Magnesium alloy properties are determined by the type and quantity of the alloy partners and impurity elements and also by the production conditions. Some effects of the alloy partners and impurity elements on the properties of the magnesium alloys are presented in C. KAMMER, Magnesium-Taschenbuch (Magnesium Handbook), p. 156-161, Aluminum Verlag Dusseldorf, 2000 first edition and are illustrate the complexity of determining the properties of binary or ternary magnesium alloys for use thereof as implant material.
The most frequently used alloy element for magnesium is aluminum, which leads to an increase in strength as a result of solid solution hardening and dispersion strengthening and fine grain formation, but also to microporosity. Furthermore, aluminum shifts the participation boundary of the iron in the melt to considerably low iron contents, at which the iron particles precipitate or form intermetallic particles with other elements.
Calcium has a pronounced grain refinement effect and impairs castability.
Undesired accompanying elements in magnesium alloys are iron, nickel, cobalt and copper, which, due to their electropositive nature, cause a considerable increase in the tendency for corrosion.
Manganese is found in all magnesium alloys and binds iron in the form of AIMnFe sediments, such that local element formation is reduced. On the other hand, manganese is unable to bind all iron, and therefore a residue of iron and a residue of manganese always remain in the melt.
Silicon reduces castability and viscosity and, with rising Si content, worsened corrosion behavior has to be anticipated. Iron, manganese and silicon have a very high tendency to form an intermetallic phase. This phase has a very high electrochemical potential and can therefore act as a cathode controlling the corrosion of the alloy matrix.
As a result of solid solution hardening, zinc leads to an improvement in the mechanical properties and to grain refinement, but also to microporosity with tendency for hot crack formation from a content of 1.5-2% by weight in binary Mg/Zn and ternary Mg/Al/Zn alloys.
Alloy additives formed from zirconium increase the tensile strength without lowering the extension and lead to grain refinement, but also to severe impairment of dynamic recrystallization, which manifests itself in an increase of the recrystallization temperature and therefore requires high energy expenditures. In addition, zirconium cannot be added to aluminous and silicious melts because the grain refinement effect is lost.
Rare earths, such as Lu, Er, Ho, Th, Sc and In, all demonstrate similar chemical behavior and, on the magnesium-rich side of the binary phase diagram, form eutectic systems with partial solubility, such that precipitation hardening is possible.
The addition of further alloy elements in conjunction with the impurities leads to the formation of different intermetallic phases in binary magnesium alloys (MARTIENSSSEN, WARLIMONT, Springer Handbook of Condensed Matter and Materials Data, S. 163, Springer Berlin Heidelberg New York, 2005). For example, the intermetallic phase Mg17Al12 forming at the grain boundaries is thus brittle and limits the ductility. Compared to the magnesium matrix, this intermetallic phase is more noble and can form local elements, whereby the corrosion behavior deteriorates (NISANCIOGLU, K, is et al, Corrosion mechanism of AZ91 magnesium alloy, Proc. Of 47th World Magnesium Association, London: Institute of Materials, 41-45).
Besides theses influencing factors, the properties of the magnesium alloys are, in addition, also significantly dependent on the metallurgical production conditions. Impurities when alloying together the alloy partners are inevitably introduced by the conventional casting method. The prior art (U.S. Pat. No. 5,055,254 A) therefore predefines tolerance limits for impurities in magnesium alloys, and specifies tolerance limits from 0.0015 to 0.0024% Fe, 0.0010% Ni, 0.0010 to 0.0024% Cu and no less than 0.15 to 0.5 Mn for example for a magnesium/aluminum/zinc alloy with approximately 8 to 9.5% Al and 0.45 to 0.9% Zn. Tolerance limits for impurities in magnesium and alloys thereof are specified in % by HILLIS, MERECER, MURRAY: “Compositional Requirements for Quality Performance with High Purity”, Proceedings 55th Meeting of the IMA, Coronado, S.74-81 and SONG, G., ATRENS, A.“Corrosion of non-Ferrous Alloys, III. Magnesium-Alloys, S. 131-171 in SCHUTZE M., “Corrosion and Degradation”, Wiley-VCH, Weinheim 2000 as well as production conditions as follows:
AlloyProductionStateFeFe/MnNiCupurenot specified0.0170.0050.01MgAZ 91pressure die castingF0.0320.0050.040high-pressure die casting0.0320.0050.040low-pressure die casting0.0320.0010.040T40.0350.0010.010T60.0460.0010.040gravity die castingF0.0320.0010.040AM60pressure die castingF0.0210.0030.010AM50pressure die castingF0.0150.0030.010AS41pressure die castingF0.0100.0040.020AE42pressure die castingF0.0200.0200.100
It has been found that these tolerance specifications are not sufficient to reliably rule out the formation of corrosion-promoting intermetallic phases, which exhibit a more noble electrochemical potential compared to the magnesium matrix.
