J Mater Sci: Mater Med (2015) 26:195 DOI 10.1007/s10856-015-5527-y

BIOMATERIALS SYNTHESIS AND CHARACTERIZATION

Original Research

Stainless steel surface biofunctionalization with PMMA-bioglass coatings: compositional, electrochemical corrosion studies and microbiological assay L. Floroian1 • C. Samoila2 • M. Badea3 • D. Munteanu2 • C. Ristoscu4 F. Sima4 • I. Negut4,5 • M. C. Chifiriuc6 • I. N. Mihailescu4



Received: 28 January 2015 / Accepted: 3 June 2015 / Published online: 18 June 2015 Ó Springer Science+Business Media New York 2015

Abstract A solution is proposed to surpass the inconvenience caused by the corrosion of stainless steel implants in human body fluids by protection with thin films of bioactive glasses or with composite polymer-bioactive glass nanostructures. Our option was to apply thin film deposition by matrixassisted pulsed laser evaporation (MAPLE) which, to the difference to other laser or plasma techniques insures the protection of a more delicate material (a polymer in our case) against degradation or irreversible damage. The coatings composition, modification and corrosion resistance were investigated by FTIR and electrochemical techniques, under conditions which simulate their biological interaction with the human body. Mechanical testing demonstrates the adhesion, durability and resistance to fracture of the coatings. The coatings biocompatibility was assessed by in vitro studies and by flow cytometry. Our results support the unrestricted usage of coated stainless steel as a cheap alternative for human implants

& I. N. Mihailescu [email protected] 1

Faculty of Electrical Engineering and Computer Science, Transilvania University of Brasov, 29 Eroilor Blvd, 500036 Brasov, Romania

2

Faculty of Materials Science and Engineering, Transilvania University of Brasov, 29 Eroilor Blvd, 500036 Brasov, Romania

3

Faculty of Medicine, Transilvania University of Brasov, 29 Eroilor Blvd, 500036 Brasov, Romania

4

National Institute for Laser, Plasma and Radiation Physics, PO Box MG-36, 77125 Magurele, Ilfov, Romania

5

Faculty of Physics, University of Bucharest, 077125 Magurele, Ilfov, Romania

6

Department of Microbiology, Faculty of Biology, University of Bucharest, 060101 Bucharest, Romania

manufacture. They will be more accessible for lower prices in comparison with the majority present day fabrication of implants using Ti or Ti alloys.

1 Introduction A significant progress has been observed in development of materials for specific applications in medicine and biology [1]. During many years of clinical experience and evaluation of the organism’s reaction to metallic biomaterial implants, their chemical and phase compositions were subject to permanent optimization. There exists a strong correlation between the biomaterials corrosion resistance and their biocompatibility [2, 3]. In general, the human body cannot be considered a hospitable environment for an implanted metal or metal alloy. Under the action of human fluids the implants corrode and the corrosion products infiltrate into biological tissues, a process called metallosis [4]. Pathomorphological changes, dependent on the type and concentration of elements, occur in tissues close to implant. Histopathological changes were observed in the detoxication organs (liver, kidneys, spleen) [1]. Good biocompatibility is observed for metal and alloys with the high anodic potential, such as titanium and its alloys, stainless steels or cobalt–chromium alloys. They may be rather safely employed for implants within a given time span, stipulating additionally for the particular physical and chemical properties of the implants surface. Titanium and its alloys have been therefore used in human implants for many years because they are more resistant to corrosion. Nevertheless, their corrosion by dissolution of Ti and alloying elements (e.g. V, Zr or Al), still remains a concern. Indeed, the metal ions liberated into the surrounding tissue may induce the release of potentially osteolytic cytokines involved in implant loosening [5], causing discoloration of

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the tissue, and even more severe complications such as inflammatory reaction of the tissue [2]. Consequently, many studies have dealt with the corrosion resistance of titanium [6–8] and Ti alloys implants [9–11] under various conditions. Also, the plastic deformation, which occurs during the pre-operative surgery shaping process, induces changes in the structure of surface layer and its electrochemical behavior in biological environments, which may change the biocompatibility of these materials [12]. Stainless steel is a very popular metallic material because of its relatively low cost, ease of fabrication and reasonable corrosion resistance [7, 13]. However, stainless steel is susceptible to a number of localized corrosions, such as crevice or pitting [14, 15], intergranular and stress corrosion cracking [16]. Shortcomings of stainless steel materials during implantation have been reported [17] due to the high nickel content and to the aggressive biological effects. The corrosion products include iron, chromium, nickel and molybdenum and because of their effects [18, 19], stainless steel is used just as temporary implants to help bone healing, as well as fixed implants like in artificial joints. Stainless steel is therefore seldom used in many countries as permanent implants, but it is still mostly used in other countries. Here, a solution is proposed to skip the inconvenience caused by corrosion of stainless steel implants in human body by coating with thin films of bioactive glasses or with composite polymer-bioactive glasses nanostructures. The composition of the coatings was investigated by FTIR and their biocompatibility was studied by dedicated microbiological assays. The samples were mechanical tested to demonstrate the adhesion, durability and resistance to failure of the coating. The resistance to corrosion of the stainless steel implants coated with bioactive glasses or polymer-bioactive glasses using high sensitive electrochemical methods has been evaluated. Electrochemical measurements involving corrosion and electrochemical impedance spectroscopy studies were carried out in physiological medium. The corrosion resistance parameters were inferred. The involved electrochemical parameters were estimated and the electrical ones of the circuits were found out by fitting the experimental data via equivalent electric circuits.

