grit blasting of medical stainless steel - implications on its corrosion behavior, ion release and...

8/13/2019 Grit Blasting of Medical Stainless Steel - Implications on Its Corrosion Behavior, Ion Release and Biocompatibility

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Grit blasting of medical stainless steel: implications on its corrosion behavior, ion

release and biocompatibility

J.C. Galvna, L. Saldaa

b,c, M. Multigner

a, A. Calzado-Martn

c,b, M. Larrea

a,

C. Serrad, N. Vilaboa

c,b, J.L. Gonzlez-Carrasco

a,b,*

(a) Centro Nacional de Investigaciones Metalrgicas (CENIM-CSIC), Avda Gregorio del

Amo n 8, 28040 Madrid, Espaa

(b) Centro de Investigacin Biomdica en Red en Bioingeniera, Biomateriales y

Nanomedicina (CIBER-BBN), Madrid, Espaa(c) Hospital Universitario La Paz-IdiPAZ, Paseo de la Castellana 261, 28046 Madrid,

Espaa

(d) CACTI Universidade de Vigo, Campus Lagoas-Marcosende 15, 36310 Vigo, Espaa

Keywords: Austenitic stainless steel; Grit blasting; Corrosion behavior; Ion release;

Biocompatibility

*Corresponding author. Tel.:+34 91 5538900 Ext 215; fax: +34 91 5347425.E-mail address:

[email protected] (J.L. Gonzalez-Carrasco)

Abstract

This study reports on the biocompatibility of 316 LVM steel blasted with small and rounded

ZrO2 particles or larger and angular shaped Al2O3 particles. The effect of blasting on the in

vitro corrosion behavior and the associated ion release is also considered. Surface of Al2O3

blasted samples was rougher than that of ZrO2 blasted samples, which was also manifested by

a higher surface area. Compared to the polished alloy, blasted steels exhibited a lower

corrosion resistance at the earlier stages of immersion, particularly when using Al2O3

particles. With increasing immersion time, blasted samples experienced an improvement of

Journal of Materials Science Materials in Medicine 23(3) (2012) 657-66.
mailto:[email protected]:[email protected]


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the corrosion resistance, achieving impedance values typical of passive alloys. Blasting of the

alloy led to an increase in Fe release and the leaching of Ni, Mn, Cr and Mo. On all surfaces,

ion release is higher during the first 24 h exposure and tends to decrease during the

subsequent exposure time. Despite the lower corrosion resistance and higher amount of ions

released, blasted alloys exhibit a good biocompatibility, as demonstrated by culturing

osteoblastic cells that attached and grew on the surfaces.

1 Introduction

Austenitic stainless steel 316 LVM (Low Vacuum Melting) is one of the most frequently used

biomaterials for internal fixation devices because of a good combination of mechanical

properties, biocompatibility and cost effectiveness [1]. One common failure encountered in

stainless steel devices arises from corrosion attack, which may decrease the structural

integrity of the implants and elicit adverse local and remote tissue responses mediated by

corrosion products [2-4]. In fact, elevated serum Cr levels were found in patients treated with

stainless steel modular femoral nails [5]. Retrieved modular nails presented signs of stainless-

steel corrosion products adherent to the junction where osteolysis, periosteal reaction, or

cortical thickening were detected. Thus, the failure of stainless steel implant devices has been

associated to inflammatory reactions in peri-implant soft tissues [5-7]. Mechanical factors

such as applied stress, wear, and micromotion may accelerate electrochemical dissolution,

leading to premature structural failure of the implant and accelerated metal ion release [8,9].

Therefore, mechanical stability of implanted stainless steel in contact with bone tissue and

body fluids is of fundamental importance to ensure the implant success.

Severe surface plastic deformation of 316 LVM by grit blasting is considered an

attractive modification to improve fatigue strength of intramedullary nails for the proximal

femur and diaphysary fractures, providing an optimal combination between high resistance



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during the consolidation period and a minimal invasive geometry. Besides roughening, this

surface treatment modifies the mechanical properties of the surface and near surface region

through the induced compressive residual stresses. Whereas a number of studies have

addressed the blasting induced effects on titanium and titanium alloys [10-14], there are only

a few reports concerning the effect of blasting on corrosion resistance of stainless steels [15]

and cytocompatibility of surface-treated stainless steel with bone-forming cells [16].

