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j.matchemphys.2018.08.060

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Accepted Manuscript
Serum albumin can influence magnesium alloy degradation in simulated blood
plasma for cardiovascular stenting
Fakiha El-Taib Heakal, Amira M. Bakry
PII:
S0254-0584(18)30725-9
DOI:
10.1016/j.matchemphys.2018.08.060
Reference:
MAC 20903
To appear in:
Materials Chemistry and Physics
Received Date: 4 May 2018
Revised Date:
15 August 2018
Accepted Date: 20 August 2018
Please cite this article as: F. El-Taib Heakal, A.M. Bakry, Serum albumin can influence magnesium alloy
degradation in simulated blood plasma for cardiovascular stenting, Materials Chemistry and Physics
(2018), doi: 10.1016/j.matchemphys.2018.08.060.
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Serum albumin can influence magnesium alloy degradation
in simulated blood plasma for cardiovascular stenting
Fakiha El-Taib Heakal*, Amira M. Bakry
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Chemistry Department, Faculty of Science, Cairo University, Giza 12613, Egypt.
ABSTRACT
Magnesium (Mg) alloys are promising materials for biodegradable cardiovascular stents due
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to their good biocompatibility and low thrombogenicity. Since dissolved Mg is unlikely to cause
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any adverse effects, magnesium materials become safe choice in humans. Herein, we report on
the in vitro degradable properties of a new bioresorbable implant for conceivable cardiovascular
stent application, namely AZ80 magnesium alloy. The degradation behavior in simulated blood
plasma (SBP) without and with 5_ 40 g/L bovine serum albumin (BSA) was assessed using open
circuit potential, electrochemical impedance spectroscopy and potentiodynamic polarization
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measurements. The results show that degradation susceptibility of the alloy at any exposure
period depends very much on the surrounding electrolyte composition. Additions of 10_20 g/L
BSA doses in SBP fluid lead to decrease the susceptibility of the alloy degradation due to
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forming an integrated protective adsorbed layer on the surface. Plausible explanations were given
for the stimulating effect induced by BSA at concentrations lower than 10 g/L or higher than
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20 g/L. The synergistic influence of the negatively charged adsorbed albumin molecules at the
physiological concentration (40 g/L) and Cl– ions in the contacting medium was also
investigated. The electrochemical results are very consistent with SEM and EDX surface
analyses, as well as the antibacterial efficiency of the alloy.
Keywords: Magnesium alloy; bovine serum albumin; impedance spectra; protein adsorption;
bactericidal.
E-mail address: feltaibheakal@gmail.com; fakihaheakal@yahoo.com: Tel.: +20 102449048.
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1. Introduction
Recently, there have been an up-surge of interest in the potential biomedical use of
magnesium and its alloys due to their advantages over traditional metallic and ceramics
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biomaterials or biodegradable polymers [1]. This is mainly attributed to their favorable strengthto-weight ratio, excellent biocompatibility and biodegradability, as well as adequate mechanical
properties [2]. Magnesium materials are chemically reactive as they can degrade naturally in the
physiological environment by corrosion and thus they are potential candidates in biodegradable
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hard-tissue implants [3]. The biocompatibility of a material is the main concern in the
development of biodegradable implants. The body should be able to resorb implanted material
and its degradation products should be non-toxic. Practically, we should choose materials that
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are abundant in the body [4], that is the case of Mg2+ as being the fourth most abundant cation in
the human body and largely stored mainly in bone tissues perhaps stimulating their formation.
The human body content of magnesium is approximately 35 g per 70 kg body weight and the
daily demand for magnesium is about 375 mg [5]. It is vital for metabolic processes such as
being co-factor in many enzymes and a key component of the ribosomal machinery that
translates the genetic information encoded by mRNA into polypeptide structures [6,7]. These
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intriguing characteristics provide magnesium-based materials with the potential to be used in
bio-implant orthopedic and cardiovascular devices [2,8].
Cardiovascular diseases refers to any pathology affecting the heart and blood vessels due to
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various etiologies such as fat deposition on vascular walls that narrow and weaken the vessels
leading to the formation of blood clots. Stent which is a small mesh-like tubular scaffold is
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inserted to open up the vessel [4] and keeping the lumen patency following its expansion inside
the artery. The current stent technology is based on using a permanent stent made from
corrosion-resistant metallic materials such as 316L stainless steel (316L SS), Ni–Ti and Co-Cr
alloy [9]. However, the role of stenting is temporary and limited to a period of 6–12 months after
which remodeling process is completed and the arterial walls heal [10]. Following this period,
the presence of in vivo stent cannot provide any beneficial effects [9]. Complementary surgery is
required to remove permanent metallic stents after healing, resulting in additional risk to the
patient and incurring additional hospital cost [4]. Therefore, biodegradable stents represent a
valid treatment [11], with their material should have at least the following characteristics: (i) It
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must be itself and its degradation products biocompatible. (ii) The material must stay in place for
several months before its complete bio absorption. (iii) The radial force of the resultant stent
must be enough for the integrity and scaffolding effect during the requested period [12]. Based
on these necessities, two classes of materials have been proposed for biodegradable stents:
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polymers [13] and metals; either Fe-based [14] or Mg-based [15] alloys. Bio absorbable
polymers, mostly polylactic and polyglycolic acid copolymers or polycaprolactone, represent a
possible approach to the development of fully degradable devices for stenting. However,
polymers can induce remarkable adverse reactions and have poor mechanical properties
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compared with metallic materials [16]. Other limitations seen in polymer stents include the
inability to expand completely with balloon dilatation along with restenosis rates similar to those
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observed for conventional bare metal stents. Additionally, polymers are typically viscoelastic
and can deform permanently over their lifetime when in use under stress, viscoelasticity is a
disadvantage of using polymers for stent applications [17]. Hence, biodegradable stents made of
metals gain more interest than their polymeric counterparts [16]. Whilst, Fe is corroded at a
sensible rate for stenting applications, but accumulates a voluminous corrosion product that
repels neighboring cells and biological matrix, and is not excreted or metabolized at an
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appreciable rate [17]. Therefore, the superior mechanical properties of Mg alloys in comparison
to polymeric materials make them good candidates for potential stent applications in small
vessels such as the coronary arteries, where a high strength to bulk ratio is required [2,15].
The use of magnesium material as a potential biodegradable stent was also based on the fact
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of being a structural constituent of the tissues and an essential element in the living organism.
Magnesium is a substantial intercellular cation, which is involved in more than 300 biological
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cell reactions. Magnesium is also regarded as a non-carcinogenic element and the generated
hydrogen gas from the Mg alloy corrosion will safely diffuse into the blood stream [18]. The first
application of magnesium in cardiovascular applications dates back to the year 1878, when Huse
[19] has used a Mg wire ligature successfully to stop bleeding vessels three times: once in a
radial artery and twice in operation for varicocele. Later in the 20th century, magnesium was
used in several investigations as biodegradable material for connectors of vessel anastomosis and
wires for aneurysm treatment [2]. Heublein et al. [15,20] were the first who investigated the idea
of using magnesium alloys for cardiovascular stents. Several Mg alloys have been investigated
for biodegradable stents including the magnesium alloy with 5.2–9.9 wt.% RE and 3.7–5.5 wt.%
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Y as claimed in a patent [21], AE21 (2 wt.% Al and 1 wt.% rare earth elements (RE)) [15],
AM60 (6 wt.% Al and 0.3 wt.% Mn) [22] and WE43 (4 wt.% Y and 3 wt.% RE) [23]. Other Mg
alloys preferably comprising 30–40 wt.% Li and 0–5 wt.% other metals [24], as well as a Mg–Li
alloy with a ratio in the range of 60 : 40 with further addition of Zn [25] were also patented.
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Recently, various coatings were applied to improve the corrosion resistance of degradable
magnesium alloys such as, sirolimus-eluting dextran and polyglutamic acid hybrid coatings on
AZ31 [26], AE42 magnesium alloy coated with silane and silane–PMMA [27], Mg-Zn-Y-Nd
alloy coated with arginine-leucine based poly (ester urea urethane) [28] and two different dual-
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phase ultra-high ductility Mg-Li-Zn alloys, LZ61 and LZ91 have been all reported [29].
