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Файл:Ординатура / Хирургия / Библиотека им академика М.И. Перельмана / Книга_5370_Библиотеки_им_академика_М_И_Перельмана
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Surface design of biodegradable magnesium alloys for biomedical applications 95
the sputtering rates of common metals vary within an order of magnitude, and thus
another point distinguishing sputter deposition from evaporation or CVD is that
sputtering preserves the stoichiomet ry of the target source because the physical
bombardment mechanism of particle ejection results in a consistent stoichiometry
on the sample surface (Mattox, 2010b; Wasa et al., 2012).
3.2.3.3 Ion plating
Ion plating is an atomistic vacuum coating process in which the deposited film is
continuously or periodically bombarded by energetic atomic inert or reactive particles
that can affect the growth and properties of the film. The depositing atoms can come
from vacuum evaporation, sputtering, or arc vaporization. Bombardment prior to
deposition is used to sputter clean the substrate surface, while bombardment during
deposition is used to modify and control the properties of the film. It is crucial that
bombardment is continuous between cleaning and deposition in the process to maintain an atomically clean interface. The bombarding species are generally ions accelerated from a plasma in the deposition chamber (ions for bombardment are extracted
from the plasma and so termed plasma-based ion plating) or ions from an ion source
(ion plating is performed in a vacuum environment and so termed vacuum-based ion
plating). Figure 3.5 shows the two variations. The individual processes in ion plating
can be separated into surface preparation, nucleation and interface formation, and film
growth. Ion plating can also be considered a special process that varies from common
Variable
leak
Insulator
Ground
shield
Substrate
Cathodic arc
vapor source
Chamber
High current
feedthroughs
Gas
Plasma
Gas / Metal
–
+
e
Vacuum
Ions
Substrate
Shutter
High voltage
supply
–
e
Current
monitor
Pump
Ion
beam
+
I
holder
Evaporant
Plasma-based ion plating
Ion
Gas inlet
gun
E-beam
Vacuum-based ion plating
Figure 3.5 Plasma-based ion plating system equipped with a cathodic arc vaporization source
using bombardment from the plasma and a vacuum-based ion plating system using thermal
evaporation and an ion gun for bombardment.
Source: Mattox (2000) with permission from Elsevier.
A
96 Surface Modification of Magnesium and its Alloys for Biomedical Applications
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deposition processes such as thermal evaporation, sputter deposition, and arc vapor
deposition because the process name comes from the source of materials being deposited. Therefore, the ion plating process is also known by a number of other names, for
example, ion vapor deposition (IVD), ion assisted deposition (IAD), bias sputtering,
sputter ion plating (SIP), energy-assisted deposition, and ion beam assisted deposition
(IBAD) (Mattox, 2000).
3.2.4 Ion implantation
Ion implantation is a process in which ions of a material are accelerated by an electrical
field to impact a solid. If the ions differ in composition from the target, namely, the
Grid I
Grid II
rc power unit
Trigger power unit
Cathode
Trigger
R1
Accel. power unit
Decel.
power unit
Oscilloscope
Substrate
R2
Figure 3.6 Picture and schematic diagram of HEMII-80 high energy metal ion implanter.
Source: Feng et al. (2012) with permission from Elsevier.
Surface design of biodegradable magnesium alloys for biomedical applications 97
specimen to be implanted, they will alter the elemental composition of the target and
possibly change the physical, chemical, and/or electrical properties of the specimen. In
particular, the use of energetic ions affords the possibility of introducing a wide range
of atomic species independent of thermodynamic factors, thus making it possible to
obtain impurity concentrations and distributions of particular interest. Ion implantation
equipment consists of an ion source to produce ions of the desired element, an accelerator to accelerate the ions to a high energy, and a target chamber. Conventional
beam-line ion implantation is a line-of-sight process in which ions are extracted
from an ion source, accelerated to high energy, and then bombard the workpiece.
Figure 3.6 shows the picture of a conventional HEMII-80 high-energy metal ion
implantation machine in City University of Hong Kong and the simplified schematic
diagram. The ion beam is usually quite small, and thus either the ion beam or sample is
raster scanned to achieve uniform implantation over a large area. For samples with a
very complicated geometry, it may be difficult to obtain conformal ion implantation by
this method ( Chu, Qin, Chan, Cheung, & Larson, 1996; Liu et al., 2004).
