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DOI: (Published online March 14th, 2013) 10.1002/adma. 201300226
Zinc Exhibits Ideal Physiological Corrosion Behavior for Bioabsorbable Stents
By Patrick K. Bowen*, Jaroslaw Drelich*, and Jeremy Goldman
[*] Patrick K. Bowen, B.S., and Jaroslaw Drelich, Ph. D.
Department of Materials Science and Engineering
Michigan Technological University
1400 Townsend Drive
512 Minerals and Materials Building
Houghton, MI 49931 (USA)
E-mails: ; [email protected] [email protected]
Jeremy Goldman, Ph.D.
Department of Biomedical Engineering
Michigan Technological University
1400 Townsend Drive
319 Minerals and Materials Building
Houghton, MI 49931 (USA)
Keywords: bioabsorbable; stent; biomaterial; zinc; corrosion
Metallic stents are commonly used in percutaneous coronary interventions in patients with
pronounced ischemic or coronary heart disease to promote revascularization and retard
possible recoil and restenosis of damaged arteries. The benefits of stents during vascular
remodeling (i.e. improved procedural success, reduced restenosis, etc.) are numerous and well
documented.[1-3]
A traditional coronary stent must remain inert in the human body for many
years, but its residence is sometimes terminated after the occurrence of serious side effects.
Three of the most prominent problems include chronic inflammation, wherein a stent causes
an intermittent immune response;[4]
late stage thrombosis, wherein a clotting response occurs
long after initial stent placement;[5]
and stent strut disruption (fracture) that can result in
perforation or damage to components of the local vasculature.[6]
To mitigate the long-term side effects associated with traditional stents, a new generation of
so-called “bioabsorbable” metal stents is currently being developed.[7-9]
In much the same
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way special stitches may be absorbed after wound healing is underway, the new generation of
coronary stents will fulfill their role as vascular scaffolding[10]
and then proceed to be
absorbed by the body. Quite stringent corrosion and mechanical benchmarks apply to any
materials considered for use in an absorbable stent for reasons related to deployment, efficacy,
and device safety. Specific design considerations that have been suggested in the scientific
literature are compiled in Table 1. It is important to note that the research community has not
yet reached a consensus on some criteria presented in this table. For example, there are at
least two different ideal grain size ranges proposed in the literature,[11, 12]
and there is still
disagreement on the subject of aluminum as an allowable alloying addition.[13, 14]
Table 1. General design constraints and criteria for a bioabsorbable metal stent
Criterion Constraints Ref.
Bioabsorption
Mechanical integrity for 3-6 months
Mechanical integrity for 4 months
Mechanical integrity for 612 months
Full absorption in 1224 months
[11]
[15]
[12]
[11, 16]
Biocompatibility
Non-toxic, non-inflammatory, hypoallergenic
No harmful release or retention of particles
No aluminum or zirconium content
[11]
[11]
[14]
Mechanical
properties
Yield strength > 200 MPa
Tensile strength > 300 MPa
Elongation to failure > 1518%
Elastic recoil on expansion < 4%
[11]
Microstructure Maximum grain size of ~30 μm
Maximum grain size of 1012.5 μm
[11]
[12]
Hydrogen evolution Evolution < 10 μL H2 cm
-2 day
-1
[14]
Corrosion rate Penetration rate < 20 μm yr-1
[17]
Research spanning the last decade has focused on iron[18-20]
and magnesium[8, 21]
and their
alloys as bioabsorbable stent materials. Over the past two years, the authors have assessed the
in vivo corrosion behavior of iron and magnesium using a specialized, rat-based method of
evaluation.[13, 22]
These analyses, in combination with characterization of samples corroded in
vitro,[23]
have proven that iron is an unsuitable material for coronary stent application. The
corrosion product not only reduces the cross sectional area of the lumen and compromises the
integrity of the arterial wall, but it appears to be stable in the physiological environment
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resulting in long-term retention. Iron was observed to generate a layered corrosion product
that pushed away any biological matter to a distance of >750 μm from the original implant
surface after 9 months with little-to-no tissue integration.[22]
Conversely, magnesium appears
to degrade relatively harmlessly in the arterial environment. However, degradation times for
magnesium and its typical alloys are in the range of one to two months,[17]
far short of the
target time of three to six months. Extensive alloying efforts have not yet yielded a material
that meets the Table 1 criteria in terms of corrosion rate.[24]
Furthermore, common
commercial magnesium alloys often have elongations to failure in the range of 1% (i.e. K1-
and AM100-series alloys) to a maximum of about 1516% (AZ21- and AZ31-series
alloys).[25]
The most ductile and corrosion resistant alloys often contain aluminum or
zirconium—both of which have been cited as undesirable elements on a toxicological
basis[14]—or components that are acutely toxic, such as thorium. Summarily, a prohibitively
large amount of sophisticated and nuanced metallurgical manipulation is required to design
completely new magnesium alloys that conform to the Table 1 criteria.
