PLASMA ELECTROLYTIC OXIDATION COATINGS
FOR IMPLANTS SURGERY
S.V. Gnedenkov, S.L. Sinebryukhov, O.A. Khrisanfova, A.V. Puz’, M.V. Nistratova
Institute of Chemistry, Far Eastern Branch, Russian Academy of Sciences,
Vladivostok, Russia
Abstract. The possibilities of plasma electrolytic oxidation (PEO) for development of
the composite coatings containing hydroxyapatite and calcium phosphates on the
titanium and nitinol (NiTi) surface were demonstrated. Such layers include pores in
the surface part of the coating which can be used as carriers for medicine (antibiotics,
hydroxyapatite or other phosphate-containing substances providing the best
compatibility of implant with bone tissues). Use of superdispersed
polytetrafluoroethylene (SPTFE) in the coatings composition enables one to increase
stability of substrate materials in the corrosion active media. PEO with subsequent
SPTFE treatment makes it possible to obtain a bioactive or bioinert implant surface.
In this case the polymer partly seals the pores where medicine previously was
inserted. It decreases the medicine diffusion from the surface layer and, therefore,
increases the duration of the therapeutic effect. Anticorrosion protective coatings
decreasing the nickel ions diffusion from nitinol substrate prevent nickel
accumulation in human tissues and its harmful after-effect. Moreover, the way of
hydroxyapatite formation in the coatings composition directly to over PEO was found.
In this case the ratio of Ca/P equals to 1.4, i.e. it is close for human bone tissue ratio
Ca/P (1.67).
INTRODUCTION
Titanium alloys have found the wide application in restorative surgery as basic
biomaterials for manufacturing of implant prostheses. Amongst the various materials
currently employed, the commercial pure titanium and alloy Ti–6Al–4V have found
extensive biomedical applications due to their good mechanical properties. These
properties combine good mechanical characteristics, high corrosion resistance and
good compatibility with biological materials. The passivity is due to the very stable
and tenaciously adherent oxide films spontaneously formed over the surface [1–3].
Besides, a bioinert material NiTi or nitinol (40–50 % at. Ti and 50–60 % at. Ni)
having a unique memory shape effect has been recently introduced in implantation
surgery [4, 5]. In particular, it is used to make holders for treating spinal traumas and
dystrophic illnesses, tack-implants used for junction of breast bone during cardiologic
surgeries etc. However, the diffusion and accumulation of nickel ions in the organism
soft tissues might have adverse effects, for example, of carcinogenic nature.
Bioactivity of titanium surfaces is not high enough to induce direct growth of
the bone tissue, and good bone fixation takes several months. Modifications of metal
surfaces are often employed as a means of controlling tissue–titanium interactions and
shortening the time for bone fixation [6]. Hydroxyapatite (HA) is a major component
of bone [7]. The hydroxyapatite-coated metallic implants show high tensile strength
and ductility of the metal, and bioactivity of hydroxyapatite. Enhanced
biocompatibility and bioactivity of titanium-base materials may be achieved by
coating them with ceramic and composite materials with using the plasma electrolytic
oxidation (PEO) method. It was shown as a result of experiments, that this method
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enables one to obtain protective layers on the surface of a material with far better
efficiency than by any other processing methods.
The PEO method is found to induce the emerging plasma micro discharges on
the electrode surface during the anodic or AC current polarization of the processed
material under the high voltage. As a result of local high energy effect, the layers
including as elements of the substrate (oxidized material) as well those of electrolyte
are formed on the surface of materials [8, 9]. The properties of such layers differ from
those of conventional anodic films. Subsequent treatment of the previously created
PEO structure (filling pores by bioactive and/or bioinert composites) allows building
composite coatings that could be prospective in terms of practical application in the
implant surgery.
At present, there is insufficient information in literature concerning the
development of protective coatings on nitinol (TiNi) regarding the use of the PEO
method. That is why it appears to be advisable to search the electrolyte compositions
and oxidation conditions to build composite structures with protective properties on
the nitinol surface to study their phase composition, surface morphology,
anticorrosion and mechanical properties.
