Advanced Biomaterials and Biodevices

By Ashutosh Tiwari

This cutting-edge book focuses on the emerging area of biomaterials and biodevices that incorporate therapeutic agents, molecular targeting, and diagnostic imaging capabilities

The design and development of biomaterials play a significant role in the diagnosis, treatment, and prevention of diseases. When used with highly selective and sensitive biomaterials, cutting-edge biodevices can allow the rapid and accurate diagnosis of disease, creating a platform for research and development, especially in the field of treatment for prognosis and detection of diseases in the early stage. This book emphasizes the emerging area of biomaterials and biodevices that incorporate therapeutic agents, molecular targeting, and diagnostic imaging capabilities.

The 15 comprehensive chapters written by leading experts cover such topics as:

• The use of severe plastic deformation technique to enhance the properties of nanostructured metals
• Descriptions of the different polymers for use in controlled drug release
• Chitin and chitosan as renewable healthcare biopolymers for biomedical applications
• Innovated devices such as “label-free biochips” and polymer MEMS
• Molecular imprinting and nanotechnology
• Prussian Blue biosensing applications
• The evaluation of different types of biosensors in terms of their cost effectiveness, selectivity, and sensitivity
• Stimuli-responsive polypeptide nanocarriers for malignancy therapeutics

• Language: English
• Publisher: Wiley
• Release date: Jun 30, 2014
• ISBN: 9781118774137

Book preview

Preface

Biomaterials are the most rapidly emerging field of biodevices. The design and development of biomaterials play a significant role in the diagnosis, treatment and prevention of diseases. Recently a variety of scaffolds/carriers have been evaluated for tissue regeneration, drug delivery, sensing and imaging. Liposomes and microspheres have been developed for sustained delivery and several anti-cancer drugs have been successfully formulated using biomaterials. Targeting of drugs to certain physiological sites has emerged as a promising tool for treatment, as it improves drug efficiency and requires reduced drug dosage. Using biodevices to target drugs may improve therapeutic success through limiting adverse drug effects, which results in better patient compliance and medication adherence. When used with highly selective and sensitive biomaterials, cutting-edge biodevices can allow the rapid and accurate diagnosis of diseases; creating a platform for research and development, especially in the field of treatment for prognosis and detection of diseases in the early stage. The emphasis of this book is the emerging area of biomaterials and biodevices that incorporate therapeutic agents, molecular targeting and diagnostic imaging capabilities.

The book is comprised of 15 chapters in total and has been divided into two major categories: Cutting-edge Biomaterials and Innovative Biodevices. The first section, Cutting-edge Biomaterials, focuses on state-of-the-art biomaterials such as nanostructures, smart polymers and nanoshells which can be used for medical applications. The first chapter, Frontiers for Bulk Nanostructured Metals in Biomedical Applications, illustrates the use of severe plastic deformation technique (SPD) to enhance the properties of nanostructured metals. This technique has been highly successful in augmenting the biomedical and mechanical properties of metals such as titanium, magnesium, cobalt and stainless steel. The second chapter, Stimuli-responsive Materials Used as Medical Devices in Loading and Releasing of Drugs, describes the potential of different polymers for use in controlled drug release. The main objective of using stimuli-responsive materials is to improve the performance of medical devices. However, the use of these materials is still in its infancy, as they are still prone to infections, inflammation and biofilm formation on their surface. Chapter three, Recent Advances with Liposomes as Drug Carriers, is a very interesting and comprehensive chapter which explains the use of artificially prepared bilayered phospholipid vesicles as a tool for drug delivery. Significant advancements in the last couple of decades have improved the efficiency of liposomes as a drug carrier and solved numerous problems related to their use. Among these are improvements in terms of the selectivity of drug carriers using engineered peptides, the use of dual-ligand combinations to reduce non-specific interactions with healthy tissues and also lowering ligand concentration using high-affinity ligands.

The chapter on Fabrication, Properties of Nanoshells with Controllable Surface Charge and Its Applications, describes the methods used to synthesize and assemble monodispersed core-shell nanoparticles. These methods are useful for improving adsorption of CNT for ultrasensitive detection using surface-enhanced Raman scattering. The chapter, Advanced Healthcare Materials: Chitosan, provides a review of chitin and chitosan as renewable healthcare biopolymers for biomedical applications such as wound healing or tissue regeneration, drug delivery and antimicrobial studies. The next chapter, Chitosan and Low Molecular Weight Chitosan: Biological and Biomedical Applications, also describes chitosan’s immunological and antioxidant properties, as well as its use for the treatment of tumors and viruses. The chapter, Anticipating Behavior of Advanced Materials in Healthcare, provides a general overview on the key aspects which need to be considered when developing advanced materials for healthcare applications.

Having advanced biomaterials is pointless if they cannot be used efficiently to reach targeted users. The reader is presented with a different point of view in the next section of the book, Innovative Biodevices, which explains how biodevices operate and how they can be used for biomedical applications. The first chapter in this section, Label-Free Biochips, illustrates a variety of miniature biodevices which can be used to measure different biomarkers for diseases. Unlike traditional optical imaging, the use of mini, dye-free sensors has the advantage of requiring less medical samples and providing noise-free measurement results. The next chapter, Polymer MEMS Sensors, illustrates another set of microelectromechanical systems (MEMS) sensors that are based on cantilevers. These miniature cantilevers can convert biological signals into different electrical signals (current, resistance and voltage).

The next chapters move away from describing devices to illustrating state-of-the-art techniques to improve them. Assembly of Polymers/Metal Nanoparticles and Their Applications as Medical Devices, demonstrates the use of polymer-coated metal nanoparticles in medical devices. Polymer-metal nanoparticles are favored due to their low toxicity and antibacterial and antiviral properties. The MEMS technologies often employ the top-down approach to build their devices. An emerging bottom-up technique uses nanostructures to form building blocks for the devices. The chapter, Combination of Molecular Imprinting and Nanotechnology: Beginning of a New Horizon, explains this new concept and its advantages such as enzyme-like and antibody-like properties, small physical size, solubility, flexibility and recognition site accessibility. The next chapter, Prussian Blue and Analogues: Biosensing Applications in Health Care, educates the readers on why Prussian blue, a transitional metal, has recently become very popular in biosensing applications. The chapter, Efficiency of Biosensors as New Generation of Analytical Approaches for the Biochemical Diagnostics of Diseases, evaluates different types of biosensors (electrochemical, optical) in terms of their cost effectiveness, selectivity and sensitivity. Nanoparticles: Scope in Drug Delivery, illustrates the use of nanoparticles (solid lipid, polymeric, liposomes, mesoporous silica) for drug-targeting to improve the efficiency of drug delivery in humans. Better drug efficacy is especially important in hazardous diseases such as cancer, which still uses toxic drugs for treatment. While having numerous advantages such as reduced dosage frequencies, versatile administration methods and better disease management, it is still too soon to know the long-term effects of these nanoparticles on humans and the environment. The final chapter, Smart Polypeptide Nanocarriers for Malignancy Therapeutics, reviews the recent advances in stimuli-responsive polypeptide nanocarriers for malignancy therapeutics.

