Keywords
dental implant - field emission scanning electron microscopy - hydroxylapatite - magnetron
sputtering - titanium - Vickers hardness
Introduction
Hydroxyapatite (HA) and β tricalcium phosphate are acceptable bioactive materials
because of their apatite formation abilities.[1]
[2]
[3] The inorganic HA, Ca10(PO4)2(OH)2, is an essential biomaterial component of human hard tissues as bones and teeth.
Owing to its inherent biocompatibility, it is used for medical and dental implant
prosthesis coatings.[4]
[5] The role of a coating material is to protect the body organs from metal ions that
may be released by metallic implant materials. Moreover, the ability to induce implant
bone bonding is an important factor to consider when selecting a coating material.
Hence, the coating material should be pore free, dense, and with a strong adhesion
to their substrate implant material.[6]
[7] Metal surfaces can be coated by bioactive materials via several methods, such as,
dip coating, electrophoretic deposition, thermal sputtering, colloidal solution deposition,
magnetron sputtering, plasma spraying, sol–gel technique, and pulsed laser deposition,
to enhance bone healing and metal implant integration.
Coating thickness is a crucial factor that will determine the success or failure of
a direct implantation procedure. The aforementioned implantation methods can achieve
a coating thickness ranging from 0.05 to 200 μm.[8] For instance, sand blasting with an acid provides a coat with numerous micropores
on the substrate surface to stimulate bone growth. However, this technique requires
a healing period of ~3 to 6 months.[9] A commonly applied method is plasma-sprayed coating of implant surface using HA
to improve biocompatibility and rapid bone tissue growth.[10] However, this technique has several disadvantages, such as it produces a thick HA
coating layer ranging from 75 to 150 µm.[11] Shortages of this coat leads to delamination and adhesion and cohesion failure with
porosity,[12] as well as a decrease in the ability of the coat to withstand shear fatigue.[13] These coat thickness-related complications can result in implant failure due to
breakage of the coating film. The radiofrequency (RF) sputtering technique is an alternative
method that involves collision of high-energy electrons with gas molecules or atoms
to form positive ions for bombarding the HA target after ionization and excitation–relaxation
reaction, which in turn deposit the target material onto the implant substrate material.[14] RF–magnetron sputtering provides elemental composition coating by sintering the
target made from a homogeneous mixture of phosphates to provide a high-purity coating.[15] This method provides a very thin coat (nanoscale films); a thin coat may degrade
rapidly, a condition that is detrimental to dental implants.[1]
[16]
[17] Therefore, the objective is to create thick, well-adhered HA coat that will allow
sufficient time for osseointegration and will not easily separate during the procedure.
An HA coating of a few micrometers in thickness is ideal because such coating may
not exfoliate as rapidly as thinner or thicker HA coatings.[18]
[19]
This study used long durations of RF–magnetron sputtering deposition to produce a
microscale HA bioceramic layer and evaluated its effects on grade 1 commercially pure
(CP) titanium (Ti) implant materials.
Materials and Methods
Sample Preparation
Grade 1 CP Ti (99.9% purity) (Baoji Jinsheng Metal Material, China) supplied as annealed
rods with a dimension of 1,000 mm × 6 mm (length × diameter) were used to prepare
the substrate. Two different substrate specimens were prepared: a disc-shaped specimen
with a dimension of 6 mm × 2 mm (diameter × thickness) was prepared by cutting the
Ti rods using a water jet machine; the other was a cylindrical specimen (root form
implant) with a dimension of 3 mm × 6 mm (diameter × length) that was fabricated using
a turning machine, to demonstrate that this technique applicable for every implant
surface design, root or screw. The disc specimens were used for in vitro coating optimization and characterization, whereas the cylindrical specimens were
utilized for in vivo biomechanical and histological evaluations. Before sputtering, all specimens were
gradually polished using a silicon carbide (SiC) polishing paper with increasing grit
sizes of 500, 800, 1,200, 2,000, and 2,400. The specimens were cleaned with absolute
ethanol in an ultrasonic cleaning device for 30 minutes and then air dried. The specimens
were etched using a solution of distilled water, 6% HNO3 and 2% HF for 3 minutes to remove the oxide layer. After etching, the specimens were
rinsed with acetone for coating.[20]
[21] A custom-made stainless steel plat holder was designed to hold the circular substrates
and fix the rod substrates (root form) vertically in the vacuum chamber during sputtering
to ensure uniform coating.