2 Experimental procedures 2.1 Materials Studies have been conducted for both bare titanium (Ti) and stainless steel (316L) samples and also for coated 316L substrates with the two types of bioactive coatings. They were namely pure bioglass (denoted BG61/316L) and

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polymer-bioactive glass nanocomposites (denoted PMMABG61/316L), respectively. A bioglass labeled BG61 belonging to the SiO2–Na2O– K2O–CaO–MgO–P2O5 system was used in experiments. It contains 61.1 wt% SiO2, 10.3 wt% Na2O, 2.8 wt% K2O, 12.6 wt% CaO, 7.2 wt% MgO and 6 wt% P2O5 and it is produced according to a protocol described in Refs. [20, 21]. Poly(methyl methacrylate)—PMMA—presents a good biocompatibility with human tissue, but nevertheless, it is softer than the glass and is easily to scratch. Our intention was to fabricate bioactive coatings more resistant to scratching by adding to PMMA BG particles which have the ability to better bond to both bone and tissue. Medical grade stainless steel 316L disks, containing 64.26 wt% Fe, 18.51 wt% Cr, 12 wt% Ni, 2.13 wt% Mo, 1.44 wt% Mn, 0.58 wt% Cu, 0.56 wt% Si, 0.0265 wt% C, 0.0036 wt% S and other elements in smaller concentration were used as deposition substrates. The 316L disks with dimensions of (12 9 12 9 1) mm3 were mechanically processed by polishing to a roughness within lm range (Rq = 2–4 lm) and then cleaned with acetone, ethanol and deionized water in an X-Tra ultrasonic bath. This procedure was applied because it was shown that for implants is beneficial a certain roughness which can increase the biocompatibility up to 10 times. An appropriate surface morphology provides a good cell adhesion and bone growth while a suitable roughness insures a long term enhanced interaction to bone-implant. The optimum values of the roughness are different for orthopedic or dentistry applications. Such optimum roughness for dental applications is in submicrometer range whilst orthopedic applications is in millimeter range. Titanium grade 4 plates were used as reference. The chemical composition (wt%) is: Fe: 0.50, C: 0.100, O: 0.40, H: 0.015, N: 0.05 and Ti up to balance. 2.2 Preparation of bioglass and composite PMMAbioglass coatings To synthesize BG61 coatings on 316L stainless steel substrates, the bioactive glass powder was pressed in a mold in pellets of 13 mm diameter and 3 mm thick, sintered at 650 °C and used in pulsed laser deposition (PLD) experiments as targets. The bioglass films were grown with a pulsed UV excimer KrF* laser source with 248 nm wavelength, 25 ns pulse duration and a deposition chamber. Both PLD and matrix-assisted pulsed laser evaporation (MAPLE) depositions were conducted with an the same excimer KrF* laser source generating pulses of 248 nm wavelength and 25 ns duration. To preserve material’s purity, the targets were submitted to a preliminary cleaning with 1000 laser pulses. A shutter was then interposed in-

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between target and substrate, to collect the flux of ablated substance containing impurities. The laser beam hits the target under an incidence of 45°. The generated plasma plume expanded normally to target. The ablated material was collected onto the 316L substrates placed parallel to the target. In order to avoid drilling and to ensure the deposition of a uniform layer, the target was continuously rotated and translated along two orthogonal axes. The laser beam was focused in a 14 mm2 spot area on target surface. For the deposition of composite polymer-bioglass coatings onto 316L stainless steel substrates, MAPLE technique was applied. MAPLE was developed after 1998 [22] as an alternative method to PLD, compulsory for the transfer of delicate (in respect to thermal and/or biological degradation and damage) materials. It essentially differs from PLD by target preparation, laser-material interaction and transfer mechanisms. Laser acts in this case as a gentle propeller for transferring different compounds, including large molecular weight species, such as polymeric or organic molecules [23–25]. The deposition chamber was modified as appropriate to MAPLE [26, 27]. The ‘‘icy’’ composite target was kept at a constant temperature with a liquid nitrogen cooler. For MAPLE experiments, a suspension of 0.5 wt% bioglass in a solution of 3 wt% PMMA in chloroform was prepared and frozen. The best deposition conditions for PLD and MAPLE techniques are collected in Table 1.

