Friction Stir Processing of Stainless Steel for Ascertaining Its

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Friction Stir Processing of Stainless Steel for Ascertaining its Superlative Performance in Bio-Implant Applications Gopinath Perumal, Aditya Ayyagari, Amrita Chakrabarti, Deepika Kannan, Soumya Pati, Harpreet Singh Grewal, Sundeep Mukherjee, Shailja Singh, and Harpreet Singh Arora ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.7b11064 • Publication Date (Web): 03 Oct 2017 Downloaded from http://pubs.acs.org on October 4, 2017

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ACS Applied Materials & Interfaces

Friction Stir Processing of Stainless Steel for Ascertaining its Superlative Performance in Bio-Implant Applications G. Perumal1, A. Ayyagari2, A. Chakrabarti3, D. Kannan3, S. Pati3, H.S. Grewal1, S. Mukherjee2, S. Singh3 ,4, # and H. S. Arora1, * 1

Department of Mechanical Engineering, School of Engineering, Shiv Nadar University, Uttar Pradesh 201314, India 2 Department of Materials Science and Engineering, University of North Texas, Denton, Texas 76203, USA 3 Department of Life Sciences, School of Natural Sciences, Shiv Nadar University, Uttar Pradesh 201314, India 4 Special Center for Molecular Medicine, Jawaharlal Nehru University, New Delhi, 110067, India *

Email: [email protected], Phone: (+91)-8130625504 # Email: [email protected], Phone: (+91)-9868512025

Abstract Substrate-cell interactions for a bio-implant are driven by substrate’s surface characteristics. In addition, the performance of an implant and resistance to degradation is primarily governed by its surface properties. A bio-implant typically degrades by wear and corrosion in the physiological environment resulting in metallosis. Surface engineering strategies for limiting degradation of implants and enhancing their performance may reduce or eliminate the need for implant removal surgeries and the associated cost. In the current study, we tailored the surface properties of stainless steel using submerged friction stir processing (FSP), a severe plastic deformation technique. FSP resulted in significant microstructural refinement from 22 µm grainsize for the as-received alloy to 0.8 µm grain-size for the processed sample with increase in hardness by nearly 1.5 times. The wear and corrosion behavior of the processed alloy was evaluated in simulated body fluid. The processed sample demonstrated remarkable improvement in both wear and corrosion resistance which is explained by surface strengthening and formation of highly stable passive layer. Methylthiazol tetrazolium assay demonstrated that processed 1 ACS Paragon Plus Environment


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sample is better in supporting cell attachment, proliferation with minimal toxicity and hemolysis. The athrombogenic characteristic of as-received and processed samples was evaluated by fibrinogen adsorption and platelet adhesion via Enzyme Linked Immuno-Sorbent Assay (ELISA) and Lactate Dehydrogenase Assay (LDH) respectively. The processed sample showed less platelet and fibrinogen adhesion compared to the as-received alloy, signifying its high thromboresistance. The current study suggests friction stir processing to be a versatile toolbox for enhancing the performance and reliability of currently used bio-implant materials. Key-words: Friction stir processing; Wear resistance; Passivation; Cell-substrate interaction; Platelet adhesion

1. Introduction Metallic biomaterials comprising austenitic stainless steel, Ti alloys and Cr-Co-Mo alloys are widely used in various medical implants such as stents, dental roots, heart valves and joints 14

. These implants play a pivotal role in replacement, repair or augmentation of lost or diseased

parts of the musculoskeletal system. In addition, implants help to restore the form and function of diseased parts for long periods with minimal failure and without showing any apoptotic or necrotic behavior. Stainless steel, SS316L, is one of the most commonly used implant materials because of its low cost, biocompatibility, ease of fabrication, and excellent mechanical properties5-6. However, SS316L shows significant wear and localized corrosion during long term use in the human body which results in accumulation of metal debris in soft tissue as well as sudden catastrophic failure of the implant itself. The cellular environment including cell-implant interface is influenced by implant characteristics such as grain size, surface topography and residual stress

2 ACS Paragon Plus Environment

7-9

. This modulates the

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biological response as well as interaction with adjacent musculoskeletal tissues at both macroscopic and microscopic length-scales. Adherence of bio-implant with blood components (such as platelets, plasma proteins, leukocytes etc.,) is also driven by its surface characteristics. It is critical to understand cell adherence behavior of an implant as it results in severe complications such as thrombosis, vasculitis, and inflammation10-11. Platelets being one of the vital components of the blood gets activated when it interacts with any foreign material, such as an implant, and triggers several tissue & plasma factors. These factors along with several clotting factors activate the formation of thrombin fibers. The thrombin further converts fibrinogen to fibrin strands, signaling the plasma protein to interact with additional clotting factors resulting in blood clots and arterial blockage. Hence, higher cellular and metabolic viabilities along with reduced platelet adhesion can assist the implant materials to serve their purpose better by increasing the service life and reducing the number of revision surgeries. There are numerous studies on reducing the degradation of implant materials and modulating the cellular response by tailoring their surface properties

12-16

. A significant fraction

of these studies reports on the development of coatings 5, 17-18 rather than surface modification of the parent material itself. Some of the limiting factors for coatings include poor adhesion, dilution effects from the substrate, porosity, residual stresses and inferior interface properties leading to delamination and cracking. Surface modification of the parent bio-implant material without introducing any compositional change and/or interface is promising in that regard. Severe plastic deformation (SPD) processes are widely used for surface modification through microstructural refinement

9, 19-21

. SPD processes not only produce refined grain structure but

may also have significant influence on the corrosion and wear behavior as well as cell proliferation phenomena. Few previous studies have reported the influence of severe plastic



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deformation on metallic biomaterials

9, 19-20

. However, these SPD processes demand multiple

steps to achieve appreciable grain refinement in addition to expensive experimental setup. Therefore, a simplistic approach to produce tailored microstructure is indispensable to ensure their realistic application. High strain-rate deformation processes such as friction stir processing (FSP) can produce highly refined microstructures in a single step

22-25

. FSP provides great flexibility in-terms of

processing conditions to tailor the localized microstructure in a material. While there are many studies on FSP of light-weight materials materials such as SS316L

30-31

26-29

, few have reported on FSP of high strength

. Further, a comprehensive study on biocompatibility, cellular

response and corrosion and wear behavior of friction stir processed stainless steel in simulated body environment have not been reported. In the current work, the surface properties of SS316L stainless steel were tailored using a novel submerged friction stir processing route. The degradation behavior of as-received and processed stainless steel was evaluated in wear, corrosion and tribo-corrosion in simulated body fluid. The processed sample showed significantly higher wear resistance even under extreme contact pressures as well as a remarkable tendency to preclude localized pits. Additionally, improved cell viability along with less cytotoxicity and hemo-compatibility has been noted in friction stir processed specimens. The processed samples demonstrated significant increase in resistance to protein adsorption and platelet adhesion. Exceptionally high resistance to degradation and favorable cellular response was explained by structural refinement, enhanced hydrophilicity, and decrease in surface rest potential obtained using high strain-rate deformation.

