Article
Comparative in vitro study and
biomechanical testing of two different
magnesium alloys
Andreas Weizbauer1,2, Christian Modrejewski1, Sabine Behrens3,
Helmut Klein4, Patrick Helmecke5, Jan-Marten Seitz3, Henning Windhagen1,
Kai Mo¨hwald3, Janin Reifenrath6 and Hazibullah Waizy1
Abstract
In this in vitro study, magnesium plates of ZEK100 and MgCa0.8 alloy similar to common titanium alloy osteosynthesis
plates were investigated as degradable biomedical materials with a focus on primary stability. Immersion tests were
performed in Hank’s Balanced Salt Solution at 37C. The bending strength of the samples was determined using the four-
point bending test according to ISO 9585:1990. The initial strength of the noncorroded ZEK100 plate was 11% greater
than that of the MgCa0.8 plate; both were approximately 65% weaker than a titanium plate. The bending strength was
determined after 48 and 96 h of immersion in Hank’s Balanced Salt Solution; both magnesium alloys decreased by
approximately 7% after immersion for 96 h. The degradation rate and the Mg2þ release of ZEK100 were lower than
those of MgCa0.8. Strong pitting and filiform corrosion were observed in the MgCa0.8 samples after 96 h of immersion.
The surface of the ZEK100 plates exhibited only small areas of filiform corrosion. The results of this in vitro study
indicate that the ZEK100 alloy may be more suitable for biomedical applications.
Keywords
In vitro, corrosion, magnesium, orthopedic, implants
Introduction
Magnesium and its alloys were rediscovered at the end
of the 20th century as promising degradable biomater-
ials for orthopedic applications.1–3 Researchers are
attracted by the prospect of avoiding a second surgery
for implant removal, which is necessary with titanium
implants.4 In addition, magnesium alloys provide desir-
able mechanical properties, close to those of natural
bone, which prevent the ‘‘stress shielding’’ phenomena.5
Magnesium reportedly stimulates new bone formation
(osteoinductive) and also supports the adherence of
bone tissue to its surface (osteoconductive).5–8
In human body fluids and blood plasma with high
Cl– concentrations, the degradation of magnesium
alloys is accompanied by a loss of mechanical stability.
The resulting surface layer consists mainly of the cor-
rosion product magnesium hydroxide (Mg(OH)2),
which is locally dissolved into soluble MgCl2. In add-
ition, calcium phosphate salts9 and MgCO3
10 precipi-
tate on the surface of magnesium alloys resulting in a
corrosive protective effect. During the corrosion pro-
cess, Mg2þ ions are released from the metal.
Magnesium is a natural component of the human
body and it is mainly stored in bone (60–65%).11
Furthermore, it is the second most abundant cation in
cells and an important cofactor for many enzymes that
are involved in metabolic pathways and DNA repair
processes.11,12 The concentration of Mg2þ in the
blood serum of human adults comprises 0.7–1.1mol/l;
Journal of Biomaterials Applications
2014, Vol. 28(8) 1264–1273
! The Author(s) 2013
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DOI: 10.1177/0885328213506758
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1Department of Orthopedic Surgery, Hannover Medical School,
Hannover, Germany
2CrossBIT, Center for Biocompatibility and Implant-Immunology,
Department of Orthopedic Surgery, Hannover Medical School,
Hannover, Germany
3Institute of Materials Science, Leibniz Universita¨t Hannover, Garbsen,
Germany
4Department of Crystallography, University of Go¨ttingen, Go¨ttingen,
Germany
5Institute of Production Engineering and Machine Tools (IFW), Leibniz
Universita¨t Hannover, Garbsen, Germany
6Small Animal Clinic, University of Veterinary Medicine Hannover,
Hannover, Germany
Corresponding author:
Andreas Weizbauer, Department of Orthopedic Surgery, Hannover
Medical School, Anna-von-Borries-Str. 1-7, Hannover 30625, Germany.
