Effect of microstructure and strain on the degradation behavior of novel bioresorbable ironmanganese alloy implants
Michael Heiden a, Andrew Kustas a, Kevin Chaput a, Eric Nauman b, c, David Johnson a, Lia Stanciu a
a
School of Materials Science and Engineering, Purdue University, West Lafayette, IN 47907,
USA b
Weldon School of Biomedical Engineering, Purdue University, West Lafayette, IN 47907,
USA c
School of Mechanical Engineering, Purdue University, 585 Purdue Mall, West Lafayette, IN
47907, USA
This article has been accepted for publication and undergone full peer review but has not been through the copyediting, typesetting, pagination and proofreading process which may lead to differences between this version and the Version of Record. Please cite this article as an ‘Accepted Article’, doi: 10.1002/jbm.a.35220
Abstract Advancing the understanding of microstructural effects and deformation on the degradability of Fe-Mn bioresorbable alloys (specifically Fe-33%Mn), will help address the current problems associated with designing degradable fracture fixation implants for hard tissues. Potentiostatic polarization tests were conducted on a wide variety of metal samples in order to examine how different deformation processes affect the instantaneous rate of degradation of Fe-Mn alloys. Large-strain machining (LSM), a novel severe plastic deformation (SPD) technique was utilized during these experiments in order to modify the degradation properties of the proposed Fe-Mn alloy. It was discovered that Fe-33%Mn after LSM with a 0° rake angle (effective strain = 2.85) showed the most promising increase in degradation rate compared to as-cast, annealed, and additional deformation conditions (rolled and other LSM parameters) for the same alloy. The mechanisms for enhancement of the corrosion rate are discussed. Keywords: Iron-manganese, Biodegradable, In vitro, Microstructure, Corrosion 1. Introduction The pursuit for a degradable biomaterial that is able to mechanically support the healing process of bone fracture for a limited amount of time and gradually be replaced by surrounding tissue has been burdened with complications. One of the biggest issues from a clinical perspective deals with tailoring the degradation rates of biodegradable implants in a diverse range of physiological environments for complete tissue reconstruction.1-4 There is clinical relevance in identifying materials that can be employed to replace existing, permanent fracture fixation devices, such as staples, suture anchors, screws, or pins, and eliminate the need of a second surgical operation for device removal.5-12 It has been established by numerous authors that the ideal biodegradable material has the following characteristics: 1) biocompatible/non-
toxic/non inflammatory, 2) fosters natural tissue ingrowth and cell proliferation, 3) has mechanical properties that support the surrounding tissue during the full reconstruction process, 4) degrades at a rate that doesn’t generate local or systemic effects in the body, and 5) releases its degradation products in a way that can be safely eliminated from the body by the excretory system.2 Currently, no such bioresorbable material fulfills all these requirements for either hard or soft tissues. For orthopedic applications, bioresorbable metallic implants have been found to have several advantages over degradable polymers. Metals tend to have higher tensile, compressive, and yield strengths. They also have high fracture toughness, do not exhibit high stress relaxation which can discourage natural tissue ingrowth, and are not radiolucent.2, 3 Bone fracture fixation materials in particular require mechanical properties that can withstand a combination of factors, including mechanical fatigue, erosion due to device insertion, and degradation due to physiological pH, temperature, and other biological intrusions in the hard tissue environment.4 A toxicological risk assessment in the body is another crucial step in designing a safe implant and it is known that the human body has exposure limits for metallic elements.13 The most commonly investigated bioresorbable metallic materials due to biocompatible degradation products include magnesium alloys and iron alloys.14, 15 Yuen et al. reported a toxicological assessment for several metals, evaluating the critical exposure limits in the human body.13 For an average 60-kg adult, iron was found to have an exposure limit of 2.55 mg/day, manganese was 0.42 mg/day, and magnesium was specified in terms of a maximum daily intake of 250-400 mg/day.13 Magnesium alloys have implant potential because their elastic moduli closely mimics that of most human cortical bones, but pure magnesium’s high degradation rate is too fast for maintaining the needed mechanical support for the full period of hard tissue reconstruction.4
Furthermore, due to this high corrosion rate, magnesium causes subcutaneous hydrogen gas bubbles to form in the body that need to be released by invasive needles.4 Unlike magnesium, pure iron has exceptional mechanical properties for supporting the rebuilding of hard tissue, however it degrades too slowly.16-18 Additionally, with a higher initial stiffness, it has the potential to cause stress-shielding at the current rate of resorbability.16 Alloys of Fe-Mn have been proposed for quite some time due to their slightly faster corrosion rates and nonmagnetic properties above 20% Mn content, which aids in the process of examining implants using magnetic resonance imaging (MRI).17 However, previously-designed Fe-Mn bioresorbable implants have degradation rates that are still too slow to be beneficial for orthopedic implantation. Hermawan et al. processed Fe-Mn bulk forms from sintering Fe and Mn powders, and found that a more porous structure helped increase the corrosion rate of the alloy immersed in modified Hank’s Solution due to the high amount of surface area.18 Conversely, Chou et al. found that 3D printing an open-pore structure helped increase the corrosion rate, but noticeably decreased the implant’s mechanical properties.19 Factors that appear to influence the corrosion rate of metals in a liquid medium include the inherent electrode potential of the metal (affects reactivity), the surface state and passivity (more passive surfaces slow corrosion rate), and the amount of surface area in contact with the medium (affects amount of ion diffusion).16-19 In order to better understand the degradation kinetics of Fe-Mn alloys in osteogenic environments, the present authors endeavored to modify the degradation rate of a Fe-33%Mn alloy by employing a variety of deformation processes, one of which included severe plastic deformation (SPD) via orthogonal plane-strain machining, in order to refine the microstructure and examine how strain plays a role in degradation behavior. It should be noted that a manganese
composition of 33% was chosen in order to further investigate how the degradation results of this alloy compared with other Fe-Mn alloy compositions presented by other authors.15-18 The use of conventional plane-strain machining is an attractive manufacturing process for fundamental studies of microstructure-property relationships in a variety of metals. The high strains, strain-rates, and temperatures achievable during machining provide unique opportunities for microstructure refinement, even in the nanometer regime, through appropriate selection of processing parameters.20 The ability to control the microstructure can have substantial effects on material properties. Furthermore, once the optimum microstructure/corrosion relationship has been established, scale-up to continuous strip or wire is possible.21, 22 2. Materials and methods 2.1 Materials (casting) Small, irregular pieces (≤12mm [0.5in]) of pure Fe (Alfa Aesar, Ward Hill, MA, 99.9%) and pure Mn (Alfa Aesar, Ward Hill, MA, 99.9% Mn) were mixed to create a Fe–33%Mn alloy composition. This specific mixture was arc melted under a flowing argon (5%H2) atmosphere and re-melted at least three times to promote homogeneity. The Fe-Mn ingot was then further processed in a vacuum induction melter (VIM). The VIM was evacuated to a pressure of ~ 10-3 atm and then purged with argon three times to reduce the oxygen content in the furnace. After the final purge, the chamber was backfilled with argon to minimize vaporization of the manganese. The alloy was melted and poured into a cylindrical copper mold resting on a water cooled copper plate to create a 1” diameter ingot 3.5” in length. The cast Fe-33%Mn alloy cylinder was sectioned into 1cm x 0.5cm rectangular pieces ~0.2cm thick. The Fe-33%Mn samples underwent a wider variety of processing techniques, including: 1) heavy cold-rolling (reductions of 50%), and 2) large strain machining (LSM).