2. Experimental details


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2.1 High strain-rate processing details The material used in the current investigation was austenitic stainless steel, SS316L, commonly used for biomedical applications. Friction stir processing (FSP) was performed to tailor the surface properties using a pin-less tool made of tungsten carbide with 16 mm shoulder diameter. FSP was performed on vertical milling machine at a constant tool rotational speed of 1800 rpm with two different conditions: (1) conventional processing under ambient cooling condition, where the sample during FSP was cooled under ambient conditions (sample referred to as 1800A) and (2) novel approach of submerged cooling condition, where the sample was completely submerged in a liquid coolant (mixture of ethanol and distilled water) bath maintained at 0 ̊C (sample referred as 1800C). A special purpose FSP fixture was fabricated for holding the sample while submerged in the liquid bath. The FSP fixture was connected to an external chiller through inlet and outlet ports for constant flow of liquid at approximately 100 ml per minute. A schematic of the FSP process is shown in Figure 1. A constant tool transverse speed of 20 mm/min and plunge depth of 0.3 mm was used during FSP. Further, no tilt was given to the FSP tool32. Following processing, all samples were polished down to 3000 grit followed by electro-polishing in 10% oxalic acid solution at 650 mV for 2 min. The grain size and phase distribution for the processed and unprocessed samples were obtained using electron back scatter diffraction (EBSD). EBSD analysis was conducted using FEI Quanta 3D FEG using step size of 0.1 µm. The optical images of the processed samples were acquired using Leica DM750M metallurgical microscope (Germany). Hardness of processed samples was evaluated using micro-hardness testing along the cross-section, while elastic modulus was obtained using nanoindentation. Surface roughness for the as-received and processed samples was obtained using atomic force microscope (AFM, WiTech 300A, Germany).



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2.2 Corrosion testing Corrosion behaviour of as-received and processed samples was investigated by potentiodynamic cyclic polarization (CP) test (using Gamry, Interface 1000E, USA electrochemical setup). A standard three electrode cell configuration was used with high density graphite rod as counter, saturated calomel electrode (SCE) as reference and sample as working electrode. Commercially available simulated body fluid (SBF), ringer solution, was used as the electrolyte. The tests were conducted after the system attained stable open circuit potential (EOCP). Cyclic polarization was done in the voltage range of -0.4 V to +1.4 V vs EOCP with forward and reverse scan rate of 0.166 mV/s as per ASTM standard G61-86. The pit density and morphology following cyclic polarization was observed using 3D optical surface profilometer. Electrochemical impedance spectroscopy (EIS) measurements were obtained at EOCP over a frequency range of 0.01 Hz to 100 kHz with a set AC voltage amplitude of 10 mV at different immersion time of 0 hrs, 12 hrs and 24 hrs for all the specimens. The Gamry Echem analyst 7.03 was used to model the electrical equivalent circuit (EEC) and to fit the Nyquist and Bode plots using simplex algorithm.

2.3 Wear testing Tribological performance was evaluated using Rtec Universal Tribometer (Rtec Instruments, San Jose, CA, USA). A 3 mm WC ball was chosen as counterface. Tests were conducted at 3 N and 5 N normal loads that correspond to hertz mean contact stress of 1.2 GPa and 1.4 GPa respectively. These stress values are nearly 10-12 times higher than those predicted for common bio-implants under static loading conditions, for instance the knee joint


33

. These

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test conditions were chosen to simulate the dynamic loading for an implant, where the stress state can be multiple times higher compared to static loading. Sliding tests were conducted with a 1 mm stroke length, 5 Hz reciprocating frequency for duration of 30 min. A fresh counterface surface was used for each test iteration. Optical interferometry was used for generating 3D profiler images. Gwydion SPM image analysis software was used for quantitative evaluation of wear volume loss from the profiler data. Similar analysis was performed on samples tested in simulated body fluid. Samples after corrosion and wear were studied for the respective degradation mechanism using FEI Quanta environmental scanning electron microscope with inbuilt energy dispersive x-ray spectroscopy. Elemental maps of the surface oxide and wear particle compositions were used for fundamental understanding of the tribological behavior.

2.4 Cell assays Polished as-received and processed samples were rinsed in ultrasonicator for 15 min with distilled water, acetone and ethanol respectively. Later, they were dried in an oven at 65 ˚C for 2 hr for carrying out following assays:

2.4.1

Cell culture

Primary MDCK-1 cells and HepG2 were obtained from National Centre for Cell Science, Pune, India. MDCK-1 cells were derived from the kidney tissue of an adult female cocker spaniel and HepG2 cells were derived from human liver carcinoma cells. These were grown in Dulbecco’s modified Eagle’s minimal essential medium (DMEM) as media in the presence of 10% (v/v) Fetal bovine serum (FBS) and penicillin-streptomycin (Gibco, Invitrogen, ThermoFisher Scientific, NY, USA) at 37ºC, humidified, 5% CO2.



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2.4.2 Cell viability MDCK-1 and HepG2 cells were seeded in cell culture treated 6 well polystyrene plates 4

separately (Nunc, Thermo Fisher Scientific NY, USA) with concentration of 3x10 cells/ml and kept in incubator (37°C and 5% CO2) for 24 hr. The cells were then washed with DMEM. 1ml of fresh media was added into each of the six wells along with as-received and both the processed samples and again maintained in incubator for 24 h. After incubation, the samples were removed and media was aspirated. Further, the cells were trypsinized with diluted with tryphan blue (1:1) dilution and counted under Nikon Ti2 eclipse microscope. Non-viable cells stained blue while viable cells remained unstained.

2.4.3 Cytotoxicity For assessing the cellular metabolic activity and toxicity, a colorimetric assay, MTT (3(4, 5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide, Thermofisher scientific) was performed. 2x104 of each MDCK & HepG2 cells were seeded in 96 wells plate. Cells were grown in the incubator (37°C and 5% CO2) for 24 h & washed twice with DMEM media. 100 µl of fresh media was added into each well along with as-received and both the processed samples. Following incubation at 37°C and 5% CO2 for 24 h, the samples were removed and 10 µl of MTT with concentration of 5 mg/ml was added in each well. After 3 h of incubation, 100 µl of Dimethyl sulfoxide (Sigma Aldrich, USA) was added to each well. Absorbance was measured by a Multi-Mode Plate Reader (Bio-Rad) at 595 nm. Percentage of cytotoxicity was calculated by following expression:


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% =

. − . . 100

2.4.4 Actin filaments staining MDCK monolayers used for fluorescent staining were grown on 18 mm cell culture treated poly-l-lysine coverslips (Neuvitro, USA). 2x104 cells were seeded on the coverslip and incubated. Cells were kept in contact with as-received and processed samples for 24 h. Upon sample removal, cells were washed twice in cold PBS at 4˚C and fixed with 2% paraformaldehyde in PBS for 5 min at room temperature. Coverslips were blocked with 2% freshly prepared Bovine serine albumin (BSA) for 1 h & labelled with polyclonal rabbit antiactin antibody (1:250) (Zymed Laboratories, USA). Cells were incubated overnight at 4˚C and washed three times with 0.1M PBS, and incubated with the appropriate fluorescein labelled secondary antibody [FITC-goat-anti-rabbit (1:250), Santa Cruz Biotechnology, USA] for actin filaments staining. Cells were incubated for 1 h at 37 ˚C and washed three times with cold 0.1M PBS. Finally, antifade-DAPI was used as mounting agent for staining nucleus. The coverslips were inverted on glass slides and cell imaging was done with Nikon Ti2 eclipse fluorescent microscope to check morphological alterations due to cytotoxicity.