HYDROXYAPATITE COATINGS ON TITANIUM
Commercially pure titanium VT1-0 (Ti – 99.4 %) plates (70 mm × 15 mm × 1
mm) were used as the substrates for PEO. As a pretreatment procedure, all samples
were ground using #400–#1000 SiC sandpaper gradually and then washed with
acetone and distilled water. The electrolyte was prepared by dissolving 30 g/l
disodium hydrogen phosphate dodecahydrous (Na2HPO4∙12H2O) and 30 g/l calcium
citrate (Ca3(C6H5O7)2∙4H2O) in bidistilled water.
PEO treatment were implemented with using the automatic control system
(ACS) consisting of the power source, control and measurement unit, computer and
software. The standard three-phase thyristor rectifier of type TEP-100.460H-22UHL4 was used as a power source. The ACS ensured the real time conditions of the
technological process parameters and detected the appearance of failures in the
system functioning. PEO process was carried out in two steps for 10 minute. First one
is the anodic polarization in the unipolar potentiostatic mode at 310 V for 400
seconds. The second step is the combined mode, which is the combination of the
anodic potentiodynamic (dU/dt = 1.25 V/s) and cathodic galvanostatic (jc = 1 A/cm2)
polarization modes for 200 seconds. The duration ratio of the anodic and cathodic
periods of polarization was τa/τc = 4.
Figure 1 shows the micrographs of the surface (Fig. 1 a) and the SEM crosssectional view (Fig. 1 b) of titanium substrate treated with PEO in the Ca- and Pcontaining solution. It could be seen that a porous network structure was formed on
the surface of titanium. The pores were well separated and homogeneously distribute
over the surface and bulk of the coating, and the pores in sizes varied from 1 to 20
m. The thickness of the coating was approximately 120 m. It was established that
the thickness of the coatings increases not proportionally during the PEO process
(Fig. 2). The coating thickness grows linearly up to 110 m during a first stage of the
process for 300 seconds and then this increases by 10 m only. According to the
calculation with the software of ImageJ 1.38x the porosity of the coating was about 10
%.
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a)
b)
Fig. 1. SEM surface morphologies (a) and cross-section view (b) of Ti sample with
coating.
Fig. 2. Dependence of the coatings thickness on the treatment time of the titanium by
PEO.
According to the X-ray data (Fig. 3) in this experiments the Ca- and Pcontaining coatings including HA were obtained. The coating formed in the Ca- and
P-containing solution contained Ca and P with Ti and O, as shown in Fig. 4. It was
implied that the elemental component in an electrolytic solution was introduced into
the coating by PEO. Relative concentrations of elements on the surface and in the
interface of titanium specimens treated with PEO and the ratios of concentrations of
calcium to those of phosphorus are not equal. The ratio of Ca to P on the surface was
1.4 while that in the interface was 0.1. The content of Ca was gradually raised and
concentration of Ti was gradually cut down while concentrations of O and P were
relatively unchanged from the titanium substrate to the surface. Thus we have in
surface region the layer which contains the Ca- and P-compounds without the
titanium oxides. The thickness of this layer on different samples varied from 10 m to
30 m. Formation of an apatite layer on the surface of metal implant provides the
living body a favorable condition for this material to bond to the living bone.
According to the detailed analyses of the surface apatite layer (Fig. 1), it was revealed
that the surface layer consisting of nano-size apatites similar to the bone mineral in its
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structure and composition. As a result, the surrounding bone can come into the direct
contact with the surface HA-layer on the implant. When this occurs, a chemical
adhesion is between the surface of apatite and bone mineral in order to reduce the
interface energy between them. It can be inferred from these that a new biomaterial is
able to form bone-like apatite on its surface in the living bone through the porous
apatite layer. The calcium phosphate is a precursor forming apatite and HA. The
porous surface of implants is beneficial to bone tissue growth and enhances the
anchorage of implants to the bone. At the same time, a defined porous structure may
be valuable as a depot for bioactive constituents such as growth factors or bone
morphogenetic proteins. Therefore, the porous and HA-containing coating on the
titanium formed by PEO method and presented in this study is expected to be
significant for medical applications.
Fig. 3. Diffractogram of titanium sample surface processed by the plasma electrolytic
oxidation method (PEO).
Fig. 4. Depth profiles of the elements in the PEO coating on the titanium.