In comparison to conventional ion implantation, plasma immersion ion implantation (PIII) is an advanced technique that can more conveniently process samples
with a complex shape. In PIII, the specimens are surrounded by a plasma and pulsebiased to a high negative potential relative to the chamber wall. Ions in the overlying
plasma are accelerated across the plasma sheat h generated around the specimens and
implanted into the surface conformally. Figure 3.7 displays the schematic illustration
of PIII. The plasma is produced in the vacuum chamber by various plasma sources
such as electron cyclotron resonance (ECR) or radio frequency (RF). If metal plasmas
are involved, plasma immersion ion implantation and deposition (PIII&D), which is a
hybrid process that involves ion implantation and deposition, can be conducted to form
an atomically intermixed layer between the substrate and coating in addition to ion
implantation and coating. It is an ion plating technique in the broad sense (Anders,
1997; Chu, Chen, Wang, & Huang, 2002; Liu, Chu, & Ding, 2010).
Chamber
Vacuumizing
t
Sample
T
Energetic
ions
–
High-voltage pulser
+
Figure 3.7 Schematic diagram of
plasma immersion ion
implantation (PIII).
98 Surface Modification of Magnesium and its Alloys for Biomedical Applications
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3.3 Surface design principles for biodegradable
magnesium alloys
3.3.1 Role of surface modification in traditional biometals
Metals play an essential role in biomedica l devices because they are suitable for loadbearing applications. Different metals, such as titanium alloys and stainless steels, have
been adopted in artificial hip joints, knee replacements, and fracture fixation devices
because they possess excellent bulk properties, such as high mechanical strength,
corrosion resistance, and superior fatigue properties (Cao, Liu, Meng, & Chu, 2011;
Wu, Liu, et al., 2008). The biocompatibility of materials is generally related to the
behavior of cells in contact, particularly cell adhesion onto the surface. Unfortunately,
most artificial materials integrate poorly with host tissues, resulting in postoperation
infection and other negative body responses. In this respect, surface modification plays
an important role in changing the response of the biological environment on the artificial medical devices. The proper techniques not only preserve the excellent bulk
attributes of the biometals, but also endow the metals with specific surface properties
required by different clinical applications (Liu et al., 2004). Therefore, surface design
and associated treatments are critical to biometal engineering.
3.3.2 Key issues in developing biodegradable magnesium alloys
Since biodegradable Mg-based implants dissolve as new tissues are formed, Mg alloys
form a dynamic interface in the physiological environment, and it is different from
those on traditional biomedical metals such as Ti alloys and stainless steels that do
not degrade and thus have a static surface, as illustrated in Figure 3.8 (Wu, Jamesh,
MgMg
H2
Magnesium-based material Hydrogen bubble
Physiological environment Surface film
Cell
Figure 3.8 Dynamic interface between the Mg-based materials and bioenvironment during
surface degradation.
Source: Wu, Jamesh, et al. (2013) with permission from Elsevier.
Surface design of biodegradable magnesium alloys for biomedical applications 99
Mechanical
performance
Mg
Corrosion
behavior
Figure 3.9 Critical factors in the design of Mg-based biomaterials.
Biological
behavior
et al., 2013). Mg corrodes in aqueous media via the following chemical reaction:
Mg þ 2H
O / Mg2þþ 2OHþ H2[ (Song, 2007; Zberg et al., 2009). Conse-
2
quently, hydrogen evolution, localized basification, and degradation occur on the
active interface to complicate the cells/biomaterials interactions. Biodegradable
Mg-based devices, which have the advantage of avoiding a second surgical operation
to remove the components thus minimizing patient trauma, are not only considered a
revolutionary concept in bioengineering but also of scientific interest. Unfortunately,
rapid degradation in the physiological environment has impeded thei r use as metallic
implants. On the one hand, improper degradation compromises the mechanical properties before the tissues have a chance to recover or heal completely. On the other
hand, the rapid degradation may lead to adverse biological response if Mg and other
elements are released into the surroundings too rapidly. Gas bubbles and subcutaneous
inflammation are possible causes of excessive Mg
2þ
or H2release in vivo, further
harming tissue growth on the surface and loosening the bonding between the tissues
and Mg-based implants (Kirkland, 2012). Owing to the dynamic interface, it is
more difficult to develop new-generation biodegradable Mg alloys compared to old
metallic biomaterials such as Ti alloys. In particular, the corrosion as well as biological
behavior and mechanical performance are critical concerns of Mg-based biomaterials
in the physi ological environment (Figure 3.9).