The biodegradable stents are surrounded with tissue fluid or blood which contains different
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types of organic compounds such as serums and proteins besides inorganic species including Cl–,
HPO42–, HCO3– etc., that may absorb on the its surface and thus influence their degradation
properties [30]. The absence of this organic ingredients may be one reason for the variation in
the corrosion rates between in vitro and in vivo corrosion tests [31]. Fibrinogen, globulin, and
albumin are the main proteins present in the blood plasma. Around 55% of blood proteins are
accounted for by serum albumin with the normal concentration ranged between 30 to 50 g/L (3.0
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to 5.0 g/dL). Bovine serum albumin (BSA) is a model protein that has been widely used to
simulate the protein in the human body. The molecular weight of BSA is reported as equal to
66,382 g/mol, consisting of a single polypeptide chain of 585 amino acid residues with 17
interchain disulphide bonds [32]. Proteins are negatively charged and metal ions are positively
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charged in the body pH of 7.4. The interactions of metal ions with proteins generate colloidal
organometallic complexes. These complexes change the pH of albumin solutions and may
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enhance the rate of metal dissolution and the corrosion rate [33], therefore the presence of
protein can either accelerate or inhibit the corrosion rate, depending on which metal and solution
is used [34,35].
To date, many studies have been conducted out on the corrosion properties of magnesium
and its alloys in solutions containing proteins [18,32,34-41]. Even so, no literature was reported
about its influence on the corrosion behavior of AZ80 alloy, however, relatively very few studies
were found on its degradation activity [42-45]. On the other hand, Al3+ ions may have
neurotoxicity in the human body, yet considerations for utilizing commercial AZ-series
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magnesium alloys are based on their better corrosion resistance with preferred mechanical
strength and lower cost [44]. It is obvious that information are deemed necessary on this
important issue, therefore the present work is focused on exploring the corrosion behavior of
AZ80 alloy in simulated blood plasma (SBP) fluid and the influence of a wide range of serum
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albumin additions on the alloy performance during the early stage of alloy immersion, especially
over the first 1 h. To fulfil these objectives, various kind of electrolytes were prepared including
SBP, SBP + BSA at different concentrations of 5, 10, 20 and 40 g/L. The synergistic influence of
the negatively charged albumin and Cl- or OH- ions was also tested by adding 40 g/L BSA to 0.7
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wt.% NaCl single solution or in distilled water (DW) alone in order to understand how each
component affects the degradation performance of AZ80 alloy. The study was performed using
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electrochemical techniques complemented with SEM and EDX surface examinations. Moreover,
magnesium materials have drawn interest as antibacterial biomaterials, because of their ability to
alkalize the surrounding medium during the degradation process [46]. Hence, the antibacterial
activity of AZ80 alloy was likewise scrutinized and reported. It is hoped that the obtained in
vitro results would be of help to provide better understanding on the in vivo behavior of AZ80
alloy as a new promising biodegradable and biocompatible candidate for potential cardiovascular
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stent applications.
2. Experimental procedures
2.1. Materials and solutions
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The tested magnesium material in the present work was extruded AZ80 alloy with a nominal
chemical composition (wt.%) of: Al 8.2; Zn 0.46; Mn 0.13; Si 0.01; Cu ˂ 0.001; Fe ˂ 0.004; Ni
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˂ 0.001 and balance Mg. The working electrode was a strip cut from the alloy sample and
embedded in Araldit epoxy resin leaving its bottom surface area of 0.245 cm2 exposed. Before
each test, the electrode surface was abraded with different grades of emery papers from coarse to
fine one (400 to 1500 grit), degreased with acetone, rinsed with deionized water and air dried,
then immersed quickly in the quiescent test solution. The following solutions were used as
testing electrolytes under normal aerated conditions.
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•
SBP stock solution of pH 7.4 with a chemical composition as indicated in Table 1
[2,37,47], was prepared from analytical grade chemicals without further purification
using deionized water.
SBP solutions containing BSA at concentrations of 5, 10, 20 and 40 g/L. The BSA was
purchased from Biowest (South America) sterile filtered.
0.7 wt.% NaCl solution.
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0.7 wt.% NaCl solution containing 40 g/L BSA.
•
Distilled water (DW) containing 40 g/L BSA.
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•
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•
The added BSA doses into the SBP fluid were chosen to include low and high concentrations
up to 40 g/L. This was based on the reference range for serum albumin of 35–50 g/L as per
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World Health Organization [48]. According to the albumin supplier, the maximum solubility
of albumin in water is 50 g/L, so the upper concentration limit used in the present work was
40 g/L to avoid its precipitation in the SBP fluid. The designations and composition of all
tested solutions are as given in Table 2.
2.2. Electrochemical measurements
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Open-circuit potential (OCP), electrochemical impedance spectroscopy (EIS) and
potentiodynamic polarization were used as electrochemical testing techniques. The
measurements were performed in a conventional three-electrode cell assembly with the alloy
sample as the working electrode, saturated calomel electrode (SCE) and a large platinum sheet of
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size 40 mm × 20 mm × 2 mm were set as a reference and counter electrodes, respectively. The
measuring instrument was an electrochemical workstation IM6e Zahner-elektrik, GmbH,
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Koronach, Germany, controlled by a computer with Thales software for I/E and impedance data
measurements/analyses. Impedance measurements were recorded at the OCP in the frequency
range 5 × 104 Hz to 10-2 Hz using a sinusoidal ac excitation signal of 10 mV peak to peak. In
each BSA-containing electrolyte, a series of three consecutive measurements were taken after 5,
30 and 60 min. from the sample immersion. Tafel polarization E-log i plots were scanned in the
potential domain from -1.7 V up to -1.0 V using a scan rate of 1 mV s-1. All electrochemical tests
were done in stagnant aerated solutions at the average human body temperature of 37 ± 0.2 oC
inside an air thermostat system. Measurements were performed in triplicate with fresh solutions
and newly abraded electrode surfaces to obtain reproducible results.
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2.3. Surface characterization
Surface microstructure morphologies were examined using FE-scanning electron
microscopy (SEM) and the chemical composition of the corrosion product layer was determined
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using energy dispersive X-ray (EDX) spectroscopy after 1 h from sample immersion in SBP at
37 oC without and with different concentrations of BSA. The micrographs were collected using
Quanta 250 FEG (Field Emission Gun) equipped with EDX unit (Energy Dispersive X-ray
Analyzer), with an accelerating voltage 30 kV and magnification 14x up to 1000 000x and
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resolution for Gun.1n.
2.4. Antimicrobial test against S. aureus
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Gram-positive Methicillin resistant Staphylococcus aureus (S. aureus, ATCC 6538) was
used as a model to evaluate in vitro antimicrobial properties of AZ80 sample estimated by the
spread plate method. This particular strain was selected based on the recommendations of the
Clinical and Laboratory Standards Institute (CLSI) [49]. The sterilized media was poured onto
the sterilized Petri dishes (20-25 ml, each petri dish) and allowed to solidify at room temperature.
Microbial suspension was prepared in sterilized saline equivalent to McFarland 0.5 standard
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solution (1.5×105 CFU/mL) and its turbidity was adjusted to OD = 0.13 using spectrophotometer
at 625 nm. Optimally, within 15 min. after adjusting the turbidity of the inoculum suspension, a
sterile cotton swab was dipped into the adjusted suspension and was flooded on the dried agar
surface then allowed to dry for 15 min. with lid in place. Wells of 6 mm diameter were made in
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the solidified media with the help of sterile borer. 100 µL of the tested electrolyte was added by
micropipette to each well. The plates were incubated at 37 °C for 24 h. This experiment was
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carried out in triplicate and zones of inhibition were measured in mm scale. In order to determine
the antibacterial activities of AZ80 alloy against S. aureus, alloy extracts were prepared by
steeping the alloy samples in SBP0, SBP10 and SBP40 solutions for 24 h at 37 oC.
3. Results and discussion
3.1. Open circuit potential measurements
The open circuit potential (OCP) was monitored over a period of 1 h in all tested electrolytes
as presented in Fig. 1(a, b). These results can provide valuable information on the natural
corrosion behavior of the system in the absence of any external current or potential. In all cases,
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the development of OCP (EOC) reveals similar variation patterns where it increases rapidly at first
and then gradually until achieving a plateau with an almost stable value. The quasi-steady
potential value is reached after a certain interval of time very dependent on the electrolyte
composition. The net positive drift in the electrode potential assumes passive film formation due
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to deposition of some corrosion products that can effectively seals active anodic sites on its
surface against further reaction, and proving that the corrosion process is under anodic control.
Despite the EOC value observed during the early initial 5 min. from electrode immersion, the
quasi steady state potential registered after 1 h allows us to rank all tested solutions based on the
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positive increase in EOC value following the descending order SBP10 > SBP20 > SBP0 >
SBP40 ≥ SBP5> NaCl > NaCl40 > DW40. This rank generally indicates that the effect of
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BSA on AZ80 alloy corrosion in SBP fluid is concentration dependent, which is not quite
straightforward, as will be further explored from the ac and dc results.