COMPOSITE PROTECTIVE COATINGS ON THE NICKEL-TITANIUM
ALLOY
Nickel-titanium alloy (NiTi) is the one of the most popular materials for
various biomedical applications. The almost equiatomic nickel-titanium alloy is
unique in that it possesses interesting properties such as shape memory effect and
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superelasticity. The nickel-titanium alloy is successfully applied in manufacturing of
special devices for the medicine due to its mechanical properties. However,
implantation of nickel containing materials in human’s body requires some caution.
The metallic implants inevitably undergo in some degree of corrosion in body fluids.
This processes lead to the releasing of the nickel from implants into the human’s
body. It is well known, that the nickel is capable to cause a toxic and allergic
responses when its concentration exceeds a certain limit.
A selection of electrolytes providing the possibility of obtaining the titanium
oxides, aluminum oxides and phosphates or spinels in the composition of
anticorrosion layers was used in our experiments. The above electrolytes included
those containing aluminates, phosphates, carbonates and vanadates. In accordance
with the X-ray analysis data, the aluminum phosphate AlPO4 and nickel-aluminium
double oxide NiAl2O4 were presented in the coating obtained by PEO method (Fig.
5). During the analysis of diffractograms of surface layers of some samples, the
presence of an oxygen-containing compound of nickel and titanium Ni3Ti3O was
detected (its concentration was negligibly small on the diffractogramm shown in Fig.
5). At the same time, the titanium oxides were not found in the surface layers
composition.
Fig. 5. Diffractogram of nitinol sample surface processed by the plasma electrolytic
oxidation method (PEO).
The ESM pictures of the coating surface (a) and the optical picture (1000) of
the sample cross-section with the surface PEO-coating are shown in Fig. 6.
a)
b)
Fig. 6. ESM (a) and cross-section photo (b) of PEO coating formed on the nitinol
surface.
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The obtained PEO-coatings were studied by electrochemical impedance
spectroscopy. This method enables one to investigate the processes occurring at the
electrode/electrolyte interface with taking into account specific features of the surface
structure. The impedance spectra presented in a Bode plot (the dependence of
impedance magnitude |Z| and phase angle theta versus the frequency) for nitinol
samples after different kinds of treatment are shown in Fig. 7.
The impedance spectra of the samples with and without SPTFE coating are
virtually identical (see Fig. 7, curves 1 and 2). The processing of nitinol surface by
SPTFE powder has small effect on the state of the electrode/electrolyte interface and
increases insignificantly the impedance. The low effect of such processing must be
related to weak adhesion of the polymer to the metal substrate and insufficient
homogeneity of the formed protective layer. The sample processing by the PEO
method results in increase of the nitinol stability in the corrosion media. The addition
of dimethylglioxyme into the electrolyte somewhat increases the coatings protective
properties (see Fig. 7, curve 5) although this difference is not very significant, as it
seen from the impedance spectra.
Fig. 7. Bode plots for the investigated nitinol samples: 1 – without coating; 2 – coated
by SPTFE (heating at 100°С, 1 h) without the preliminary treatment; 3 – with PEOcoatng formed in unipolar mode (electrolyte: Na3PO4·12H2O – 10 g/l, NaAlO2 – 20
g/l, Na2CO3 – 10 g/l); 4 – with PEO-coatng formed in bipolar mode (electrolyte:
Na3PO4·12H2O – 10 g/l, NaAlO2 – 20 g/l, Na2CO3 – 10 g/l); 5 – with PEO-coating
formed in unipolar mode (electrolyte: Na3PO4·12H2O – 10 g/l, NaAlO2 – 20 g/l,
Na2CO3 – 10 g/l, dimethylglioxyme – 1 g/l); 6 – as the sample № 3 and treatment by
54
SPTFE; 7 – as the sample № 4 and treatment by SPTFE. The equivalent circuit
simulating the experimental data is presented in the insert.
The only appreciable increasing of the impedance was detected on the
polarization curves (see Fig 7, curve 5). One can suggest that a chelate compound –
nickel dimethylglioxymate – deposited in pores of the oxide layer stimulates the
increasing of the oxide layer protective properties. However, its concentration in the
film is not high (less than 10 %), because the lines attributed to this phase were not
detected in the diffractogram. Being a thermally unstable compound, nickel
dimethylglioxymate is partly decomposed under the effect of PEO that is known to
involve attainment of high temperatures in short-lived plasma channels and due to
thermolysis in the coating area adjacent to the plasma channel. That is why low
content of the nickel dimethylglioxymate in the coating pores does not allow attaining
the efficient corrosion protection under significant field shifts.