3.3.3 Aims of surface design for biodegradable magnesium
alloys
Biodegradation of Mg alloys is a dynamic process and often concomitant with continuously changing surface properties and loss of mechanical strength. In addition, the
human body is a very complex environment for biodegradation, thus making degradation unpredictable in the healing stage. The aim is thus to ensure safety when using
Mg-based implants in vivo. A proposed degradation mode is shown in Figure 3.10.
It is obviously easier if the Mg-based implants have an approximately intact initial contour in the serving period to retain the designed load-bearing capacity. With enough
100 Surface Modification of Magnesium and its Alloys for Biomedical Applications
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Mechanical strength
Healing stage
Biodegradation
Time
Figure 3.10 Schematic of desirable biodegradation and mechanical strength as a function
of time.
corrosion resistance, slow degradation can ensure the designed mechanical strength
and structure integrity in the healing stage. The dynamic surface becomes relatively
stable to foster cell/tissue growth without producing unacceptable hydrogen evolution
and other side effects. After the healing stage, faster degradation is allowable, and the
mechanical strength of the implant is permitted to decrease gradually because the surrounding tissues can bear the load gradually. Adverse effects on human physiology are
usually not expected, and thus it is still important to be concerned with degradation
from the perspective of the actual clinical requirements. This aspect mainly depends
on the alloy design and selection.
To achieve this designed degradation process, the surface design must aim to
construct a temporary surface on Mg alloys to render the proper mecha nical properties,
corrosion behavior, and biocompatibility. This temporary surface should effectively
impede corrosion in the healing stage, but gradually change its role as a corrosion barrier during healing. This temporary surface must satisfy some special requirements.
First, compared to bulk properties, the surface needs better resistance against corrosion
during the initial period after implantation and must be biocompatible. Second, the surface is biodegradable and expected to disappear eventually via biological or chemical
reactions. Third, the designed surface is further required to allow controlled corrosion
after the healing stage and must fully degrade afterward without introducing or leaving
harmful materials.
3.3.4 Strategies in surface design
In the surface design of biodegradable Mg alloys, it is important to realize that their
primary applications are in bone fixation devices and cardiovascular stents. It is also
essential to fathom the bulk properties of Mg alloys, including mechanical strength
and degradation behavior, before taking the next step to select the proper materials
Surface design of biodegradable magnesium alloys for biomedical applications 101
for the temporary surface as mentioned above. In this process, it is necessary to establish the surface composition, microstructure, and associated preparation methods to
achieve the desirable corrosion resistance, surface mechanical performance, and
biocompatibility according to clinical requirements.
Biocompatibility, which usually encompasses tissue biocompatibility and hemolytic biocompatibility, is the primary factor in designing new biomaterials, and different
applications require different properties. Tissue biocompatibility refers to cytotoxicity,
inflammatory reactions, and mutagenic or carcinogenic reactions, whereas hemolytic
biocompatibility is associated with hemolytic response and blood coagulation. Generally speaking, materials with good biocompatibility induce a minimal level of cytotoxicity, hemolysis, blood coagulation, and inflammatory reaction while imposing no risk
of inducing mutagenic or carci nogenic reactions (Wang, Cao, Qiu, & Bi, 2011).
Conventional corrosion-resistant biometals can be more easily rendered biocompatible
to meet clinical needs, but the issue is more complicated for biodegradable Mg alloys.
Some alloying elem ents such as zinc (Zn), calcium (Ca), and manganese can cause
deleterious effects if the release rate is too high (Kirkland, 2012). Therefore, a temporary surface is required to produce good biocompatibility in the initial stage and control
the degradation process to ensure acceptable release from the bulk materials. CalciumPhosphorus (Ca-P) coatings such as hydroxyapatite (HA) are good candidates to
obtain excellent biocompatibility in osteosynthesis.
