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128 Surface Modification of Magnesium and its Alloys for Biomedical Applications
vivo milieu (Levesque, Hermawan, Dube, & Mantovani, 2008; Xin et al., 2011)
(see Tables 4.2 and 4.3). These mixtures do not contain all the degradation products
of ma gne sium or one of its alloys. During the production of such mixes, the
hydrogen evaporates and particles are separated out in the centrifuging or precipitation process (Gu et al., 2009; Witte, Hort, et al., 2008). A further problem (as
described by Lorenz et al., 2009) associated with in vitro testing is mismatch be-
tween the small quantities of the medium and the surface of the test material.
This im balance may lead investigators to think corrosion is far greater than it really
is. Purnama, Hermawan, Couet, and Mantovani (2010) recommend that biocompatibility testing for degradable materials should not be derived directly from tests for
nonbiodegradable materials.
Magnesium corrosion involves the formation of a degradation layer on the s urface, which exhibits a corrosion-inhibiting effect (Shaw, 2003). However, this layer
is stable only in an alkaline environment, so that corrosion takes place a good deal
more rapidly at a stable pH value (Witte et al., 2005). If it is presumed that, in the
human body, diverse buffer systems keep the pH level constant, then corrosion
would have to proceed faster in vivo than in vitro. However, as early as 1910,
Lespinase assumed that the corrosion rate in vitro does not correspond to that in in
vivo experiments (Lespinase, 1910). Witte et al. (2006) showed in their study that
the degradation behaviour of AZ91 and LAE442 differs between in vitro and in
vivo tests. Corrosion in vivo was considerably slower than that in vitro, with in
Table 4.2 Ion concentrations in five common solutions
0.9% NaCl PBS Hanks’ DMEM c-SBF
þ
(mmol/L) 153 157 142 127.3 142
Na
þ
K
(mmol/L) e 4.1 5.9 5.3 5.0
2þ
Ca
(mmol/L) ee1.3 1.8 2.5
2þ
Mg
(mmol/L) ee0.8 0.8 1.5
HCO
(mmol/L) ee4.2 44.1 4.2
3
(mmol/L) 153 140 145 90.8 147
CI
2
HPO
(mmol/L) e 11.5 0.8 0.9 1
4
2
SO
(mmol/L) ee0.8 0.8 0.5
4
Tris (g/L) eeee6.069
Protein (g/L) eeeee
Amino acids (g/L) eee1.6 e
Glucose (mmol/L) ee1 4.5 e
Hepes (g/L) eee5.96 e
Xin, Hu, and Chu (2011).
Bioabsorbable behaviour of magnesium alloys e an in vivo approach 129
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Table 4.3 Constituents and concentrations of buffering agents
in several test solutions
Concentration
(mmol/L) HCO
Plasma 27 0.5 16e18 ee 43.5e45.5
0.9% NaCl ee eee 0
PBS e 11.5 eee 11.5
Hanks’ 4.2 0.8 eee 5
DMEM 44.1 0.9 ee25 70
c-SBF 4.2 0.5 ee40 44.7
Xin et al. (2011).
L
3
HPO
2L
HPr Hepes TriseHCl Total
4
vivo experiments revealing that the corrosion rate of AZ91 and LAE442 exhibited a
mutually opposing trend in favour of the latter (Witte et al., 2006). A number of
studies indicate that changes over time in volume, form, and density are of importance in this regard (Jo et al., 2011; Wang et al., 2011). Difficulties with comparing
in vivo and in vitro experiments are compounded by pitting corrosion (Mueller, De
Mele, Nascimento, & Zeddies, 2009). Moreover, it was demonstrated that proteins in
particular are instrumental in slowingtherateofcorrosion(Kirkland et al., 2010;
Yamamoto & Hiromoto, 2009). These findings show that it is not yet possible to
broadly transfer in vitro findings to in vivo tests. Several authors posit that new
testing systems that more closely imitate the in vivo conditions would be the most
suitable (Gu et al., 2010; Levesque et al., 2008; Yamamoto & Hiromoto, 2009).
These findings illustrate the problem of establishing such a system that reflects the
biological, chemical and physical attributes of the human body.