Select cast and cold-rolled Fe-33%Mn samples were then annealed at 1300°C for 1 hour to provide additional microstructure variation. Initial mechanical properties such as elastic modulus, hardness, and stiffness were also measured using nanoindentation (Hysitron TI-950 Triboindenter, Minneapolis, MN) with a Berkovich tip with a nominal radius of 150 nm. 2.2 Deformation Processing Plane-strain machining (Figure 1) employs a sharp, wedge tool with rake angle (α) set at a preset depth (ao) into the work piece (bulk), which is travelling at constant surface velocity (Vo). Material is removed from the bulk piece in the form of a chip-extrudate with deformed thickness (ac). Three variables (α, ao, Vo) control most of the deformation characteristics, creating very large shear strains (1-15), at high strain rates (104 s-1) and high temperatures during cutting.23 The severe deformation occurs within a very narrow zone (∆ ~ 50 µm wide), resulting in microstructure refinement of the extrudate through simultaneous accumulation of large strain and localized (adiabatic) heating.22-24 The narrow width of the deformation zone is most commonly approximated as a single shear plane with orientation angle ϕ from the surface of the work piece. Using this approximation, the shear strain can be calculated simply using an upperbound model as23, 25: (1)
=
() ( ) ( )
Where α is the tool rake angle and ϕ is the shear plane angle, which is determined from measuring ao and ac and through geometry: (2)
tan(φ) =
() ()
By definition, the von-Mises effective strain for plane-strain machining is expressed as:
(3)
=
√
Effective strain provides a convenient metric for the direct comparison of different deformation processes (e.g. rolling and machining) by reducing complex straining situations to a single component.25 Equations 1-3 23, 25 show that a variety of strains are possible in machining through the selection of different tool rake angles. In this study, rectangular chips of Fe-33%Mn were produced from cutting tools at three increasingly negative rake angles (+20°, 0°, -20°) to increase both stored strain energy and microstructural refinement. Effective strains for the three machining parameters range from “low” (~1.5) to “high” (~3.5).
2.3 Solution preparation and microstructural analysis All samples were ground to 600 grit, polished with 1µm alumina powder, rinsed with deionized water, rinsed with ethanol, and then air dried before immersion into osteogenic cell culture media for electrochemical experiments. Osteogenic cell culture media was prepared according to the procedure for hard tissue fluid outlined by Dickerson et al.26 For 300 mL total media: 1032.3 µL cell culture media (low glucose), 15 µL ascorbic acid, 2.7 µL Bglycerophosphate 3mM, 11.76 µL Dexamethasone, 30 mL FBS (fetal bovine serum), 3 mL Penicillin Streptomysin, and 265.938 mL D-MEM media.26 All Fe-33%Mn processed samples were mounted in Bakelite, mechanically ground and polished with 0.02µm silica, and then chemically etched using Nital solution (2% HNO3 in ethanol). Optical microscopy was performed to examine microstructures after each processing
step and scanning electron microscopy (SEM; FEI Philips XL-40) coupled with energydispersive x-ray spectroscopy (EDS) was used to examine the surface further while also acquiring an elemental analysis. X-ray diffraction (XRD) with Cr Kα radiation was also employed (Bruker D8 Discover X-Ray Diffractometer with GADDS (General Area Detector Diffraction System)) for characterization of microstructural phases present in the cast Fe-Mn alloy. Density was approximated by the Archimedes Principle. In determination of microstructure, the average grain size/diameter was measured for each processing technique applied by counting grains from three micrographs for each sampling, taken from different places on the specimens. The images were processed for grain size measurement using ImageJ software and measured according to standard ASTM E112.27, 28 2.4. Linear polarization test Rectangular samples of Fe-33%Mn for each processing technique were immersed in osteogenic media and tested in triplicate. Tests were conducted according to ASTM G 59L.29 Electrochemical impedance spectroscopy (EIS) tests were carried out under the conditions of 37±0.5 °C, with a pH of 7.4±0.1; using a potentiostat (SP-150 Biologic Science Instrument, Claix, France). The experimental setup was a standard three-electrode system with the sample as the working electrode, a saturated calomel electrode (SCE) as the reference electrode, and a platinum wire as the counter electrode. The results of the measurements were analyzed using EC-Lab v-10.19 BioLogic Science Instruments software. Each sample was submerged in 5mL of osteogenic media and the media was heated and maintained at 37±0.5 °C. Media was changed out for each test. The open circuit potential (OCP) was recorded after 55 minutes of immersion to maintain steady-state conditions of stable free corrosion potential. The sweep rate of the polarization curve measurement was set to 0.1667mVs-1 and the potential range was -150 to
+150 mV vs. OCP. Linear polarization curves were recorded until the potentials reached the ends of the designated range. Ecorr and Icorr were recorded for each polarization test. The generated Tafel plots based on the cathodic and anodic curves were fitted linearly to measure the intersection of the corrosion potential (Ecorr) and total corrosion current (Icorr). The corrosion current density (icorr) for each sample was then calculated by dividing Icorr by the exposed surface area of each specimen. The analyzed icorr was further converted into the corrosion rate (mm/year) using29, 30: (4)
Corrosion Rate (mm/year) = 3.27x103
∗ !
where icorr is the corrosion current density (A/cm2), D is the sample standard density (g/cm^3), and K is a constant; where 3272 was used to obtain units in (mm/yr), EW is the equivalent weight according to the percent of alloying elements (Fe33Mn: 27.77 g*eq-1). Finally, polarization curves were plotted together for comparison of every type of processing technique imparted upon the Fe-33%Mn samples, while average corrosion rates were calculated and then plotted in a bar graph with confidence intervals. 3. Results and Discussion 3.1. Mechanical properties of the Fe-Mn alloy Data from EDS shown in Table 1 indicate the average compositions of the as-cast alloy were homogeneous throughout the samples near the desired compositions with some microsegregation observed, typical of common solidification processes. Additionally, nanoindentation results are summarized in Table 2. After casting the Fe-Mn alloy in the VIM, the microstructure displayed a columnar grain morphology with grains growing radially inward towards the centerline of the ingots. Figure 2
shows the dendritic structure to be relatively coarse. When examining the binary phase diagram for Fe-Mn, it can be found that the partition coefficient (k) is less than one, indicating the interdendritic liquid during solidification is expected to be richer in manganese while the dendrites will be depleted due to solute rejection.32 Figure 2 indeed confirms this expectation, i.e. regions between dendritic arms are higher in manganese than the dendrites themselves. The electrochemically degraded Fe-33%Mn cast sample shown in Figure 3 reveals the media’s preferential corrosion on the manganese rich, inter-dendritic regions. It was observed that the elastic modulus and density of the Fe-33%Mn alloy (Table 1) were lower than that of the pure iron19, but higher than that of 316L stainless steel.33 An elastic modulus closer to that of bone is ideal to improve implant interface/tissue interactions, decreasing chance of failure. However, the results from Table 3 show that besides looking into simply minimizing the material’s Young’s Modulus, it is also important to characterize and at least qualitatively assess the interrelationship between deformation, composition (percent Mn), and degradation rate. For example, Hermawan et al. found both UTS and YS to decrease with increasing manganese content in four porous Fe-Mn alloys, while elongation at fracture increased.30 For a more complete analysis of Fe-Mn alloys being employed as a clinically viable bioresorbable implant material, it would be beneficial for a wider range of Fe-Mn alloy compositions to have their degradation rates tested while measuring the changes in all of their mechanical properties. 3.2. Influence of microstructural evolution and processing method on corrosion rate Linear polarization curves obtained for the Fe-33%Mn samples that were tested in the osteogenic media are shown in Figure 4. Additionally, the legend lists the curves in order from higher obtained Ecorr values to lower Ecorr values for enhanced comparisons since there is a
