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310 Surface Modification of Magnesium and its Alloys for Biomedical Applications
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Testing the mechanical properties
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of surface-modified magnesium
10
and magnesium alloys for
biomedical applications
T. Imwinkelried
Formerly Synthes GmbH, Oberdorf, Switzerland (Presently at Robert Mathys Foundation)
10.1 Introduction
10.1.1 Biomedical applications of magnesium implants
The concept of biodegradable magnesium implants is not new. The history of biodegradable magnesium implants has been recently reviewed by Witte (2010). He reports
a multitude of biomedical applications that were investigated about a century ago, but
fell into oblivion after the 1940s.
The goal of a degradable implant is to fulfil a biomedical function and to disappear
within a defined duration with acceptable side effects. Degradable implants are particularly attractive if a temporary (nondegradable) implant would be difficult to remove
or if a permanent remainder of the implant would present an increased complication
risk for the patient.
A use of bioabsorbable materials should allow tissues to regain their natural state
and ability to transform and, in the case of children, to continue to grow.
Magnesium implants could offer additional benefits compared to existing implant
materials. They might replace permanent metallic implants for indications in which
degradation is a significant advantage. They could also allow the treatment of indications in which degradable polymers are limited by their mechanical properties.
There might be additional benefits like the renouncement on antiplatelet medication,
which is needed when permanent cardiovascular stents are used (Haude et al., 2013).
10.1.2 The life cycle of a degradable implant
The requirements for a degradable implant can vary considerably during its life
cycle e from the implantation through the fulfilment of the desired function until complete degradation.
During implantation, strength and ductility, for example, can be decisive parameters. Strength is important to avoid interoperative failure of the implant (e.g., breaking
of a screw head), and ductility is important if the implant is plastically deformed during
surgery (e.g., expansion of a stent).
Surface Modification of Magnesium and its Alloys for Biomedical Applications. http://dx.doi.org/10.1016/B978-1-78242-077-4.00010-3
Copyright © 2015 Elsevier Ltd. All rights reserved.
312 Surface Modification of Magnesium and its Alloys for Biomedical Applications
During the fulfilment of the implant function, parameters like strength retention and
fatigue strength are important for a safe and effective achievement of that function. An
osteosynthesis plate/screw construct, for example, should maintain the reduction of the
bone fracture to allow healing and should not break before the healing process is
sufficiently advanced (R€uedi & Murph y, 2000).
Finally, during the degradation process, the magnesium implant should disappear
within an acceptable time frame without causing significant side effects. The degradation products should be bioabsorbed (i.e., transported away or metabolized).
One should note that these life-cycle phases might overlap, as some degradation
could already occur during implantation. In many cases, however, implantation time
is short compared to the total life time of an implant.
10.1.3 Goals of mechanical testing
The goal of mechanical testing is to gain knowledge about the performance of a
magnesium alloy or of a magnesium implant until fulfilme nt of its function. It is aimed
at simulating worst-case conditions and at exploring the limits of a safe use. With
regard to regulatory approval of an implant system, mechanical testing should supply
supportive data for proving the safety and efficacy of the medical device.
10.1.4 Assumptions about mechanical anisotropy of
magnesium
Magnesium and magnesium alloys have a hexagonal close packed atom lattice. At
ambient temperature (i.e., in the operating room), the plastic deformation of magnesium is governed by a limited number of slip systems and the possible formation of
twins. As a consequence, yielding in tension or in compression will occur at different
stresses for a magnesium single crystal. The degree of anisotropy might be reduced in
polycrystalline magnesium by reducing the grain size. Magnesium alloys with grain
sizes less than 4 mm have been reported to have very little tension-compression yield
stress anisotropy (Gunde, H€anzi, Sologubenko, & Uggowitzer, 2011). In the
following, it will be assumed that magnesium behaves as an isotropic material and
can thus be described by mechanical values that are not direction-dependant (e.g.,
single values for the young’s modulus, the shear modulus, the yield strength).
Anisotropic mechanical properties would make predictions of implant behaviour
more difficult.
