human body fluid and the metal ions released leading to inflammation in the tissue surrounding the implants, decreasing the pH and then the toxicity. Corrosion ...
Republic of Iraq Ministry of Higher Education and Scientific Research University of Technology Materials Engineering Department
Corrosion Behavior of Some Implant Alloys in Simulated Human Body Environment A thesis Submitted To
The Department of Materials engineering, University of Technology in Partial Fulfillment of the Requirements for the Degree of Master in Science in Materials Engineering By
Zina Noori Abdulhameed B.Sc. In Materials Engineering -2004 Under Supervision Prof. Dr. Khahtan K. Al-Khazraji
Assist. Prof. Dr. Rana A. Anaee
1432 A.H
2011 A.D 1
Abstract The subject of this work had involved the investigation of the polarization behaviour of two implants that included SS 316L and Co – Cr – Mo alloy in simulated human body fluid. The polarization behaviour had included the study in the Tafel region to predict the cathodic and anodic reactions, where the cathodic reaction represents the reduction of oxygen to produce hydroxyl ions (OHˉ) because of the neutralization of human body environment at pH=7.4, while the anodic reaction represents the dissolution of metals from alloys in human body fluid and the metal ions released leading to inflammation in the tissue surrounding the implants, decreasing the pH and then the toxicity. Corrosion parameters of polarization behaviour had included corrosion potentials (Ecorr), corrosion current densities (icorr), and cathodic and anodic Tafel slopes (bc & ba). These data were necessary to calculate the polarization resistance (Rp) and corrosion rate (CR) to compare between the behaviour of the two selected alloys at pH=7.4 and Temp. =37oC. Co – Cr – Mo alloy showed higher resistance and lower corrosion rate than SS 316L in human body fluid without any addition due to higher charge transfer resistance and lower capacitance which means thicker passive films on the Co-Cr-Mo alloy than SS 316L, in addition to the data of Tafel slope interpreting the rate – determining step which is due to proton discharge step (bc ≈ 120 mV.decade-1). Microstructure had enhanced this result, it shown more pits on the surface of SS 316L compared with Co-Cr-Mo alloy. The effect of three anti-inflammatory drugs on polarization behaviour of the two implants was studied. The anti-inflammatory drugs had included Aspirin, Paracetamol and Mefenamic acid with three concentrations of each drug (0.00303, 0.00606 and 0.01212), (0.0086, 0.0172 and 0.0344), and (0.00111, 0.00156 and 0.00201) g/300 mL respectively and according to response of different patients. Generally, results of corrosion parameters and the relationship between 2
polarization resistance and concentrations of drugs had shown the inhibitive action of these drugs in addition to the therapeutic agents of first choice for the treatment of inflammation, pain, and fever. This behaviour may have been due to the organometallic complexes, produced between the metal ions released and the drug molecules that cover the surface of implant and acts as protective layers for SS 316 L, while these drugs act as corrosive environment for Co-Cr-Mo alloy. This work had focused on the abnormal concentrations of uric acid, which is not a medical condition, but is associated with a variety of medical conditions. Excess serum accumulation of uric acid can lead to a type of arthritis, known as gout.0. 7 g/L uric acid (a maximum level in male) was added to study the effect of combination of uric acid with three anti-inflammatory drugs in human body fluid after surgical implant. The results had showed that the drugs act as inhibitors for SS 316 L with three drugs and for Co-Cr-Mo alloy only the presence of Aspirin, while Paracetamol and Mefenamic acid act as corrosion media. In mixture of 1.2 mg/L uric acid and drugs in simulated HBF, all drugs had behaved as inhabitants for both SS 316L and Co-Cr-Mo alloy.
3
List of Symbols & Abbreviations
Symbol
Description
As.
Aspirin
Unit -
ba
Anodic Tafel slope
V.decade-1
bc
Cathodic Tafel slope
V.decade-1
CR
Corrosion rate
mm/y
Conc.
Concentration of Drug
g/ml
dL
Deci l = 0.1 L
0.1 L
E
Potential
V
Ecorr
Corrosion potential
V
Eeq
Equilibrium potential
V
Eoc
Open circuit potential
V
e
equivalent weight
g
F
Faraday constant (96500)
C.mol-1
ΔG
Gibbs free energy change
kJ.mol-1
ΔH
Enthalpy change
kJ.mol-1
HBF
Human body fluid
icorr
Corrosion current density
A.cm-2
I
Current
A
i0
Exchange current density
A.cm-2
Me.
Mefenamic acid
-
M
Atomic weight
gm/mole
4
n
charge number
NSAIDS
Non-Steroidal Anti-Inflammatory Drugs
-
Pa.
Paracetamol
-
Rp
Polarization resistance
Ω.cm-2
SS 316 L
low carbon Stainless steel 316
η
Over potential
mV
ρ
Density
g/cm3
θ
Angle
Degree
U.A
Uric acid
5
Contents
Chapter One: General Introduction and Literature Survey 1-1 1-2 1-3
An Introduction The Aim of the Presence Work Literature Survey
1 2 3
Chapter Two: Theoretical Part 2-1 2-2 2-3 2-4 2-4-1 2-4-2 2-4-2-1 2-4-2-2 2-4-2-3 2-5 2-5-1 2-5-2 2-5-3 2-5-4 2-5-4-1 2-5-4-2 2-5-4-3 2-5-4-4 2-6 2-6-1 2-6-2 2-6-3
General Concepts Related to Corrosion Electrochemical Corrosion History of Implants Requirements of Implants Mechanical Properties Non- mechanical Requirements High Corrosion Resistance Biocompatibility Osseointegration Types of Implants Polymer Ceramic Composites Metallic Implants Titanium and its Alloys Noble Metals Cobalt – based Alloys Stainless steel Types of Corrosion in Implants Pitting Corrosion Crevice Corrosion Fretting Corrosion
2-6-4
Galvanic Corrosion 6
15 16 19 23 24 24 24 25 26 26 26 27 28 28 30 30 31 33 36 36 36 37 37
2-6-5 2-7 2-8 2-9 2-10 2-11 2-12 2-13 2-13-1 2-13-2 2-13-3 2-14
Corrosion Fatigue Corrosion Parameters Calculation of Corrosion Rate from Corrosion Current Why Metal Corrodes in Human Body? Drug Inflammation Non Steroidal Anti-inflammatory Drugs Some Anti-inflammatory Drugs Aspirin Paracetamol Mefenamic Acid Case study of Arthritis ( High Uric Acid)
38 38 40 41 44 45 46 47 47 48 49 49
Chapter Three: Experimental Part 3-1 3-1-1 3-1-2 3-2 3-3 3-3-1 3-3-2 3-3-3 3-3-4 3-4 3-4-1 3-4-2 3-5
Materials And Chemicals Materials Chemical Solution Adjustment of pH and Temperature Specimen Preparation Cutting Mounting Grinding and Polishing Etching Corrosion Test Electrochemical Cell Corrosion Instrument Metallographic Examination
7
51 51 52 53 54 54 54 55 56 56 56 58 60
Chapter Four: Results and Discussion Subject 4-1 4-2 4-2-1 4-2-2 4-3 4-3-1 4-3-2 4-4 4-4-1 4-4-2 4-5
page Corrosion Behaviour of Implants in Human Simulated Body Fluid
Effect of Anti-inflammatory Drugs on the corrosion behaviour Stainless Steel 316L Co-Cr-Mo Alloy Effect of Arthritis on the Corrosion Behaviour Behaviour of SS 316L with the Arthritis Disease Behaviour of Co-Cr-Mo Alloy with Arthritis Disease High Uric Acid up to 1.2 g/L Behaviour of SS 316L at 1.2 g/L Uric Acid Behaviour of Co-Cr-Mo Alloy at 1.2 g/L Uric Acid Microstructure of Implants
61 65 66 71 76 76 80 83 84 88 91
Chapter Five: Conclusions and Suggestion for Further Studies 5-1 5-2
Conclusions Suggestions for Further Studies
96 96
References 98 Dictionary of Medical Words
8
Chapter One 1-1 An Introduction Corrosion is one of the major processes that cause problems when metals and alloys are used as implants in the human body Kruger, 1979. It is realized from this definition that although biomaterials deteriorate over time, metals and alloys are the only ones highly susceptible, while polymers also suffer degradation. The metallic biomaterials in bioliquids are a function of many parameters, related to surface preparation and environment specific composition including the special influence of chlorine or fluorine anion or the effect of organic compounds. Corrosion of implants in the aqueous medium of body fluids takes place via electrochemical reactions Shreir, 1994 and it is necessary to appreciate and understand the electrochemical principles that are most relevant to the corrosion processes. The electrochemical reactions that occur on the surface of the surgically implanted alloys are identical to those observed during exposure to seawater (namely, aerated sodium chloride). The metallic components of the alloy are oxidized to their ionic forms and the dissolved oxygen is reduced to hydroxyl ions [1]. Corrosion is the unwanted chemical reaction of a metal with its environment, resulting in its continued degradation to oxides, hydroxides or other compounds. Tissue fluid in the human body contains water, dissolved oxygen, proteins, and various ions such as chloride and hydroxide. As a result, the human body presents a very aggressive environment for metals used for implantation [2]. The use of biomaterials did not become practical until the advent of an aseptic surgical technique developed by Dr. J. Lister in the 1860s. Earlier surgical procedures involved biomaterials or not, were generally unsuccessful as 9
a result of infection. Problems of infection tend to be exacerbated in the presence of biomaterials, since the implant can provide a region inaccessible to the body’s immunologically competent cells. The earliest successful implants, as well as a large fraction of modern ones, were in the skeletal system. Bone plates were introduced in the early 1900s to aid in the fixation of long bone fractures. Many of these early plates broke as a result of unsophisticated mechanical design, they were too thin and had stress concentrating corners. Also, materials such as vanadium steel, which was chosen for its good mechanical properties, corroded rapidly in the body and caused adverse effects on the healing processes. Better designs and materials soon followed. Following the introduction of stainless steels and cobalt chromium alloys in the 1930s, greater success was achieved in fracture fixation, and the first joint replacement surgeries were performed. As for polymers, it was found that warplane pilots in World War II who were injured by fragments of plastic (polymethyl methacrylate PMMA) aircraft canopy did not suffer adverse chronic reactions from the presence of the fragments in the body [ 2 ].
1-2 The Aim of the Present Work The aim of this work is to study and compare the corrosion resistance of S.S. 316L and Co-Cr-Mo alloys which may be used as Implant alloys. The main objectives of this work include: 1- Corrosion behaviour of SS 316L and Co – Cr – Mo alloys in Ringer’s solution at pH=7.4 and temperature of 37oC. 2- Corrosion behaviour in the presence of three anti-inflammatory drugs include Aspirin (C9H8O4), Paracetamol (C8H9NO2), and Mefenamic acid (C15H15NO2) with three concentrations of each drug (0.00303, 0.00606, and 0.01212), (0.0086, 0.0172, and 0.0344), and (0.00111, 0.00156, and 0.00201) gm/300ml respectively.
10
3- Corrosion behaviour of in the presence of three anti-inflammatory drugs in addition to increasing percentage of uric acid in the blood, where it rises to 0.7 and 1.2 mg/L in female and male respectively.
4- Evaluate the corrosion parameters in different cases by potentiodynamic polarization tests, such as corrosion potentials, corrosion current densities, and cathodic and anodic Tafel slopes, in addition to measuring corrosion rates and polarization resistance. 5- Microstructure investigation of 316L and Co – Cr – Mo alloys in Ringer’s solution at pH=7.4 and temperature of 37oC, in the absence and presence of three anti-inflammatory.
1-3 Literature Survey Corrosion of surgical implant materials in human body fluid has attracted the interest of many researchers, mainly in the presence of many aggressive materials that can produce adverse effects in the human body by using various techniques.
Gurappa I., (2002) studied the characterization of different materials for corrosion resistance under simulated body fluid conditions. A systematic characterization study has been carried out on different materials such as commercial purity titanium, Ti–6Al–4V, 316L stainless steel, and a cobalt-based alloy under simulated body fluid conditions at 37oC. Breakdown potential, corrosion rates, pitting/crevice corrosion resistance, and the ability to form protective oxide scales were evaluated and compared .An attempt has also been made to study the suitability of titanium alloy for 11
biomedical applications. AC impedance measurements were also carried out in order to provide supportive evidence for the above results [3].
Chun-Che Shih et al., (2003) studied the characterization of the thrombogenic potential of surface oxides on stainless steel for implant purposes. Marketed stents are manufactured from various metals and passivated with different degrees of surface oxidation. Related properties of these oxide films were studied by open - circuit potential, current density detected at open circuit potential, the electrochemical impedance spectroscopy, transmission electron microscopy, Auger spectroscopy (AES), X-ray photoelectron spectroscopy (XPS), and scanning electron microscopy. Experimental evidence showed that blood clot weight after a 30-min follow-up was significantly lower for the stainless steel wire passivated with amorphous oxide (AO) compared to the wire passivated with polycrystalline oxide (PO) or commercial as-received wire coils (AS). Surface characterizations showed that a stable negative current density at open - circuit potential and a significant lower potential were found for the wire surface passivated with AO than for the surface passivated with PO [4]. Tomohiko Yoshioka et al., (2003) studied the preparation of alginic acid layers on stainless-steel substrates for biomedical applications. This study is concerned with the blood compatibility of alginic acid layers immobilized on gaminopropyltriethoxysilane (γ-APS) - grafted stainless-steel. The surfaces were characterized with contact angle measurement and Xray photoelectron spectroscopy (XPS). The blood compatibility was evaluated in 12
terms of platelet adhesion and blood clotting time. An in vitro platelet adhesion assay indicated that only a small number of platelets adhered to substrate surfaces modified with γ-APS and subsequently within alginic acid [5].
Chun-Che Shih et al., (2004) studied the effect of surface oxide properties on corrosion resistance of 316L stainless steel for biomedical applications. Surface passivation is a promising technique for improving the corrosion resistance both in vitro and in vivo as well as the biocompatibility of 316L stainless steel. The effect of different passivation processes on the in vitro corrosion resistance of 316L stainless steel wire was studied. Characterization techniques such as anodic polarization test, scanning electron microscopy, Auger electron spectroscopy, X-ray photoelectron spectroscopy, and transmission electron microscopy were employed to correlate the corrosion to various surface characteristics and surface treatments [6]. Hanawa T. , (2004) studied the metal ion release from metallic materials implants such as stainless steel, cobalt–chromium alloy, titanium, and titanium alloys, implanted into human body. Surface oxide films on metallic materials play an important role as an inhibitor of ion release and it change with the release in vivo. Low concentration of dissolved oxygen, inorganic ions, proteins, and cells may accelerate the metal ion release. The regeneration time of the surface oxide film after disruption also governs the amount of released ion. In addition, preferential release of specific elements during wear and fretting of metallic materials occurs. 13
The metal ion release into biofluid is governed by the electrochemical rule. Released metal ions do not always combine with biomolecules to yield toxicity because active ion immediately combine with a water molecule or an anion near the ion to form an oxide, hydroxide, or inorganic salt. Thus, there is only a small chance that the ion will combine with biomolecules to cause cytotoxicity, allergy, and other biological influences [7].
Laure Duisabeau et al., (2004) studied the environmental effect on fretting of metallic materials for orthopaedic implants; the environment of orthopaedic implants sometimes induces vibrations at the contact of the modular prostheses components. These microdisplacements contribute to the total failure of the implant. The necessary optimisation of orthopaedic device life requires a better knowledge of the damages induced by fretting corrosion. This work describes the damage mechanism at the head–neck contact of a total hip joint with a neck in Ti–6Al–4V alloy and a head in austenitic stainless steel (AISI 316L SS). Simple tests were performed at the ambient air and in an artificial physiologic medium in order to reveal the damage induced by the physiological medium. The presence of a solution containing chloride ions activates a localised corrosion phenomenon which leads to the modification of the displacement accommodation regime [8]. Robert Wen-Wei Hsu et al., (2005) studied the electrochemical corrosion studies on Co–Cr–Mo implant alloy in biological solutions. The electrochemical corrosion of Co–Cr–Mo implant alloys in different biological solutions including urine, serum and joint fluid, was studied by using potentiodynamic scan method, cyclic voltammeter (CV), and impedance spectroscopy. The corrosion characteristic properties of Co–Cr–Mo implant alloys were investigated in terms of corrosion potential (Ecorr), corrosion current density 14
(icorr), and polarization resistance (Rp). Based on the result of CV, the Co–Cr– Mo implant alloy only exhibits small passive region in joint fluid and serum, but a much large region for urine. However, the corrosion resistance of Co–Cr–Mo implant alloys in urine (5128 Ω/cm2) was slightly lower than that in joint fluid (6513 Ω/cm2) and serum (6691 Ω/cm2) at Ecorr and 37oC based on the result of analyses.
