MICRO-ARC OXIDATION ENHANCES
MECHANICAL PROPERTIES AND
CORROSION RESISTANCE OF TI-6AL-7NB
ALLOY
Qabas Khalid Naji
Biomedical Engineering Department, AL-Mustaqbal University Collage, Babil, Iraq.
qabas.khalid@mustaqbal-college.edu.iq
Jassim Mohammed Salman
Department of Metallurgical Engineering, College of Materials Engineering,
University of Babylon, Iraq.
mat.jassim.mohammed@uobabylon.edu.iq
Nawal Mohammed Dawood
Department of Metallurgical Engineering, College of Materials Engineering,
University of Babylon, Iraq.
nawalmohammed2018@gmail.com
Reception: 19/01/2022 Acceptance: 07/01/2023 Publication: 02/02/2023
Suggested citation:
K. N., Qabas, M. S., Jassim and M. D., Nawal. (2023). Micro-Arc Oxidation
Enhances Mechanical Properties and Corrosion Resistance Of Ti-6Al-7Nb
Alloy. 3C Tecnología. Glosas de innovación aplicada a la pyme, 12(1), 262-280.
https://doi.org/10.17993/3ctecno.2023.v12n1e43.262-280
https://doi.org/10.17993/3ctecno.2023.v12n1e43.262-280
3C Tecnología. Glosas de innovación aplicadas a la pyme. ISSN: 2254-4143
Ed.43 | Iss.12 | N.1 January - March 2023
262
MICRO-ARC OXIDATION ENHANCES
MECHANICAL PROPERTIES AND
CORROSION RESISTANCE OF TI-6AL-7NB
ALLOY
Qabas Khalid Naji
Biomedical Engineering Department, AL-Mustaqbal University Collage, Babil, Iraq.
qabas.khalid@mustaqbal-college.edu.iq
Jassim Mohammed Salman
Department of Metallurgical Engineering, College of Materials Engineering,
University of Babylon, Iraq.
mat.jassim.mohammed@uobabylon.edu.iq
Nawal Mohammed Dawood
Department of Metallurgical Engineering, College of Materials Engineering,
University of Babylon, Iraq.
nawalmohammed2018@gmail.com
Reception: 19/01/2022 Acceptance: 07/01/2023 Publication: 02/02/2023
Suggested citation:
K. N., Qabas, M. S., Jassim and M. D., Nawal. (2023). Micro-Arc Oxidation
Enhances Mechanical Properties and Corrosion Resistance Of Ti-6Al-7Nb
Alloy. 3C Tecnología. Glosas de innovación aplicada a la pyme, 12(1), 262-280.
https://doi.org/10.17993/3ctecno.2023.v12n1e43.262-280
https://doi.org/10.17993/3ctecno.2023.v12n1e43.262-280
ABSTRACT
Investigation results of micro-arc coating on the (Ti-7Nb-6Al) alloy were presented. It has
potential clinical value in applications such as dental implant, knee, and hip prostheses. An
electrolyte solution of (Na2CO3 + Na
2SiO3). The micro-arc oxidation (MAO) technique was
employed for in situ oxidation of Ti-6Al-7Nb surface. The wettability of a porous TiO2 covering
made up of anatase and rutile phases was investigated. The test findings revealed that the
possibility of deposition of ceramics coatings on the surface of Ti-6Al-7Nb alloy by using
voltages (400V )at different deposition times (7, 15, and 30) min. The results indicate that
ceramics layer of titanium oxide (TiO2) which is formed during coating porous and
homogenous distribution. The bioactive composition of the oxide layers can be suitable for use
as advanced biomedical implants. The coatings also revealed an increased surface roughness,
porosity, microhardness, surface wettability and corrosion resistance of the Ti-6Al-7Nb
substrate reaches to (CR= 0.1114× mpy) in Ringer’s solution and (CR= 1.03× mpy) in
Saliva’s solution with increased deposition time.