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and quantity. A batch of samples were immersed in SBF at 37 °C and investigated after different immersion times. 2.4 Electrochemical studies The corrosion measurements were performed with a PalmSens potentiostat (Palm Instrument BV, Hoten Netherlands) controlled by PSTrace software, in a three electrodes configuration with an Ag/AgCl reference electrode (3.0 mol/kg KCl, ?210 V potential) and a platinum wire as counter electrode. The time elapse between completing electrical connections and start of measurement was 30 s. The plots were recorded with 0.002 V/s scan rate and the working potential was varied from -1 to ?2 V versus Ag/AgCl while the samples are immersed in SBF used as electrolyte. The electrodes are square and have 0.8 cm2 active surface area. The corrosion extension of program PS Trace ensures the possibility to conduct specific types of corrosion measurement and analysis of recorded curves. The Stern–Geary

Table 3 Reagents and quantities used in preparing of 1L SBF with pH 7.4

2.3 Surface analysis To get information about sample surface composition, FTIR analyses were conducted using a Nicolet 380 apparatus equipped with an orbit ATR (diamond crystal), wave number range 7800–350 cm-1, spectral resolution 0.4 cm-1, S/N ratio 20,000:1. The spectra were acquired in the absorbance mode. The analysis was carried out in simulated body fluid (SBF) with an ionic composition identical to that of blood plasma (Table 2). SBF is prepared after Kokubo formula [28] by mixing the reagents in Table 3, respecting the order

Order

Reagent

Quantity

1

H2O2

750 mL

2

NaCl

7.996 g

3

NaHCO3

0.350 g

4

KCl

0.224 g

5

K2HPO4 ? 3H2O

0.228 g

6

MgCl2 ? 6H2O

0.305 g

7

1 kmol/m3 HCl

40 cm3

8

CaC2

0.278 g

9

Na2SO4

0.071 g

10

(CH2OH)3CNH2

6.057 g

11

1 kmol/m3 HCl

Appropriate amount for adjusting pH

Table 1 Optimum deposition parameters Deposition regime

Sample code

Substrate temp. (°C)

Laser fluence (J/cm2)

Ambient pressure (Pa)

Separation distance (m)

Target rotation frequency (Hz)

No. of subsequent laser pulses

PLD

BG61/316L

400

2.8

13

0.04

0.4

10,000

MAPLE

PMMA-BG61/316L

RT

0.55

2.7

0.03

1.3

5000

Table 2 Ion concentrations of used SBF versus plasma blood [28]

Ions

Na?

K?

Mg2?

Ca2?

Cl-

HPO2 4

SO2 4

HCO 3

Composition (mM)

142

5

1.5

2.5

147.8

1

0.5

4.2

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equation was used for the analysis of the measured curve and the values of the corrosion parameters were optimized so that the sum of squares of the differences between the experimental and theoretical values is the lowest possible. The best fit is shown in the plot and the optimum corrosion parameters are given in a table. In order to obtain reliable statistics, all corrosion measurements were carried out in triplicate for each type of samples. The mean value of corrosion parameters (corrosion potential Ecorr and corrosion current Icorr) along with the standard deviation values were calculated. Corrosion current is an important parameter because the corrosion rate is calculated using Icorr. In addition, electrochemical impedance spectroscopy (EIS) analysis was conducted three times for each sample in the same freshly filled with SBF electrochemical cell using PC-controlled system Autolab PGSTAT 100 Eco Chemie by means of FRA 4.9 software. An alternating voltage with 0.01 V rms amplitude was applied, scanning from 10,000 to 0.1 Hz with 10 points per frequency decade, and auto-integration time of 5 s. The software calculates and records the real and imaginary parts of electrochemical impedance (Zreal and Zimag) together with the phase and represents them in Nyquist and Bode diagrams. Based upon the principles of electrochemical spectroscopy, the equivalent electric circuit that best fits the experimental data was found with the help of FRA 4.9 software [29] and optimum electrical parameters were inferred: electrical resistance of the solution, charge transfer resistance and constant phase element. 2.5 Mechanical testing A nanoindentation device produced by CSM Instruments (NHT-2) equipped with a Berkovich diamond tip was used for nanoindentation studies. To minimize substrate contributions, the indentation experiments were performed controlling the penetration depth of the indenter, which was limited to maximum 20 % of the total film thickness. In order to minimize the creep effect, a pause of 2 s was set after the loading stage of the measurement. On each sample, 5 progressive multicycles have been performed, with the following protocol: approach speed 2 lm/min, loading rate 1 mN/min, unloading rate 20 mN/min. The adhesion of the films to the silicon substrate was assessed using a Micro Scratch Tester (CSM Instruments) with a 100Cr6 steel tipped indenter with a Rockwell geometry (tip radius = 100 lm). The dynamic friction coefficient values were obtained using a ball-on-disc tribosystem (CSM Instruments). The counter-body was a 6 mm diameter steel ball (AISI 100Cr6) hardened and low tempered. Taking into account that the practical aim of these kinds of coatings is the tribological functional character, the normal load applied by