In surface design, emphasizing only the properties of the surface is often not
enough, and the substrate must also be considered with the surface as one entity.
This concept is not only confined to biocompatibility and can be generalized to corrosion and wear resistance. Using anticorrosion coatings as an example, pinholes, pores,
and cracks are inevi tably present in the coatings and the surrounding electrolyte in the
bioenvironment can penetrate the coating via these cracks and pores. Since Mg is
chemically active and has a smaller standard potential than many permanent metals
such as Ti alloys, Al alloys, and stainless steels (Song & Atrens, 1999), the galvanic
effect must be considered in materials selection. Figure 3.11 depicts the schematic
diagram of the corros ion failure mechanism of the coated Mg-based materials. If the
coating is conductive, a galvanic cell is formed between the coating and substrate
Crack
Physiological environment
Galvanic cell
R
p(C)
R
Mg-C
Mg-based material
R
R
s
p(Mg)
Figure 3.11 Schematic diagram illustrating
the corrosion failure mechanism on coated
Mg-based materials.
Source: Wu, Jamesh, et al. (2013) with
permission from Elsevier.
Coating
Pore
102 Surface Modification of Magnesium and its Alloys for Biomedical Applications
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and the galvanic current is given by the following formula (Song, Johannesson,
Hapugoda, & StJohn, 2004; Wu, Jamesh, et al. (2013)):
EC E
I ¼
R
pðMgÞ
þ R
pðCÞ
Mg
þ Rsþ R
MgC
(3.1)
where E
R
p(Mg)
of the cathode; R
and EMgare the corrosion potentials of the cathode and anode, respectively;
C
is the polarization resistance of the anode; R
is the electrical resistance of the electrolyte; and R
s
is the polarization resistance
p(C)
MgC
is the
electrical resistance between the anode and cathode. When a Mg-based implant is in a
fixed bioenvironment, R
anode and cathode depends on E
galvanic current is generated by decreasing E
and Rsdo not change easily, and the current between the
p(Mg)
EMg, R
C
, and R
p(C)
EMgand increasing R
C
. Obviously, a small
MgC
MgC
and R
p(C)
and it is better to use insulating and chemically inert coatings such as microarc
oxidation (MAO) coatings.
For Mg alloys used in artificial joints, friction and wear raise some unwanted issues
such as possible inflammation caused by the debris. To improve the wear resistance, a
surface with a small friction coefficient and high hardness is called for. Although
harder coatings can improve the surface mechanical properties, the hardest materials
may not always be the best because Mg alloys are relatively soft. Only when sufficient
adhesion is ensured can hard coatings deliver the desirable performance. For example,
diamond-like carbon (DLC) with high hardness does not adhere to Mg well unless an
interlayer such as chromium is inserted ( Wu, Sun, Dai, Song, & Wang, 2010). It
should also be mentioned that the conductive interlayer may be dangerous when the
coating system is damaged if its electrode potential is higher than that of the Mg alloy.
It is necessary to select other appropriate interlayers or alternative methods. Ion implantation or laser treatment is a moderate way to modify the surface hardness.
Although its improvement is finite compared to that rendered by hard coatings, it
has other advantages. For example, the interfacial bonding strength is not an issue
because of the metallurgical bond.
There are many surface modification techniques that can be used when designing
Mg biomaterials, for example, microarc oxidation, physical vapor deposition, electrodeposition, ion implantation, etc. By using these techniques, it is possible to construct
an appropriate temporary surface on Mg-based biomaterials, and in the following
sections, some recent research examples pertaining to surface modification of biomedical Mg alloys will be discussed.
,
3.4 Research examples
3.4.1 Microarc oxidation coatings
MAO is a simple way to fabricate thick, porous, and hard ceramic coatings on metals.
The properties of MAO coatings mainly depend on the processing parameters,
substrate, and chemi cal composition of the electrolyte. In the case of Mg, these
Surface design of biodegradable magnesium alloys for biomedical applications 103
properties should be further fine-tuned to adapt to different clinical applications, and
some recent studies are presented here.
