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320 Surface Modification of Magnesium and its Alloys for Biomedical Applications
10.5 Testing mechanical properties in vitro
10.5.1 Mechanical tests without medium
Mechanical tests without medium are important to characterize magnesium alloys and
to compare the performance of medical devices under well-defined load situations, but
excluding the influence of corrosion.
For the material characterization in the ‘dry’ state, basic uniaxial tests for metals
according to international standards can be carried out using standardized sample
geometries and test configurations (compression: DIN 50,106; tension: ASTM
E 8M, EN ISO 527).
The degree of mechanical anisotropy in compression and tension can therefore be
checked. As (crystalline) magnesium alloys are not particularly temperature sensitive
at 37
C, it is usually sufficient to carry out these tests at room temperature. Stiffness,
yield strength (YS), ultimate tensile or compressive strength (UTS, UCS), and the
ductility of the material can be determined.
Additional tests lik e bending, torsion, or shear might also be used to analyse the
performance of a material. A bending test might be a fast and easy way to characterize
a material as compressive and tensile stresses occur simultaneously. Figure 10.1
shows a simple 3-point-bending test of metal wires. The example illustrates that a
magnesium wire has to be about 1.5 times thicker than titanium or steel to achieve
similar flexural properties. The magnesium wire with 3 mm diameter has comparable
bending stiffness to the steel wire with 2 mm diameter, i.e., both wires exhibit similar
140
120
100
80
60
Force (N)
40
20
0
02468101214
Displacement (mm)
Figure 10.1 Comparison of the bending behaviour of wires with 150 mm length by a threepoint-bending test with 60 mm span and a vertical displacement of 5 mm/min. Plunger and
supports have a radius of 5 mm. The steel wire (Synthes art. nr. 292.200) and the titanium wire
(Synthes art. nr. 492.200) are 2 mm in diameter, whereas the magnesium wire (alloy WE43,
as extruded) has a diameter of 3 mm.
2 mm steel wire
3 mm magnesium wire
2 mm titanium wire
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elastic behaviour. The maximum bending force for the magnesium wire (alloy WE43)
lies between the values of the stainless steel wire (316L) and of the titanium wire
(TiAl6V4 alloy).
Flexural properties can be measured according to the standard ISO 178 (for plastics). The deformed specimens can be used to roughly estimate the ductility of the
material by a subsequent folding test (EN ISO 7435).
For coated magnesium, further tests are necessary to characterize the mechanical
properties of the coating in the dry state (e.g., adherence to the substrate, scratch resistance). For shear testing of calcium phosphate coatings and metallic coatings, the
ASTM F1044 can be applied.
For established medical devices, existing standardized tests can be used to evaluate
their mechanical behaviour. In the area of osteosynthesis, for example, standards exist
for the testing of:
• metallic bone plates (e.g., bending tests according to ASTM F382, ISO 9585)
• metallic bone screws (e.g., torsion test according to ASTM F543)
• intramedullary rods and nails (bending and torsion tests according to ASTM F383, ASTM
F1264)
• angled orthopaedic fracture devices (ASTM F384)
• bone staples (pull out and bending according to ASTM F564).
10.5.2 Testing of in vitro degraded samples without load
For degradable polymers, a typi cal way of in vitro testing is described in the ASTM
F1635 and ISO 13781 standards. Semifinished components or finished implants are
immersed in buffered saline solution at physiological temperature (i.e., PBS at
37
C) for specific periods of time. The samples are periodically removed and tested
for material or mechanical proper ties (e.g., percent weight loss, inherent viscosity,
strength loss).
A limitation of this way of testing is that the used saline solution only roughly simulates the in vivo environment. It does not account for the local blood flow, the water or
fat content, or the microbiologic interaction with cells in a specific tissue. Another
limitation is that the combined effect of stress and degradation is not considered.
For example, the polymer might experience a considerable degree of relaxation, i.e.,
the implant cannot retain the initial load as it gradually yields to the imposed stress
due to the imposed temperature (creep) and due to the presence of water (increased
mobility of polymer chains).
For magnesium, a similar prolonged (static) immersion in saline solution is not
advised because the conditions cannot be kept as constant as in the human body.
The release of hydroxide ions increases the pH (as soon as the buffering capacity of
the solution is consumed), and the release of magnesium ions increases the osmolality
of the solution. The released hydrogen gas might form bubbles and could locally mask
the surface. As a consequence, regular medium changes or a dynamic flow system are
advised to approximate conditions to a physiological range. Figure 10.2 shows the
example of a lag screw that has been degraded in simulated body fluid (SBF) for
322 Surface Modification of Magnesium and its Alloys for Biomedical Applications
Figure 10.2 Example of a generic magnesium screw after 12 weeks of immersion in 250 ml
simulated body fluid with weekly medium change (according to Imwinkelried et al. (2012)).