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the ball on the sample surface was progressive between 0.03 and 30 N, with loading rate of 10 N/min and 1 mm/ min speed while the maximum distance covered was 2000 m. 2.6 Biological assay The cell morphology was examined using an inverted Observer. D1 Carl Zeiss microscope. 5 identical specimens were prepared for each biological test. 2.6.1 Cytotoxicity assay The biological compatibility of the obtained material was assessed by growing human Wharton’s Jelly-derived Mesenchymal Stromal Cells (WJ-MSCs) on surface. The obtained specimens were sterilized by UV irradiation and placed in 35 mm diameter Petri dishes. In each Petri dish 3 9 105 mesenchymal cells were incubated. The monolayer morphology was evaluated after 24 h, by fixing the cells with 70 % alcohol and staining the monolayer with 5 lg/mL propidium iodide (PI). The stained specimens were examined by fluorescent microscopy and photographed in UV field according to the protocol in Ref. [30]. 2.6.2 Proliferation assay Each specimen was placed in 35 mm diameter Petri dishes. In each Petri dish 3 9 105 mesenchymal cells were added and incubated for 72 h, to allow the proliferation over tested devices. Thereafter, the cells were removed from the device by tripsinisation and analyzed by flow cytometry for the cellular cycle. To this purpose, the cells were washed twice in PBS, and then incubated 15 min, at 37 °C, with RNAse A (100 lg/mL), and for 1 h with PI (100 lg/mL). After contact of cells with PI, the acquisition was performed using Epics Beckman Coulter flow cytometer. Data were analyzed using FlowJo software and expressed as fraction of cells in the different cell cycle phases.

3 Results and discussion 3.1 FTIR studies The spectrum of initial BG61/316L film (noted with 0 in Fig. 1a) shows peaks at 768, 982 and 1008 cm-1 which belong to Si–O bending vibration modes. After 1 day of immersion in SBF, one notices for BG61/ 316L samples a higher amplitude of all peaks compared to initial spectrum (curve 1 in Fig. 1a). In accordance with

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(a) 40

Page 5 of 14

Si-O

(PO4)

35

40

21 1

OH

3 BG61 / 316L

2-

20 15

7

10

7

Si-O

35

(CO3)

25

PMMA-BG61 / 316L

30 25

Si-O

3

20

(PO4)

15

3-

35

10

5 0

(b) 45

Absorbance (a.u.)

Absorbance (a.u.)

30

3-

195

0 600

5

Si-O 700

800

900 1000 1100 1200 1300 1400 -1

Wavenumber (cm )

0 600

21

0 700

800

900

1000 1100 1200 1300 1400 -1

Wavenumber (cm )

Fig. 1 FTIR spectra of BG61/316L (a) and PMMA-BG61/316L (b) after different immersion days in SBF

literature, the glass reacts with SBF and both chemical and structural changes occur on glass surface in time [31]. First, there is a rapid exchange of Ca2? and Na? ions from BG61 with H3O? or H? from SBF: SiONaþ þ Hþ OH ! SiOH þ Naþ ðaqÞ þ OH This is indicative for the formation of a detectable superficial layer where all elements have a larger concentration. After 3 and 7 days of immersion in SBF, the cations exchange increases the hydroxyl concentration in the solution which leads to the attack of the silica glass network: SiOSi þ H2 O ! SiOH þ OHSi Accordingly, major transformations could be clearly observed on the surface of the coatings (curves 3 and 7 in Fig. 1a). The bioglass peaks at 982 and 1008 cm-1 decrease and new peaks appear at 1045 cm-1 that belong to asymmetric stretching of P-O bond in (PO4)3-, at 602 cm-1 assigned to the vibrational mode of OH and at 813 cm-1 corresponding to (CO3)2- group. All these peaks belong to hydroxyl carbonate apatite (HCA), which is the main component of the bone [32]. Thus, after 7 days of immersion in SBF, the spectrum indicates a large drop of silica content and the start of a new HCA layer growth. After 21 days of immersion in SBF, the amorphous CaO-P2O5 film crystallizes by incorporating OH- and CO2 ions from solution to form a hydroxyl carbonate 3 apatite layer. In spectra recorded after 21 days of immersion, one can recognize HCA peaks only, while all bioglass peaks vanish.

The surface of PMMA-BG61/316L samples exhibited similar transformations as BG61/316L ones, but in a slower process, the HCA new layer appearing after 21 days (Fig. 1b). 3.2 Electrochemical polarization measurements The electrochemical corrosion studies of individual samples were performed in electrochemical cell by immersing Ti, 316L, BG61/316L or PMMA-BG61/316L samples in SBF as working electrode while cyclic voltammograms were recorded at 37 °C (Fig. 2). All the samples were immersed in SBF and analyzed after 0, 3, 7, 14 and 21 days of immersion. The corrosion parameters inferred using PSTrace software were collected in Table 4. As known, the main corrosion parameters are: the corrosion potential (Ecorr—the potential where the sum of the anodic and cathodic reaction rates on the electrode surface is zero), corrosion current (Icorr given by the anodic current density at the potential of Ecorr) and density of corrosion current (icorr) obtained from the potentiodynamic polarization curves. All current densities were normalized to the surface area. Low values of standard deviation have been obtained, which signify that the results are quite reproducible. Tafel plots for bare Ti, bare 316L and 316L covered with BG61 and PMMA-BG61 layers before immersion are shown in Fig. 3a. The dependence of the corrosion behavior on electrode immersion time in SBF for all electrodes is given in Fig. 3b. In case of titanium work electrode, icorr rapidly decreases from 1.32 to 0.84 lA/cm2 after the first 3 days of immersion in SBF. This is a rather low corrosion current density pointing to a high corrosion resistance of titanium,

