Biomedical Testing of Magnesium-Based Biodegradable Stents

Introduction

The study of biodegradable materials has drawn a lot of attention both in research as well as in publications. These materials are likely to be used in for temporal function implants; for example, coronary stents. Biomaterials used for implantations include metals, polymers, as well as composites. Meta biomaterials have high ductility, impact strength, resistance to wear, as well as ability to absorb high energy of stress. This has increased the popularity of metal biomaterials. Metals were used first in the 18th century during which iron, silver, platinum, and gold were used to fix fractures in bones as wires and pins.

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A stent is a tiny mesh of tubular scaffold placed and later expanded in the coronary artery to keep the lumen unblocked (Moravej & Mantovani 2011, p. 2). Stents are used to provide an opening support that is mechanical and stops recoiling to the vessel. Since the remodeling of the wall of the artery is likely due to the mechanical stress that the stent generates, stenting is temporal; the affected artery with time regains equilibrium because of the stent’s stress, and thus the stent becomes unnecessary. Therefore, after the degradation of the stent, a healed vessel remains.

A biodegradable stent should compromise the integrity of its degradation as well as mechanical at the period of implantation. The remodeling process requires about 6 to 12 months before attaining completion, after which the stent’s mechanical integrity declines and degradation progresses tolerably without accumulating products of degradation on the site of implantation. Total degradation requires about 12-24 months after implantation (Hermawan, Dube, & Mantovani 2010, p. 1693). The stent is built up on the balloon, then expanded into position by inflation of the balloon. Stent expansion pushes it up alongside the arterial wall, but remains in place when the balloon deflates and keeps the artery open.

Biodegradable stents are those stents that have been made of degradable biomaterials. The first magnesium implantation in 1878 was used as a ligature to prevent bleeding of vessels among three patients (Moravej & Mantovani 2011, p. 1). Contemporarily, implants that are developed are aimed at ensuring strength, durability, and compatibility to the body. One of the most recent advances in the clinical field was dedicated to the use of magnesium-based stents from alloys of magnesium to manage two conditions of innate heart disease in children and critical limb ischemia (CLI) in adults (Zartner et al. 2005, p. 592). Besides, the results from a clinical test of magnesium stents to manage coronary artery diseases in adults were very encouraging (Erbel et al. 2007, p. 1869).

The popularity of magnesium-based biomaterials is due to their mechanical properties, and their compatibility; magnesium and their alloys are remarkably light-weighted with a density varying from 1.7 to 2.0 grams per cubic centimeter, which is close to bone density that is about 1.8 o 2.1 grams per cubic centimeter. Compared to other ceramic biomaterials, the fracture toughness associated to magnesium is greater and its modulus elasticity (41-45GPa) is close to the elastic modulus of the bone. This enables the shielding effect of the stress. In addition, magnesium is important in human metabolism, as well as the fourth most plentiful cation in the body with storage content of approximately 25 grams in the body, of which about half of it is stored in the bones. Magnesium acts as a co-factor for numerous enzymes, as well as the stabilized forms of DNA and RNA (Saris 2000, p. 12). The standard electrode potential of magnesium is about -2.37 V, and has a poorer resistance to corrosion in chlorine ions with physiologic environment (Gu & Zheng 2010, p. 111).

Most alloys of commercial magnesium comprise aluminum as well as rare earth. While aluminum is a neurotoxicant, severe hepatotoxicity is observable on administration of rare earth (Gu & Zheng 2010, p. 111). Consequently, next research emphasis is the study of the latest alloy of magnesium with nontoxic or low elements that are toxic. Nevertheless, there is deficiency of a uniform criterion for assessment of the biomedical properties of the alloys of magnesium. Generally, a good degradable biomaterial for stents should have adequate strength and a rate of degradation that is equivalent to that of the healing rate of the tissue. The biomaterial should as well possess fine biocompatibilities (Gu & Zheng 2010, p. 111).

Mechanical property of magnesium alloys

Magnesium indicates a very close elastic modulus to that of the cortical bone in comparison to that of the traditional Ti6A14V. The elasticity modulus Magnesium, as well as its commercial alloys, as mentioned above is about 45GPa, but in shear elasticity, modulus or rigidity modulus is approximately 16GPa. The elasticity modulus is indirectly proportional to temperature. Repeated working to stresses beyond the yielding strength of magnesium alloys both in tension as well as in compression can decrease their elasticity modulus (Gu & Zheng 2010, p. 111).