2.4.5 Hemolysis assay Hemolysis assay was performed as described by Lale et al

34

. 3 ml of human blood was

centrifuged at 1500 rpm for 10 min. Red blood corpuscles (RBC) pellet was washed thrice with 0.01M PBS and diluted to 6 ml to prepare RBC stock solution. As-received and processed SS316L samples were sterilized and incubated with RBC stock solution (300 µl each) at 37 °C for 3 h. Post incubation, supernatant was collected and centrifuged at 2500 rpm for 5 min. 100 µl 9 ACS Paragon Plus Environment


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of hemoglobin released in the supernatant was analyzed at 540 nm using UV−visible spectrophotometer (Bio-RAD 680, USA). Percent hemolysis was estimated with respect to hemolysis caused by negative control (PBS) and positive control (De-ionized water), as given by the following equation: % =

!" #$% &' − ( #$% &' * 100 ) #$% &' − ( #$% &'

2.4.6 Fibrinogen adsorption To determine fibrinogen adsorption, as-received and processed samples were sterilized and incubated overnight with platelet poor plasma (PPP). The fibrinogen in PPP was allowed to coat on the metal surface overnight at 4°C. Polystyrene 24 well plates coated with fibrinogen were used as positive control while well without PPP was used as reference control. The fibrinogen binding potential for the samples was determined by ELISA (Enzyme Linked Immuno-Sorbent Assay). Plasma pre-adsorbed metal samples were rinsed (3x5min) with 0.01 M phosphate buffered saline, PBST, (0.05% tween-20) blocked for 1 h (5% skimmed milk in 0.01M PBS) to eliminate non-specific protein binding. Rinsed samples were then incubated in primary antibody for 1 h at room temperature (anti-FGA produced in rabbit 1:1000). Samples were again rinsed with 0.01M PBST (3x5min) and incubated in secondary antibody for 1 h (antirabbit IgG HRP conjugated 1:10000), rinsed with 0.01M PBST (3x5min). Finally, the sample surfaces were incubated with substrate (3, 3’, 5, 5’-Tetramethyl benzidine) (O.1M TMB) for 30 min. at 37°C in dark. Solutions were transferred to 96 wells plate and absorbance values were read at 450 nm.


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2.4.7 Platelet isolation Platelet isolation was performed as explained by Watson35-36with few alterations. Human whole blood was centrifuged at 200 g (without brake) for 20 min to remove erythrocytes. Plasma rich platelet (PRP) in the supernatant was centrifuged at 200 g (without brake) for 10 min. Remnant erythrocytes were precluded and PRP was centrifuged at 2000 g for 10 min. Platelets obtained in the pellet were re-suspended in Tyrode’s buffer (freshly prepared 134 mM NaCl, 12mM NaHCO3, 2.9 mM KCl, 0.34 mM Na2HPO4, 1 mM MgCl2, 10 mM HEPES (4-(2hydroxyethyl)-1-piperazineethanesulfonic acid), 5 mM Glucose, set pH 7.4, freshly added BSA 3 mg/ml) and count was estimated in hemocytometer.

2.4.8 Platelet adhesion assay Platelet density was approximated at different platelet concentrations (0-200x103)37. Briefly, 100 µl of each concentration were prepared in Tyrode’s buffer as diluent. 70 µl of each concentration was treated with 2% (v/v) Triton X-100 in 0.1M PBS buffer for 30 min at 37°C. 50 µl of the lysed platelet solution was incubated with 50 µl of LDH substrate (CytoTox 96® NonRadioactive Cytotoxicity Assay from Promega Corp., WI, USA) for 30 min at 37°C. Reaction was stopped with 50 µl stop solution. The absorbance was read at 490 nm and different platelet concentrations (0-200x103) were used to plot the standard curve. For evaluating platelet adhesion property, the as-received, and both processed samples were incubated with platelet concentration (100 x103) for 1 h at 37°C. Post hour incubation buffer was removed and adhered platelets were treated with 0.1M PBS containing 1% (v/v) Triton X-100 for 30 min at 37°C. 50 µl of the lysed platelet solution was incubated with 50 µl of LDH substrate for 30 min at 37°C. Reaction was stopped with 50 µl of stop solution and the absorbance was read at 490 nm.



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2.4.9 SEM protocol Platelet adhesion on 1800C specimen was examined by scanning electron microscope (SEM) as reported38 with slight modification obtained from Fischer et.al

39

. Briefly, platelets

isolated from human citrated blood were co-incubated with the specimen for 1 h at 37 °C, 5% CO2. Platelets incubated on coverslips were used as control. Samples were then washed with 0.1M Phosphate buffer (pH 7.4) and fixed in 2.5% glutaraldehyde in 0.1 M PB at 4 °C for 2-3 h. Then, samples were post fixated in 1% osmium tetroxide for 1 h and dehydrated by gradient acetone concentration (50-100%), 20 min each. Thereafter, samples were treated with 100% HMDS at room temperature for 5 min and mounted on aluminum stubs with adhesive carbon tape. Prior to SEM application, a thin gold layer was coated by means of a sputter-coater (SC7640, Polaron Equipment, England, UK). The samples were observed under an environmental, variable pressure Scanning Electron Microscope (Carl Zeiss EV0-40, Cambridge, UK) at a voltage of 20 kV and a working distance of 10 mm. The adhered platelets were characterized based on their morphologies as described previously40.

2.5 Statistical analysis All the experiments were performed in triplicates and the standard deviation is reported with error bars. The analysis of variance (ANOVA) method followed by a student’s - t test has been performed to determine the statistical significance. Differences were considered significant at p < 0.05.

3. Results


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3.1 Microstructure evolution- EBSD analysis The optical images of the cross-section for both the processed specimens is shown in Figure 2. The processed region comprises a highly refined zone (nugget) of nearly 80-100 µm thickness and thermomechanical affected zone (TMAZ) zone of nearly 70-80 µm. Thus, the total thickness of the processed region is nearly 170-180 µm for both the processed specimens. The grains in TMAZ are coarser and elongated which is attributed to the shearing action of the FSP tool. In contrast to 1800A, the depth of nugget zone is slightly larger while TMAZ is slightly smaller for 1800C specimen which is attributed to lower peak temperature and rapid cooling for 1800C specimen. Heat affected zone (HAZ) is not evident in the processed region which is likely due to lower heat generation during FSP with a pin-less tool. The EBSD maps from the nugget region of the top surface showing the grain structure and grain size distribution for the asreceived alloy and both the processed specimens are given in Figure 3. The EBSD map for asreceived alloy (Figure 3(a)) shows the presence of twin boundaries, which is attributed to low stacking fault energy (SFE) for austenitic steels. The average grain size for as-received stainless steel, as calculated from the grain size distribution, is nearly 22 µm (Figure 3(b)). Figure 3(c) to Figure 3(f) show the microstructural details for the processed samples. Both the processed samples showed significant grain refinement. The average grain size for 1800C sample was 0.8 µm (Figure 3(e)), while for 1800A sample was found to be nearly 2.5 µm (Figure 3(c)). Further, 1800C sample showed significant fraction of extremely fine equiaxed grains on the order of few nanometers (Figure 3(f)). The strain rate during FSP was estimated to be nearly 700 s-1 using the equation +́ = -' 2/ ⁄

41

where -' is the material flow rate, taken as half of the tool rotational

speed, is the radius of recrystallized zone and is the depth of recrystallized zone which is assumed to be equal to plunge depth for a pin-less FSP tool used in the current study. Further, the



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sample processed under cooling condition, 1800C, show a more significant log normal distribution (Figure 3(f)) compared to that processed under ambient condition, 1800A (Figure 3(d)).

3.2 Mechanical properties Figure 4(a) shows the variation in microhardness for both the processed samples as a function of depth from the top surface. For both the processed samples, top surface shows the maximum hardness and it decreases along the sample depth. Maximum hardness for 1800A sample is nearly 1.3 times the hardness of the unprocessed steel, while it is nearly 1.5 times for the 1800C sample. The hardness of both the processed samples decreases to the value for asreceived alloy at about 175 µm, which corresponds to the depth of the processed zone. The elastic modulus for all the samples was obtained using nano-indentation. The variation of peak hardness and elastic modulus for the as-received alloy and both the processed specimens are shown in Figure 4(b). It is seen that hardness significantly increased after processing, while elastic modulus decreased slightly. X-ray diffraction analysis (Figure S1 (a) of supporting information) indicate no secondary phase formation during processing. In contrast, EBSD phase map, shown in Figure S1 (b), indicates formation of small fraction (~3% for 1800A and ~8 % for 1800C) of martensite phase in the processed specimens32. Therefore, the significant rise in hardness after processing can be primarily attributed to grain boundary strengthening in accordance with the well-known Hall-Petch relation. In contrast, the change in elastic modulus post-processing can be understood by texture changes during friction stir processing. Figure 4(c) shows normal direction (ND) inverse pole figures for the as-received as well as both the processed samples. The as-received alloy showed a strong (111) parallel ND texture while