Hydroxyapatite Target Preparation
The targets for magnetron sputtering (50 mm × 3 mm, diameter × thickness) were custom
fabricated and sintered from HA powder (99.99% purity, the particle size was > 30
nm, ~2.13 g/cm in density, and had a stoichiometric calcium [Ca]/phosphorus [P] ratio
of 1.6). The targets were produced by vacuum hot-press sintering (Jiangyin Entret
Coating Technology Co.). The pressure was 50 MPa, and the sintering temperature was
1,100°C.[22] A protective copper cover was used to protect the targets from cracks during sputtering
process because the power of device suddenly breakdown to zero when plasma strikes
to the broken area (cracks) of the target as sparking occurs in that area.
Sputtering Process
Several sputtering trials were conducted in a pilot study to determine the suitable
sputtering parameters for optimal coating deposition for 90 minutes. A magnetron sputtering
device (Torr International Inc., United States) was used with different parameters,
including magnetron power, working pressure, and target-to-substrate distance ([Table 1]
[Fig. 1]). Thickness was calculated at six different target-to-substrate distances (20, 30,
40, 50, 60, and 70 mm) for each specific magnetron power that was used. The optimum
sputtering parameters were identified by measuring coat thickness on a square quartz
microscope slide (25.4 mm × 25.4 mm × 1 mm thick; Ted Pella, Inc., United States)
using a laser ellipsometer (SE 800, Sentech, Germany). The quartz slides provide extremely
high purity, optical, and ultraviolet transparency, all these are important when a
laser ellipsometer was used to measure coat thickness.
Table 1
Sputtering process parameters
|
Sample
|
Magnetron power (W)
|
Working pressure (Torr)
|
The target-to-substrate distance (mm)
|
Coat thickness (nm)
|
Observation results
|
|
A
|
100
|
3 x 10–3
|
20–70
|
90 ± 11
|
|
|
B
|
120
|
3 x 10–3
|
20–70
|
120 ± 13
|
|
|
C
|
130
|
3 x 10–3
|
20–70
|
180 ± 19
|
|
|
D
|
150
|
3 x 10–3
|
20–70
|
192 ± 22
|
Coat was uniform, dense, and compact ([Fig. 1A]
[1B])
|
|
E
|
>150
|
|
|
|
Crack of the target and change in stoichiometry
|
|
F
|
<100
|
|
>70
|
|
Coat was delaminated, adhesive and cohesive failure with porosity ([Fig. 1C]
[1D])
|
Fig. 1 FESEM images for substrates surfaces after sputter processing (A, B) as a function
of sputter at power level of 150 W and the target-to-substrate distance was 40 mm
(C, D) as a function of sputter at power level of 70 W and the target-to-substrate
distance was 80 mm. FESEM, field emission scanning electron microscopy.
Final Coating Process
The results of the pilot study showed that sample D was the best sample to obtain
the optimum coat. Its sputtering parameters are listed in [Table 1].
The pressure of the sputtering chamber was set to less than 1.5 × 10–5 Torr as the base pressure, and the working pressure was set 1.5 × 10–3 Torr. Argon gas was then introduced into the device as sputtering gas at a flow rate
of 150 cm /min. The coating process was performed by RF sputtering at a power level
of 150 W. During sputtering, the magnetron power applied to the targets was gradually
increased by 10 W every 3 minutes and maintained at 150 W until the end of the sputtering
process. The deposition heat was fixed at 100°C; the temperature of the substrate
during the sputtering procedure was 100°C. During the sputtering process, water was
ejected to cool the targets. The entire procedure (deposition time) lasted for 22
hours to obtain a coat more than 5 µm in thickness. The substrates were moved in a
rotary motion during deposition at 10 rpm. The target-to-substrate distance was 30
to 40 mm. Afterward, all HA substrates were sintered in a sintering oven (Nabertherm,
Sintering Furnace HTCT01-16, United States). The heat treatment was conducted in air,
and the temperature was gradually increased by 10°C/min up to 550°C for 1 hour ([Fig. 2]).[23]
Fig. 2 (A) HA target, (B) holder of the target, (C) schematic diagram of experimental sputtering
setup, and (D) Ti- and HA-coated substrates. HA, hydroxyapatite; Ti, titanium.