This study deals with austenitic stainless steel 316 LVM modified by blasting of the

surface with small and rounded ZrO2particles or larger and angular shaped Al2O3particles.

The influence of surface blasting of austenitic stainless steel 316 LVM on osteoblastic cell

adhesion and proliferation, functions that play a key role during cell colonization of the

implant, was evaluated. Other surface dominated events such as corrosion behavior and ion

release were also addressed. The effects of blasting on the subsurface residual stresses and

mechanical properties have been reported elsewhere [17-19].

2 Experimental procedures

2.1 Materials

Austenitic stainless steel 316 LVM, which chemical composition (wt%) is Cr 17.48, Ni 14.13,

Mo 2.87, Mn 1.62, Si 0.53, C 0.024, Cu 0.067, N 0.061, S 0.001, and Fe in balance, was

supplied by the implant manufacturer (Surgival SL, Valencia, Spain). Discs of 20 mm

diameter and 2 mm thick, hereafter PL samples, were grinded and polished by conventional

metallographic techniques. A final finishing was applied with silica gel. A set of samples was

blasted with small and rounded ZrO2particles or larger and angular shaped Al2O3particles,

hereafter BL-ZrO and BL-AlO samples, under the same experimental conditions.18

Blasted

and polished samples were finally passivated in citric acid (20%) at 40 C during 30 minutes.



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For cell culture and ion release studies, all the samples were routinely sterilized under UV

light in a laminar flow hood for 12 h on each side and stored until use.

2.2 Surface characterization

Topographic surface analysis was performed with an interferometer optical profilometer

NT1100 (Wyco-Veeco, Santa Brbara, CA, USA) using a Vertical Scanning Interferometry

mode. This experimental setting provides with a vertical resolution of < 1 nm and a lateral

resolution of 400 nm.Ra(m),Rq(m),Rz(m),Rt(m), and real surface area (mm2) were

determined at 5X, 20X, and 50X magnifications, which yields fields of view of 1.092 mm2,

0.068 mm2, and 0.011 mm

2, respectively. Surface Skewness, Ssk, which can be interpreted as

the degree of asymmetry of a surface height distribution, was determined from 5X images.

Average values correspond to 10 fields of view.

The geometric surface area (A) was 6.41 cm2for all the discs. The area increase after

blasting corresponds to the area ratio (%) of measured surface / scanned surface and allows

determining an index area.

Microstructural characterization of surfaces and cross-sectional views were performed

by using a scanning electron microscope (SEM) Jeol JSM-6500F (Japan) equipped with a

field emission gun (FEG) emitter coupled with an energy dispersive X-ray (EDX) system for

chemical analysis. Depending of the analysis, secondary (SEI) or backscattered (BEI) electron

images were selected.

2.3 Cell culture assays

In vitrobiocompatibility of the samples was evaluated by using human osteoblastic Saos-2

cells (ECACC, Salisbury, Wiltshire, UK). Cells were grown in Dulbeccos modified Eagles

medium (DMEM) (Lonza, Barcelona, Spain) supplemented with 10% (v/v) heat-inactivated



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fetal bovine serum (FBS), 500 UI/ml of penicillin and 0.1 mg/ml of streptomycin and

maintained at 37C in 5% CO2in a humidified incubator.

For adhesion assays, cells were seeded on the investigated surfaces in 12-well plates (5

104 cells/well) and incubated for 2, 4 and 6 h. Cell adhesion was assessed using the

alamarBlue assay (Biosource, Nivelles, Belgium) which incorporates a redox indicator that

fluoresces in response to cellular metabolic reduction. After washing extensively with PBS,

attached cells were incubated in DMEM containing 10% alamarBlue dye for 4 h. After

excitation at 530 nm, the fluorescence emitted at 590 nm was quantified using a microplate

reader Synergy 4 (BioTek Instruments, Winooski, VT, USA). For viability assays, cells were

seeded on the surfaces in 12-well plates (1.5 104cells/well) and cultured for 1, 4 and 7 days.

Cell viability was determined using the alamarBlue assay as described above. Cells cultured

on tissue culture-treated polystyrene 12-well plates (PS) (Nunc, Roskilde, Denmark) were

used as controls for cell adhesion and viability assays. All experiments were carried out at

least twice, each in duplicate, with similar results.