Nevertheless, the obtained results can be explained alongside Brasher’s rule [50] which
stated that corrosive or inhibitive properties of an electrolyte are a function of its concentration in
the solution. Thus, at a concentration level of 10 to 20 g/L the present results showed that BSA
could play a good role in inhibiting the corrosion attack of AZ80 sample in SBP fluid. On the
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other hand, BSA concentration limits lower than 10 g/L and higher than 20 g/L give more
negative EOC values than that recorded in the blank SBP0 electrolyte proposing an increase in the
corrosion susceptibility of the sample. This negative shift is possibly due to chelating effect and
forming soluble metal complexes [51] at the low BSA dose (< 10 g/L), and/or to protein
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aggregation [52] at the higher dose (> 20 g/L). Additionally, at the same BSA content of 40 g/L
(which mimics the protein level in human blood plasma), the recorded EOC in SBP40 is higher
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than in NaCl40 or in DW40 solution. The results demonstrate that, not only the BSA dose which
determines the corrosion behavior of the implant, but relevant ions prevailing in the surrounding
physiological environment are also of such importance as will be discussed in the next sections.
3.2. EIS measurements
3.2.1. Behavior in SBP blank solution
EIS is a powerful non-destructive sensitive tool widely apply to provide valued information
about the corrosion kinetics of electrochemical systems and deduce their mechanisms [53]. In the
present work, EIS was performed at the OCP to investigate the time-dependency of the
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degradation behavior for AZ80 alloy in SBP fluid over a time period extended to 3 h. Fig. 2(a, b)
presents the obtained Bode and Nyquist diagrams for the impedance spectra as a function of the
immersion time. The Bode plots (Fig. 2a) confirm the presence of a resistive region at the low
frequency (LF) limit from ~1 Hz down to 10-2 Hz. The linear impedance diagonal shown at the
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medium frequency (MF) range (> 1.0 Hz) indicates a clear capacitive behavior commensurate
with a primary phase maximum at MF and a second peak maximum appeared at the high
frequency end (~ 104 Hz). It is also evident that the magnitude of the impedance modulus (|Z|) at
the LF range increases with the immersion time up to 1 h indicating self-passivation of the
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sample due to the growth of the partially protective film on its surface. Afterward, this film
suffers from dissolution in the environment with a concomitant decrease in the impedance of the
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solid/liquid interface when the exposure time is protracted more than 1 h.
Meanwhile, the impedance diagrams on the Nyquist format (Fig. 2b) are characterized by
two merged depressed semicircles at the high and medium frequency regions. The size of both
HF and MF merged capacitive loops first increases with time up to 1 h indicating an enhanced
capacitive behavior, and then continuously diminishes after that. The reduction in the diameter of
the capacitive loop after 1 h immersion implies an increase in the corrosion rate due to possible
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rupture in the quasi-passive film initially formed on the alloy surface [2]. This trend corroborates
well with the behavior of the Bode plots, where the corrosion resistance of the alloy firstly
increases resulting from the rapid growth of a partially protective Mg(OH)2 layer upon
immersion. However, this quasi-passive layer formed on the alloy surface cannot withstand
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longer and its resistance is decreased after a certain time period (> 1 h), due to a possible attack
by aggressive Cl– ions in the medium [39]. Hara et al. [54] reported that surface films on Mg and
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its alloy mainly composed of Mg(OH)2 rapidly dissolve in solutions containing Cl− and SO42−
ions. It follows that for AZ80 alloy in SBP fluid the following reactions can prevail during the
early stage of electrode immersion:
Mg → Mg2+ + 2e
(anodic reaction)
(1)
2H2O + 2e → 2OH- + H2↑
(cathodic reaction)
(2)
Mg2+ + 2OH- → Mg(OH)2
(net corrosion product)
(3)
This suggests that corrosion performance of AZ80 alloy in SBP fluid should be attributed to the
dissolution-precipitation process of Mg(OH)2 film layer. Additionally, HCO3– and HPO42– ions in
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solution can also take part in the surface reactions on the biodegradable Mg alloy, which limit
the advance of the corrosion attack. On prolonging the immersion time more than 1 h the
adsorbed notorious chloride ions can transfer Mg(OH)2 into soluble MgCl2 and thus stimulating
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the alloy degradation [42,55]:
Mg(OH)2 + 2 Cl− → MgCl2 + 2 OH−
(4)
The small amount of sulfate in the medium (0.5 mM) helps also to stimulate the alloy dissolution
[56].
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Several equivalent circuit (EC) models have been suggested to describe and analyze EIS
measurements using Thales software provided with the electrochemical workstation. The twotime constant EC model shown in Fig. 2b inset was found to be appropriate giving good
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conformity between experimental and theoretical data. The simulation results, which are also
embedded with the experimental data in Fig. 2, reveal excellent agreement with an average error
of less than 0.7%. The proposed EC consists of a network from the solution resistance (RS) in
series with two RC parallel combinations. In this adopted model, the first time constant (R1C1)
defines the resistance and capacitance behavior at the LF range due to the formed quasi-passive
film on the alloy sample. The appearance of this LF time constant (R1C1) indicates that the
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formed corrosion products manifest certain degree of porosity as corrosion attack occurs on spots
where electrolyte can enter through the pores of the film to the metal surface [57]. On the other
hand, the second time constant (R2C2) expresses the HF range behavior related to charge transfer
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resistance and double layer capacitance at the base of pores or defects in the partially protective
surface film [58].
To achieve optimum-fit results with the experimental EIS data, a constant phase element
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(CPE) is considered instead of the pure capacitor (C1) to account for the non-ideal capacitive
behavior of the system due to surface roughness and non-homogeneity [2]. CPE is a special
electrical element whose impedance is a function of the angular frequency (
=2
rad s-1, f
being the applied frequency in Hz or s-1) and related to the ideal capacitor impedance as follows:
/(
where
/(
) =
(Ω
)
(5)
) is the frequency independent admittance of the CPE.
the idealized capacitance (C) of the film layer at
would be identical to
= 1 [42]. The complex operator
is the
square root of negative one. The deviation parameter α has a value varies between unity for
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purely capacitive behavior associated with ideally homogeneous surface and 0.5 for a porous
electrode [42], and its deviation from unity is an indication of deviation of Yo from C. The EC
impedance parameters in all tested electrolytes were estimated and organized in Table 3. At any
immersion time, it is noticeable that
meantime, both
indicating that resistance of partially protective
) dominates the corrosion rate during AZ80 degradation in SBP medium. In the
and
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surface film (
>
values increase with time up to 1 h and decrease afterward. Thus, the
total resistance of the surface film (RT = R1 + R2) typically increases from 3.56 kΩ cm2 after
3 h from continuous immersion in SBP0 solution.
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5 min. reaching a maximum value of 37.1 kΩ cm2 after 1 h then decreases to 25.8 kΩ cm2 after
Based on these results, one can envisage that partially protective films on Mg alloy grow in
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two sequential stages. During the first one film thickness increases continuously with immersion
time and hence its protective efficacy increases, whilst in the second stage the growth rate
decreases with time due to film dissolution. Once the growth rate becomes less than the
dissolution rate, the corrosion resistance of the quasi-passive film decreases [40]. This
conclusion accords well with the white precipitate observed on the reacted surface. This is being
likely composed from Mg3(PO4)2 (Ksp= 1.04 × 10-24) and Ca3(PO4)2 (Ksp= 2.07 × 10-33) or
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Ca10(PO4)6(OH)2 (Ksp= 6.3 × 10-54), when the solubility limit is exceeded, as a product of the
surface reaction with HCO3–, Ca2+ and HPO42– ions prevailing in the medium. However, due to
the low HPO42– concentration (1 mM) in SBP0 solution, the formed insoluble deposit is loose
and cannot offer sufficient protection to the metal substrate for longer time. Hence, after 1 h from
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sample immersion RT value starts to decrease as Cl- ions near the surface can ingress through
pores and defects of the quasi-passive layer leading to increase the alloy degradation by forming
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the more soluble MgCl2 compound.
3.2.2. Effect of bovine serum albumin
Real blood plasma fluid contains beside inorganic ions many organic components such as
serums and proteins. The largest portion of the proteins in the human body is albumin that can
affect the dissolution properties of the metallic implant. Therefore, the tested electrolytes herein
are close to the real conditions by combining inorganic compounds with BSA as a model protein.
The influence of BSA addition on the corrosion performance of AZ80 alloy in SBP medium was
scrutinized in the present experiments at different protein doses of 5, 10, 20 and 40 g/L as a
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function of immersion time at 37 oC. EIS experimental results which are presented as Bode
diagrams in Fig. 3(a − d) demonstrate two resistive regions at low and high frequency ranges.