Metallic zinc appears to be one of the few physiologically acceptable metals that have not yet
been considered for application in a bioabsorbable stent. Zinc is widely acknowledged as an
essential element for basic biological function, as it participates in nucleic acid metabolism,
signal transduction, apoptosis regulation, and gene expression in addition to interacting with a
variety of organic ligands.[26, 27]
The recommended daily value of zinc ranges from 10 mg
day-1
for adult males to 2 mg day-1
for infants.[28]
Aside from being physiologically essential,
a 1996 review by Hennig and coworkers indicated that zinc exhibited strong antiatherogenic
properties.[29]
The authors posited that this behavior arises from the role of ionic zinc as an
antioxidant and endothelial membrane stabilizer. By these two interactions, the integrity of
the endothelium is enhanced and the constituent cells are protected from lipid- or cytokine-
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induced perturbation. This feature adds significantly to the possible health benefits of a
bioabsorbable zinc stent.
Though concerns may be raised about the use of metallic zinc in the body, the toxic potential
of a zinc stent should be negligible. A cut, polished zinc stent may comprise, to a first
approximation, ~50 mg of the pure metal. Assuming complete degradation within one year,
the expected daily dose of zinc would be ca. 150 μg day-1
. This is still a fraction of the
recommended daily value for infants and a mere 4% of the recommended daily intake for
adult males.[28]
Hence, the systemic toxicity of a zinc stent should be nonexistent. Rapid
transport of ionic zinc in living tissue[30]
should prohibit zinc enrichment, cytotoxicity, or
necrosis in the vicinity of a stent, though cytotoxicity studies will be required in the future.
The selection of zinc for this study was largely based on the crucial physiological role played
by this metal, but was also inspired by the success of Mg-Zn-Ca bulk metallic glasses
pioneered by Zberg et al.[31]
Considerable resources have been dedicated to evaluating these
bulk metallic glasses as a breakthrough material.[32-34]
While these materials appear to be
efficacious in terms of corrosion rate and mechanical properties, major processing challenges
remain before they may be manufactured into a tubular stent preform. One of the primary
barriers is the requirement of mechanical homogeneity,[11]
implying the need for negligible
( 1%) porosity to prevent any microscale mechanical anomalies. This introduces difficulties
in processing glassy materials via powder metallurgy. Nevertheless, the apparent ideality of
biocorrosion in high-zinc metallic glasses is promising. These results prompted the authors to
question what would be possible if one approached this formulation from the other end
composition: from metallic zinc.
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Here, it is reported for the first time that metallic zinc shows great potential as a base material
for the next generation of bioabsorbable stents. Zn-Mg alloys were previously considered for
orthopedic application,[35]
though concerns remain regarding the systemic toxicity of a large
Zn implant. The current contribution addresses the suitability of zinc for cardiac stent
application based on its near-ideal biocorrosion behavior and other advantageous properties
(i.e. antiatherogenicity). Its favorable corrosion behavior is demonstrated by a series of four
medium-term wire samples, which spent either 1.5, 3, 4.5, or 6 months in the abdominal aorta
of a Sprague-Dawley rat.
After explantation, healthy arterial tissue clung firmly to the wires, with each successive
explant exhibiting more attached biological matter than the last. The tissue unfortunately
precluded surface characterization by optical microscopy, electron microscopy, or other,
similar techniques. Instead, the degradation behavior of the zinc wires was elucidated by
creating a series of cross sections from each sample. Representative backscattered electron
section images from a section of each explant are presented in Figure 1. Remaining metallic
zinc is visible as a bright feature in the center, surrounded (successively) by corrosion product,
tissue, and epoxy. The wires that remained in the biological milieu for 1.5 and 3 months
showed signs of relatively uniform corrosion. The remaining metallic cores of both explants
had ragged edges in cross section, which are posited to correspond to the removal of material
by semi-localized dissolution.