A simple Randles circuit model could be used to approximate the corrosion interface of Co–Cr–Mo implant alloy in three biological solutions. It was also experimentally observed that the corrosion interface of Co–Cr–Mo implant alloy in biological solutions showed a characteristic of being capacitive. Finally, the experimental results of Tafel plot analyses were found in good agreement with those of impedance analyses [9]. Sudhakar K.V., (2005) studied the metallurgical investigation of a failure in 316L stainless steel orthopaedic implant. An orthopaedic implant (nail for shinbone) made of 316L stainless steel (SS) that failed prematurely was examined to determine the root cause for the fracture. Detailed scanning electron microscopy was carried out to conclusively establish the evidence. Based on the results of extensive fracture surface analysis as well as the background information provided on the implant, it was determined that the implant (stainless steel nail) failed by the mechanism of predominantly ductile fracture facilitated by the presence of non-metallic inclusions [10]. Yoshimitsu Okazaki et al.,
(2005) studied the comparison of metal
release from various metallic biomaterials in vitro using SS 316L, Co–Cr–Mo casting alloy, commercially pure Ti grade 2, and Ti–6Al–4V, V-free Ti–6Al– 7Nb and Ti–15Zr–4Nb–4Ta alloys , immersed in various solutions containing 15
calf serum, 0.9% NaCl, artificial saliva, 1.2 w% 1-cysteine, 1 w% lactic acid, and 0.01 w% HCl for 7days [11]. Yee-Chin Tang et al., (2006) studied the electrochemical behaviour of type 304 and 316L stainless steels in simulated body fluids and cell cultures. The electrochemical corrosion behaviour of type 304 and 316L stainless steels was studied in Hank's solution, Eagle's minimum essential medium (MEM), serum containing medium (MEM with 10% of fetal bovine serum) without cells, and serum containing medium with cells over a 1-week period. Polarization resistance measurements indicated that the stainless steels were resistant to Hank's and MEM solutions. Type 304 was more susceptible to pitting corrosion than type 316L in Hank's and MEM solutions. The uniform corrosion resistance of stainless steels, determined by RP, was lower in culturing medium than in Hank's and MEM solution. The low corrosion resistance was due to surface passive film with less protection to reveal high anodic dissolution rate. When cells were present, the initial corrosion resistance was low, but gradually increased after 3 days, consistent with the trend of cell coverage. The presence of cells was found to suppress the cathodic reaction, that is, oxygen reduction, and increase the uniform corrosion resistance as a consequence. On the other hand, both type 304 and 316L stainless steels became more susceptible to pitting corrosion when they were covered with cells [12]. Burstein G.T.
et al., (2007) studied the nucleation of corrosion pits in
Ringer's solution containing bovine serum. The effects of the presence of bovine serum on the nucleation of corrosion pits on 316L stainless steel and commercially pure titanium in Ringer’s physiological solution at 37 0C were studied. The experiments involved measurement of current transients generated on microelectrodes under potentiostatic control below the pitting potential.
16
Results show that the presence of the organic components of the serum stimulates the nucleation of pits on 316L stainless steel. A similar but smaller effect is shown in commercially pure titanium [13]. Sargeant A. and Goswami T. , (2007) studied the toxic concentrations of ions that can lead to many adverse physiological effects, including cytotoxicity, genotoxicity, carcinogenicity, and metal sensitivity. There is need to map ion concentrations establishing boundaries between normal and toxic levels; which however, does not exist.
This work reviews the concentrations of ions released from different alloys, including cobalt, chromium, nickel, molybdenum, titanium, aluminium, and vanadium, also reviews clinical data on metal ion concentrations in patients with metal joint prostheses, and laboratory data on the physiological effects of the metals [14]. Huajuan Yang et al., (2007) studied the pitting corrosion resistance of lanthanum (La), a rare earth element with anticoagulative and antiphlogistic function, added into the medical grade 316L stainless steel in order to improve its biocompatibility. The corrosion resistance of the La added 316L steel in two different simulated body fluids, simulated blood plasma and Hank's solution was evaluated. The result showed that the addition of La in the steel could largely affect the corrosion behaviour of the steel. The steel with 0.01% La showed the widest passive region and the best resistance to pitting attack within the addition range of La from 0.01% to 0.08%. The corrosion resistance improvement of La added 316L stainless steel is probably due to the effect of La on the purification of the steel, the modification of inclusions, and the passive film formation in the simulated body fluids. [15]. Donatella Granchi et al., (2008) studied the sensitivity to implant materials in patients with total knee arthroplasties. Materials used for total knee 17
arthroplasty (TKA), may elicit an immune response whose role in the outcome of the arthroplasty is still unclear. The aim of this study was to evaluate the frequency of sensitization in patients who had undergone TKA, and the clinical impact of this event on the outcome of the implant. Ninety-four subjects were recruited, including 20 patients who had not yet undergone arthroplasty, 27 individuals who had a wellfunctioning TKA, and 47 patients with loosening of TKA components [16]. Dania´n Alejandro Lo´pez et al., (2008) studied the electrochemical characterization of AISI 316L stainless steel in contact with simulated body fluid under infection conditions. Titanium and cobalt alloys, as well as some stainless steels, are among the most frequently used materials in orthopaedic surgery. In industrialized countries, stainless steel devices are used only for temporary implants due to their lower corrosion resistance in physiologic media when compared to other alloys. However, due to economical reasons, the use of stainless steel alloys for permanent implants is very common in developing countries The implantation of foreign bodies is sometimes necessary in the modern medical practice. However, the complex interactions between the host and the implant can weaken the local immune system, increasing the risk of infections. Therefore, it is necessary to further study these materials as well as the characteristics of the superficial film formed in physiologic media in infection conditions in order to control their potential toxicity due to the release of metallic ions in the human body. Study of the superficial composition and the corrosion resistance of AISI 316L stainless steel and the influence of its main alloying elements when they are exposed to an acidic solution that simulates the change of pH that occurs when an infection develops was investigated. Aerated simulated body fluid (SBF) was employed as working solution at 370C. The pH was adjusted to 7.25 and 4 in order to reproduce normal body and disease state respectively. Corrosion resistance was measured by means of 18
electrochemical impedance spectroscopy (EIS) and anodic polarization curves [17]. Kajzer W. et al., (2008) studied the corrosion behaviour of AISI 316L steel in artificial body fluids. The tests were carried out on samples of the following surfaces: grinded – average roughness Ra = 0.31 μm and electropolished and chemically passivated average roughness Ra = 0.10 μm. The corrosion tests were realized by recording of anodic polarization curves with the use of the potentiodynamic method. The tests were carried out in electrolyte simulating urine (pH = 6-6.4), Tyrode’s physiological solution (pH = 6.8-7.4) and plasma (pH = 7.2-7.6) at the temperature of 37±1°C. Surface condition of AISI 316L stainless steel determines its corrosion resistance. The highest values of breakdown potentials were recorded for all electropolished and chemically passivated samples in all simulated body fluids. The highest values of anodic current density were recorded for samples tested in artificial urine; the lowest values were recorded for samples tested in Tyrode’s physiological solution. The obtained results are the basis for the optimization of physicochemical properties of the AISI 316L stainless steel [18]. Valero Vidal C. and Lgual Munoz, (2008) studied the electrochemical characterisation of biomaterial alloys for surgical implants in simulated body fluids and compared the electrochemical behaviour of two biomedical alloys AISI 316 L and Co-Cr-Mo in simulated body fluids, this comparison is focused on the influence of solution chemistry and immersion time on passive behaviour using electrochemical techniques which include potentiodynamic curve and potentiostatic tests. Influence of albumin on both biomaterials depends on the nature of the alloy it decreases the corrosion resistance of AISI 316L while increases the corrosion resistance of Co-Cr-Mo. Although it is known that it adsorbs in both cases, properties of the passive layer modifies the effect of albumin. Precipitation of phosphate ions could explain the highest resistance values in the phosphate solutions on both cases. [19]. 19
Virtanen S. et al., (2008) studied the Stainless steels and Cobalt- based alloys of corrosion under physiological and simulated physiological conditions. The aim of their article is to review those aspects of corrosion behaviour that are most relevant to the clinical application of implant alloys. The resistance of the different materials against the most typical corrosion modes (pitting corrosion, crevice corrosion and fretting corrosion) is compared, together with observations of metal ion release from different biomaterials. A short section is dedicated to possible galvanic effects in cases when different types of materials are combined in a biomedical device [20]. Landoulsi J. et al., (2009) studied the enzyme-induced ennoblement of AISI 316L stainless steel. The use of purified enzymes in microbial influenced corrosion (MIC) studies is increasingly recognized as a powerful tool to understand electrochemical interfacial processes, especially the ennoblement of stainless steels (SS) in natural waters. The ennoblement has attracted the interest of many researchers as the consequences in terms of pitting corrosion are still not well understood. The ennoblement of AISI 316L SS was induced by glucose oxidase (Gox) catalyzed reaction or by adding hydrogen peroxide (H 2O2) in synthetic fresh water was studied. The corrosion behaviour of the sample was studied using potentiodynamic and galvanostatic polarization tests.Results obtained using this enzymatic system enable us to reappraise the commonly acknowledged hypothesis that the ennoblement increases the risk of localized attacks [21]. Tavares S.S.M. et al., (2010) studied the characterization of prematurely failed stainless steel orthopaedic implants. Metallic surgical implants are structural components used to accelerate bone consolidation after fracture. A group of implants consists of compression plates fixed to the bone by bolts and nuts. This is particularly useful when the excessively long period of consolidation by traditional methods (without implants) would probably provoke the atrophy of cartilages and articulations of the human body. 20
Surgical implants are submitted to aggressive working conditions such as static and dynamic mechanical loading and exposed to the biochemical and dynamic environments of the human body that contributes to accelerate wear. The load on implant varies with position in walking cycle and reaches a peak of about four times the body weight at the hip and three times the body weight at the knee. Larger loads are assumed by the hip and knee joints during activities such as running and jumping. Austenitic stainless steel has been widely used as osteosynthesis implants because of the excellent mechanical properties, corrosion resistance and cost benefit. Therefore, the high chloride concentration plus the regular temperature of the human body might create localized corrosions like pitting, crevice corrosion and fretting fatigue. The studies of failure analyses help develop better implant devices [22]. Arash Shahryari et al., (2010) studied the response of fibrinogen, platelets, endothelial and smooth muscle cells to an electrochemically modified SS 316L surface towards the enhanced biocompatibility of coronary stents. Modification of a biomedical-grade stainless steel 316L surface by electrochemical cyclic potentiodynamic passivation (CPP) and the response of fibrinogen (Fg), platelets, endothelial cells (ECs) and smooth muscle cells (SMCs) to this surface was investigated. Polarization modulation infrared reflection absorption spectroscopy revealed a significant difference between the secondary structure of Fg adsorbed on the unmodified and CPP surface, the latter being closer to that of native Fg. This was postulated as the origin of the significantly lower surface density of attached platelets on the CPP surface. The competitive interaction of ECs and SMCs with the surface showed that the ECs/SMCs surface density ratio is significantly higher on the CPP surface over the first 2 h of attachment, suggesting faster initial attachment kinetics of ECs on the CPP surface. The presented results thus clearly demonstrate an increase in biocompatibility of the CPP 316L surface [23]. 21
Antunes R.A. et al., (2010) studied the corrosion resistance and in vitro biocompatibility of physical Vapour Deposition titanium carbonitride (TiCN) coated AISI 316 L austenitic stainless steel for orthopaedic applications. The electrochemical behaviour was assessed using potentiodynamic polarization and electrochemical impedance spectroscopy. Cytotoxicity and genotoxicity tests were performed to evaluate the potential biocompatibility of the specimens. TiCN morphology was investigated using scanning electron microscopy (SEM). Bare 316L specimens were also evaluated for comparison. The results showed that the film morphology strongly influences the electrochemical behaviour of the coated underlying metal. TiCN-coated specimens presented neither cytotoxicity nor genotoxicity [24].
22
Chapter Two 2-1 General Concepts Related to Corrosion Corrosion is the interaction of a material with its environment, it is also considered as any process that involves the transfer of atoms (metallic) to ionic state. Corrosion involves the destructive attack of metal by chemical or electrochemical reaction with its environment. Usually, the corrosion process consists of a set of redox reactions which are electrochemical in nature. Thus, the metal is oxidized to corrosion products at anodic sites and some species are reduced at cathodic sites [25]. Thermodynamic considerations determine whether or not a reaction can occur. However, in spite of this limitation, thermodynamic is very important to understand the electrochemistry of corrosion. The kinetics of electrochemical reaction is based largely on the mixed potential theory of electrode kinetics as started by Wagner and Traud [26]. The basic assumptions of the theory are quite simple: (a) The kinetics of the various partial reactions can be treated separately, and (b) no net current flows from an electrode which is in equilibrium or at steady state. The condition of no net current flow means that the total rate of reduction must equal the total rate of oxidation on the electrode surface. When a reaction is forced away from equilibrium, when one direction of the reaction is favored over the other, the potential at which the reaction occurring changes [25]. The amount by which the potential change is the over voltage which is defined as:
E Eeq
……. (2-1)
η= over voltage, Eeq= equilibrium potential, E= polarized potential, which corresponds to corrosion potential of the metal.
2-2 Electrochemical Corrosion
23
Most metal corrosion occurs via electrochemical reactions at the interface between the metal and an electrolyte solution. A thin film of moisture on a metal surface forms the electrolyte for atmospheric corrosion. Corrosion normally occurs at a rate determined by equilibrium between opposing electrochemical reactions. The first is the anodic reaction, in which a metal is oxidized, releasing electrons into the metal. The other is the cathodic reaction, in which a solution species (often O 2 or H+) is reduced, removing electrons from the metal. When these two reactions are in equilibrium, the flow of electrons from each reaction is balanced, and no net electron flow (electrical current) occurs. The two reactions can take place on one metal or on two dissimilar metals (or metal sites) that are electrically connected as shown in Fig, (2-1). The theoretical current for the anodic and cathodic reactions are shown as straight lines. The curved line is the total current, the sum of the anodic and cathodic currents [27].
Fig. (2-1): Corrosion process showing anodic and cathodic current components[ 27]. Steels and other iron-based alloys are the metallic materials most commonly exposed to water. Metal ions go into solution at anodic areas in an amount chemically equivalent to the reaction at cathodic areas (Fig. 2-2). [28] 24
Fig . (2-2): Simple model describing the electrochemical nature of corrosion processes [28]. In the cases of cobalt- based alloys and iron- based alloys, the following reaction usually takes place at anodic areas: M →M+n + ne …..(2-2) This reaction is rapid in acidic media, as shown by the lack of pronounced polarization when iron is made as an anode, employing an external current. When iron corrodes, the rate is usually controlled by the cathodic reaction, which in general is much slower (cathodic control). In deaerated solutions, the cathodic reaction is: 2H+ + 2e → H2 ……. (2-3) This reaction proceeds rapidly in acids, but only slowly in alkaline or neutral aqueous media. The corrosion rate of iron in deaerated neutral water at room temperature, for example, is less than 5 mm/year [28]. The rate of hydrogen evolution at a specific pH depends on the presence or absence of low-hydrogen overvoltage impurities in the metal. For pure iron, the metal surface itself provides sites for H2 evolution; hence, high-purity iron continues to corrode in acids, but at a measurably lower rate than does commercial iron.