The magnesium alloys poses a big range of ultimate strength of tensile plus elongation; about 86.8 to 280 MPa and 3 percent to 21 percent correspondingly. Due to the quick loss of strength on early stages of degradation in vivo, the inherent strength of alloys of magnesium is not high enough even then. Nevertheless, the highest strength and stiffness is not necessary for internal fixation since its role is to provide only a temporal prop up and not a permanent replacement of the bone (Ruedi & Murphy 2001). A previous study by Gu and Zheng (2010, p. 113) indicated that the addition of aluminum, silicon, zinc, tin, and silver could greatly improve the strength as well as the elongation of magnesium. The strength of magnesium can further be attributed to the hot rolling and extruding, as well as ECAP, though they at times deteriorate ductility (Zartner et al. 2005, p. 592)

In regard to resistance of impact, Charpy and Izod impact standard tests performed on notched bars provide a way to gauge the comparative sensitivity notch of the alloys of magnesium subjected to impact loading. The same impact tests performed on bars that are not notched indicate much elevated values and may be tailored to designate resilience (ability of materials to absorb shock within the range of elasticity) or hardiness. Resilience increases with decrease in modulus elasticity. This means that magnesium-based materials have a high resilience owing to the fact that they have a relatively low elasticity modulus. Toughness or hardiness is dependent on the strength as well as ductility, and is measured by a calculation of the area under the curve (curve of strain against stress). Toughness may be estimated as a function of tensile strength and elongation (Zartner et al. 2005, p. 592)

Corrosion property of magnesium alloys

The rate at which biomaterials degrade is very crucial in performance. The rate of corrosion of magnesium is, thus, a matter to be considered when it comes to making stents. The rate of degradation affects both the period of healing of the tissue and affects the biomaterial mechanical properties loss during degradations. Zartner et al. (2005, p. 592) implanted a metal stent that is absorbable in a preterm kid with unintentional ligation of a pulmonary artery (left). Results indicated reperfusion of left lung, which continued through a 4-month follow-up. The gradual degradation of the magnesium was completed during this period of time (Gu & Zheng 2010, p. 112).

While magnesium alloys are generally thought to be rapidly corroding, the corrosion of magnesium alloys can be sustained within acceptable ranges depending on the temperature and particular design considerations. An evaluation of the AMS stent was performed, as well, in a study that was both multicenter as well as systematic; the study involved 63 participants in nine clinics with a single de novo native lesion of a the coronary artery. The outcome established that 71 stents, 10 to15 mm long, 3.0 to 3.5 mm diameter, and approximately 3 mg heavy, achieved an instantaneous angiographic effect equal to that of the effects of other stents of metal, which were safely degradable within four months (Gu & Zheng 2010, p. 112).

Biomaterials in orthopedics require approximately three to four months from the formation of the fracture callus to the formation of the new bone, and ultimately to the healing of the solid bone, which reinstates most of the original strength of the bone (Ruedi & Murphy 2001). Therefore, alloys of magnesium ought to retain their mechanical property beyond 3 months to evade the occurrence of a second occurrence that is bound to result from the quick magnesium implant degradation. The in vitro and in vivo rates of corrosion vary with physiological solutions in magnesium alloys. From the most recent researches, magnesium alloys show extremely rapid rates of corrosion and, therefore, more improvement in magnesium resistance to corrosion is indispensable (Ruedi & Murphy 2001).

Rational biocorrosion occurs in the body on biomaterials, for example implants and stents. The immune system bars colonization of bacterial as well as fungi in blood vessels and bones, thus coherent biocorrosion is due to oxidation reactions in the physiological media of electrolytes in the presence of bacteria. Corrosion rate ought to be cautiously managed to avoid absolute dissolution of implants before the patient heals. Severe corrosion is possible in aqueous physiological surroundings, particularly in the presence of chloride ions at levels of about 0.15mol/L. For example, Mg(OH)2 reacts with chloride ions forming a highly soluble chloride of magnesium and hydrogen (Ghali & Revie 2010, p. 423).