KEYWORDS
MAO; Contact Angle; Clinical Application; Oxidation Time; porosity; and Corrosion
Resistance
PAPER INDEX
10−3
10−3
ABSTRACT
KEYWORDS
1. INTRODUCTION
2. MATERIALS AND METHODS
3. RESULTS AND DISCUSSION
3.1. CHARACTERIZATION OF OXIDE SURFACE
3.2. MECHANICAL PROPERTIES:
3.3. CONTACT ANGLE TEST
3.4. ELECTROCHEMICAL BEHAVIOR OF THE ALLOY/OXIDE SYSTEMS
4. CONCLUSION
ACKNOWLEDGEMENTS
REFERENCES
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1. INTRODUCTION
Metallic are the most important technical materials, and because of their great heat
conductivity and mechanical properties, they are used as biomaterials [1]. The most
important characteristic of a metal as a biomaterial is that it does not cause an
adverse reaction when used in service, which is known as biocompatibility [2]. For
load-bearing implants and inner fixing systems, metallic materials are the most
frequently used. The primary functions of orthopedic implants systems are to restore
the load-bearing joints function that undergo to elevate levels of mechanical stress,
wear, and fatigue during ordinary activity [3]. Important orthopedic implants are
prostheses for ankle, knee, hip, shoulder, elbow joints and also need equipment like
cables, screws, plates, pins, etc. that used in the fixation of fracture [4]. Metals are
powerful, and most of them are capable to be formed into complicated forms. During
or after final formation, the required mechanical characteristics of metals can be
accomplished by heat and mechanical processing. In addition, the correct treatment of
components produced from chosen metal compositions can achieve a degree of
corrosion and wear resistance. The high tensile strength, high yield strength, fatigue
resistance and corrosion resistance are some of the features of metallic materials [5].
In medicine, titanium and its alloys have specific advantages over steels, such as low
weight, high corrosion resistance, and a wide range of applications,, low density, low
thermal conductivity, non-magnetism, processing workability, and other properties that
make it a highly appealing material [6]. Because the modulus of elasticity of titanium
and its alloys is closer to that of bone than that of stainless steels and cobalt-based
alloys, stress shielding is less of a problem [7]. Because of a TiO2 solid oxide layer, Ti
alloys are one of the most common choices in biomedical applications due to their
main characteristics. On the other hand, have poor tribological characteristics due to
their low resistance to plastic shearing, low work hardening, and lack of surface oxide
protection [2]. This titanium surface oxide layer, which is generally a few nanometres
thick, has high passivity and resistance to chemical attack [8]. Due to the coarse
microstructure of cast alloys (as seen by a high coefficient of friction), weak shear
strength, low fatigue strength, and restricted elongation compared to wrought alloys,
titanium and its alloys have a high price tag as well as a significant sensitivity to
friction and wear. As a result, extra microstructural modification is often required to
improve mechanical qualities while maintaining the product's form [9]. The surface of
biomedical implants is frequently modified to increase corrosion resistance, wear
resistance, surface roughness, and biocompatibility [10]. In addition to increasing
other desirable features, all revised surfaces should be evaluated for corrosion
behavior. In order to get implants that can survive in the human system for longer
periods of time, a thorough understanding of the interactions that occur at the atomic
level between the surface of the implant, the host, and the biological environment, as
well as all types of micromotions of the implants retained inside the human system,
should be researched further [11]. The material surface has a significant impact on the
biological environment's response to artificial medical devices [12]. Surface
modification does more than simply change the appearance of the surface; it also
enhances adhesion properties, micro cleaning, functionalization of amine, and
biocompatibility [13]. Many types of surfaces may be created using the surface
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1. INTRODUCTION
Metallic are the most important technical materials, and because of their great heat
conductivity and mechanical properties, they are used as biomaterials [1]. The most
important characteristic of a metal as a biomaterial is that it does not cause an
adverse reaction when used in service, which is known as biocompatibility [2]. For
load-bearing implants and inner fixing systems, metallic materials are the most
frequently used. The primary functions of orthopedic implants systems are to restore
the load-bearing joints function that undergo to elevate levels of mechanical stress,
wear, and fatigue during ordinary activity [3]. Important orthopedic implants are
prostheses for ankle, knee, hip, shoulder, elbow joints and also need equipment like
cables, screws, plates, pins, etc. that used in the fixation of fracture [4]. Metals are
powerful, and most of them are capable to be formed into complicated forms. During
or after final formation, the required mechanical characteristics of metals can be
accomplished by heat and mechanical processing. In addition, the correct treatment of
components produced from chosen metal compositions can achieve a degree of
corrosion and wear resistance. The high tensile strength, high yield strength, fatigue
resistance and corrosion resistance are some of the features of metallic materials [5].