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Figure 1. Representative backscattered electron images from zinc explant cross sections after
1.5, 3, 4.5, and 6 months’ time in vivo.
After 4.5 and 6 months in vivo, the shallow, evenly distributed features yielded to relatively
severe, localized corrosion. Some locations appeared to be heavily attacked at 4.5 months, as
in Figure 1. Similar behavior was observed at 6 months. Instances of localized degradation
did not appear to be more severe at 6 months than at 4.5 months, but the cases of local
corrosion were more numerous. In all cases, tissue adhered preferentially to points of
localized attack. A limited amount of microscale corrosion particulate was firmly embedded
in the still-attached arterial tissue surrounding the metallic zinc core. The particles appeared
to have migrated no more than 250300 μm from their respective points of generation. The
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observed biocorrosion behavior of zinc is much preferred to that of pure iron reported
earlier.[22]
A quantitative evaluation of zinc biocorrosion was necessary to compare results against the 20
μm yr-1
design criterion (Table 1). Image analysis of all sample sections enabled
measurement of cross sectional area reduction and average penetration rates presented in
Figure 2 and Figure 3, respectively. The error bars in both figures correspond to the sample
standard deviation with four samples (n = 4) at 1.5 and 3 months and nine (n = 9) at 4.5 and 6
months. The area reduction and penetration rate corresponding to 20 μm yr-1
appears as a
dashed line in both plots for reference.
Figure 2. Measured values for cross sectional area reduction of zinc wires upon explantation.
The dashed red line corresponds to the projected area reduction at the target value of 20 μm
yr-1
.
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Figure 3. Average (lifetime) corrosion rates calculated from measured cross sectional areas.
The dashed red line shows the target value of 20 μm yr-1
.
The cross sectional area reduction in Figure 2 appears nonlinear, trending in a concave-up
direction, and the average rates in Figure 3 seem to rise gradually. Both observed trends
suggest a gradual acceleration of biocorrosion. To a first approximation, the corrosion rate is
zero at the time of surgery (p-value = 0.739), but increases linearly with implantation time (p-
value = 0.025, R2
= 0.95). A corrosion rate that trends in the first order corresponds to a
fourth order relationship between time and cross sectional area reduction. The trendline
presented in Figure 2 follows this correlation, and suggests that corrosion accelerates at a rate
of ~200 μm yr-2
. It should be noted that the acceleration value does not apply directly to the
average rates in Figure 3, but would rather apply to instantaneous penetration rate values that
are not discussed in this contribution. It is unclear at this time why the penetration rate
appears to increase over time, but the reason for this behavior merits further investigation.
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While proposed times of residence vary widely—from three months to one year, as in Table
1—previous clinical experiments with bioabsorbable magnesium stents have shown that
providing absorbable vascular support for about four months allowed substantial healing and
prevented vascular recoil.[15]
Figure 2 shows that pure zinc retained about 70% of its original
cross sectional area after four months in vivo, after which degradation was observed to
increase rapidly. Accelerating corrosion is advantageous in a bioabsorbable stent application;
zinc can remain mechanically stable for a four month period and disintegrate quickly
thereafter. The nearly ideal corrosion of zinc as a base material offers ample opportunity for
metallurgical and surface treatments to enhance bioactivity and/or biocompatibility. This is
superior to magnesium alloys; very few Mg alloys have been shown to possess biocorrosion
rates in the tens of micrometers per year,[24]
thus limiting the potential for creative biomedical
modification.
The products present on the corroded wires are a subject of interest in this exploratory work,
as identification can guide future efforts in alloying and/or surface treatment. Elemental
mapping with energy dispersive spectroscopy was performed to this end, and is presented in
Figure 4. At early times, thin layers of zinc oxide were the only product observed (data not
shown). As corrosion accelerated via local action at 4.5 and 6 months, the corrosion layer
thickened and the apparent compositional nature of the layer began to change. In the section
presented in Figure 4, four phases are observed. The lightest phase (possessing the highest
average atomic number) is the remaining metallic zinc. A calcium/phosphorus phase appears
on the elemental maps on the exterior surface near the top of the section, but it does not
appear to have formed a true bulk product. As a result, the calcium/phosphorus layer is not
thought to play a significant role in zinc biocorrosion.