25
The cathodic reaction can be accelerated by the reduction of dissolved oxygen in accordance with the following reaction, a process called depolarization: 4H+ + O2 + 4e → 2H2O …. (2-4) Dissolved oxygen reacts with hydrogen atoms adsorbed at random on the iron surface, independent of the presence or absence of impurities in the metal. The oxidation reaction proceeds as rapidly as oxygen reaches the metal surface. Adding equations (2-2) and (2-4), making use of the reaction H2O ↔ H+ + OHˉ, leads to the following reaction: 2Fe + 2H2O + O2 → 2Fe (OH) 2 … (2-5) Hydrous ferrous oxide (FeO.nH2O) or ferrous hydroxide [Fe(OH)2] composes the diffusion-barrier layer next to the iron surface through which O2 must diffuse. The pH of a saturated Fe (OH)2 solution is about 9.5, so that the surface of iron corroding in aerated pure water is always alkaline. The color of Fe (OH)2, although white when the substance is pure, is normally green to greenish black because of incipient oxidation by air. At the outer surface of the oxide film, access to dissolved oxygen converts ferrous oxide to hydrous ferric oxide or ferric hydroxide, in accordance with: 4Fe (OH)2 + 2H2O + O2 → 4Fe(OH)3 ….(2-6) Hydrous ferric oxide is orange to red-brown in colour and makes up most of ordinary rust. It exists as nonmagnetic Fe2O3 (hematite) or as magnetic Fe2O3, the form having the greater negative free energy of formation (greater thermodynamic stability). Saturated Fe (OH)3 is nearly neutral in pH. A magnetic hydrous ferrous ferrite, Fe3O4.nH2O, often forms a black intermediate layer between hydrous Fe2O and FeO. Hence rust films normally consist of three layers of iron oxides in different states of oxidation [28].
2-3 History of Implants Development of biomaterials science from historical point of view has been thoroughly described by Popp (1939), Weinberger (1948), Harkins and 26
Koepp Baker (1948), Baden (1955), Sivakumar (1999), Ratner and Bryant (2004), Staiger et al.(2006), Park and Lakes (2007). In 1565 Alexander Petronius described palatine obturators (De morbo Gallico). He used wax [29]. Steel materials were used in the nineteenth century as bone plates and screws to fix fractures. Fixing fractures with screws allowed a stronger fixture than the earlier method of fixing with metallic wires. Steel made from nickel-plating steel and vanadium steel later replaced carbon steel materials as steel corrodes easily in the human body. However, these newer materials were not sufficiently corrosion resistant. It also became clear that they become toxic inside the human body. Historically speaking, until Dr. J. Lister’s aseptic surgical technique was developed , attempts to implant various metal devices such as wires and pins constructed of iron, gold, silver, platinum were largely unsuccessful due to infection after implantation [29]. Sherman and Pittsburgh modified the Lane plate to reduce the stress concentration by eliminating sharp corners. He used vanadium alloy steel for its toughness and ductility. Subsequently, Stellite (Co–Cr-based alloy) was found to be the most inert material for implantation by Zierold in 1924[30]. Soon 18 w% Cr, 8 w% Ni and 2–4 w% Mo stainless steels were introduced for their corrosion resistance, with 18-8 SS Mo being especially resistant to corrosion in saline solution. Later, another alloy (19 w% Cr, 9 w% Ni) named Vitallium was introduced into medical practice [29]. The first use of magnesium was reported by Lambotte in 1907, who utilized a plate of pure magnesium with gold-plated steel nails to secure a fracture involving the bones of the lower leg . The attempt failed as the pure magnesium metal corroded too rapidly in vivo, disintegrating only 8 days after surgery
and
producing
a
large
amount
of
gas
beneath
the skin [31]. Albee and Morrison first studied calcium phosphate (CaP) compounds in 1920, injecting tricalcium phosphate (TCP) into animals to test its efficacy as a bone substitute. A noble metal, tantalum, was introduced in 1939, but its poor 27
mechanical properties and difficulties in processing it from the ore made it unpopular in orthopaedics, yet it found wide use in neurological and plastic surgery[32]. Smith-Petersen designed the first nail with protruding fins to prevent rotation of the femoral head. He used stainless steel but soon changed to Vitallium R .Thornton attached a metal plate to the distal end of the SmithPetersen nail and secured it with screws for better support. Later in 1939, he used an artificial cup over the femoral head in order to create new surfaces to substitute for the diseased joints. Also he used glass, Pyrex, and Vitallium. The latter were found more biologically compatible, and 30–40% of patients gained usable joints [29]. Similar mold arthroplastic surgeries were performed successfully by the Judet brothers of France, who used the first biomechanically designed prosthesis made of an acrylic (methyl methacrylate) polymer. In particular, the idea that the release of toxic leachables from biomaterials will adversely affect healing was formalized—this toxicology idea is implicit in today’s definition of biocompatibility [29]. As developments took place in biology and materials science, biomaterials researchers were quick to incorporate these new ideas into biomaterials. By the time of the 1950s–1960s, blood vessel replacements were in clinical trials and artificial heart valves and hip joints were in development. Thus, till the polymer industry was developed in 1950s, the metallic materials were mainly used. The first quarter century, 1950–1975, of biomaterials development was dominated by the characteristics of the materials intended for prostheses and medical devices. Blood vessel implants were attempted with rigid tubes made of polyethylene [29]. Heart valve implantation was made possible only after the development of open-heart surgery in the mid-1950s. Starr and Edwards in 1960 made the first commercially available heart valve, consisting of a silicone rubber ball poppet in 28
a metal strut. Concomitantly, artificial heart and heart assist devices have been developed [33]. Important in the early days was the long-term integrity of the biomaterial as well as its non-toxic nature. Biological interactions that were considered included the non-toxic nature of the biomaterial as well as its normal inflammatory and wound healing responses when implanted. Many materials were described as being inert, but this was a confusing descriptor as it did not adequately and appropriately describe material changes following implantation or cell and tissue responses to the implanted biomaterial. It eventually became clear that materials could change without adversely affecting the function and interaction of the biomaterial, prosthesis, or medical device. Likewise, modulation of the inflammatory and wound healing responses could occur without altering the function of the biomaterial, prosthesis, or medical devices. From 1970 to 2000, biological interactions with biomaterials started to be more extensively investigated [29]. The discovery by Hench and co-workers that a range of compositions of modified phosphosilicate glasses has the ability to form a stable chemical bond with living tissues (bone, ligament, and muscle) opened a completely new field in biomedicine . Since then, many artificial biomaterials based on, or inspired by, Hench’s glasses have been developed and successfully employed in clinical applications for repairing and replacing parts of the human body. This field is continuously expanding: new processing routes have extended the range of applications toward new and exciting directions in biomedicine, many of which still rely on the original Hench’s base formulation, 45S5 Bioglass, which has now become the paradigm of bioactive materials. Advances in our knowledge of biological mechanisms, for example, the coagulation, thrombosis, and complement pathways, led to a better understanding of biological interactions with biomaterial surfaces. In the 1980s, the revolution in techniques for the study of cell and molecular biology led to their application to the investigation of interactions occurring at biomaterial interfaces [34]. More recently, with the 29
advent of the areas of tissue engineering and regenerative medicine, heavy emphasis has been placed on biological interactions with biomaterials. What is the state of the art today? Surprisingly, gold is still quite popular! Recently, it was shown that implants of pure metallic gold release gold ions which do not spread in the body, but are taken up by cells near the implant. It was hypothesized that metallic gold could reduce local neuron inflammation in a safe way [35]. The modern biomaterials science is defined and explained through the introduction of biotechnology and advances in the understanding of human tissue compatibility. Developing from bio-inert materials to biodegradable materials, biomaterials are widely used in medical devices, tissue replacement, and surface coating applications [29]. Improved patient benefits form the most important factor stimulating market growth for biomaterials, where major segments are as usual ceramics, metals, polymers, and composites. Reconstructive surgery and orthobiologics are the dominant segments in orthopaedic biomaterials today. Placement of end osseous implants has improved the quality of life for millions of people. It is estimated that over 500,000 total joint replacements, primarily hips and knees, and between 100,000 and 300,000 dental implants are used each year in the United States alone [36]. Total joint arthroplasty relieves pain and restores mobility to people such as those afflicted with osteoarthritis, and dental implants provide psychological and aesthetic benefits in addition to improving masticatory function for edentulous patients. Modern biomaterials found applications not only in orthopaedic, cardiovascular, gastrointestinal, wound care, urology, and plastic surgery, but in such directions as brain repair [37]. Recently Orive et. al. In 2009, developed biomaterials that can enable and augment the targeted delivery of drugs or therapeutic proteins to the brain, allow cell or tissue transplants to be effectively delivered to the brain, and help to rebuild damaged circuits. Similarly, biomaterials are being used to promote regeneration and to repair damaged neuronal pathways in combination with stem 30
cell therapies. Many of these approaches are gaining momentum because nanotechnology allows greater control over material–cell interactions that induce specific developmental processes and cellular responses including differentiation, migration and outgrowth [38].
2-4 Requirements of Implants An implant should possess some important properties for long-term usage in the body without rejection. The design and selection of biomaterials depend on their mechanical and non- mechanical characteristics:
2-4-1 Mechanical Properties The mechanical properties such as hardness, tensile strength, modulus, elongation , fracture resistance and fatigue strength or life play an important role in material selection for application in the human body. Fatigue strength is related to the response of the material to repeated cyclic loads [39]. Teoh pointed out in his paper that fatigue fracture leads some of major problems associated with implant loosening, stress-shielding and ultimate implant failure [40]. For major applications such as total joint replacement, higher yield strength is basically coupled with the requirement of a lower modulus close to that of human bones. The magnitude of bone modulus varies from 4 to 30 GPa depending on the type of the bone and the measurement direction [39]. Au et al. [41] and Geetha et. al. [42] emphasized about the modulus and described that the large difference in the Young’s modulus between implant material and the surrounding bone can contribute to generation of severe stress concentration, namely load shielding from natural bone, which may weaken the bone and deteriorate the implant/bone interface, loosening and consequently failure of implant. The modulus is considered as a main factor for selection of total knee replacement (TKR) materials [39]. 31
2-4-2 Non-mechanical Requirements In addition to the mentioned mechanical properties, some non-mechanical requirements which have significant role in performance of the material in the human body are as follows: 2-4-2-1 High Corrosion Resistance Singh and Dahotre [43] researched on corrosion resistance as an important issue in selection of metallic biomaterials because the corrosion of metallic implants due to the corrosive body fluid is unavoidable. The implants release undesirable metal ions which are non biocompatible. Corrosion can reduce the life of implant device and consequently may impose revision surgery. In addition, the human life may be decreased by the corrosion phenomenon. Okazaki and Gotoh [44] expressed the fact that dissolved metal ions (corrosion product) can either accumulate in tissues, near the implant or they may be transported to other parts of the body. They revealed for example, replacement of 20 stainless steel Charnley hip arthroplasties in the human body after 10–13 years showed a considerably higher metallic concentration in body fluid in comparison with that without implant. 2-4-2-2 Biocompatibility One of the most important non-mechanical requirements of orthopaedic biomaterials is the biocompatibility. Biocompatibility is the ability to exist in contact with tissues of the human body without causing an unacceptable degree of harm to the body. It is not only associated with toxicity, but to all the adverse effects of a material in a biological system [39]. Navarro et al. [45] supported the study of Smallman and Bishop [46] and with retrospect to the last 60 years, categorized three generations for evolution of biomaterials: bioinert materials, bioactive and biodegradable materials and materials designed to stimulate specific cellular responses at the molecular level. 32
Bioinert is related to reducing the body reaction to the implant to a minimum. Bioactivity is defined as the ability of the material to interact with the biological environment to enhance the biological response. The third generation refers to the capability of the material to stimulate specific cellular responses at the molecular level.
Williams defined the biomaterial requirements of total joint replacements in terms of biocompatibility as, optimizing the rate and quality of bone opposition to the material, minimizing the release rate of corrosion and the tissue response to the released particles, minimizing the release rate of wear debris and the tissue reaction to this debris and optimizing the biomechanical environment in order to minimize disturbance to homeostasis in the bone and surrounding soft tissue [47]. 2-4-2-3 Osseointegration Osseointegration is fundamental in orthopaedic. Several literatures explained the integration of the implant with adjacent bone and tissue. Osseointegration is defined as the process of formation of new bone and bone healing. The incapability of an implant surface to join with the adjacent bone and other tissues due to micromotions results in formation of a fibrous tissue around the implant and promote loosening of the prostheses. Thus, materials with a proper surface are extremely essential for the implant to integrate well with the surrounding bone. Surface chemistry, roughness and topography are all parameters that influence both the osseointegration and biocompatibility. It should be considered that in addition to properties of the implanted biomaterial, the characteristics and regenerative capability of the host bone affect the osseointegration of biomaterial [48]. 33
2-5 Types of Implants 2-5-1 Polymer Polymers are organic materials that form large chains made up of many repeating units. Polymers are extensively used in joint replacement components. Currently the polymers most widely used in joint replacements are: [49] • Ultrahigh molecular weight polyethylene (UHMWPE) • Acrylic bone cements • Thermoplastic polyether ether ketone (PEEK) • Bioabsorbables. 2-5-2 Ceramic Ceramics are polycrystalline materials. The great majority are compounds made up of metallic as well as non-metallic elements; they generally have ionic bonds or ionic bonds with some covalent bonds. The main characteristics of ceramic materials are hardness and brittleness. Ceramic-on-ceramic materials (Figure 2-3) have behaved more predictably in the laboratory than metal-onmetal devices. In addition, ceramic is an inert material that does not raise the concerns of future biological consequences specific to metal-on-metal designs [49]. The main ceramics in orthopaedic surgery and their applications are: • Alumina, Al2O3, used for acetabular and femoral components • Zirconia, ZrO2, used for acetabular and femoral components • Hydroxyapatite, Ca10(PO4)6(OH)2, used for coating stem femoral components to integrate the surface material to the bone.
Fig.(2-3): ceramic implant [49]. 34
2-5-3 Composites Composite biomaterials are made with a filler (reinforcement) addition to a matrix material in order to obtain properties that improve every one of the components. This means that the composite materials may have several phases. Some matrix materials may be combined with different types of filler. Polymers containing particulate filler are known as particulate composites [49]. The following composites are considered in the orthopaedics devices: • Fiber-reinforced polymers • Aggregates of polymethyl methacrylate (PMMA). 2-5-4 Metallic Implants Metallic implants are often used to support and/or replace components of the skeleton. They are used, e.g. as artificial joints, bone plates, screws, intramedullary nails, spinal fixations, spinal spacers, external fixators, pace maker cases, artificial heart valves, wires, stents, and dental implants. They possess greater tensile strength, fatigue strength, and fracture toughness when compared to polymeric and ceramic materials. Most widely used metallic biomaterials for implants devices are 316L stainless steels, cobalt alloys, commercially pure titanium and Ti-6Al-4V alloys. Originally, these materials were developed for industrial purposes. Their excellent corrosion resistance, which results in very small release of harmful toxins when exposed to bodily fluids, is the main reason for these materials and can be left inside the body for a longer period of time and are therefore appropriate for medical uses as shown in Fig. (2-4), [50]. Fig. (2-5) shows some shape of metallic biomaterials. These review metallic biomaterials will be divided into four subgroups: titanium alloys, noble metals, cobalt alloys and stainless steels [51].
35
-Stainless steel 316 L - Ti, Ti alloys -Co- Cr- Mo alloys - Stainless steel 316L L
-Stainless steel 316L -Co-Cr- Mo alloys - Ti,Ti alloys
-Stainless steel 316 L -Co- Cr – Mo alloys
-Co – Cr – Mo alloys -Stainless steel 316 L
-Stainless steel 316 L - Co-Cr-Mo alloys - Ti- Al – V alloys
Fig (2-4): Metallic implants in human body [51].
36
Fig. (2-5): Some of metallic implants [50]. 2-5-4-1 Titanium and its Alloys Titanium and its alloys have been increasingly used in medical implants because of their excellent biocompatibility, corrosion resistance and relatively low density. Like stainless steel, titanium alloys from passive films on their surface, in this case a TiO2 film that is the source of their corrosion resistance. The mechanical properties depend on the condition of cold working and annealing for a given grade of interstitial element levels. Fully annealed with equiaxed grains exhibits lower strength than the cold worked ones [52]. 2-5-4-2 Noble Metals 1- Gold Gold is an inert metal that has high resistance to bacterial colonization. Gold and its compounds have been historically used in oriental cultures for the treatment of ailments. It has been one of the first materials to be used as an implantable material (dental tooth implant). Due to its high malleability, it is used in restorative dentistry for crown and permanent bridges. Gold possesses excellent electrical conductivity and biocompatibility and is used in wires for pacemakers and other medical devices [53].
2- Platinum 37
Platinum possesses excellent corrosion resistance, biocompatibility, and stable electrical properties. It is used for manufacture of electrodes in devices such as cardiac pacemakers and electrodes in cochlear (cavity within the inner ear) replacement for the hearing impaired [53]. 3- Silver Silver is used in surgical implants and as a sanitizing agent. They are used as studs of earring to prevent infection of newly pierced ears [53].