4.3.1 Ex vivo test on bovine udder via microdialysis
Microdialysis is a well-established technique for analysing metabolic processes in
extracellular space. The most important aspects of microdialysis catheters to consider
are components such as length, pore size, and membrane material (Horal, Ungerstedt,
Persson, Westgren, & Marcus, 1995); the driving force involved is diffusion. Ultrafil-
tration and osmosis, which may occur under physiological conditions, have an adverse
effect on recovery (Kehr, 1993). To keep ultrafiltration and osmosis to a minimal level,
the membrane pores need to be as small as possible, with flow rate and catheter length
also reduced to a minimum. The concentration gradient can be influenced by varying
flow rate (Benveniste & Huttemeier, 1990). One criticism of this method is that, in
most cases in which it is applied in vivo, metabolic exchange between the extracellular
space and the perfusion medium does not fully occur, preventing equilibrium from being achieved (Pasnik, Moll, Cywinska-Bernas, Sysa, & Zeman, 2007). The efficiency
130 Surface Modification of Magnesium and its Alloys for Biomedical Applications
Tyrode solution
(reservoir)
Peristaltic
pump
Water bath
Fraction collector
Figure 4.2 Schematic diagram of the isolated perfused bovine udder (Kietzmann et al., 1993).
of microdialysis is described in terms of relative and absolute recovery, with recovery
determined by means of calibration (Jacobson, Sandberg, & Hamberger, 1985; Kehr,
1993; Lonnroth, Jansson, & Smith, 1987). Use of relative-recovery rates obtained by
in vitro microdialysis for assessing extracellular concentration in vivo frequently results in underestimation of the actual concentration, because the extent of in vivo relative recoveries tends to be far lower than those in vitro (Grubb, Chadburn, & Boucher,
2002; Kovar, Nolting, & Derendorf, 1997).
The bovine udder model is an example of an ex vivo model. Both efferent and
afferent vessels serving this organ had to be cannulated as soon as the animal was
euthanised, and supplied with oxygenated Tyrode’s solution. Controlled perfusion
of the organ allows the tissue to remain vital for up to eighth (Kietzmann, L€oscher,
Arens, Maass, & Lubach, 1993) (see Figure 4.2). Microanalysis, enabled by the im-
plantation of magne sium material in the bovine udder, makes it possible to analyse
metabolic processes in close proximity to the materialetissue interface. This method
also allows the concentration of the individual components to be determined (de
Lange, de Boer, & Breimer, 2000). In their study, Schumacher et al. (2011) demon-
strated that pure magnesium exhibits excellent biocompatibility and does not lead to
an increase in proinflammatory cells. These findings are in line with those of the histological examinations.
4.4 In vivo biodegradation of magnesium alloys
4.4.1 Biodegradation e general considerations
To understand biodegradation of magnesium and magnesium alloys, it is vital to
appreciate the biological importance of magnesium in humans. Magnesium is one
Bioabsorbable behaviour of magnesium alloys e an in vivo approach 131
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of the essential elements and is involved in numerous metabolic enzyme processes
(Kretz & Sch€affer, 2008; M€uller, Westenh€ofer, Bosy-Westphal, Loser, & Selberg,
2007; Saris, Mervaala, Kappanan, Khawaja, & Lewenstam, 2000). Approximately
half of all the magnesium in the body is concentrated in bone (Topf & Murray,
2003; Wacker, 1980), with the remainder stored predominantly in the muscles and or-
gans, mainly the liver. Blood plasma accounts for only around one percent of the total
(Niestroj, 2000; Vormann, 2003; Wacker, 1980). Magnesium occurs in the body in
three different forms: ionised (60%), protein-bound (30%), and bound to serum anions
(10%) (Topf & Murray, 2003). Uptake of magnesium in the gastrointestinal tract and
its elimination via the kidney plays a crucial part in the regulation of magnesium levels,
with the gall bladder and sweat having a more marginal role (Beyenbach, 1990;
Niestroj, 2000; Wacker, 1980) (see Figure 4.3)
Hypermagnesaemia and hypomagnesaemia are generally rare and lead to
impaired muscle excitability (Iannello & Belfio re , 2001). A number of cellular
transport mechanisms are involved in magnesium homeostasis, primarily the sodium/magnesium exchanger (Gunther, Vo rm ann, & Forster, 1984). The increase
in intracellular magnesium has a positive effect on protein synthesis, although
this process is inhibited at excessive concentrations (Rubin et al. 1997).