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Fig. 2 Cyclic voltammograms for 316L (a) and 316L coated with BG61 and PMMA-BG61 (b) samples

consistent with previous results [6–11, 33–36]. When a freshly polished Ti surface is exposed to moist air, a thin Ti-oxide film is spontaneously growing on its surface. After immersion of titanium electrode into the electrolyte, a titanium oxide film starts to grow on the electrode surface. The shielding effect of the oxide film, and therefore the corrosion resistance of the electrode, increases and promotes the decrease of the anodic dissolution current of titanium. During next weeks, icorr gradually increases reaching a value of about 1.12 lA/cm2 at 14 days, which points to a slow corrosion process of titanium allowing for the release of metallic ions in the body. This process is confirmed by the decrease of corrosion potential Ecorr, from -346 mV/Ag/AgCl to -398 mV/Ag/AgCl. Because stainless steel work electrode has an initial value of icorr = 13.41 lA/cm2, much higher than the value for titanium, it follows that its corrosion resistance is much weaker. In the same time, the corrosion potential Ecorr = -507 mV/Ag/AgCl, inferior to that of titanium, decreases with immersion time down to -695 mV/Ag/ AgCl. The corrosion current density increases, reaching a value of 18.20 lA/cm2. It results that both passivity and corrosion resistance of 316L are lower than for Ti. This is why Ti is preferred for implants manufacture. Significant changes appear for 316L samples behavior after coating with bioactive glass films. From Fig. 3a and Table 3 one can see that Ecorr for 316L coated with BG61 is higher than for Ti, while corrosion current density is initially with one order of magnitude lower than for Ti. The corrosion current density has a small rise in first 3 days of immersion in SBF and then significantly decreases reaching after 14 days of immersion a value of icorr = 0.10 lA/ cm2, much lower than for Ti. The changes are more important for 316L coated with polymer-bioglass nanocomposite layer. The corrosion current density drops one order of magnitude as compared

123

to bare Ti, and continues to slowly decrease with immersion time, reaching values 2 orders of magnitude lower after 14 days of immersion. 3.3 Electrochemical impedance spectroscopic measurements The electrochemical evolution of the samples under conditions which simulate their biological interaction with the human body was analyzed. To this purpose, samples of Ti, 316L and 316L coated with BG61 or PMMA-BG61 thin films were introduced in SBF at 37 °C and the electrochemical impedance spectra were recorded after different immersion times. All measurements were carried out in triplicate, in the same single compartment electrochemical cell, freshly filled with appropriate solution. The EIS plots (the Nyquist diagrams) for all electrodes were recorded in the open circuit potential (OCP) configuration presented in Fig. 4 The equivalent electrical circuits that fit the plots of electrochemical impedance were found using FRA 4.9 software. The Nyquist diagram of Ti electrode in the open circuit potential given in Fig. 4a is characterized by only one time constant. Consequently, congruent to other references in literature for titanium [37], the Randles electrical-equivalent circuit with only one time constant (Fig. 5a) was used to model the experimental spectra. The electric elements have the following meaning: Rs is the resistance of the electrolyte between the working and reference electrodes, R1 is the charge transfer resistance related to the rate of corrosion reactions at the OCP, and it is inversely proportional to the corrosion current, and Q is the capacitance represented by the constant-phase element (CPE). As a rule, the use of CPE is required to take into account the relaxation times distribution as a result of microscopic inhomogeneities presence at the oxide–

279 112 0.05 -262 267 357 0.08 -276 430 620 0.08 -288 143 214 -301 PMMABG61/ 316L

0.09

166

895 2084

333 0.10

18.20 -695

-316 48

177 330

253 0.21

14.71 -620

-325 132

135 230

450 0.75

14.17 -552

-336 105

99 194

208

-507

-349

316L

BG61/316L

13.41

Page 7 of 14

0.32

397 427 1.12 -398 426 567 0.97 -369 398 439 0.84 -346 279 1.32 -357 Ti

112

Ecorr (mV/ Ag/AgCl) Ecorr (mV/ Ag/AgCl) icorr (lA/ cm2) Ecorr (mV/ Ag/AgCl) icorr (lA/ cm2) Ecorr (mV/ Ag/AgCl) Tafel slopes

ba (mV/ dec)

bc (mV/ dec)

3 days 0 days Time

Table 4 Corrosion parameters for electrodes at different immersion time

ba (mV/ dec)

bc (mV/ dec)

7 days

icorr (lA/ cm2)

ba (V/ dec)

bc (V/ dec)

14 days

icorr (lA/ cm2)

ba (mV/ dec)

bc (mV/ dec)