2-5-4-3 Cobalt- based Alloys Cobalt alloys may be generally described as nonmagnetic, corrosion and heat resistant which exhibit high strength even at elevated temperature and are also resistant to wear. Many of their properties originate from the crystallographic nature of cobalt, and formation of extremely hard carbides and the corrosion resistance imparted by chromium. Cobalt alloys are difficult to fabricate, therefore, their use has been limited, but continuous work led to the development of specialized casting method and recently to selective laser sintering. Due to cobalt alloys, excellent resistance to degradation in the oral environment, the first medical use was in the cast of dental implants. Various invitro and in-vivo tests have shown that the alloys are biocompatible and suitable for use as surgical implants. Today the use of Co alloys for surgical applications is mainly related to orthopedic prostheses for the knee, shoulder and hip as well as to fracture fixation devices. Joint end prostheses are typical long-term implants, and the applied implant material must therefore meet extremely high requirements with regard to biocompatibility with the surrounding body tissue material and corrosion resistance to body fluids [54]. Now days, the Co-Cr-Mo cast and wrought versions of alloys are highly biocompatible materials and are widely used as orthopedic implant materials in clinical practice, such as hip joint and knee replacement. The biocompatibility of 38
Co-Cr-Mo alloy is closely related to its excellent corrosion resistance due to the presence of an extremely thin passive oxide film that spontaneously forms on the alloy surface. Similar to AISI 316L stainless steel, predominant oxide film is Cr2O3 with some minor contribution from Co and Mo oxides. In spite of the alloys excellent corrosion resistance, there is still some concern about metal ion release from orthopedic implants into the human body environment. Implant components fabricated from Co-Cr based alloys have been reported to produce elevated Co, Cr and Ni concentrations in the surrounding tissue. Other Co alloys used in medicine are MP35N or Co-Ni-Cr-Mo (ASTM F 562) with a nickel content of 35 % used for cardiovascular pacing leads, styles, catheters and orthopedic cables. Increased content of nickel exhibits an improved resistance to stress-corrosion cracking in aqueous solution, but also increase the possibility of nickel allergy reactions. Therefore, these alloys are not ideal for orthopedic applications. The biocompatibility of the wear particles produced can be troublesome because of the increased surface area of these small particles which are in direct contact with the surrounding medium or tissue material. In work-hardened or workhardened and aged conditions, this alloy has very high tensile properties which are among the strongest available for implant applications. Other Co-based alloy is L-605 cobalt alloy or Co-Cr-W-Ni (ASTM F 90) which is used for heart valves and in an annealed condition (ASTM F 1091) for surgical fixation wires. Its mechanical properties are approximately the same as those of Co-Cr-Mo alloys, but after the material is cold worked the mechanical properties increase more than double [54].
2-5-4-4 Stainless Steel Stainless steel, used for medical implants, is mainly austenitic type 316L due to its resistance to corrosion, together with a wide range of other physical and mechanical properties coupled with inert, easily to clean surfaces. The 39
chemical composition of type 316L stainless steel was developed to obtain stable austenitic structure which has numerous advantages, namely: Austenitic stainless steel has a face centered cubic structure and is characterized by very low yield strength to tensile strength ratio and high formability. To increase strength, cold working and successive strain aging treatment can be applied. Austenitic stainless steel is superior to ferritic stainless steel in corrosion resistance because the crystallographic atomic density of the former is higher than that of the latter. Austenitic stainless steel is essentially nonmagnetic. The disadvantages of austenitic stainless steels generally are higher sensitivity toward pitting corrosion and stress corrosion cracking. Pitting corrosion causes deep pits on the metal surface. It is initiated when an oxidant such as dissolved oxygen, reacts with chloride ions. Pitting is further accelerated by the existence of an oxygen concentration cell at the early growth stage. The chemical composition of type 316L (ASTM F138, F139) austenitic stainless steel where ―L‖ denotes low carbon content is as follows: ≤0.030 % C, ≤1.0 % Si, ≤2.0 % Mn, ≤0.045 % P, ≤0.030 % S, 12.0- 15.0 % Ni, 16.0-18.0 % Cr, and 2.0-3.0 % Mn . Its corrosion resistance is improved by adding molybdenum, increasing nickel and reducing carbon to less than 0.030 %. This steel has less than 0.03 wt. % carbon in order to reduce the possibility of in vivo corrosion. If the carbon content of the steel significantly exceeds 0.03 %, there is increased danger of formation of carbides such as Cr23C6 [55]. These tend to precipitate at grain boundaries when the carbon concentration and thermal history have been favorable to the kinetics of carbide growth. In turn, this carbide precipitation depletes the adjacent grain boundary regions of chromium, which has the effect of diminishing formation of the protective chromium- based oxide Cr2O3. The presence of molybdenum as an 40
alloying element in stainless steel reduces both the number and the size of nucleations and metastable pits. This is because bonds in the oxide film are strengthened and active sites caused by the formation of molybdates or of molybdenum oxyhyroxides are eliminated. Due to high content of chromium, 316L stainless steel forms a protective, adherent and coherent oxide film that envelops the entire outer surface. This oxide film, often called passive layer, is intentionally formed when device is manufactured as chromium in the surface layer reacts with oxygen creating Cr2O3. The passive film serves as a barrier to corrosion processes in alloy systems that would otherwise experience very high corrosion rates and has ability of self-healing, when damaged, as chromium in the steel reacts with oxygen and moisture in the environment to reform the protective oxide layer [55]. Cobalt and chromium have the same affinity for proteins, but nickel significantly competes for cobalt and chromium binding areas. In the metal ions – mechanisms, biological risks of metal ions include wear debris, colloidal organometallic complexes, free metal ions, and inorganic metal salts or oxides formed. Organometallic complexes are formed by metal ions binding to proteins. Since proteins are Zwitter ions, most are negatively charged in the body’s pH of 7.4. Positively charged metal ions, including iron, cobalt, chromium, and nickel bind to proteins, changing the pH of solutions. Proteins increase corrosion rate of an implant by increasing the dissolution of metals, especially cobalt and chromium [56]. Figs (2-6) and (2-7) show sample of stainless steel 316 L that was implanted and corroded respectively, taken at Al-Karkh hospital by Dr. Faried Salman.
41
Fig. (2-6): Corroded surgical implant sample.
Fig. (2-7): Corroded sample after removal.
2-6 Types of Corrosion in Implants 2-6-1 Pitting Corrosion Probably the most common type of localized corrosion is pitting, in which small volumes of metal are removed by corrosion from certain areas on the surface to produce craters or pits that may culminate in complete perforation of a pipe or vessel wall. Pitting corrosion may occur on a metal surface in a stagnant or slow-moving liquid. It may also be the first step in crevice corrosion. Pitting is considered to be more dangerous than uniform corrosion damage because it is more difficult to detect, predict, and design against. A 42
small, narrow pit with minimal overall metal loss can lead to the failure of an entire engineering system. Only a small amount of metal is corroded, but perforations can lead to costly repair of expensive equipment [57]. 2-6-2 Crevice Corrosion Crevice corrosion occurs in cracks or crevices formed between mating surfaces of metal assemblies, and usually takes the form of pitting or etched patches. Both surfaces may be of the same metal or of dissimilar metals. It can also occur under scale and surface deposits and under loose fitting washers and gaskets that do not prevent the entry of liquid between them and the metal surface. Crevices may proceed inward from a surface exposed to air, or may exist in an immersed structure [57].
2-6-3 Fretting Corrosion Fretting corrosion refers to corrosion damage at the asperities of contact surfaces. This damage is induced under load and in the presence of repeated relative surface motion, as induced, for example, by vibration. Pits or grooves and oxide debris characterize this damage, typically found in machinery, bolted assemblies, and ball or roller bearings. Contact surfaces exposed to vibration during transportation are exposed to the risk of fretting corrosion [57]. 2-6-4 Galvanic Corrosion Galvanic or two metal corrosion takes place when two different metals are in physical contact in an ionic conducting fluid medium such as serum or interstitial fluid. The differential composition or process variables of a plate and
43
the adjoining screws is responsible for the set-up of a galvanic couple, which results in galvanic corrosion [58]. Galvanic corrosion depend on a large number of complicating factors including the relative areas of electronic and ionic contact, as well as the actual metal pair involved. However, it is safe to assume that some galvanic corrosion will occur in any dissimilar metal pair in acidic pH. In many practical applications, the contact of dissimilar materials is unavoidable. In surgical implants, galvanic corrosion can occur if bone plate and bone screws are made of dissimilar metals or alloys. Corrosion is likely to occur between the plate and bottom side of the screw holes [58]. 2-6-5 Corrosion Fatigue Corrosion fatigue is a fracture failure of metal that occurs because of the combined interaction of electrochemical reactions and cyclic loading. Corrosion fatigue resistance is an important factor of consideration for load-bearing surgical implant metals or for metals used in cyclic motion applications. Normally, a failure may not occur, but cracks can initiate from hidden imperfections, surface damage, minute flaws, chemical attack and other causes. The corrosive environment may result in local corrosive attack that accentuates the effect of the various imperfections. The corrosive attack will be influenced by solution type, solution pH, oxygen content and temperature. The body fluid environment may well decrease the fatigue strength of the implant. Fatigue striations are observed on the fractured surface of the device with coloured ―beach marks‖ are indicative of corrosion fatigue. The presence of corrosion pit or pits could induce the fatigue to develop (Sivakumar et al 1994). Failures of mechanical origin in orthopaedic implants are most commonly due to fatigue or environmentally assisted fatigue [58].
44
2-7 Corrosion Parameters The corrosion current density (icorr) cannot be measured directly; it can be estimated from current versus voltage data. It can be measure a log current versus potential curve over a range of about one half volt. The voltage scan is centered on open circuit potential (Eoc), and then fit the measured data to a theoretical model of the corrosion process. The model, used for the corrosion process, assumes that the rates of both the anodic and cathodic processes are controlled by the kinetics of the electron transfer reaction at the metal surface. This is generally the case for corrosion reactions. An electrochemical reaction under kinetic control obeys the Tafel equation [59]:
E Eeq I I o exp ……….. (2-7) b where Io is the equilibrium exchange current, b is the Tafel slope (cathodic or anodic) and Eeq= equilibrium potential, E= polarized potential which corresponds to corrosion potential of the metal.
The Tafel equations for both the anodic and cathodic reactions in a corrosion system can be combined to generate the Butler-Volmer equation [59] Ecorr. Eeq Eeq Ecorr. I corr. I o exp I o exp …… (2-8) ba bc
where Ecorr. is corrosion potential. At Ecorr, each exponential term equals one, the cell current is therefore zero. Near Ecorr, both exponential terms contribute to the overall current. 45
Finally, as the potential is driven far from Ecorr by the potentiostat, one exponential term predominates and the other term can be ignored. When this occurs, a plot of log current versus potential becomes a straight line. A log I versus E plot is called a Tafel Plot. In practice, many corrosion systems are kinetically controlled and thus obey Equation 2-8. A log current versus potential curve that is linear on both sides of E corr is indicative of kinetic control for the system being studied. Equation 2-8 can be further simplified by restricting the potential to be very near to Ecorr. Close to Ecorr, the current versus voltage curve approximates a straight line. The slope of this line has the units of resistance (ohms). The slope is, therefore, called the polarization resistance (Rp). An Rp value can be combined with an estimate of the Beta coefficients (b) to yield an estimate of the corrosion current [28]. icorr.
(1 / R p )babc
2.303(ba bc )
………. (2-9)
In a polarization resistance experiment, it records a current versus voltage curve as the cell voltage is swept over a small range of potential that is very near to Eoc (generally ± 10 mV). A numerical fit of the curve yields a value for the polarization resistance (Rp). Polarization resistance data does not provide any information about the values for the Beta coefficients. Therefore, to use Equation 3-3, it must provide Beta values.
2-8 Calculation of Corrosion Rate from Corrosion Current As mentioned in application notes, most corrosion phenomena are of electrochemical nature and consist of reactions on the surface of the corroding metal. Therefore, electrochemical test methods can be used to characterize corrosion mechanisms and predict corrosion rates. The corrosion rate depends 46
on the kinetics of both anodic (oxidation) and cathodic (reduction) reactions. According to Faraday's law, there is a linear relationship between the metal dissolution rate and corrosion rate (CR) and the corrosion current density (icorr) [28]: CR
M icorr. ……..(2-10) nF
Where M is the atomic weight of the metal, ρ is the density, n is the charge number which indicates the number of electrons exchanged in the dissolution reaction and F is the Faraday constant, (96.485 C/mol). The ratio M/n is also sometime referred to as equivalent weight (e), and then equation (214) becomes [28]: CR
e …………..(2-11) icorr. F
We are interested in corrosion rates in the more useful units of rate of penetration, such as millimeters per year (mm/y), where the corrosion rate e
equation becomes [ CR (mm / y) 3.271 icorr. ], when e is in (g), ρ in (g/cm3), and
icorr. in (μA.cm-2) units.
2-9 Why Metal Corrodes in Human Body? Corrosion, the gradual degradation of materials by electrochemical attack is of great concern particularly when a metallic implant is placed in the hostile electrolytic environment of the human body. The implants face severe corrosion environment which includes blood and other constituents of the body fluid which encompass several constituents like water, sodium, chlorine, proteins, plasma, amino acids along with mucin in the case of saliva. The aqueous medium in the human body consists of various anions such as chloride, phosphate, and bicarbonate ions, cations like Na+, K+, Ca2+, Mg2+ 47
etc., organic substances of low-molecular-weight species as well as relatively high molecular- weight polymeric components, and dissolved oxygen . The biological molecules upset the equilibrium of the corrosion reactions of the implant by consuming the products due to anodic or cathodic reaction. Proteins can bind themselves to metal ions and transport them away from the implant surface upsetting the equilibrium across the surface double layer that is formed by the electrons on the surface and excess cations in the solution. In addition, proteins that are absorbed on the surface are also found to reduce the diffusion of oxygen at certain regions and cause preferential corrosion at those regions. Hydrogen which is formed by cathodic reaction acts as a corrosion inhibitor, however, the presence of bacteria seems to change this and enhance corrosion by absorbing the hydrogen present in the vicinity of the implant. Changes in the pH values also influence the corrosion [60]. Though, the pH value of the human body is normally maintained at 7.0, this value changes from 3 to 9 due to several causes such as accidents, imbalance in the biological system due to diseases, infections and other factors and after surgery, the pH value near the implant varies typically from 5.3 to 5.6. In spite of the fact that most of the materials used are protected by the surface oxide layers from the environmental attack, there is clinical evidence for the release of metal ions from the implants and this leaching has been attributed to corrosion process. The most common forms of corrosion that occur are uniform corrosion, intergranular, galvanic and stress corrosion cracking, pitting and fatigue corrosion. Even though new materials are continuously being developed to replace implant materials used in the past, clinical studies show that these materials are also prone to corrosion to a certain extent. The two physical characteristics which determine implant corrosion are thermodynamic forces which cause corrosion either by oxidation or reduction reaction and the kinetic barrier such as surface oxide layer which physically 48
prevents corrosion reactions. In some cases, though the material will not fail directly due to corrosion, it is found to fail due to accelerated processes such as wear and fretting leading to tribocorrosion. Fretting results in the rupture of protective oxide layer, initiation of cracks and formation of reactive metal atoms on the surface that are susceptible to corrosion [61]. In order to limit further oxidation, initially formed passive films must have certain characteristics; i) non - porous ii) atomic structure that will limit the migration of ions and electrons across the metal oxide - solution interface and iii) high abrasion resistance. Hence, when a material is developed for implant application, it should not only be subjected to basic corrosion screening test, but also has to be tested for other tests such as reciprocatory wear, stress corrosion and fretting etc. as shown in Fig.(2-8), depending upon their applications. Corrosion is accelerated in the presence of wear and also simultaneous corrosion and wear are often encountered in biomedical implants. Dearnley et al. have evaluated the corrosion of the scratched coated specimens to determine the wear accelerated corrosion of the coatings and also suggested a methodology to measure the simultaneous corrosion and wear of a material. However it is important to note that there are no standards available to test the tribocorrosion of the implants [62].