10.1.5 Chapter outlook and methodology
An accurate simulation of the combined effects of metabolism (i.e., degradation) and
of biomechanics (i.e., load condition) by a testing apparatus is a challenging task. As
the mechanical testing of a degradable implant system is more demanding than for a
permanent implant system, an appropriate and magne sium-specific testing methodology should be adopted.
Testing the mechanical properties of surface-modified magnesium 313
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Table 10.1 Examples of mechanical properties which might be
relevant during the life cycle of a biodegradable magnesium implant
and potential influence of surface modifications on these properties
Influence of surface
modification
Life cycle Material property
(Chapter 10.4)
Implantation related
(Chapter 10.2)
Degradation related
(Chapter 10.3)
Stiffness
Strength
Ductility
Hardness
Toughness
Stress corrosion resistance
Strength retention
Fatigue resistance
Abrasion resistance
No
No
No
Yes
No
Yes
Yes
Yes
Yes
Table 10.1 gives an overview of the theoretical part of the chapter, in which the
potentially relevant material properties are introduced in Sections 10.2 and 10.3.
The influence of surface modifications on these properties is discussed in Section 10.4.
Table 10.2 introduces a general methodology for the testing of magnesium alloys
under increasingly demanding test conditions. It relates to examples discussed in the
practical Sections 10.5 and 10.6.
10.2 Implantation-related mechanical properties
10.2.1 Material requirements during implantation
The mechanical requirements of magnesium as an implant material depend on their
medical indication. During the surgical procedure of implantation, the material might
be submitted to peak loads or deformations that can exceed the subsequent values
during the function as an implant.
For example, a cardiovascular stent is significantly deformed during the expansion
of the balloon catheter. The struts of the stent are stretched, exceeding the limit of elasticity, and will remain permanently deformed. However, this plastic deformation
should not negatively affect the subsequent degradation behaviour.
Another exa mple of plastic deformation during implantation is osteosynthes is
plates that are used to fix bone fractures. The surgeon usually needs to adapt the shape
of th e plate to the contour of the bone. Pliers are used to prebend the bone plates and
might leave impressions on the surface of the implant. Beside s geometrical factors of
the plate, the material needs to be ductile enough to allow contouring without
cracking.
314 Surface Modification of Magnesium and its Alloys for Biomedical Applications
Table 10.2 Methodology for the testing of magnesium alloys and
implants under increasingly severe conditions
Example
Test condition In vitro In vivo (or ex vivo)
(Reference)
Without medium at
room temperature
With medium at
37
C, without
load
With medium at
37
C, static load
With medium at
37
C, cyclic load
With medium
at 37
C,
pre-deformation &
load
Uniaxial or
multiaxial tests
(tension,
compression,
shear, torsion,
bending)
Immersion in
medium for
various periods,
removal and
mechanical
testing
Elastic pre-
tensioning of a
device,
immersion in
medium until
failure
Immersion under
cyclic load until
failure
Immersion of
previously
deformed
specimen under
load conditions
simulating the
in vivo situation
Ex vivo:
Biomechanical
tests with
implants using
harvested tissues
(i.e., from
cadavers)
Implant in animals
for various
periods, explant
and test
mechanically
Implant device in
healthy tissue
with controlled
preload
Implant in healthy
tissue with known
cyclic load
Implant in
pathological or
modified healthy
tissue to restore
function
(according to
intended use)
In vitro
3-point-bending
test of
magnesium,
titanium, and
steel wires
(Chapter 10.5.1)
Strength retention
test with in vivo
and in vitro
degraded
specimens
(Chapter 10.6.2)
In vitro stress
corrosion test of
a lag screw fixed
with a given
torque (Chapter
10.5.3)
In vitro cyclic
bending test of a
plate/screw
device for bone
fixation (Chapter
10.5.5)
Preclinical study of
a magnesium
stent in a porcine
coronary model
(Wittchow et al.,
2013)
The in vitro and in vivo columns give an overview of possible procedures for the testing of biomedical magnesium devices.
The text in italic refers to in vivo test for which no examples could be found. The referenced examples show how these
procedures can be concretized.