2.4 Corrosion experiments

Electrochemical impedance spectroscopy (EIS) tests were performed in a conventional

electrochemical cell filled with the Ringers solution (8.36 g of NaCl, 0.3 g of KCl, and 0.15 g of

CaCl2in each 1000 ml of distilled water) and using the sample as working electrode. A counter

electrode of platinum and a reference electrode of Ag/AgCl saturated in a potassium chloride

solution were used. The EIS measurements were performed using a potentiostat/galvanostat

AutoLab EcoChemie PGSTAT30 (Eco Chemie, Utrecht, The Netherlands) equipped with a

FRA2 frequency response analyzer module. Frequency scans were carried out close to the

corrosion potential. Sinusoidal wave perturbations of 10mV in amplitude were applied in the

frequency range of 100 kHz to several mHz. Five impedance sampling points were registered per



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frequency decade. The EIS measurements were made after 5 min and 24 h of immersion. The

impedance data were analyzed by using the EQUIVCRT program[20].

2.5 Ion release

Samples were incubated in 30 ml of the Ringers solution in a humidified 5% CO2

atmosphere at 37C for up to 4 days. At every 24 h period, 23 ml of solution were removed

and replaced by fresh one, in order to simulate the fluid exchange of an adult human by the

excretion of urine, as described elsewhere [21,22]. Released ions in the solutions were

quantitatively analyzed using an Inductively Coupled Plasma Optical Emission Spectrometer

(ICP-OES PerkinElmer Model Optima 3300 DV, Palo Alto, CA, USA). Calibration

solutions of Cr, Fe, Mn, Mo and Ni were prepared by appropriate dilution of 1000 mg.L-1

multielementalCertiPur grade (Merck, Darmstadt, Germany) standards solutions. The

solutions were prepared using the Ringers solution as matrix. All solutions were 1% (v/v) in

nitric acid. All experiments were performed by triplicate. The amount of released ions was

calculated as mass per volume unit (g.ml-1) or per unit area (g.cm-2). Detection limits of the

ICP for the investigated metals were


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used for blasting, thus Zr, Al and Si-rich oxides are found in sample BL-ZrO and Al-rich

oxide in BL-AlO.

Topographic parameters shown in Table 1 reveal a higher roughness (Ra) for the

samples blasted with alumina, which evidences the higher erosion and plastic deformation of

this surface. The roughness increase is obviously accompanied with a significant increase in

the surface area (up to about 170 % for the BL-AlO sample). The observation of a number of

irregular intrusions and protrusions with sharp ridges in the alumina blasted surfaces is

consistent with their higherRt values, which denotes the average distance between the higher

tenth picks and the deeper tenth valleys. From the analysis of the Skewness parameter it

follows that the polished surface, with an average values well above 0, is a flat surface with

peaks. BL-AlO presents values well below 0, which indicates is a surface mainly composed

with holes. The nearly 0 value found for the BL-ZrO samples denotes a surface with holes

having the most symmetric height distribution respect PL and BL-AlO samples. Since Ssk

values are numerically below 1.0, the presence of extreme holes or peaks on the blasted

surfaces are not expected. All these parameters calculated from the three fields of view shows

a strong consistency.

Cross sectional examination confirmed the presence of remnant of the blasting

particles embedded at the surface and the development of a narrow zone (10-15 m) with an

ultrafine grain size (Fig. 2). Taking into consideration the gradients in hardness and

compressive residual stresses, blasting affected zones of about 150 and 200 m were

determined for the BL-ZrO and BL-AlO samples, respectively [18].

3.2 Biocompatibility

Next, we investigated whether blasting of 316 LVM affects osteoblast-like Saos-2 cells

adhesion and viability. The number of attached cells increased with time on the three tested



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surfaces (Fig. 3A). The process of grit blasting did not significantly affect the number of

attached cells at any tested time. Cell growth increased over time on both polished and rough

surfaces, as indicated by measurements of metabolic activity (Fig. 3B). Cell viability was

similar on the three studied surfaces at days 1 and 4. At day 7, the number of viable Saos-2

cells on the BL-AlO samples was lower than on PL and BL-ZrO samples (Fig. 3B).