Email: weizbauer.andreas@mh-hannover.de
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serum Mg2þ levels exceeding this amount are excreted
by the kidneys.11
In this in vitro study, two different magnesium alloys
were investigated: one containing rare earth elements
(REEs) (ZEK100) and MgCa0.8. Both alloys were
chosen because of their promising biomechanical prop-
erties, and both had been evaluated in vivo in previous
studies.13–21 ZEK100 pins have shown promising initial
mechanical stability.15,20 Lensing et al. reported good
biocompatibility and an osteoconductive effect of
ZEK100 in the middle ear of rabbit.17 In contrast,
some studies have shown that the biocompatibility of
the ZEK100 alloy appears to be questionable.16,22
Pure magnesium possesses insufficient mechanical
properties.23 The alloying of magnesium with a large
variety of elements was introduced to solve this issue.
These alloying elements influence the corrosion film
composition and therefore the corrosion properties.24
Magnesium alloyed with passivating elements, such as
the REE (yttrium and neodymium), exhibits enhanced
corrosion resistance.18,24 Magnesium alloys containing
REE were first investigated at the beginning of the
1980 s.25 The addition of REE contributes to solid solu-
tion strengthening and grain boundary strengthening,
and therefore enhances the ductility of the alloy.1,25
Several studies have investigated the properties and
corrosion behavior of binary magnesium–calcium
alloys.13,14,18,19,26–30 The addition of calcium provides
solid solution strengthening and contributes to grain
refinement owing to grain boundary strengthening.1
It was observed that the addition of calcium above
the solubility limit of approximately 1.34wt% Ca
enhanced the corrosion rate.18,19,26 The addition of
less than 1wt% Ca demonstrated a low corrosion
rate with an appropriate elastic modulus.18
Previous studies have demonstrated good biocom-
patibility of MgCa0.8 in vivo.13,28 The production of
high amounts of gas in a short time due to high corro-
sion rates is undesirable for clinical applications. The
accumulation of gas associated with MgCa0.8 implants
was mild to moderate in vivo, with no harmful effects
toward the animal.14
In this in vitro study, we investigated the corrosion
behavior and mechanical properties of two different
alloys (MgCa0.8 and ZEK100) in Hank’s Balanced
Salt Solution (HBSS). The results were compared
with a titanium alloy plate (Ti6Al4V) of similar
sample size and shape. This study is focused on primary
stability and the effects of degradation during the first
days of immersion. In addition, Mg2þ release from the
alloy was determined, and the corrosion rate was eval-
uated according to the mass loss method. Alloy sur-
faces were investigated using scanning electron
microscopy (SEM), and high-energy synchrotron radi-
ation was applied for microstructural characterization.
Methods
Material and preparation
Two different magnesium alloys were examined in this
study: ZEK100 (1wt% zinc, 0.1wt% zirconium, and
0.1wt% rare earth metals) and MgCa0.8 (0.8wt% cal-
cium). The plates were 50mm long, 8mm wide, and
2mm thick. They were manufactured by face slab
milling. Five screw holes were drilled and counterbored
in each plate. The resulting screw holes had an inner
diameter of 3mm on the bone-facing side and a
counterbore diameter of 5mm (Figure 1). The samples
had a surface of 898mm2. Two titanium alloy plates
with the same dimensions were manufactured from a
Ti6Al4V rod (Goodfellow GmbH, Bad Nauheim,
Germany).
The surfaces of the magnesium samples were pickled
prior to immersion testing with a solution of 5ml 65%
nitric acid (Merck, Darmstadt, Germany), 20ml
85% glycerol (Merck, Darmstadt, Germany), and
5ml 100% acetic acid (Merck, Darmstadt, Germany)
for 30 s. Next, the samples were ultrasonically cleaned
for 2min with isopropanol and then dried in a vacuum
drying oven (Heraeus Vacuum oven, Thermo Scientific,
Bonn, Germany) at 50C for approximately 20min.