Given the diversity of topics covered in this book, it can be read both by university students and researchers from various backgrounds such as chemistry, materials science, physics, pharmacy, medical science and biomedical engineering. The interdisciplinary nature of its chapters and simple tutorial nature make it suitable as a textbook for both undergraduate and graduate students, and as a reference book for researchers seeking an overview of state-of-the-art biomaterials and devices used in biomedical applications. We hope that the chapters of this book will give its readers’ valuable insight into alternative mechanisms in the field of advanced materials and innovative biodevices.

Editors

Ashutosh Tiwari, PhD, DSc

Anis Nurashikin Nordin, DSc.

Part 1

CUTTING-EDGE BIOMATERIALS

Chapter 1

Frontiers for Bulk Nanostructured Metals in Biomedical Applications

T.C. Lowe¹,* and R.Z. Valiev²,³

¹Colorado School of Mines, Golden, CO, USA

²Ufa State Aviation Technical University, Russia

³Laboratory for Mechanics of Bulk Nanomaterials, Saint Petersburg State University, Saint Petersburg, Russia

*Corresponding author: [e-mail]

Abstract

In recent decades, the nanostructuring of metals by severe plastic deformation (SPD), aimed at enhancing their properties, has become a promising area of modern materials science and engineering. With regard to medical applications, the creation of nanostructures in metals and alloys by SPD processing can improve both mechanical and biomedical properties. This chapter describes in detail the results of the investigations relating to titanium and its alloys, cobalt-based alloys, magnesium alloys, and stainless steels, which are the most extensively used to fabricate medical implants and other articles. The examples demonstrate that nanostructured metals with advanced properties pave the way to the development of a new generation of medical devices with improved design and functionality.

Keywords: Nanostructured metals, ultrafine grains, severe plastic deformation, mechanical and biomedical properties, orthopedic implants, biomaterial, biocompatibility, titanium, Co-Cr alloys, magnesium, stainless steel

1.1 Introduction to Nanostructured Metals

1.1.1 Importance of Nanostructured Biomedical Metals

The development of advanced materials for biomedical applications continues to enable superior solutions to improve human health. While new engineered materials impact most product sectors, their development for biomedical applications in particular has been rapidly expanding. This is partly a result of the convergence of nanoscale science and biological science over the past decade. Nanoscience, as applied to materials, addresses the same size scales of physical phenomena that are critical in living tissues. Consequently, Nanostructured Materials are now being engineered at a scale that matches the size range of attributes and physiological processes associated with human cells. New nanostructured soft and hard materials are being introduced every year. As of May 2013, 1,164 patents have been issued worldwide that reference nanomaterials.

Soft material structures, such as polymers and polymer-based composites, are the most prominent class of biomedical materials. This is partly because they are similar to soft tissues that predominate in human physiology. They are readily tailored to physiological applications since their nano/micro/macro-scale internal structures and surfaces can be functionalized for specific biomedical environments. They can be made biodurable for long-time use through surgical implantation, or biodegradable for temporary functions such as aiding drug delivery.

Aside from wood and other nature-made substances, metal is the oldest class of engineered biomaterial. Gold was used by the Greeks for fractures around 200 B.C. and iron and bronzes were used in sutures as early as the 17th century [1]. Silver, gold, and platinum were used as pins and wires for fractures in the 19th century. Steel was introduced for use in bone plates and screws at the beginning of the early 20th century, and in an ever growing number of orthopedic devices in the latter half of the 20th century [1]. The metals that are most prominently used in biomedical applications today are stainless steel, titanium, and cobalt-chromium (Co-Cr) alloys. Stainless steel, invented and produced first between 1908 and 1919, was used in bone plates by 1926. Co-Cr first appeared in bone plates 10 years later. Tantalum, a refractory metal, appeared in prostheses by 1939 and has since been used as radiographic markers, vascular clips, stents, and in repair of cranial defects [2]. Titanium and its alloys appeared in bone plates and hip joints by 1947. The well-known NiTi alloy Nitinol, discovered in 1958 found its way into orthodontic applications in the 1970s and cardiovascular stents in 1991 [1, 3].

Biomedical applications have traditionally required only small volumes of metal relative to the high tonnage production volumes that are most common in the metals manufacturing industry. Consequently, the alloys used in medical applications have typically been selected from those available for high volume non-medical applications, such as aerospace. However, during the past 20 years the attention to biomedical applications of metals has continued to grow, driven in part by increasing attention to quality of life, increasing longevity of populations worldwide, and the overall advancement of diagnostic and surgical procedures in medicine. Consequently, the demand for medical grades of alloys has grown as well. In addition, metal production techniques have evolved to support more economical production of small lot sizes. This has enabled the development of new alloys and surface modifications of existing alloys that are optimized for biomedicine.

This chapter addresses a new class of metals that have emerged over the past 20 years: bulk nanostructured metals [4–6]. Nanostructured metals are by definition metallic solids that have been deliberately engineered to have nanometer scale features (grains, precipitates, etc.) within the range between 1 nm to 100 nm that impart desirable physical, mechanical, electrical, and biological properties. We focus in particular on metals that can be produced in bulk forms such as rod, wire, sheet or plate. We will not address thin film technology (<100 nm thickness), compaction of nanosized powders that includes such techniques as hot isostatic pressing (HIP), and serial 3–dimensional fabrication methods such as Selective Laser Sintering (SLS), Laser-assisted Chemical Vapor Deposition (LCVD), and Laser-Based Additive Manufacturing (LBAM) [7, 8]. Instead, here we are interested in bulk nanostructured metals for which the mechanical and other properties can be customized, particularly for structural biomedical applications. Such metals can be produced by severe plastic deformation (SPD).

1.1.2 Brief Overview of the Evolution of Bulk Nanostructured Metals

Since the Second World War (1939–1945), researchers recognized that desirable characteristics such as improved strength and formability could be achieved in metals with fine grain sizes, in the range of 1 to 10 microns. The relationship between grain size and strength was published by E.O. Hall in a series of papers in the Proceedings of the Royal Physical Society in 1951 [9, 10]. In parallel, N.J. Petch from the University of Leeds independently published the results of his experimental work from 1946–1949 showing the relationship between fracture strength and ferritic grain size [11]. Also from 1945, researchers increasingly recognized the importance of fine grain size for enabling superplastic shaping and forming [12–17]. The earliest studies of SPD processing were enabled by the development of Bridgman’s anvils to impart very large shear strains via high pressure torsion [18]. From the mid-1970s researchers increasingly examined the behaviors of grain boundary structures in fine grain size metals in connection with superplastic deformation [19, 20]. This research provided the foundation for subsequent work on the processing to produce even smaller ultrafine grain sizes and focusing on the nanoscale structural characteristics of grain boundaries.