In medicine, titanium and its alloys have specific advantages over steels, such as low
weight, high corrosion resistance, and a wide range of applications,, low density, low
thermal conductivity, non-magnetism, processing workability, and other properties that
make it a highly appealing material [6]. Because the modulus of elasticity of titanium
and its alloys is closer to that of bone than that of stainless steels and cobalt-based
alloys, stress shielding is less of a problem [7]. Because of a TiO2 solid oxide layer, Ti
alloys are one of the most common choices in biomedical applications due to their
main characteristics. On the other hand, have poor tribological characteristics due to
their low resistance to plastic shearing, low work hardening, and lack of surface oxide
protection [2]. This titanium surface oxide layer, which is generally a few nanometres
thick, has high passivity and resistance to chemical attack [8]. Due to the coarse
microstructure of cast alloys (as seen by a high coefficient of friction), weak shear
strength, low fatigue strength, and restricted elongation compared to wrought alloys,
titanium and its alloys have a high price tag as well as a significant sensitivity to
friction and wear. As a result, extra microstructural modification is often required to
improve mechanical qualities while maintaining the product's form [9]. The surface of
biomedical implants is frequently modified to increase corrosion resistance, wear
resistance, surface roughness, and biocompatibility [10]. In addition to increasing
other desirable features, all revised surfaces should be evaluated for corrosion
behavior. In order to get implants that can survive in the human system for longer
periods of time, a thorough understanding of the interactions that occur at the atomic
level between the surface of the implant, the host, and the biological environment, as
well as all types of micromotions of the implants retained inside the human system,
should be researched further [11]. The material surface has a significant impact on the
biological environment's response to artificial medical devices [12]. Surface
modification does more than simply change the appearance of the surface; it also
enhances adhesion properties, micro cleaning, functionalization of amine, and
biocompatibility [13]. Many types of surfaces may be created using the surface
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modification approach to control correct biological response in a specific cell/tissue
scenario, with the goal of reducing healing time and limiting harmful reactions [14].
Because titanium and its alloys have poor tribological qualities, such as low wear
resistance, they aren't recommended for use in vehicles, the implant's service life is
shortened. Surface coatings can help to solve this problem to a considerable extent.
Surface engineering can significantly improve the performance of titanium orthopedic
devices, allowing them to outperform their inherent capabilities [15]. examples of
surface modification processes: physical and chemical method, laser cladding,
thermal oxidation, plasma spray, and ion implantation [16].
2. MATERIALS AND METHODS
In the test, Ti-6Al-7Nb alloy with element composition of 6.3Al, 67Nb, 0.47Ta,
0.23Fe, 0.18O, 0.077C, 0.046N,0.0088H, and the balance Ti (wt%) were used as raw
materials. The substrate was sliced into 13 mm x 3 mm round wafers and polished
using SiC abrasive sheets ranging from 150 to 5000 grit. After that, ultrasonic cleaning
with acetone, alcohol, and deionized water was performed. The ceramic coatings
were deposited using a DC-AC homemade MAO deposition device with a voltage of
(0-500) V and a current of (0-5) A MAO with an impulse frequency of 500 Hz, current
density of 20 A/cm2, duty cycle of 10%, and oxidation durations of 7 minutes, 15
minutes, and 30 minutes at voltage 400V. Deionized water and 10 g/L sodium
carbonite and 2 g/L sodium silicate were used to make the electrolyte solution.
Following the ultrasonic processing of the MAO test sample, the sample was dried
and set aside.
3. RESULTS AND DISCUSSION
3.1. CHARACTERIZATION OF OXIDE SURFACE
In Fig.1 (a) The XRD results proved the deposition of titanium oxide layer after
MAO on the surface of the Ti-6Al-7Nb alloy substrate at 7min. The formation of TiO2
layer on the surface of specimen A3 has crystalline phases: rutile (tetragonal) and
anatase (tetragonal) phases also the (α-HCP) and (β-BCC) return to the Ti-alloy. The
peaks of rutile TiO2 (200), (211), and (202) at 2ϴ° (39.3, 54.2, and 76.0) and those of
titania crystals structures (anatase) (101), (103), and (200) at 2ϴ
° (25.9, 37.9, and
48.3) strength of the Ti-6Al-7Nb alloy peaks reduced compared to the untreated Ti
sample. This is due to the crystal structure of both types, the energy gaps for anatase
are more than those of rutile, this makes the anatase more pores and it’s used in
optical application while the rutile is with low energy gape and more stable at high
temperatures and more important for medical application [17]. Limiting voltage
increased, perhaps due to oxide layer formation as illustrated in Fig.1 (b) at 15min.
For the highest deposition time in Fig.1 (c). the presence of anatase indicates that
throughout the MAO process, a significant oxidation reaction took place on a titanium
surface. As a result, the combination of anatase and rutile crystal phases in the coated
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Ti-alloys specimen developed in this work is expected to have a positive influence on
Ti-alloy bioactivity by enhancing their osteogenic properties. It is also suggests that
predominantly anatase is created at lower forming voltages, however because
anatase, as a metastable phase, gradually converts into rutile at higher temperatures
as dielectric breakdown processes increase, the mixture of anatase and rutile phases
develops at increased deposition time [18].