Magnesium alloys application as biodegradable implants in orthopedics is vitally dependent on the mechanical integrity of the stent during the period of service. The susceptibility of sand-cast magnesium-aluminum-zinc alloy was subjected to a stress corrosion cracking (SCC) examination in a modified-stimulated fluid of the body using a slow rate test (SSRT), which had the same ion concentration as blood plasma and excellent stability of storage. The results indicated 4.7% elongation towards fracture, and 120 MPa eventual strength of tensile in the air.

The mechanical property reduced by 17% and 21% when examined in m-SBF in comparison to that of the air. The formation of hydroxyapatite on magnesium alloy in this solution appeared to enhance biocompatibility, as well as corrosion resistance. The study indicated that the corrosion was not significant, and thus should not bar the use of magnesium for biodegradable implant (Ghali & Revie 2010, p. 424).

Biocompatibility

In vitro biocompatibility

The biocompatibility of magnesium-based materials have been tested in many cells, especially fibroblasts cell lines, primary cells, and bony or vascular-origin cell lines, depending on the desired application field. Favorable biocompatibility has been reported many times in magnesium and in different alloys of magnesium (Zhang et al. 2007). It was established that adhesion of the bone cell can be enhanced by magnesium on biomaterials through expression of integrin, as well as the pathway of MAP kinase. An assessment of biocompatibility of magnesium-based implants on cardiovascular system indicated that the endothelial cells of humans did not manifest toxic effects in high concentrations of the salts of magnesium, and the same kind of living cells observed near magnesium-calcium alloys also showed no colonization of the materials (Schumacher 2011, p. 24). An investigation on hemocompatibility of the alloys of magnesium manifested varied results depending on the alloy used (Zhang et al. 2007).

The extracts of magnesium-calcium for dendritic cells and L929 fibroblasts showed good biocompatibility (Zhang et al. 2007). The studies did not indicate any DNA induction damage or aberrations of chromosomes, nor mutations of genes when subjected to the cement extracts of phosphate bone. Pure extracts of magnesium were discovered to be slightly damaging to all types of cells used, except the L929 cells in which this impact was very pronounced. Pure magnesium, as well, established the worst results as regards to hemolysis and adhesion of platelets. These results were due to the high pH due to corrosion of pure magnesium rather than the content of Mg2+ ion (Zhang et al. 2007).

Almost all of the biocompatibility evaluations above were performed by use of MTT assay, which is tetrazolium salt-based, and it measures the viability as well as the proliferation through the cells’ metabolic activities, leading to tetrazolium conversion to formazan. The observable change in color is proportional to metabolically intact cell number (Schumacher 2011, p. 28).

In vivo biocompatibility

A study by Witte et al. (2005) investigated four alloys of magnesium (AZ31, AZ91, WE43 and LAE442) compared to degradable polylactide. Implants of rods of about 20 mm long and 1.5 mm diameter were put into guinea pig femura and then degradation process was evaluated. Degradation indicated great formation of the bone near the magnesium alloys and a corrosion layer with formless calcium phosphate. It was further observed that subcutaneous gas pockets were formed in the area of implantation, though no adverse clinical impacts were indicated and they did not reappear after removal via aspiration. It was established that LAE442 was the most gradually corroding, but later a report indicated more rapid degradation than AZ91D in vitro.

Besides, elevations of magnesium concentration in serum or evidences of impaired function of organs are yet to be performed, both histologically and blood chemistry alterations (Zhang et al. 2007). The inflammatory response towards degradation was either undetectable or was in form of an apt placid reaction of a foreign body. Recently, a study by El-Rahman (2003, p. 189) examined a surrounding tissue where they implanted MgCa0.8 screws in rabbits and observed a very minimal to moderate inflammation of the surrounding tissue. The study noted the fact inflammation increased in six to eight weeks, and as such the researchers suggested that longer-term investigations were necessary. Lastly, Witte et al. (2008) did not observe any potential of alloy skin sensitizing (AZ31, AZ91, WE43 and LAE442) in guinea pigs, leading to a conclusion that magnesium and its alloys have a generally good biocompatibility.