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The remaining two phases include zinc oxide and zinc carbonate, both of which are present in
the compact corrosion layer. Local alkalization of the matrix during corrosion is suspected to
play a role in the development of the oxide layer.[36]
The precise stoichiometry of the zinc
oxide is unknown at this time, though this question warrants investigation; a zinc-deficient
oxide may have enhanced mass transport properties[37]
and partially explain the gradual
acceleration of corrosion. At this stage, the oxide appears in formations isolated from one
another in this cross section by an apparent zinc carbonate phase. The precise composition,
state of hydration, and mechanism of formation of the carbonate is unknown but merits
further study. Further work is also required to determine if the carbonate mineral is the final
product of biocorrosion. Other zinc minerals, such as hydrated phosphates,[38]
are more
thermodynamically stable than zinc carbonate. In fact, a hydrated zinc phosphate was
predicted by Zberg et al. to be thermodynamically preferred in an aqueous environment with
physiological composition.[31]
Therefore, it may be possible that the observed carbonate is
simply an intermediate phase that will ultimately react to form a different terminal product.
Figure 4. Schematic phase map (left), backscattered electron (BSE) image (center/top), and
individual elemental maps (right) of the 4.5 month section presented in Figure 1. The 50 μm
scale applies to the BSE image and all elemental maps.
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An important conclusion that can be drawn from compositional mapping is that zinc’s
biocorrosion behavior is distinct from that of magnesium and iron. A similarly degraded iron
wire was shown to exhibit sheet-like iron oxide surrounding a core of metallic iron [22]
. A
faint calcium/phosphate layer was seen surrounding the outer layer of iron oxide, and there
was little sign of tissue integration. Magnesium, on the other hand, had a dense, compact
corrosion layer that had integrated carbon, oxygen, phosphorus, and calcium from the arterial
milieu.[13, 39]
The magnesium also had highly visible deposits of a calcium/phosphorus phase
on the surface.[13]
The obvious similarity of zinc to iron is the rate of degradation previously
discussed: zinc corrosion is far slower than magnesium and thus provides a platform for
biologically beneficial modification. Unlike iron, however, zinc biocorrosion does not appear
to operate by a harmful mode of degradation. In fact, the corrosion layer observed on zinc is
mostly compact, like that of magnesium. Also like magnesium, there is no obvious tissue
necrosis or repulsion adjacent to the implant. Hence, zinc seems to corrode in a manner
distinct from both iron and magnesium, but simultaneously exhibiting the best aspects of the
other materials.
The results presented heretofore indicate that pure zinc is a viable bioabsorbable material.
However, the mechanical properties of zinc per se are not sufficient to allow application in a
bioabsorbable stent. While pure zinc exhibits an exceptional elongation to failure of 6080%,
it possesses a tensile strength of only ~120 MPa.[40]
Alloying to increase strength while
retaining ductility and corrosion resistance should be uncomplicated, however, as commercial
zinc alloys possess ~300 MPa tensile strength and simultaneously exhibit >20% elongation to
failure.[40]
The general classification of aluminum as an undesirable element makes this
alloying process more complex than it would be otherwise. Alloying of zinc will undoubtedly
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change its corrosion characteristics, and so experimental alloys should be produced with the
secondary objective of maintaining or increasing the uniformity of corrosion without
significantly shifting its rate from the native value. Electrochemically active additions or
elements promoting passivation may not produce the desired biodegradation in the final alloy.
The limited reports on zinc alloy corrosion—generally constrained to Zn-Al and Zn-Fe
alloys[41]
—will necessitate broad investigations on alloying element effects. Despite the
challenging and complex nature of this work, the significant ductility of pure zinc should
make this undertaking more practical than shifting magnesium ductility from its native 6% [40]
to a more acceptable value of ~15% (Table 1) while concurrently reducing the penetration
rate.