Two time constants can be seen on the diagram of the phase angle dependence
versus the frequency for the samples with PEO coating (Fig. 7) as compared to the
samples without coating and containing SPTFE only. The above data indicate to the
two-layer coating structure: the upper layer is porous while the lower one is pore-free.
It is in a good agreement with the earlier suggested model of the PEO-coating [10].
As a rule, the porous layer has crater-like cavities with the diameters up to few
micrometers. So the pore sizes are larger than the size of SPTFE powder (nearly 1
µm) applied for surface treatment. The development of the porous structure can be an
additional advantage of the PEO method, since the developed surface promotes the
best overgrowing of implant by bone tissue and allows filling pores with bioinert or
bioactive composites.
After the processing of the PEO-coated sample with SPTFE powder the
modulus of impedance increases significantly. Its value is of an order of magnitude
higher than for the sample without coating. This fact confirms that processing with
SPTFE powder allows filling the coating pores with the polymer on the surface and,
therefore, forming an additional barrier layer hindering the metal ions release into the
solution.
The deviation of the phase angle theta from 90° characterizes the degree of
“imperfection” (in other words, heterogeneity) of the object under investigation.
Therefore, the constant phase element CPE was used instead of the capacitance
during fitting of the experimental impedance spectra. The impedance of a constant
1
phase element is defined as ZCPE 
, where 1  n  1 ,   2 f is the
Q( j )n
angular frequency, and Q is a frequency-independent parameter. The circuit presented
on the insert in Fig. 4 provides an adequate simulation of the experimental data
presented and takes into account the particularities of the composite coating structure.
The time constant described by the elements CPE1 and R1 is responsible for
the porous part of the composition layer, its morphological structure, and roughness.
This constant is clearly revealed for PEO layers (curves 3 and 4) on the dependence of
the phase angle versus the frequency in the frequency range 105–106 Hz. Moreover,
for the PEO layers processed by SPTFE this constant was somewhat transformed due
to equalizing of their surface (decrease of their roughness) as a result of the
proportional distribution of SPTFE powder under thermal treatment conditions and
transformation of resistive and capacitive components of the time constant.
As CPE1 and R1 characterize the geometrical capacitance and resistance of the
surface porous part, respectively, the pores filling by polymer and the increasing of
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the composition layers thickness are responsible for changes of the phase angle in the
high frequencies range.
For the samples 1 and 2 this time constant is absent. The elements CPE2 and
R2 characterize nonporous layers of the PEO coating (Fig. 7, curves 3 and 4) and such
surface layers as natural oxide and natural oxide coated by the SPTFE surface layer
(curves 1 and 2, respectively). The frequency range in which the above particularities
of the studied objects are observable is rather wide: from 10-2 up to 103 Hz. The
elements CPE3 and R3, in accordance with similar model suggested for the description
of impedance spectra of the anodic coatings on aluminum [11], provide the
information on contribution of the pores that are compactly sealed up by polymer
(with air between the polymer “plug” and the pore bottom). Such a situation is more
likely for the pores of small sections always presenting in the PEO layer. The above
time constant characterizes only the PEO coatings processed by SPTFE and is
detected on the curve of the phase angle (Fig. 4, curves 6 and 7) in the range of
intermediate frequencies (10–103 Hz).
As was shown by the experimental results of previous studies [8], after pores
filling with SPTFE powder the thermal treatment is necessary to use the polymer to
build smooth hydrophobic surface layers. In this case the polymer seals up the pores,
where medicine can be introduced beforehand, and, therefore, prevents its diffusion
from the surface layer. So the effect of the therapeutic medicine presence in a pore
could be extended. Besides, the formed anticorrosion protective coatings decrease
significantly the nickel ions release from nitinol and prevent nickel accumulation in
human tissues.