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(a)
(b)
(c)
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Ti-alloys specimen developed in this work is expected to have a positive influence on
Ti-alloy bioactivity by enhancing their osteogenic properties. It is also suggests that
predominantly anatase is created at lower forming voltages, however because
anatase, as a metastable phase, gradually converts into rutile at higher temperatures
as dielectric breakdown processes increase, the mixture of anatase and rutile phases
develops at increased deposition time [18].
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(a)
(b)
(c)
Figure 1. XRD of MAO Process with different time (a) coating at 7min, (c) coating at 15min,
and (d) coating at 30min.
The FESEM results of microstructure coated specimen from Fig.2 which show that
for surface morphology of the oxide layer TiO2 to the Ti-6Al-7Nb alloys at different
magnifications treated by MAO process relatively rougher and exhibited a grainy
structure with limited amount of pores with different sizes by the spark discharges.
Micro-pores and submicron-pores were visible in the MAO coating, with the micro-
pores having a roughly round or elliptical form like a volcanic vent [19].
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(c)
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Figure 2. FESEM Micrographs of TiO2 coating MAO process at different magnification and
time.
A typical porous structure was also found in coating of sample at 15min. Pores with
miximum diameters and homogenous distribution can be observed on the surface of
the Ti-alloy. The diameters of the such holes and the surface roughness grew as the
voltage rose; after 30 min of treatment, the pores diameters increased, and the coated
surface progressively became rough. The oxide layer on both materials is formed by
several micro-protrusions with uniformly scattered pores with diameters varying from
sub-micron to few microns. When compared to a polished surface that hasn’t been
covered, the presence of this porosity improves osseointegration because the pores
function as sites for bone tissue formation, hence improving anchoring [20]. The
FESEM cross-sectional morphology of TiO2 coating layer has a regular thin film
structure in thickness with more compact, homogeneity, and full adhesion between the
coating and the underlying substrate, as shown in Fig. 3 (a). each sample’s coating
layer has a compact diffusion layer in contact with the substrate and an external
porous conversion zone with discharge channels make up the two sections. The
average thickness of the diffusion layer remains constant throughout the procedure.
the average thickness of the external porous conversion layer rises as the deposition
duration increased, from 1.94µm at 7 min to 5.54µm at 30 min, as shown in Fig.3 (b
and c).
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Micro-pores
(d)
(e)
(f)
Nested number
(i)
(h)
(g)
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Figure 2. FESEM Micrographs of TiO2 coating MAO process at different magnification and
time.
A typical porous structure was also found in coating of sample at 15min. Pores with
miximum diameters and homogenous distribution can be observed on the surface of
the Ti-alloy. The diameters of the such holes and the surface roughness grew as the
voltage rose; after 30 min of treatment, the pores diameters increased, and the coated
surface progressively became rough. The oxide layer on both materials is formed by
several micro-protrusions with uniformly scattered pores with diameters varying from
sub-micron to few microns. When compared to a polished surface that hasn’t been
covered, the presence of this porosity improves osseointegration because the pores
function as sites for bone tissue formation, hence improving anchoring [20]. The
FESEM cross-sectional morphology of TiO2 coating layer has a regular thin film
structure in thickness with more compact, homogeneity, and full adhesion between the
coating and the underlying substrate, as shown in Fig. 3 (a). each sample’s coating
layer has a compact diffusion layer in contact with the substrate and an external
porous conversion zone with discharge channels make up the two sections. The
average thickness of the diffusion layer remains constant throughout the procedure.
the average thickness of the external porous conversion layer rises as the deposition
duration increased, from 1.94µm at 7 min to 5.54µm at 30 min, as shown in Fig.3 (b
and c).
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Micro-pores
(d)
(e)
(f)
Nested number
(i)
(h)
(g)
Figure 3. Cross section of MAO coatings: (a) at 7min, (b) 15min, and (c) 30min.