Gu et al. (2011) used MAO to prepare a coating on the Mg-Ca (1 wt%) alloy. The
microarc oxidation process was conducted in a 10 kW microarc oxidation setup
comprising a high-power plasma source with a frequency of 700 Hz and a duty cycle
of 30% as well as a stirring and cooling system. The aqueous electrolytes were
prepared from solutions of 10 g/L sodium silicate with 3.5 g/L sodium hydroxide.
MAO was conducted for 10 min at a fixed applied voltage in the range of
300e400 V and the effects of the applied voltages on the surface morphology, phase
constituents, hydrogen evolution, pH variation in the solution, and in vitro biocompatibility were studied. The thickness and pore size of the MAO coating increased with
increasing applied voltages as shown in Figure 3.12. Immersion in Hank’s solution
for 50 days revealed that the MAO coating produced at 360 V had the best
Figure 3.12 Surface and cross-sectional morphologies of MAO-treated Mg-Ca alloys at
different voltages: (a) 300 V, (b) 360 V, and (c) 400 V.
Source: Gu et al. (2011) with permission from Elsevier.
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long-term corrosion resistance. Adhesion, proliferation, and differentiation of
MG63 cells were promoted on the MAO coatings because of reduced Mg ion release
and pH value variation. Besides the voltage, the pulsing frequency was crucial to the
performance of the MAO coatings. Gu et al. (2012) deposited MAO coatings on AZ31
Mg alloys using 30 g/L Na
aqueous solution and a constant DC voltage of 325 V
3PO4
was applied for 5 min in the coating preparation. Four different pulse frequencies at a
constant pulse ratio of 0.3 (300 Hz, 500 Hz, 1000 Hz and 3000 Hz) were investigated,
and it was found that the MAO coating produced using 3000 Hz exhibited the best
corrosion resistance.
Fischerauer et al. (2013) studied the in vivo degradation behavior of MAO-modified
ZX50 Mg alloy. They implanted the samples into the femoral legs of 20 male SpragueeDawley rats and monitored them using microcomputed tomography over an
observation period of 24 weeks. The MAO-modified samples showed almost no corrosion in the first week, but after 3 weeks the degradation rate increased and it was even
higher than that of the untreated sample. Based on the fact that MAO implant s
degraded inhomogeneously via localized corrosion attacks, it was considered that
this increase was due to an increase in the surface-area-to-volume ratio of the MAO
implants. Histological analysis performed after 4, 12, and 24 weeks showed that the
initially improved corrosion resistance observed from the MAO implants had a positive effect on bone and tissue response. The reduced hydrogen evolution increased
osteoblast apposition from the very beginning, thus generating a stable
boneeimplant interface. Since MAO can delay initial degradation after implantation,
it improves the fracture stabilization, minimizes the burden on the postoperatively irritated surrounding tissues, and generates good boneeimplant bonding. Actually, if a
good boneeimplant interface is formed in the initial stage of bone healing, accelerated
degradation in the later stage will be of interest to osteosynthetic applications.
Wang et al. (2011) used MAO coatings to tailor the blood compatibility on Mg
alloys. They found that the MAO Mg-1.0 wt% Zn-1.0 wt% Ca alloy exhibited favorable blood compatibility. In their experiments, the MAO Mg alloy showed a decreased
hemolytic ratio (2.25%) compared to the untreated one (24.58%). The MAO Mg alloy
also showed significantly shorter prothrombin and thrombin time and significantly
longer activated partial thromboplastin time than the untreated Mg alloy. Arachidonic
acid- and adenosine diphosphate-induced platelet aggregation was significantly
reduced by the untreated Mg alloy extract, but it was less affected by the extract of
the MAO-treated Mg alloy.
Good antimicrobial properties are also very important to surgical operation due to
the risk of bacterial infection (Wu, Liu, et al., 2011), and an antibacterial surface to
inhibit bacterial colonization is of practical interest. Silver and silver-bearing compounds are well known to exhibit antimicrobial activity and have been incorporated
into the surfaces of a variety of medical devices. Ryu and Hong (2010) developed
an approach to fabricate MAO coatings on the AZ31 magnesium alloy using
AgNO
-containing electrolytes. The Ag-containing MAO coatings exhibited higher
3
corrosion resistance than the Ag-free MAO coatings and also excellent antibacterial
activity of over 99.9% against two strains of bacteria, Staphylococcus aureus and
Escherichia coli.
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