\act\Shape Memory Alloys

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Rev C Medical Handbook T. Duerig

\act\Shape Memory Alloys

\bla\Tom Duerig

This chapter is not intended as a treatise of Nitinol or the shape memory effect, but rather focuses on specific aspects of Nitinol that are of interest to medical-device designers, highlighting general principles, potential pitfalls, and key properties. For example, there is a great deal of emphasis on superelastic shape memory rather than thermal shape memory, and on fatigue and corrosion, often limiting factors in medical-device design. For those seeking a more complete discussion of Nitinol, Ref 1-4 are recommended.

\a\A Brief History

The discovery of metallic phase transformations that lead to unusual macroscopic shape changes can be traced to the Ölander’s study of Au-Cd in 1932 [Ref 5]; the discovery of the shape memory effect itself, however, is more fairly attributed to the 1951 work of Chang and Read [Ref 6] also studying the Au-Cd system, or to Reynolds and Bever’s studies of Cu-Zn in 1952 [Ref 7]. Rachinger [Ref 8] first observed superelasticity in 1958 and even coined the term still in use today. The observation of a shape memory effect in NiTi (Nitinol) is properly credited to Buehler, Gilfrich, and Wiley [Ref 9] in 1963.

Despite extensive research and hundreds of application patents, few commercial Nitinol products were launched during the twenty years following Nitinol’s discovery. A variety of factors were responsible for this, perhaps paramount of which were high costs and a lack of reliable material supply. There was also a consistent lack of understanding regarding what could be done with the materials versus what should be done—many ideas were technically workable, but commercially unviable. Nitinol quickly became a pariah, known as a “solution looking for a problem.”

The first Nitinol application (or any shape memory application, for that matter) to be sold rather than merely developed were fluid fitting couplings in the early 1970s [Ref 10]. The first application of Nitinol in medicine was as superelastic orthodontic archwires, extensively sold in the late 1970s. The first “for-sale,” agency-approved medical device was likely the Homer Mammalok, produced by Mitek Corporation [Ref 11], and the first commercially marketed permanent implant was likely Mitek’s bone anchor [Ref 12]. All these early medical devices employed the superelastic properties of Nitinol rather than the thermal shape memory properties. The 1990s brought about a demand to make smaller, less invasive medical devices, which in turn led to a rapid growth of the Nitinol industry as a whole. At the heart of this growth was the birth of the self-expanding stent, a device used to scaffold diseased arteries, first explored in 1982 [Ref 13]. Even though less than twenty percent of all stents are made from Nitinol, it likely remains the most commercially-important Nitinol application (as of 2010).

\a\Physical Metallurgy

The term Nitinol refers to the equiatomic, intermetallic nickel-titanium compound (Fig. 1). (This chapter disregards Nitinol-60, a highly nickel-rich composition with scattered applications as corrosion resistant tools, but not exhibiting useful shape memory properties.) Central to the properties of Nitinol is a reversible martensitic phase transformation: at high temperatures, the NiTi compound adopts a simple cubic structure structure (Fig. 2a) referred to as austenite. When the compound is cooled below its transformation temperature (approximately 80°C (176°F) in the fully annealed state), it transforms to a monoclinic structure referred to as martensite (see Fig. 2d).

As with nearly all phase transformations, the relative stabilities of austenite and martensite are influenced by both temperature and stress (leading to the thermal shape memory effect and superelasticity, respectively). Transformation between the two phases upon heating and cooling is defined by four characteristic temperatures, Ms, Mf, As, and Af (see top of Fig. 3). There are several ways to define “start” and “finish” (the subscripts “s” and “f” in Fig. 3) but unless there is a compelling reason to do otherwise, industry has settled on the intersection of tangents (demonstrated for the As temperature in Fig. 3). Because almost all physical properties change during the transformation, almost any property acts as a proxy for the volume fraction of martensite: electrical resistivity, density, etc. It is also common to determine transformation temperatures by Differential Scanning Calorimetry (DSC), which measures the heat content differences between the phases—it measures the rate at which the transformation is occurring (see bottom figure in Fig. 3). While this is particularly useful in characterizing ingots and fully annealed conditions, it loses resolution and engineering relevance when applied to product forms or materials with retained cold work. For these conditions, a Shape Recovery Test (SRT) is generally preferred, consisting simply of deforming martensite, and monitoring the temperature at which the original shape begins and finishes restoration.