Fretting
Fig. (2-8): Fretting corrosion [62]. 49
2-10 Drug Drug is a substance that induces certain biological reposes in the human body. There are three phases that a drug undergoes after administration: pharmaceutical phase, pharmacokinetic phase and pharmacodynamic phase. Pharmaceutical phase is the initial phase of the drug in the body after administration, the active ingredient of the formulation dissolves. Drug delivery concerns how the drug is introduced to the body, and it is indeed an old topic. Egyptian physicians employed pills for treatment over4000 years ago. Intravenous injections were first performed in humans in 1665, only a few years following Wren’s infusion of opium into dogs and a few decades after Harvey’s description of the circulatory system in 1616. Subcutaneous injections were introduced by Wood in 1853 and the modern hypodermic syringe was developed by Luer in 1884 [63]. A body contains many biological barriers that serve to protect its interior from a variety of external invaders. All drug products are ―foreign‖ and thus challenged throughout the delivery process. Pharmacokinetics (PK) is the study of the action that body takes towards a therapeutic and gent. It can be divided into four basic processes: absorption (A), distribution (D), metabolism (M) elimination (E) [63]. • Absorption refers to the transfer of the drug from the site of administration to the general circulation. • Distribution deals with the transfer of the drug from the general circulation into the different organs of the body. • Metabolism concerns the biotransformation process (chemical modification) that drugs undergo before elimination. • Elimination is the process by which the drug is removed from the body. This may involve excretion and/or metabolism.
2-11 Inflammation
50
Inflammation is a nonspecific physiological response to tissue damage in animal systems, it arises as a response to trauma, infection, intrusion of foreign materials, local cell death, or as in adjunct to immune response. An inflammatory process is initiated consisting of a complex series of reactions tending to prevent the ongoing tissue damage, isolate and destroy the foreign material and activate the repair processes. The aim of inflammatory response is the same whether it is induced by mechanical trauma, microbial infection, foreign antigens or electrical, chemical or radiological energy. Depending on the types of protein and the degree of denature of the adsorbed proteins, the material may cause inflammatory responses [64]. Inflammation is the reaction of living vascularised tissue to local damage. It serves many functions, including the isolation, dilution or neutralization of the process or agent causing the injury. Inflammation initiates a series of events that may lead to healing. In particular, it may lead to replacement of the damaged tissue by regeneration of nieghbouring parenchymal cells, or to the formation of scar tissue, so called fibrosis. Immediately as the damage has occurred, blood flow and vascular permeability are altered. This causes fluid that is rich in proteins and blood cells escape from the adjacent blood vessels in a process known as exudation. This triggers a series of cellular events that characterize the inflammatory response. Among them is blood clotting and possibly thrombosis formation.Following injury, there is a progressive change in the cells which predominate at the site. Initially, neutrophils are present. These are attracted to the site early on as a result of chemotactic factors related soon after the initial injury. They are short- lived, and they disappear or disintegrate within some 24-48 hours. They are then replaced by monocytes. At the site of injury, monocytes undergo a differentiation process to become macrophages, and these cells may live for several months. Emigration of the undifferentiated monocytes towared the site of injury may continue for days or even weeks, depending on the extent of damage, partly because chemotactic factors for monocytes are activated over relatively long 51
periods of time.Although triggered by the injury, i.e. in the case the surgical procedure to place the implant, the nature of the biomaterial itself has an influence on the progress of the inflammatory response. Depending on its geometry, surface chemistry and mechanical properties, the implant can cause variations in the intensity and duration of the inflammatory response [64].
2-12 Non Steroidal Anti-inflammatory Drugs The NSAIDS are a group of chemically dissimilar agents that differ in their antipyretic, analgesic, and anti-inflammatory activities. They act primarily by inhibiting the cyclooxygenase enzymes that catalyze the first step in prostanoid biosynthesis. This leads to decreased prostaglandin synthesis with both beneficial and unwanted effects. Detection of serious cardiovascular events associated with COX-2 inhibitors has led to withdrawal of rofecoxib and valdecoxib from the market. The U.S. Food and Drug Administration (FDA) has required that the labelling of the traditional NSAIDS and celecoxib be updated to include 1) a warning of the potential risks of serious cardiovascular thrombotic events, myocardial infarction, and stroke, which can be fatal, additionally, a warning that the risk may increase with duration of use and that patients with cardiovascular disease of risk factors may be at greater risk, 2) a warning that use is contraindicated for the treatment of perioperative pain in the setting of coronary artery bypass graft surgery, and 3) a notice that there is increased risk of serious gastrointestinal adverse events, including bleeding, ulceration, and perforation of the stomach or intestines, which can be fatal [65].
2-13 Some Anti-Inflammatory Drugs 2-13-1 Aspirin
52
Aspirin also known as acetylsalicylic acid is a salicylate drug, often used as an analgesic to relieve minor aches and pains, as an anti-inflammatory medication. Antipyretic to reduce fever, Salicylic acid, the main metabolite of aspirin is integral part of human and animal metabolism, while much of it is attributable to diet a substantial part is synthesized endogenously. Aspirin also has an antiplatelet effect by inhibiting the production of thromboxane, which under normal circumstances binds platelet molecules together to create a patch over damaged walls of blood vessels. Because the platelet patch can become too large and also block blood flow, locally and downstream, aspirin is also used long-term, at low doses, to help prevent heart attacks, strokes, and blood clot formation in people at high risk for developing blood clots. It has also been established that low doses of aspirin may be given immediately after a heart attack to reduce the risk of another heart attack or of the death of cardiac tissue Aspirin was the first discovered member of the class of drugs known as nonsteroidal anti-inflammatory drugs (NSAIDs), not all of which are salicylates, although they all have similar effects and most have inhibition of the enzyme cyclooxygenase as their mechanism of action. Today, aspirin is one of the most widely used medications in the world, with an estimated 40,000 tones of it being consumed each year, the chemical structure of aspirin is shown below [66].
2-13-2 Paracetamol
53
Paracetamol or acetaminophen is a widely used over-the-counter analgesic (pain reliever) and antipyretic (fever reducer). It is commonly used for the relief of headaches, other minor aches and pains, and is a major ingredient in numerous cold and flu remedies. In combination with opioid analgesics, paracetamol can also be used in the management of more severe pain such as post surgical pain and providing palliative care in advanced cancer patients. The onset of analgesia is approximately 11 minutes after oral administration of paracetamol, and its half-life is 1–4 hours. While generally safe for use at recommended doses (1,000 mg per single dose and up to 4,000 mg per day for adults, up to 2,000 mg per day if drinking alcohol, acute overdoses of paracetamol can cause potentially fatal liver damage and, in rare individuals, a normal dose can do the same, the risk is heightened by alcohol consumption. It is the active metabolite of phenacetin, once popular as an analgesic and antipyretic in its own right, but unlike phenacetin and its combinations, paracetamol is not considered to be carcinogenic at therapeutic doses. The words acetaminophen (used in the United States, Canada, Japan, Hong Kong, Iran and Colombia) and paracetamol (used elsewhere) both come from chemical names for the compound: para-acetylaminophenol and para-acetylaminophenol. In some contexts, it is simply abbreviated as APAP, for acetyl-paraaminophenol.The chemical structure of paracetamol is shown below [67].
2-13-3 Mefenamic Acid Mefenamic acid is a non-steroidal anti-inflammatory drug used to treat pain, including menstrual pain. It is typically prescribed for oral administration. 54
Mefenamic acid is marketed in the USA as Ponstel. Mefenamic acid decreases inflammation (swelling) and uterine contractions by a still unknown mechanism. However, it is thought to be related to the inhibition of prostaglandin synthesis. There is also evidence that supports the use of mefenamic acid for perimenstrual migraine headache prophylaxis, with treatment starting 2 days prior to the onset of flow or 1 day prior to the expected onset of the headache and continuing for the duration of menstruation. Since hepatic metabolism plays a significant role in mefenamic acid elimination, patients with known liver deficiency may be prescribed lower doses. Kidney deficiency may also cause accumulation of the drug and its metabolites in the excretory system. Therefore, patients suffering from renal conditions should not be prescribed Mefenamic acid. The chemical structure of Mefenamic acid is shown below [68].
2-14 Case study of Arthritis (High Uric Acid) Uric acid (or urate) is a heterocyclic compound of carbon, nitrogen, oxygen, and hydrogen with the formula C5H4N4O3. Uric acid is a chemical created when the body breaks down substances called purines. Purines are found in some foods and drinks, such as liver, anchovies, mackerel, dried beans and peas, beer, and wine. Purines are also a part of normal body substances, such as DNA. In human blood plasma, the reference range of uric acid is between (3.6 mg/dL) and (8.3 mg/dL). This range is considered normal by the American Medical Association. Uric acid concentrations in blood plasma above and below the normal range are known, respectively, as hyperuricemia and hypouricemia. Similarly, uric acid concentrations in urine above and below normal are known 55
as hyperuricosuria and hypouricosuria. Such abnormal concentrations of uric acid are not medical conditions, but are associated with a variety of medical conditions. Excess serum accumulation of uric acid can lead to a type of arthritis known as gout [69]. This painful condition is the result of needle-like crystals of uric acid precipitating in joints and capillaries. Elevated serum uric acid (hyperuricemia) can result from high intake of purine-rich foods, and/or impaired excretion by the kidneys. Saturation levels of uric acid in blood may result in one form of kidney stones when the urate crystallizes in the kidney. Gout can occur where serum uric acid levels are as low as (6 mg/dL), but an individual can have serum values as high as (9.6 mg/dL) and not have gout [70].
56
Chapter Three 3-1 Materials And Chemicals 3-1-1 Materials The biomaterials, used in this work, were stainless steel 316L and Co – Cr – Mo alloy, and their chemical compositions are shown in the tables (3-1a) and (3-1b), which were obtained by Bruker advanced X-ray solutions D8-advance. The mechanical properties are shown in table (3-2a) and (3-2b) [71].
Table (3-1a): Chemical composition of Stainless Steel 316L obtained by XRF. Element
C
N
Mo
Ni
Mn
Cr
S
P
Si
Fe
0.03 0.05 3.00 12.0 1.50 16.0 0.01 0.03 0.75 Remained
Wt%
Table (3-1b): Chemical composition of Co – Cr – Mo alloy obtained by XRF. Cr
Mo
C
28.0
6.00
0.35
Element Wt%
Si
Co
1.00 Remained
Table (3-
2a): Mechanical properties of Stainless Steel 316L.
Material
SS 316 L
Material
Elastic
Tensile
Elongation
Fracture
Fatigue
Modulus
Strength
(%)
toughness
Strength
(GNm-2)
(MNm-2)
(MNm-2)
(MNm-2)
100
200-250
200
200-1100
40
Elastic
Tensile
Elongation
Fracture
Fatigue
Modulus
Strength
(%)
toughness
Strength
(GNm-2)
(MNm-2)
(MNm-2)
(MNm-2)
57
Co-Cr-Mo
230
450-1000
10-30
100
600
Table (3-2b): Mechanical properties of Co – Cr – Mo alloy 3-1-2 Chemical Solution The basic experimental solution was simulated human body fluid, which prepared by adding Ringer tablet to 0.5 liter of distillated water and heating the solution to temperature 120 0C for 15 min. and leaving it to cool. Then Na2HCO3 was added to obtain pH of 7.4. Ringer tablets were obtained from Merck Company Germany. The anti – inflammatory drugs, used in this work, were Aspirin (Acetylsalicylic Acid), Paracetol (Paracetamol), and Ponastan (Mefenamic acid) with three concentrations of each drug that were added by using Satorius BS/BT electronic with 0.1 mg accuracy balance (Fig. 3-1). Chemical formulas and concentrations of these drugs are shown in table (3-3). The concentrations of drugs were taken according to references [72, 73 and 74].
Fig. (3-1): Sartorius balance.
58
Table (3-3): Chemical formula and concentrations of anti – inflammatory drugs. [72,73and74] Drug
Aspirin
Paracetamol
Mefenamic acid
Chemical formula
Concentrations
Concentrations
g/ L
g/300 mL
0.00101
0.00303
0.0202
0.00606
0.0404
0.01212
0.01433
0.0086
0.02875
0.0172
0.115
0.0344
0.0037
0.00111
0.0052
0.00156
0.0067
0.00201
C9H8O4
C8H9NO2
C15H15NO2
3-2 Adjustment of pH and Temperature The normal pH and temperature of human body is 7.4 and 37 oC respectively, hence all experiments were adjusted to pH at 7.4 by using Eutech instruments Pte Ltd (Fig. 3-2), calibrated before any use by buffers.
Fig. (3-2): pH instrument. 59
Adjustment of temperature was performed by using MSC 400 Ceramic Hotplate Stirrer (Fig. 3-3). The MSC 400 is the perfect combination of a heating and stirring tool. The stirrer comes with 170x170mm ceramic coated hotplate suitable for temperatures of up to 400 oC. A specially coated layer of protection ensures safe and effective heating, as well as high resistance to chemicals.
Fig. (3-3): Heating instrument.
3-3 Specimen Preparation 3-3-1 Cutting Surface condition of specimen considers as an important role in corrosion resistance, hence, it is necessary to prepare uniform surface and careful specimen preparation. The stainless steel specimens were cut out to cylinder shape with dimensions of (1.266 mm) diameter and (10 mm) long but Co-Cr-Mo alloy had cubic shape with (1mm2) surface area and (1 mm) high
for
electrochemical tests.
3-3-2 Mounting The shaped specimens were molded using hot mounting with 2cm height and 2.226 cm diameter with leaving the topside of the specimen exposed as shown in Fig. (3-4).
60
The mounting process was performed by using XQ-2B mounting press, where the specimen was placed with phenolic resin in mold and heated up to 140oC under pressure of 3500 – 4000 psi, for 5 – 10 minutes. The instrument is shown in Fig. (3-5). for electrochemical studies, suitable provision was made on the other side for electrical contact.
Hole to electrical connection Epoxy resin 2.226 cm
Hole to electrical connection Epoxy resin 2.226 cm Exposed surface 2 cm
Exposed surface 2 cm
SS 316L
Co-Cr-Mo alloy
Fig. (3-4): Mounted specimens.
Fig. (3-5): XQ-2B mounting instrument.
3-3-3 Grinding and Polishing The mounted specimens were ground with SiC emery papers in sequence of 100, 180, 400, 600, 800, 1000, 1200, and 2000 grit to get flat and scratchfree surface.
61
The specimens were polished using polish cloth and alpha alumina 0.5µm and 1µm, and then washed with distilled water. The polished specimens were degreased with acetone trichloroethylene and cleaned in the same solution. The degreased specimens were washed with deionized water, dried and kept in a dissector over a silica gel pad and used for microstructure evolution and electrochemical investigation.
3-3-4 Etching Kroll’s reagent, containing 45ml of Glycerol, 15 ml of HNO3 and 30 ml of HCl, was used for etching the surface of stainless steel 316L for optical observation, while an etching solution for Co – Cr – Mo alloy contained 60 ml HCl, 15 ml Water, 15 ml acetic acid, and 15 ml HNO3 with time of exposure of 30 second for two alloys. [75, 76]
3-4 Corrosion Test 3-4-1 Electrochemical Cell The electrochemical standard cell is a versatile cell assembly kit, which allows performing a wide range of common electrochemical tests and experiments with the same base setup, supported by a variety of electrodes, and different accessories suitable for thermostating the electrolyte, passing gasses through the cell, or other special treatment. Either a 1 liter or a half-liter cell beaker, equipped with flat bottom and plane flange, is the core. It contains a variety of sleeve bores with different diameters, which may be used to insert the electrodes, electrolyte flow lines, gas sparging frits, thermometers or auxiliary electrodes. All parts except the electrodes are made of Duran heat resistant laboratory glass, PTFE (Teflon), or PEEK (Ryton), except the sealings (Fig. 3-6). 62
The main part of an electrochemical cell is the working electrode which consists of conductor rod and the mounted specimens. The reference electrode has to be connected using a Haber – Luggin – capillary. The accurate position of the capillary tip is essential to obtain proper results as shown in Fig. (3-7). A special saturated calomel electrode (SCE) fits into an integrated Harber – Luggin – capillary, giving best performance also if low – conductive electrolytes are used. Finally, Pt electrode is used in this research as an auxiliary electrode.
Fig. (3-6): Electrochemical cell.
Fig. (3-7): Auxiliary, Reference and Working electrodes [77]. 3-4-2 Corrosion Instrument 63
One of the important features of Autolab electrochemical instruments is the modularity (Fig. 3-8). The basis of the instrument is a potentiostat, galvanostat, which can be further configured to any needs by adding one or more of the available modules. All instruments come with the latest version of Nova software which allows performing a wide variety of electrochemical techniques as well as sophisticated data analysis and fit and simulation software as shown in Fig. (3-9). The entry level member of the modular Autolab instruments family, the PGSTAT128N is a low current, low noise and fast potentiostat/galvanostat capable of measuring maximum 800 mA, with a compliance voltage of 12V. The PGSTAT 128 N is designed for a wide range of electrochemical applications, from corrosion measurements to the characterization of energy storage devices. Table (3-4) shows the features of Autolab electrochemical instrument.