Other degradation and resorption processes are to be expected when magnesium is
used in implants. Magnesium is a light metal with a density of 1.74 g/cm
3
and a higher
Daily Mg intake:
300 mg
50% of total body Mg
is found in bone
Dietary absorption can
vary from 24–76%
Typical Mg
absorption: 120 mg
Net GI absorption 100 mg
GI secretion of Mg: 20 mg
Unabsorbed Mg: 180 mg Renal excretion: 100 mg
Figure 4.3 Magnesium metabolism. Of the 300 mg of magnesium ingested, approximately
120 mg are absorbed from the gut. A total of 20 mg are lost in gastrointestinal secretions, leaving
a net absorption of 100 mg. Patients in magnesium balance excrete all of this absorbed
magnesium in the urine. Bones provide a large magnesium buffer (Topf & Murray, 2003).
132 Surface Modification of Magnesium and its Alloys for Biomedical Applications
fracture toughness than that of ceramic biomaterials such as hydroxyapatite (Staiger,
Pietak, Huadmai, & Dias, 2006) (see Table 4.4). In human bodily fluids, magnesium
and its alloys exhibit relatively rapid corrosion (Song & Atrens, 1999, 2003). Solid
magnesium hydroxide, magnesium chloride, and gaseous hydrogen are formed during
the corrosion process (Li et al., 2008; Makar & Kruger, 1993; Song & Atrens, 1999;
Staiger et al., 2006; Wang, Wei, Gao, Hu, & Zhang, 2008) (see Figure 4.4). As degra-
dation progresses, a thin but stable layer of magnesium hydroxide is subsequently
formed beneath the magnesium hydroxide coating, leading to a local increase in pH
value (Barnett, 2007; Song, 2007). The magnesium oxide coating, i.e., the degradation
layer, is a white, crystalline, nonclosed outer layer that forms directly on the implant. In
vivo tests on animals show that the degradation products include magnesium hydroxide (Mg[OH]
magnesium calcium apatite in the form (Ca
), magnesium oxide (MgO) and magnesium chloride (MgCl2), as well as
2
1-xMgx)10
(PO4)6OH2(Erdmann et al.,
2011; Kuwahara, Al-Abdullat, Mazaki, Tsutsumi, & Aizawa, 2001; Li et al., 2008;
Staiger et al., 2006; Thomann et al., 2010; Wang et al., 2008; Witte et al., 2005; Xu
et al., 2 007) (see Figure 4.5). Part of the degradation layer (magnesium and alloy el-
ements) is derived from the implant itself and other parts (calcium and phosphate),
i.e., constituents of the body, are stored there (Witte, Nellesen, Crostack, & Beckmann,
2002; Xu et al., 2007). Its thickness, which tends to increase, depends on the alloy’s
composition (Krause et al., 2010). This degradation layer delays the initially rapid
corrosion process and also provides a certain degree of corrosion resistance (Makar &
Kruger, 1993). The resulting pH values close to the surface, which (temporarily) are
relatively high as corrosion progresses, may lead to the formation of magnesiumcontaining apatite in the form of (Ca
1-xMgx)10
(PO4)6OH2, which is deposited at the
implant surface (Liu, Huang, Shen, & Cui, 2001; Qi et al., 2008). In addition, gas production around the implant, the quantity of which corresponds to the degradation
speed, has often been observed during in vivo experiments (Li et al., 2008; von der
H€oh et al., 2006; Witte et al., 2005; Staiger et al., 2006). Nonbound Mg
2þ
ions can
be involved in normal physiological processes of metabolism. Hydrogen is not formed
in the human body, but may have a positive antioxidative effect on the cells. Alkalosis
and acidosis are compensated in the body e primarily via the kidneys and lungs e by
both bicarbonate and haemoglobin buffering systems (van den Berg, 2005).
The corrosion pattern, which is a function of the alloy composition and environ-
mental conditions, is usually initiated as a pitting-corrosion process (Song & Atrens,
1999).
The corrosion speed depends to a large extent on the purity of magnesium, alloy
components, and the fabrication process (Li et al., 2008; Pardo et al., 2008; Song,
2007). Even low levels of impurity in magnesium result, owing to galvanic effects,
in higher rates of degradation (Ren et al., 2007; Song, 2007; Song & Atrens, 1999;
Witte, Hort, et al., 2008). Less than a tenth of one percent of calcium can substantially
increase the corrosion resistance of magnesium (Kaese, 2002).