The formation of crossing pores and big pores distributed along the whole
thickness. Generally, the coating thickness is increased with increasing deposition
time because the voltage on the sample could not reach the sparking threshold, and a
thin layer of oxide film quickly formed on the sample surface because of anodic
oxidation. When the oxidation time was increased, the sparking voltage was reached
and the energy rose; consequently, some discharge channels on the specimens
became evident. Oxide film formed on the inner and outer surfaces of the discharge
channel as the reaction product erupted along the channel. The oxide coating
thickened when the oxidation duration, and energy were increased. Furthermore, the
molten oxide spilled over the discharge tube, immediately cooled, and was deposited
on the surface. The process was repeated until the end of the oxidation reaction,
causing incessant growth of the oxide film [21]. The Presents of schematic data of
EDS results for MAO TiO2 coatings with different times on containing Ti, O, Al, and Nb
ions. EDS analysis showed that increasing of time up to 30 min had it effects on the
content of oxide layer as shown in Fig.4 coated with different times. As a result of the
presence of Ti and O2 components in the coatings, TiO2 layers with varied weights of
these modification elements.
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(a)
1.94μm
(b)
2.67μm
(c)
5.45μm
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Full Area
(b)
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(a)
Full Area
(b)
Full Area
Figure 4. EDS result of (a) coating at 7min, (b) coating at 15min, and (c) coating at 30min.
Results of AFM analysis are given in Fig.5, the differences in surface topography
between the substrate and different coatings in 2D and 3D, where observed an
increase in the roughness of the TiO2 coatings because of the phenomena of micro-
discharge resulting from the nature of the MAO process [22]. The Ra of coatings were
more than those of the substrates and increased with time roughness also increased
from (7.19nm) for 7min, (12.6nm) for 15min, and (18.8nm) for 30min because
increase oxidation time.
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(c)
(a)
(b)
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Figure 5. AFM results of samples (a) base, (b) coating at 7min, (c) coating 15min, and (d)
coating 30min.
3.2. MECHANICAL PROPERTIES:
Results of micro-hardness at TiO2 of the coated samples at a load of 50gm (0.49N)
using a fixed loading duration of (15sec.) in Fig.6. In general, it can be observed that
the hardness of coated samples is improved by MAO process. The value of hardness
equal to (268.55 HV at 7min) are significantly higher than that of Ti-6Al-7Nb alloy, and
the increase hardness value with increase the deposition time to the (311.5 at 15min)
because the production of dense oxide layer, which is attributed to formation of
thermal micro arcs during MAO and increase thickness of ceramic coating [17]. The
MAO treated sample’s greater standard deviations might also be attributed to their
higher surface porosity. Which resulted in lower leading hardness at the interface
(287.3HV at 30min).
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(c)
(d)
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Figure 5. AFM results of samples (a) base, (b) coating at 7min, (c) coating 15min, and (d)
coating 30min.
3.2. MECHANICAL PROPERTIES:
Results of micro-hardness at TiO2 of the coated samples at a load of 50gm (0.49N)
using a fixed loading duration of (15sec.) in Fig.6. In general, it can be observed that
the hardness of coated samples is improved by MAO process. The value of hardness
equal to (268.55 HV at 7min) are significantly higher than that of Ti-6Al-7Nb alloy, and
the increase hardness value with increase the deposition time to the (311.5 at 15min)
because the production of dense oxide layer, which is attributed to formation of
thermal micro arcs during MAO and increase thickness of ceramic coating [17]. The
MAO treated sample’s greater standard deviations might also be attributed to their
higher surface porosity. Which resulted in lower leading hardness at the interface
(287.3HV at 30min).
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(c)
(d)
Figure 6. Relationship between the deposition time and the micro-hardness of coatings at
TiO2 by MAO process.
Wear resistance is one of the most significant implant mechanical properties that
wear failure contributes from all implant’s failure reasons. Wear mass loss of test
specimens under 10N stress and several times (5, 10, 15, 20, 25, and 30) min were
used to evaluate the wear rates produced by pin-on disc sliding wear tests. Generally,
it can be observed that the weight loss increased with increasing of loading time. Fig.7
shows that high wear rate of Ti-6Al-7Nb alloy substrate comparison, with coating. The
ceramic oxide layers were found to have high wear resistance, resulting in a lower
wear rate in the samples. ceramic coatings deposited TiO2 by MAO process, gave the
best wear resistance and low wear rate at 7min. Due to improved
hardness by presence of the alpha and beta titanium phases and presence of
modified elements could be reduce the friction and increase wear resistance of
coating by reduce wear rate. The intensity of micro-discharges rises as the applied
duration increases (30min), resulting in an increase in coating porosity owing to a
decrease in coating electric resistance. The coating porosity distribution has an impact
on both mechanical and tribological properties [23]. This is supported by the current
findings, which reveal that the coating has greater wear rates equal
after passing 30min.