In 1981 Vladimir Segal patented and published an original method for imposing very large plastic deformations to bulk metals by simple shear [21]. The method entailed pushing a cylindrical or rectangular billet through a die built with an entrance channel and exit channel with essentially identical cross sectional dimensions, but differing in orientation by a fixed angle. Intense shear and accompanying rotations occur in the billet material as it passes through the channel intersection. Today this method, known by the label Equal Channel Angular Pressing (ECAP) or Equal Channel Angular Extrusion (ECAE) is one of the most popular techniques developed for imposing severe plastic deformation.

In the early 1990s, Valiev and co-workers made the first demonstrations of how severe plastic deformation leads to the continuous refinement of grain size and formation of ultrafine grain structure [22–24]. By 1999 there was enough interest in grain refinement by severe plastic deformation within the scientific community that Lowe and Valiev organized the first international workshop through NATO on the subject [25].

Of the various approaches for fabricating bulk nanostructured metals, methods involving severe plastic deformation have become among the most widely recognized. This is due in part to the fact that while SPD offers a cost effective means for grain refinement, it also enhances other properties of metals as well. For example, physical properties such as solid state diffusivity, radiation damage resistance, and acoustic dampening are enhanced [4, 26]. Within the context of biomaterials, one particularly distinctive property of SPD-processed metals stands out: living cells readily attach and proliferate on their nanostructured surfaces [27–37]. The rate of proliferation of osteoblast cells on nanostructured titanium has been reported to be as much as 19 times greater than on conventional titanium [38]. We will explore this and other distinctive properties of bulk nanostructured metals in the sections that follow.

1.1.3 Desirable Characteristics of Nanostructured Metals for Medical Applications

1.1.3.1 Good Manufacturability

The intrinsic advantage of producing bulk nanostructured metals by severe plastic deformation is that the process is predominantly mechanical and can be economically implemented in a manner that is fundamentally similar to extrusion or rolling. Thus severe plastic deformation can be conveniently inserted into conventional manufacturing production flow as one or several additional process steps. Over 50 continuous SPD processing methods that are suitable for manufacturing bulk nanostructured metals in the form of rod, bar, wire, plate or sheet have been published in the academic or patent literature. Of these, the methods that appear to be most promising for economical full scale commercial implementation include ECAP-Conform [39–43], variants of severe rolling [44–50], and multi-axis forging [51–60]. For medical applications, the use of SPD processing methods is favored by the fact that the cost of the constituent metal is typically a very small fraction of the total cost of most medical devices. For example, the value of commercial purity titanium that is typically used in a dental implant is on the order of $0.30 USD. Yet a single dental implant may be sold to oral surgeons for prices ranging from $50 to $400 USD.

Material reliability and consistency, more than material cost, are critically important in medical device applications. Reliability is one of the great intrinsic advantages of nanostructured metals produced by SPD: they typically possess relatively uniform microstructures and predictable grain size distributions. SPD processing incrementally imparts changes to microstructure in proportion with the increasing magnitude of imposed strain. Thus the degree of refinement can be carefully controlled. In contrast, grain refinement through recrystallization depends upon control of highly non-linear kinetics of nucleation and growth processes. Grain size refinement through severe plastic deformation processing is nearly time independent for simple alloy systems. Thus mechanical processing, rather than using elevated temperature thermal processing, reinforces reliability and product consistency. This is particularly important since as metals attain ever higher strength levels, their susceptibility to failure due to minor perturbations in microstructure can become unacceptable.

Metal processing and medical device production are governed through the implementation of specific quality standards. The ISO 13485 is the most widely applicable standard governing the quality requirements for design and manufacture of medical devices. It complements European medical device directives 93/42/EEC, 90/385/EEC, 98/79/EEC, and MDEG-2009-12-01 MSOG Class I Guidance for manufacturers of medical devices. These standards or guidance documents specify quality control procedures, including specification of raw materials and manufacturing methods. They apply to nanostructured metals that are used in medical devices. In principle bulk nanostructured metals can be substituted directly for conventional metals in existing medical device applications. However, there are subtle differences in their manufacturability.

For example, nanostructured metals are most commonly shaped by machining of bulk metal into their final form. The studies that have addressed machining of nanostructured metals generally show that nanostructured metals possess superior machinability. Documented advantages include reduced tool wear and superior surface finish [61–63]. Cutting forces for ultrafine grained copper and its conventional metal counterpart have been shown to be undifferentiated [62]. Lapovok et al. [61] noted that the thermal conductivity of nanostructured metals decreases with decreasing grain size, thereby reducing the length of machining chips during turning and enhancing machinability. However, it should be noted that the chips formed by machining can be significantly stronger and harder than derived from machining of conventional metals. Thus drilling or machining of internal cavities can be more difficult due to the need to remove the extra hard chips. In addition, the higher yield strength of nanostructured metals results in higher elastic stresses during boring or machining of interior cavities. This can accelerate tool wear under these circumstances.

There are distinct advantages to shaping and forming of bulk nanostructured metals. Their ultrafine grain size enables them to deform by grain boundary sliding, and therefore they can be formed superplastically at lower temperatures and higher strain rates than conventional fine grain metals [4, 26, 64–69]. The availability of grain boundary sliding as a deformation mechanism aids formability even during conventional intermediate temperature and high rate forging. One consequence of this advantage is that multi-step forging operations, for example, to produce high strength hip implant structures, can be accomplished in fewer steps. Similarly, in closed die forging it is easier to achieve complete die fill at lower temperatures and with lower forces.

Metals used in medical devices are commonly subject to surface modifications or coating processes during manufacturing. A significant body of knowledge has emerged on the viability of coating bulk nanostructured metals [70–76]. For example, hybrid oxide coatings or hydroxyapatite adhere readily to nanostructured titanium, providing enhanced biocompatibility and osseointegration [73, 77, 78]. While nanostructured surfaces have been shown to have intrinsically superior biocompatibility [28, 29, 36, 37, 79–82], the additional enhancement of their biological properties through surface treatments are notable [27, 83–91].

1.1.3.2 Superior Physical and Mechanical Properties

Perhaps the most distinctive characteristic of nanostructured metals is their superior mechanical properties compared to their conventional coarse-grained counterparts. The superior strength is particularly important for medical applications such as orthopedic devices. Drivers for additional strength in medical applications include the growing average weight and prevalence of obesity in adults [92], the need to achieve greater structural functionality using smaller volumes of metal, the need for greater fracture resistance in high load applications, and the need to place implants in confined spaces in the human body that cannot be adressed otherwise [93].