Fig. (3-8): Autolab electrochemical instrument.
Fig. (3-9): Autolab software 64
Table (3-4): Features of Autolab electrochemical instrument Key features: Electrode connections Potential range Compliance voltage Maximum current
2, 3 and 4 +/- 10 V +/- 12 V +/- 800 mA (10 A with BOOSTER10A)
Current ranges
10 nA to 1 A (100 pA with ECD module)
Potential accuracy Potential resolution Current accuracy Current resolution Potentiostat bandwidth Computer interface Control software
+/- 0.2 % 0.3 µV +/- 0.2 % 0.0003 % (of current range) 500 kHz USB NOVA
3-5 Metallographic Examination The tested specimens were examined microscopically using an optical microscope to examine the surface texture as a function of the change in current density in order to ensure that either the potentiostatic method represents reliable method for corrosion test or not. Some of the specimens were microstructure using the optical microscope Olympus BX2M, magnification used was 50 X.
65
Fig. (3-10): Optical microscope.
Chapter Four 4-1Corrosion Behaviour of Implants in Simulated Human Body Fluid Corrosion is a natural phenomenon where higher energy states of a metal attain equilibrium by transforming to such constituents as natural ore, which has lower energy states. As metal degrades, particles released initiate acute and chronic inflammatory responses and necrosis. The particles contaminate surrounding joint tissues, with major concentrations seen near the head of the prostheses in the trochanteric insertion of muscle and along the stem as an effect of bone circulation. Contamination of adjacent tissues depends on the type of implant and the state of the tissue. Since wear debris released from metal components ranges from 10 to 50 nm, it has a high surface area that increases the rate of corrosion. Their small size allows the metal particles to migrate large distances from the joint tissues surrounding the prosthesis, leading to systemic effects of the metal ions. Corrosion takes place in many known mechanisms, including crevice, fatigue, stress, fretting, galvanic corrosion, intergranular corrosion and pitting. Intergranular corrosion is the result of mechanical and heat treatments, crevice corrosion is due to sulphur in amino acids, and impurities in manufacturing of metal components of implants lead to galvanic corrosion. Corrosion is an important process since it contributes to the release of ions into the body. Different articulating surfaces have been used to find combinations that result in less corrosion, and therefore, less release of ions into the body [78]. Figure (4-1) shows the polarization curve for SS 316L and Co-Cr-Mo alloy in simulated human body fluid at pH=7.4 and temperature of 37oC. The 66
polarization behaviour shows two main regions including cathodic and anodic Tafel region, cathodic section (ab) represents the reduction reaction where reduction of oxygen can occur according to the following reaction [78]: O2 + 2H2O + 4e → 4OHˉ
…… (4-1)
While the anodic section, (bc) the dissolution of metals that can occur, as shown in the following reactions: The main anodic reaction in SS 316L: Fe → Fe2+ + 2e
(E=-0.44 Volt) ……..(4-2)
Ni → Ni2+ + 2e
(E=-0.25 Volt) ……..(4-3)
Cr → Cr3+ + 3e
(E=-0.75 Volt) ……..(4-4)
The main anodic reaction in Co-Cr-Mo alloy: Co → Co2+ + 2e Cr → Cr3+ + 3e
(E=-0.277 Volt) ……..(4-5) (E=-0.75 Volt) ……..(4-6)
Stainless steel and Co- based alloy materials are widely used as implant materials in clinical practice with each material having its own advantages. Since the metal release from implants is an important subject, numerous studies, including long-term clinical studies, have been conducted on metal release from orthopedic implants into body fluids (serum, urine, plasma, etc.). This metal release has been associated with clinical implant failure, osteolysis, cutaneous allergic reactions, and remote site accumulation. The increase in the incidence of allergy, and the necessity for prolonged use require implants having less metal release [79].
67
200 100
S.S alloy in HBF
vs SCE
c
Co-Cr-Mo alloy in HBF
0
c
-100 -200
E (mV)
-300 -400
b
-500 -600 -700
b
-800 -900 -1000
a
a
-1100 0.000100
0.001000
0.010000
0.100000
1.000000
log i (mA/cm^2)
Fig. (4-1): Potentiodynamic curve of SS 316L and Co-Cr-Mo alloys in simulated HBF The corrosion potential (Ecorr) of a material in a certain medium at a constant temperature is a thermodynamic parameter which is a criterion for the extent of the corrosion feasibility under the equilibrium potential (in opposite sign) of the cell consisting of the working electrode and the auxiliary electrode when the rate of anodic dissolution of working electrode material becomes equal to the rate of the cathodic process that takes place on the same electrode surface [80]. When (Ecorr) becomes more negative, the potential of the Galvanic cell becomes more positive and hence the Gibbs free energy change (ΔG) for the corrosion process becomes more negative. The corrosion reaction is then expected to be more spontaneous on pure thermodynamic ground. When the measured value of (Ecorr) becomes less negative, the potential of the corresponding Galvanic cell becomes less positive, hence the (ΔG) value for the corrosion process becomes less negative, and the process is thus less spontaneous at the same alloy [80].
68
The corrosion parameters were calculated by Tafel extrapolation method for two implant alloys in simulated human body fluid , as listed in Tables (4-1) and (4-2). The comparison between two alloys shows that the corrosion potential (Ecorr) for Co-Cr-Mo alloy is more active than corrosion potential of SS 316L. The corrosion current density (icorr) is a kinetic parameter and represents the rate of corrosion under specified equilibrium condition. Any factor that enhances the value of (icorr) results in an enhanced value of the corrosion rate on pure kinetic ground. The data of corrosion current density in tables (4-1) and (42) show that the value of (icorr) for Co-Cr-Mo alloy is lower than that for SS 316L.This means that the corrosion rate for Co-Cr-Mo alloy less than that obtained for SS 316L [80]. From deep analysis of the cathodic and anodic regions of the polarization curves which have been obtained for two implants in simulated human body fluid, it was possible to derive data concerning the cathodic (b c) and anodic (ba) Tafel slopes. Values of Tafel slopes (bc or ba) for both cathodic and anodic reactions were generally close to (≈ 120 mV.decade-1). A value of the cathodic Tafel slope of (-120 mV.decade-1), may be diagnostic of a proton discharge – chemical desorption mechanism in which the proton discharge is the rate – determining step. If the chemical desorption is the rate – determining step, the rate would be independent of the overpotential since no charge transfer occurs in such a step and the rate becomes directly proportional to the concentration or the coverage (θ) of adsorbed hydrogen atoms. On the other hand, if the discharge process is followed by a rate – determining step involving chemical desorption, the expected value of Tafel slope, in such step, would then be (≈ -30 mV.decade-1) [81]. When electrochemical desorption becomes the rate – determining step for hydrogen
evolution
reaction
on
the
cathode,
the
expected
value
of Tafel slope, in such step, would then be (≈ -50 mV.decade-1). Generally, the 69
values of bc for SS and Co-Cr-Mo alloy in tables (4-1) and (4-2) respectively show that the rate- determining step is proton discharge step. The value of Rp for Co-Cr-Mo alloy is higher than that observed for SS 316L in simulated human body fluid (HBF) without any other additives. This result was in agreement with the study by C. Valero [19], who deduced through the electrochemical behaviour of Co-Cr-Mo alloy, that this alloy has higher charge transfer resistance and lower capacitance which means thicker passive films than SS 316 L. The Co-Cr-Mo alloys are important implants because of their corrosion resistance, can be an order of magnitude greater than stainless steels and high mechanical properties [82, 83]. On the other hand, inactive ions, e.g. nickel and copper ions, do not immediately combine with water molecules and inorganic anions, and survive as an ionic state for relatively long time. Therefore, these ions have more chance to combine with biomolecules and reveal toxicity [7].
4-2 Effect of Anti-inflammatory Drugs on the Corrosion Behaviour Most drugs that are in clinical use are either weak bases or weak acids or their salts, due to their pH-dependent solubility and dissociation. Weak bases and weak acids have pH-dependent rates of drug dissolution substantial dissolution rate variations may become a problem in peroral drug delivery, if the compound has poor solubility in its un-ionized state and much higher solubility in its ionized state. PH varies considerably in the different parts of gastrointestinal tract (between 2 and 7) and, thus, dissolution rate may also change during the transit of the dosage form in the gastrointestinal tract. Osteolysis and aseptic loosening, or loosening in the absence of infection, are the major causes of failure of total hip replacements (THRs). Particles released from the implant as the femoral head articulates against the acetabular cup during movement have been found to be of a clinically relevant size (0.1-10 μm) that activates macrophages. The reaction to wear debris particles is type of 70
cell-mediated hypersensitivity reaction as result of excessive responses against foreign antigens. Because wear particles cannot be destroyed after phagocytosis, chronic inflammation results [84]. The pro inflammatory mediators produced by macrophages add to weakening of the connective tissue and bone surrounding the implant and results in osteolysis and aseptic loosening [85]. Metal – on – metal implant models release little debris but yield increased amounts of ions in the body. All metals in a biological environment like that of the body undergo corrosion, which results in the release of ions [86]. Duration of the implant and health of the patient depend on the fixation materials and articulating surfaces used in the prosthesis [87]. The behaviour and performance of materials are affected by the chemical, mechanical, biological, and bioelectrical events that occur in the environment of the body [88].Biofunctionality and biocompatibility of materials must be determined to predict their effect in the body. Non steroidal anti-inflammatory drugs (NSAIDs) are the therapeutic agents of first choice for the treatment of inflammation, pain, and fever. 4-2-1 Stainless Steel 316L Figure (4-2) shows effect of addition of three concentrations of Aspirin (C9H8O4) in simulated human body fluid (HBF) at pH=7.4 and Temp. =37 oC on corrosion behaviour of SS 316L. This figure shows that addition of Aspirin with three concentrations (0.00303, 0.00606, and 0.01212 g/300ml) shifts the corrosion potential (Ecorr) toward more negative direction (active) , but shift the corrosion current density (icorr) toward less values. Effect of presence of Paracetamol (C8H9NO2) on corrosion behaviour of SS 316L in simulated HBF is shown in Fig. (4-3), where it is noted that the lower and higher concentrations (0.0086 and 0.0344 g/300ml) shift the (Ecorr) value in the active direction, while the middle concentration 71
(0.0172 g/300ml) shifts in noble direction. All additive concentrations shift (icorr) to lower values. Fig. (4-4) shows the effect of addition Mefenamic acid (C15H15NO2) on the corrosion behaviour of SS 316L in HBF. This behaviour indicates that addition of Mefenamic acid shift (Ecorr) toward more active values and (icorr) toward lower values. Corrosion parameters in the presence of anti-inflammatory drugs and their effect on the SS 316L are listed in Table (4-1).The most accurate understanding of effect of drugs was noted through the values of polarization resistance (Rp) and corrosion rates (CR). Table (4-1): Corrosion parameters of SS 316L in the absence and presence of three anti-inflammatory drugs. Drugs
As. Pa. Me.
Conc. (g/300ml) HBF 0.00303 0.00606 0.01212 0.0086 0.0172 0.0344 0.00111 0.00156 0.00201
-Ecorr (mV)
icorr (A.cm-2)
437 677 576 632 632 224 667 671 606 630
1.693 e-5 1.413e-8 9.021 e-6 1.789 e-5 6.414 e-6 1.022 e-6 1.524 e-5 4.397 e-6 1.11 e-5 8.134 e-6
-bc ba -1 (mV.dec ) (mV.dec-1) 075 168 134 137 09 175 189 063 143 157
333 162 238 397 222 359 256 102 294 248
Rp (Ω.cm-2)
CR (mm/y)
5.229e2 2.508e6 1.251e3 1.081e3 1.063e3 2.112e4 1.091e3 5.219e2 1.341e3 1.643e3
1.752e-1 1.461e-4 9.331e-2 1.851e-1 6.634e-2 1.057e-2 1.558e-1 4.548e-2 1.148e-1 8.414e-2
Effect of three drugs on the corrosion of SS 316L in terms of polarization resistance is shown in Fig. (4-5 a, b, c), which indicates, that the drugs behave as inhibitive agents, in general, and have high polarization resistance values. This result is in agreement with the LeeAnn and co-workers [89] who studied the quantification of cellular viability and inflammatory response to stainless steel alloys. However, the success or failure of the implant is largely dependent on the design of the implant, surgical expertise, and corrosion resistance. Severe corrosion results in the release of toxic ions such as nickel which compromises the structural integrity of the device resulting in deleterious 72
clinical effects such as inflammation and the formation of fibrous capsules [90, 91]. Degrading materials also release wear debris, which are usually micrometer-sized particles that can be phagocytosed by monocytic cells [92]. 200 S.S alloy in HBF
100
S.S alloy in HBF+ 0.00303 As. S.S alloy in HBF+ 0.00607 As.
vs SCE
0
S.S alloy in HBF+ 0.01215 As.
-100 -200
E (mV)
-300 -400 -500 -600 -700 -800 -900 -1000 -1100 0.00000
0.00001
0.00010
0.00100
0.01000
0.10000
1.00000
log i (mA/cm^2)
Fig. (4-2): Potentiodynamic curve of SS 316L in simulated HBF in presence of Aspirin with three concentrations. 200 100
vs SCE
0
S.S alloy in HBF S.S alloy in HBF+ 0.0086 Pa. S.S alloy in HBF+ 0.0172 Pa. S.S alloy in HBF+ 0.0344 Pa.
-100 -200
E (mV)
-300 -400 -500 -600 -700 -800 -900 -1000 -1100 0.00001
0.00010
0.00100
0.01000
0.10000
1.00000
log i (mA/cm^2)
Fig. (4-3): Potentiodynamic curve of SS 316L in simulated HBF in presence of Paracetamol with three concentrations . 73
200
vs SCE
100
S.S alloy in HBF S.S alloy in HBF +0.00111Me.
0 -100
S.S alloy in HBF +0.00156Me. S.S alloy in HBF +0.00201Me.
-200
E (mV)
-300 -400 -500 -600 -700 -800 -900 -1000 -1100 0.00010
0.00100
0.01000
0.10000
1.00000
log i (mA/cm^2)
Fig. (4-4): Potentiodynamic curve of SS 316L in simulated HBF in presence of Mefenamic acid with three concentrations .