Immersion testing
The static immersion test was performed in HBSS
(Biochrom AG, Berlin, Germany) to simulate the
normal ion concentration under physiological condi-
tions at 37.5C. The corrosion media contained
8.00 g/l NaCl, 0.4 g/l KCl, 0.04 g/l Na2HPO4, 0.06 g/l
KH2PO4, 0.20 g/l MgSO47H2O, 0.14 g/l CaCl2,
1.00 g/l glucose, and 0.35 g/l NaHCO3. The plates
were placed into 500ml HBSS by a nonresorbable
Ti6Al4V
MgCa0
.8
ZEK10
0
Figure 1. Picture of the plates (length: 50mm, width: 8mm,
thickness: 2mm) used in this in vitro study.
Weizbauer et al. 1265
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wire (EthiconTMProleneTM, 4-0, Johnson & Johnson
GmbH, Neuss, Germany). The pH of the solution
was monitored during the immersion test using a pH
meter (Schott Instruments Lab850, SI Analytics
GmbH, Mainz, Germany). After immersion, the sam-
ples were dried and the corrosion layer was removed
with chromic acid. Two samples were used per interval
of each magnesium alloy.The degradation rate was
evaluated according to the weight loss method31
W ¼Wbefore Wafter
At
where W is the weight loss, Wbefore (mg) is the weight
of each sample before corrosion, and Wafter (mg) is the
weight of each sample after corrosion. A is defined as
the sample’s surface area (cm2) and t is the immersion
time (d). This weight loss rate was converted into an
average corrosion rate (mm/y) using31,32
PW ¼ 2:1W
Detection of Mg2þ release through inductively
coupled plasma optical emission spectroscopy
(ICP-OES)
The ICP-OES (Spectro Ciros Vision EOP, Spectro
Analytical Instruments, Kleve, Germany) was applied
to quantitatively determine the release of magnesium
ions from the plates during the corrosion process.
Two plates of each alloy were immersed in 100ml
HBSS at 37C for 158 h. At fixed time intervals (the
first after 45min), 1ml was taken for analysis and
replaced with the same amount of fresh corrosion
media. The pH of every analysis sample was monitored.
Biomechanical testing
To specify the mechanical properties of the specimens,
they were tested according to ISO 9585:1990. The four-
point bending test setup (Figure 2) was designed for
tests with a uniaxial material testing machine (Mini
Bionix 858, MTS Systems in Minneapolis, MN,
USA). The test system consisted of four rollers with a
diameter of 8mm. The distance between the lower and
outer rollers was 32mm, and the distance between the
upper and inner rollers was 10mm. The plates were
positioned on the lower rollers according to the proto-
col of ISO 9585:1990. The distance between the outer
and inner rollers was 11mm, including a screw hole of
the demonstrators. The central screw hole and another
screw hole were positioned between the inner rollers.
The test rig was produced by the Research Workshop
of the Medical School, Hannover, Germany.
Microstructural characterization
After the corrosion products were removed from the
surface with chromic acid, the magnesium plates were
analyzed with a Zeiss EVO 60 VP SEM.
Diffraction experiments were performed using high-
energy synchrotron radiation (E100 keV,
¼ 0.12561 A˚). A Perkin–Elmer area detector
(400mm 400mm, 200 mm pixel size) was used with a
sample–detector distance of 1430mm.
Results
Immersion test and corrosion properties
The pH of HBSS was measured during the immersion
of the plates (Figure 3). The pH of the HBSS increased
rapidly during the first hours of exposure. After 12 h of
Figure 2. Picture of the four-point bending setup. The samples were tested according to ISO 9585:1990.
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immersion, the alkalization of the corrosion medium
progressed at a slower pace. However, the immersion
of MgCa0.8 plates resulted in a higher HBSS pH than
the immersion of ZEK100 plates. After 96 h of immer-
sion, the pH of HBSS rose to 9.9 for MgCa0.8 plates
compared with 8.8 for ZEK100 plates.
The corrosion rate was determined after 48 and 96 h
using the weight loss method (Table 1). ZEK100
demonstrated a five-fold lower corrosion rate than
MgCa0.8.