In general, one can increase the strength of virtually any metal or alloy by 20% to as much as a factor of four via SPD [94]. One can also improve ductility via SPD processing, with increments in the elongation to failure of up to 5 times reported [95–102]. Fracture toughness can also be increased in most alloy families [103–106]. Resistance to fracture under cyclic load may increase or decrease in SPD metals [107–116]. Generally, bulk nano-structured metals have superior fatigue properties. However, cyclic softening of SPD-induced microstructures subject to large inelastic cyclic strains can lead to diminished low cycle fatigue resistance [117]. Threshold stress levels for fatigue crack growth are generally higher in SPD metals, but can be lower in some cases [118]. This is due in part to the fact that SPD can cause the formation of textures that are deleterious to fatigue resistance [111, 119]. The localized shear that occurs during SPD can also alter second phase or precipitate morphologies so as to diminish the resistance to fatigue crack growth [120].

The corrosion resistance of most alloys is enhanced by nanostructuring [27, 71, 79, 121–137]. The improved corrosion performance of nanostructured metals has been attributed to their reduced grain size [138, 139], greater uniformity of microstructure [139, 140], and higher polarization resistance [141]. For alloys such as stainless steel that form tenacious oxides the polarization resistance and stability of the oxide layer increases with decreasing grain size [122, 142–144].

For most medical applications the range of temperatures experienced in service is comparable to the ambient temperatures experienced by humans. However, sterilization processes used to prepare metals for medical applications can expose metals to temperature of 100 °C. Low temperature storage of medical materials, as low as −80 °C must also be considered. Within the range of −80 °C to 100 °C, the microstructures present in bulk nanostructured metals are highly stable. These microstructures possess a diverse range of features, including for example high dislocation densities, high angle grain boundaries, dense dislocation walls, micro- and macro- shear bands, stacking faults, microtwins, and vacancy clusters [145–157]. Because these structures are mechanically induced, often at lower temperatures than would be possible via conventional deformation modes, they are commonly regarded as being metastable [158–160]. In addition, solid solutions can be supersaturated and second phases, especially intermetallics, can be amorphized by SPD. Thus, exposure of SPD-processed metals and alloys to elevated temperatures can enable transformation of mechanically induced structures to lower energy equilibrium states. Conversely, SPD can enhance thermal stability. For example, Efros et al. [161] showed that the formation of nanocrystals in an iron-manganese alloy retards the reverse martensitic transition and stabilizes the epsilon phase. Similarly, Srinivasarao, et al. [162] showed that high pressure torsion of body centered cubic magnesium-lithium alloys stimulates the precipitation of hexagonal close packed phases that remain stable under ambient conditions.

Creep resistance of metals and alloys is seldom a concern for medical metals since their long term exposure is only to the low temperature of the human physiology (37 °C). However, since creep resistance can be enhanced or diminished by SPD it is worthwhile considering the prospect of low temperature creep. In general SPD metals undergo creep via the same mechanisms that occur in conventional metals [163]. The refined grain structure of SPD metals and SPD-induced homogenization of second phase distributions can impart superior creep resistance [164, 165]. However, at very low loads grain boundary sliding [166–169] and Coble creep [163, 169–172] mechanisms may become active at lower temperatures than in conventional metals. At very high stresses Ti-6Al-4V can undergo creep relaxations, even at room temperature [173]. However, at the highest stresses the creep relaxation rates are lower by an order of magnitude for ultrafine grain Ti-6Al-4V compared to conventional Ti-6Al-4V. The stability and thermal tolerance depends upon multiple factors that must be evaluated for each alloy system under the loads and temperatures to which they would be subjected during use. The results of studies and developments done for different metals subjected to nanostructuring by SPD techniques are considered in section 1.2.

1.2 Nanostructured Metals as Biomaterials for Medical Applications

In general, nanostructured metals provide superior functionality compared to conventional metals. Their ongoing development is preparing them for more extensive use in diverse biomedical applications. Because of their distinctive interaction with cells, bulk nanostructured metals are increasingly becoming a significant class of biomaterials. Jonathan Black offers this definition of a biomaterial: a material intended to interface with biological systems to evaluate, treat, augment, or replace any tissue, organ, or function of the body [174]. Nanostructured metals are already finding use in dental implants and are being evaluated for other medical applications such as intramedullary nails, as described below. For example, the Nanoimplant® dental implant has been manufactured and marketed since 2006 by Timplant in the Czech Republic (see Section 1.2.1). This was the first medical device to be made from nanostructured titanium. One of the next nanostructured titanium products, also a dental implant, was manufactured and marketed by BASIC Dental Implant Systems under the trademark Biotanium in the USA beginning in 2011. The nanostructured titanium for both these products was fabricated by NanoMet LLC. (Ufa, Russia). We will examine prospective new applications in the sections that follow.

1.2.1 Nanostructured Titanium and its Alloys

Titanium and its alloys are widely used for medical implants in trauma surgery, orthopedic and oral medicine [3, 175, 176]. Successful incorporation of these materials in design, fabrication and application of medical devices requires that they meet several critical criteria. Of paramount importance is their biocompatibility, as determined in part by their relative degree of reactivity with human tissues. Another criterion for medical devices is their ability to provide sufficient mechanical strength, especially under cyclic loading conditions. Finally the material should be machinable and formable, thereby enabling device fabrication at an affordable cost. Recent studies have shown that nanostructuring of titanium and its alloys by severe plastic deformation (SPD) opens new avenues and concepts for medical implants, providing benefits in multiple areas of medical device technology. Results of processing these materials, including their properties relevant to advanced medical applications are presented below.

1.2.1.1 Commercially Pure Titanium

Numerous clinical studies of medical devices fabricated from commercial purity (CP) titanium for trauma, orthopedic and oral medicine have proven its excellent biocompatibility [3]. However, the mechanical strength of CP titanium is relatively low compared to other metals used in biomedical devices. Whereas the strength of this material can be increased by either alloying or secondary processing, for example by rolling or drawing, these enhancements normally come with some degradation in biometric response and fatigue behaviour. Recently it has been shown that nanostructuring CP titanium by SPD processing can provide a new and promising alternative method for improving the mechanical properties of this material [24, 177–180]. This approach also has the benefit of enhancing the biological response of the CP titanium surface [181].

Valiev et al. [182] summarized the results of the earliest development of long rods of nanostructured titanium (n-Ti) to provide superior mechanical and biomedical properties for making dental implants. The effort was conducted using commercially pure Grade 4 titanium [C − 0.052 %, O2 − 0.34 %, Fe − 0.3 %, N − 0.015 %, Ti-bal. (wt. pct.)]. Nanostructuring involved SPD processing by equal-channel angular pressing [26] followed by thermo-mechanical treatment (TMT) using forging and drawing to produce 7 mm diameter bars with a 3 m length. This processing resulted in a large reduction in grain size, from the 25 μm equiaxed grain structure of the initial titanium rods to 150 nm after combined SPD and TMT processing, as shown in Figure 1.1. The selected area electron diffraction pattern, Figure 1(c), further suggests that the ultrafine grains contained predominantly high-angle non-equilibrium grain boundaries with increased grain-to-grain internal stresses [4].