Toxicity

Whether or not the magnesium ion or other metal ions released by magnesium-based stents implanted in the body will cause cell damage is determined by the concentration of the ion that is released (Waksman 2009, p. 95). Nevertheless, it is not easy for one to approximate the concentration since the amount released differs with the shape; that is, size, as well as the surface area of the apparatus. Therefore, an attempt to estimate the amount of ion that is released by a stent whose normal size was taken to be 3 mm diameter, 20 mm long, and the ratio of the metal to that of the blood vessel was taken as 3:20. The thickness of the stent was estimated at 0.15 mm, and the width of the strut was 0.1 mm. Thus, a calculation of the total surface area resulted to 113 mm2. Taking the pure magnesium concentration to be 25.67 mg/L, the initial degradation rate would be 15.46 mg/L. But on the other hand, evaluation of the toxicity of a solution of one Molar magnesium chloride hexahydrate was done using the Japanese pharmacopoeia, whereby a murine fibroblast L929 was used. The results led to a conclusion that the concentration that led to cell growth inhibition by the magnesium ion was equal to 0.01 M (=243.1 mg/L) or thereabouts (Yamamoto et al. 1998, p. 332)

Thus, one can understand with precision from the experiment above that, for instance, a stent of pure magnesium with a 5 μm size that can give a significant degradation rate does not possess toxicity while being degraded in a live body. It is known that the concentration that lead to slowed proliferation of half murine fibroblast L929 gives 0.254 mM for YCl30.0132 Molar-

LiCl; 4.18 mM - Al(NO3)3; 0.254 mM - YCl3; and 0.145 mM - InCl4. Consequently, it was established that even in the alloys that have In, considering the shape of the apparatus as well as that of the alloy’s degradation rate, In makes it possible to use the alloys for medical applications as biodegradable alloys (Yamamoto et al. 1998, p. 334).

Waksman et al. (2006, p. 609) conducted an investigation of bioabsorbable WE43 magnesium safety in alloy stents in porcine artery of coronary in three months. The results indicated no embolization of stent particle or thrombosis, excessive inflammation, or deposition of fibrin. A less neointimal area was substantially observed in segments of magnesium stent in comparison to segments of steel stents. For that matter, it was concluded that stents of magnesium alloys are safe for use and have little formation of neointima. In larger lumens, there was no observable neointima.

Clinical Tests

In vivo

Peeters et al. (2005, p. 3) explains an initial follow-up of treatment of CLI with magnesium stents (AMS) in 20 people (adults) who manifested a high major clinical patency of about 89.5%, but without a major/minor amputation. This yielded a 100% salvage rate of the limb. Later, Bosiers et al. (2009, p. 424) performed a random study for 6 months, involving 60 patients with 74 lesions. The study established that the rate of angiographic patency for lesions managed with AMS was substantially lesser than in those treated with PTA. Notably, the stents degraded very fast, resulting to restenosis within the period of study, and these were associated with the early recoil as well as formation of neointima.

A multi-centre as well as an organized prospective study by Erbel et al. (2007 p. 1869) was performed whereby 71Mg alloy stents was implanted in 63 patients and the outcomes indicated that after a 4-month period, the rate of target lesion revascularization induced by ischemia was 23.8%. However, after one year, the overall target rate was 45%. On the other hand, angiography at the close of the four months indicated an increased stenosis diameter by approximately 17%. The results established that Mg stents are capable of attaining instant angiographic response that is the same as that of other metallic stents. Nevertheless, the limitations of the stent were where restenosis was higher compared to that reported in stents of bare metals. Thus, the study proposed a crucial modification to the stent characteristics to prolong time of degradation.

The most recent study by McMahon et al. (2007, p. 736) indicated early restenosis following the implantation of the stents of magnesium in aorto-pulmonary of a kid. Here, to begin with, there was a substantial vessel diameter increase, which later resulted to significant restenosis after four months after replacement of the stent. This further indicates that refinement of magnesium stents can be suitable for implants in children, which was not considered possible at first.

In vitro

The immersion (ASTM G31) test and the electrochemical (ASTM G59 and ASTM F 2129) test are among the common methods used for studies of magnesium implants, particularly to investigate their degradation. Moreover, a dynamic immersion test that simulated conditions in the coronary artery of humans was employed, and this showed the importance of the role of stress on degradation. In vitro tests are used to screen materials suitable for further studies (Gu et al. 2009, p. 491).