THE INFLUENCE OF PEO ON THE MECHANICAL CHARACTERISTICS
OF THE NITINOL
The plasma termochemical reactions realize on the sample surface at high
discharge temperatures and high pressures. In the spark discharge channel the
temperature reaches about 103–104 K and the pressure equals to about 102 MPa. As a
result of these processes an inorganic glass-ceramic-like coating structures obtain on
the metal surface. Local effects of the high temperatures may influence on mechanical
properties of the substrate. It is very important for the implant materials at least to
save the mechanical properties after any preliminary processing. As a result of
thermal oxidation on the air medium, for instance, the microhardness of surface layers
on titanium due to dissolving of oxygen in the metal sharply increases and therefore
increase of fragility and decrease of the durability of the metal construction in whole.
The aim of present investigation was to study the influence of the plasma
electrolytic oxidation on the mechanical properties of the films obtained on surface
NiTi as well the lying near to the coating layers of the substrate.
NiTi wire samples diameter of 1.3 mm were used. Two types of nickeltitanium alloys were studied in present investigation. There were in the austenitic state
Ni50.7Ti49.3 (at. %) and martensitic state (Ni50Ti50) (at. %) at the room temperature.
Preliminary the samples were polished with graded SiC paper down to 1000-grit and
finally washed with distilled water. Cross-sections of the coatings were prepared by
the metallurgical method of the samples processing. Load-unload tests on the
investigated materials were carried out on a Dynamic Ultra-micro Hardness Tester
DUH-W201 (Shimadzu, Japan). As the indenter the Berkovich triangular pyramid
with 110 tip angle was used.
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From the experimental data an average values of the microhardness and elastic
modulus were calculated for tested materials (for nickel-titanium in austenitic and
martensitic modifications: Ha = 2.6 ± 0.1 GPа, Hm = 2.0 ± 0.1 GPа, Ea = 64 ± 2 GPа,
Em = 57 ± 2 GPа; for coating: Hcoat= 1.6 ± 0.2 ГПа, Ecoat = 30 ± 2 ГПа). It should be
noted that coating has less microhardness and elastic modulus as compared the
substrate. It is certain advantage of the obtained coating material because these
meanings are situated some near to values of the natural bone's tissue. Figure 8 shows
the distribution of the microhardness (a) and elastic modulus (b) depending on the
distance of the indentation point from the resin/coating interface. Linear
approximation of the experimental data separately within of the coating area and
within the substrate material both austenitic and martensitic modification was
performed (Fig. 8). It is seen naturally that the hardness and elastic modulus of the
austenitic substrate is higher than martensitic one. But the hardness values of both
coatings obtained on the austenite as well as on the martensite are equal practically.
Average microhardness of the coating is less, than the values of this parameter for the
substrate. Nevertheless, on the both modifications of nitinol the microhardness of the
boundary surface layers of substrate materials contacting with coating has values of
microhardness some smaller than volume layers.
Fig. 8. Plot of the dependence of microhardness (a) and Young's modulus (b) versus
the distance of the indentation point from the coating/resin interface for the austenitic
and martensitic substrate with PEO coating (linear trend approximation was used).
As a result of the detailed analysis of the experimental data, the conclusion can
be done: the surface hardness and elastic modulus have an equal reducing in the
region of the coating/substrate interface. It confirms that the surface layers of alloy
nearby to a coating have a microhardness and elastic modulus smaller than volume
layers of the alloy. It should be noted that there are some inhomogeneous zones
(outliers) in both a coating and alloy immediately. It is explained by cluster’s
construction of the coating and its heterogeneous composition. These outliers may be
caused by the inhomogeneous inclusions characterized by different values of the
microhardness and elastic modulus in the alloy directly (Fig. 8).
CONCLUSIONS
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Titanium alloys surfaces can be modified by electrochemical treatment for
better biomimetic coatings. In this study, the hydroxyapatite-containing was obtained
by PEO method in Ca- and P-containing electrolyte solution. The surface layer of the
PEO coating has a porous structure and consists of the calcium phosphate including
hydroxiapatite. This work introduces a simple method to make titanium implant
surface bioactive and porous.
Nitinol may be treated by the PEO method with purpose to form the protective
coating which decreases a release and accumulation in human tissues of the nickel
ions. Such coatings have good mechanical properties also. It was shown the absence
of the negative influence of the plasma electrolytic oxidation on the mechanical
properties of the nickel-titanium alloy.
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