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 TecnologÃa. 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 TecnologÃa. 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 TecnologÃa. 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
https://doi.org/10.17993/3ctecno.2023.v12n1e43.262-280
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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
https://doi.org/10.17993/3ctecno.2023.v12n1e43.262-280
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 2Ï´
° (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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(b)
(c)
(a)
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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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(a)
Full Area
(b)
Full Area
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