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1800A and 1800C showed strong (100) parallel ND and (211) parallel ND texture respectively. Face-centred-cubic materials show higher elastic modulus along (111) plane compared to other crystallographic planes due to the closest atomic packing for (111)42. In the AFM measurements, all samples showed similar roughness values: 2.16 ± 0.25 nm for the as-received alloy, 2.01 ± 0.19 nm for 1800A and 1.99 ± 0.2 nm for 1800C. This indicates that surface roughness is not significantly influenced by grain size. Similar results have been reported in many previous studies as well9, 43

3.3 Corrosion behavior 3.3.1

Tafel and cyclic polarization studies

The cyclic polarization behavior comprising of cathodic and anodic polarization curves for the as-received and both the processed specimens are shown in Figure 5. The corrosion current (Icorr) was evaluated using Tafel exploration from the polarization curves and values are shown in Table 1. 1800C showed lowest Icorr value (Figure 5(c)) followed by 1800A (Figure 5(b)) and the as-received alloy (Figure 5(a)). Figure 5(d) shows the corrosion rate dependence on grain-size for as-received and processed samples. Interestingly, corrosion rate shows a linear relationship with d-1/2 (d is the grain size), similar to Hall-Petch equation relating hardness/strength with grain size. Such a relationship has been demonstrated for a wide range of materials44. The relation signifies higher corrosion rate for fine grained austenitic steel in simulated body fluid condition. The hypothesis of higher chemical activity of grain boundaries implies that corrosion current should depend on the grain boundary length (gl). Ralston et al

44

proposed a generalized relation for grain boundary length in terms of grain size: = + 234 where A is the area term, B is the scale term, gs is the grain size and α lies between 0.5



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and 1. For random grain structures, value of α is 0.5, which supports the linear dependence of corrosion current on 35/7 . Localized pitting behavior of all the samples was obtained by evaluating the pitting potential (Epit) and protection potential (Epp) from cyclic polarization test (Figures 5(a) to 5(c)). Epit-Ecorr and Epp-Ecorr are two important parameters that describe the pitting behavior of a material. While Epit-Ecorr is a measure of resistance to pit initiation, Epp-Ecorr quantifies the resistance to pit propagation. The values of both these parameters for all specimens are given in Table 1. 1800C specimen showed highest resistance to pit initiation as well as pit propagation followed by 1800A and the as-received alloy. Further, the severity of pitting corrosion relates to the loop area above Epp which is again largest for the as-received alloy and least for 1800C specimen (Table 1). Thus, both the processed specimens show stronger passivation compared to the as-received alloy. The results of EIS testing are shown as Nyquist and Bode plot in Figure 6.

The

impedance spectra are significantly higher for both the processed samples (Figure 6(c) and 6(e)) compared to the as-received alloy (Figure 6(a)) indicating higher corrosion resistance of the former. Further, the impedance increases with the immersion time for all the specimens. The Bode plot for the as-received alloy suggest existence of two time constants while both the processed specimen show a single time constant. Therefore, a two-time-constant EEC for asreceived alloy and single-time-constant EEC for both the processed specimens was used to model the current electrochemical system. The EEC are shown as an inset in Figure 6. This electrical circuit has been previously shown to give a best fit for austenitic stainless steels45. Here, Rs is the solution or electrolyte resistance, CPE1 and R1 are the electrochemical double layer capacitance and charge transfer resistance respectively for the high frequency part of the


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spectrum, while CPE2 and R2 are the corresponding elements for the low frequency spectrum. CPE, a frequency dependent constant phase element with exponent n is used instead of pure capacitance to account for surface in-homogeneities such as roughness, adsorption or diffusion. The values of different elements of the fitted electrochemical equivalent circuit for different immersion times is shown in Table 2. The total resistance, R1 for 1800C and 1800A and summation of R1 and R2 for the as-received alloy, is highest for 1800C, followed by 1800A and the as-received alloy. This is exactly in line with Tafel and cyclic polarization results supporting higher corrosion resistance for the processed samples. In addition, the total polarization resistance increases with the immersion time while capacitance decrease for all the specimens indicating growth of the passive oxide layer. Further, the n value for both the processed samples is higher compared to the as-received alloy indicating better homogeneity of the passive layer for the former. An SEM image of the specimen before the corrosion test is shown in Figure S2 (supporting information) while SEM images after the corrosion test are shown in Figure 7. Figure 7(a) and 7(b) show low magnification SEM image for the as-received and 1800C specimen after the corrosion test. As-received alloy show extensive pitting all over the surface while the density and severity of pits is significantly lower for 1800C specimen. This agrees well with the higher pit-initiation resistance, high polarization resistance and smaller loop area of the latter. 1800A showed similar features as 1800C, therefore not shown here. High magnification image of a pit for the as-received alloy and 1800C is shown in Figure 7(c) and 7(d) respectively. Figure 7(e) and 7(f) show 3D profiler image of the pits for the as-received and 1800C specimen respectively. Depth profiles for pits from as-received alloy and 1800C specimen are shown in Figure 7(g) and Figure 7(h) respectively. Pit depth for 1800C is nearly one tenth of the as-



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received alloy which is attributed to its higher pit-propagation resistance. Such an unusual decrease in pitting may be attributed to the formation of highly stable oxide layer.

3.4 Wear behavior Wear volume loss (WVL) in dry test as well as in ringer’s solution for two different normal loads are shown in Figure 8(a) and 8(b) respectively. WVL is larger at higher loads, irrespective of the test condition and the specimen type. The WVL is significantly lower for the processed specimens under all test conditions. 1800C sample showed highest wear resistance with at-least 40% lower wear loss in dry and 50% lower in ringer solution. For the same load, the overall material loss is lower in ringer’s solution which is likely due to the formation of passive layer and lubricating effect of the solution. The depth profile of the wear tracks for all samples was obtained as shown in Figures 8(c) and 8(d) for the as-received and processed specimens at 5N load under dry test and in ringer solution respectively. In-line with the wear volume loss, wear depth was lowest for 1800C followed by 1800A and the as-received alloy. In contrast to dry wear, depth of the wear track for 1800C is unusually smaller in ringer solution implying its larger wet wear resistance (Figure 8(d)). Figures 8(e) and 8(f) show the correlation of WVL in dry wear with hardness (H) and hardness to modulus ratio (H/E). WVL scales linearly with both H and H/E and shows a strong correlation, but not with modulus (E, not shown here). Representative scanning electron micrographs of the wear tracks for 1800C steel after sliding reciprocating condition are shown in Figure 9. Low magnification SEM images of the wear tracks at 5 N and 3 N are shown in Figure 9(a) and 9(b) respectively. The corresponding high magnification images highlighted are shown in Figures 9(c) and 9(d). The wear track at 5 N is seen to undergo severe oxidation and subsequent delamination of top layers along with signs of micro-scuffing. The 3 N wear track (Figure 9(d)) does not show extensive oxidation and 18 ACS Paragon Plus Environment

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delamination. Material removal mechanism is dominated by long abrasion seen as scuff marks in the wear track. The 5 N sample shows plastic deformation and smearing of material along the periphery of the wear track. Such extensive plastic deformation is not seen on the 3 N wear track. To quantify the nature of wear tracks, EDS mapping was performed and spectrum for oxygen is shown in Figure 9(e) for 5 N and in Figure 9(f) for 3 N test condition. The EDS maps show uniform oxide formation in the wear tracks and the wear debris to be oxides of iron and chromium, formed due to frictional heating. The wear morphology of 1800A sample was similar to that of 1800C. In contrast to the dry wear samples, the samples tested with ringer solution had significantly lower wear volume loss as seen from the sizes of the wear tracks in Figure 9(g-l). Representative wear tracks on 1800C sample at 5 N and 3 N load conditions in SBF are shown in Figures 9(g) and 9(h) respectively. The wear track is shallow and devoid of deformation signs such as micro-cutting and extensive plastic deformation. Faint micro-scuffing signs are seen for the 5 N sample at high magnification (Figure 9(i)), which are minimal for the 3 N sample (Figure 9(j)). Element specific EDS mapping shows the surface to have oxides of calcium and phosphate (Figure 9(k) and 9(l)), precipitating from the ringer electrolyte. Besides these, no other metal oxides were observed. This is likely due to the lubricating effect from the ringer electrolyte, which prevented temperature rise and metal oxide formation.