In conclusion, zinc has been examined for the first time as a bioabsorbable cardiac stent
material. Zinc, aside from being necessary in myriad biological processes, exhibits
antiatherogenic properties and possesses outstanding ductility. Early indications are that the
critical aspects of biocorrosion—the rate of penetration and the immediate effects of
generated products—satisfy the requirements for stent application (Table 1). The rate of
penetration has been shown to increase linearly with additional residence time in the murine
aorta. It is also apparent that pure zinc remains intact for four months or more in a small
animal model, after which time corrosion accelerates thus ensuring timely degradation of the
implant (Figure 2). The corrosion products on zinc after 4.5 and 6 months in vivo are largely
compact and comprise zinc oxide interspersed in zinc carbonate, though further investigation
is needed to identify the terminal corrosion product definitively. Zinc degradation was shown
to combine the desirable aspects of iron and its alloys, namely in vivo longevity, with the
harmless degradation of magnesium and its alloys. While it requires further development to
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achieve the desired mechanical properties, these results indicate zinc is a promising material
that could supplant magnesium as the favored base metal for bioabsorbable cardiac stents.
Experimental
Zinc wires of 99.99%+ purity (Goodfellow; Oakdale, Pennsylvania) were cut into 15 mm
segments and the ends smoothed to avoid any burrs that would hinder implantation. The
wires were cleaned in a ProCleaner™ device (BioForce Nanosciences; Ames, Iowa) for 30
minutes prior to the surgery. This device is equipped with a high-intensity mercury vapor
lamp generating UV illumination which attacks organic contamination at the molecular level
on small samples. Adult, male Sprague-Dawley rats were anesthetized with isoflurane in O2
and the 15 mm wires were placed in the abdominal aorta by puncturing the arterial adventitia
and leading the wire in. This procedure was reviewed and approved by Michigan
Technological University’s Internal Review Board. This procedure firmly embeds the wire
within the arterial media. After a period of 1.5, 3, 4.5, or 6 months, the rats were euthanized
in a manner approved by the Internal Review Board and in accordance with the Panel on
Euthanasia of the American Veterinary Medical Association. The wires were then harvested
for analysis and the state of the artery documented during necropsy. To preserve any
corrosion layer on the wires’ surfaces, the attached tissue on the 4.5 and 6 month explants was
dehydrated by immersing the samples for a short time in absolute ethanol. All samples were
stored in the low-humidity environment of a desiccator prior to analysis.
To elucidate the penetration rate of pure zinc, it was necessary to prepare multiple cross
sections of the corroded wires. The samples were held by a plastic sample clip, placed into a
12.5 mm diameter silicone tube, adjusted to ensure vertical alignment, and a two-part epoxy
resin was added into the tube. After curing, the mounts were cut transversely to expose the
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wire in cross section. To eliminate the small amount of deformation induced by cutting, the
wires were ground with 800 grit SiC, 1200 grit SiC, and polished with 1μm alumina.
Sonication in absolute ethanol ensured removal of all alumina from the surface. Another cut
was made below the polished surface to produce a section of approximately 1.01.5 mm
thickness, which was attached to an aluminum mount with carbon tape. The newly exposed
wire surface in epoxy was ground, polished, and cut to produce another section. This process
was repeated to produce the desired number of cross sections. The mounted wire sections
were coated with a thin layer of carbon to improve conductivity prior to imaging with the
scanning electron microscope.
A JEOL (Peabody, MA) JSM-6400 (research-grade, thermionic emission) scanning electron
microscope equipped with a dSpec automation system (Geller MicroÅnalytical Laboratory;
Topsfield, MA) was used for examining the sections. Imaging of the coated specimens was
conducted at 10 kV accelerating voltage at a reduced working distance using a backscattered
electron detector. The acquired backscattered electron images were analyzed with imageJ
(National Institute of Mental Health; Bethesda, Maryland) to yield cross sectional
measurements. The bright zinc portion of the image was selected by thresholding, in which
only the area containing the brightest pixels was measured. The original outline of the
implant wire was approximated by an ellipse. From these cross sectional area measurements,
a penetration rate was calculated for each image. The resulting measurements were averaged
to yield an estimated penetration rate, and their standard deviation was taken as the resulting
error.
Elemental maps were produced using the JSM-6400 with the attached energy dispersive x-ray
spectroscopy system (4pi Analysis; Hillsborough, NC). The maps presented in this
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contribution were acquired at a resolution of 512 pixels square with each pixel having a dwell
time of 15 ms. An accelerating voltage of 15 kV was used to improve the x-ray yield.
Acknowledgements
P. K. B. was supported by the DeVlieg Foundation and an American Heart Association
Fellowship during the time this work was conducted. Owen Mills is acknowledged for his
assistance in preparing the cross sectional specimens for analysis. The authors thank Emily
Shearier and Adam Drelich for their assistance in editing and proofreading.
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