For bridging the fracture, the osteosynthesis plate needs to be fixed to the bone
pieces. This is commonly done using screws. During the insertion of the screw, the
friction between bone and the screw threads needs to be overcome. The material needs
to withstand multiaxial loads (e.g., torsion, shear, compressio n) that are transmitted by
the screw driver.
Testing the mechanical properties of surface-modified magnesium 315
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These examples show how different the mechanical requirement for a material can
be during implantation.
10.2.2 Strength and stiffness
In most appli cations, the strength and stiffness of the material play an important role
for the design of the implant. If degradation can be neglected during the surgical
implantation procedure, ‘classical’ mechanical engineering might be applied to
determine a material-specific geometry of the implant.
The compliance of a material towards a mechanical stress that does not permanently
deform the material is generally expressed as rigidity or stiffness of the material. For
metals, the deformation caused by a mechanical load can be described by a linear
stressestrain relation. The proportionality constant reflects the stiffness of the material
in tension/compression (Young ’ s modulus E), uniform compression (bulk modulus
K), or in shear (shear modulus G). As magnesium is more ‘compliant’ than steel or
nitinol, the struts of a cardiovascular magnesium stent, for example, need to be thicker
if a comparable recoil has to be achieved. The AMS-3.0 magnesium stent (Wittchow
et al., 2013), for instance, has a cross section of 130 120 mm, whereas the stainless
steel reference stents have strut thicknesses of 97 and 85 mm.
The limit of elasticity is known as yield strength and usually is defined for a given
percentage of permanent deformation. As many metallic alloys harden during the
permanent deformation (‘strain hardening’), the maximum strength is generally above
the yield strength. The characteristic of the stressestrain curve can be important if the
implant is deformed during implantation (e.g., for the deployment of a cardiovascular
stent).
For the design of magnesium screws, the shear strength of the material might be a
limiting factor because it dictates the maximum torque that can be transmitted during
implantation. Titanium or steel osteosynthesis screws are often self-tapping, which
means that they can be inserted directly into a drilled bone hole (R€uedi & Murphy,
2000). For a magnesium screw with ident ical diameter, pretapping of the screw thread
might be needed to avoid premature failure.
10.2.3 Ductility
The ability to plastically deform is a major benefit of metallic implants. A permanent
deformation of the implant might be necessary to adapt to the anatomy of the patient.
For example, metallic osteosynthesis plates are often prebent to embrace the shape of
the bone pieces that need to be fixed. Cold bending is usually not possible with polymeric implants for which preheated water baths can be used to adjust the shape of the
plates.
In case of a metallic stent, ductility is of uppermost importance. During the inflation of the balloon inside a blood vessel, the mesh implant is subjected to plastic deformations that exceed the ductility of many common magnesium alloys. As a
consequence, a magnesium alloy used for a stent should be as ductile as
possible e a goal that can be achieved by appropri ate alloyin g and thermom echa nical
processing.
316 Surface Modification of Magnesium and its Alloys for Biomedical Applications
10.2.4 Other properties
Besides stiffness, strength, and ductility, other ‘static’ mechanical properties can be of
importance during and after the implantation.
The hardness of a material is its ability to resist a permanent local deformation
of its surface (e.g., by indentation or by scratching). For example, H€anzi et al.
(2011) report a fully degradable wound-closing rivet for tissue joining and
fixing via endoscopic surgery in the gastrointestinal tract consisting of a
fast-degrading magnesium tip for tissue penetration and of a slow-degrading
polymeric part.
Similarly, the tip of a Kirschner wire (so-called K-wire used in osteosynthesis)
should be harder than cortical bone in order to drill into it. The penetration of a
magnesium K-wire into bone will be less easy than for harder ma terials like stainless
steel. However, a hard coating of the magnesium wire might improve the insertion
into bone.
The toughness is indicative of the amount of energy a material can absorb before it
fails. In fracture mechanics, the resistance towards crack propagation is usually
measured using precracked compact tension specimens (Kuhn & Medlin, 2000).