Additionally, all spectra are characterized by a capacitive contribution associated with a single
broad peak in the middle frequency range that corresponds to the maximum in phase shift (θ) vs.
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log f plots. Meanwhile, at any given BSA concentration the Bode diagrams reveal a significant
increase in both |Z| and θmax values with immersion time up to 1 h. The acquired Nyquist
diagrams shown in Fig. 3(e − h) have all similar pattern, where each plot exhibits an obvious
tendency to form two scarcely depressed capacitive loops with increasing size from the
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beginning of immersion up to 1 h. This feature indicates that the corrosion process proceeds by
the same mechanism with a decreasing rate over this exposure period. The difference in
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impedance results of the alloy in SBP solutions without and with BSA might be mainly
attributed to the difference in the kinetics behavior of BSA adsorption. It is worth noting that,
BSA molecule is mostly composed of amino acid linked by peptide bonds, where the amino acid
basic structure includes an amino group (–NH2), a carboxyl group (–COOH), a central carbon
and a side chain. In SBP physiological fluid of pH 7.4, albumin molecule undergoes a neutral–
acidic transition and becoming negatively charged as its isoelectric point laying at pH 4.5 – 4.7
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[38]. Protein is known to be adsorbed on the alloy surface through a complex process including
van der Waals, hydrophobic and electrostatic interactions, and hydrogen bonding [59]. The
presence of divalent cations such as Mg2+ or Ca2+ can serve as bridging agents to enhance
adsorption of albumin molecules on the implant surface during its immersion. The effects of
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adding 40 g/L albumin in either DW only or in 0.7 wt.% NaCl single solution were also
examined and presented in Fig. 4 as Bode and Nyquist diagrams. Fig. 4(c, d) discloses that
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Nyquist diagrams in NaCl40 and DW40 consist of two capacitive loops in addition to a very
small inductive loop in the lower frequency range. The diameter of the high frequency capacitive
loop at any immersion period > 5 min. is larger in DW40 solution than in NaCl40, while it is
much lower in neat 0.7 wt.% NaCl solution (Fig. 4c inset). This result corroborates well with the
Bode diagrams depicted in Fig. 4(a, b).
Comparison of all Nyquist diagrams for the impedance spectra recorded after the first hour of
AZ80 alloy exposure in the different tested electrolytes is shown in Fig. 5. As can be obviously
seen, the corrosion performance of the alloy in protein-containing single NaCl solution or DW is
inferior to that in the other tested complex physiological fluid. Nevertheless, serum albumin does
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inhibit effectively the alloy degradation induced by Cl− ions alone (see Fig. 5 inset). On the other
hand, in SPB fluid the size of the two merged capacitive loops allows us to rank the tested
electrolytes as this size are increased in the following descending order SBP10 > SBP20 > SBP0
> SBP40 > SBP5, being in good agreement with the ranking from OCP results (Section 3.1).
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These findings disclose the non-straightforward influence of BSA concentration on the alloy
degradation as well as the synergetic and mutual effects of HCO3−, HPO42−, SO42− and Ca2+ ions
in the physiological environment [56,60].
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To gain further insight into the influence of BSA on the corrosion susceptibility of AZ80
alloy in SBP fluid, as well as in simple NaCl solution or DW alone, the impedance results were
fitted to theoretical data using an appropriate EC model. Based on the above considerations, Fig.
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3e inset depicts the proposed EC Voigt model, which was found to give better fitting results to
the experimentally obtained impedance spectra with an average error of 0.2 − 0.8%. This fiveelement model with two series lumped time constants R1C1 and R2C2 is necessary for modeling
faradaic processes involving the adsorption of two different species, namely, corrosion products
and protein molecules [61]. In case of SBP containing BSA solutions, the first time constant
(R1C1) accounts for the behavior of the corrosion product layer coating the alloy surface as for
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the case in SBP0 solution (Fig. 2b inset). While, the second time constant (R2C2) describes the
behavior of the top adsorbed protein layer. The relevant equivalent circuit parameters were all
estimated and listed in Table 3 too. Careful inspection of the results in Table 3 and the
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representations shown in Fig. 6(a − d) reveals the following interesting inferences.
1. For all solutions containing BSA, both R1 and R2 values increase with increasing immersion
time, where R1is always much higher thanR2 (Table 3). This indicates that the corrosion product
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layer on AZ80 surface is more resistant and thus more protective than the outer adsorbed layer.
Furthermore, the increase in R1 and R2 values with time-lapse denotes that the two surface layers
grow up gradually and become more compact and hence more protective over the time. In
addition, as it can be noted in Fig. 6 (a and b) the total resistance (RT) of the surface layered film
decreases in the following order SBP10 > SBP20 > SBP0 > SBP40 > SBP5, which is very
consistent with the OCP results. Generally, protein adsorption from SBP solution on the alloy
surface occurs via a complex dynamic process during which protein may bind, rearrange and
detach. Protein-surface interaction are known to be an important factor in the corrosion behavior
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of biomaterial in physiological environment. However, protein can be adsorbed in different
orientations depending on its concentration, as well as on the physical and chemical
characteristics of the surface [37].
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2. At BSA concentration range of 10-20 g/L, the obtained results indicate that protein
collision is fast, thus decreases the time and/or room for attaining the optimized interaction
between BSA and the corrosion products on the alloy surface [33]. This will spawn in forming
an integrated protective adsorbed layer on the surface, which can effectively inhibit the
degradability of AZ80 sample in SBP fluid by blocking the reactive sites on the surface and thus
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increasing its RT value [39,40]. In this context, adsorption of protein has been previously
exploited to develop protective films in order to control Mg alloys degradation in simulated body
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fluid environment [62,63]. The adsorbed protein integrated layer could act as a physical barrier
between the metallic surface and the surrounding environment, hence inhibit the ingress of
aggressive ions [38,40] resulting in RT value increase.
3. At low BSA concentration (< 10 g/L), slow structural rearrangement with time and space
are allowable for the adsorbed albumin to accommodate the surface [33], which is prone to
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accelerate the alloy degradation via adsorption and a chelating-binding effect [32]. Actually,
BSA at the low-dose level in SBP5 solution seems to be insufficient to smear the entire alloy
surface. Instead, it is possible that negatively charged albumin molecules can be adsorbed
preferentially on the more electropositive reactive sites of the β-phase in AZ80 microstructure.
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This will disrupt the continuous network distribution of the second phases around the main α-Mg
matrix [42,43] and will decrease the cathodic-to-anodic area ratio, leading to stimulate its
degradation. The released Mg2+ ions interact with BSA to generate colloidal organometallic
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formulation [33,64], consequently decreases the protection ability of the surface layer and thus
lowering its RT value relative to the control RT in SBP0 solution.
4. Conversely, at high BSA concentration (> 20 g/L) there is a consensus that albumin may
aggregate in SBP40 solution leaving most of the alloy surface uncovered and exposed to the
physiological fluid. As such, this trait will enhance the anodic reaction rate and the cathodic one
by increasing the number of released electrons, based on the electroneutrality theory.
Accordingly, acceleration of the alloy corrosion reduces its RT value, which is in agreement with
the recent work by Hedberg et al. [52] who demonstrated that at physiological concentration of
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BSA (40 g/L) in PBS solution, there is a significant protein aggregation due to metal release and
subsequent precipitation during implant degradation process.
5. Regarding the two C1 and C2 capacitance values and their corresponding resistances R1 and
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R2, a reverse trend is clearly noticed. Thus, Fig. 6(c and d) shows that the inner layer capacitance
C1 is always much smaller than the adsorbed outer layer C2 suggesting that the inner layer is
much thicker than the outer one at any immersion time. Meanwhile, the higher C2 value in the
presence of BSA compared with its value in albumin devoid SBP0 solution (Fig. 6d) appears to
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be the reason of negatively charged adsorbed proteins in the outer layer at the prevailing
physiological conditions of temperature and pH. Furthermore, Table 3 and Fig. 6c reveal that in
the presence and absence of BSA, C1 values are all comparable with each other. This confirms
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that protein plays its critical role by interacting merely with the outermost layer as illustrated
schematically in Fig. 7.
3.3. Surface examination
Both surface film morphology and composition can provide us with significant information
concerning efficacy of the metallic implant under specific physiological environment. Fig. 8
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clearly proves that SEM images of AZ80 alloy exhibit quite different degradation properties in
SBP solutions without and with different BSA concentrations at 37 oC for 1 h. In fact, the
configuration of the corroded surfaces backings well the above obtained results from OCP and
EIS measurements. In absence of BSA, the sample immersed in SBP0 solution (Fig. 8a) shows a
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smooth surface with only some few detected local pits that can act as corrosion initiation sites.