3.5 Biocompatibility and cellular response 3.5.1

Cellular viability and cytotoxicity

Several mitochondrial enzymes are analyzed to determine the viability of cells. NAD(P)H-dependent cellular oxidoreductase is one such enzyme which is detected by MTT assay. MTT assay is based on enzymatic release of NAD(P)H-dependent cellular oxidoreductase



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from live cells, reducing the tetrazolium dye MTT to its insoluble formazan crystal of purple colour. Using optical density value, the cellular toxicity under defined conditions can be evaluated which signifies the presence of number of viable cells and dead cells simultaneously. The viable cell count on MDCK line, shown in Figure 10(a), reveal minor variation among the samples when trypan blue exclusion assay was performed. Similar results are shown for HepG2 cell line (Figure S3 (a), supporting information). 1800C specimen showed better viability as the count of living cells after 24 hrs is higher. Figure 10(b) shows the results of MTT-cytotoxicity assay which clearly indicates that both as-received and processed samples are nontoxic to the cells. The result for HepG2 cell line were found to be similar, (Figure S3 (b), supporting information). However, 1800C shows higher cellular optical density close to the untreated sample. It is seen that cytotoxicity for 1800C is nearly 4 times lesser compared to the as-received alloy. 1800A also performed relatively better compared to the as-received steel. The statistical significance (p*< 0.05) in variation w.r.t positive control is reported in 1800A and as-received. Thus, 1800C remains closer to the positive control in cell viability assay indicating its better cellular response.

3.5.2

Morphological aberrations of the cells

MDCK-I cells were cultured and stained against actin and DAPI to initially study the cell cytoskeleton structure and the morphological aberration, when kept in contact with treated samples. In comparison to the increase in elongation and spreading of MDCK-I cells on the control sample (Figure 10(c)), cells on as-received (Figure 10(d)) appear to acquire a predominantly round shape with decreased fluoresce of actin cytoskeletal staining. This is not the case in both the FSP processed samples (Figure 10(e) & (f)). The cells on 1800C sample has


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proved to sustain the normal elongated and segregated morphology (as control cells) compared to slightly strained and clustered morphology in 1800A.

3.5.3

Hemo-compatibility

Figure 11(a) depicts the potential of all tested specimens in direct blood contact applications. It is seen that both as-received and processed samples showed minimal hemolysis after treatment with human RBC. While total hemolysis was observed in the positive control, 1800A > 1800C. The corrosion behavior of a material is a complex function of number of interdependent factors such as alloy composition, surface topography, type of electrolyte, electrolyte pH etc. Therefore, corrosion behavior cannot be explained solely based on a single parameter. However, Ralston et al 44 showed a direct correlation between corrosion behavior and the grain size. Based on the corrosion behavior of the as-received alloy with coarse grain structure, these authors classified materials broadly into two categories: (1) those that show passive corrosion behavior (Icorr ≤ 10 µA. cm-2) and (2) those that show active corrosion behavior (Icorr ≥ 10 µA. cm-2). Grain refinement in passive materials results in further reducing the corrosion rate while it leads to increase in corrosion for active materials44. In the current study, the corrosion rate for the as-received stainless steel was found to be 10 µA. cm-2, which as per



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the above classification suggests its passive nature in simulated body fluid. Further, grain refinement for this alloy resulted in decrease in the corrosion rate, in concurrence with observations by Ralston et al

44

. Thus, grain refinement exaggerates the corrosion behavior for

similar alloy chemistry in a given electrolyte. Therefore, lower corrosion rate for processed samples is attributed to faster passivation kinetics due to high grain boundary density. In addition, finer grain structure is known to promote stability of the passive oxide layer. This is attributed to faster Cr diffusion along the grain boundaries, the density of which is significantly higher for processed samples. Cr is the primary element of the protective oxide layer in stainless steel 9. In addition, grain boundaries tend to anchor the protective passive oxide layer by the oxide pegging effect

47

. Therefore, faster Cr diffusion and pegging of oxide layer for refined

grain structure in processed samples likely provides more stability towards pitting by electrochemical dissolution. This effect is manifested by higher pitting potential, larger resistance to pit initiation and propagation and lower area under the cyclic polarization curve for the processed specimens. Localized pitting is one of the major reasons for failure of stainless steel implants whereby corrosion pits serves as the initiation point for the failure due to corrosion fatigue. Therefore, higher pitting resistance shown by the processed specimens is believed to be of particular significance in context of implant applications. In addition to localized corrosion, wear debris spawned during relative motion between different moving parts and joints, acetabular cup and femoral head of hip joint for instance, can lead to sudden premature implant failure. Wear rate depends on number of factors such as peak load and contact stresses, loading rate and loading time. The intensity of contact stresses and the corresponding wear rate can be extremely high under dynamic loading conditions. In general, stainless steels have lower specific strength and have relatively poor wear properties compared to


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other commonly used implant materials such as Ti-alloys and Co-Cr alloys

48

. Therefore, it is

critical to enhance its wear resistance for efficient utilization in implant applications. In the simplistic Archard model

49

, the wear resistance scales linearly with the material’s hardness.

However, some earlier studies have shown the dependence of wear resistance on both hardness and elastic modulus

50-51

. In the current study, the WVL scales linearly with both hardness as

well as with hardness to elastic modulus ratio (H/E) (Figure 8(e) and 8(f))). While hardness signifies resistance to abrasion, hardness to modulus ratio quantifies the extent of plastic deformation. Plasticity index, A, defined as A = BC 35/7 = D ⁄EF⁄G

51

combines the

material and topographic properties of surfaces in contact to explain their deformation behavior. Here, BC represent the critical value of elastic displacement for plastic flow, E is the elastic modulus, H is the hardness, and σ and β are variables defining topography of contact surfaces5253

. The critical load required for transition from elastic to plastic deformation is a function of

hardness as well as elastic modulus. A material with low H/E ratio shows transition from elastic to plastic deformation at lower stress levels compared to those with high H/E values. In other words, deformation under contact for a high H/E ratio material is mostly elastic. In the current study, both the processed samples have higher H/E compared to the as-received alloy with highest value for 1800C specimen. Thus, the processed specimens are likely to undergo lower plastic deformation and therefore lower wear rate. The dominant wear mechanisms observed in the current study include abrasion and plastic deformation, as discussed in the preceding section. Abrasion behavior is typically controlled by hardness, while hardness to modulus ratio dictates plastic deformation. This explains the linear dependence of WVL on H and H/E ratio. Further, the formation of oxide layer, as shown by the EDS analysis (Figure 9(e) and (f)), prevents metal to metal contact and reduces the wear loss. Thus, the remarkable improvement in the wear



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resistance after processing (nearly 2 times higher) can be explained based on microstructural refinement and its effect on mechanical properties. Cyto-compatibility is of primary importance for implant materials. In the current study, cytotoxicity was determined using primary MDCK and HepG2 cell lines and all samples provided appropriate response to in vitro viability and cytotoxicity. However, 1800C demonstrated better response compared to the as-received alloy. This is likely attributed to highly stable passive oxide layer for 1800C specimen. The leaching out of metal ions from the implant surface during reaction with the biological fluid and corrosion can result in toxicity. In addition, the released ions can be absorbed by macrophages that facilitate the release of cytokines causing inflammatory reactions and implant failure 54. As shown in preceding section, 1800C showed higher pitting potential, higher impedance and pit propagation resistance. Stable passive layer limits leaching out of metal ions and ensures better cyto-compatibility for 1800C specimen. Also, an increase in cell proliferation by grain refinement has been reported previously

