A review on manufacturing processes of cobalt-chromium alloy implants and its impact on corrosion resistance and biocompatibility.

Cobalt-Chromium (CoCr) alloys are currently used for various cardiovascular, orthopedic, fracture fixation, and dental implants. A variety of processes such as casting, forging, wrought processing, hot isostatic pressing, metal injection molding, milling, selective laser melting, and electron beam melting are used in the manufacture of CoCr alloy implants. The microstructure and precipitates (carbides, nitrides, carbonitrides, and intermetallic compounds) formed within the alloy are primarily determined by the type of manufacturing process employed. Although the effects of microstructure and precipitates on the physical and mechanical properties of CoCr alloys are well reviewed and documented in the literature, the effects on corrosion resistance and biocompatibility are not comprehensively reviewed. This article reviews the various processes used to manufacture CoCr alloy implants and discusses the effects of manufacturing processes on corrosion resistance and biocompatibility. This review concludes that the microstructure and precipitates formed in the alloy are unique to the manufacturing process employed and have a significant impact on the corrosion resistance and biocompatibility of CoCr alloys. Additionally, a historical and scientific overview of corrosion and biocompatibility for metallic implants is included in this review. Specifically, the failure of CoCr alloys when used in metal-on-metal bearing surfaces of total hip replacements is highlighted. It is recommended that the type of implant/application (orthopedic, dental, cardiovascular, etc.) should be the first and foremost factor to be considered when selecting biomaterials for medical device development.

Triphasic and quadriphasic waveforms are superior to biphasic waveforms for synchronized beating of cardiomyocytes.

BACKGROUND AND PURPOSE: Within pacemaker research few attempts have been made to find an optimal waveform phase sequence that synchronizes beating of cardiomyocytes at an electrode. Multielectrode arrays (MEAs) offer electrophysiological screening of cardiomyocytes serving as a system for preliminary screening of pacing waveform design. MATERIALS AND METHODS: The HL-1 cell line was cultured in MEAs until confluence and stimulated with biphasic, triphasic, and quadriphasic waveforms. The amplitudes required for synchronized beating of the cells were determined. RESULTS: Triphasic and quadriphasic waveforms were more efficient in eliciting synchronized beating of the HL-1 cells compared with the biphasic waveform because it allows significant reductions in synchronizing voltage amplitudes and reductions in supplied stimulus. CONCLUSION: The MEA system allows for a straightforward manner to investigate effects of waveform design on synchronized beating in cardiomyocytes in vitro. Increased number of phase changes in a pacing waveform seems to be the major reason for the reduction in synchronizing amplitudes.

Surface finishing of Nitinol for implantable medical devices: A review.

Nitinol (NiTi), a nickel-titanium alloy, has been used for various cardiovascular, orthopedic, fracture fixation, and orthodontic devices. As with most other metallic biomaterials, the corrosion resistance and biocompatibility of NiTi are primarily determined by the properties of the surface oxide layer such as thickness, chemical composition, structure, uniformity, and stability. Currently, a number of finishing methods are used to improve the properties of surface oxide of NiTi with an ultimate goal to produce a defect-free, impurity-free, thin homogeneous oxide layer that is stable and composed of only titanium dioxide (TiO2 ) with negligible amount of Ni species. This review discusses the effects of various surface finishing methods such as mechanical polishing, electropolishing, magnetoelectropolishing, heat treatments at different temperatures, passivation, chemical etching, boiling in water, hydrogen peroxide treatment, and sterilization techniques (steam autoclave, ethylene oxide, dry heat, peracetic acid, and plasma-based treatments) on the properties of a surface oxide layer and how it impacts the corrosion resistance of NiTi. Considering the findings of the literature review, a checklist has been provided to assist with choosing finishing/sterilization methods and relevant rationale and recommendations to consider when selecting a surface finishing process for NiTi used in implantable medical devices.

A comprehensive review of biological and materials properties of Tantalum and its alloys.