Figure 1.1 Microstructure of Grade 4 CP Ti: a) the initial coarse grained rod; b, c) after ECAP + TMT (Optical and electron photomicrographs).

A similar structure for CP Ti can also be produced using a continuous SPD method, ECAP-Conform, combined with further drawing into long rods [180]. It was essential to produce homogeneous ultrafine-grained structure along the entire three-meter rod lengths to enable economical pilot production of implants and provide sufficient material for thorough testing of the mechanical and bio-medical properties of the nanostructured titanium.

Table 1 illustrates mechanical property benefits attainable by nanostructuring of CP titanium. Note, for example, that the strength of the nanostructured titanium is nearly twice that of conventional CP titanium. This improvement has been achieved without the drastic ductility reductions (to below 10% elongation to failure) normally seen after rolling or drawing.

Laboratory fatigue studies of nanostructured and conventional CP titanium conducted in air at room temperature were performed per ASTM E 466–96 at a load ratio R (rmin/rmax) = 0.1 and loading frequency of 20 Hz. Table 1.1 also shows that the fatigue strength of nanostructured CP titanium after one million cycles is nearly two times higher than conventional CP titanium and exceeds that of the Ti-6Al-4V alloy [175, 176].

Table 1.1 Mechanical properties of conventionally processed and nanostructured CP Grade 4 titanium.

Cytocompatibility tests utilizing mouse fibroblast cells L929 were undertaken to verify the previously reported benefits of nanostructured CP titanium vis à vis conventional coarse-grained CP Ti. This study was performed, as described elsewhere [183], with hydrofluoric acid surface etching being performed prior to cell exposure. Figure 1.2 shows the etched conventional and nanostructured titanium surfaces, respectively. The differences in surface roughness of these materials are easily seen, a homogeneous and nanometer-sized roughness being apparent for nanostructured titanium compared with the much coarser structure for etched CP Grade 4 titanium.

Figure 1.2 Surface relief after hydrofluoric acid treatment of nanostructured (left) and CP Grade 4 titanium (right) surfaces.

The cell attachment investigation shows that fibroblast colonization of the CP Grade 4 titanium surface dramatically increases after nanostructuring, as shown in Figure 1.3. For example, the surface cell occupation for conventional CP Ti was 53.0% after 72 hrs in contrast to 87.2% for nanostructured CP Grade 4 (Table. 1.2). The latter observations also confirm the previous studies [181, 184, 185], showing that cell-adhesion on nanostructured titanium is greater than on conventional CP Grade 4 titanium. This result suggests that a high osseointegration rate should be expected with nanostructured CP Grade 4 titanium when compared to conventional titanium.

Figure 1.3 Occupation of the mice fibroblast cells L929 after 24 hours; Nanostructured (left) and conventional (right) CP Grade 4 titanium.

Table 1.2 Surface cell occupation for conventional and nanostructured CP Grade 4 titanium.

In work by Estrin, et al. [38] it is also shown that nanostructuring of titanium leads to increase of adhesion and proliferation of bone cells in comparison to conventional CG titanium. The authors explain this fact by surface topography change at a nano-scale. At the same time the influence of nanotitanium surface characteristics –relief topography, chemical state of the oxide film, electrical charge state – on osseoinduction, osseoconduction, and osseointegration requires more detailed investigation.

One objective of the effort in [182] was to design, fabricate and implant nanostructured CP Grade 4 titanium dental posts to clinically demonstrate the benefits associated with nanostructuring outlined previously. Toward this end, a reduced diameter implant post Nanoimplant® was designed and fabricated. This implant sustains the same load as a conventional 3.5 mm-diameter titanium implant, the former having the added capability of being used as a pillar in cases of insufficient thickness of the alveolar bone. The certified system of Timplant® manufactured to standard EN ISO 13485:2003 was used during development of the Nanoimplant® implant. The implants are shown in Figure 1.4, the nanoimplant intraosseal diameter 2.4 mm, having a strength equivalent to the conventional of 3.5 mm diameter implant.

Figure 1.4 3.5 mm diameter Timplant® (above) and 2.4 mm diameter Nanoimplant® (below).

To date over 250 Nanoimplants® have been implanted, most of them as immediate load implants, with all results indicating the excellent primary stability of Nanoimplants® when compared to other implant types [http://www.timplant.cz/e_stomatolog.asp]. For example, a 55-year-old male with edentulous mandible and maxilla was treated by insertion of conical implants laterally and Nanoimplants® in the narrow anterior part. Primary retention of all implants was very good; on the day of surgery the patient received a complete provisional bridge. Post-operation healing at the surgery site occurred without complications, with subsequent attachment of a definitive metalloceramic bridge completing the treatment.

The clinical tests, performed jointly with the R&D Institute of tooth transplantation Vitadent, Ufa, confirm also advantages of implantation of nanotitanium over conventional titanium. The dynamic densitometric investigations showed that in 3 months after implantation of nanotitanium the new generated bone tissue has the radiological density of 6.8 units, which exceeds bone density before implantation (4.7 units). For conventionally used implants this result, as a rule, is achieved only in 6 months after implant placement. This result substantiates a higher degree of nanotitanium osseointegration.

Thus, nanostructuring of titanium by SPD processing provides significantly superior mechanical performance when compared to conventional CP Grade 4 titanium. Furthermore, cytocompatibility studies with mouse fibroblast cells L929 have indicated that the nanostructured Ti surface has significantly higher cell colonization, suggesting more rapid osseointegration. Nanostructured (Nanoimplants®) implants have been successfully designed and fabricated. Clinical trials with over 250 patients, most of them receiving immediate load implants, have shown no adverse effects, with preliminary results being extremely encouraging [http://www.timplant.cz/e_stomatolog.asp].

Further clinical studies are presently underway with a population of 1000 patients.

1.2.1.2 Titanium Alloys

Titanium alloys are attractive materials for biomedical applications due to their light weight, high strength, relatively low Young’s modulus and good biocompatibility. Currently Ti-6Al-4V (Ti64) and Ti-6Al-7Nb are the most widely used commercial Ti alloys for dental and orthopedic applications [186–189]. The alloys consist of both hexagonal close-packed a and body-centered cubic β phases, with a Young’s modulus of ~ 110 GPa. Although Ti64 exhibits only half the Young’s modulus of either stainless steel or Co-Cr alloys, it is still about 4 times stiffer than cortical bone (20–30 GPa) [186, 190–192]. The difference in the modulus between artificial biomedical alloys and cortical bone creates a ‘stress shielding’ effect that undermines normal bone remodeling and maintenance and results in low bone density, loosening of implants, implant failure, and an increased likelihood for revision surgery [186, 193]. Furthermore, the passive film of Ti64 can slowly leach-out toxic vanadium ions [194], which have been linked to lower in-vitro cultured cell viability compared with pure titanium [195]. Therefore, recently research has started to develop a new generation of titanium alloys that have enhanced strength, lower Young’s modulus, and better biocompatibility that Ti64. The new approach based on nanostructuring of titanium alloys by SPD techniques plays a key role to provide these improvements.