An immersion was done according to ASTM-G31-72 in SBF, as well as Hank’s solution. The experimental samples were immersed in 50 ml solutions and the temperature is maintained at 37°C in a water bath. The samples were removed after different immersion times and then the differences on the surface morphology, as well as the microstructure of the samples both before as well as after immersion were evaluated using environmental scanning electron microscopy. The microscope had an attachment of energy-disperse spectrometer (EDS), plus an X-ray diffractmeter. The results indicated that magnesium has an inherent corrosion property with the release of hydrogen, which may affect the cells in vivo (Gu et al. 2009, p. 491)

Fibroblast cells of murine, that is, L-929 and NIH3T3, the calvarial preosteoblasts of murine, that is, MC3T3-E1, endothelial cells of human umbilical vein (ECV304), as well as vascular smooth muscle cells (VSMC) of a rodent were used to examine cytotoxicity of pure magnesium and its alloys. The murine fibroblasts and preosteoblasts, VSMC plus ECV304 cells were grown in the Dulbecco’s modified Eagle’s medium (DMEM), the bovine serum of fetus (FBS) (10%), penicillin and streptomycin (100 mgml each) at a temperature kept at 37°C in an atmosphere that was humidified with 5% carbon dioxide. The cells were then observed by optical spectrometer. It was observed that the concentration of Mg++ ions produced was much lower than the IC50 of magnesium ions (Gu et al. 2009, p. 491).

Sarah, et al (1) performed a study involving pure Mg, Mg4Y, plus AZ31 stents of magnesium. A superficial midline incision was made on the trachea and then exposed. They used trachea of a rat as a bypass model for intraluminal stents. The researchers introduced two small defects and then a fresh donor trachea was prepared. Animals were then biweekly imaged by a μCT . Then an extra luminal stent was assessed in a canine model and the animals imaged weekly using X-ray. The results indicated that all animals survived beyond the predetermined endpoints. There was little degradation in MicroCT and X-Ray, but small fibrous capsules that were hollow were observed due to gas evolution from corrosion (Gu et al. 2009, p. 491).

Sterilization of Magnesium Alloy Stents

The magnesium alloy stents need to be sterilized before they can be used since they are implanted into the human tissue, and are likely to cause damage instead of healing if not sterilized. A number of the studies in this paper have indicated some toxicity, though slight, and corrosion associated with magnesium stents, among other reactions; thus sterilization to ensure purity is indispensable here. Thus, following stent fabrication, a stent normally undergoes sterilization to lessen the bioburden of that particular stent to a sterility assurance level (SAL) that is acceptable. Sterilization is normally measured via sterility assurance level (SAL) that indicates the likelihood of a viable microorganism that are present on a product (stent) after it has been sterilized (Hermawan 2012).

There exists many methods to sterilize the stents, and the commonest is treatment with ethylene oxide as well as ionization radiation treatment, like use of an electron beam or a gamma radiation. Other broad sterilization methods include heat, chemicals, and filtration. Autoclaving is an example of using heat to sterilize instruments. Numerous chemicals are used for sterilizing stents including the ozone, sodium hypochlorite, formaldehyde, ethylene oxide (mentioned above as the commonest), hydrogen peroxide, as well as per-acetic acid. It is important that the sterilization procedure pose little or no adverse impact on the stent’s material properties (Hermawan 2012).

Stents are normally terminally sterilized, which includes sterilization that is performed by fabrication, mounting on the device for delivery, as well as attachment to a delivery system. The stent is exposed to a chemical for sterilizing at ambient temperature. Process operations are then performed on the stent after which it is packaged, wherein the sterility assurance level (SAL) of the stent that is packaged must be less than a particular value (Hermawan 2012).

Conclusion

Apparently, biodegradable alloys of magnesium have attracted a lot of attention by both researchers and doctors owing to their suitability and advantages as presented in this paper. Nevertheless, their use is confronted with some challenges too; they have high rates of corrosion and degradation under physiological environment. To curb this challenge, procedures, including those mentioned in this article have been put forward to slow down these rates including modification with heat treatment with alkaline micro arc oxidation, treatment with phosphates, electro deposition, as well as coating with polymer. Other procedures include exploring the biomedical alloy of magnesium alloys using a new structure-like bulk metallic glass (BMG).