large grouping of polarization curves around Ecorr = -0.7V. Among the Fe-33%Mn alloys, the highest electrode potential was found to be the samples that underwent LSM at α= -20°, which interestingly underwent the highest amount of effective strain. In contrast, the lowest Ecorr was the annealed sample (1300°C) after CW50%. The BioLogic software’s tafel line fit program was used to measure icorr values from the polarization curves, which was then utilized to calculate the corresponding corrosion rates for each type of processed sample. Figure 5 provides the resulting corrosion bar graphs after analysis with standard error bars. Table 3 shows the actual averaged Ecorr, icorr, corrosion rate (CR), and standard error for CR values for all the samples. All processed samples were etched using 2% Nital and then optically imaged in order to examine the microstructural evolution shown in Figure 6. The cast structures were all dendritic in nature and the cold-rolled samples show an elongated grain structure along the rolling direction, which is typical of a rolled microstructure. Upon annealing at 1300°C, the rolled sample exhibits a relatively equiaxed recrystallized structure with grains on the order of ~50µm in diameter. The LSM samples all showed heavily sheared microstructures. Analysis on an etched, cast Fe-33%Mn sample using EDS show the initial dendritic structures are richer in manganese compared to the cooled inter-dendritic liquid. Figure 2 displays the compositions obtained for the thick dendrites and the thin interdendritic regions. Linear polarization scans showed interesting variations in Ecorr and icorr for different processing techniques. The shifts in polarization curves seen in Figures 4-5 demonstrate that both Ecorr and icorr increased simply by alloying manganese with iron, comparable to what Hermawan et al found for sintered Fe-Mn structures.30 Pure iron and zinc were used for
comparison purposes only and clearly had too low of Ecorr, icorr, and degradation rates for bioresorbable fracture fixation applications. It was found that cold rolling Fe-33%Mn to 50% increased the corrosion rate compared to the cast structure. Furthermore, annealing Fe-33%Mn at 1300 for an hour after cold rolling dampened the corrosion rate. Interestingly, annealing the sample with no cold work increased its corrosion rate slightly. This phenomenon potentially resulted from homogenization of the microstructure, effectively reducing microsegregation, e.g. more diffusion of solute atoms from solute-rich regions to solute-depleted regions. All LSM conditions increased the degradation rate of the cast structure substantially, but LSM at α=0° (ε~2.85) showed the most improvement in corrosion rate. When examining the microstructures of the machined chips (Figure 6 e-g), the dendritic structure has clearly undergone large deformation under each condition. Unlike the rolled microstructures with dendrites elongated along the direction of rolling, chip microstructures show the dendrites sheared and oriented in a manner mimicking the angle of the shear deformation plane. Interestingly, degradation rates do not continuously increase with increasing deformation strain. This indicates that larger deformation strains do not necessarily lead to higher corrosion rates as assumed in a linear correlation. The degradation rate/grain size relationship may be more of a bimodal distribution in nature as suggested by Moravej et al’s data.16 Moravej et al. showed that annealed Fe with an average grain size of 40µm had a higher degradation rate than Fe with nanocrystalline-sized grains, but a lower degradation rate than electroformed Fe with 2-8µm grains.16 It has been established by others that decreasing grain size increases grain boundary area and thus creates a high internal energy in a material.31 Since corrosion tends to target grain boundaries first, it is assumed that decreasing grain-size will promote higher corrosion rates. This effect was partially
observed in the current study. The results from Table 4 indicate that grain size appears to influence the degradation behavior more than induced plastic strain. However, it should be noted that it is difficult to discern the grain size of the LSM samples since etching the surfaces revealed dendritic and interdendritic bands rather than grain boundaries, as expected in a homogenized microstructure. This deformed microstructure is analogous to banded structures found after cold rolling of a cast part.34 This deformed structure makes it difficult to discern the grain size for these LSM samples, however it is possible to compare the LSM dendritic band thicknesses (Figure 7) to the original dendritic spacing (Figure 8). It can be seen that the cast structure’s average primary dendrite arm thicknesses (15-40 um) decrease to 2-10um after LSM at α=0°. These measurements along with the degradation results suggest that preferential attack on the interdendritic areas occurs during degradation of these alloys, leading to an overall increase in corrosion rate due to the increase of interdendritic surface area per unit volume with decreasing band thickness. It can also be concluded that the degradation rate is controllable, to some extent, by the amount of deformation strain induced in these classes of biodegradable materials. It appears that increasing the amount of deformation strain in the material can increase the electrode potential of the material, yet the degradation rate increases only up to a certain point in terms of induced strain. Beyond this point, increasing deformation strain does not appear to enhance the corrosion rate. It was observed that too large of a strain (LSM α = -20°) had adverse effects on increasing the degradation rate (Table 5). However, it appears that simply reducing grain size doesn’t necessarily lead to a lower degradation rate, as evidenced by the low corrosion rates of the smallgrained, annealed structure. Based on these results, there appears to be an interrelated grainsize/plastic strain relationship on material degradation rate. The precise role of each mechanism
is currently unknown; however it has been well-documented that continually increasing deformation strain can lead to continuous recrystallization, thereby developing a refined microstructure.35 Therefore, at the highest deformation strain, it is possible some amount of continuous recrystallization occurred during deformation, which would decrease the total stored energy and consequently, the degradation rate relative to the lower strain conditions. From these results, there appears to be a significant influence of deformation strain on the degradation rate of the material. Moreover, it appears high effective strains exhibit relatively high degradation rates. Comparatively, the Fe-33%Mn LSM sample at α= +20° had an effective strain roughly double that of the cold rolled sample that underwent 50% CW, but its corrosion rate was only increased by 0.0054mm/year.
5. Conclusion In this study, the in vitro degradation rates of potential bioresorbable materials were tailored by varying the processing routes for a Fe-33%Mn alloy. The novel LSM technique at a rake angle of 0° showed a 140% increase in corrosion rate for Fe-33%Mn compared to a cast structure of the same alloy. The authors suggest that in order to increase the degradation rate further for Fe-Mn alloys, 1) tailored shear-based deformation processing should be employed to increase the kinetic effects of corrosion up to a critical value, and 2) the surface area of the implants should be increased to allow for more diffusion of the osteogenic media into the Fe-Mn bulk alloy, which would also provide more adhesion and ingrowth of the hard tissue environment. To better tailor the properties of these potential bioresorbable Fe-Mn hard tissue implants, it is essential to balance the mechanical properties needed for orthopedic support with the degradation rate of the material. It appears that in order to achieve higher degradation rates,
refinement of the grains is needed up to a critical value, while also being careful to avoid increasing the material’s elastic modulus and strength too much, thereby leading to poorer interface/tissue bonding. Other authors’ findings and the results found above seem to indicate that an open-porous, Fe-Mn fine-grained structure deformed by shear might be more beneficial for achieving the clinically desired rate of bioresorbability for hard tissue implants. In either case, the study garners additional evidence that Fe-Mn alloys have potential for being highly degradable for orthopedic applications. Acknowledgements The authors would like to thank the help of Julia Aspaugh for aiding in osteogenic media preparation, Kaitlyn Jarry for aiding in the electrochemical experiments, Samantha Lawrence for aiding in nanoindentation, and Kristin Paulson for XRD training. Guidance covering the nature of material corrosion was gratefully provided by Dr. Robert Spitzer of Purdue University. Furthermore, the authors would like to thank Dr. Srinivasan Chandrasekar of Purdue University for use of his machining equipment used towards LSM the Fe-Mn alloy. The authors acknowledge the financial assistance from Fort Wayne Metals, IN in supporting this work. References 1. Peuster M, Wohlsein P, Brügmann M, Ehlerding M, Seidler K, Fink C, et al. A novel approach to temporary stenting: degradable cardiovascular stents produced from corrodible metal-results 6-18 months after implantation into New Zealand white rabbits. Heart (British Cardiac Society) 2001; 86:563–9. 2. Hofmann GO, Wagner FD. New implant designs for bioresorbable devices in orthopaedic surgery. Clinical Materials 1993; 14:207–15.