Indeed, the alloyed Al metal in the sample can form a semi-protective Al-rich oxide layer
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improving its corrosion resistance followed by a reduction in the corrosion progression, near to
the lamellar β-phase (Mg17Al12) aggregate [65]. In addition, SBP fluid containing phosphates and
carbonates may form stable compounds covering the surface and helping to decrease the alloy
corrosion rate. On the other hand, in presence of BSA, the best dense and compact
microstructures in SBP10 and to lesser extent in SBP20 are shown in (Fig. 8c) and (Fig. 8d)
respectively. This means that addition of 10 g/L and 20 g/L BSA to SBP can decrease pitting
susceptibility and peeling of the oxide film. Both of them are beneficial to the corrosion-resistant
properties based on the partially protective film mechanism [40]. In comparison, the
microcorrosion morphologies after immersion in SBP5 (Fig. 8b) and SBP40 (Fig. 8e) solutions
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reveal severe corrosion attack with complete surfaces deterioration, especially the one immersed
in SBP5 solution containing 5 g/L BSA. As it is well established, second phase precipitations in
the alloy and/or impurities are highly cathodic to the alloy matrix phase rendering most
magnesium alloys vulnerable to pitting corrosion both in vitro and in vivo [66].
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Moreover, EDX elemental analysis allows us to compare between the corrosion products
detected on individual sample surface. The EDX spectra reveal that these products are mainly
composed of Mg, Al, Ca, P, Na, Cl, O, N and C. In absence and presence of BSA, the
appearance of carbon in the surface layer might stem from SBP solution and the added BSA.
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This leads to gradual increase in its content with increasing BSA dose in the medium as found
experimentally. As to Mg and Al peaks originate from AZ80 alloy itself, while those for O, Ca, P
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along with Mg are attributed to Mg(OH)2/CaP film formation [32]. It is also noticeable that Ca/P
ratio is almost close to one, confirming that calcium deposition in the corrosion layer only occurs
if phosphates are present in the physiological solution [67]. The presence of Ca and P in the
corrosion product may be indicative of the alloy corrosive behavior, such that increasing Ca and
P contents is related to decreasing in both Mg2+ evolution and alloy degradation [18]. It is well
known that calcium phosphate compounds generally show good compatibility and improved
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corrosion resistance [8]. A good observation in favor of that is the presence of a white product
coating, likely magnesium calcium apatite, which is clearly seen macroscopically on all samples
at the end of their exposure periods.
Likewise, the increasing trend of RT values as obtained from impedance measurements (Fig.
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6b) can be further confirmed by the decreasing order of both oxygen and chlorine contents in the
corrosion layer as follows: SBP5 > SBP40 > SBP0 > SBP20 > SBP10. Additionally, albumin
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adsorption can be possibly identified with EDX spectra by detecting some nitrogen in the
corrosion layer, because of the amino groups in BSA [30]. Nevertheless, in SBP10 and SBP20
solutions (Fig. 8c and d) nitrogen cannot be detected in the surface layers. This is probably due
to the growth of a protective layer over the adsorbed protein molecules that can shield N
appearance in the EDX spectra, as previously reported [67]. Accordingly, one can assume that
adsorbed negatively charged protein molecules can form an integrated protective layer with the
corrosion product Mg(OH)2/CaP deposit. This leads to formation of a compact mixed layer with
much less defects and higher RT value offering a good barrier to mitigate overall alloy
degradation [32,39,40].
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3.4. Potentiodynamic polarization measurements
Fig. 9a illustrates the potentiodynamic polarization (E-log i) behavior of AZ80 alloy in the
tested SBP solutions containing different BSA concentrations (0-40 g/L) after 1 h exposure
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under aerated conditions at 37 oC. For all cases, the plots disclose very small variation in the
shape of the cathodic polarization branches, while a distinct change in the anodic branches is
observed indicating anodic control behavior during metal dissolution and passive film formation.
In the presence of albumin, it should be especially noted a brief passivation zone ended by a
breakdown potential (Ebd), beyond which a sudden increase in the anodic current density occurs.
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The value of this transition potential is much dependent on the electrolyte composition, such that
a more positive Ebd value at a given BSA concentration means that the formed anodic passive
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film on the alloy is more stable in terms of its localized corrosion performance [55]. The
corrosion current density (icorr) value, as a direct parameter of the corrosion rate, was determined
by extrapolation of the cathodic Tafel region for each curve to the corrosion potential (Ecorr)
value [42,43]. Accordingly, the corrosion rate (Pi) was assessed from the correlation (Eq. (6)):
Pi = 22.85 icorr.
(6)
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Corrosion and passivation kinetic parameters derived from Fig. 9a are all listed in Table 4 as
a function of BSA concentration in the test electrolyte. It is noticeable that values of cathodic
Tafel slope (βc) are almost similar indicating same electrochemical reaction occurring in all
cases. This validates the generally accepted notion that hydrogen evolution reaction is the
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dominant local cathodic process during magnesium materials corrosion in neutral or alkaline
aqueous solutions through water reduction [55,68]. Further, the obtained results show a good
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correlation between the trend of Ecorr value obtained from the polarization method and the quasisteady Eoc value recorded from the OCP measurements after 1 h exposure. Clearly seen that, the
corrosion potential (Ecorr) value does not exhibit much variation upon albumin addition, in
agreement with Mueller et al. [37] observation, along with Kirkland and Waterman [69] results.
Thus, comparing with Ecorr of -1.436 V in SBP0 solution, this value increases to -1.417 V and
-1.420 V in SBP10 and SBP20 electrolytes, respectively. Padilla and Broson [70] have indicated
that the fractional coverage of the adsorbed albumin on the alloy surface usually increased with
increasing potential. Therefore, in case of 10 and 20 g/L BSA additions, the positive shift in Ecorr
may result in more adsorption of albumin molecules. In this regard, the small positive shift in
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Ecorr relative to Eoc value in any tested solution indicates that the overall kinetics of AZ80 alloy
corrosion in these electrolytes are simply under anodic control. An acceptable reason for this
effect is that Mg and its alloys exhibit the so-called negative difference effect (NDE), which is
defined by the difference between the hydrogen evolution (HE) rate on the metal at Ecorr and
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during anodic polarization. Since HE rate increases with increasing anodic polarization in case of
Mg and its alloys, this difference is negative number. Therefore, HE continues at a significant
rate even when Mg sample is anodically polarized above its Ecorr. The evolving hydrogen gas
acts as an insulator [71], so as it can be isolating the corroding part on the sample causing Ecorr to
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be more anodic than Eoc as found experimentally (see Table 4). Interestingly, the difference (∆E)
between these two potential values exactly coincides with the increasing trend being observed
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for icorr values as shown in Table 4. This confirms that increasing corrosion rate leads to more
anodic shift in Ecorr relative to OCP value because of the increase in HE on the sample surface,
leading the corrosion process to become more anodically controlled, such as the case in neat 0.7
wt% NaCl solution.
Compared to the corrosion rate exhibited by AZ80 alloy in SBP0 solution in term of icorr value
(2.82 µA cm-2), results in Table 4 show lower corrosion rates for the samples soaked in SBP10
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solution (1.01 µA cm-2) and in SBP20 solution (1.78 µA cm-2). On the other hand, higher
corrosion rates are obtained for the samples soaked in SBP5 solution (3.98 µA cm-2) and in
SBP40 solution (3.37 µA cm-2). This trend corroborates well with the reverse trend of RT values
as given in Table 3 and Fig. 6b indicating that corrosion rate of the tested alloy has a non-linear
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relationship with BSA concentration in SBP solution. In fact, BSA has a dual action on the
degradation susceptibility of the alloy depending on its concentration in the medium. Thus,
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within 10-20 g/L range, BSA inhibits AZ80 dissolution, while at concentrations < 10 g/L or > 20
g/L BSA can stimulate the corrosion process. On reviewing the literature, careful comparison
can be made between AZ80 alloy tested herein and Fe, Mg and their alloys tested for stent
applications are summarized in Table 5. The comparative data demonstrate that AZ80 alloy
exhibits low corrosion rate and good mechanical properties.
Polarization behavior of AZ80 alloy is also studied in 0.7 wt.% NaCl solution without and
with the physiological relevant protein concentration (40 g/L BSA), as well as in DW40 solution.