(a) Screw with degradation products. (b) Screw after removal of the degradation products by
brushing.
12 weeks with weekly medium changes as described by Imwin kelried et al. (2012).
Because the screw was fixed inside a tensioning device, some of the broad threads
at the tip are not degraded. The degraded neighbouring threads and the screw shaft
have kept their initial shape (Figure 10.2(a)). The outer shape of the screw remained
unchanged as the metallic magnesium has been replaced by a thick conversion layer.
These white degradation products are then removed with a common hand brush, and
the nondegraded magnesium core becomes visible (Figure 10.2(b)). The ductility of
the remaining magnesium core can be examined with an additional bending test.
Gas release during immersion can be collected and compared to the weight loss of
the screw with and without degradation products.
This example shows that magnesium specimens can be degraded in vitro for various
periods of time, then be removed and tested outside the ‘bath’ e just as for the tests
without medium. Depending on the test, the degradation products might be dehydrated
first (e.g., by rinsing with ethanol and subsequent drying) or even removed mechanically or with a chemical agent that does not attack magnesium but dissolves the formed
calcium phosphate conversion layer (e.g., hydrofluoric acid).
10.5.3 Testing of in vitro degrading samples with load
After implantation, a medical device might be subject to a certain base load. F or
example, the above-mentioned lag screws are used in traumatology to put broken
bone pieces under compression. The screw itself is experiencing tensile stresses
between the screw threads and the screw head. The higher the insertion torque
of the screw, the larger are the tensile stresses along the screw axis. The screw
shown in Figure 10.2 was tested at a low insertion torque and did not fail within
12 weeks in SBF. With higher insertion torques, the combined effect of tensile
stress and corrosion attack might lead to premature failure due to stress corrosion
cracking. A coating could be helpful in preventing stress corrosion cracking
because it might keep the corrosive medium away from the magnesium surface.
However, it is important to make sure that the coating is also effective when the
implant i s under load.
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10.5.4 Testing of in vitro degrading samples after plastic
deformation
If implants are submitted to plastic deformation during the implantation, it is necessary
to test the magnesium samples in vitro after deform ing them permanently. This kind of
test is particularly important if coatings or other surface modifications are used. Due to
the plastic deformation of the base material, a coating might locally crack or even
delaminate and thus induce a completely different degradation behaviour (e.g., with
more pronounced localized corrosive attack). In addition to the plastic deformation,
a given load might be applied to test worst-case conditions.
An example for in vitro testing of plastically deformed samples with applied load is
shown in Figure 10.3. A rectangular plate is first deformed in a bending template with a
defined curvature (Figure 10.3(a)). In a second step, the plate is elastically deformed to
fit inside a polymer implant holder. Figure 10.3(b) shows how the exerted force can be
measured before immersing the floating device into SBF. This measurement can be
repeated at each medium change to track the force decrease over time and to characterize the strength retention of the alloy. Figure 10.3(c) shows the tensioned magnesium specimen with the implant holder floating at the (SBF) surface of a
gas-collecting tube.
Premature failure of the sample might occur before the sample has degraded to a
point where no load is left (Figure 10.3(d)). In addition, the released gas can be
collected. The amount of released gas can be compared to the weight loss of the
samples.
10.5.5 Fatigue testing under cyclic stress
Besides a given base load as mentioned above, implants will, in most cases, be submitted to cyclic and/or dynamic loads. As an example, a cardiovascular stent is deployed
by the expansion of the balloon and exerts a radial pressure on the artery wall to keep
the vessel open. In addition to this base load, the pulsating blood flow puts the stent
under cyclic stress. An adequate in vitro test is therefore needed to determine the
Figure 10.3 Immersion test of a rectangular magnesium specimen after plastic deformation and
tensioning. (a) Controlled deformation in a bending template. (b) Force measurement of the
spring type specimen before fixation in the implant holder. (c) Immersion in simulated body fluid
with a floating PE polymer holder. (d) Example of a broken specimen after a immersion.
324 Surface Modification of Magnesium and its Alloys for Biomedical Applications
Figure 10.4 In vitro testing of a magnesium plate/screw construct with a simulated fracture gap.
(a) Static bending test. (b) Cyclic bending test in distilled water.
failure behaviour of the stent in such harsh conditions (i.e., corrosive environment,
plastic deformation, cyclic stresses).