3.3 In vitro corrosion behavior

Figure 4 shows the Nyquist plots of the impedance data for the polished and blasted samples.

After 5 min of immersion, all the impedance plots tended to describe a semicircle,

approaching to the behavior of a non-passive alloy. With progressing immersion time,

semicircles were also described but with larger diameter, which indicates that all samples

reached a spontaneous passivation when they are exposed to the Ringers solution. Figures 5A

and 5B show the same impedance data plotted in the Bode format into the domain of 105-10

-3

Hz and Figure 5C shows a detail of the impedance modulus plots at low frequencies (1-10-3

Hz) using a linear-linear scale.

The experimental impedance data can be modeled using a complex non-linear least-

square (CNLS) fit analysis and suitable electrical equivalent circuits (EECs). In a first

approach, the EEC proposed by Mansfeld and Wang [23], Figure 4D, was used. This circuit

consists of a resistance R1 and a parallel CPE-R2 couple. R1 corresponds to the electrolyte

resistance and R2 to the polarization resistance of the passive surface. In reference[23], C

represented the capacitance of the passive surface but in this discussion Cis implemented as a

Constant Phase Element (CPE). The CPE should be used instead of a pure capacitance to

account for a non-ideal capacitive response. This CPE arises because microscopic material

properties are themselves often distributed. For example, the solid electrode/electrolyte

interface on the microscopic level contains a large number of surface defects, local charge



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inhomogeneities, adsorbed species, variations in composition and stoichiometry, etc[24]

Following the equation used by Boukamp in the EQUIVCRT program [20], the impedance of

CPE is defined by:

Z(CPE) = (j)-n/Yo

where j = 1, is the angular frequency = 2f, and f is the frequency in Hz. The

exponential factor n is related to a non-uniform current distribution due to the surface

roughness or other distributed properties, and varies between 0 and 1.

Table 2 shows the parameters obtained by using CNLS fit analyses from the Boukamp

program [20]. The chi-square values and the error calculus describe the quality of the fitting.

An exception is observed for the BL-ZrO samples, where the relative error found for the R2

values after 1 day of testing was extremely high (506%). Moreover, the R2value obtained for

this sample is also anomalous. It is believed that impedance spectra can be totally masked by

experimental noise and uncertainties, thus the capacitance components of the response may

not give exactly a CPE behavior [24].

A classification based on the values of polarization resistance of the passive surface

(R2) leads at the 5 minutes of immersion to values that are similar to those obtained with

criteria based on the impedance modulus reached at the lowest frequencies. As can be seen,

polarization resistance for polished samples is higher than that for BL-AlO samples.

However, in the case of BL-ZrO samples this circuit leads to anomalous values when

considering 24 h of immersion. Complexity of the blasted surfaces makes difficult to get more

quantitative information for these samples.

3.4 Ion release

Figure 6 shows the accumulative amount of ions released from the surfaces along 24, 48, 72,

and 96 h of immersion in the Ringers solution. We only detected small amounts of Fe



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released from polished samples within the first 24 h that remained constant over time.

Blasting of the alloy led to elevated Fe ion release, but also the leaching of Ni, Mn, Cr, Mo,

and Al ions. In general, ion release increased with immersion time in both blasted surfaces.

The total amount of ions significantly increases from 0.08 g.ml-1for the polished condition

to about 1.3 g.ml-1for the blasted samples, being Fe preferentially released compared to Cr,

Mn and Ni. The relative accumulative amounts of ions for BL-AlO samples was Fe>> Ni

Mn> Cr> Al > Mo whereas for BL-ZrO samples was Fe >> Mn> Ni > Cr > Mo. (p


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Blasting of stainless steel has been developed to provide roughness in a micrometrical

range that bone cells can recognize. Biocompatibility tests were set to address the effect of

surface roughness on osteoblastic behavior. Data presented herein show that blasted surfaces

are biocompatible, irrespective of the particle used for blasting. However, some

inconsistencies related to the roughness dependence observed on other metallic biomaterials

are found. At first, it should be considered that changes in surface topography and chemistry

of metallic materials may modulate cell behavior, such as initial cell attachment, proliferation

or differentiation [16,26-29]. In particular, a number of previous studies indicate that blasting

of metallic materials produced a detrimental effect on initial osteoblast adhesion [27,29,30]