74
RP*103 (Ω.cm-2)
3000
a
2000 1000 0
CD (g/300ml)
RP*103 (Ω.cm-2)
RP*103 (Ω.cm-2)
b 30 20 10 0
c
2 1.5 1 0.5 0
0
0.0086
0.0172
0.0344
CD (g/300ml)
CD (g/300ml)
Fig. (4-5): Relationship between polarization resistance and Concentration of drugs for SS 316L in simulated HBF: (a) Aspirin, (b) Paracetamol, and (c) Mefenamic acid
75
4-2-2 Co – Cr – Mo Alloy Effect of addition of Aspirin, Paracetamol and Mefenamic acid on the corrosion behaviour of Co-Cr-Mo alloy in simulated HBF at pH=7.4 and temperature of 37oC is shown in Figs. (4-6), (4-7), and (4-8) respectively. Generally, this behaviour indicates that the addition of drugs shifts the (Ecorr) toward active direction except in some cases. Anti-inflammatory drugs, added in case of Co-Cr-Mo alloy in simulated HBF increases the value of (icorr) with all concentrations of drugs. The most accurate understanding of effect of drugs is shown in Fig. (4-9a, b, c), which indicates the relationship between polarization resistance and concentration of drugs. Table (4-2) includes the corrosion parameters of Co-Cr-Mo alloy in the absence and presence of anti-inflammatory drugs in simulated HBF. Table (4-2): Corrosion parameters of Co-Cr-Mo in the absence and presence of three anti-inflammatory drugs. Drugs
Conc. (g/300ml)
HBF 0.00303 As. 0.00606 0.01212 Pa. 0.0086 0.0172 0.0344 Me. 0.00111 0.00156 0.00201
-Ecorr (mV)
icorr (A.cm-2)
-bc (mV.dec-1)
ba (mV.dec-1)
Rp (Ω.cm-2)
CR (mm/y)
677 870 829 592 808 458 775 827 776 779
6.311e-7 2.785e-5 3.234e-5 1.369e-6 1.333e-5 8.287e-6 7.949e-6 1.554e-5 1.307e-6 5.507e-6
101 229 301 149 178 119 158 22 17 181
07 079 131 137 084 395 094 115 107 096
4.859 e3 2.839 e2 5.28 e2 6.471 e3 4.848 e2 2.46 e3 8.152 e2 7.076 e2 6.028 e3 1.366 e3
6.946 e-3 3.065 e-1 3.522 e-1 1.491 e-2 1.467 e-1 9.121 e-2 8.748 e-2 1.71 e-1 1.438 e-2 5.998e-2
Where the data indicates in general, increases of corrosion rate and different behaviour of polarization resistance in presence of drugs for Co-CrMo alloy in simulated HBF. These data indicate that the SS 316L has more resistance to corrosion in simulated human body fluid than Co-Cr-Mo alloy in the presence of drugs. This mean that there are interaction (affinity) between released metal ions from Co-Cr-Mo alloy and drugs molecules more than that observed for SS 316 L. All metal implants, however, corrode in the body at a rate determined, in part, 76
by their surface area [93]. In Co based alloy, cobalt is transported from tissue to the blood and eliminated in the urine within 48h, while chromium builds up in the tissues and red blood cells [94]. The only ion, taken up intracellular by red blood cells following corrosion of alloy is Cr6+, is then rapidly converted to Cr3+. Morais et al. [95] found that chromium and nickel are retained in bone marrow. Nickel is very small and has a low affinity for blood cells. Cobalt binds to both red blood cells and white blood cells. Although only very small quantities of Cr3+ bind to cells, Cr6+ binds very strongly to red blood cells and white blood cells [96]. In reversible electrochemical reactions, reactants are transformed to products and then the latter reacts together to produce the reactants. At equilibrium the rate of cathodic reaction (Rc) is equal to the rate of anodic reaction (Ra). This mean that consuming Mn+ in the reactions with any material in human body leads to the disequilibrium in the above equation and then increasing the reaction in the forward direction, i.e. increasing the dissolution of metals of implant, and namely increasing of corrosion rate. In this work Cr3+ ions represent the Mn+ and the materials in human body act as drug molecules which can produce organometallic complexes. This result is same as the result obtained by Eddie et al. [99] about cobalt complexes as antiviral and antibacterial agents through Co 2+complexes containing N, O donor ligands (CTC), which are used as class of drugs, as performed using a rabbit eye model infected with Herpes simplex Virus type 1 (HSV-1) . All complexes inhibited HSV-1 replication in vitro with as little as 5µg/ml required for strong antiviral activity. Structure of CTC- type Co3+ complex is shown below.
77
200 100
vs SCE
0 -100
Co-Cr-Mo alloy in HBF Co-Cr-Mo alloy in HBF+ 0.00303 As. Co-Cr-Mo alloy in HBF+ 0.00606 As. Co-Cr-Mo alloy in HBF+ 0.01212 As.
-200
E (mV)
-300 -400 -500 -600 -700 -800 -900 -1000 -1100 0.00001
0.00010
0.00100
0.01000
0.10000
1.00000
log i (mA/cm^2)
Fig. (4-6): Potentiodynamic curve of Co-Cr-Mo alloy in simulated HBF in presence of Aspirin with three concentrations . 200
vs SCE
100 0
-100
(14)
Co-Cr-Mo alloy in HBF Co-Cr-Mo alloy in HBF+0.0086 Pa. Co-Cr-Mo alloy in HBF+0.0172 Pa. Co-Cr-Mo alloy in HBF+0.0344 Pa.
-200
E (mV)
-300 -400 -500 -600 -700 -800 -900 -1000 -1100 0.00010
0.00100
0.01000
0.10000
1.00000
log i (mA/cm^2)
Fig. (4-7): Potentiodynamic curve of Co-Cr-Mo alloy in simulated HBF 78
in presence of Paracetamol with three concentrations .
200 Co-Cr-Mo alloy in HBF
100
vs SCE
Co-Cr-Mo alloy in HBF+ 0.00111 Me.
0
Co-Cr-Mo alloy in HBF+ 0.00156 Me. Co-Cr-Mo alloy in HBF+ 0.00201Me.
-100 -200
E (mV)
-300 -400 -500 -600 -700 -800 -900 -1000 -1100 0.00010
0.00100
0.01000
0.10000
1.00000
log i (mA/cm^2)
RP*103 (Ω . cm-2)
Fig. (4-8): Potentiodynamic curve of Co-Cr-Mo alloy in simulated HBF in presence of Mefenamic acid with three concentrations .
8 6 4 2 0
a
5 4 3 2 1 0
b
7 6 5 4 3 2 1 0
RP *103 (Ω.cm-2)
RP*103 (Ω.cm-2)
CD (g/300ml)
c
CD (g/300ml)
CD (g/300ml)
Fig. (4-9): Relationship between polarization resistance and concentration of drugs for Co-Cr-Mo alloy in simulated HBF: 79
(a) Aspirin, (b) Paracetamol, and (c) Mefenamic acid
4-3 Effect of Arthritis on the Corrosion Behaviour Figure (4-10) shows corrosion behaviour of SS 316L and Co-Cr-Mo alloy in simulated HBF in the presence of (0.7 g/L) uric acid, where this concentration represents the maximum of uric acid in male. The polarization behaviour indicates that the corrosion process in Co-Cr-Mo alloy is lower than in SS 316L according to thermodynamic and kinetic standpoint, i.e. corrosion potential and corrosion current density values. The data of polarization resistance and rate of corrosion enhances the above result of decreasing the corrosion in Co-Cr-Mo alloy compared with SS 316L in the presence of 0.7g/L uric acid. 200 100
S.S alloy in HBF+ 7mg/dL U.A Co-Cr-M0 alloy in HBF+ 7mg/dL U.A
vs SCE
0 -100
E (mV)
-200 -300 -400 -500 -600 -700 -800 -900 -1000 0.00010
0.00100
0.01000
0.10000
1.00000
log i (mA/cm^2)
Fig. (4-10): Potentiodynamic curve for SS 316L and Co-Cr-Mo alloy
[
in simulated HBF in presence of 0.7 g/L uric acid . 4-3-1 Behaviour of SS 316L with the Arthritis Disease Figs. (4-11) to (4-13) show the polarization behaviour of SS 316L in HBF in the presence of 0.7g/L U.A. (uric acid) in addition to the three antiinflammatory drugs Aspirin (As.), Paracetamol (Pa.), and Mefenamic acid (Me.). 80
These figures show the variation of Ecorr and icorr with concentrations of drugs. Generally, the polarization resistance increases in the presence of three drugs (Fig. 4-14a, b, c), i.e. decreasing of the corrosion rate. 200 100
vs SCE
0 -100
S.S alloy in HBF S.S alloy in HBF+7mg/dL U.A+ 0.00303 As. S.S alloy in HBF+7mg/dL U.A+ 0.00606 As. S.S alloy in HBF+ 7mg/dL U.A +0.001212 As.
-200
E (mV)
-300 -400 -500 -600 -700 -800 -900 -1000 -1100 0.00010
0.00100
0.01000
0.10000
1.00000
log i (mA/cm^2)
Fig. (4-11): Potentiodynamic curve for SS 316L in simulated HBF In presence of 0.7 g/L uric acid and Aspirin . 200 100
S.S alloy in HBF+ 7mg/dL U.A S.S alloy in HBF+ 7mg/dL U.A + 0.0086 Pa.
vs SCE
S.S alloy in HBF+ 7mg/dL U.A + 0.01725 Pa
0
S.S alloy in HBF+ 7mg/dL U.A + 0.0345 Pa
-100 -200
E (mV)
-300 -400 -500 -600 -700 -800 -900 -1000 -1100 0.00100
0.01000
0.10000
1.00000
log i (mA/cm^2)
Fig. (4-12): Potentiodynamic curve for SS 316L in simulated HBF in presence of 0.7 g/L uric acid and Paracetamol . 81
200 S.S alloy in HBF+ 7mg/dL U.A
100
S.S alloy in HBF+ 7mg/dL U.A + 0.00111Me.
0
S.S alloy in HBF+ 7mg/dL U.A + 0.00156Me. S.S alloy in HBF+ 7mg/dL U.A + 0.00201Me.
E (mV) vs SCE
-100 -200 -300 -400 -500 -600 -700 -800 -900 -1000 -1100 0.00010
0.00100
0.01000
0.10000
1.00000
log i (mA/cm^2)
Fig. (4-13): Potentiodynamic curve for SS 316L in simulated HBF in presence of 0.7 g/L uric acid and Mefenamic acid .
a
RP*103 (Ω.cm-2)
2 1.5 1 0.5 0
CD (g/300ml)
c
b
RP*103 (Ω.cm-2)
0.6
RP*103 (Ω.cm-2)
1.5 1 0.5
0.4 0.2 0
0
CD (g/300ml) CD (g/300ml)
Fig. (4-14): Relationship between polarization resistance and concentration of drugs for SS 316L in simulated HBF and 0.7g/L U.A: 82
(a) Aspirin, (b) Paracetamol, and (c) Mefenamic acid
Table (4-3): Corrosion parameters of SS 316L in the absence and presence of three anti-inflammatory drugs and 0.7g/L uric acid. Drugs
Conc. (g/300ml)
HBF +Uric acid 0.00303 As. 0.00606 0.01212 Pa. 0.0086 0.0172 0.0344 Me. 0.00111 0.00156 0.00201
-Ecorr (mV)
icorr (A.cm-2)
579 701 481 475 510 421 155 565 539 526
5.817e-6 1.46e-5 4.426e-6 5.63e-6 3.774e-6 5.967e-6 3.96e-6 3.898e-6 2.364e-6 2.76e-6
-bc ba -1 (mV.dec ) (mV.dec-1) 057 071 11 147 065 108 108 073 054 053
128 22 152 168 112 108 123 075 056 082
Rp (Ω.cm-2)
CR (mm/y)
4.288e2 3.668e2 1.293e3 1.503e3 6.641e2 6.699e2 1.146e3 4.826e2 4.372e2 5.442e2
6.017e-2 1.51e-1 4.578e-2 5.824e-2 3.903e-2 6.172e-2 4.096e-2 4.032e-2 2.445e-2 2.855e-2
4-3-2 Behaviour of Co-Cr-Mo Alloy with Arthritis Disease Figs. (4-15) to (4-17) indicate effect of three anti-inflammatory drugs in human body fluid (HBF) and 0.7 g/L uric acid. The data in table (4-4) indicates that the corrosion potential values, in general, shift toward more negative direction and variation of corrosion current values in the presence of three drugs. Polarization resistance values show that the presence of Aspirin in HBF and 0.7 g/L U.A increases the resistance to corrosion, but the presence of Paracetamol and Mefenamic acid decreases the polarization resistance. Fig. (418a, b, c) illustrates this behaviour. Generally, the resistance of Co-Cr-Mo alloy in HBF and 0.7 g/L is higher than the resistance of SS 316L as shown from the data listed in Tables (4-3) and (4-4).
83
vs SCE
200 100 0 -100 -200
E (mV)
-300 -400 -500 -600 -700 Co-Cr-Mo alloy in HBF+ 7mg/dL U.A
-800
Co-Cr-Mo alloy in HBF+ 7mg/dL U.A+ 0.00303 As.
-900
Co-Cr-Mo alloy in HBF+ 7mg/dL U.A+ 0.00606 As. Co-Cr-Mo alloy in HBF+ 7mg/dL U.A+ 0.01212 As.
-1000 -1100 0.00010
0.00100
0.01000
0.10000
1.00000
log i (mA/cm^2)
Fig. (4-15): Potentiodynamic curve for Co-Cr-Mo alloy in simulated HBF in presence of 0.7 g/L uric acid and Aspirin .
200
vs SCE
100 0 -100 -200
E(mV)
-300 -400 -500 -600 -700 -800
Co-Cr-Mo alloy in HBF+ 7mg/dL U.A Co-Cr-Mo alloy in HBF+ 7mg/dL U.A+ 0.0086 Pa.
-900 -1000
Co-Cr-Mo alloy in HBF+ 7mg/dL U.A+ 0.0172 Pa. Co-Cr-Mo alloy in HBF+ 7mg/dL U.A+ 0.0344 Pa.
-1100 0.00010
0.00100
0.01000
0.10000
1.00000
log i (mA/cm^2)
Fig. (4-16): Potentiodynamic curve for Co-Cr-Mo alloy in simulated HBF in presence of 0.7 g/L uric acid and Paracetamol .
84
200
vs SCE
100 0 -100 -200
E (mV)
-300 -400 -500 -600 -700 -800
Co-Cr-Mo alloy in HBF+ 7mg/dL U.A
-900
Co-Cr-Mo alloy in HBF+ 7mg/dL U.A+ 0.00111Me.
-1000
Co-Cr-Mo alloy in HBF+ 7mg/dL U.A+ 0.00201Me.
Co-Cr-Mo alloy in HBF+ 7mg/dL U.A+ 0.00156 Me.
-1100 0.00001
0.00010
0.00100
0.01000
0.10000
1.00000
log i (mA/cm^2)
RP*103 (Ω.cm-2)
Fig. (4-17): Potentiodynamic curve for Co-Cr-Mo alloy in simulated HBF in presence of 0.7 g/L uric acid and Mefenamic acid .
3 2 1 0
a
CD (g/300ml)
b
RP*103 (Ω.cm-2)
RP*103 (Ω.cm-2)
2
2
1
c
1 0
0 CD (g/300ml) CD (g/300ml)
Fig. (4-18): Relationship between polarization resistance and concentration of drugs for Co-Cr-Mo alloy in simulated HBF and 0.7 g/L U.A: (a) Aspirin, (b) Paracetamol, and (c) Mefenamic acid 85
Table (4-4): Corrosion parameters of Co-Cr-Mo alloy in the absence and presence of three anti-inflammatory drugs and 0.7 g/L uric acid. Drug s
Conc. (g/300ml)
HBF +Uric acid 0.00303 As. 0.00606 0.01212 Pa. 0.0086 0.0172 0.0344 Me. 0.00111 0.00156 0.00201
-Ecorr (mV)
icorr (A.cm-2)
-bc (mV.dec-1)
ba (mV.dec-1)
Rp (Ω.cm-2)
CR (mm/y)
066 223 215 223 344 331 205 700 187 215
3.652 e-6 2.803 e-6 3.839 e-6 2.803 e-6 2.661 e-6 5.202 e-6 4.832 e-6 1.213 e-5 3.812 e-6 4.688 e-6
092 076 118 076 07 115 097 071 068 101
157 156 164 156 067 129 193 091 165 124
1.731 e3 1.843 e3 2.181 e3 1.843 e3 7.57 e2 1.241 e3 1.691 e3 2.317 e2 1.283 e3 1.168 e3
3.977e-2 3.052 e-2 4.181 e-2 3.052 e-2 2.898 e-2 5.665 e-2 5.262 e-2 1.321 e-1 4.151 e-2 5.106 e-2
4-4 High Uric Acid up to 1.2 g/L Uric acid (weak acid), which has the chemical formula shown below, may be increased to high concentration in the human body. Fig. (4-19) illustrates the behaviour of SS 316L and Co-Cr-Mo alloy in human body fluid in the presence of 1.2 g/L uric acid where it is shown that the Co-Cr-Mo alloy has more resistance than SS 316L and has more noble Ecorr , less value of icorr and lower corrosion rate.
86
200 100
S.S alloy in HBF+12 mg/dL U.A Co-Cr-Mo alloy in HBF+12 mg/dL U.A
vs SCE
0 -100 -200
E(mV)
-300 -400 -500 -600 -700 -800 -900 -1000 -1100 0.00010
0.00100
0.01000
0.10000
1.00000
log i (mA/cm^2)
Fig. (4-19): Potentiodynamic curve for SS 316L and Co-Cr-Mo alloy in simulated HBF in presence of 1.2 g/L uric acid . 4-4-1 Behaviour of SS 316L at 1.2 g/L Uric Acid Behaviour of SS 316L in the presence of three drugs with 1.2 g/L uric acid is illustrated in Figs. (4-20) to (4-22) and table (4-5). The data of corrosion indicates that corrosion potentials in the presence of drugs shift toward more noble direction and variation of icorr and corrosion rate values, while the values of polarization resistance increases and the comparison among the resistances of drugs are shown in Figs. (4-23 a, b, c). It is noted there is directional proportionality between concentration of drugs and polarization resistance. 200 100 0 -100
S.S alloy in HBF+12 mg/dL U.A S.S alloy in HBF+12 mg/dL U.A+ 0.00303 As. S.S alloy in HBF+12 mg/dL U.A+ 0.00606 As. S.S alloy in HBF+12 mg/dL U.A+ 0.01212 As.