Release of Mg2þ from the magnesium plates
The Mg2þ concentration dissolved into the HBSS from
the magnesium samples was determined using ICP-OES
(Figure 4(a)). The MgCa0.8 plates exhibited rapid
Mg2þ release during the first 20 h of immersion. The
ZEK100 plates exhibited a more consistent increase of
Mg2þ concentration throughout the exposure time. The
rate of Mg2þ release from the MgCa0.8 plates during
the first 24 h of immersion was 1.082mg/h (1.2
10–3mg/hmm2), nearly eight-fold higher than the rate
of ion release from the ZEK100 plates: 0.143mg/h
(1.592 10–4mg/hmm2) (Figure 4(b)). Moreover, the
pH of HBSS increased more rapidly as a result of the
immersion of the MgCa0.8 samples than the ZEK100
samples. After 20 h, the MgCa0.8 corrosion media
reached a stable pH of approximately 10 (Figure 4(c)).
The effect of corrosion on bending strength
Figure 5 shows the bending strengths of the magnesium
and titanium plates. Uncorroded samples of ZEK100
(0.306 0.003Nm) were stronger than MgCa0.8
(0.272 0.00005Nm). In contrast, the bending strength
of titanium plates with the same dimensions was higher
(0.803 0.001N m).
The strength of both magnesium alloys decreased
with immersion time. The bending strength after 96 h
of immersion was approximately 7% lower than the
initial bending strength (0 h immersion).
Microstructural characterization and
phase identification
Figure 6 shows the surface of the magnesium plates
after removal of the surface corrosion products that
resulted from 96 h of immersion. Filiform and pitting
corrosion were observed on the surface of the MgCa0.8
alloy. The surface of the ZEK100 plates showed only
small areas of filiform corrosion (Figure 6(a) and (b)).
The diffraction diagram and the diffraction rings of
both alloys are displayed in Figure 7. The diffraction
rings of both alloys demonstrated an inhomogeneous
intensity distribution (Figure 7(b) and (e)). The diffrac-
tion diagram of the MgCa0.8 alloy is shown in Figure
7(c) and (d). In addition to the measured intensities
10.50
10.00
9.50
9.00
8.50
8.00
7.50
0 20 40 60
Immersion time (h)
pH
 v
al
ue
80 100
MgCa0.8
ZEK100
Figure 3. pH value versus immersion time of MgCa0.8 and ZEK100 plates in 500ml Hank’s solution. The dots represent mean values
and standard deviation. Sample size for each alloy: n¼ 4 for 0–48 h, n¼ 2 each for 72 and 96 h.
Table 1. Corrosion rates of the magnesium alloy plates
according to the mass loss method.31 Sample size: Two of each
alloy per interval.
Mean corrosion rate (mm/y)
48 h immersion 96 h immersion
MgCa0.8 3.62 2.89
ZEK100 0.70 0.63
Weizbauer et al. 1267
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(black line), simulated diffraction diagrams are also
provided for Mg (solid solution, red line) and Mg2Ca
(intermetallic compound, blue line). Both simulated
phases fit the position of the measured reflection lines
quite well, but not the intensity. The measured intensi-
ties of the ZEK100 samples also differed strongly
because of the material’s texture. No intermetallic
phases were detected.
60(a) (b)
(c)
50
40
30
20
10
0
0 10 20 30 40 50 60 70 80 90 100 110 120 130 140 150
MgCa0.8
ZEK100
MgCa0.8
ZEK100
Linear regression line MgCa.8
y =0.143x + 15.029
y =1.082x + 10.846
Linear regression line ZEK100
MgCa0.8
ZEK100
0 10 20 30 40 50 60
0 4 8 12 16 20 24
70 80
Immersion time (h)
Immersion time (h) Immersion time (h)
90 100110120130140150
45
40
35
30
25
20
15
5
10
0
10.5
pH
 v
al
ue
M
g2
+  
-c
on
ce
nt
ra
tio
n 
(m
g/
l)
M
g2
+  
-c
on
ce
nt
ra
tio
n 
(m
g/
l)
10
9.5
9
8.5
8
7.5
Figure 4. (a) Mg2þ release from the ZEK100 and MgCa0.8 plates as a function of time. (b) Average Mg2þ release during the first 24 h.