Results shown in Fig. 9b and Table 4 reveal a clear negative shift in Ecorr value of the alloy for
both NaCl40 and DW40 solutions. This is relative to that in the neat NaCl medium indicating
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cathodic control behavior in the presence of BSA. Besides, the synergistic effect between the
protein and Cl– ions is marked by the positive shift of Ebd value in NaCl40, as well as by the
decrease in both icorr and ipass in comparison to their values in the absence of protein. In DW40,
the apparent negative shift of Ebd for AZ80 can be the reason of the highly negative shift in its
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Ecorr value, being devoid of any Cl– ions. Even though, the icorr value in DW40 solution decreases
further to 4.56 µA cm-2, i.e. becomes almost one fifth its value in NaCl solution (20.50 µA cm-2),
while it decreases by a factor of one fourth only in NaCl40 medium (5.73 µA cm-2).
At this point, it is of interest to emphasis the synergistic role and mutual effect of the
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biologically relevant inorganic ions; Cl–, HPO42–, HCO3–, SO42– and Ca2+ present in the blood
plasma fluid on AZ80 corrosion products. As it has been previously reported, when AZ-type Mg
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alloy is exposed to an aqueous environment, the internal second phase being ~ 300 mV more
positive than the α-phase matrix acts as efficient cathode for hydrogen evolution while anodic
Mg dissolution is taking place. In this process, Mg(OH)2 film forms on the surface of the alloy
due to significant alkalization near the surface [72]. However, in the presence of Cl– ions, being
the most important part in SBP fluid, their size is small enough to displace water molecules in
the hydration sheath. Hence, Cl– ions can preferentially complex with Mg2+ cations and largely
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accelerate the Mg implant degradation by breaking the surface film integrity. The key factors
limiting progression of corrosion attack are the relatively fine β-phase (Mg17Al12) network and
Al enrichment produced on the corroded AZ80 surface [65]. Practically, the SO42– ions can also
attack the active anodic sites on Mg alloy and forming the fairly soluble MgSO4 salt and thus
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deteriorate the sample surface [73].
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However, HPO42– ions can decrease the degradation rate of the implant due to formation of
a dense magnesium phosphate layer (Mg3(PO4)2) which also delays the formation of pitting
corrosion [43,56,60]. In addition, Ca2+ ions in the SBP can precipitate as calcium phosphate
(Ca3(PO4)2) or hydroxyapatite (HAp) (Ca10(PO4)6(OH)2) when the solubility limit is exceeded
[43]. In this scenario, covering the reactive spots on the alloy surface is finally achieved leading
to reduction of its corrosion rate. On the other hand, rapid surface passivation can be induced via
MgCO3 formation when an appropriate amount of HCO3– ions is present after protracted time of
exposure that is more likely to occur in presence of Cl– ions [34].
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3.5. Antimicrobial measurements
The antibacterial efficiency of AZ80 immersed for 24 h in SBP0, SBP10 and SBP40
solutions was investigated against S. aureus. Fig. 10 shows the inhibition zones of S. aureus for
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these solutions where clear areas of inhibition zone can be seen for the three tested solutions.
Inhibition zone diameters for solutions extracted from the alloy immersion in SBP0 and SBP40
media are slightly larger than the one for SBP10 solution. This result may be related to the slow
degradation rate of AZ80 in SBP containing 10 g/L BSA. Thus, less degradation products will be
found in its extract solution. Whilst, fast degradation of the alloy should lead to an increase of
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Mg ion concentration and solution pH. Lock et al. [79] with other authors [80,81] stated the
bactericidal effect during magnesium alloying degradation process as being attributed to many
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factors among which; increased alkalinity and concentration of Mg2+ ions as well as magnesiumbased nanoparticles such as magnesium hydroxide and magnesium oxide particles.
In fact, alongside their good antibacterial activity magnesium and its alloy can fight bacterial
proliferation [82,83], adhesion and biofilm formation [84]. Robinson et al. [82] studied the
growth of Staphylococcus aureus, Escherichia coli and Pseudomonas aeruginosa in vitro, and
found that addition of Mg turnings to the culture broth resulted in a lower number of bacteria as
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compared to the stainless steel. Also, Lock et al. [83] evaluated the inhibitiory effect of Mg on
bacterial cell proliferation by seeded E. coli on pure Mg strips incubated for 16 h. The results
revealed decreased bacterial cell density as compared with polyurethane stents and glass slides.
On the other hand, Li et al. [84] demonstrated that Mg implant can reduce bacterial adhesion and
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prevent biofilm formation, most likely due to increased local alkalinity caused by metal
degradation. Likewise, He et al. [85] exhibited that all the Mg alloys used in their work not only
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inhibited the bacterial growth but also decreased the number of the adhered bacteria on the alloy
surface. Furthermore, the corrosion products formed during the alloy degradation process
contributed to bacterial death. Brooks et al. [86] supported these results by observing the in vitro
antimicrobial properties of AZ91 alloy. The results showed that the ability of AZ91 alloy to
reduce bacterial adhesion and subsequent biofilm formation. Admittedly, pH plays an important
antibacterial role due to the alkaline environment created during degradation of magnesium.
Such mechanism responsible for this process are explored and illustrated in details by Qin et al.
[87], confirming that the high alkalinity rather than release of Mg2+ ions would inhibit bacterial
growth and colonization where most bacteria can only survive in an appropriate pH range (from
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5.5–9). The high alkalinity can also affect the chemical and/or physical bacterial surface
properties and decrease bacterial surface hydrophobicity leading to inhibition of bacterial
adhesion and biofilm formation. Moreover, Rahim et al. [88] found that the influence of original
extracts of pure Mg on infectious bacteria, bioluminescent S. aureus or P. aeruginosa displayed
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good antibacterial activity and the proliferation of both types of bacteria was suppressed.
Furthermore, there was a dramatic decrease of antibacterial effect correlated with increased pH
levels of supernatant solution proving that alkaline pH alone was necessary and sufficient for the
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antibacterial effects of magnesium and its alloys.
4. Conclusions
In this work, in vitro degradation conditions closer to the physiological environment are
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implemented by the addition of different BSA doses to SBP fluid at pH 7.4 and temperature
37 oC. This was to explore the corrosion behavior of AZ80 alloy as a potential bioresorbable
implant using various electrochemical techniques. The conclusions drawn from the findings are:
1. In absence of BSA, corrosion resistance of the alloy in SBP0 initially increases up to 1 h due
to the rapid growth of partially protective Mg(OH)2 layer on the alloy surface. However, this
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quasi-passive layer cannot withstand longer. Hence, the corrosion resistance decreases rapidly
when immersion time is progressed to 3 h. Comparison of the present AZ80 alloy with Fe, Mg
and their alloys investigated for stent applications confirm that AZ80 alloy is a potential good
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choice due to its low corrosion rate (Pi = 77 µm/y) and better mechanical properties.
2. In presence of BSA, the total film resistance of the alloy (RT) increases at 10 g/L and 20 g/L
BSA additions and the OCP value moves to the positive direction. This is likely due to adsoption
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of negatively charged albumin molecules forming an integrated protective compact layer, which
acts as a physical barrier to mitigate the alloy degradation.
3. However, 5 g/L and 40 g/L BSA additions in SBP fluid lead to lowering the RT value and
stimulating the alloy corrosion susceptibility with a negative shift in the OCP value. These
findings imply that the blocking effect of the adsorbed protein layer becomes weaker, possibly
due to chelating effect and forming soluble metal complexes at lower BSA dose (< 10 g/L) and
to protein aggregation at higher BSA dose (> 20 g/L).
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4. The synergistic effects of negatively charged albumin molecules with Cl-, and OH- ions are
evaluated by testing the behavior in NaCl40 simple solution or in DW40 alone. The results
indicate that the competition between albumin adsorption molecule, OH-, and Cl- for sites on the
alloy surface can inhibit the in-leakage of Cl- and its corrosion attack on the alloy substrate under
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Mg(OH)2 layer.
5. Good agreement is observed between the results obtained from ac impedance and dc
potentiodynamic polarization techniques. Surface examination with FE-SEM images and EDX
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spectra as well as the antibacterial efficacy of AZ80 supports well the electrochemical results.
Acknowledgement
used to accomplish this work.
References
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The authors would like to thank Faculty of Science at Cairo University for providing facilities
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Table 1. Chemical composition of simulated blood plasma (SBP) solution [2,37,47].
Concentration (mM)
Ion
Reagent
Concentration (g/L)
Human blood plasma Simulated blood plasma
HPO42–
142
5
1.5
2.5
125
27
1
0.5
NaCl
KCl
MgCl2.6H2O
CaCl2
MgSO4. 7H2O
NaHCO3
Na2HPO4
---
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SC
SO42-
142
5
1.5
2.5
103
27
1
0.5
Table 2. Designation and composition of solutions tested in the present work.