The discs were molded vertically, and the yielded blocks were ground using a 400 grit
SiC paper and then finished and polished with a 2,000 grit SiC paper to expose the
substrate–coat area.[20]
[24]
Coating thickness and surface morphology were investigated via field emission scanning
electron microscopy (FESEM) (MIRA3 Tescan, Czech Republic). SEM–energy dispersive
X-ray (EDX) analysis was conducted for elemental analysis. Elemental distribution
was determined via EDX mapping. Vickers hardness (VH) test (Buehler Micromet 5103,
United States) was performed to measure coat hardness.
Results
Surface Microstructure
Ti- and HA-coated substrates were examined via FESEM. The HA coating surface was continuous,
crack free, and the HA particles were uniformly distributed with few aggregates of
different sizes ([Fig. 3]).
Fig. 3 FESEM surface topographic image: (A) uncoated Ti and (B–I) FESEM surface topographic
images of coated substrate at different magnifications of x 500, x 1,000, x 5,000,
x 10,000, x 20,000, x 35,000, x 70,000, and x 140,000. FESEM, field emission scanning
electron microscopy; Ti, titanium.
Chemical Analysis of Ti and Hydroxyapatite-Coated Substrates
EDX analysis of the control (uncoated) Ti surface revealed that the surface consisted
of Ti and oxygen (O). By comparison, the analysis revealed the presence of HA particles
on the surface of the HA-coated substrate and that it comprised Ca, Ti, P, and O.
Moreover, the coating was not contaminated by other elements ([Table 2]).
Table 2
EDX results of the coated substrate
|
Element
|
Mass%
|
Atom%
|
|
Abbreviations: Ca, calcium; EDX, energy dispersive X-ray; O, oxygen; P, phosphoru s;
Ti, titanium.
|
|
O
|
52.43
|
71.60
|
|
P
|
16.04
|
11.31
|
|
Ca
|
30.41
|
16.58
|
|
Ti
|
1.13
|
0.51
|
|
100
|
100
|
Ca/P Ratio of Hydroxyapatite Coating
The Ca/P ratio in the coat layer (16.58/11.31) was ~1.5 ([Table 2]), which was close to that in Ca10(PO4)6·2H2O (10/6)[25] because the HA target had a stoichiometric Ca/P ratio of 1.6. Therefore, RF–magnetron
sputtering did not alter the Ca/P ratio ([Fig. 4]).
Fig. 4 SEM-EDX maps of coated Ti substrate surface. SEM-EDX, scanning electron microscopy–energy
dispersive X-ray; Ti, titanium.
The cross-section line elements at the HA–Ti interface were analyzed via EDX line
scans of the interface at different areas ([Fig. 5]).
Fig. 5 (I) FESEM cross-section image of the coat, (II) EDX compositional line scan data of the Ti–HA layer interface at the selected areas;
(A) HA coat, (B) interface, and (C) Ti substrate. EDX, energy dispersive X-ray; FESEM, field emission scanning electron
microscopy; HA, hydroxyapatite; Ti, titanium.
EDX elemental analysis of the coat area revealed abundant Ca, P, and O atoms in the
coating. EDX elemental analysis showed that the interface consisted of Ti, Ca, P,
and O atoms. EDX elemental analysis revealed that the Ti substrate comprised abundant
Ti atoms. From the Ti substrate area to the coating area, the EDX line scans showed
that the levels of Ca and P increased and that of Ti decreased.