3. Levesque J, Dube D, Fiset M, Mantovani D. Materals and properties for coronary stents. Advanced Materials & Processes. 162 (9) 2004; 45-48. 4. Zeng R, Dietzel W, Witte F, Hort N, Blawert C. Progress and Challenge for Magnesium Alloys as Biomaterials. Advanced Engineering Materials 2008; 10: B3–B14. 5. Wierer, M., P. Biberthaler, W. Mutschler, and S. Grote, [Removal of a bent intramedullary tibia nail. Case report and review of literature]. Unfallchirurg, 2011; 114(7): 629-33. 6. Harrison, M.R., S. Hamilton, and A.J. Johnstone, Pseudo-rupture of Extensor Pollicis Longus following Kischner wire fixation of distal radius fractures. Acta Orthop Belg, 2004; 70(5): 492-4. 7. Velich, N., Z. Nemeth, C. Suba, and G. Szabo, Removal of titanium plates coated with anodic titanium oxide ceramic: retrospective study. J Craniofac Surg, 2002; 13(5): 636-40. 8. Schultz, J.H., D. Wolter, G. Ortel, and B. Fink, [Fracture treatment in the area of the tibia]. Unfallchirurg, 1992; 95(11): 537-40. 9. Velazco, A. and L.L. Fleming, Open fractures of the tibia treated by the Hoffmann external fixator. Clin Orthop Relat Res, 1983; (180): 125-32. 10. Rehm, A., A. Divekar, and M.E. Conybeare, External fixation for femoral derotation osteotomy in developmental dysplasia of the hip. J Pediatr Orthop B, 2003; 12(5): 319-27. 11. Landry, P.S., A.A. Marino, K.K. Sadasivan, and J.A. Albright, Effect of soft-tissue trauma on the early periosteal response of bone to injury. J Trauma, 2000; 48(3): 479-83. 12. Wachter, R. and P. Stoll, [Can steel screws be combined with titanium plates? Hard polishing technique and SEM in animal experiments]. Dtsch Z Mund Kiefer Gesichtschir, 1991; 15(4): 275-84 13. Yuen CK, Ip WY. Theoretical risk assessment of magnesium alloys as degradable biomedical implants. Acta Biomaterialia 2010; 6:1808–12. 14. Denkena B, Lucas a. Biocompatible Magnesium Alloys as Absorbable Implant Materials – Adjusted Surface and Subsurface Properties by Machining Processes. CIRP Annals - Manufacturing Technology 2007; 56:113–6.
15. Schaffer JE, Nauman E a, Stanciu L a. Cold drawn bioabsorbable ferrous and ferrous composite wires: An evaluation of in vitro vascular cytocompatibility. Acta Biomaterialia 2012:1–11. 16. Moravej M, Mantovani D. Biodegradable metals for cardiovascular stent application: interests and new opportunities. International Journal of Molecular Sciences 2011; 12:4250–70. 17. Hermawan H, Purnama A, Dube D, Couet J, Mantovani D. Fe-Mn alloys for metallic biodegradable stents: degradation and cell viability studies. Acta Biomaterialia 2010; 6:1852–60. 18. Hermawan H, Alamdari H, Mantovani D, Dubé D. Iron–manganese: new class of metallic degradable biomaterials prepared by powder metallurgy. Powder Metallurgy 2008; 51:38–45. 19. Chou D-T, Wells D, Hong D, Lee B, Kuhn H, Kumta PN. Novel processing of iron-manganese alloybased biomaterials by inkjet 3-D printing. Acta Biomaterialia 2013. 20. Swaminathan S, Ravi Shankar M, Rao BC, Compton WD, Chandrasekar S, King AH, et al. Severe plastic deformation (SPD) and nanostructured materials by machining. Journal of Materials Science 2007; 42: 1529–41. 21. Efe M, Moscoso W, Trumble KP, Dale Compton W, Chandrasekar S. Mechanics of large strain extrusion machining and application to deformation processing of magnesium alloys. Acta Materialia 2012; 60: 2031–42. 22. Shaw, Milton. Metal Cutting Principles. 1. New York: Oxford Series on Advanced Manufacturing, 1996. 18-46. Print. 23. Sagapuram D, Efe M, Moscoso W, Chandrasekar S, Trumble KP. Controlling texture in magnesium alloy sheet by shear-based edeformation processing. Acta Materialia 2013. 24. Guo Y, Efe M, Moscoso W, Sagapuram D, Trumble KP, Chandrasekar S. Deformation field in largestrain extrusion machining and implications for deformation processing. Scripta Materialia 2012; 66:235–8. 25. Hosford, William, and Robert Caddell. Metal Forming Mechanics and Metallurgy. 2. Upper Saddle River, NJ: PTR Prentice Hall, 1993. 29-48. Print.
26. Dickerson D a, Misk TN, Van Sickle DC, Breur GJ, Nauman E a. In vitro and in vivo evaluation of orthopedic interface repair using a tissue scaffold with a continuous hard tissue-soft tissue transition. Journal of Orthopaedic Surgery and Research 2013; 8:18. 27. W.S. Rasband, ImageJ, U.S. National Institutes of Health, Bethesda, ML, USA, 2014 28. ASTM E112, Annual Book of ASTM Standards, American Society of Testing and Materials, West Conshohocken, PA, 2013, Standard Test Methods for Determining Average Grain Size, doi:10.1520/E0112-96R04E02. 29. ASTM Standard G59-97: Standard Test Method for Conducting Potentiodynamic Polarization Resistance Measurements, West Conshohocken, PA: ASTM International, 2009 30. Hermawan H, Dubé D, Mantovani D. Degradable metallic biomaterials: design and development of Fe-Mn alloys for stents. Journal of Biomedical Materials Research Part A 2010; 93:1–11. 31. Witte F, Hort N, Vogt C, Cohen S, Kainer KU, Willumeit R, et al. Degradable biomaterials based on magnesium corrosion. Current Opinion in Solid State and Materials Science 2008; 12:63–72. 32. Dantzig, J.A, and M Rappaz. Solidification. 1. Switzerland: EPFL, 2009. 395-425. Print. 33. Schaffer JE, Nauman E a., Stanciu L a. Cold-Drawn Bioabsorbable Ferrous and Ferrous Composite Wires: An Evaluation of Mechanical Strength and Fatigue Durability. Metallurgical and Materials Transactions B 2012; 43:984–94. 34. Grange, R. A. (1971). Effect of Microstructural Banding in Steel, 2(February), 417–426. 35. N Tsuji, Y Ito, Y Saito, Y Minamino, Strength and ductility of ultrafine grained aluminum and iron produced by ARB and annealing, Scripta Materialia, 2002; 47:893-899
Figure 1. Plain-strain machining diagram. Wedge tool uses a rake angle (α) set at a preset depth (ao) into the work piece (bulk), which is travelling at constant surface velocity (Vo). Chipextrudate has a deformed thickness (ac). ɸ is the angle from the shear plane to the horizontal preset depth. Figure 2. EDS “single spot” analysis on dendrite arm (a) and spacing between arms (b-c) of the cast Fe-33%Mn alloy at 8000X after Nital etching.
Figure 3. Surface of the cast Fe-33%Mn sample at 10X after experiencing electrochemical polarization.
Figure 4. Polarization curves for pure iron, pure zinc, and a variety of processed Fe-33%Mn alloys after being immersed in osteogenic media.
Figure 5. Average electrochemical corrosion rates of all tested samples while immersed in osteogenic media. Error bars are in terms of standard error.
Figure 6. Etched microstructures of Fe-33%Mn: (a) cast, (b) rolled CW50%, (c) annealed at 1300C (no CW), (d) annealed at 1300C (CW50%), (e) LSM α = +20, (f) LSM α = 0, (g) LSM α = -20. Figure 7. Dendritic (white) and interdendritic (black) band structures for the LSM α = -20. Figure 8. Primary and secondary dendrite arms of the cast sample, oriented from the edge of the circular ingot (top) to the center (bottom).