Increasing aluminium content has a corrosion-protective effect (Huang, Ren,
Jiang, Zhang, & Yang, 2007; Kaese, 2002). Recent studies have shown that alloy-
ingwithzinc(Huang et al., 2007; Pardo et al., 2008; Song, 2007) or rare-earth el-
ements (Kannan & Raman, 2008; Witte et al., 2006; Wu, Fan, Zhai, & Zhou, 2005)
Table 4.4 Summary of the physical and mechanical properties of various implant materials in comparison
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to natural bone
Synthetic
Properties Natural bone Magnesium Ti alloy Co-Cr alloy Stainless steel
Density (g/cm
3
) 1.8e2.1 1.74e2.0 4.4e4.5 8.3e9.2 7.9e8.1 3.1
Elastic modulus (Gpa) 3e20 41e45 110e117 230 189e205 73e117
Compressive yield
130e180 65e100 758e1117 450e1000 170e310 600
strength (Mpa)
Fracture toughness
Staiger et al. (2006).
(MPam
1/2
)
3e615e40 55115 N/A 50200 0.7
hydroxyapatite
Bioabsorbable behaviour of magnesium alloys e an in vivo approach 133
134 Surface Modification of Magnesium and its Alloys for Biomedical Applications
35
2
30
as-cast Mg-1Ca
as-extruded Mg-1Ca
25
20
15
10
5
Volume of evolved hydrogen/ml/cm
0
0 50 100
150 200 250
Immersion time/h
Figure 4.4 The hydrogen evolution volumes of as-cast and as-extruded Mge1Ca alloy samples
as a function of the immersion time in simulated body fluid (Li et al., 2008).
Figure 4.5 Schematic diagram of the alloy/solution biocorrosion interface: (a) the galvanic
corrosion between Mg and Mg
Mg-Ca alloys, (c) the adsorption of chloride ions to transform Mg(OH)
hydroxyapatite formation by consuming Ca
Ca phase, (b) the partially protective film covering the surface of
2
2þ
and PO
3
, and (e) the disintegrated particle-shape
4
into MgCl2, (d) the
2
residues falling out of the bulk substrate (Li et al., 2008).
Bioabsorbable behaviour of magnesium alloys e an in vivo approach 135
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results in a considerably slowed degradation process as compared with pure
magnesium.
Apart from the alloy compo nents, the fabrication process itself has a significant influence on the mechanical properties of magnesium (Li et al., 2008; Ma et al., 2009).
By reducing the grain size, it is possible to achieve both improved elasticity and higher
ductility (Koike, Ohyama, Kobayashi, Suzuki, & Maruyama, 2003; Zelin, Yang,
Valiev, & Mukherjee, 1992). Extrusion provides the materials with a dded corrosion
protection and enhanced mechanical properties (Kaese, 2002; Li et al., 2008).
Extruded magnesium-calcium alloys consisting of one percent calcium by weight
have excellent properties (Li et al., 2008).
Surface treatment and surface coating are further approaches to influencing
degradation behaviour (Gray & Luan, 2002; Zhang, Xu, & Yang, 2005). von der
H€oh et al. (2006) were able to demonstrate that the smooth implant surface of
magnesium-calcium alloys has a negative effect o n corrosion, unlike the case
with blasted surfaces (see Figures 4.6, 4.7, and 4.8). The findings of Gogolewski
(2000) were similar. Increased corrosion resistance was achieved by using gas
displacement to apply magnesium onto magnesium alloys (Yamamoto, Watanabe,
Sugahara, Tsubakino, & Fukumo to, 2001). von Staesche developed an inexpensive
procedure to coat implants with magnesium fluoride and was able to show a
reduced rate of degradation in the specimens (Staesche, 1948). When an implant
made of magnesium alloy has a fluoride coating, fluoride is deposited in the natural
magnesium hydroxide layer, which compresses and stabilises the natural layer
(Gnesca et al., 1996) and thus increases corrosion resistance (Chiu, Wong, Cheng, &
Man, 2007; Staesche, 1948). A coating of bioactive hydroxyapatite (Song, Shan, &
Han, 2008), the technique of plasma immersion (Liu, Xin, Tian, and Chu (2007),
anodizing of the specimen surface (Song, 2007) or alkali-heat treatment (Li, Gao,
& Wang, 2004; Lorenz et al., 2009) followed by slow degeneration has also
been described in the literature.
Figure 4.6 m-Computed tomography of a smooth implant after 3 months (a) and 6 months
(b) implantation duration: a small homogeneous resorption layer and a very close bone to
implant contact layer is recognisable (scale, 1 mm) (Von der H€oh et al., 2006).