-200
E (mV)
-300 -400 -500 -600 -700 -800 -900 -1000 -1100 0.00001
0.00010
0.00100
0.01000
log i (mA/cm^2)
87
0.10000
1.00000
Fig. (4-20): Potentiodynamic curve for SS 316L in simulated HBF in presence of 1.2 g/L uric acid and Aspirin . 200 S.S alloy in HBF+12 mg/dL U.A
100
S.S alloy in HBF+12 mg/dL U.A+ 0.0086 Pa.
vs SCE
0
S.S alloy in HBF+12 mg/dL U.A+ 0.0172 Pa. S.S alloy in HBF+12 mg/dL U.A+ 0.0344 Pa.
-100 -200
E (mV)
-300 -400 -500 -600 -700 -800 -900 -1000 -1100 0.00010
0.00100
0.01000
0.10000
1.00000
log i (mA/cm^2)
Fig. (4-21): Potentiodynamic curve for SS 316L in simulated HBF in presence of 1.2 g/L uric acid and Paracetamol.
200
vs SCE
100 0 -100
S.S alloy in HBF+12 mg/dL U.A S.S alloy in HBF+12 mg/dL U.A+ 0.00111 Me. S.S alloy in HBF+12 mg/dL U.A+ 0.00156 Me. S.S alloy in HBF+12 mg/dL U.A+ 0.00201 Me.
-200
E (mV)
-300 -400 -500 -600 -700 -800 -900 -1000 -1100 0.00001
0.00010
0.00100
0.01000
0.10000
1.00000
log i (mA/cm^2)
Fig. (4-22): Potentiodynamic curve for SS 316L in simulated HBF 88
in presence of 1.2 g/L uric acid and Mefenamic acid .
a
RP*103 (Ω.cm-2)
2.5 2 1.5 1 0.5 0
b
2 RP*103 (Ω.cm-2)
RP*103 (Ω.cm-2)
CD (g/300ml)
1.5 1
c
1.5 1 0.5 0
0.5 0 CD (g/300ml)
CD (g/300ml)
Fig. (4-23): Relationship between polarization resistance and concentration of drugs for SS 316L in simulated HBF and 1.2 g/L U.A: (a) Aspirin, (b) Paracetamol, and (c) Mefenamic acid
Table (4-5): Corrosion parameters of SS 316L in the absence and presence of three anti-inflammatory drugs and 1.2 g/L uric acid. Drugs
Conc. (g/300ml)
HBF +Uric acid 0.00303 As. 0.00606 0.01212 Pa. 0.0086 0.0172 0.0344
-Ecorr (mV)
icorr (A.cm-2)
650 502 511 493 499 531 513
6.129 e-6 4.478 e-6 5.648 e-6 8.544 e-6 6.454 e-6 5.254 e-6 8.492 e-6
-bc ba -1 (mV.dec ) (mV.dec-1) 09 11 142 161 087 154 185 89
133 154 2 332 23 151 25
Rp (Ω.cm-2)
CR (mm/y)
6.741 e2 1.289 e3 1.727 e3 2.145 e3 1.059 e3 1.513 e3 1.868 e3
6.34e-2 4.632 e-2 5.842 e-2 8.837 e-2 6.675 e-2 5.435 e-2 8.783 e-2
Me.
0.00111 0.00156 0.00201
598 540 566
5.379 e-6 6.291 e-6 3.328e-6
129 113 113
147 208 113
1.21 e3 1.284 e3 1.31e3
5.564 e-2 6.507 e-2 3.443 e-2
4-4-2 Behaviour of Co-Cr-Mo Alloy at 1.2 g/L Uric Acid Behaviour of Co-Cr-Mo alloy in the presence of three drugs with 1.2 g/L uric acid is illustrated in Figs. (4-24) to (4-26) and table (4-6). The data of corrosion indicates that corrosion potentials in the presence of drugs shift toward more noble direction and, in general, increases the i corr values. while the values of polarization resistance increases and the comparisons between the resistances of drugs are shown in Figs. (4-27 a, b, c). The comparison between the two alloys shows that the Co-Cr-Mo alloy is more resistant than SS 316 L.This means that the Anti- Inflammatory drugs behave as inhibitors for biomaterials when the pH is decreased by metal ions releases and infection with an arthritis disease. These results are in agreement with the results observed by J.A. Von et al. [98] who indicated that oxytetracycline at addition levels of 0.01-1.0 mg/ml acts as an anodic corrosion inhibitor when they studied the antibiotic- metal interactions in saline medium. In another research, J.A. Von et al. [99] studied the effect of fused- ring antibiotics on metallic corrosion, where they observed that the variable depended upon the nature of the metal and its surface condition. Thomas Dorner studied the implant- related inflammatory arthritis, he found immune activation by an implant could have a key role in the development of arthritis. Implants have also been impacted in the development of allergic dermatitis and urticaria . In studies of metal sensitivity, immunesystem activation after implantation was seen either in response to the metallic implant itself or to particulate debris generated by the implant, metal particles are known to be related from miniplates into tissues. The degradation products of metallic biomaterials include particulate wear debris, organometallic complex, free metallic ions, inorganic metal salts or oxides and precipitated organometallic storage forms [100]. 90
200 100
vs SCE
0 -100 -200
E (mV)
-300 -400 -500 -600 -700 -800 Co-Cr-Mo alloy in HBF+ 12mg/dL U.A
-900
Co-Cr-Mo alloy in HBF+ 12mg/dL U.A+ 0.00303 As. Co-Cr-Mo alloy in HBF+ 12mg/dL U.A+ 0.00606 As.
-1000
Co-Cr-Mo alloy in HBF+ 12mg/dL U.A+ 0.001212 As.
-1100 0.00001
0.00010
0.00100
0.01000
0.10000
1.00000
log i (mA/cm^2)
Fig. (4-24): Potentiodynamic curve for Co-Cr-Mo alloy in simulated HBF in presence of 1.2 g/L uric acid and Aspirin .
200 100
Co-Cr-Mo alloy in HBF+ 12mg/dL U.A
vs SCE
Co-Cr-Mo alloy in HBF+ 12mg/dL U.A+0.0086 Pa.
0
Co-Cr-Mo alloy in HBF+ 12mg/dL U.A+0.0172 Pa. Co-Cr-Mo alloy in HBF+ 12mg/dL U.A+0.0344 Pa.
-100 -200
E (mV)
-300 -400 -500 -600 -700 -800 -900 -1000 -1100 0.00001
0.00010
0.00100
0.01000
0.10000
1.00000
log i (mA/cm^2)
Fig. (4-25): Potentiodynamic curve for Co-Cr-Mo alloy in simulated HBF in presence of 1.2 g/L uric acid and Paracetamol .
91
200 100
Co-Cr-Mo alloy in HBF+ 12mg/dL U.A
vs SCE
Co-Cr-Mo alloy in HBF+ 12mg/dL U.A+ 0.00111 Me.
0
Co-Cr-Mo alloy in HBF+ 12mg/dL U.A+ 0.00156 Me. Co-Cr-Mo alloy in HBF+ 12mg/dL U.A+ 0.00201Me.
-100 -200
E (mV)
-300 -400 -500 -600 -700 -800 -900 -1000 -1100 0.00001
0.00010
0.00100
0.01000
0.10000
1.00000
log i (mA/cm^2)
Fig. (4-26): Potentiodynamic curve for Co-Cr-Mo alloy in simulated HBF in presence of 1.2 g/L uric acid and Mefenamic acid .
RP*103 (Ω.cm-2)
4
a
2 0
4
b
RP*103 (Ω.cm-2)
RP*103 (Ω.cm-2)
CD (g/300ml)
3 2 1
3
c
2 1 0
0
CD (g/300ml)
CD (g/300ml)
Fig. (4-27): Relationship between polarization resistance and concentration of drugs for Co-Cr-Mo alloy in simulated HBF and 1.2 g/L U.A: (a) Aspirin, (b) Paracetamol, and (c) Mefenamic acid
92
Table (4-6): Corrosion parameters of Co-Cr-Mo alloy in the absence and presence of three anti-inflammatory drugs and 1.2 g/L uric acid. Drugs
Conc. (g/300ml)
HBF+Uric acid 0.00303 As 0.00606 0.01212 Pa. 0.0086 0.0172 0.0344 Me. 0.00111 0.00156 0.00201
-Ecorr (mV)
icorr (A.cm-2)
327 220 234 234 229 247 262 373 260 259
3.484 e-6 5.417 e-6 5.597 e-6 4.183 e-6 3.465 e-6 5.869 e-6 3.009 e-6 5.461 e-6 3.217 e-6 3.285e-6
bc ba (mV.dec-1) (mV.dec-1) 109 119 149 105 083 176 128 177 115 111
111 199 208 17 162 249 103 135 15 159
Rp (Ω.cm-2)
CR (mm/y)
1.507 e3 1.900 e3 2.410 e3 1.848 e3 1.682 e3 3.248 e3 1.900 e3 1.899 e3 2.320 e3 2.327e3
3.794e-2 5.899 e-2 6.095 e-2 4.555 e-2 3.773 e-2 6.392 e-2 3.277 e-2 5.947 e-2 3.503 e-2 3.577 e-2
4-5 Microstructure of Implants In order to obtain useful information about effect of anti – inflammatory drugs on the corrosion behaviour of implant biomaterials as well as showing the type of corrosion, the alloys surface were micrographed by using optical microscopy. Figs. (4-28) to (4-31) show the microstructure of SS 316L in human body fluid before and after addition of drugs and with and without etching. These figures show the presence of uneven shaped pits, distributed along the whole surface with appearance of austenite phase. Also one notes a difference in the size of pits and the color of the surface. This result shows the inhibition action of drugs. Figs. (40-32) to (4-35) show the microstructure of CoCr-Mo alloy in human body fluid before and after addition of drugs, and with and without etching. These figures show the presence of uneven shaped pits and appearance dendrites structure, but the drugs behave as corrosive materials.
93
Pits
𝛾-phase pits
(a) After immersion in HBF (b) After etching Fig.( 4-28): Microstructure of SS 316L in simulated HBF (X=50). Pits
𝛾-phase pits
(a) After immersion in HBF (b) After etching + Aspirin Fig.( 4-29): Microstructure of SS 316L in simulated HBF with Aspirin
(X=50). Pits
𝛾-phase pits
(a) After immersion in HBF (b) After etching + Paracetamol Fig.( 4-30): Microstructure of SS 316L in simulated HBF with Paracetamol (X=50).
94
Pits
𝛾-phase pits
(a) After immersion in HBF (b) After etching + Mefenamic acid Fig.( 4-31): Microstructure of SS 316L in simulated HBFwith Mefenamic acid (X=50).
Pits
𝐷𝑒𝑛𝑑𝑟𝑖𝑡𝑒𝑠 stricture phase)
(𝛾
(a) After immersion in HBF (b) After etching Fig.( 4-32): Microstructure of Co-Cr-Mo alloy in simulated HBF ( X=50).
95
Pits
𝛾- phase 𝐹𝑖𝑛𝑒 𝐷𝑒𝑛𝑑𝑟𝑖𝑡𝑒𝑠
(a) After immersion in HBF (b) After etching + Aspirin Fig.( 4-33): Microstructure of Co-Cr-Mo alloy in simulated HBF with Aspirin (X=50).
Pits
𝐹𝑖𝑛𝑒 𝑑𝑒𝑛𝑑𝑟𝑖𝑡𝑣𝑒𝑠 phase
𝛾
(a) After immersion in HBF (b) After etching +Paracetamol Fig.( 4-34): Microstructure of Co-Cr-Mo alloy in simulated HBFwith Paracetamol (X=50).
96
Pits
𝑐𝑜𝑎𝑟𝑠𝑒𝑟 𝑑𝑒𝑛𝑑𝑟𝑖𝑡𝑣𝑒𝑠 𝛾 phase
(a) After immersion in HBF + Mefenamic acid
(b) After etching
Fig.( 4-35): Microstructure of Co-Cr-Mo alloy in simulated HBF with Mefenamic acid (X=50).
97
Chapter Five 5-1 Conclusions The conclusions drawn from the research work are as follows: 1. Study of the corrosion resistance of SS 316 L and Co-Cr-Mo alloys in simulated human body fluid at pH=7.4 and Temp.=37oC show that Co-CrMo alloy is more resistant than SS 316 L alloy. 2. This work show that the drugs behave as inhibitors and increases the corrosion resistance of SS 316L while decrease in corrosion resistance of Co-Cr-Mo alloy. 3. The study shows that the drugs with 0.7 g/L uric acid behave as inhibitors for SS 316L while in the case of Co-Cr-Mo alloy the aspirin only act as inhibitor for corrosion. 4. In the presence of drugs with 1.2 g/L uric acid both alloys showed increase in corrosion resistance.
5-2 Suggestions for Further Studies 1. Study the corrosion behaviour in human body fluid by using various biomaterials such as cp Ti, Ti based alloy, and Ni-Cr alloy. 2. Addition alloying elements to improve the corrosion resistance. 3. Study the corrosion behaviour of SS 316L and Co-Cr-MO alloy under same condition of this work but in vivo. 4. Study the corrosion behaviour of SS 316L and Co-Cr-Mo alloy in human body fluid at various time of immersion. 5. Study the formation of passive layer on 316L and Co-Cr-Mo alloy in the absence and presence of drugs. 98
6. Study the corrosion behaviour in human body by using various drugs and chronic diseases. 7. Study some of mechanical properties such as wear, fatigue, and hardness which enhance the corrosion. 8. Study cyclic polarization under conditions of this work to predict the pitting and crevice potentials. 9. Study the corrosion behaviour for Implant alloys in mixture of two or more drugs.
99
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Dictionary of medical words Word Albumin
Translate الزالل
Word Masticatory
Translate مضغي
Anticoagulation
منع ختثر الدم
Mitral
خاص بالصمام التاجي
Antiphlogistic
مضاد االلتهاب
Mucin
سلاطني
Arthritis
التهاب ادلفاصل
Necrosis
موت اخلاليا
Arthroplasties
عملية تقومي ادلفاصل
Neurological
عصبية
Artificial cup
كاس اصطناعي
Nieghbouring
رلاورة
Aseptic
معقم
Orthobiologics
تقومي بايولوجي
Biocompatibility
التوافق احليوي
Orthopedic
جتبريي
Bovine serum
ادلصل البقري
Osseointegration
تكامل العظام
Cardiovascular
جهاز القلب واالوعية الدموية
Osteoporosis
هشاشة العظام
Culture Cell
خلية زرع
Osteoarthritis
التهاب ادلفاصل والعظام
Clot
جلطة
Osteolysis
حتلل العظام
Coagulation
ختثر
Palatine obturators
خلل والدي يف سقف الفم
cornea
قرنية
Parenchymal cell
متين اخللية
Crown
تاج
Peroral
االدوية عن طريق الفم
Dermatitis
التهاب اجللد
Physiological
فسيولوجي
Endentulous
ليس لديهم اسنان
Platelet
الصفائح الدموية
Endothelial
غشائي
prostheses
عضو اصطناعي
Exudation
نضح
Rupture
كسر
Femoral head
راس الفخذ
Saliva
لعاب
Gastrointestinal
اجلهاز اذلضمي
Scar tissue
نسيج نديب
Genototoxcity
السمية اجلينية
Secure
آمن
Gout
النقرس
Serum
مصل الدم
Germicide
مبيد اجلراثيم
Shinbone
الساق
Healing
الشفاء
Skeletal system
اذليكل العظمي
Hip joint
مفصل الورك
Spinal fixation
عقد شوكية
Immunology
علم ادلناعة
Subcutaneous
حتت اجللد
Infection
التهاب
Therapies
عالجات
Inflammation
التهاب
Thrombosis
ختثر
Tissue
نسيج
Toxicity
مسية
Trauma
رض- صدمة
Intramedullary nails مسامري داخل احلبل الشوكي Intravenous وريدي Intrusion
دخول
111
Ligament
رباط
Urine
البول
Macrophages
الضامة Translate جوار
Urology
علم امراض ادلسالك البولية Translate
Word Vicinity
وسط زلضر
Vitro Vivo
Word
وسط حي
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