but in other publications such effects have not been observed [31,32]. Although the reason for

these discrepancies remains unclear, multiple evidence indicates that cell attachment is

strongly influenced by the process used to prepare the surface and hence by physicochemical

surface characteristics [31]. Since initial attachment of Saos-2 cells was unaffected by surface

roughness on stainless steel, we speculate that the effect exerted by blasting of 316 LVM

surfaces do not affect short-term adhesion. In this regard, it has been suggested that initial

non-specific electrostatic forces established between cells and substrates and passive

formation of ligand-receptor bonds are more influenced by surface chemistry than by surface

topography [31]. Moreover, it has been recently reported that nanograined/ultrafine-grained

structure of the stainless steel enhances early interactions of fibroblasts [33]. Since blasting

develops an ultrafine grain size at the outermost blasted affect zone of about 1015 m thick,

with grain sizes ranging between 50 and 500 nm, we hypothesize that the detrimental effect of

surface roughness on cell attachment can be circumvented by these ultrafine structures.

However, the short-term adhesion does not always reflect the further behavior of cells on the

substrates [28,34]. In fact, we observed that Saos-2 viability decreased when cultured for one

week on BL-AlO samples, with pronounced roughness (Rahigher than 5 m), as compared to



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polished or BL-ZrO surfaces. In principle, this effect cannot be related to increased ion

release that resulted in higher cell toxicity, as total ions released from both blasted samples

were found to be rather similar. Micromolar doses of Al ion, which was released from BL-

AlO surfaces, have been reported not to affect metabolic activity and proliferation of bone

forming cells [35]. More likely, the effects observed on samples blasted with Al 2O3 support

the idea that changes of surface topography of metallic alloys in the micrometric range, also

including 316 LVM, influence long-term cell adhesion and proliferation.

The corrosion resistance of biocompatible materials in body fluids is one of the

essential factors in the determination of the lifetime of medical implants. Blasting the surface

of stainless steel may affect its tendency to corrode when implanted. Figure 5A shows that the

highest impedance modulus corresponds to the polished samples and the lowest to the blasted

samples, which indicate a decrease in the corrosion resistance following blasting. Corrosion is a

surface dominated process, thus it could be argued that this different behavior is an artifact

related to the area increase of the surface. For sake of clarity it is worth mentioning that Figure

5A shows the typical log-log plot, thus from his analysis it is difficult to assess the specific

weight of the area increase for the blasted samples. Using a linear-log scale, Figure 5C, it can be

clearly seen that for a given frequency the difference in the impedance modulus between

polished and blasted surfaces is nearly the same despite the area increase is much higher for the

BL-AlO samples. Thus, the larger area increase following blasting cannot explain differences in

the corrosion behavior between blasted surfaces and additional effects such as surface

contamination, strain induced -martensite formation, residual stresses, and grain size

refinement are next considered.

On the one hand, remnant of blasting particles embedded at the surface would play a

detrimental role since they are non-conducting and could act as cathodic zones, enhancing

dissolution at the non-contaminated zones [36]. On the other hand, blasting with alumina



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particles (largest size, density and hardness) lead to more cumulative plastic deformation,

which implies more strain hardening and formation of -martensite at deeper regions from

the surface [18]. Since the plastically deformed subsurface region was constrained elastically

by the material beneath the blasted affected zone, a compressive residual stress zone was

developed. The angular shape of the alumina particles, however, causes severe erosion that

grinds down the material and yields a more heterogeneous deformation forming large pits, as

deduced from the Sskvalues (Table 1). Besides partial removal of -martensite, magnitude

of the maximum compressive residual stress for the BL-AlO samples (470 MPa) is lower than

for the BL-ZrO samples (670 MPa) [18]. Interestingly, this value decreases both to the

interior and to the blasted surface, likely changing into tensile residual stresses at a certain

depth, as required to achieve a zero macroscopic residual stress on the specimen. At the

sample surface, therefore, slight tensile stresses rather than compressive stresses are expected,

which will make the surface more reactive [37]. Besides, it is known that the corrosion

behavior of stainless steels can deteriorate substantially when the strain-induced -martensite

is present because the structural non-homogeneities would increase the density of the

localized states [38]. Magnetic measurements indicated that both type of samples have

approximately the same quantity of -martensite [18].