Tantalum (Ta) and its alloys have been used for various cardiovascular, orthopedic, fracture fixation, dental, and spinal fusion implants. This review evaluates the biological and material properties of Ta and its alloys. Specifically, the biological properties including hemocompatibility and osseointegration, and material properties including radiopacity, MRI compatibility, corrosion resistance, surface characteristics, semiconductivity, and mechanical properties are covered. This review highlights how the material properties of Ta and its alloys contribute to its excellent biological properties for use in implants and medical devices.

Influence of strain on the corrosion of magnesium alloys and zinc in physiological environments.

During implantation load-bearing devices experience stress that may influence its mechanical and corrosion profile and potentially lead to premature rupture. The susceptibility to stress corrosion cracking (SCC) of the Mg-Al alloy AZ61 and Zn was studied in simulated body fluid (m-SBF) and whole blood by slow strain rate (SSR) testing in combination with electrochemical impedance spectroscopy (EIS) and further ex situ analysis including scanning electron microscopy (SEM) and Fourier transform infrared spectroscopy. AZ61 was found to be highly susceptible to SCC. EIS analysis show that although the majority of cracking occurred during the apparent plastic straining, cracking initiation occurs already in the elastic region at ∼50% of the ultimate tensile strength (UTS). Shifts in EIS phase angle and open circuit potential can be used to detect the onset of SCC. Zinc demonstrated a highly ductile behavior with limited susceptibility to SCC. No significant decrease in UTS was observed in m-SBF but a decrease in time to failure by ∼25% compared to reference samples indicates some effect on the mechanical properties during the ductile straining. The formation of micro cracks, ∼10µm deep, was indicated by the EIS analysis and later confirmed by ex situ SEM. The results of SSR analysis of zinc in whole blood showed a reduced effect compared to m-SBF and no cracks were detected. It appears that formation of an organic surface layer protects the corroding surface from cracking. These results highlight the importance of considering the effect of biological species on the degradation of implants in the clinical situation. STATEMENT OF SIGNIFICANCE: Strain may deteriorate the corrosion properties of metallic implants drastically. We study the influence of load on the corrosion properties of a magnesium alloy and zinc by a combination of electrochemical impedance spectroscopy (EIS) and slow strain rate analysis. This combination of techniques has previously not been used for studying degradation in physiological relevant electrolytes. EIS provide valuable information on the initial formation of cracks, detecting crack nucleation before feasible in slow strain rate analysis. This sensitivity of EIS shows the potential for electrochemical methods to be used for in situ monitoring crack formation of implants in more applied studies.

The influence of buffer system and biological fluids on the degradation of magnesium.

The influence of frequently used buffer system 4-(2-hydroxyethyl)-1-piperazineethanesulfonic acid (HEPES) compared to CO2 /HCO3- on the corrosion of magnesium is investigated. Samples were immersed in simulated body fluid (m-SBF) while monitored by electrochemical impedance spectroscopy (EIS) for up to 30 days. In CO2 /HCO3- the initial corrosion rate was 0.11 mm yr-1 . An inner protective layer of magnesium oxide was formed within the first 30 min exposure and subsequently covered by an outer layer of apatite within 24 h. The corrosion mechanism thereafter is best described as passive pitting with a porosity of ∼10%. Using HEPES as buffer agent increased the corrosion rate to 3.37 mm yr-1 . Cross sectional microscopy show a porous outer corrosion layer allowing rapid diffusion of aggressive ions through the film. Here the EIS results are best described by an active pitting model with an inner layer 5 to 10 times less protective compared to the inner layer formed without HEPES. Further the suitability of human whole blood and plasma as in vitro models for Mg degradation was evaluated. Mg corrosion caused coagulation after 24 h in both biological fluids. The corrosion during the first 24 h is similar to the corrosion in m-SBF with HEPES. © 2016 Wiley Periodicals, Inc. J Biomed Mater Res Part B: Appl Biomater, 105B: 1490-1502, 2017.