136 Surface Modification of Magnesium and its Alloys for Biomedical Applications
Figure 4.7 m-Computed tomography of a sand-blasted implant after 3 months (a) and 6 months
(b) implantation duration: the degradation process affected the whole implant, after 6 months the
core body of the cylinder no longer existed. The bone to implant contact is less than in smooth
implants and threaded cylinders (scale, 1 mm) (Von der H€oh et al., 2006).
Figure 4.8 m-Computed tomography of a threaded implant after 3 months (a) and 6 months (b)
implantation duration: the hole shaped degradation at the edges and close bone contact around
the implant is shown (scale, 1 mm) (Von der H€oh et al., 2006).
The earliest investigations into pure magnesium’s suitability as an implant material date back to the beginning of the 20th century (Lambotte, 1932). The magnesium degraded very rapidly during these experiments, and gas production was
clearly evident. As early as 1932 (Lambotte, 1932), either the use of implants
made of pure magnesium was viewed critically or pure magnesium was deemed unsuitable for these purposes. However, no systemic adverse effects were discovered
(Lambotte, 1932; McBride,B 1938; Verbrugge, 1934). The first magnesium alloys
were produced in 1932. In vivo experiments with magnesiumealuminium alloys,
however, indicated rapid degradation and gas production, although the latter had
no obvious harmf ul effects on the organism, and the gas was completely resorbed
(Verbrugge, 1934 ).
Bioabsorbable behaviour of magnesium alloys e an in vivo approach 137
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Magnesium saw something of a renaissance at the beginning of the 21st century.
Various alloys were developed with the aim of controlling the pattern and speed of
the corrosion. Different alloy elements have an effect on degradation behaviour and/
or mechanical properties (Song, 2007; Pardo et al, 2008).
4.4.2 Volumetry by means of m-computed tomography
Several methods e with different approaches, and various advantages and disadvantages e are suitable for investigating the in vivo degradation process of
magnesium-containing implants. These techniques are presented here, and their benefits and drawbacks outlined.
This method enables volumetric analysis of the implant during the course of animal
experiments (see Figures 4.9 and 4.10). In this way, volume loss and the extent of
corrosion on the alloy can thus be determined in a noninvasive procedure, and the
duration of the experiment can be optimised (Holdsworth & Thornton, 2002; Paulus,
Gleason, Easterly, & Foltz, 2001). The limiting factors in the performance of the inves-
tigation are twofold: the spatial distribution of the gantry in clinical m-computed tomography (CT) scanners of up to around 70 cm and the size of the test animal. This
procedure takes considerably longer than with conventional CT scanning, and the animals must be anaesthetised. Artefacts often occur when using metallic implants (J€akel
& Reiss, 2007; Shalabi, Wolke, Cuijpers, & Jansen, 2007; Stoppie, Wevers, & Naert,
2007), but are less of a problem with magnesium implants (Witte et al. 2005, 2007a).
Micro-CT provides a means of presenting two-dimensional images of the implant
using three-dimensional geometry (Kiba et al., 2003; Stoppie et al., 2007). The latest
devices enable in vivo measurements in animals with a local resolution of 10e20 mm
(Brouwers, van Rietbergen, & Huiskes, 2007). Implants should be measured prior to
implantation to obtain comparable outcomes. The results serve as reference values for
intraoperatively performed measurements. Depending on the contrast ratios, the
implant can be scanned either automatically using software or manually. It should
be noted that specific thresholds must be determined and kept at a constant level during
the series of tests (Erdmann et al., 2011). Volume loss can be established using a volume subtraction technique.
4.4.3 Determination of weight loss
The determination of weight loss is another method for evaluating the degenerative
behaviour of magnesium implants (Li et al., 2008; Song, Bowles, & Stjohn, 2004;
Xu, Zhang, Yin, Zeng, & Yang, 2008). For this purpose, specimens are weighed prior
to implantation and, generally, also following euthanasia of the animal. Li et al. (2008)
investigated the weight loss of magnesiumecalcium implants in vivo and were able to
show that the weight of the magnesiumecalcium implants was increasingly reduced
over time subsequent to implantation.
Prior to implantation, the initial weight of each individual specimen must be verified using precision analytical scales. The specimen is reweighed after explantation
(although it must be borne in mind that the weight obtained includes all adherent corrosion products in the form of oxides and hydroxides). These corrosion products are not
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