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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: ; pkbowen@mtu.edu jwdrelic@mtu.edu

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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10

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.

Accepted to

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