Microhardness HV
100
125
150
175
200
225
250
275
300
325
350
Number of samples
1
2
3
4
Base
7min
15min
30min
(2.78 ×10−5)
(9.46×10−6)
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Figure 7. Relationship between wear rate and test time for Ti-6Al-7Nb alloy substrate and
MAO process with different deposition time.
3.3. CONTACT ANGLE TEST
The contact angle, which is an essential measuring tool for determining material
surface wettability was also discovered to be a key factor in increasing the bioactivity
of titanium surfaces. Fig.8 show the contact angles tested of TiO2 by MAO coatings
prepared at various deposition times in Ringer’s and Saliva’s solution. With increased
surface roughness and porosity, the specimen’s contact angle reduces considerably
following MAO treatment. The contact angle reached value to (56.74° at 7min in
ringer’s solution and 54.7° in saliva’s solution) and decreased with increase deposition
time reached to (11° at 30min in ringer solution and 13.1° in saliva solution). The MAO
treatment resulted in an uneven coating surface, increased roughness, increased
absorbability, and decreased contact angles, all of which together affected the surface
energy; and the OH and O2 oxygen-containing groups produced on the coated
surfboard. These factors combined to increase the wettability of the MAO treated
because a large number of micro/nano-pores formed on the oxidation coating surface
caused its specific surface area to oxidation coating increase, which benefited water
retention [24].
Wear rate (g/m)
0
0,0001
Time (min)
0
8
15
23
30
base
7min
15min
30min
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Figure 7. Relationship between wear rate and test time for Ti-6Al-7Nb alloy substrate and
MAO process with different deposition time.
3.3. CONTACT ANGLE TEST
The contact angle, which is an essential measuring tool for determining material
surface wettability was also discovered to be a key factor in increasing the bioactivity
of titanium surfaces. Fig.8 show the contact angles tested of TiO2 by MAO coatings
prepared at various deposition times in Ringer’s and Saliva’s solution. With increased
surface roughness and porosity, the specimen’s contact angle reduces considerably
following MAO treatment. The contact angle reached value to (56.74° at 7min in
ringer’s solution and 54.7° in saliva’s solution) and decreased with increase deposition
time reached to (11° at 30min in ringer solution and 13.1° in saliva solution). The MAO
treatment resulted in an uneven coating surface, increased roughness, increased
absorbability, and decreased contact angles, all of which together affected the surface
energy; and the OH and O2 oxygen-containing groups produced on the coated
surfboard. These factors combined to increase the wettability of the MAO treated
because a large number of micro/nano-pores formed on the oxidation coating surface
caused its specific surface area to oxidation coating increase, which benefited water
retention [24].
Wear rate (g/m)
0
0,0001
Time (min)
0
8
15
23
30
base
7min
15min
30min
https://doi.org/10.17993/3ctecno.2023.v12n1e43.262-280
Figure 8. Results of contact angle in TiO2 at 400V in Ringer’s and Saliva’s solution.
3.4. ELECTROCHEMICAL BEHAVIOR OF THE ALLOY/OXIDE
SYSTEMS
The potentiodynamic polarization curves for the Ti-6Al-7Nb alloy substrate and
TiO2 coated by MAO process samples in Ringer's solution at 37°C±1 at various times
in Fig. 9. Tafel extrapolation is used to calculate the corrosion current densities (icorr.)
and corrosion potentials (Ecorr.) using potentiodynamic curves, and corrosion rates
(CR) were also included in Table.1. It can be seen from the results obtained in the
uncoated substrate has a greater corrosion current density (icorr= 6.8284 A/cm2) of
thus lowest corrosion resistance because of the occurrence of metal ions dissolution
on the surface of the uncoated substrate. The corrosion current density and corrosion
rate of all coated samples by MAO decrease after TiO2 coating, indicating that the
TiO2 coating offers a protective layer on the substrate surface that reduces corrosion
rate. The lowest corrosion current (icorr. = 2.8161 A/cm2) and increased corrosion
potential of the specimens are achieved when coating for 7 min, and this result has a
reduced corrosion rate equal to (CR= 3.48×
10-3 mpy), indicating that corrosion
resistance is improved. Furthermore, the film's surface structure influences the
material's corrosive qualities. Materials with denser and thicker oxide layers have a
lower corrosion current density and a lower corrosion rate (icorr. = 0.0902A/cm2), and
a lower corrosion rate (CR=0.111×10-3 mpy) at (15min).
Contect angle degree
0
10
20
30
40
50
60
70
Times min
0
7
15
30
Ringer's solution
saliva's solution
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Figure 9. Potentiodynamic polarization curves of TiO2 coated by MAO process and base at
different time in Ringer’s solution.