10.3 Degradation-related mechanical properties
10.3.1 Material requirements during degradation
The integrity of an implant is particularly important when mechanical loads are
involved. During degradation, the implant should not loose its integrity prior to the
fulfilment of its mechanical function. If failure occurs, it should be in a controlled
and reproducible manner without causing unwanted side effects. The risk for the
patient should always remain acceptable.
To avoid failure of the degradable medical device, a material-specific design has to
be adopted. A reliable degradation behaviour is one of the basic material requirements
that need to be fulfilled before starting to design an implant. Ideally, the material would
degrade in a homogenous way, irrespective of inflicted plastic deformation and elastic
load condition. In practice, most magnesium alloys degrade by localized pitting.
Designing an implant becomes possible if the pitting does not lead to large preferential
degradation and if a threshold for the failure load of the degrading material can be
found.
Before determining the strength retention and the fatigue strength during degradation, the premature failure due to a combination of stress and corrosive attack e
hereafter referred to as environmentally assisted cracking or stress corrosion
cracking e should be avoided.
A magnesium alloy with a high susceptibility towards stress corrosion cracking
should not be considered for application as an implant material (even if a coating is
applied).
Testing the mechanical properties of surface-modified magnesium 317
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10.3.2 Environmentally assisted cracking
Environmentally assisted cracking is also known as stress corrosion cracking and
refers to premature failure of a part due to the combined effects of load and corrosion
attack. Its phenomenology and mechanisms are described in Chapter 4.
For the in vitro testing, human blood pl as ma can be c ons ide red as a sa li ne
solution that contains a variety of ions within certain compositional ranges.
Some of these ingredients are more aggressive to magnesium than others. In particular, the content of chlorine ions is a mon g the mo st imp orta nt sp ec ies t o cons ider .
Chlorine ions are presumably att acki ng the passivation layer on the surface of magnesium alloys and can therefore enhance the aggressiveness of the saline solution
(Song, 2011).
Besides human blood plasma, proteins, fat, and other metabolic species
contained in the respective tissue at the implantation site contribute to the specific
environment. A stent, for example, will initially be subject to the blood stream and
then become covered by endothelial cells. A osteosynthesis screw inserted into
cortical bone will encounter a different environmentthanifinsertedintocancellous
bone.
The surface condition is important for the occurrence of environmentally assisted
cracking. The surface roughness can have an influence on crack initiation. A smooth
surface, for example, can delay the occurrence of cracks and therefore delay the failure
of the implant due to environmentally assisted cracking. Ways to prevent or delay
stress corrosion cracking, for example, by biocompatible coatings, are discussed in
Chapter 4.
10.3.3 Strength retention
Strength retention is an important characte ristic of a degradable material, provided
that an implant does not fail prematurely and that degradation proceeds in a rather
homogenous way . It is conc ise information for the surgeon about how fast a degradable material loses its holding power and how much strength an implant can retain
during the fulfilment of its function. As an example, resorbable polylactide glycolide (PLGA) polymer implants for cranio-maxillofacial s urgery are expected to
retain at least 60% of their initial strength for up to 8 weeks after implantation, a
period long enough to allow for sufficient healing of fractured or osteotomized
bones.
In the case of magnesium, degradation proceeds from the surface to the inside and
is not a bulk phenomenon as for many polymers. As a consequence, the strength
retention depends on the thickness of the implant. Thinner implants will loose their
strength much faster than thicker implants. For a rectangular plate with 0.5 mm initial
thickness as an example, a degradation depth of 0.1 mm will reduce the bending
strength by a much greater extent than for a plate with 1.5 mm thickness. Strength
retention is therefore not a material property, but must be evaluated for each implant
individually.
318 Surface Modification of Magnesium and its Alloys for Biomedical Applications
10.3.4 Fatigue strength
Beyond strength retention, the fatigue strength becomes important in applications with
cyclic loads and might become the limiting factor for the design of an implant.