Journal of Biomedical Materials Research: Part A
Table 1. Chemical compositions of the cast Fe-Mn alloy Element Concentration Wt % Mn 33.08 ± 0.85 Fe 67.10 ± 0.97
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Table 2. Mechanical properties of cast Fe-33%Mn alloy Elastic Modulus (GPa)
201.47 ± 49.13
Density (g/cm^2)
7.64 ± 0.13
Vickers Microhardness (GPa)
6.64 ± 0.53
Stiffness (µN/nm)
182.54 ± 49.13
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Table 3. Electrochemical corrosion rates of samples immersed in osteogenic media
Fe33Mn LSM α= 0°
Ecorr (mV) -685.469
icorr (A/cm^2) 7.25E-05
Average CR (mmpy) 0.8354
Standard Error 0.056
Fe33Mn LSM α= -20°
-644.658
5.15E-05
0.5936
0.030
Fe33Mn LSM α= +20°
-727.762
4.46E-05
0.5148
0.041
Fe33Mn Rolled CW 50% Annealed Fe33Mn (1300C) CW 50% Annealed Fe33Mn (1300C) No CW
-687.245 -737.181 -675.052
4.42E-05 4.13E-05 3.80E-05
0.5094 0.4763 0.4378
0.056 0.042 0.060
Cast Fe33Mn
-698.345
3.02E-05
0.3486
0.094
Pure Iron
-272.138
4.38E-06
0.0508
0.016
Pure Zinc
-308.451
1.10E-06
0.0165
0.003
Processed Sample Type
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Microstructure
Table 4 D (um)
General Grain Shape
Rolled CW 50% 7595.16 ± 93.86 Elongated Cast 4687.14 ± 18.58 Blocky Annealed 1300°C 1032.79 ± 16.40 Elongated blocky No CW LSM α = +20° *45.56 ± 0.75 Dendritic band-like LSM α = -20° *37.42 ± 0.44 Dendritic band-like CW 50% & 32.01 ± 0.36 Equiaxed Annealed 1300°C LSM α = 0° *16.01 ± 0.41 Dendritic band-like *Dendrite band diameters rather than grain diameter
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Table 5. Effective strains for a variety of processed Cast-33%Mn samples LSM α (degrees) Effective Strain -20 3.45 0 2.85 +20 1.50 Cold-rolling (CW 50%) 0.80
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Michael Heiden a, Andrew Kustas a, Kevin Chaput a, Eric Nauman b, c, David Johnson a, Lia Stanciu a
a
School of Materials Science and Engineering, Purdue University, West Lafayette, IN 47907,
USA b
Weldon School of Biomedical Engineering, Purdue University, West Lafayette, IN 47907,
USA c
School of Mechanical Engineering, Purdue University, 585 Purdue Mall, West Lafayette, IN
47907, USA
This article has been accepted for publication and undergone full peer review but has not been through the copyediting, typesetting, pagination and proofreading process which may lead to differences between this version and the Version of Record. Please cite this article as an ‘Accepted Article’, doi: 10.1002/jbm.a.35220
Abstract Advancing the understanding of microstructural effects and deformation on the degradability of Fe-Mn bioresorbable alloys (specifically Fe-33%Mn), will help address the current problems associated with designing degradable fracture fixation implants for hard tissues. Potentiostatic polarization tests were conducted on a wide variety of metal samples in order to examine how different deformation processes affect the instantaneous rate of degradation of Fe-Mn alloys. Large-strain machining (LSM), a novel severe plastic deformation (SPD) technique was utilized during these experiments in order to modify the degradation properties of the proposed Fe-Mn alloy. It was discovered that Fe-33%Mn after LSM with a 0° rake angle (effective strain = 2.85) showed the most promising increase in degradation rate compared to as-cast, annealed, and additional deformation conditions (rolled and other LSM parameters) for the same alloy. The mechanisms for enhancement of the corrosion rate are discussed. Keywords: Iron-manganese, Biodegradable, In vitro, Microstructure, Corrosion 1. Introduction The pursuit for a degradable biomaterial that is able to mechanically support the healing process of bone fracture for a limited amount of time and gradually be replaced by surrounding tissue has been burdened with complications. One of the biggest issues from a clinical perspective deals with tailoring the degradation rates of biodegradable implants in a diverse range of physiological environments for complete tissue reconstruction.1-4 There is clinical relevance in identifying materials that can be employed to replace existing, permanent fracture fixation devices, such as staples, suture anchors, screws, or pins, and eliminate the need of a second surgical operation for device removal.5-12 It has been established by numerous authors that the ideal biodegradable material has the following characteristics: 1) biocompatible/non-
toxic/non inflammatory, 2) fosters natural tissue ingrowth and cell proliferation, 3) has mechanical properties that support the surrounding tissue during the full reconstruction process, 4) degrades at a rate that doesn’t generate local or systemic effects in the body, and 5) releases its degradation products in a way that can be safely eliminated from the body by the excretory system.2 Currently, no such bioresorbable material fulfills all these requirements for either hard or soft tissues. For orthopedic applications, bioresorbable metallic implants have been found to have several advantages over degradable polymers. Metals tend to have higher tensile, compressive, and yield strengths. They also have high fracture toughness, do not exhibit high stress relaxation which can discourage natural tissue ingrowth, and are not radiolucent.2, 3 Bone fracture fixation materials in particular require mechanical properties that can withstand a combination of factors, including mechanical fatigue, erosion due to device insertion, and degradation due to physiological pH, temperature, and other biological intrusions in the hard tissue environment.4 A toxicological risk assessment in the body is another crucial step in designing a safe implant and it is known that the human body has exposure limits for metallic elements.13 The most commonly investigated bioresorbable metallic materials due to biocompatible degradation products include magnesium alloys and iron alloys.14, 15 Yuen et al. reported a toxicological assessment for several metals, evaluating the critical exposure limits in the human body.13 For an average 60-kg adult, iron was found to have an exposure limit of 2.55 mg/day, manganese was 0.42 mg/day, and magnesium was specified in terms of a maximum daily intake of 250-400 mg/day.13 Magnesium alloys have implant potential because their elastic moduli closely mimics that of most human cortical bones, but pure magnesium’s high degradation rate is too fast for maintaining the needed mechanical support for the full period of hard tissue reconstruction.4
Furthermore, due to this high corrosion rate, magnesium causes subcutaneous hydrogen gas bubbles to form in the body that need to be released by invasive needles.4 Unlike magnesium, pure iron has exceptional mechanical properties for supporting the rebuilding of hard tissue, however it degrades too slowly.16-18 Additionally, with a higher initial stiffness, it has the potential to cause stress-shielding at the current rate of resorbability.16 Alloys of Fe-Mn have been proposed for quite some time due to their slightly faster corrosion rates and nonmagnetic properties above 20% Mn content, which aids in the process of examining implants using magnetic resonance imaging (MRI).17 However, previously-designed Fe-Mn bioresorbable implants have degradation rates that are still too slow to be beneficial for orthopedic implantation. Hermawan et al. processed Fe-Mn bulk forms from sintering Fe and Mn powders, and found that a more porous structure helped increase the corrosion rate of the alloy immersed in modified Hank’s Solution due to the high amount of surface area.18 Conversely, Chou et al. found that 3D printing an open-pore structure helped increase the corrosion rate, but noticeably decreased the implant’s mechanical properties.19 Factors that appear to influence the corrosion rate of metals in a liquid medium include the inherent electrode potential of the metal (affects reactivity), the surface state and passivity (more passive surfaces slow corrosion rate), and the amount of surface area in contact with the medium (affects amount of ion diffusion).16-19 In order to better understand the degradation kinetics of Fe-Mn alloys in osteogenic environments, the present authors endeavored to modify the degradation rate of a Fe-33%Mn alloy by employing a variety of deformation processes, one of which included severe plastic deformation (SPD) via orthogonal plane-strain machining, in order to refine the microstructure and examine how strain plays a role in degradation behavior. It should be noted that a manganese
composition of 33% was chosen in order to further investigate how the degradation results of this alloy compared with other Fe-Mn alloy compositions presented by other authors.15-18 The use of conventional plane-strain machining is an attractive manufacturing process for fundamental studies of microstructure-property relationships in a variety of metals. The high strains, strain-rates, and temperatures achievable during machining provide unique opportunities for microstructure refinement, even in the nanometer regime, through appropriate selection of processing parameters.20 The ability to control the microstructure can have substantial effects on material properties. Furthermore, once the optimum microstructure/corrosion relationship has been established, scale-up to continuous strip or wire is possible.21, 22 2. Materials and methods 2.1 Materials (casting) Small, irregular pieces (≤12mm [0.5in]) of pure Fe (Alfa Aesar, Ward Hill, MA, 99.9%) and pure Mn (Alfa Aesar, Ward Hill, MA, 99.9% Mn) were mixed to create a Fe–33%Mn alloy composition. This specific mixture was arc melted under a flowing argon (5%H2) atmosphere and re-melted at least three times to promote homogeneity. The Fe-Mn ingot was then further processed in a vacuum induction melter (VIM). The VIM was evacuated to a pressure of ~ 10-3 atm and then purged with argon three times to reduce the oxygen content in the furnace. After the final purge, the chamber was backfilled with argon to minimize vaporization of the manganese. The alloy was melted and poured into a cylindrical copper mold resting on a water cooled copper plate to create a 1” diameter ingot 3.5” in length. The cast Fe-33%Mn alloy cylinder was sectioned into 1cm x 0.5cm rectangular pieces ~0.2cm thick. The Fe-33%Mn samples underwent a wider variety of processing techniques, including: 1) heavy cold-rolling (reductions of 50%), and 2) large strain machining (LSM).