Composition
SBP0
SBP5
SBP10
SBP20
SBP40
NaCl
NaCl40
DW40
Blank SBP solution
SBP solution + 5 g/L BSA
SBP solution + 10 g/L BSA
SBP solution + 20 g/L BSA
SBP solution + 40 g/L BSA
0.7 wt.% NaCl
0.7 wt.% NaCl solution + 40 g/L BSA
Pure deionized water + 40 g/L BSA
EP
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Designation
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C
6.6
0.37
0.20
0.28
0.12
2.27
0.14
---
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PT
Na+
K+
Mg2+
Ca2+
Cl–
HCO3–
ACCEPTED MANUSCRIPT
Table 3. Equivalent circuit parameters of AZ80 alloy in the different tested electrolytes as a function
of immersion time
C1
(µF cm-2)
R2
(kΩ cm2)
SBP0
RS
(Ω cm2)
1.01
0.34
0.24
0.16
0.14
15.7
12.3
12.7
12.2
16.0
0.32
4.92
10.4
23.9
15.3
15.9
8.2
7.4
6.9
5
30
60
120
180
3.51
32.0
36.9
29.7
25.2
0.16
1.41
1.24
0.97
0.83
0.74
0.72
0.70
0.68
0.67
5
30
60
2.08
8.33
15.6
4.65
2.41
2.19
0.71
0.59
0.59
5
30
60
6.18
34.7
58.6
3.18
2.00
1.79
0.73
0.70
0.69
0.77
5.10
10.6
29.6
28.7
25.8
10.7
9.7
9.5
5
30
60
4.44
16.6
37.5
3.05
2.09
1.72
0.68
0.64
0.63
0.86
8.21
12.0
23.9
17.9
17.6
10.6
10.0
9.7
SBP10
SBP40
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D
SBP20
0.05
0.12
0.24
0.42
0.60
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SBP5
3.58
17.0
20.6
3.13
2.05
1.64
0.69
0.66
0.64
0.83
6.12
7.99
25.50
20.70
20.33
9.4
7.6
6.5
1.07
0.19
0.19
9.34
10.9
10.8
0.58
0.82
0.83
0.77
0.83
0.82
11.38
11.57
11.29
9.1
13.2
13.2
5
30
60
2.22
3.35
4.53
8.79
7.59
7.11
0.66
0.70
0.71
1.03
1.56
1.93
10.73
11.45
11.01
13.4
14.7
15.8
5
30
60
1.61
4.22
5.91
5.13
4.79
4.40
0.64
0.68
0.67
0.64
1.39
2.33
11.18
14.19
13.36
236.9
253.6
265.6
5
30
60
NaCl
AC
C
EP
5
30
60
NaCl40
C2
(µF cm-2)
RI
PT
Time
R1
(min.) (kΩ cm2)
SC
Electrolytes
DW40
ACCEPTED MANUSCRIPT
Table 4. Electrochemical corrosion parameters of AZ80 magnesium alloys after 1 h immersion
in the different tested electrolyte at 37 oC; ∆E = Ecorr - Eoc.
( V)
( V)
-1.512
-1.542
-1.464
-1.471
-1.525
-1.565
-1.571
-1.607
-1.436
-1.454
-1.420
-1.417
-1.445
-1.305
-1.429
-1.464
icorr
∆E
(mV)
|βc|
-2
Ebd
-1
(µA cm )
(mV decade )
2.82
3.98
1.01
1.78
3.37
20.50
5.73
4.56
136
130
132
130
146
165
84
189
76
88
44
54
80
260
142
143
(V)
ipass
Pi
-2
(µA cm )
-1.392
-1.345
-1.384
-1.405
-1.241
-1.089
-1.352
3.46
1.02
1.23
2.24
50.5
18.1
5.29
(µm y-1)
64.4
90.9
23.1
40.7
77.0
468.4
130.9
104.2
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SBP0
SBP5
SBP10
SBP20
SBP40
NaCl
NaCl40
DW40
Ecorr vs. SCE
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Eoc vs. SCE
SC
electrolyte
Table 5. Mechanical and degradation properties of Fe, Mg and their alloys investigated for
cardiovascular stent applications in comparison with AZ80 alloy.
Fe
Fe-35Mn
Electrocasted
heated at 550 oC
annealed
S solution treated
Fe-30Mn-6Si
-
Elongation Solution Pi
(mm/y)
(%)
40
SBP
0.19
Ref.
[17,74]
54
270
290
18
Hank’s 0.46-1.22 [17,75]
-
230
430
30
SBP
0.43
[17,74]
-
180
450
16
Hank’s
0.30
[9,76]
SBP
407 ]
[17,74]
Cast
1.74
41
20
86
13
AZ31
extruded
1.78
45
125-135
235
7
HBSS
2.96
[26,77]
AM60B
die cast
1.78
45
-
220
6-8
SBP
8.97
[4,74,77]
AZ91
die cast
1.81
-
150
230
3
SBP
147.7
[4,74,77]
extruded
extruded T5
extruded
extruded
1.84
1.80
1.80
42
44
44.8
44.8
170
195
165
165
280
280
285
285
19
10
6
6
SBF
Hank’s
SBP00
SBP40
Mg-6Zn
WE43
AZ80
AZ80
AC
C
Pure Mg
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D
Armco Fe
Metallurgy Density Young modulus
N Yield strengthUltimate tensile
(g/cm3)
(GPa)
(MPa)
strength (MPa)
annealed
200
150
200
EP
Materials
7
7
0.16
1.35
0.07
0.09
[17]
[4,78]
Present work
Present work
ACCEPTED MANUSCRIPT
-1.40
(b)
-1.50
-1.50
-1.55
-1.60
-1.65
SBP0
SBP5
SBP10
SBP20
-1.70
-1.75
-1.55
-1.60
-1.65
-1.70
RI
PT
EOC vs. SCE / V
EOC vs. SCE / V
-1.45
(a)
-1.45
-1.75
-1.80
-1.85
-1.80
-1.90
0
5 10 15 20 25 30 35 40 45 50 55 60 65 70
0
Time / min
SBP40
NaCl
NaCl 40
DW40
5 10 15 20 25 30 35 40 45 50 55 60 65 70
SC
Time / min
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Fig. 1. Variation of OCP with the exposure time for AZ80 alloy in the different tested electrolytes.
-35
(b)
-60
-50
5 min
-40
0.0
-30
-0.5
-20
5 min
30 min
1h
2h
3h
-10
EP
-1.0
-1.5
-2.0
-2
-1
0
1
2
3
0
-25
z'' / kΩ cm2
3h
1.0
-70
5 min
phase / degree
30 min
0.5
log ( IzI /
1h
5 min
30 min
1h
2h
3h
-30
TE
D
(a)
1.5
-80
-20
-15
0.27 Hz
-10
30 min
0.72 Hz
-5
10
20
3h
0
4
1h
2h
5 min
0.01 Hz
0.01 Hz
log (f / Hz)
5
5
10
15
20
z' / kΩ
Ω
25
30
2222
0
m
c
AC
C
cm2)
2.0
35
Fig. 2. EIS spectra of AZ80 alloy in SBP0 fluid as a function of the exposure time: (a) Bode
diagrams and (b) Nyquist diagrams. Inset (b): Equivalent circuit model used for impedance data
fitting.