Interestingly, impedance values increases with increasing the immersion time despite the

high concentration of Cl- ions of the medium, which denotes an improvement of the corrosion

protection likely due an increase in the thickness of the passive film as consequence of the

equilibrium with the surrounded medium. The outermost ultrafine-grained layer would yield

good corrosion resistance because the high amounts of grains boundaries would enable fast

diffusion of Cr to the oxide passive film covering the surface [39]. However, although the

impedance increase at the lowest frequencies for the BL-AlO samples (up to about 46%)

approaches to the values found for the polished surfaces (about 56 %), differences between both



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type of surface becomes slightly higher at 24 h. As the magnitude of the impedance achieved for

blasted surfaces are typical of passive alloys, it could be concluded that blasting does not

seriously challenge the passive behavior expected for an austenitic stainless steel.

While ion release from austenitic stainless steel 316 L have been investigated in

various physiologic media and different surface conditions [40-43], studies on the blasted

condition are almost lacking. Reliable data are relevant to assess whether there is any

potential for risk of adverse effects arising from the element contained in the blasted alloy.

Results for the polished steel agree with previous studies that found preferential Fe release for

all exposure periods [40,41,43]. The Fe ion release is higher during the first 24 h exposure

and tends to decrease during the subsequent exposure time.

Relevant for this investigation is that blasting of the alloy yields an increase in the Fe

content but also the leaching of new ions, irrespective the immersion time. This fact correlates

with the decrease in the corrosion resistance of the blasted samples. The thinner oxide films

on the surfaces rather than their area increases may account for higher ion release from

blasted surfaces. In fact, a direct correlation cannot be made between the surface area of

blasted surfaces and the ion release increase. While real surface area increase from 8.2 cm2for

BL-ZrO to 13.3 cm2 for BL-AlO, total ion release per real area reveals values of 5.07 and

3.02 g.cm-2 for the BL-ZrO and BL-AlO samples, respectively. The lower ion release from

BL-AlO samples, exhibiting the worse corrosion behavior, is somewhat confusing and a more

detailed analysis of the oxide film formed on each metal surface seems to be necessary. Ion

release is a process related to the dissolution of elements forming the passive film and

therefore its thickness, chemical composition and element distribution could be different.

Determination of these features on a rough surface, however, would be rather complex.

Relevant for the intended application is the increase in the Ni release since the earliest

stage of immersion, which agrees with results of a recent work of Reclaru et al. [44] that has



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shown that cold working (23%) of austenitic stainless steel significantly increases the release

of Ni. Although every metal has its own intrinsic effect and potential toxicity, Ni ions are of

particular interest since they may cause biological effects and also are the origin of the most

widespread contact dermatitis [45]. Consequently, the European Union set from 2001 a

regulation limiting the Ni release of utensil and jewellery items at 0.5 g.cm-2during a week

for at least two years [46]. This would equal 0.071 g.cm-2 on a daily basis, which is

overcome by the BL-ZrO (0.112 g.cm-2) and the BL-AlO (0.095 g.cm-2) samples (values

determined using the real areas). The results of this study should not be viewed as conclusive

evidence since immersion times were rather short, thus more realistic average values would

be obtained after longer times of immersion. Moreover, metal release is strongly influenced

by the biological environment [47] and thus further experiments are needed to elucidate the

influence of biomolecules contained in physiological fluids on the corrosion behavior of

blasted steels.

5 Conclusions

Topographical analysis reveals that surface of the alumina blasted samples is rougher

(Ra~ 5 m) than the zirconia blasted samples (Ra~ 1 m), which is also manifested by a

higher surface area increase (~170%).

Corrosion tests reveal a lower corrosion resistance of the blasted surfaces at the earlier

stages of immersion, particularly when using Al2O3particles. With increasing immersion

time, blasted alloys experience an improvement of the corrosion resistance achieving after

24 h exposure impedance values typical of passive alloys. The benefits of the ultrafine

grained structure beneath the blasted surfaces must balance the detrimental role played by

the embedded blasting particles and the strain induced -martensite reaching the surface.



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Blasting of the alloy yields an increase in the Fe release, but also to the leaching of new

ions (Ni, Mn, Cr, Mo, and Al).