As pointed out above, the relative stability of martensite and austenite is affected by stress as well as temperature, in this case, by the deviatoric stress rather than hydrostatic stress, or pressure; the application of a stress favors the martensitic phase due its ability to easily adapt its shape. One can, therefore, stress-induce martensite as well as thermally induce martensite. The relationship between temperature and stress in governing the transformation is given by the Clausius-Clapeyron equation:

\eq\ d/dT = H/tTo (Eq 1)

\btn\where is the applied stress, Hthe latent heat of transformation,t the transformation strain, and To is the transformation temperature. Values for d/dT, often called the stress rate, range from 4 to 15 MPa/°C, but for superelastic alloys commonly used in medical devices, that range is reliably 5-6 MPa/°C. As will be shown below, this turns out to be a very important consideration in product design. For example, tensile properties measured at room temperature rather than body temperature may be incorrect by 75-80 MPa, and a stent constrained in a catheter exposed to 60°C (140°F) will increase pressure against the polymer by as much as 200 MPa (29 ksi) when compared to room temperature. Yet another key descriptor of an alloy is the Md temperature, the temperature above which it is impossible to stress-induce martensite—here, the stress needed to stabilize the martensite phase exceeds the stress needed to plastically deform austenite.

As one can surmise from the phase diagram, transformation temperatures are unaffected by increases in titanium, but they can be depressed by adding extra nickel to the equiatomic composition. In fact, that relationship becomes extremely sensitive in compositions richer than about 50.5 at. % nickel (Fig. 4). It is important to note that this relationship is not an absolute; it is highly dependent on the carbon and oxygen content of the alloy, both of which form titanium-rich compounds, preferentially removing titanium from the NiTi compound matrix. Third element additions such as Cr, Fe, Al, and Co strongly suppress transformation temperatures, and additions of Ta, Hf, Pd, and Pt can increase transformation temperatures, but not without complex side effects. The majority of medical devices today are simple binary Ni-Ti alloys, with nickel contents ranging from 50.6 to 51.0 atomic percent, allowing users to easily “tune” the Af temperatures to 15-35°C (59-95°F).

Because the martensitic unit cell in Fig. 2 is of a different shape than the austenitic, nucleation of one phase within the other can only take place if either the nucleating phase or the surrounding phase is able to adapt its shape to accommodate the new structure. In the case of Nitinol, that is accomplished by twinning in the martensitic phase. While the details of martensite twinning are beyond the scope of this short review, Fig. 5 demonstrates the critical role of twinning in a simplified two-dimensional model. Once martensite is formed, the twinned structure has served its purpose and twins are free to move, grow, or shrink. Twin boundaries are highly mobile and of very low energy, and consequently, the mechanical properties of martensitic Nitinol are very unusual, feeling much like lead-tin solder rather than an intermetallic compound.

Armed now with the three states shown in Fig. 5, one can easily grasp the origin of both the shape memory and superelastic effects in Nitinol (Fig. 6). The thermal shape memory effect occurs when austenite is cooled to form twinned martensite, an applied stress deforms the martensite by moving twin boundaries, then subsequent heating reverts the martensite to austenite, “forgetting” all of the deformation attributable to moving martensite twin boundaries—due to the greater symmetry of austenite, all variants of martensite must revert to the single parent variant of austenite. Superelasticity is exactly the same, but in this case the martensite is stress-induced rather than thermally induced, thus the twinned martensite condition is transitory, occurring only at the austenite-martensite interface. Unloading, of course, reverts to the phase that is stable without load, austenite, again returning the original, undeformed shape. (To be precise, our definition of superelasticity is an inflection point upon unloading, thus one could thermally induce martensite and observe superelasticity, so long as unloading resulted in the reversion of martensite to austenite.) Of course there are limits to these two scenarios: first of all, the strain that can be harbored within the martensite phase is limited to the transformational strain of the austenite-martensite transformation (typically 7% in polycrystalline Nitinol), and secondly, both effects only occur within a limited temperature range (shape memory below the Af temperature, and superelasticity above Af and below Md.) Both of these points will be quantified below.

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