The biologically degradable implants presuppose a load-bearing function and therefore strength in conjunction with a sufficient extension capability during its physiologically required support time. The known magnesium materials however fall far short of the strength properties provided by permanent implants, such as titanium, CoCr alloys and titanium alloys. The strength Rm for permanent implants is approximately 500 MPa to >1,000 MPa, whereas by contrast that of the magnesium materials was previously <275 MPa or in most cases <250 MPa.
A further disadvantage of many commercial magnesium materials lies in the fact that they is have only a small difference between the strength Rm and the proof stress Rp. In the case of plastically formable implants, for example cardiovascular stents, this means that, once the material starts to deform, no further resistance opposes the deformation and the regions already plastically deformed are deformed further without a rise in load. This can lead to overstretching of parts of the component and fracture may occur.
Many magnesium materials, such as the alloys in the AZ group, also demonstrate a considerably pronounced mechanical asymmetry, which manifests itself in contrast to the mechanical properties, in particular the proof stress Rp under tensile or compressive load. Asymmetries of this type are produced for example during forming processes, such as extrusion, rolling, or drawing, for production of suitable semifinished products. If the difference between the proof stress Rp under tensile load and the proof stress Rp under compressive load is too great, this may lead, in the case of a component that will be subsequently deformed multiaxially, such as a cardiovascular stent, to inhomogeneous deformation with the result of cracking and fracture.
Generally, due to the low number of crystallographic slip systems, magnesium alloys may also form textures during forming processes, such as extrusion, rolling or drawing, for the production of suitable semifinished products as a result of the orientation of the grains during the forming process. More specifically, the semifinished product has different properties in different spatial directions. For example, after the forming process, there is high deformability or elongation at failure in one spatial direction and reduced deformability or elongation at failure in another spatial direction. The formation of such textures is likewise to be avoided, since, in the case of a stent, high plastic deformation is impressed and a reduced elongation at failure increases the risk of implant failure. One method for largely avoiding such textures during forming is the setting of the finest possible grain before the forming process. At room temperature, magnesium materials have only a low deformation capacity characterized by slip in the base plane due to their hexagonal lattice structure. If the material additionally has a coarse microstructure, i.e., a coarse grain, what is known as twin formation will be forced in the event of further deformation, wherein shear strain takes place, which transfers a crystal region into a position axially symmetrical with respect to the starting position.
The twin grain boundaries thus produced constitute weak points in the material, at which, specifically in the event of plastic deformation, crack initiation starts and ultimately leads to destruction of the component.
If implant materials have a sufficiently fine grain, the risk of such an implant failure is then highly reduced. Implant materials should therefore have the finest possible grain so as to avoid an undesired shear strain of this type.
All available commercial magnesium materials for implants are subject to severe corrosive attack in physiological media. The prior art attempts to confine the tendency for corrosion by providing the implants with an anti-corrosion coating, for example formed from polymeric substances (EP 2 085 100 A2, EP 2 384 725 A1), an aqueous or alcoholic conversion solution (DE 10 2006 060 501 A1), or an oxide (DE 10 2010 027 532 A1, EP 0 295 397 A1).