Despite the lower corrosion resistance and higher amount of ions released, blasted alloys

exhibit a good biocompatibility. As total ions released from both blasted samples were

found to be similar, the slight decrease in the cell viability observed on the alumina

blasted samples, having the highest roughness, support the idea that changes of surface

topography in the micrometric range influence long-term cell adhesion and proliferation.

Acknowledgements The authors wish to express their thanks for the financial support of

Spanishs Projects from Ministerio de Ciencia e Innovacin (MAT2009-14695-C04-02 and -

04) and Fundacin Mutua Madrilea. NV is supported by program I3SNS from Fondo de

Investigaciones Sanitarias (Spain).

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22

Table captions

TABLE 1. Topographic surface parameters of the investigated surfaces as a function of the scanned

area.

TABLE 2. EEC parameter values, relative errors and chi-square values obtained by modeling

the experimental impedance data showed in Figures 4 and 5 for specimens immersed in the

Ringers solution during 5 minutes (a) and 1 day (b).

Figure captions

FIGURE 1. BEI images corresponding to selected areas of 316 LVM specimens blasted with

A) alumina and B) zirconia particles.

FIGURE 2. BEI images corresponding to cross sectional views of BL-AlO specimens. The

inset is a close up of the ultrafine grain size zone.

FIGURE 3. Cell attachment (A) and viability (B) on stainless steel surfaces. Saos-2 cells were

cultured on polished ( ), BL-ZrO ( ) and BL-AlO ( ) samples for the indicated incubation

periods. The results are expressed as the percentage of the fluorescence measured on PS at 2 h

or 1 day, which was given the arbitrary value of 100. Each data represent the mean S.D. of

four independent experiments. * p < 0.05 compared to PL specimens.

FIGURE 4. Nyquist plots of; (A) polished specimens, (B) specimens blasted with ZrO2, and (C) the

specimens blasted with Al2O3, after 5 minutes and 1 day of immersion in Ringers solution. D)

Electrical equivalent circuit used to analyze the experimental impedance data.

FIGURE 5. Impedance modulus (A) and phase angle (B) of the Bode plot for the polished (o), BL-

ZrO () and BL-AlO () specimens after 5 minutes (full symbol) and 1 day (empty symbol) of

immersion in the Ringers solution. (C) is an inset of (A) using a linear-log graph.

FIGURE 6. Accumulative amount of ions released from the surfaces along 24, 48, 72, and 96 h of

immersion in the Ringers solution.



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50

100

150

200

250

300

2h 4h 6h

Adhes

ion(%)

CellViability(%)

200

400

600

800

1000

1200

1 Day 4 Days 7 Days

*

polished BL-ZrO BL-AlO

50

100

150

200

250

300

50

100

150

200

250

300

2h 4h 6h

Adhes

ion(%)

CellViability(%)

200

400

600

800

1000

1200

200

400

600

800

1000

1200

1 Day 4 Days 7 Days

*

polished BL-ZrO BL-AlO

Fig. 3. Cell attachment (A) and viability (B) on stainless steel surfaces.Saos-2 cells

were cultured on polished ( ), BL-ZrO ( ) and BL-AlO ( ) samples for the indicated

incubation periods. The results are expressed as the percentage of the fluorescence

measured on PS at 2 h or 1 day, which was given the arbitrary value of 100. Each data

represent the mean S.D. of four independent experiments. * p < 0.05 compared to PL

specimens.



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Figure 4. Nyquist plots of polished specimens (A), specimens blasted with ZrO2(B), and the

specimens blasted with Al2O3 (C), after 5 minutes and 1 day of immersion in Ringers

solution. D) Electrical equivalent circuit used to analyze the experimental impedance data.

0,0 5,0x105

1,0x106

1,5x106

2,0x106

0,0

5,0x105

1,0x106

1,5x106

2,0x106

A)

-jZimag

/ohms

Zreal/ ohms

PL_ 5 minutesFit data, 5 minutes

PL_1 dayFit data, 1 day

0,0 2,0x105

4,0x105

0,0

2,0x105

4,0x105

B)

BL-ZrO_ 5 minutesFit results, 5 minutesBL-ZrO_ 1 dayFit results, 1 day

-jZimag

/ohms

Zreal/ ohms

0,00 2,50x104

5,00x104

7,50x104

0,00

2,50x104

5,00x104

7,50x104

-jZimag

/ohms

Zreal/ ohms

BL-AlO_5 minutes

Fit results, 5 minutesBL-AlO1 dayFit results, 1day

C)D)



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Figure 5. Impedance modulus (A) and phase angle (B) of the Bode plot for the polished (o),

BL-ZrO () and BL-AlO () specimens after 5 minutes (full symbol) and 1 day (empty

symbol) of immersion in the Ringers solution. (C) is an inset of (A) using a linear-log graph.