During the last decade a number of studies have been performed aiming at increasing mechanical and functional properties of titanium alloys. Key results are described below.

In [196, 197] complex studies of the microstructure and mechanical properties of Ti-6Al-4V ELI (extra low interstitial alloys are used for medical applications) processed by SPD were conducted. The processing was performed using rods 40 mm in diameter made from the Ti-6Al-4V ELI alloy (Intrinsic Devices Company USA) of the following composition: Ti − base, Al − 6.0%; V − 4.2%; Fe − 0.2%; C − 0.001%; O2 − 0.11%; N2 −0.0025%; H2 − 0.002% (wt.%). The temperature of polymorphic transformation (TPT) in the alloy is 960 °C. The microstructure of the alloy in the as-received state was a granular type with a grain size of 8 μm in a cross-section, 20 μm in a longitudinal section. According to the X-ray analysis, the corresponding volume fractions of a and β phases was about 85% and 15%, respectively. The rods, 250 mm in length, were subjected to processing in 2 stages: ECAP in a die-set with channel intersection angle Θ = 120° at a temperature of 600 °C via route B and multicycle extrusion with total elongation ratio of 4.2 [198]. As a result of this processing, rods 18 mm in diameter and up to 300 mm in length were produced. The extrusion was carried out at 300 °C, and the last pass was conducted at room temperature in order to impart a high density of the crystal defects in the lattice structure. Subsequent annealing was implemented within the temperature range of 200 °C to 800 °C for 1 hour, followed by air cooling.

Microhardness of the samples was measured using a Buehler Omnimet machine with a load of P = 100 g for 10 seconds. The microstructures of these samples were studied by means of optical microscopy (OM) and TEM.

Figure 1.5 illustrates the dependence of microhardness of the alloy subjected to combined SPD-processing on annealing temperature. Changes of the microhardness are seen to be non-monotonic and its increase from 4200 up to 4780 MPa takes place when the annealing temperature is increased to 500 °C. Figure 1.5 displays the increase in microhardness with increasing annealing temperature up to 500 °C and decrease in microhardness at higher temperatures. The high microhardness observed at 500 °C in [199] after annealing of the UFG Ti-6Al-4V alloy is unusual. The authors associated these values with relative structure stability and changes of the proportions and structures of the α and β phases, with the β phase volume fraction being slightly increased. In this case strengthening of the UFG Ti-6Al-4V ELI alloy after annealing at 500 °C can be associated with the aging process as well. This process is known to be accompanied by decay of the metastable β-phase and precipitation of secondary particles of α-phase of various morphologies in conventional alloys [200]. Herewith, any deformation more or less increases the possibility of metastable phase breaking up due to high crystal lattice distortions. However, investigation of the nature of this aging in the UFG alloy requires more thorough study.

Figure 1.5 Temperature dependence of microhardness of the UFG Ti-6Al-4V ELI alloy samples after heating for 1h.

Figures 1.6a-c represent the alloys structure after ECAP and extrusion. SPD processing leads to considerable refinement and formation of a complex UFG structure with grains and subgrains having a mean size of about 300 nm. These grains have irregular form, a great number of various defects in the crystal lattice, and a high level of internal elastic stresses. The elongated form of structural elements created by extrusion straining is seen in the longitudinal section in Figure 1.6a and Figure 1.6c. Grain boundaries are not clear in these images due to large crystal lattice microdistortions as a result of severe plastic deformation.

Figure 1.6 Microstructure of the UFG Ti-6Al-4V ELI alloy before (a, b, c) and after annealing at 500 °C, 1 hour (d, e, f). Longitudinal section. a, d and c, f – bright-field and dark-field images respectively, b, e – diffraction patterns. TEM.

Figures 1.6d-f demonstrate that annealing at 500 °C of the alloy subjected to ECAP and extrusion leads to significant structural changes, particularly in the rod’s longitudinal section, which is characterized by formation of more equiaxed grains with a mean grain size of about 250 nm and thin boundaries (Figure 1.6d). The decrease in azimuthal spot spreading seen in SAED patterns (Figure 1.6e) shows the considerable decrease in internal elastic stresses as a result of processes of dislocation redistribution and recovery during annealing.

Figure 1.7 displays typical stress–strain curves for the CG and UFG alloy, which show the significant strengthening of the alloy after SPD processing due to microstructural refinement. Tensile elongation of the UFG alloy (curve 2) is reduced from 17% down to 9% compared to the as-received state (curve 1). Figure 1.7 demonstrates that subsequent annealing at the 500 °C increased strength and ductility (up to 12%), with uniform elongation of about 4%. The results of tensile tests are consistent with the data on microhardness measurement (Figure 1.5). Ductility enhancement in the UFG alloy after annealing is obviously conditioned by such factors as decrease in internal elastic stresses and dislocation density. As mentioned above, additional strengthening of the alloy can be associated with decay of metastable β-phase during cooling from the annealing temperature. Its volume fraction in the structure of the UFG alloy at 500 °C can be higher than before annealing, as has been shown in [199], using quenching from the annealing temperature. Though there are no visible particles of any second phase, aging processes could lead to the formation of grain boundary segregations that could additionally contribute to the enhancement of properties of the UFG alloy subjected to annealing [201, 202].

Figure 1.7 Elongation stress–strain curves of the Ti-6Al-4V ELI alloy: as-received state (1); UFG state before (2) and after annealing at 500 °C (3).

Investigations of fatigue properties of the UFG Ti-6Al-4V ELI alloy revealed that high strength and enhanced ductility after SPD processing and additional annealing at 500 °C (1370 MPa and 12%) resulted in fatigue limit enhancement on the basis of 10⁷ cycles up to 740 MPa in comparison with 600 MPa in the initial coarse-grained (CG) state (Figure 1.8).

Figure 1.8 Fatigue test results of the smooth samples out of CG and UFG alloy after annealing at 500 °C, 1 hour.

The fatigue limit for the UFG Ti-6Al-4V alloy, observed in [196] under the conditions of rotating bending, slightly exceeded the value reported previously, [197, 203], which testifies to the fact that the level of fatigue properties depends on the measurement technique.