The use of polymeric passivation layers is controversial, since practically all corresponding polymers sometimes also produce high levels of inflammation in the tissue. On the other hand, structures without protective measures of this type do not achieve the necessary support times. The corrosion at thin-walled traumatological implants often accompanies an excessively quick loss of strength, which is additionally encumbered by the formation of an excessively large amount of hydrogen per unit of time. This results in undesirable gas enclosures in the bones and tissue.
In the case of traumatological implants having relatively large cross sections, there is a need to selectively control the hydrogen problem and the corrosion rate of the implant over its structure.
Specifically in the case of biologically degradable implants, there is a desire for maximum body-compatibility of the elements, since, during degradation, all contained chemical elements are received by the body. Here, highly toxic elements, such as Be, Cd, Pb, Cr and the like, should be avoided in any case.
Degradable magnesium alloys are particularly suitable for producing implants that have been used in a wide range of embodiments in modern medical engineering. For example, implants are used to support vessels, hollow organs and vein systems (endovascular implants, for example stents), to fasten and temporarily fix tissue implants and tissue transplants, but also for orthopedic purposes, for example as pins, plates or screws. A particularly frequently used form of an implant is the stent.
In particular, the implantation of stents has become established as one of the most effective therapeutic measures in the treatment of vascular diseases. Stents are used to perform a supporting function in a patient's hollow organs. For this purpose, stents of conventional design have a filigree supporting structure formed from metal struts, which is initially provided in a compressed form for insertion into the body and is expanded at the site of application. One of the main fields of application of such stents is the permanent or temporary widening and maintained opening of vascular constrictions, in particular of constrictions (stenoses) of the coronary vessels. In addition, aneurysm stents are also known for example, which are used primarily to seal the aneurysm. The supporting function is provided in addition.
A stent has a main body formed from an implant material. An implant material is a non-living material, which is used for an application in the field of medicine and interacts with biological systems. Basic preconditions for the use of a material as implant material that comes into contact with the bodily environment when used as intended is its compatibility with the body (biocompatibility). Biocompatibility is understood to mean the ability of a material to induce a suitable tissue response in a specific application. This includes an adaptation of the chemical, physical, biological and morphological surface properties of an implant to the receiver tissue with the objective of a clinically desired interaction. The biocompatibility of the implant material is also dependent on the progression over time of the response of the biosystem into which the material has been implanted. Relatively short-term irritation and inflammation thus occur and may lead to tissue changes. Biological systems therefore respond differently according to the properties of the implant material. The implant materials can be divided into bioactive, bioinert and degradable/resorbable materials in accordance with the response of the biosystem.
Conventional implant materials include polymers, metal materials and ceramic materials (for example as a coating). Biocompatible metals and metal alloys for permanent implants include stainless steels for example (such as 316L), cobalt-based alloys (such as CoCrMo cast alloys, CoCrMo forged alloys, CoCrWNi forged alloys and CoCrNiMo forged alloys), pure titanium and titanium alloys (for example cp titanium, TiAl6V4 or TiAl6Nb7) and gold alloys. In the field of biocorrodible stents, the use of magnesium or pure iron as well as biocorrodible master alloys of the elements magnesium, iron, zinc, molybdenum and tungsten is recommended.
The use of biocorrodible magnesium alloys for temporary implants having filigree structures is in particular hindered by the fact that the implant degrades very rapidly in vivo. Various approaches are under discussion for reducing the corrosion rate, that is to say the degradation rate. Modified alloys and coatings represent categories of approaches to reduce the corrosion rate of magnesium alloys. Modified allows are produced to slow down the degradation on the part of the implant material as a result of suitable alloy development. Coatings are used to temporarily inhibit the degradation. Some approaches were very promising, but it has not yet been possible to produce a commercially obtainable product to the knowledge of the inventors. Rather, irrespective of the previous efforts, there is still an ongoing need for solution approaches that enable at least temporary reduction of the in vivo corrosion with simultaneous optimization of the mechanical properties of magnesium alloys.