Coat Thickness and Interface Microstructure
The average thickness of the coating layer was 7 µm ([Fig. 6]). FESEM images of the interface showed that the Ti substrate embedded the HA particles
well such that the HA particles were trapped on the surface layer of Ti and bonded
(no gap) with the Ti substrate. The thickness of 7 ± 0.9 µm was obtained as a result
of long sputtering time of 22 hours at 150 W. A coating thickness of 1 to 50 µm does
not affect the fatigue strength of the substrate.[11] Aside from the problems commonly encountered by various coating techniques, a study
reported that coating with pure HA results in poor adhesion with the substrate and
deterioration of several mechanical properties over time when the coating is immersed
in simulated body fluids.[26] The FESEM image showed that the coat layer was highly uniform, dense, and compact,
without visible defects (microcracks and pores) and with good adhesion to the Ti substrate
([Fig. 6]).
Fig. 6 FESEM images of cross-section substrate at Ti–HA interface area at the sputtering
power of 150 W for 22 hours, at different magnifications of 5,000 and 10,000. FESEM,
field emission scanning electron microscopy; HA, hydroxyapatite; Ti, titanium.
EDX mapping of the cross-section of the substrate showed that the HA particles were
embedded in the Ti substrate at the interface area with no gaps ([Fig. 7]). Elemental analysis of the deep structure (not the interface line) of a typical
cross-section of the HA-coated and Ti interfaces revealed the presence of Ca and P,
which are the main elements of HA.
Fig. 7 SEM-EDX maps of HA-coated cross-section Ti substrate surface. HA, hydroxyapatite;
SEM-EDX, scanning electron microscopy–energy dispersive X-ray; Ti, titanium.
Surface Roughness
In this study, three-dimensional roughness parameters (nm) were investigated via atomic
force microscopy (Nanoscope, California, United States). The mean values of five substrates
were each tested at three different positions ([Table 3]). The HA coat of Ti had surface roughness significantly higher than uncoated Ti
substrates.
Table 3
Topographic analyses roughness (nm) of the Ti substrate and HA coat
|
Substrates
|
Sa (nm)
|
Sdr (%)
|
Sdq (/nm2)
|
Sq (nm)
|
|
Abbreviations: HA, hydroxyapatite; Sa, roughness average; Sdq, slop root mean square; Sdr, increment of the interfacial surface area relative to a flat plane baseline; Sq, height root mean square of the surface; Ti, titanium.
Note: Result indicated highly significant difference between uncoated Ti and HA coat
using t-test at p-value < 0.05.
|
|
Untreated Ti
|
9.36 ±1.2
|
0.5 ± 0.06
|
0.01 ± 0.005
|
0.27 ± 0.03
|
|
HA coat
|
18.28 ± 2.01
|
3.07 ± 0.1
|
0.261 ± 0.06
|
3.69 ± 0.4
|
Vickers Hardness of the Coating
The mean values of the hardness of the HA-coated Ti substrate was (266.7 VH), which
was significantly higher than that of the Ti substrate (1,118 VH). The difference
between the coated and uncoated substrates was highly significant (p-value < 0.05; [Fig. 8]
[Table 4]).
Table 4
Descriptive statistic and t-test for hardness test in Vickers hardness number of the
Ti substrate and HA coat
|
Group
|
Minimum
|
Maximum
|
Mean
|
SD
|
t-test
|
p-Value
|
|
Abbreviations: HA, hydroxyapatite; HS, highly significant; SD, standard deviation;
Ti, titanium.
|
|
Ti
|
255
|
280
|
266.7
|
2.765
|
32.22
|
<0.0001 (HS)
|
|
HA
|
980
|
1,210
|
1,118
|
26.26
|
Fig. 8 Plot of VH values for control substrate and the HA-coated substrate. HA, hydroxyapatite;
VH, Vickers hardness.
Discussion
A common limitation of metallic implants is its inability to pair harmoniously with
the surrounding bone. This mismatch between metallic implants and human bones is attributed
to differences in chemical compositions of the two materials. A surface modification
commonly applied to address this problem is coating the implant with a bioactive material,
such as HA, with a chemical composition similar to that of bone. Surface coating proves
to be successful in enhancing osseointegration and controlling the limitations of
metal-based implants, such as corrosion, wear, and release of harmful ions.[27] Surface coating is the solution in dental implantology. A bioactive coat masks the
adverse effects of metal surfaces (dental implants). A continuous and crack-free coat
indicates positive biological interactions between living tissues and artificial substitutes.