20, 55-56

. In addition, there are evidences of cellular elongation, increased viability and

proliferation on the cells grown on 1800C compared to that of 1800A where roundness and clustered morphology is prominent (Figure 10). Surface topography, roughness and wettability have been reported to significantly influence the cell-substrate interactions including cellular morphology and cell proliferation9, 19, 57. In the current study, surface roughness was found to be similar for all the specimens and therefore cannot explain differences in cell morphology. To determine the influence of surface wettability, water contact angle for all samples was measured and was found to be 60˚ ± 3˚ for the as-received alloy, 55˚ ± 2˚ for 1800A and 46˚ ± 2˚ for 1800C. Thus, the processed samples were found to be more hydrophilic (polar) compared to the as-received alloy. This is attributed to higher surface free energy of processed samples due to


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their refined grain structure. Increased hydrophilicity with grain refinement has been demonstrated in few previous studies as well

57-59

. Hydrophilic surfaces have been reported to

promote cell spreading and reduce interfacial reactions60. Improved hydrophilicity enables the cells to release the proteins & extend their actin, myosin cytoskeletal fibers resulting in improved cell migration & spreading. Thus, 1800C being more hydrophilic shows elongated and segregated morphology with extended filopodia. On contrary, the as-received alloy and 1800A being less hydrophilic promotes cellular adhesion but substantially lesser cellular spreading rendering clustered morphology. Besides cyto-compatibility, biomaterials for blood contacting applications should be validated for their hemo-compatibility. While total hemolysis was observed in positive control, negligible hemolysis of less than 5% occurred in all the tested samples and can be classified as blood-compatible as per ASTM F756-00. However, a major disadvantage for blood-contacting metal implants has been their tendency to adsorb plasma proteins such as thrombin, fibrinogen etc. Presumably, when a bio-implant is placed in human body, continuous contact with the blood components triggers some of the blood cells such as platelets to adhere on metal surface. Adhered platelets further get activated secreting cytoplasmic components like enzymes, transcription factors, clotting factors etc. Together these factors assist the conversion of inactive plasma proteins such as fibrinogen into active fibrin strands. Adsorption of these fibrins on the metal implant co-initiates platelet aggregation and promotes the coagulation cascade resulting in adverse cellular responses such as vasculitis, thrombosis & activation of immune responses. Subsequent immunological reaction initiates various foreign body cascades attenuating the in vivo application of the implant. Therefore, it is imperative to fabricate the mechanistic property of these bio-implant to restrain cellular & protein interactions thereby eliminating further immunological complications. Hence, the athrombogenic feature of



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the bio-implant was evaluated by protein adsorption and platelet adhesion property via ELISA and LDH assay respectively. Since blood plasma comprises pool of proteins, a sensitive method to detect the presence of abundant fibrinogen is requisite. During adsorption, binding of any protein to a surface could be in variable conformations. Therefore, to increase the sensitivity of the study antibody binding to the specific epitope region of the fibrinogen was used. The antigen antibody reaction was measured based on the absorbance. Thus, difference in the antibody binding w.r.t the protein adsorption suggested processed samples to be more resistant. In LDH assay study, release of Lactate Dehydrogenase (LDH), a cytosolic enzyme is measured upon platelet lysis. Defined concentrations of platelets were studied for their relative LDH enzyme release. Release of enzyme from the platelets adhered on implant samples was estimated w.r.t the defined platelet concentrations. 1800C showed significantly lower fibrinogen adsorption and platelet adhesion compared to the as-received and 1800A alloy. The dynamics of protein and platelet adsorption is a complex process and cannot be related to a single chemical or physical factor. One of the key factors that play a vital role in protein adsorption is the wettability characteristics of the implant material. Protein being amphiphilic, the hydrophilic residue of protein does not take part much in the reaction with the surface, and hydrophobic (non-polar) residue interaction plays a vital role in attracting protein to the metal surface

61

. The protein

adsorption on hydrophobic surface occurs by hydrophobic interactions. This is later followed by Vroman effect resulting in increase in protein adsorption. In contrast, protein adsorption on hydrophilic surface necessitates substitution of adsorbed water molecules. This process offers a significant energy barrier and lowers the probability of protein adsorption on a hydrophilic surface. Thus, lower fibrinogen adsorption and platelet adhesion can be attributed to more hydrophilic nature of 1800C specimen. Lower fibrinogen adsorption on hydrophilic surface has


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been showed in few previous studies as well62-63. However, the protein adsorption behavior is also driven by the type of protein and influence of surface wettability may differ with the nature of protein. Additionally, surface charge has also been reported to influence the protein adsorption on implant surface64. Generally, negatively charged surfaces tend to show more resistance to protein and platelet adsorption 65. As most of the platelets and proteins possess negative charge, there exists an electrostatic repulsion with the negatively charged implant material making it difficult to adhere on implant surface and more thromboresistant. In the current study, FSP processed samples tends to be more electronegative (-663 mV, surface rest potential) compared to as-received (-490mV) (Figure 5). Thus, processed samples were more resistant for protein adsorption, subsequently, reducing the platelet activation as fibrinogen is the key element for platelet adhesion activity

66-68

. Decreased platelet adhesion in turn inhibits platelet activation

impeding protein adsorption and further vascular complications. The results obtained were consistent and proved using SEM micrographs for 1800C. Ideal dendritic morphology of the fully spread activated platelets was absent. The metal surfaces were deprived of any aggregation of platelets. As a result, the key complication of platelet adhesion followed by protein adsorption onto the surface of blood-contacting implant was forestalled. Based on the results of the current study, 1800C appears to be a promising material for biomedical applications with substantially improved overall performance, longer life-span and reliability.

5. Conclusions Surface properties of austenitic stainless steel were significantly enhanced using friction stir processing with different cooling conditions. The average grain size decreased by an order of magnitude from roughly 20 µm to 0.9 µm after processing. The hardness of processed samples



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increased by nearly 1.5 times while elastic modulus decreased slightly. Corrosion behavior of asreceived and processed samples was evaluated in simulated body fluid, while wear properties were investigated both in dry condition and in simulated body fluid. The processed samples showed significantly higher pitting resistance which was attributed to stronger passivation owing to pegging of oxide layer to the refined grain structure. Processed samples also demonstrated substantial increase in wear resistance. The improved tribological behavior of processed samples was attributed to higher hardness, greater resistance to onset of plastic deformation and stronger passivation. Similarly, the processed samples show better cellular response and cell viability. The processed sample showed good hemo-compatibility with hemolysis less than 5%, very low fibrinogen adsorption (~13% compared to 55% for as-received) and platelet adhesion (~7% compared to 43% for as-received). Superior properties and enhanced performance of processed samples was explained by microstructural refinement, wettability, surface charge and texture changes. This study provides fundamental insights into surface engineering of stainless steel to enhance its performance as a bio-implant material.

Acknowledgements H.S. Arora thankfully acknowledge the financial assistance provided by Science and Engineering Research Board (SERB), Department of Science and Technology, under the project titled “Tailoring the Surface Properties of Crystalline and Amorphous Metals for Advanced BioImplants” (File no. YSS/2015/000678). S. Singh thankfully acknowledge the financial support received from DST-PURSE.