(c) pH of the analyzed corrosion media sample. The dots represent mean values and standard deviation. Two samples were tested per
alloy.
0.805 titanium plates without immersion
MgCa0.8
ZEK100
0.305
0.295
0.275
0.285
Be
nd
in
g 
st
re
ng
th
 (N
m)
0.265
0.255
0.245
0 48
Immersion time (h)
96
Figure 5. Bending strength versus immersion time. The bending strength of both alloys decreased with immersion time. The dots
represent mean values and standard deviation. Two samples were tested per alloy and immersion time.
1268 Journal of Biomaterials Applications 28(8)
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0 200 400 600 800 1000 1200 1400 1600 1800 2000
0 200 400 600 800 1000
Columns
Columns
1200 1400 1600 1800 2000
0 200 400 600 800 1000 1200 1400 1600 1800 2000
0
0
200
200
400
400
600
600
800
800
1000
1000
1200
1200
1400
1400
1600
1600
1800
1800
2000
2000
0
200
400
600
800
1000
1200
1400
1600
1800
2000
0
200
400
600
800
1000
1200
1400
1600
1800
2000
0
200
400
600
800
1000
R
ow
s
R
ow
s
1200
1400
1600
Mg
Measurement
Mg
Measurement
Mg
2
Ca
1800
2000
1 2 3 4 5 6 7 8 9 10
1 2 3 4 5
2 Theta/°
2 Theta/°3
0
4
2 Theta/°
In
te
ns
ity
In
te
ns
ity
In
te
ns
ity
6 7 8 9 10
(a) (b)
(d)(c) (d)
(d)
Figure 7. Diffraction diagram and diffraction rings of ZEK100 and Mg2Ca0.8 alloy samples. (a) and (b): ZEK100 and (c) to (e):
Mg2Ca0.8.
Figure 6. SEM pictures of the plates after 96 h immersion and removal of the corrosion layer. (a) and (b): ZEK100 and (c) and (d):
MgCa0.8.
Weizbauer et al. 1269
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Discussion
Fracture healing is dependent on the fracture gap and
achieved stabilization. Plates are a common orthopedic
implant and are usually attached to the bone with
screws.33 The fixation of a plate should supply sufficient
stability to enhance fracture healing. High primary sta-
bility is important and should not decrease greatly
during the first days after the placement of the plate.
Therefore, in this in vitro study we focused on the ini-
tial strength of the magnesium implants versus titanium
implants and investigated the implant stability specific-
ally during the first days of immersion. The currently
used plates are made of stainless steel or titanium
alloys5 and are available in a wide range of sizes.
In this study, five-hole plates from two different mag-
nesium alloys were examined after immersion in HBSS
and compared with titanium plates.
The biomechanical results of the magnesium-based
plates were compared with those of a plate made from
the titanium alloy Ti6Al4V, which is commonly used in
orthopedic trauma surgery. Titanium alloys exhibit
minor biocorrosion with the release of metallic ions,
which normally do not compromise plate stability.34
The corrosion products, released metallic ions, have
been associated with proinflammatory, cytokine-
induced, enhanced osteoclastic activity, which in turn
causes bone resorption.34
We assume that the loss of stability of titanium
alloys is negligible within the short period of time
(96 h) observed in our study. Therefore, the biomech-
anical examinations of the titanium plates were per-
formed to their initial conditions without immersion
tests.
The bending strength of the titanium plates is com-
parable to that of a 2mm mini-fragment plate, which
was also tested according to ISO 9585:1990.35 In this
study a bending strength of 0.53Nm was determined.
The biomechanical parameter of the tested 2mm mini-
fragment plate is lower than the bending strength of the
titanium plates evaluated in our study (0.803Nm). The
reason for this slight deviation might be explained by
the difference in plate design.