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195

electrolyte interface and beneath the oxide phase. This is a non-ideal capacitor of capacitance Q, with ZCPE = [Q(ix)n]-1, where ZCPE is the impedance of a CPE. n parameter describes the extent of time constant distribution. A normal distribution of the time constant can arise from porosity or changes in the conductivity of surface layers (i.e. oxide layers). The values of Rs, R1, Q and n obtained from fits were collected in Table 5. The maximum phase angle was inferred in each case using Bode plots (not shown here) and is given in the same table. One should mention that the parameters in the table were normalized to the surface area of each specimen. A remarkable good agreement between fitted and experimental data was obtained. In case of Ti electrode, the maximum phase angle has varied after first day of immersion in SBF from 69° to 73°, because of passive TiO2 film growing up on the titanium surface. After the next 2 days, the maximum phase angle falls to 65°, that it is indicative for destruction of the weak passive layer by Cl- ions from SBF, resulting in localized corrosion, in agreement with our corrosion studies on Ti introduces in the previous section. The local breakdown of this passive layer leads to increase of Q value (the layer thickness and capacitance are inverse proportional) and to decrease of R1 value. In contrast with the Ti, for 316L electrode, no passivation in SBF was observed. The maximum phase angle continuously decreases from 67° to 50°, the charge transfer resistance decreases and capacitance increases (Fig. 4). All together is indicative for the electrode corrosion process in SBF. For BG61/316L sample, initial maximum phase angle is 78° suggesting the presence on surface of a highly stable film, behaving like pure capacitive impedance. The Nyquist diagram (Fig. 4c) is a beginning of a semicircle with a very large radius, the charge transfer resistance reaches a high value of 8302.30 kX cm2, while the capacitance has a very small value of 22.43 lF cm2. The n factor inferred with the software FRA 4.9 is 0.84. This is because the 316L surface has numerous inhomogeneities, introduced by mechanical polishing applied before bioglass deposition to increase the surface active area and bioactivity. The equivalent circuit is the same Randles circuit from Fig. 5a with the electrical elements given in Table 5. After 1 day of immersion in SBF at 37 °C, ion exchange between BG61 layer and electrolyte induces a decrease of the charge transfer resistance and an increase of the capacitance which reach values of 737.97 kX cm2 and 63.32 lF cm2, respectively. This certifies the formation of a detectable superficial layer where all elements have a larger concentration, in accordance with results of FTIR studies mentioned in paragraph 3.1. The Nyquist diagram is characterized by two well defined time constants and two peaks are present in the Bode plot at

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Fig. 3 a Tafel plots for bare Ti (A), bare 316L (B) and 316L covered with BG61 (C) and PMMA-BG61 (D) layers before immersion; b Changes of Icorr values for bare Ti (A), bare 316L (B) and 316L covered with BG61 (C) and PMMA-BG61 (D) layers after different immersion times in SBF

(a)

(b)

60

3

Ti

50

40

2

-Z" (kohm cm )

7

21

10 kHz

30

0.1 Hz

20

1

0

10

0

0

10

20

30

40

50

60

2

Z' (kohm cm )

(c)

(d)

250 10kHz

0

BG61/316L

200

10kHz

0

PMMA-BG61/316L

35

200

21

2

-Z" (kohm cm )

2

-Z" (kohm cm )

250

21

150

1 3

100

150

7

100

1

7

50

50

3

0.1Hz

0

0

50

100

150 2

Z' (kohm cm )

200

250

0

0

50

100

150

0.1Hz

200

250

2

Z' (kohm cm )

Fig. 4 EIS plot of a bare Ti, b 316L, c BG61/316L and d PMMA-BG61/316L thin films after 0, 1, 3, 7, 21 and 35 days of immersion in SBF

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Page 9 of 14

195

Fig. 5 The equivalent circuit a Rs(Q1R1) and b Rs(Q1[R1(R2Q2)]) used for fitting the EIS spectra of the samples

Table 5 Electrochemical impedance spectroscopy data of Ti, 316L and BG61/316L samples after various immersion times in SBF at 37 °C

Sample

Time Equivalent circuit

0 days Rs(Q1R1)

1 day Rs(Q1R1)

3 days Rs(Q1R1)

7 days Rs(Q1R1)

21 days Rs(Q1R1)

Ti

Rs (X cm2)

60.65

58.12

61.11

60.18

61.21

R1 (kX cm2)

98.21

372.11

77.42

37.92

11.51

2

316L

BG61/316L

Q (lF cm )

169.72

131.23

140.07

223.78

291.86

n

0.84

052

0.67

0.48

036

Max phase angle (deg)

69

73

65

61

60

Rs (X cm2)

63.20

66.13

61.08

62.31

63.22

R1 (kX cm2) Q (lF cm2)

621.45 135.12

151.19 150.17

22.53 118.66

15.01 119.32

10.97 127.79

n

0.80

0.78

0.56

0.42

0.40

Max phase angle (deg)

67

63

57

52

50

Rs (X cm2)

841.19

810.15

821.49

894.38

880.75

R1 (kX cm2)

8302.30

737.97

680.21

125.16

17,157.22

Q (lF cm )

22.43

63.32

78.22

95.23

24.14

n1

0.84

0.80

0.78

0.76

0.82

R2 (kX cm2)

422.54

450.11

712.65

Q2 (lF cm2)