The power was fixed according to the result of pilot study. When we used power more
than 150 W, this result in the crack of the target and change in stoichiometry. When
we used power less than 100 W, coat was delaminated, adhesive, and cohesive failure
with porosity ([Fig. 2C]
[2D]). The temperature was 100°C to improve the adhesion.
FESEM revealed that the roughness of the HA coating surface was favorable. This roughness
is crucial in ensuring that the bone tissues and implants will bond, thereby enhancing
osseointegration and improving the prognosis of implant treatments.[19]
The chemical integrity of HA affects its biocompatibility. The molar ratio of Ca/P
is a key indicator of HA.[25] Based on crystallographic analysis, the HA coat resembles natural bones in terms
of chemical structure. Hence, the HA coat stimulates cell adhesion and improves cell
anchorage.
This result indicated that HA was the main constituent phase in the coating. Moreover,
the chemical composition of the coating material did not change during the sputtering
process. Despite long sputtering time, the device power of 150 W did not induce HA
to undergo thermal decomposition, a process that may affect its biocompatibility.
The whole section of the EDX map including the substrate and the coat revealed penetration
of Ca and P indicates HA inclusion within the superfacial thickness of Ti. If this
is correlated with dental implant applications, it enhances the bioactivity of Ti
fixtures that in turn may help in osseointegration.[28]
In this study, coat layer thickness was measured from the cross-sectional FESEM images
of Ti-coated specimens. A thick coating layer is beneficial that will allow sufficient
time for osseointegration, but previous studies reported that the coating layer cracks
when the thickness is 30 to 80 µm. The optimal coat thickness to enhance osseointegration
is 5 to 50 µm.[29]
[30] To achieve the desired osteoconductive effect of HA coating and prevent cracking
of the coating layer, researchers proposed an implant coating layer of only few micrometers
in thickness.[30]
[31] The excellent properties of the coat ensure proper fixture insertion during implantation
without the coat crumbling. Furthermore, a dense, thin HA layer is more dissolution
resistant than a porous, thick HA layer.[32] The 7-µm-coated thickness is ideal because such coating may not exfoliate as rapidly
as thinner or thicker HA coatings.
EDX mapping of the cross-section of the substrate, if this is correlated with dental
implant applications, it enhances the bioactivity of Ti fixtures that in turn may
help in osseointegration.[28]
Regarding implant surface roughness, increase surface roughness helps in mechanical
interlocking for hard and soft tissues.[33] Such rough coating surfaces are advantageous because they enhance bonding to bone
tissue in vivo.[34]
[35] Applying bioactive ceramics and increasing roughness are the main causes to improve
surface properties for Ti implants. Implant surface textures at values <100 nm promoted
cell adhesion and longevity,[36] while surfaces textures decreased cell attachment at values > 100 nm. Similar trends
have been found in other studies using different cells.[37] The removal torque values directly related to rough implant surface aid in mechanical
interlocking and improve the surface area.[33] The literature confirms that surface texture and chemistry are important. Other
technique which has the same coating criteria is long durations of pulse laser deposition
technique.
Surface roughness of the implant is another important surface, and high Vickers hardness
indicates that the coat layer is less likely to degrade over time. When bones remodel
at the implant site, the HA coating slowly degrades without loss of adhesion. The
slow release of HA ions at the interface probably stimulates the formation of new
bones.[38]
[39] An extended osseointegration will result from slow resorption of implant coating
and bone remodeling. Coated Ti dental implants will be covered with three-layered
cell sheets that successfully form using collagen grafts as a scaffold[40] to construct biohybrid implants.
Conclusion
Magnetron sputtering can be used to coat Ti dental implants with HA. A desirable coat
thickness can be safely achieved over a prolonged sputtering time of up to 22 hours.
Proper coat adhesion and interconnection and suitable coat thickness and hardness
can be achieved via magnetron sputtering. This technique is applicable for every implant
surface design, root or screw.