Select cast and cold-rolled Fe-33%Mn samples were then annealed at 1300°C for 1 hour to provide additional microstructure variation. Initial mechanical properties such as elastic modulus, hardness, and stiffness were also measured using nanoindentation (Hysitron TI-950 Triboindenter, Minneapolis, MN) with a Berkovich tip with a nominal radius of 150 nm. 2.2 Deformation Processing Plane-strain machining (Figure 1) employs a sharp, wedge tool with rake angle (α) set at a preset depth (ao) into the work piece (bulk), which is travelling at constant surface velocity (Vo). Material is removed from the bulk piece in the form of a chip-extrudate with deformed thickness (ac). Three variables (α, ao, Vo) control most of the deformation characteristics, creating very large shear strains (1-15), at high strain rates (104 s-1) and high temperatures during cutting.23 The severe deformation occurs within a very narrow zone (∆ ~ 50 µm wide), resulting in microstructure refinement of the extrudate through simultaneous accumulation of large strain and localized (adiabatic) heating.22-24 The narrow width of the deformation zone is most commonly approximated as a single shear plane with orientation angle ϕ from the surface of the work piece. Using this approximation, the shear strain can be calculated simply using an upperbound model as23, 25: (1)
=
() ( ) ( )
Where α is the tool rake angle and ϕ is the shear plane angle, which is determined from measuring ao and ac and through geometry: (2)
tan(φ) =
() ()
By definition, the von-Mises effective strain for plane-strain machining is expressed as:
(3)
=
√
Effective strain provides a convenient metric for the direct comparison of different deformation processes (e.g. rolling and machining) by reducing complex straining situations to a single component.25 Equations 1-3 23, 25 show that a variety of strains are possible in machining through the selection of different tool rake angles. In this study, rectangular chips of Fe-33%Mn were produced from cutting tools at three increasingly negative rake angles (+20°, 0°, -20°) to increase both stored strain energy and microstructural refinement. Effective strains for the three machining parameters range from “low” (~1.5) to “high” (~3.5).
2.3 Solution preparation and microstructural analysis All samples were ground to 600 grit, polished with 1µm alumina powder, rinsed with deionized water, rinsed with ethanol, and then air dried before immersion into osteogenic cell culture media for electrochemical experiments. Osteogenic cell culture media was prepared according to the procedure for hard tissue fluid outlined by Dickerson et al.26 For 300 mL total media: 1032.3 µL cell culture media (low glucose), 15 µL ascorbic acid, 2.7 µL Bglycerophosphate 3mM, 11.76 µL Dexamethasone, 30 mL FBS (fetal bovine serum), 3 mL Penicillin Streptomysin, and 265.938 mL D-MEM media.26 All Fe-33%Mn processed samples were mounted in Bakelite, mechanically ground and polished with 0.02µm silica, and then chemically etched using Nital solution (2% HNO3 in ethanol). Optical microscopy was performed to examine microstructures after each processing
step and scanning electron microscopy (SEM; FEI Philips XL-40) coupled with energydispersive x-ray spectroscopy (EDS) was used to examine the surface further while also acquiring an elemental analysis. X-ray diffraction (XRD) with Cr Kα radiation was also employed (Bruker D8 Discover X-Ray Diffractometer with GADDS (General Area Detector Diffraction System)) for characterization of microstructural phases present in the cast Fe-Mn alloy. Density was approximated by the Archimedes Principle. In determination of microstructure, the average grain size/diameter was measured for each processing technique applied by counting grains from three micrographs for each sampling, taken from different places on the specimens. The images were processed for grain size measurement using ImageJ software and measured according to standard ASTM E112.27, 28 2.4. Linear polarization test Rectangular samples of Fe-33%Mn for each processing technique were immersed in osteogenic media and tested in triplicate. Tests were conducted according to ASTM G 59L.29 Electrochemical impedance spectroscopy (EIS) tests were carried out under the conditions of 37±0.5 °C, with a pH of 7.4±0.1; using a potentiostat (SP-150 Biologic Science Instrument, Claix, France). The experimental setup was a standard three-electrode system with the sample as the working electrode, a saturated calomel electrode (SCE) as the reference electrode, and a platinum wire as the counter electrode. The results of the measurements were analyzed using EC-Lab v-10.19 BioLogic Science Instruments software. Each sample was submerged in 5mL of osteogenic media and the media was heated and maintained at 37±0.5 °C. Media was changed out for each test. The open circuit potential (OCP) was recorded after 55 minutes of immersion to maintain steady-state conditions of stable free corrosion potential. The sweep rate of the polarization curve measurement was set to 0.1667mVs-1 and the potential range was -150 to
+150 mV vs. OCP. Linear polarization curves were recorded until the potentials reached the ends of the designated range. Ecorr and Icorr were recorded for each polarization test. The generated Tafel plots based on the cathodic and anodic curves were fitted linearly to measure the intersection of the corrosion potential (Ecorr) and total corrosion current (Icorr). The corrosion current density (icorr) for each sample was then calculated by dividing Icorr by the exposed surface area of each specimen. The analyzed icorr was further converted into the corrosion rate (mm/year) using29, 30: (4)
Corrosion Rate (mm/year) = 3.27x103
∗ !