There has been a lot of research lately on biomaterials field, particularly biodegradable metal implants and specifically magnesium based alloys for stents. Their introduction has spurred a paradigm where metals must be corrosion-resistant. As discussed above, magnesium as a biodegradable biomaterial has quickly advanced into preclinical trials in humans. In common is the idea to recognize that less emotional but more scientific ways are needed and extensive studies on material and stent properties is indispensable before going into implantation research. This field, therefore, requires structured and long-term as well as a multidisciplinary investigations to ensure substantial improvement aimed at benefitting patients.

Reference List

Bosiers, M, Peeters, P, D’Archambeau, O, Hendriks, J, & Pilger, E et al. 2009, ‘Absorbable metal stent implantation for treatment of below-the knee critical limb ischemia’, Cardiovasc Intervent Radiol, vol. 32, no. 3, pp. 424–435.

El-Rahman, S 2003, ‘Neuropathology of aluminum toxicity in rats (glutamate and GABA impairment)’, Pharmacological Research, vol. 47, no. 3, pp. 189–194

Erbel, R, Di Mario, C, Bartunek, J, Bonnier, J & de Bruyne, B et al. 2007, ‘Temporary scaffolding of coronary arteries with bioabsorbable magnesium stents: a prospective, non-randomised multicentre trial’, Lancet, vol. 369, no. 9576, pp. 1869–1875

Ghali, E & Revie, WR 2010, Corrosion resistance of aluminum and magnesium alloys, John Wiley & Sons, Hoboken, NJ

Gu, X, Zheng, Y, Cheng, Y, Zhong, S & Xi, T 2009, In vitro corrosion and biocompatibility of binary magnesium alloys’, Biomaterials, vol. 30, no. 4, pp. 484-98

Gu, XN & Zheng, YF 2010, ‘A review on magnesium alloys as biodegradable materials’, Front Mater Science China, vol. 4, pp. 111–115.

Hermawan, D, Dubé, D. & Mantovani 2010, ‘Developments in metallic biodegradable stents’, Acta Biomaterialia, pp. 1693-1697

Hermawan, H 2012, ‘Biodegradable metals, state of art’, SpringerBriefs in Materials, DOI:

10.1007/978-3-642-31170-3_2

McMahon, CJ, Oslizlok, P, & Walsh KP 2007, Early restenosis following biodegradable stent implantation in an aortopulmonary collateral of a patient with pulmonary atresia and hypoplastic pulmonary arteries’, Catheter Cardiovascular Intervention, vol. 69, pp. 735–

738

Moravej, M & Mantovani, D 2011, ‘Biodegradable metals for cardiovascular stent application: Interests and new opportunities’, International Journal of Molecular Sciences, vol. 12, pp. 4250-4270; doi:10.3390/ijms12074250

Peeters, P, Bosiers, M, Verbist, J, Deloose, K, & Heublein, B 2005, Preliminary results after application of absorbable metal stents in patients with critical limb ischemia’, Journal of Endovascular Ther. Vol. 12, pp. 1–5

Ruedi, TP, & Murphy, WM 2001, Principle of fracture management, Thieme Medical Publishers, New York, NY

Saris, NEL 2000, ‘Magnesium: An update on physiological, clinical and analytical aspects’, Clinica Chimica Acta, vol. 294, no. 1–2, pp. 1–26

Schumacher, S, 2011, ‘In-vitro-studies of the tissue compatibility of magnesium-based implants’, University of Veterinary Medicine Hannover pp 1-98

Waksman, R 2009, ‘Current state of the absorbable metallic (magnesium) stent’, European Intervention Supply, vol. 5, pp. F94–F98

Waksman, R, Pakala, R, Kuchulakanti, PK, Baffour, R, & Hellinga, D et al. 2006, ‘Safety and efficacy of bioabsorbable magnesium alloy stents in porcine coronary arteries’, Catheter. Cardiovasc. Interv, vol. 68, pp. 607–617

Witte, F, Fischer, J, Nellesen, J, Crostack, H, Kaese, V, & Pischd, A, et al. 2006, ‘In vitro and in vivo corrosion measurements of magnesium alloys’, Biomaterials, vol. 27, pp. 1013–1028

Yamamoto, A, Honma, R, & Sumita, M 1998, ‘Cytotoxicity evaluation of 43 metal salts using murine fibroblasts and osteoblastic cells’, Mater Res, vol. 39, no. 2, pp. 331–340.