermidis)
Significantly reduction in bacterial adhesion
[366]
Ultraviolet light irradiation
Osteoblastic cell culture and
antibacterial test (Staphylococcus
epidermidis and Staphylococcus
aureus)
Cell adhesion was not altered; bacterial adhesion was
reduced
[367]
Coated with hyaluronic acid/chitosan
polyelectrolyte multilayers, coupled with
surface-immobilized cell-adhesive
arginine-glycine-aspartic acid
Osteoblast adhesion and
antibacterial test
Enhanced beneficial host cell responses; reduced
pathogenic microbial adhesion
[368]
Ti–6Al–4V
CP–Ti
Ti–6Al–4V
CP–Ti
treatments and surface severe plastic deformation (e.g. friction
stir processing). An adherent or coherent layer with high wear
resistance and corrosion resistance are often prepared to
enhance the surface properties of Ti alloys.
Although Ti and its alloys have been widely used as
biomedical materials, deeper investigations are still requested
to increase their reliability in the human body during long-term
service. Based on the review above, a Ti alloy with non-toxic
elements, satisfied mechanical properties, and excellent biocompatibility is more expected. Therefore, porous β-Ti alloys with
refined microstructures coated by multifunctional materials may
be the next generation of biomedical implants.
Overseas Research & Training Program for University Prominent Young &
Middle-aged Teachers. The authors are grateful to C.D. Rabadia, S.E.
Haghighi, Y.J. Liu, S.F. Jawed, H.B. Lu, L.Q. Wang, G.H. Cao, and T.B.
Sercombe for the collaboration.
Conflict of Interest
The authors declare no conflict of interest.
Keywords
additive manufacturing, processing, properties, surface modification,
titanium alloys
Acknowledgements
This research was partially supported by the Australian Research Council
Discovery Projects (DP110101653, DP130103592), National Natural
Science Foundation (51601075), China Postdoctoral Science Foundation
(2017M611751), and Jiangsu University of Science and Technology
Adv. Eng. Mater. 2019, 21, 1801215
1801215 (23 of 29)
Received: November 15, 2018
Revised: February 6, 2019
Published online: March 20, 2019
© 2019 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
www.advancedsciencenews.com
www.aem-journal.com
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