Another example can be found in osteosynthesis. When a bone fracture is fixed
using plates and screws, the plate is usually prebent to follow the contour of bone,
and the screws are fixed to the bone in such a way that the fracture is well reduced
and eventually put under compression (R€uedi & Murphy, 2000). As a consequence,
the plastically deformed plate is submitted to a basic load. Cyclic loads come in addition as early mobilization of the patient is sought to accelerate bone healing.
A major difficulty in fatigue testing is to define adequate testing conditions. The
forces occurring in vivo are often not well known and can be multiaxial. The aggressiveness of the used medium and exposure time are important parameters to consider in
defining adequate cycle times.
For uncoated magnesium, distilled water might be a good medium to start with. On
one side, distilled water does not give deposits on the magnesium surface. On the other
side, the magnesium ions that are leached out of the material usually do not change the
pH as Mg(OH)
is formed. The important effect of distilled water is that it can alter the
2
occurrence and growth of cracks. Premature failure therefore will occur at lower load
levels and/or at shorter periods than if no medium is used.
Figure 10.4(a) shows a static bending test of a magnesium plate that is fixed to two
polymer tubes with screws to simulate a bone fracture. For cyclic bending tests in an
aqueous medium, it is important to avoid contact with other metals to prevent galvanic
corrosion, which could falsify the degradation conditions. In the example shown in
Figure 10.4(b), a carbon fibre rod (black) is used instead of the metal plunger used
without the medium.
10.6 Testing implant performance in vivo
10.6.1 Suitability of animal models
The suitability of animal models in testing the performance of an implant in vivo
depends on how well the human environment of a particular indication can be simulated in an animal tissue. Small animals such as mice, rats, or rabbits are often used for
Testing the mechanical properties of surface-modified magnesium 325
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first screening tests of degradable implant materials. For testing the performance of a
human implant, however, small animal testing might not be sufficient. Frink et al.
(2011) recommends the use of large animals if physiological processes are to be inves-
tigated, which is the case for magnesium degradation.
The metabolism of an animal is of considerable importance when degradation is
involved. Degradation rate might depend on the size of the animal. Witte et al.
(2008) report that the blood flow in the human body, for instance, is an order of magni-
tude higher than in a small animal (rat, rabbit). In a rat study with the slow-degrading
magnesium alloy WE43, Tschegg et al. (2011) revealed a significant increase of push
out force of a pin implanted in the distal femur compared to titanium and/or PLGA
polymer pins of same geometry. Because the degradation-related gas release was
slow enough to avoid the formation of gas bubbles, it did not interfere with bone apposition, and the receding implant was replaced by bone. The same material implanted on
top of the nasal bone of a minipig, however, lead to the formation of gas bubbles
(Imwinkelried, Beck, Iizuka, & Schaller, 2012).
The blood flow in miniature pigs is more similar to humans than that in small animals. Because the physiology of the miniature pig’s heart resembles the human heart,
these animals are often used to test the in vivo performance of cardiovascular stents.
Wang, Liu, Fang, and Shi (2007) consider the miniature pig to be a useful large
animal model for dental and orofacial research. Minipigs are not ruminants (like sheep)
and have a temporomandibular joint similar to the humans. If bone healing is studied,
the use of skeletally mature animals is advised.
10.6.2 Testing of in vivo degraded samples
One method to test the in vivo degradation behaviour of magnesium is to place an
implant inside a given tissue of an animal and to let it degrade for a defined period.
After explantation, mechanical testing of the partially degraded implant s can be done.
Thomann et al. (2010) implanted 25-mm-long McCa0.8 pins with 2.5 mm diameter
in the intramedullar cavity of rabbit tibiae for 3, 6 and 12 months. After explantation,
three-point-bending tests of the degraded pins were carried out, and the strength retention of the material could be determined.
Imwinkelried et al. (2012) used large rectangular plates that were implanted on top
of the nasal bone of miniature pigs for up to 24 weeks. Figure 10.5(a) shows the
removal of magnesium plate after 12 weeks of implantation. Figure 10.5(b) illustrates
the three-point-bending testing of the in vivo degraded rectangles. To compare the
strength retention of in vivo and in vitro degradation, identical plates were immersed
in SBF (SBF27) for up to 12 weeks. In vitro degradation was found to occur about four
times faster than in vivo.