Thus, the results show that high strength can be achieved in UFG Ti-6Al-4V ELI alloy through ECAP and additional mechanical and thermal treatment. Herewith, varying the SPD parameters, in particular temperature, strain rate, strain, provides the opportunity to control grain boundary structure in UFG materials, and consequently produce the best combinations of strength and ductility, as well as increasing the fatigue endurance limit of up to 740 MPa, well beyond the 600 MPa level measured in the coarse-grained alloy.

Another recent work [204] is devoted to enhancement of strength and ductility of the Ti-6Al-7Nb alloy. This alloy is regarded as being less harmful to humans from the medical point of view in comparison to Ti-6Al-4V. It has been demonstrated that formation of an ultrafine-grained structure in this alloy via equal channel angular pressing in combination with heat and deformation treatments results in high strength (UTS=1400 MPa) and ductility (elongation 10%). These levels of properties are very attractive for new applications and fabrication of medical implants with enhanced service properties.

Titanium alloys consisting of mainly the β phase have drawn substantial attention because they exhibit Young’s moduli ranging between 55 GPa and 90 GPa, and thus result in less stress shielding [195, 205–208]. In addition, these Ti alloys contain only non-toxic elements such as Nb, Zr, and Ta. Unlike Ti64 where V can leach out from the surface passive oxide film into the human body, the addition of Nb stabilizes the film, thus improving the passivation and corrosion resistance of titanium alloys in the body. However, high hardness and low Young’s modulus are desirable but hardly coexist in this group of materials. This is because the single phase β-Ti alloys, which exhibit the lowest Young’s modulus, are generally obtained after solution treatment, and so are relatively soft. Substantial strengthening can be achieved by ageing treatments that induce a fine and uniform precipitation of ω and α phase components, but this inevitably increases the Young’s modulus of the alloy [195, 205, 209, 210]. Consequently, there is a critical need to devise strategies to produce β-Ti alloys with low Young’s modulus, and high strength, making them more suitable for use in dental and orthopedic applications.

The results of the recent study suggest that it is possible to design nanocrystalline β-Ti alloys that meet the simultaneous requirements of high strength, low modulus of elasticity and excellent bio compatibility. Notably, all of the alloying elements (Nb, Ta, Zr, and O) in the β-Ti alloy are nontoxic and non-allergenic [195]. The nano-grain nature of the material leads to improved bulk mechanical properties, plus the nanotopography on the surface contributes to improved biological responses. Higher strength is evident by the superior hardness which arises from grain refinement [211]. Lower rigidity was achieved, which is attributed to the nanocrystalline structure and the complete elimination of the ω phase. In addition to these desirable mechanical properties, the nanocrystalline β-Ti alloy also displays excellent in vitro biocompatibility, indicated by enhanced cell attachment and proliferation. This novel nanocrystalline β-Ti alloy has a significant potential as a new generation of implant material with significant promise in load bearing biomedical applications.

1.2.2 Stainless Steels

Stainless steels are the most widely used family of alloys for medical applications. They contain 17–21% chromium which imparts good corrosion resistance due to the adherent chromium oxide film that forms and heals in the presence of oxygen. ASTM standards F138, F139, F1314, F1586, and F2229 define the chemical, mechanical, and microstructural requirements for various types of stainless steels that are used for medical applications. Austenitic stainless steels (American Iron and Steel Institute 300 series) such as 304L or 316L are used in mainly in temporary implant applications such as bone screws, bone plates, and intramedullary nails. Despite their passive oxide layer, they nevertheless corrode, releasing chromium and nickel into the body. Only minute quantities of these metals can be tolerated in the blood. Some stainless steels contain high amounts of nitrogen (per ASTM F2229) so that the nickel content can be reduced to less than 0.05 wt.%, minimizing the risks associated with nickel allergy reactions. These steels fall in the AISI 200 series, and are used for bone screws, plates, and fracture fixation. Martensitic stainless steels from the AISI 400 series are exceedingly strong and hard. However, they are ferromagnetic and are generally used outside the body, for example, for surgical instruments.

The austenitic stainless steels (AISI 300 series) are single phase and are often strengthened through cold or hot working. Austenitic stainless steel alloys are also readily strengthened by severe plastic deformation. For example, Idell et al. used SPD to increase the strength of 316L stainless steel from 515 MPa to 1647 MPa [212]. Similarly, Chen et al. obtained a yield strength of 1460 MPa in a duplex 32304 stainless steel after just 4 ECAP passes [213]. High nitrogen variants of stainless steel (ASTM F1586 and F2229) may also be cold worked to have ultimate tensile strengths above 1400 MPa.

While stainless steels are highly responsive to SPD, the adoption of SPD-processed stainless steels for commercial applications has been slow, in part because SPD introduces complex changes in microstructure that need to be more thoroughly understood. SPD alters the phase compositions from those normally expected in stainless steel. This occurs through mechanisms including stress-induced and strain-induced martensite formation. For example strain-induced martensite was nucleated during ECAP of 301 stainless [214] and 304 stainless steels [215]. SPD introduces microstructural features such as nano-twins, micro-twins, micro-shear bands, very high dislocation densities, and diffuse subboundary structures. SPD also alters the formation and distribution of carbides [216]. These microstructural effects in combination can significantly alter the annealing and recrystallization behaviors of stainless steels [217, 218].

As found for titanium, the ultrafine grained and nanostructured surfaces of stainless steel provide enhanced corrosion resistance [122, 142] and cell growth and proliferation [219, 220]. These results are encouraging indicators of the potential for nanostructured stainless steels for biomedical applications.

1.2.3 Cobalt-Chromium Alloys

Cobalt-Chromium alloys are sought for medical applications because of their combination of corrosion resistance, wear resistance, and high strength. Increasing amounts of chromium added in solid solution to cobalt, up to 35 weight percent, enhances corrosion resistance through the presence of a passive chromium oxide film. The general corrosion resistance of a typical Co-Cr alloy is an order of magnitude greater than stainless steels [1]. With Vickers hardness as high as 450 Hv, Co-Cr alloys are particularly suited for applications where sliding friction demands high wear resistance. Wrought cobalt-chromium-nickel alloys can have ultimate tensile strengths over 1800 MPa when cold worked and aged [174, 221]. Molybdenum is added to Co-Cr alloys to achieve finer grain sizes and further increase strength. For instance, MP35N contains nominally 10 weight percent Mo and can have an ultimate tensile strength of 2025 MPa when cold worked 53% and then aged at 565 °C for 4 hours [221].

Co-Cr superalloys generally fall into two categories: castable alloys containing some Mo and wrought alloys containing additions of Ni, plus Mo, W or Ta. The castable Co-Cr alloys were the first used in medical applications, particularly for dental applications [222–224]. Their minimum properties are specified in the ISO 5832-4 and ASTM F75 quality standards. Wrought Co-Cr based superalloys generally contain 3 to 35 weight percent nickel. The ASTM F90, F562, F563, F1091–12 and ISO 5832–6 standards specify the minimum properties for these alloys. Wrought Co-Cr alloys are used in knee, hip, and shoulder orthopedic prostheses, fracture fixation, and surgical wire [224–226].