where icorr is the corrosion current density (A/cm2), D is the sample standard density (g/cm^3), and K is a constant; where 3272 was used to obtain units in (mm/yr), EW is the equivalent weight according to the percent of alloying elements (Fe33Mn: 27.77 g*eq-1). Finally, polarization curves were plotted together for comparison of every type of processing technique imparted upon the Fe-33%Mn samples, while average corrosion rates were calculated and then plotted in a bar graph with confidence intervals. 3. Results and Discussion 3.1. Mechanical properties of the Fe-Mn alloy Data from EDS shown in Table 1 indicate the average compositions of the as-cast alloy were homogeneous throughout the samples near the desired compositions with some microsegregation observed, typical of common solidification processes. Additionally, nanoindentation results are summarized in Table 2. After casting the Fe-Mn alloy in the VIM, the microstructure displayed a columnar grain morphology with grains growing radially inward towards the centerline of the ingots. Figure 2
shows the dendritic structure to be relatively coarse. When examining the binary phase diagram for Fe-Mn, it can be found that the partition coefficient (k) is less than one, indicating the interdendritic liquid during solidification is expected to be richer in manganese while the dendrites will be depleted due to solute rejection.32 Figure 2 indeed confirms this expectation, i.e. regions between dendritic arms are higher in manganese than the dendrites themselves. The electrochemically degraded Fe-33%Mn cast sample shown in Figure 3 reveals the media’s preferential corrosion on the manganese rich, inter-dendritic regions. It was observed that the elastic modulus and density of the Fe-33%Mn alloy (Table 1) were lower than that of the pure iron19, but higher than that of 316L stainless steel.33 An elastic modulus closer to that of bone is ideal to improve implant interface/tissue interactions, decreasing chance of failure. However, the results from Table 3 show that besides looking into simply minimizing the material’s Young’s Modulus, it is also important to characterize and at least qualitatively assess the interrelationship between deformation, composition (percent Mn), and degradation rate. For example, Hermawan et al. found both UTS and YS to decrease with increasing manganese content in four porous Fe-Mn alloys, while elongation at fracture increased.30 For a more complete analysis of Fe-Mn alloys being employed as a clinically viable bioresorbable implant material, it would be beneficial for a wider range of Fe-Mn alloy compositions to have their degradation rates tested while measuring the changes in all of their mechanical properties. 3.2. Influence of microstructural evolution and processing method on corrosion rate Linear polarization curves obtained for the Fe-33%Mn samples that were tested in the osteogenic media are shown in Figure 4. Additionally, the legend lists the curves in order from higher obtained Ecorr values to lower Ecorr values for enhanced comparisons since there is a
large grouping of polarization curves around Ecorr = -0.7V. Among the Fe-33%Mn alloys, the highest electrode potential was found to be the samples that underwent LSM at α= -20°, which interestingly underwent the highest amount of effective strain. In contrast, the lowest Ecorr was the annealed sample (1300°C) after CW50%. The BioLogic software’s tafel line fit program was used to measure icorr values from the polarization curves, which was then utilized to calculate the corresponding corrosion rates for each type of processed sample. Figure 5 provides the resulting corrosion bar graphs after analysis with standard error bars. Table 3 shows the actual averaged Ecorr, icorr, corrosion rate (CR), and standard error for CR values for all the samples. All processed samples were etched using 2% Nital and then optically imaged in order to examine the microstructural evolution shown in Figure 6. The cast structures were all dendritic in nature and the cold-rolled samples show an elongated grain structure along the rolling direction, which is typical of a rolled microstructure. Upon annealing at 1300°C, the rolled sample exhibits a relatively equiaxed recrystallized structure with grains on the order of ~50µm in diameter. The LSM samples all showed heavily sheared microstructures. Analysis on an etched, cast Fe-33%Mn sample using EDS show the initial dendritic structures are richer in manganese compared to the cooled inter-dendritic liquid. Figure 2 displays the compositions obtained for the thick dendrites and the thin interdendritic regions. Linear polarization scans showed interesting variations in Ecorr and icorr for different processing techniques. The shifts in polarization curves seen in Figures 4-5 demonstrate that both Ecorr and icorr increased simply by alloying manganese with iron, comparable to what Hermawan et al found for sintered Fe-Mn structures.30 Pure iron and zinc were used for
comparison purposes only and clearly had too low of Ecorr, icorr, and degradation rates for bioresorbable fracture fixation applications. It was found that cold rolling Fe-33%Mn to 50% increased the corrosion rate compared to the cast structure. Furthermore, annealing Fe-33%Mn at 1300 for an hour after cold rolling dampened the corrosion rate. Interestingly, annealing the sample with no cold work increased its corrosion rate slightly. This phenomenon potentially resulted from homogenization of the microstructure, effectively reducing microsegregation, e.g. more diffusion of solute atoms from solute-rich regions to solute-depleted regions. All LSM conditions increased the degradation rate of the cast structure substantially, but LSM at α=0° (ε~2.85) showed the most improvement in corrosion rate. When examining the microstructures of the machined chips (Figure 6 e-g), the dendritic structure has clearly undergone large deformation under each condition. Unlike the rolled microstructures with dendrites elongated along the direction of rolling, chip microstructures show the dendrites sheared and oriented in a manner mimicking the angle of the shear deformation plane. Interestingly, degradation rates do not continuously increase with increasing deformation strain. This indicates that larger deformation strains do not necessarily lead to higher corrosion rates as assumed in a linear correlation. The degradation rate/grain size relationship may be more of a bimodal distribution in nature as suggested by Moravej et al’s data.16 Moravej et al. showed that annealed Fe with an average grain size of 40µm had a higher degradation rate than Fe with nanocrystalline-sized grains, but a lower degradation rate than electroformed Fe with 2-8µm grains.16 It has been established by others that decreasing grain size increases grain boundary area and thus creates a high internal energy in a material.31 Since corrosion tends to target grain boundaries first, it is assumed that decreasing grain-size will promote higher corrosion rates. This effect was partially
observed in the current study. The results from Table 4 indicate that grain size appears to influence the degradation behavior more than induced plastic strain. However, it should be noted that it is difficult to discern the grain size of the LSM samples since etching the surfaces revealed dendritic and interdendritic bands rather than grain boundaries, as expected in a homogenized microstructure. This deformed microstructure is analogous to banded structures found after cold rolling of a cast part.34 This deformed structure makes it difficult to discern the grain size for these LSM samples, however it is possible to compare the LSM dendritic band thicknesses (Figure 7) to the original dendritic spacing (Figure 8). It can be seen that the cast structure’s average primary dendrite arm thicknesses (15-40 um) decrease to 2-10um after LSM at α=0°. These measurements along with the degradation results suggest that preferential attack on the interdendritic areas occurs during degradation of these alloys, leading to an overall increase in corrosion rate due to the increase of interdendritic surface area per unit volume with decreasing band thickness. It can also be concluded that the degradation rate is controllable, to some extent, by the amount of deformation strain induced in these classes of biodegradable materials. It appears that increasing the amount of deformation strain in the material can increase the electrode potential of the material, yet the degradation rate increases only up to a certain point in terms of induced strain. Beyond this point, increasing deformation strain does not appear to enhance the corrosion rate. It was observed that too large of a strain (LSM α = -20°) had adverse effects on increasing the degradation rate (Table 5). However, it appears that simply reducing grain size doesn’t necessarily lead to a lower degradation rate, as evidenced by the low corrosion rates of the smallgrained, annealed structure. Based on these results, there appears to be an interrelated grainsize/plastic strain relationship on material degradation rate. The precise role of each mechanism
is currently unknown; however it has been well-documented that continually increasing deformation strain can lead to continuous recrystallization, thereby developing a refined microstructure.35 Therefore, at the highest deformation strain, it is possible some amount of continuous recrystallization occurred during deformation, which would decrease the total stored energy and consequently, the degradation rate relative to the lower strain conditions. From these results, there appears to be a significant influence of deformation strain on the degradation rate of the material. Moreover, it appears high effective strains exhibit relatively high degradation rates. Comparatively, the Fe-33%Mn LSM sample at α= +20° had an effective strain roughly double that of the cold rolled sample that underwent 50% CW, but its corrosion rate was only increased by 0.0054mm/year.
5. Conclusion In this study, the in vitro degradation rates of potential bioresorbable materials were tailored by varying the processing routes for a Fe-33%Mn alloy. The novel LSM technique at a rake angle of 0° showed a 140% increase in corrosion rate for Fe-33%Mn compared to a cast structure of the same alloy. The authors suggest that in order to increase the degradation rate further for Fe-Mn alloys, 1) tailored shear-based deformation processing should be employed to increase the kinetic effects of corrosion up to a critical value, and 2) the surface area of the implants should be increased to allow for more diffusion of the osteogenic media into the Fe-Mn bulk alloy, which would also provide more adhesion and ingrowth of the hard tissue environment. To better tailor the properties of these potential bioresorbable Fe-Mn hard tissue implants, it is essential to balance the mechanical properties needed for orthopedic support with the degradation rate of the material. It appears that in order to achieve higher degradation rates,
refinement of the grains is needed up to a critical value, while also being careful to avoid increasing the material’s elastic modulus and strength too much, thereby leading to poorer interface/tissue bonding. Other authors’ findings and the results found above seem to indicate that an open-porous, Fe-Mn fine-grained structure deformed by shear might be more beneficial for achieving the clinically desired rate of bioresorbability for hard tissue implants. In either case, the study garners additional evidence that Fe-Mn alloys have potential for being highly degradable for orthopedic applications. Acknowledgements The authors would like to thank the help of Julia Aspaugh for aiding in osteogenic media preparation, Kaitlyn Jarry for aiding in the electrochemical experiments, Samantha Lawrence for aiding in nanoindentation, and Kristin Paulson for XRD training. Guidance covering the nature of material corrosion was gratefully provided by Dr. Robert Spitzer of Purdue University. Furthermore, the authors would like to thank Dr. Srinivasan Chandrasekar of Purdue University for use of his machining equipment used towards LSM the Fe-Mn alloy. The authors acknowledge the financial assistance from Fort Wayne Metals, IN in supporting this work. References 1. Peuster M, Wohlsein P, Brügmann M, Ehlerding M, Seidler K, Fink C, et al. A novel approach to temporary stenting: degradable cardiovascular stents produced from corrodible metal-results 6-18 months after implantation into New Zealand white rabbits. Heart (British Cardiac Society) 2001; 86:563–9. 2. Hofmann GO, Wagner FD. New implant designs for bioresorbable devices in orthopaedic surgery. Clinical Materials 1993; 14:207–15.