The porosity affected the corrosion behavior of highly porous where decreased
corrosion resistance because increased (icorr. =1.3643 µ
A/cm2) and (CR
=1.686×
10-3mpy). The corrosion behavior of porous metallic materials has in the
present work, the oxygen/air entrapped in the most inner pores neither the difficulty of
electrolyte penetration into these pores, may result in various passive states on the
native oxide surface. The surface area, on the other hand, has minimal influence on
the corrosion rate of porous materials. Although Ti is known for its strong resistance to
localized corrosion, and localized breakdown of its passive coating occurs at
substantially higher potentials, the increased corrosion density with greater porosity
level is due to the larger surface area in contact with the electrolyte, but fissures or
restrictions to the flow of species into the connected pores can result in corrosion
rates that are not proportional to the real contact surface area. Although these
challenges may result in faster corrosion rates, good passivation qualities for porous
Ti structures have been documented in the literature. On the other hand, increased
porosity was also associated with a decreased susceptibility to corrosion (less
negative Ecorr.), as interconnected pores encouraged the free flow of ionic species,
whereas isolated pores trapped the electrolyte and depleted the oxygen supply,
leading to a thinner oxide film. Aside from species free movement, air entrapment, and
electrolyte penetration [25].
Potenial (µV)
-0,95
-0,1
Log i (μA)
0
0
0
0
0,0001
0,001
Base
7min
15min
30min
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Ed.43 | Iss.12 | N.1 January - March 2023
276
Figure 9. Potentiodynamic polarization curves of TiO2 coated by MAO process and base at
different time in Ringer’s solution.
The porosity affected the corrosion behavior of highly porous where decreased
corrosion resistance because increased (icorr. =1.3643 µA/cm2) and (CR
=1.686×10-3mpy). The corrosion behavior of porous metallic materials has in the
present work, the oxygen/air entrapped in the most inner pores neither the difficulty of
electrolyte penetration into these pores, may result in various passive states on the
native oxide surface. The surface area, on the other hand, has minimal influence on
the corrosion rate of porous materials. Although Ti is known for its strong resistance to
localized corrosion, and localized breakdown of its passive coating occurs at
substantially higher potentials, the increased corrosion density with greater porosity
level is due to the larger surface area in contact with the electrolyte, but fissures or
restrictions to the flow of species into the connected pores can result in corrosion
rates that are not proportional to the real contact surface area. Although these
challenges may result in faster corrosion rates, good passivation qualities for porous
Ti structures have been documented in the literature. On the other hand, increased
porosity was also associated with a decreased susceptibility to corrosion (less
negative Ecorr.), as interconnected pores encouraged the free flow of ionic species,
whereas isolated pores trapped the electrolyte and depleted the oxygen supply,
leading to a thinner oxide film. Aside from species free movement, air entrapment, and
electrolyte penetration [25].
Potenial (µV)
-0,95
-0,1
Log i (μA)
0
0
0
0
0,0001
0,001
Base
7min
15min
30min
https://doi.org/10.17993/3ctecno.2023.v12n1e43.262-280
Table.1. Electrochemical parameters of base and TiO2 coated MAO at different time in
Ringer’s solution.
The corrosion behavior in Saliva’s solution of the alloy and coated samples TiO2 by
MAO process in Fig.10 respectively, and Table.2 it is clear, that the specimen showed
relatively similar behavior to that observed in saliva's solution such as MAO process
for alloy improved corrosion resistance because of reduction in corrosion current
[26,27]].
Figure 10. Potentiodynamic polarization curves of TiO2 coating by MAO process and base at
different time in Saliva’s solution.
Table 2. Electrochemical parameters of base and TiO2 coated MAO at different time in
Saliva’s solution
Parameters of coating icorr. (µA/cm2) Ecorr. (mV) Rate of Corrosion
(mpy)×10-3
Enhancement percentage
(%)
base 6.8284 -533 8.437 /
7min 2.8161 -342 3.480 58.7
15min 0.0902 -341 0.111 98
30min 1.3643 -390 1.686 80
Potianial (mV)
-0,9
-0,1
Log i (µA)
0
0
0
0
0,0001
0,001
0,01
Base
7min
15min
30min
Parameters of coating icorr. (µA/cm2) Ecorr. (mV) Rate of Corrosion
(mpy)×10-3
Enhancement
percentage (%)
base 7.103 -579 8.777 /
7min 3.479 -491 4.3 51
15min 0.833 -472 1.03 88.26
30min 1.475 -392 1.8 79.5
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Ed.43 | Iss.12 | N.1 January - March 2023
277
4. CONCLUSION
In the current study, the TiO2 coating has been deposited on the surface of
Ti-6Al-7Nb alloy successfully by using micro-arc oxidation process for biomedical
applications.