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Figure Captions Figure 1. Schematic diagram explaining the novel submerged friction stir processing (FSP). The surface of as-received (SS316L) alloy was modified using experimental setup specially designed for submerged processing. Figure 2. Optical microscope images of the cross section at 100X magnification for (a) SS316L, friction stir processed under submerged cooling (1800C) and (c) SS316L, friction stir processed under ambient cooling (1800A); at 200X magnification for (b) 1800C and (d) 1800A. The processed region comprises a highly refined zone (nugget) of nearly 80-100 µm thickness and thermomechanical affected zone (TMAZ) zone of nearly 70-80 µm with total thickness of the processed region as 170-180 µm. Figure 3. EBSD map for (a) as-received austenitic steel SS316L, (c) friction stir processed SS316L under ambient cooling (1800A) and (e) friction stir processed SS316L under submerged cooling (1800C); Grain size distribution for (b) as-received SS316L, (d) 1800A and (f) 1800C. Figure 4: (a) Hardness as function of depth for the as-received austenitic stainless steel, SS316L, SS316L friction stir processed under ambient cooling (1800A) and SS316L friction stir processed under submerged cooling (1800C); (b) Variation of peak hardness and elastic modulus for the as-received austenitic stainless steel, SS316L, 1800A and 1800C specimens; (c) Normal direction (ND) inverse pole figure (IPF) for as-received, 1800A and 1800C specimens. The color bar shows the texture intensity in each case.



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Figure 5: Cyclic polarization curves for (a) the as-received austenitic stainless steel, SS316L, (b) SS316L friction stir processed under ambient cooling (1800A), (c) SS316L friction stir processed under submerged cooling (1800C), The pitting potential (Epit), corrosion potential (Ecorr), and protection potential (Epp) for all specimens are marked on the polarization curves, (d) Correlation between corrosion current and grain size for as-received SS316L, 1800A and 1800C. Figure 6: Electrochemical impedance measurements showing Nyquist plot at different immersion times of 0 hrs, 12 hrs, and 24 hrs for (a) the as-received austenitic stainless steel, SS316L, (c) SS316L friction stir processed under ambient cooling (1800A), (e) SS316L friction stir processed under submerged cooling (1800C); Bode plot at different immersion times of 0 hrs, 12 hrs, and 24 hrs for (b) the as-received austenitic stainless steel, SS316L, (d) 1800A and (f) 1800C. The symbols represent the measured data while the lines represent the fit data obtained by employing the electrical equivalent circuit as shown in the inset. Figure 7: Low magnification scanning electron microscope (SEM) image showing the corroded surface for (a) the as-received austenitic stainless steel, SS316L, (b) SS316L friction stir processed under submerged cooling (1800C); High magnification SEM image of a pit in (c) asreceived SS316L, (d) 1800C. 3D profiler images of the pits for (e) as-received SS316L, (f) 1800C. As-received specimen showed significantly higher pit density compared to 1800C specimen. Depth profiles of pits from as-received alloy and 1800C specimen are shown in (g) and (h) respectively. Color bar shows depth at different locations in the pit. 1800A shows similar features as 1800C and is not shown here. Figure 8: Wear volume loss at 5N and 3N normal load for the as-received austenitic stainless steel, SS316L, SS316L friction stir processed under ambient cooling (1800A) and SS316L friction stir processed under submerged cooling (1800C) in (a) dry condition, (b) in ringer solution; Wear track profiles at 5N load for as-received steel, 1800A and 1800C specimens in (c) dry wear, (d) ringer solution; Correlation of wear volume loss for all specimens with (e) hardness and (f) hardness to modulus ratio (H/E). Figure 9: Low magnification scanning electron microscope (SEM) images for 1800C subjected to dry wear testing at (a) 5 N normal load, (b) 3 N normal load; (c) and (d) High magnification SEM images of (a) and (b) respectively; (e) and (f) EDS mapping of the images shown in (c) and (d) respectively. Low magnification scanning electron microscopy (SEM) images for 1800C subjected to wear testing in ringer solution at (g) 5 N normal load, (h) 3 N normal load; (i) and (j) High magnification SEM images of (g) and (h) respectively; (k) and (l) EDS mapping of the images shown in (i) and (j) respectively. Figure 10: (a) Tryphan blue exclusion assay, percentage of living MDCK cells evaluated after 1 day of culture; (b) MTT-Cytotoxicity assay, percentage of dead cells after MDCK Assay. Fluorescence micrographs of MDCK cells after 1 day of culture on (c) Polystyrene (Control), (d) as-received, (e) 1800A, and (f) 1800C samples. (Phalloidin488, AlexaFluor594 and DAPI stained actin cytoskeleton (red), and cell nuclei (blue), respectively). Data are presented as the mean ± SD, Error bars represent standard deviation, Statistical significance - p* < 0.05, p** < 0.001.


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Figure 11: Hemocompatibility of the studied alloys: (a) hemolysis percentages, (D.I –De-ionized water used as positive control), (b) protein adhesion percentages by ELISA assay, (c) standard used to record platelet adhesion activity at an absorbance of 490 nm, (d) platelet adhesion percentages by LDH assay; SEM micrographs of platelets adhering to (e) control polystyrene plate and (f) 1800C. Data are presented as the mean ± SD, Error bars represent standard deviation, Statistical significance - p* < 0.05, p** < 0.001.

Table caption Table 1: Electrochemical parameters obtained from cyclic polarization test of as-received alloy, 1800A and 1800C specimens. Icorr, Ecorr, Epit, Epp represents corrosion current, corrosion potential, pitting potential, and passivation potential respectively. Table 2: Electrochemical parameters obtained after fitting the EIS data of as-received alloy, 1800A and 1800C specimens in Ringers solution measured at open circuit potentials at different immersion time. Rs is the solution or electrolyte resistance, CPE1 and R1 are the electrochemical double layer capacitance and charge transfer resistance respectively for the high frequency part of the spectrum, while CPE2 and R2 are the corresponding elements for the low frequency spectrum.

References: (1) Park, J.; Lakes, R. S., Biomaterials: Metallic Implant Materials, Springer New York: New York, NY: 2007 (2) Basu, B.; Katti, D. S.; Kumar, A., Adv. Biomater. : Fundamentals, Processing, and Applications. John Wiley & Sons: 2010. (3) Ratner, B. D.; Hoffman, A. S.; Schoen, F. J.; Lemons, J. E. Biomater. Sci. : An Introduction to Materials in Medicine, Academic press: 2004. (4) Geetha, M.; Singh, A.; Asokamani, R.; Gogia, A. Ti based Biomaterials, The Ultimate Choice for Orthopaedic Implants–a Review. Prog. Mater. Sci. 2009, 54 (3), 397-425. (5) Nagarajan, S.; Mohana, M.; Sudhagar, P.; Raman, V.; Nishimura, T.; Kim, S.; Kang, Y. S.; Rajendran, N. Nanocomposite Coatings on Biomedical Grade Stainless Steel for Improved Corrosion Resistance and Biocompatibility. ACS Appl. Mater. Interfaces 2012, 4 (10), 5134-5141. (6) Muley, S. V.; Vidvans, A. N.; Chaudhari, G. P.; Udainiya, S. An Assessment of Ultra Fine Grained 316L Stainless Steel for Implant Applications. Acta Biomater. 2016, 30, 408-419. (7) Misra, R.; Nune, C.; Pesacreta, T.; Somani, M.; Karjalainen, L. Understanding the Impact of Grain Structure in Austenitic Stainless Steel from a Nanograined Regime to a Coarse-Grained Regime on Osteoblast Functions using a Novel Metal Deformation–Annealing Sequence. Acta Biomater. 2013, 9 (4), 6245-6258. (8) Misra, R.; Thein-Han, W.; Pesacreta, T.; Hasenstein, K.; Somani, M.; Karjalainen, L. Cellular Response of Preosteoblasts to Nanograined/Ultrafine-grained Structures. Acta Biomater. 2009, 5 (5), 1455-1467.