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15. Schaffer JE, Nauman E a, Stanciu L a. Cold drawn bioabsorbable ferrous and ferrous composite wires: An evaluation of in vitro vascular cytocompatibility. Acta Biomaterialia 2012:1–11. 16. Moravej M, Mantovani D. Biodegradable metals for cardiovascular stent application: interests and new opportunities. International Journal of Molecular Sciences 2011; 12:4250–70. 17. Hermawan H, Purnama A, Dube D, Couet J, Mantovani D. Fe-Mn alloys for metallic biodegradable stents: degradation and cell viability studies. Acta Biomaterialia 2010; 6:1852–60. 18. Hermawan H, Alamdari H, Mantovani D, Dubé D. Iron–manganese: new class of metallic degradable biomaterials prepared by powder metallurgy. Powder Metallurgy 2008; 51:38–45. 19. Chou D-T, Wells D, Hong D, Lee B, Kuhn H, Kumta PN. Novel processing of iron-manganese alloybased biomaterials by inkjet 3-D printing. Acta Biomaterialia 2013. 20. Swaminathan S, Ravi Shankar M, Rao BC, Compton WD, Chandrasekar S, King AH, et al. Severe plastic deformation (SPD) and nanostructured materials by machining. Journal of Materials Science 2007; 42: 1529–41. 21. Efe M, Moscoso W, Trumble KP, Dale Compton W, Chandrasekar S. Mechanics of large strain extrusion machining and application to deformation processing of magnesium alloys. Acta Materialia 2012; 60: 2031–42. 22. Shaw, Milton. Metal Cutting Principles. 1. New York: Oxford Series on Advanced Manufacturing, 1996. 18-46. Print. 23. Sagapuram D, Efe M, Moscoso W, Chandrasekar S, Trumble KP. Controlling texture in magnesium alloy sheet by shear-based edeformation processing. Acta Materialia 2013. 24. Guo Y, Efe M, Moscoso W, Sagapuram D, Trumble KP, Chandrasekar S. Deformation field in largestrain extrusion machining and implications for deformation processing. Scripta Materialia 2012; 66:235–8. 25. Hosford, William, and Robert Caddell. Metal Forming Mechanics and Metallurgy. 2. Upper Saddle River, NJ: PTR Prentice Hall, 1993. 29-48. Print.
26. Dickerson D a, Misk TN, Van Sickle DC, Breur GJ, Nauman E a. In vitro and in vivo evaluation of orthopedic interface repair using a tissue scaffold with a continuous hard tissue-soft tissue transition. Journal of Orthopaedic Surgery and Research 2013; 8:18. 27. W.S. Rasband, ImageJ, U.S. National Institutes of Health, Bethesda, ML, USA, 2014 28. ASTM E112, Annual Book of ASTM Standards, American Society of Testing and Materials, West Conshohocken, PA, 2013, Standard Test Methods for Determining Average Grain Size, doi:10.1520/E0112-96R04E02. 29. ASTM Standard G59-97: Standard Test Method for Conducting Potentiodynamic Polarization Resistance Measurements, West Conshohocken, PA: ASTM International, 2009 30. Hermawan H, Dubé D, Mantovani D. Degradable metallic biomaterials: design and development of Fe-Mn alloys for stents. Journal of Biomedical Materials Research Part A 2010; 93:1–11. 31. Witte F, Hort N, Vogt C, Cohen S, Kainer KU, Willumeit R, et al. Degradable biomaterials based on magnesium corrosion. Current Opinion in Solid State and Materials Science 2008; 12:63–72. 32. Dantzig, J.A, and M Rappaz. Solidification. 1. Switzerland: EPFL, 2009. 395-425. Print. 33. Schaffer JE, Nauman E a., Stanciu L a. Cold-Drawn Bioabsorbable Ferrous and Ferrous Composite Wires: An Evaluation of Mechanical Strength and Fatigue Durability. Metallurgical and Materials Transactions B 2012; 43:984–94. 34. Grange, R. A. (1971). Effect of Microstructural Banding in Steel, 2(February), 417–426. 35. N Tsuji, Y Ito, Y Saito, Y Minamino, Strength and ductility of ultrafine grained aluminum and iron produced by ARB and annealing, Scripta Materialia, 2002; 47:893-899
Figure 1. Plain-strain machining diagram. Wedge tool uses a rake angle (α) set at a preset depth (ao) into the work piece (bulk), which is travelling at constant surface velocity (Vo). Chipextrudate has a deformed thickness (ac). ɸ is the angle from the shear plane to the horizontal preset depth. Figure 2. EDS “single spot” analysis on dendrite arm (a) and spacing between arms (b-c) of the cast Fe-33%Mn alloy at 8000X after Nital etching.
Figure 3. Surface of the cast Fe-33%Mn sample at 10X after experiencing electrochemical polarization.
Figure 4. Polarization curves for pure iron, pure zinc, and a variety of processed Fe-33%Mn alloys after being immersed in osteogenic media.
Figure 5. Average electrochemical corrosion rates of all tested samples while immersed in osteogenic media. Error bars are in terms of standard error.
Figure 6. Etched microstructures of Fe-33%Mn: (a) cast, (b) rolled CW50%, (c) annealed at 1300C (no CW), (d) annealed at 1300C (CW50%), (e) LSM α = +20, (f) LSM α = 0, (g) LSM α = -20. Figure 7. Dendritic (white) and interdendritic (black) band structures for the LSM α = -20. Figure 8. Primary and secondary dendrite arms of the cast sample, oriented from the edge of the circular ingot (top) to the center (bottom).
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Table 1. Chemical compositions of the cast Fe-Mn alloy Element Concentration Wt % Mn 33.08 ± 0.85 Fe 67.10 ± 0.97
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Table 2. Mechanical properties of cast Fe-33%Mn alloy Elastic Modulus (GPa)
201.47 ± 49.13
Density (g/cm^2)
7.64 ± 0.13
Vickers Microhardness (GPa)
6.64 ± 0.53
Stiffness (µN/nm)
182.54 ± 49.13
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Table 3. Electrochemical corrosion rates of samples immersed in osteogenic media
Fe33Mn LSM α= 0°
Ecorr (mV) -685.469
icorr (A/cm^2) 7.25E-05
Average CR (mmpy) 0.8354
Standard Error 0.056
Fe33Mn LSM α= -20°
-644.658
5.15E-05
0.5936
0.030
Fe33Mn LSM α= +20°
-727.762
4.46E-05
0.5148
0.041
Fe33Mn Rolled CW 50% Annealed Fe33Mn (1300C) CW 50% Annealed Fe33Mn (1300C) No CW
-687.245 -737.181 -675.052
4.42E-05 4.13E-05 3.80E-05
0.5094 0.4763 0.4378
0.056 0.042 0.060
Cast Fe33Mn
-698.345
3.02E-05
0.3486
0.094
Pure Iron
-272.138
4.38E-06
0.0508
0.016
Pure Zinc
-308.451
1.10E-06
0.0165
0.003
Processed Sample Type
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Microstructure
Table 4 D (um)
General Grain Shape
Rolled CW 50% 7595.16 ± 93.86 Elongated Cast 4687.14 ± 18.58 Blocky Annealed 1300°C 1032.79 ± 16.40 Elongated blocky No CW LSM α = +20° *45.56 ± 0.75 Dendritic band-like LSM α = -20° *37.42 ± 0.44 Dendritic band-like CW 50% & 32.01 ± 0.36 Equiaxed Annealed 1300°C LSM α = 0° *16.01 ± 0.41 Dendritic band-like *Dendrite band diameters rather than grain diameter
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Table 5. Effective strains for a variety of processed Cast-33%Mn samples LSM α (degrees) Effective Strain -20 3.45 0 2.85 +20 1.50 Cold-rolling (CW 50%) 0.80
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70x48mm (600 x 600 DPI)
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50x16mm (600 x 600 DPI)
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70x48mm (300 x 300 DPI)
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81x52mm (600 x 600 DPI)
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