10.6.3 Extrapolation of mechanical properties from residual
volumes
To avoid the sacrifice of an animal for the in vivo degradation of an implant, the
residual volume of an implant might be determined using in vivo imaging
326 Surface Modification of Magnesium and its Alloys for Biomedical Applications
Figure 10.5 Mechanical testing of in vivo degraded magnesium plate. (a) Removal of plate from
the nasal bone of a miniature pig after 3 months. (b) Three-point-bending test of in vivo degraded
plate.
techniques (e.g., live computed tomography (CT)). Numerical modelling might
then be used to estimate the remaining strength and stiffness of the implants. In
practice, it may not be possible to determine the residual volume of an implant
due to the limited image resolution. Furthermore, the conversion zone (i.e., the
volume of the implant that has been transformed into degradation products)
supposedly has a similar density to the magnesium because it can hardly be distinguished from nondegraded material.
Witte et al. (2010) were able to distinguish between metallic magnesium and the
corrosion layer by using synchrotron-radiation-based micro-CT (SRmCT) on
explanted rat femurs. They were also able to determine the corrosion rate based on
the pit depths of coated and noncoated cylinders. However, they did not extrapolate
the remaining mechanical proper ties from these residual volumes.
Cho et al. (2012) have used conventional postmortem micro-CTs (mCTs) on
explanted bone blocks from rabbit femurs containing magnesium screws. On strongly
degraded samples, they were able to determine the residual equivalent diameter of the
screws and to determine the remaining bending and torsion strength using numerical
modelling.
As can be seen by these two examples, the determination of residual magnesium
volumes either needs sophisticated imaging techniques or a sufficiently thick conversion zone. An in vivo determination of residual implant volumes remains very challenging, even if image resolution could be further improved.
10.6.4 Testing of implants using functional animal models
The registration process of a medical device requires the proof of its safety and performance. As a consequence, the intended use of an implant system has to be tested using
a functional animal model. The results gained from such preclinical tests are the basis
for the regulatory approval of an implant system or for the respective human clinical
trial required for getting such approval from the authorities.
Cardiovascular magnesium stents, for instance, have been implanted into the coronary arteries of miniature pigs exactly as in the human procedure. Wittchow et al.
Testing the mechanical properties of surface-modified magnesium 327
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(2013) showed that the efficacy and healing characteristic of a magnesium stent was
comparable to permanent metalli c stents in the porcine model used.
For the testing of osteosynthesis implants, few functional animal models are available because the conflict among clinical comparability, standardization, and reproducibility remains a major problem (Frink, Andruszkow, Zeckey, Krettek, & Hildebrand,
2011).
A common way to simulate a bone fracture is to apply an osteotomy (i.e., to cut a
bone in two) and fix the bone pieces with the implant system being tested. The use of
large animals is important not only because a similar metabolism should be sought, but
also has the advantage that human-size implants can be used.
Furthermore, biomechanics in the animal often represent a worst case, but static and
dynamic forces occurring are often unknown and are difficult to measure.
10.6.5 Likely future trends
The future in testing of mechanical properties of magnesium alloys will depend on
which biomedical applications will be successful. The requirements of the regulatory
bodies for the registration of magnesium implants will also influence the amount and
the quality of the test data.
The requirements on test planning, documentation, and statistical analysis as well as
for the validation of experimental methods and setups are likely to increase. A standardization of the in vitro testing of mechanical properties of magnesium alloys might
be reached.
Accelerated in vitro degradation might become a way to simulate the in vivo
behaviour of magnesium implants and to reduce the amount of animal testing.
The use of sophisticated in vivo imaging techniques, the implantation of sensors,
and numerical modelling of the biomechanics will improve our knowledge of the
animal models used.
10.7 Sources of further information
The following books by Woodhead Publishing can provide further information about:
• Fundamentals of magnesium alloy metallurgy (Pekguleryuz, Kainer, & Kaya, 2013)
• Metals for biomedical devices (Niinomi, 2010)
• Corrosions of magnesium alloys (Song, 2011)
• Coatings for biomedical applications (Driver, 2012)
Further information about mechanical testing can be found in:
ASM Handbook Volume 08: Mechanical Testing and Evaluation (Kuhn & Medlin, 2000)
Since 2009, a dedicated symposium on biodegradable metals for biomedical applications has taken place every year (Berlin 2009, Maratea 2010 and 2012, Quebec City
2011, Umang Island 2013). The related abstracts and publications can provide further
information about testing media and methods.
328 Surface Modification of Magnesium and its Alloys for Biomedical Applications
Acknowledgements
All figures in this chapter are a courtesy of Synthes GmbH, Oberdorf, Switzerland, and have not
been published previously. I thank my colleagues and superiors from the biomaterials department for their support and in particular Dr Andrea Montali for the revision of the text. I also
acknowledge the support of the prototype shop and of the materials testing department for
manufacturing and testing of the magnesium specimens.
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