As an example, a cardiovascular stent is exposed to cyclic blood flow changes that
submit blood vessels to swelling/expansion and contraction movements. A strut of a
stent has to withstand several million of such load cycles without failing (1 year
corresponds to about 30 million heart beats).
Another example to illustrate cyclic loading is to imagine a broken rib that is fixed
with osteosynthesis plates and screws. On the one hand, the plated rib would be
exposed to the expansion and contraction movement induced by the breathing. On
the other hand, less frequent loads from laying down on the side would also have to
be considered. As a consequence, a combination of low-cycle and high-cycle fatigue
might have to be taken into account for the fatigue testing.
Fatigue testing in a corrosion environment is sometimes also referred to as corrosion fatigue. The effect of surface treatment on the fatigue life of biomedical magnesium devices is the topic of Chapter 8.
10.3.5 Other properties
In addition to the above-mentioned properties, the interaction of magnesium implants
with biological tissues could lead to a local increase or decreas e of degradation. As a
hypothetical example, consider a tendon that is frequently sliding over a magnesium
implant. Local abrasion of the metal surface and/or of the degradation products might
occur and eventually compromise the performance of the implant system. If such
effects would become clinically relevant, the abrasion resistance of a magnesium
implant could be improved by an adequate surface coating.
Wear resistance is another property that should be considered when the implant
system consists of several parts that are in contact and move with respect to each other.
If the relative movements are small, the wear between contact surfaces is also referred
to as fretting corrosion. The combination of friction and adhesion can also lead to a
wear phenomena called galling (i.e., when sliding surfaces are microscopically roughened and thus strongly adhere to each other).
10.4 Effect of surface modifications on implant
performance
10.4.1 Goals of surface modifications
The main goal of surface modification of magnesium implants is the improvement of
their performance during the application. Surface modification(s) can influence degradation behaviour, stress corrosion resistance, fatigue strength, wear resistance, optical
appearance of the implant, and other properties. Surface modifications can also be used
to fulfil additional functions such as local drug delivery.
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Wang et al. (2012) have reviewed surface modifications of magnesium alloys for
orthopaedic applications. A number of chemical and physical surface modifications
as well a s combinations of such methods are available.
10.4.2 Effect of surface structure
For titanium implants, increasing the surface roughness can improve or accelerate the
adhesion of bone cells (e.g., in dentistry, osteosynthesis). On the other hand, an
increased surface roughness might have a negative effect on the fatigue resistance
and/or on the stress corrosion cracking tendency of the implant. Surface roughness
can locally act as a stress raiser and therefore be responsible for the initiation of surface
cracks.
For a degradable implant, the surface roughness might change during the course of
degradation.
A polishing treatment can help to avoid or delay the formation of such cracks.
Mechanical, chemical or electrochemical polishing techniques might be applied to
smooth the surface. However, contact with the polishing media can lead to a contamination of the surface and thus potentially compromise the biocompatibility or the
degradation behaviour.
If a coating is applied, the surface roughness might be increased, and the positive
effect of delaying the initial degradation might therefore be compromised by the negative effect of increased surface roughness.
10.4.3 Effect of surface chemistry
The surface chemistry of a magnesium implant might be influenced by the shaping
operations (e.g., extrusion, machining, rolling, forging, laser cutting) and eventual
postprocessing operations (e.g., handling, polishing, storage) used.
Contaminations of the surface can modify the interaction of the implant with the
biological environment. For example, fingerprints on a newly machi ned magnesium
surface or the storage in a humid atmosphere can alter the surface chemistry.
Intended surface modifications usually reduce the reactivity of the blank magnesium surface. A coating, for instance, could keep away the corrosive medium from
the bulk magnesium material and thus delay the occurrence of stress corrosion.
10.4.4 Effect of internal stresses
Thermomechanical treatments of the surfa ce can lead to local plastic deforma tion and/
or residual internal stresses. A surface modification can modify existing or induce new
internal stresses in the surface layer of the implant. Metallurgical factors like grain
size, grain boundaries, and dislocation density can therefore differ from the bulk
material.
These internal stresses might have positive or negative effects on the properties of
the implant as they eventually add to the externally applied stresses.
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