The initial strengths of both magnesium alloys were
approximately 65% lower than that of the titanium
alloy. Of the two magnesium alloys, the ZEK100
plates exhibited greater initial stability. Therefore,
ZEK100 plates may be considered more suitable for
osteosynthesis applications than MgCa0.8 plates. The
degradation of magnesium alloys is known to be high in
the first hours of exposure to the corrosion media. This
is mainly due to the fact that the corrosion protective
layer needs some time for formation. Therefore, we
particularly investigated the first hours of corrosion
and its effect on the stability of the magnesium plates.
The study of Waizy et al. showed that between week 2
and week 6 of immersion of ZEK100 in HBSS no con-
siderable change in bending strength was reported.36
Our study showed that in the first 96 h of degradation
the bending strength of both magnesium alloys
decreased about 7%. ZEK100 still have a higher stabil-
ity and is therefore based on the biomechanical results
of the preferred magnesium alloy. Many studies have
shown that the degradation of magnesium alloys in
in vitro experiments is higher than in comparable in vivo
studies.37–39 A decreased loss of stability in vivo com-
pared to the results in this study is assumed.
Alloy plates are used in clinical practice as internal
splints to hold fractured ends of bone together; add-
itional fracture stabilization with lag screws is
common. According to the biomechanical results, the
preferred use of these magnesium-based plates is the
fixation of non-weight-bearing bones (e.g. fracture of
the metacarpal bones). Sufficient stability of degradable
plates in vivo must be ensured during the first 6–8
weeks after implantation.40
The magnesium alloy ZEK100 contains REE and
zirconium in low concentrations. The alloying elements
are used to improve the mechanical and corrosion
properties of the biomaterial.41 These alloying elements
are known to have adverse effects and potential long-
term risks,42–45 although some studies have reported
good biocompatibility.46,47 Allergic hypersensitivity,
reduction of hemoglobin oxygen affinity, and fatty
liver are some potential adverse effects of these elements
at toxic doses.43,45 Chelated REE are easily excreted in
the urine, and mortality studies have revealed that REE
are not highly toxic.43 However, bone is one of the
main target organs of REE because REE are of a
radius similar to Ca2þ ions, which may lead to REE
accumulation in the environment directly surrounding
the implant.43 Because implants made of degradable
magnesium alloys are not removed after implantation,
in vivo biocompatibility studies should focus on the
long-term effects of these implants. In an animal
study with 12 months of observation time, a good bio-
logical response to a REE-containing alloy
(MgYREZr, similar to WE43) and its degradation
products was observed.48
The degradation of magnesium and its alloys in bio-
fluids is described by the following reaction
Anodic reaction: Mg ! Mg2þþ 2 e-
Cathodic reaction: 2H2Oþ 2 e– ! H2"þ 2OH–
Mg2þ þ 2OH !Mg OHð Þ2 sð Þ
Degradation is accompanied by an alkalization of the
corrosion media due to the production of hydroxide
ions (OH–). These experiments were performed without
CO2 gas application. HCO3
–/CO2 is the main buffering
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system in blood and in HBSS. The high proportion of
hydroxide ions supports the formation of magnesium
hydroxide, which in turn acts as a protective layer
against corrosion. The greater tendency toward film
formation is reportedly associated with local
pH> 10.5 and is believed to have a corrosion-inhibiting
effect.49 Magnesium hydroxide is disrupted by chloride
ions with the release of OH–
Mg OHð Þ2þ2Cl !MgCl2 þ 2OH
In response to the immersion of both magnesium alloys
(ZEK100 and MgCa0.8), the pH of the HBSS increased
sharply during the first hours of exposure. An acceler-
ated pH increase during the first hours of immersion
was reported by several in vitro studies.50–53 The decel-
erated increase of pH after 6 h may be attributed to the
formation of magnesium hydroxide and other phos-
phate- and calcium-containing salts, which precipitated
on the surface of the magnesium samples. Interestingly,
the pH of the HBSS increased more with the immersion
of MgCa0.8 plates than ZEK100 plates. This finding