1
40
ACCEPTED MANUSCRIPT
-35
-50
-40
0.5
log ( IzI /
30 min
5 min
-30
0.0
-20
-0.5
5 min
30 min
1h
-1.0
-10
0
-1.5
-25
-20
-15
-10
0
1
2
3
5
4
-40
0.5
-30
0.0
-20
5 min
30 min
1h
-10
0
-1.5
-40
-30
0.163 Hz
-20
0
1
2
3
4
10
-50
5 min
-40
-30
-20
5 min
30 min
1h
-1.0
-1.5
60
70
5 min
30 min
1h
-10
0
10
-2.0
-30
cm2
-60
30 min
z'' /
0.0
-0.5
50
(g)
phase / degree
0.5
40
-70
AC
C
1.0
30
-40
EP
1.5
20
z' / kΩ
Ω
-80
5 min
1h
10
2222
(c) SBP20
2.0
0
m
c
log (f / Hz)
2.5
0.01 Hz
0
TE
D
-1
40
-10
20
-2
35
5 min
30 min
1h
10
-2.0
30
SC
cm2
-50
5 min
log ( IzI /
25
-50
z'' /
30 min
1.0
-1.0
phase / degree
-60
-0.5
20
(f)
-70
M
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cm2)
5 min
1h
1.5
15
-60
-80
2.0
10
z' / kΩ
Ω
2.5
(b) SBP10
5
2222
0
m
c
-1
log (f / Hz)
cm2)
0.01 Hz
0
20
-2
log ( IzI /
0.189 Hz
-5
10
-2.0
5 min
30 min
1h
RI
PT
1.0
-60
cm2
1.5
-70
5 min
1h
z'' /
(a) SBP5
2.0
(e)
-30
-80
phase / degree
cm2)
2.5
-20
-10
0.531 Hz
0.01 Hz
0
20
-2
-1
0
1
2
3
4
10
10
20
z' / kΩ
Ω
2
30
2222
0
m
c
log (f / Hz)
40
50
ACCEPTED MANUSCRIPT
-30
1.0
1h
-60
30 min
-50
5 min
-40
-30
0.0
-20
-0.5
5 min
30 min
1h
-1.0
-10
0
-1.5
10
-2.0
cm2
5 min
1.5
-20
z'' /
-70
phase / degree
2.0
-10
-15
0.32 Hz
-5
0
0.01 Hz
20
-1
0
1
2
3
5
4
5
10
15
SC
z' / kΩ
Ω
20
2222
0
log (f / Hz)
m
c
-2
5 min
30 min
1h
-25
RI
PT
(d) SBP40
0.5
25
30
35
TE
D
M
AN
U
Fig. 3. Impedance spectra of AZ80 alloy in SBP containing: (a, e) 5 g/L BSA, (b, f) 10 g/L BSA, (c,
g) 20 g/L BSA, and (d, h) 40 g/L BSA over exposure time of one hour: (a-d) Bode diagrams, and
(e-h) Nyquist diagrams. Inset (e): Equivalent circuit model used for impedance data fitting.
EP
log ( IzI /
(h)
-80
AC
C
cm2)
2.5
3
ACCEPTED MANUSCRIPT
-7
(c)
-5
-50
-40
0.0
-30
-0.5
-20
-1.0
-10
0
-1.5
-2.0
1
2
3
-2
-1
0
10
1
20
2
4
0.01 Hz
0
1
2
3
4
z' / kΩ
Ω
log (f / Hz)
0.031 Hz
0.031 Hz
SC
0
-3
5
6
2222
-1
-4
m
c
-2
NaCl (5 min)
NaCl (30 min)
NaCl (1 h)
NaCl40 (5 min)
NaCl40 (30 min)
NaCl40 (1 h)
RI
PT
1.0
cm2
-70
-60
0.5
log ( IzI /
-6
z'' /
(a)
1.5
-80
phase / degree
cm2)
2.0
7
8
9
-8
M
AN
U
-70
-50
5 min
-40
0.0
-30
-20
-0.5
-1.0
-1.5
-10
TE
D
5 min
30 min
1h
0
1
-2
0
2
3
4
2
0
EP
log (f / Hz)
0.01 Hz
0.01 Hz
2
4
z' / kΩ
Ω
6
2222
0
-4
m
c
-1
1h
10
20
-2
cm2
0.5
z'' /
-60
30 min
5 min
30 min
-6
phase / degree
1h
1.0
8
10
Fig. 4. Impedance spectra of AZ80 alloy in: (a, c) NaCl and NaCl40, and (b, d) DW40 over
exposure time of one hour. (a, b) Bode diagrams, and (c, d) Nyquist diagrams. Inset (c) Nyquist
diagrams in neat NaCl solution.
AC
C
cm2)
(b) DW40
log ( IzI /
(d)
-80
1.5
4
ACCEPTED MANUSCRIPT
-60
SBP0
SBP5
SBP10
SBP20
SBP40
NaCl
NaCl40
DW40
RI
PT
-50
-30
SC
-20
-10
0
10
20
30
40
z' / kΩ
Ω
2222
0
m
c
10
M
AN
U
z'' /
cm2
-40
50
60
70
AC
C
EP
TE
D
Fig. 5. Comparison of all Nyquist diagrams for the impedance spectra of AZ80 alloy in the different
tested electrolytes after 1 h immersion.
5
ACCEPTED MANUSCRIPT
(a)
(b)
80
50
30
SBP20
SBP0
SBP40
SBP5
20
DW40
10
RI
PT
cm2
60
40
RT /
SBP10
SBP0
SBP5
SBP10
SBP20
SBP40
NaCl
NaCl40
DW40
70
NaCl40
0
NaCl
0
10
20
30
40
50
60
70
Time / min
(c)
(d)
8
NaCl40
DW40
4
SBP10
SBP5
2
SBP20
SBP40
SBP0
0
0
10
20
30
40
50
70
SBP10
SBP40
SBP20
SBP5
DW40
NaCl
NaCl40
0
TE
D
Time / min
60
C2 /
6
C1 /
cm-2
10
33
30
27
24
21
18
15
12
9
6
3
0
SC
NaCl
M
AN
U
cm-2
12
10
20
SBP0
30
40
50
60
70
Time / min
AC
C
EP
Fig. 6. (a) The total film resistance (RT) in all tested electrolytes as a function of immersion time,
(b) Dependence of RT value on BSA concentration in SBP fluid after 1 h immersion, (c) and (d)
Variation with time of C1 and C2 for the inner and outer layers of the surface film on AZ80 alloy.
Fig. 7. Schematic illustrations for BSA adsorption mechanism on AZ80 alloy surface at different
BSA doses in SBP fluid.
6
ACCEPTED MANUSCRIPT
(a) SBP0
Element
CK
OK
Na K
Mg K
Al K
PK
Cl K
Ca K
Element
CK
NK
OK
Na K
Mg K
Al K
PK
Cl K
Ca K
SC
AC
C
EP
TE
D
M
AN
U
(c) SBP10
(d) SBP20
(e) SBP40
Wt %
4.39
1.53
28.32
1.05
52.87
7.74
0.72
2.71
0.67
RI
PT
(b) SBP5
Wt %
3.89
13.10
1.12
73.61
6.83
0.47
0.52
0.46
Element
CK
NK
OK
NaK
Mg K
Al K
PK
Cl K
Ca K
Wt %
5.72
2.78
1.29
85.56
3.89
0.32
0.12
0.32
Element
CK
NK
OK
NaK
Mg K
Al K
PK
Cl K
Ca K
Wt %
7.62
5.84
1.02
79.77
4.68
0.33
0.41
0.33
Element
CK
NK
OK
NaK
Mg K
Al K
PK
Cl K
Ca K
Wt %
9.65
0.85
20.11
0.92
59.68
7.08
0.52
0.65
0.54
Fig. 8. SEM images and EDX spectra of the corrosion layers formed on AZ80 in the different
electrolyte after 1 h immersion at 37 oC: (a) SBP0, (b) SBP5, (c) SBP10, (d) SBP20, and (e)
7
SBP40.
ACCEPTED MANUSCRIPT
0
(a)
-3
-4
3
1
-5
(1) SBP0
(2) SBP5
(3) SBP10
(4) SBP20
(5) SBP40
-6
-7
25
-8
4
-9
-1.7
-1.6
-1.5
-1.4
-1.3
-1.2
-1.1
(b)
-2
-3
-4
-5
-6
-7
8 7
-8
-9
-1.7
-1.0
E (vs. SCE) / V
SC
-2
log (i / A cm -2 )
-1
-1.6
-1.5
-1.4
(6) NaCl
(7) NaCl40
(8) DW40
6
-1.3
-1.2
-1.1
-1.0
M
AN
U
E (vs. SCE) / V
EP
TE
D
Fig. 9. Potentiodynamic polarization curves of AZ80 alloy after 1 h immersion at 37 oC under
aerated conditions: (a) in SBP without and with different BSA concentrations, and (b) in neat 0.7%
NaCl solution, and in 0.7% NaCl solution or DW each containing 40 g/L BSA.
AC
C
log (i / A cm -2)
-1
RI
PT
0
Fig. 10. The inhibition zones of S. aureus after 24 h incubation with different extracts from AZ80
alloy immersed for 24 h at 37 oC in: (A) SBP0, (B) SBP10, and (C) SBP40 electrolytes.
8
ACCEPTED MANUSCRIPT
Highlights
Electrochemical techniques show that AZ80 degradation in SBP depends on BSA amount.
•
10_20 g/L BSA can inhibit the alloy degradation via an integrated adsorbed layer.
•
At doses < 10 g/L BSA accelerates the alloy dissolution via metal chelating effect.
•
At doses > 20 g/L BSA stimulates the alloy degradation due to protein aggregation.
•
SEM, EDX spectra and alloy antibacterial efficiency all support the obtained data.
AC
C
EP
TE
D
M
AN
U
SC
RI
PT
•
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