10-3

10-2

10-1

100

101

102

103

104

105

0

30

60

90PL_5min

PL_1day

BL-ZrO_5min

BL-ZrO_1day

BL-AlO_5min

BL-AlO_1day

Phase/degrees

Frequency / Hz

(B)

10-3

10-2

10-1

100

101

102

103

104

105

102

103

10

4

105

106 PL_5min

PL_1day

BL-ZrO_5min

BL-ZrO_1day

BL-AlO_5min

BL-AlO_1day

(A)

|Z|/o

hms

Frequency / Hz

10-3

10-2

10-1

100

5.0x105

1.0x106

1.5x106

2.0x106

2.5x106

PL_5min

PL_1day

BL-ZrO_5min

BL-ZrO_1day

BL-AlO_5min

BL-AlO_1day

(C)

|Z|/ohmscm

2

Frequency / Hz



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0.0

0.4

0.8

1.2

Polished24 h

48 h

72 h

96 h

Element released

0.0

0.4

0.8

1.2

BL-ZrO

Al Cr Fe Mn Mo Ni

0.0

0.4

0.8

1.2BL-Al

2O

3

A

mountreleased(g.m

l-1)

Figure 6. Accumulative amount of ions released from the surfaces along 24, 48, 72, and 96 h

of immersion in the Ringers solution.



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Table 1

Topographic surface parameters of the investigated surfaces as a function of the scanned area.

Polished

Scanned

area

(mm2)

Ra

(m)

Rq

(m)

Rz

(m)

Rt

(m)

Ssk Real surface

area

(mm2)

Area

Index

Area

increase

(%)

1.092 0.005 0.007 0.155 0.275 0.55 1.097 - -

0.068 0.004 0.005 0.114 0.282 0.068 - -

0.011 0.003 0.004 0.098 0.224 0.011 - -

BL-ZrO

1.092 1.3 1.6 16.8 24.2 0.09 1.209 0.012 1.11 11

0.068 1.2 1.5 12.8 16.2 0.087 0.002 1.28 28

0.011 0.9 10.1 8.0 8.9 0.014 0.001 1.27 27

BL-AlO

1.092 10.5 13.3 101.7 115.7 -0.32 1.225 0.278 1.12 12

0.068 7.9 9.9 63.3 71.4 0.140 0.019 2.06 106

0.011 5.2 6.5 37.6 49.0 0.030 0.005 2.73 173



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Table 2. EEC parameter values, relative errors and chi-square values obtained by

modeling the experimental impedance data showed in Figures 4 and 5 for specimens

immersed in the Ringers solution during 5 minutes (a) and 1 day (b).

(b) R1/

% Rel

Err

R2 / % Rel

Err

Y0-CPE /

-1sn% Rel

Err

n % Rel

Err

Chi-

squarevalues

Polished 56.0 0.70 9.36106 8.79 3.1010-5 0.55 0.923 0.16 6.4010-4

BL-ZrO 62.9 0.71 1.28108 506.0 8.6810-5 0.58 0.836 0.20 7.4210-4

BL-AlO 57.3 0.37 2.03106 13.60 3.0610-4 0.30 0.846 0.12 3.4710-4

(a) R1/

% Rel

Err

R2 / % Rel

Err

Y0-CPE /

-1sn

% Rel

Err

n % Rel

Err

Chi-

squarevalues

Polished 68.5 1.08 1.78106 4.37 3.6110-5 0.92 0.913 0.28 2.5910-3

BL-ZrO 60.1 1.30 1.13106 7.38 7.1610-5 1.07 0.833 0.36 2.3310-3

BL-AlO 61.3 0.80 2.74105 8.75 2.6610-4 0.76 0.762 0.33 9.7410-4

grit blasting of medical stainless steel - implications on its corrosion behavior, ion release and...

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