Co-Cr alloys are strengthened by the presence of multiple phases, solid solution hardening, precipitates, intermetallic dispersoids, and inclusions. Carbides are an important constituent of superalloys, and are commonly present in grain boundaries, though in lower concentration in wrought alloys. Their presence helps control grain size. They also provide some degree of matrix strengthening. Chromium enables the formation of chromium carbides. Additions of tungsten, for example as found in L-605 (15 wt.% W), Stellite 6B (4.5 wt.% W), or Haynes 188 (14 wt.% W) allow the formation of tungsten carbides. Consequently, they are notoriously wear resistant, hard, and difficult to machine or forge. Forging of Co-Cr alloys is commonly conducted above 1000 C with only small reductions initially, just sufficient to cause recrystallization and grain size refinement upon reheating [227]. Such constraints on deformation processing make Co-Cr superalloy less amenable to nanostructuring by severe plastic deformation. There are only a few instances of SPD of superalloys reported in the academic literature [228–235]. Surface-confined SPD techniques such as Friction Stir Processing (FSP) [236], shot peening [237], and machining [235, 238, 239] are the primary SPD approaches that have been investigated. Conventional powder metallurgical techniques, including Hot Isostatic Pressing, have also been used to prepare cobalt-based superalloys [227, 240–242]. There is at least one instance in which a variant of severe plastic deformation, Equal Channel Angular Pressing and Torsion (ECAPT) has been used to compact Mo powders and refine grain size [243]. The ECAPT process reduced the powder size by 200% and enhanced mechanical properties. High strain deformation is particularly difficult in powder metallurgy produced compacts because of susceptibility to cracking [227]. However, the superposition of high back pressure during processes such as ECAP may reduce cracking. Back pressure has been shown to be effective for ECAP of other difficult-to-deform alloys [244–246].

Another consideration in SPD of superalloys is that the range of temperature for deformation generally needs to be closely controlled to manage the various precipitation reactions that can occur during deformation. Since SPD produces significant self-heating, strain rates need to be carefully controlled to avoid excessive heating and associated temperature excursions. One variant of SPD that may be particularly well suited for maintaining temperature control during nanostructuring Co-Cr superalloys is incremental ECAP (I-ECAP) [247–249] or incremental angular splitting (I-AS) [250, 251]. In these techniques shearing deformation is imposed discontinuously in small increments, but under conditions that are comparable to ECAP. These approaches eliminate heating associated with die-billet friction and allow removal of deformation-induced heating. In addition, there are no challenges with frictional die wear as are present in ECAP and other SPD processes. Aware of the challenges of SPD of multiphase high strength alloys such as Co-Cr superalloys, Yamanaka et al. [252, 253] have developed an approach to achieve ultrafine grain sizes in Co-Cr-Mo without the need for severe plastic deformation. They instead rely upon conventional forging using conditions that induce a novel mechanism of dynamic recrystallization.

Current uses of Co-Cr superalloys in biomedical applications include artificial heart valves, dental prosthesis, orthopedic fixation plates, artificial joint components, and vascular stents [254, 255]. Co-Cr is often chosen for applications in which there may be sliding wear because of its high hardness and wear resistance. The development of nanostructured Co-Cr superalloys for medical applications offers the prospect of higher strength and the possibility of even higher wear resistance through the effects of high shear on the morphology of precipitates and homogeneity of their distribution. Improving wear resistance is particularly important in in Co-Cr alloys since wear can create particulate debris. The cobalt and chromium particles have been shown to have some toxicological effects in humans [256, 257].

New medical applications of nanostructured Co-Cr alloys will appear only as processes to nanostructure are developed and validated. Surface oriented techniques or powder-based appear most promising. One technique of particularly high potential is Surface Mechanical Attrition Treatment (SMAT) [258, 259]. SMAT induces large deformation in surfaces by recurring impact of hard spheres. These deformations result in the formation of microstructures containing high densities of strengthening defects, including twins and the intersecting twin systems, dislocation walls, microbands, highly disoriented polygonal submicronic grains, and randomly oriented nanograms [258]. The SMAT technique has been successfully demonstrated for cobalt [260] and other difficult to deform metals [261–264]. The application of SMAT for enhancing the biocompatibility of titanium surfaces has been documented [265, 266].

1.2.4 Magnesium Alloys

Magnesium is highly promising for medical applications because of its very light weight and its ability to be bioabsorbed [267–272]. As the lightest of all structural metals (excepting beryllium) magnesium enables light-weighting of many medical structures, ranging from wheelchairs and stretchers to surgical tools, to vascular stents, to orthopedic implants [273–279]. Magnesium is also among the most biocompatible of metals. It is essential to human health, with approximately 35 grams distributed between bones, tissues, and blood in an average size human adult [270, 280]. The normal level of magnesium in the blood is 1.7 to 2.2 mg/dL [254]. Magnesium is necessary for many biochemical processes in our bodies, aiding in over 300 enzymatic reactions that help maintain normal functioning of muscles and nerves and regulation of heart beat, blood sugar levels, and the immune system [270]. The normal daily demand by the human body for magnesium is about 375 mg/day [280].

Increasing investigation of magnesium as a biomaterial has been stimulated by multiple trends. Magnesium alloys have become increasingly available worldwide because of the increasing demands for fuel savings through weight reduction in motor vehicles. At the same time, the cost of magnesium has declined to become comparable to aluminum, partly because of the availability of low priced alloys from China and improvements in the efficiency of primary magnesium production [279]. Global production of primary magnesium has increased from 260,800 tons in 1990, to 479,000 tons in 2000, to 809,000 tons in 2010 [281]. Magnesium and its alloys offer high specific strength, but has been limited in its use because of challenges associated with poor low room temperature workability and poor elevated temperature properties [282]. However, magnesium alloys have good machinability, weldability, castability, and formability (at least at high temperature), supporting their commercial use in a wide range of applications [279].

Compared to the major classes of alloys in medical use (stainless steel, titanium, Co-Cr) magnesium has by far the lowest strength. The ultimate tensile strength of magnesium alloys falls in the range 160 MPa to 380 MPa, while austenitic stainless steels range from 515 MPa to 1275 MPa, titanium alloys range from 240 MPa to 1380 MPa, and Co-Cr alloys range from 600 MPa to 2025 MPa [221]. The highest strength medical alloys, such as Co-Cr superalloys MP35N and MP159 have tensile elongations to failure of 8% − 10% at room temperature. Though much lower in strength, the elongation to failure for all magnesium alloys is low, ranging between 1% and 16%. This is due in part to the limited number of slip systems available in the hexagonal close packed crystal structure of magnesium. For orthopedic applications, magnesium has the distinct advantage of having an elastic modulus of 45 GPa, closer than