indicated the more progressed and faster corrosion of
the MgCa0.8 plates. These observations were sup-
ported by the determined corrosion rates for both mag-
nesium alloys according to the mass loss method (Table
1). The corrosion rates were approximately five-fold
greater for the MgCa0.8 plates than for the ZEK100
plates. The corrosion rate was slower after 96 h of
immersion than after 48 h, which can be explained by
the protective effect of the formed surface films (result-
ing from the precipitation of calcium phosphate salts
and magnesium hydroxide). Song et al. reported the
formation of a magnesium hydroxide film on the sur-
face of an AZ31 within 2 h after immersion.23 Similar to
our results, Zhu et al. reported a reduction of the cor-
rosion rate with progressing immersion time for a AZ31
alloy.54 A study by Walker et al. reported lower corro-
sion rates for MgCa0.8 than were observed in the pre-
sent study (1.937mm/y in MEMp, 1.291mm/y in
MEM, and 0.795mm/y in EBSS).38
Several studies have shown that the corrosion rates
in vivo differ from those determined in vitro.39,55 Many
different parameters influence the in vitro corrosion
rate (e.g. choice of corrosion media, temperature, CO2
gas application, etc.).3,56,57 Therefore, the alteration of
parameters makes it difficult to compare different
in vitro tests.
The release of Mg2þ from the metal was measured
using ICP-OES analysis (Figure 4). Only a few studies
have analyzed the release of Mg2þ from magnesium
alloys.51,52 These results confirm the observations of
the immersion test and the determination of the corro-
sion rate. The release of Mg2þ decelerates with immer-
sion time as a consequence of protective surface layer
formation. This analysis is influenced by the formation
and degradation of Mg(OH)2. Some of the dissolved
Mg2þ ions from the metal form insoluble magnesium
hydroxide Mg(OH)2, and therefore are excluded from
the analysis, although as mentioned previously,
Mg(OH)2 can be transformed into soluble MgCl2 by
Cl– ions and is therefore again accessible for ICP-
OES analysis. However, the speed of this accumulation
and degradation process is unknown. Notably, the high
rate of MgCa0.8 corrosion appears to result in a faster
loss of biomechanical stability.
The microstructural characterization of the
MgCa0.8 plates revealed the presence of the intermetal-
lic Mg2Ca phase, which is consistent with other
studies.26,58 Intermetallic phases form during the solidi-
fication of the cast alloys and mainly distribute at the
grain boundaries.59 Most of the intermetallic second
phases reportedly possess a corrosion-accelerating
effect.60 This reduction of corrosion resistance is
mainly attributed to galvanic corrosion between the
anodic a-Mg and the intermetallic phases, which act
as a cathode.60 Kirkland et al. report that this phase
is very reactive, with faster dissolution rates than Mg.26
The Mg2Ca phase is described as a more efficient anode
than a-Mg.26
Analysis of the ZEK100 plate revealed no interme-
tallic phases. Here, a combined Rietveld texture ana-
lysis would have given very precisely the amounts of the
present phases and the texture (to calculate the aniso-
tropic behavior of the material). Small amounts of
other phases might also be found using this method,
because of the improved alignment of the measured
intensities. MgZn and Mg2Zn are the most prominent
intermetallic phases that would be expected according
to an analysis of a similar magnesium alloy.51,61
Conclusion
Magnesium plates made of different magnesium alloys
were investigated after in vitro corrosion in HBSS.
Biomechanical tests demonstrated a continuous loss
of mechanical strength by osteosynthesis plates made
of magnesium alloys. Although the initial bending
strength of the uncorroded ZEK100 plate was greater
than that of the MgCa0.8 plate, both were approxi-
mately 65% weaker than a titanium plate. The degrad-
ation rate and the Mg2þ release of ZEK100 were lower
than those of MgCa0.8.
Acknowledgments
The authors gratefully acknowledge the financial support pro-
vided by the German Research Society (DFG) within the col-
laborative research project (SFB 599). We thank Markus
Badenhop for excellent technical support and Christopher
Mu¨ller for the design of Figure 1.
Weizbauer et al. 1271
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