35.83

29.61

11.14

n2

0.56

0.59

066

61

62

66

37

35

28

2

Max phase angle (deg)

61° and 37°. The equivalent circuit is now Rs(Q1[R1(R2Q2)]) from Fig. 5b. It refers to a layer (Q1) with charge transfer at the bottom of the pores (R2//Q2), at the stainless steel-solution interface. A similar electrical equivalent circuit describes the processes which take place at film-electrolyte interface after 3 days of sample immersion in SBF, with slightly different electrical parameters. After 7 days of immersion in SBF, the Nyquist diagram shows two well defined time constants and the maximum phase angles become 66° and 28° because two processes take place simultaneously: the BG61 dissolution in SBF and the adsorption of some electrolyte ions on surface. The BG61 dissolution causes the flat surface degradation and the n factor gets a much smaller value of 0.66. The equivalent circuit is now Rs(Q1[R1(R2Q2)]) from Fig. 5b where for the porous layer (Q1) the charge transfer resistance decreases and the capacitance increases, reaching values of 125.16 kX cm2 and 95.23 lF cm2, respectively. These are indicative for a continuous degradation of the BG61 coating. The second process, with R2 and Q2 parameters, points to the formation of a CaO–P2O5 film.

78

69

After 21 days of immersion in SBF, the amorphous CaO-P2O5 film crystallizes by incorporating OH- and CO2 ions from solution to form a hydroxyl carbonate 3 apatite layer. One gets a single time constant Nyquist diagram, and the growth of the HCA layer only takes place. The new apatite layer has an n factor of 0.82 and presents a charge transfer resistance of 17,157.22 kX cm2 and a double layer capacitance of 24.14 lF cm2. The EIS fitting generated an equivalent circuit Rs(Q1R1). PMMA-BG61/316L exhibited the same behavior as BG61/316L samples. Initially, the sample acts as pure capacitive impedance suggesting the presence on surface of a highly stable film. The Nyquist diagram (Fig. 4d) is a beginning of a semicircle with a very large radius, the charge transfer resistance reaches a high value of 9752.17 kX cm2, while the capacitance has a very small value of 23.59 lF cm2. The equivalent circuit is the Randles circuit in Fig. 5a with the electrical elements given in Table 6. After 1 day of immersion in SBF at 37 °C, ion exchange between PMMA-BG61 layer and electrolyte induces a

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Table 6 Electrochemical impedance spectroscopy data of PMMA-BG61 samples after various immersion times in SBF at 37 °C

Time Equiv. circ.

0 days Rs(Q1R1)

1 day Rs(Q1R1)

3 days 7 days Rs(Q1[R1(R2Q2)])

21 days

35 days Rs(Q1R1)

Rs (X cm2)

69.13

67.26

68.15

67.32

63.56

67.54

R1 (kX cm2)

9752.17

724.11

107.56

36.23

21.34

21,296.25

Q (lF cm2)

23.59

51.98

77.15

115.82

123.14

38.19

n1

0.90

0.86

0.86

R2 (kX cm2)

0.84

0.72

0.68

429.23

1014.81

1822.13

Q2 (lF cm2)

63.08

38.14

13.23

n2

0.38

0.44

0.56

64

62

58

31

28

37

Max phase angle

80

67

(deg)

70

decrease of the charge transfer resistance and an increase of the capacitance. They reach values of 724.11 kX cm2 and 51.98 lF cm2, respectively, while the layer behavior is described by the same Randles circuit. After 3, 7 or 21 days of immersion in SBF, the Nyquist diagram shows two semicircles that suggest two simultaneous processes: the BG61 dissolution in SBF and the adsorption of some electrolyte ions on surface. The equivalent circuit is Rs(Q1[R1(R2Q2)]) from Fig. 5b where, in case of the porous layer (Q1) the charge transfer resistance decreases and the capacitance increases, reaching values of 21.34 kX cm2 and 123.14 lF cm2, respectively. This behavior suggests the continuous degradation of the BG61 coating. The second process, with R2 and Q2 parameters, supports the formation of a CaO–P2O5 film. After 35 days of immersion in SBF, we got a single time constant Nyquist diagram and the growing of HCA layer. The new apatite layer has an n factor of 0.86 and exhibits a high charge transfer resistance (21296.25 kX cm2) and a double layer capacitance of 38.19 lF cm2. The EIS fitting generated an equivalent circuit Rs(Q1R1).

During the scratch tests it was not either delamination or cracks. The determined critical loads are given in Table 7. Scratch test showed a very good adhesion to the stainless steel substrate of thin films of BG61 and PMMA-BG61. These values are comparable with those found in literature [39, 40]. Tribological tests demonstrate good resistance to friction for both types of films, proper to static implants, without moving parts. All of these suggest that BG61 and PMMA-BG61 thin films are suitable from mechanical point of view to be used for implant coating.

3.4 Mechanical characterization

4 Discussion

In nanoindentation test the hardness Hit and indentation elastic modulus Eit for BG61/316L and PMMA-BG61/ 316L samples were determined following the model of Oliver and Pharr [38]. The obtained values showed that BG61 has a higher hardness than PMMA-BG61 because of its larger Si content, but it exhibits a lower elasticity (Table 7).