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Sample Asreceived 1800A 1800C

Icorr (µA/cm2)

Ecorr (mV)

Epit (mV)

Epp (mV)

Epit-Ecorr (mV)

Epp-Ecorr (mV)

10.68

-295

407

13.35

702

308.35

Area under the curve x 10-4 (V.A/cm2) 7.65

6.75 5.53

-353 -345

395 475

37.6 72.82

748.6 820

390.6 417.82

5.74 5.64

Table 1: Electrochemical parameters obtained from cyclic polarization test of as-received alloy, 1800A and 1800C specimens. Icorr, Ecorr, Epit, Epp represents corrosion current, corrosion potential, pitting potential, and passivation potential respectively.



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Sample

Time

Rs (ohm)

R1 (k ohm cm2)

R2 (k ohm cm2)

Total R (k ohm cm2)

CPE1 (µ S sn cm2 )

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n1

CPE 2 (µ S sn cm-2)

Table 2: Electrochemical parameters obtained after fitting the EIS data of as-received alloy, 1800A and 1800C specimens in Ringers solution measured at open circuit potentials at different immersion time. Rs is the solution or electrolyte resistance, CPE1 and R1 are the electrochemical double layer capacitance and charge transfer resistance respectively for the high frequency part of the spectrum, while CPE2 and R2 are the corresponding elements for the low frequency spectrum.


n2

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0hr

54.4

0.975

6.300

7.275

49.02

0.66

741.2

0.5

12hr

54.2

1.979

15.62

17.559

41.22

0.67

589.3

0.5

24hr

54.7

2.736

18.26

20.996

39.98

0.67

537.9

0.5

0hr

54.4

258.9

--

258.9

40.66

0.852

--

--

12hr

54.0

335.3

--

335.3

30.29

0.88

--

--

24hr

55.9

598.1

--

598.1

11.45

0.87

--

--

0hr

55.6

394.8

--

394.8

23.04

0.82

--

--

12hr

56.2

568.0

--

568.0

11.19

0.86

--

--

24hr

54.6

737.5

--

737.5

11.93

0.86

--

--

Asreceived

1800A

1800C



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413x251mm (96 x 96 DPI)

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Figure 1. Schematic diagram explaining the novel submerged friction stir processing (FSP). The surface of as-received (SS316L) alloy was modified using experimental setup specially designed for submerged processing. 254x220mm (96 x 96 DPI)



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Figure 2. Optical microscope images of the cross section at 100X magnification for (a) SS316L, friction stir processed under submerged cooling (1800C) and (c) SS316L, friction stir processed under ambient cooling (1800A); at 200X magnification for (b) 1800C and (d) 1800A. The processed region comprises a highly refined zone (nugget) of nearly 80-100 µm thickness and thermomechanical affected zone (TMAZ) zone of nearly 70-80 µm with total thickness of the processed region as 170-180 µm. 228x169mm (96 x 96 DPI)


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Figure 3. EBSD map for (a) as-received austenitic steel SS316L, (c) friction stir processed SS316L under ambient cooling (1800A) and (e) friction stir processed SS316L under submerged cooling (1800C); Grain size distribution for (b) as-received SS316L, (d) 1800A and (f) 1800C. 225x274mm (96 x 96 DPI)



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Figure 4: (a) Hardness as function of depth for the as-received austenitic stainless steel, SS316L, SS316L friction stir processed under ambient cooling (1800A) and SS316L friction stir processed under submerged cooling (1800C); (b) Variation of peak hardness and elastic modulus for the as-received austenitic stainless steel, SS316L, 1800A and 1800C specimens; (c) Normal direction (ND) inverse pole figure (IPF) for asreceived, 1800A and 1800C specimens. The color bar shows the texture intensity in each case. 251x200mm (96 x 96 DPI)


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Figure 5: Cyclic polarization curves for (a) the as-received austenitic stainless steel, SS316L, (b) SS316L friction stir processed under ambient cooling (1800A), (c) SS316L friction stir processed under submerged cooling (1800C), The pitting potential (Epit), corrosion potential (Ecorr), and protection potential (Epp) for all specimens are marked on the polarization curves, (d) Correlation between corrosion current and grain size for as-received SS316L, 1800A and 1800C. 255x200mm (96 x 96 DPI)



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Figure 6: Electrochemical impedance measurements showing Nyquist plot at different immersion times of 0 hrs, 12 hrs, and 24 hrs for (a) the as-received austenitic stainless steel, SS316L, (c) SS316L friction stir processed under ambient cooling (1800A), (e) SS316L friction stir processed under submerged cooling (1800C); Bode plot at different immersion times of 0 hrs, 12 hrs, and 24 hrs for (b) the as-received austenitic stainless steel, SS316L, (d) 1800A and (f) 1800C. The symbols represent the measured data while the lines represent the fit data obtained by employing the electrical equivalent circuit as shown in the inset. 248x273mm (96 x 96 DPI)


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Figure 7: Low magnification scanning electron microscope (SEM) image showing the corroded surface for (a) the as-received austenitic stainless steel, SS316L, (b) SS316L friction stir processed under submerged cooling (1800C); High magnification SEM image of a pit in (c) as-received SS316L, (d) 1800C. 3D profiler images of the pits for (e) as-received SS316L, (f) 1800C. As-received specimen showed significantly higher pit density compared to 1800C specimen. Depth profiles of pits from as-received alloy and 1800C specimen are shown in (g) and (h) respectively. Color bar shows depth at different locations in the pit. 1800A shows similar features as 1800C and is not shown here. 251x391mm (96 x 96 DPI)



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Figure 8: Wear volume loss at 5N and 3N normal load for the as-received austenitic stainless steel, SS316L, SS316L friction stir processed under ambient cooling (1800A) and SS316L friction stir processed under submerged cooling (1800C) in (a) dry condition, (b) in ringer solution; Wear track profiles at 5N load for asreceived steel, 1800A and 1800C specimens in (c) dry wear, (d) ringer solution; Correlation of wear volume loss for all specimens with (e) hardness and (f) hardness to modulus ratio (H/E). 249x275mm (96 x 96 DPI)


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Figure 9: Low magnification scanning electron microscope (SEM) images for 1800C subjected to dry wear testing at (a) 5 N normal load, (b) 3 N normal load; (c) and (d) High magnification SEM images of (a) and (b) respectively; (e) and (f) EDS mapping of the images shown in (c) and (d) respectively. Low magnification scanning electron microscopy (SEM) images for 1800C subjected to wear testing in ringer solution at (g) 5 N normal load, (h) 3 N normal load; (i) and (j) High magnification SEM images of (g) and (h) respectively; (k) and (l) EDS mapping of the images shown in (i) and (j) respectively. 225x542mm (96 x 96 DPI)



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Figure 10: (a) Tryphan blue exclusion assay, percentage of living MDCK cells evaluated after 1 day of culture; (b) MTT-Cytotoxicity assay, percentage of dead cells after MDCK Assay. Fluorescence micrographs of MDCK cells after 1 day of culture on (c) Polystyrene (Control), (d) as-received, (e) 1800A, and (f) 1800C samples. (Phalloidin488, AlexaFluor594 and DAPI stained actin cytoskeleton (red), and cell nuclei (blue), respectively). Data are presented as the mean ± SD, Error bars represent standard deviation, Statistical significance - p* < 0.05, p** < 0.001. 242x274mm (96 x 96 DPI)


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Figure 11: Hemocompatibility of the studied alloys: (a) hemolysis percentages, (D.I –De-ionized water used as positive control), (b) protein adhesion percentages by ELISA assay, (c) standard used to record platelet adhesion activity at an absorbance of 490 nm, (d) platelet adhesion percentages by LDH assay; SEM micrographs of platelets adhering to (e) control polystyrene plate and (f) 1800C. Data are presented as the mean ± SD, Error bars represent standard deviation, Statistical significance - p* < 0.05, p** < 0.001. 251x270mm (96 x 96 DPI)


Friction Stir Processing of Stainless Steel for Ascertaining Its

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