1. The TiO2 layer formed on Ti-6Al-7Nb alloy substrate material using MAO methods
has circular micro holes in rough and volcanic structures because of continuous
micro discharges occurring during the process. Rutile TiO2 and anatase TiO2
phases are determined on the material surface following the XRD analysis.
2.
The substrate's surface roughness plays has an important role to improving
coating-substrate adhesion. AFM topography shows homogeneous and dense at
(30 min).
3. EDS results showed that the ratio of Ti/O increased with time at coating.
4. The apparent contact angle somewhat reduces following treatment at various times.
The surface morphology and composition of the MAO coatings may be the cause of
the MAO coatings' considerable shift in apparent contact angle. Due to its smaller
pores, the wettability of the MAO coating created at lower roughness may be
underestimated. It's possible that the wettability of the MAO coating generated at
reduced roughness is overestimated. The wettability of the MAO coating generated
at higher roughness may be overstated since no gas is trapped and the liquid/solid
interface is rougher.
5.
The potentiodynamic polarization results that Ti-6Al-7Nb base alloy substrate and
TiO2 at different times in Ringer's and Saliva's solutions; the best result equal
(icorr.=0.0902µA/cm2) in Ringer's solution and (icorr.= 0.833µ
A/cm2) in Saliva’s
solution compared to the uncoated sample.
ACKNOWLEDGEMENTS
The Authors are grateful for the University of Babylon for their help. Special thanks
for, AL Mustaqbal University Collage, Biomedical Engineering Department.
REFERENCES
(1) L. Thair, U. K. Mudali, N. Bhuvaneswaran, K. G. M. Nair, R. Asokamani, and B.
Raj. (2002). Nitrogen ion implantation and in vitro corrosion behavior of as-
cast Ti-6Al-7Nb alloy. Corros. Sci., 44(11), 2439-2457.
(2) G. Wu, P. Li, H. Feng, X. Zhang, and P. K. Chu. (2015). Engineering and
functionalization of biomaterials via surface modification. J. Mater. Chem.
B, 3(10), 2024-2042.
(3) A. T. Sidambe. (2014). Biocompatibility of advanced manufactured titanium
implants-A review. Materials (Basel), 7(12), 8168-8188.
(4) M. kawano, Y. Takeda, K.Ogasawara. (2015). Pathological Analysis of Metal
Allergy to Metallic Materials, 305-321.
https://doi.org/10.17993/3ctecno.2023.v12n1e43.262-280
3C Tecnología. Glosas de innovación aplicadas a la pyme. ISSN: 2254-4143
Ed.43 | Iss.12 | N.1 January - March 2023
278
4. CONCLUSION
In the current study, the TiO2 coating has been deposited on the surface of
Ti-6Al-7Nb alloy successfully by using micro-arc oxidation process for biomedical
applications.
1. The TiO2 layer formed on Ti-6Al-7Nb alloy substrate material using MAO methods
has circular micro holes in rough and volcanic structures because of continuous
micro discharges occurring during the process. Rutile TiO2 and anatase TiO2
phases are determined on the material surface following the XRD analysis.
2. The substrate's surface roughness plays has an important role to improving
coating-substrate adhesion. AFM topography shows homogeneous and dense at
(30 min).
3. EDS results showed that the ratio of Ti/O increased with time at coating.
4. The apparent contact angle somewhat reduces following treatment at various times.
The surface morphology and composition of the MAO coatings may be the cause of
the MAO coatings' considerable shift in apparent contact angle. Due to its smaller
pores, the wettability of the MAO coating created at lower roughness may be
underestimated. It's possible that the wettability of the MAO coating generated at
reduced roughness is overestimated. The wettability of the MAO coating generated
at higher roughness may be overstated since no gas is trapped and the liquid/solid
interface is rougher.
5. The potentiodynamic polarization results that Ti-6Al-7Nb base alloy substrate and
TiO2 at different times in Ringer's and Saliva's solutions; the best result equal
(icorr.=0.0902µA/cm2) in Ringer's solution and (icorr.= 0.833µA/cm2) in Saliva’s
solution compared to the uncoated sample.
ACKNOWLEDGEMENTS
The Authors are grateful for the University of Babylon for their help. Special thanks
for, AL Mustaqbal University Collage, Biomedical Engineering Department.
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Allergy to Metallic Materials, 305-321.
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