Our studies demonstrated that bare 316L is not appropriate for implant use. Indeed, the aseptic loosening after implantation is induced. We have shown that this undesired effect during immersion in SBF can be removed by covering 316L substrates with BG61 or PMMA-BG61 as protective layers. Thus, it was observed an intense corrosion process of 316L bare electrode (Fig. 2a), a very slow corrosion process for BG61/316L (ten times slower) and a negligible corrosion process in case of PMMA-BG61/316L (Fig. 2b) due to PMMA ? BG61 protection. We have observed that the thin BG61 and PMMA-BG61 layers behave as good insulators. These prove that PMMA-BG61 nanocomposite coatings confer an enhanced protection of 316L against corrosion than simple bioglass layer.

Table 7 Mechanical parameters of samples Sample

H (MPa)

E (GPa)

Lc (mN)

l

BG61/316L

3451.93

96.57

86.14

0.56

989.12

71.32

90.04

0.45

PMMA-BG61/316L

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3.5 Microbiological assay The morphology and growth of the WJ-MSCs on obtained structures was not affected as against control (Fig. 6). The microscopic results were confirmed by the flow cytometry assay of the cellular cycle, showing no changes in the distribution of the growing phases (Fig. 7).

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Fig. 6 Fluorescence microscopy images of WJMSCs grown on different substrates. Magnification: 9100 (left column), 9630 (right column)

(a) Control cells

(b) Control cells

(c) 316L

(d) 316L

(e) PMMA/316L

(f) PMMA/316L

(g) BG61

(h) BG61

(i) PMMA-BG61/316L

(j) PMMA-BG61/316L

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Fig. 7 Effects of obtained specimens on cell cycle progression assessed by flow cytometry

On the other hand, the FTIR data after immersion in SBF suggest the dissolution of the bioglass and/or the presence on surface of a freshly growing layer, similar  to carbonated hydroxyapatite Ca10x ðPO4 Þ6x CO2 3 x  OH 2x which is a prevalent mineral component of vertebrate bones [32]. The growth of this layer demonstrates

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the ability of the material to firmly bind to tissue via a bioactive fixation by a chemical bond at the bone-implant interface. Next, our results are consistent with other experimental observations on bioglasses [31, 39] pointing to a loss of soluble silica in SBF. The cations exchange increases the

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hydroxyl concentration in the solution, which leads to the attack of the silica glass network: SiOSi þ H2 O ! SiOH þ OHSi: The condensation and repolymerisation of a SiO2-rich layer on surface, the migration of PO43- and Ca? ions through this layer are followed by adsorption on surface and a CaO–P2O5-rich film growth. This is also confirmed by the parameters of the equivalent circuit (Fig. 5) supporting the formation on surface of a highly stable compact film, with characteristics close to pure capacitive impedance that protects implant against environmental agents. PMMA was not dissolved in SBF but remained on 316L surface. Our systematical biological assays have shown that all deposited films are fully biocompatible. In the opinion of the authors, the reported physical chemical results and in-vitro assessment support the possible use of MAPLE coating method for long-term implantation with rather cheap and easy access material implants. Nevertheless, systematic in vivo studies are mandatory before a definite decision. One may conclude that after coating with BG61 or better with PMMA-BG61 films, the cheap 316L exhibits superior anticorrosive protection and can therefore efficiently replace Ti for implants.

5 Conclusions The formation of a hydroxyl carbonate apatite superficial layer was put in evidence by FTIR studies of BG61/316L and PMMA-BG61/316L samples after immersion in SBF at 37 °C as an effect of ion exchange with the fluid, concurrently with the bioglass decomposition. To our opinion, PMMA was not released in SBF but rather preserved on 316L surface. Electrochemical impedance spectroscopy measurements demonstrated that, as an effect of immersion in biological fluids, the stainless steel corrosion is slowed down and the metal ions release into tissues is inhibited. More precisely, the nanostructured PMMA-BG61/316L coating ensures the preservation of the stainless steel surface bioactivity and serves as an efficient shield against corrosion. The full biocompatibility of MAPLE deposited nanostructures was confirmed by both microscopic and flow cytometry investigations. Moreover, mechanical tests demonstrated the adhesion, durability and resistance to failure of the coatings. Our results support the use of stainless steel coated with bioglass-based and especially with PMMA-bioglass laser deposited thin films as a challenging alternative for fabrication of reliable and cheap human implants and prostheses.

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Acknowledgments CR, FS, IN and INM acknowledge with thanks the financial support by UEFISCDI under the contracts ID 304/2011, and PCCA 244/2014. This work is also supported by the Sectorial Operational Programme Human Resources Development (SOP HRD), financed from the European Social Fund and by the Romanian Government under the Project Number POSDRU/159/1.5/S/134378.

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Stainless steel surface biofunctionalization with PMMA-bioglass coatings: compositional, electrochemical corrosion studies and microbiological assay.

A solution is proposed to surpass the inconvenience caused by the corrosion of stainless steel implants in human body fluids by protection with thin f...
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