**2. Biocompatible metallic materials**

Medical parts as implants, stents, scaffolds and fixation devices are used when a bone is broken, for supportive structural purposes, bearing the body's load until the tissue can properly regenerate. These parts must present the following properties:


Biomaterials currently used for this application can be polymers, metals, ceramics or composites. Advantage of metals for load-bearing application is the combination of better mechanical strength compared to polymers and better tenacity when compared to ceramics. The most common metals used are titanium alloys, stainless steel and chromium-cobalt alloy, that are bioinert. These metals have high strength to support body load and can keep a good integrity, staying longer in the body due to their high-corrosion resistance [5]. Iron, magnesium, and zinc are also biocompatible metals, with the advantage of being biodegradable in physiological environment. Other metallic biomaterials include shape memory alloys, tantalum and some precious metals [6].

**Table 1** summarizes some mechanical properties, density, pros and cons of the usual biocompatible metals [6, 7]. For comparison, the properties of the cortical bones are: a density around 1.8 g/cm3 , a modulus of elasticity of the order of 17.0–20.0 GPa in longitudinal direction and 6.0–13.0 GPa in the transversal direction, the tensile strength in the range of 78.8–151.0 MPa in longitudinal direction and 51.0–56.0 MPa in transversal direction, and a fracture toughness in the range of 2–12 MPa.m1/2 [8].


*Mg-Based Composites for Biomedical Applications DOI: http://dx.doi.org/10.5772/intechopen.95079*

#### **Table 1.**

*Characteristics of usual metallic implant materials. Values taken from Ref. [6].*

It is possible to notice a considerable mismatch of mechanical properties of these bioinert metals and the cortical bones. When the stress transfer between the implant and the tissue is not homogeneous, stress shielding of the bone may occurs. This takes place by differences between Young's moduli of the implant and the bone, what may interfere the regeneration process, cause pain, bone atrophy with resorption of the bone surrounding the implant, loosening of the implant and possibly inducing a new fracturing [9]. Almost 10% from overall operations would undergo for revision surgery for replacing a previously implant, and 79% of all revisions were due to implant loosening [10]. The most common cause of implant loosening is the loss of bone mass due to stress shielding [10].

In the cases which the tissue can regenerate, the implant loses its function after the healing process, and additional surgery may be required to remove it. Besides, an implant which stay in the body for a long period of time can induce some complications, i.e. release toxic elements, permanent physical irritation, a chronic inflammatory response, or infection caused by debris and corrosion products [11, 12].

Biodegradable implants seem to be a clever choice for application where the body can eventually regenerate, since this would dispense the need of a secondary surgery to remove it, diminishing risks of inflammation, pain, surgery risks and costs [2, 12]. So, it is of great interest to develop devices which can be gradually degraded and naturally absorbed by the body while assist tissue healing. Yet, it has been challenging to obtain the combination of a controllable degradation rate, good biocompatibility, and suitable mechanical properties.

## **3. Magnesium as a biodegradable material**

Among the available biodegradable metals, magnesium is very advantageous since it offers necessary stiffness, toughness and tenacity, high damping capacity, very low density and high biocompatibility. It is the fourth most abundant mineral in the body, being present mostly as surface substituents of the bone's hydroxyapatite, and in skeletal muscle and soft tissue [13]. Also, Mg is involved in more than 300 essential metabolic reactions, being a cofactor for several enzymes that stabilize RNA and DNA structures. It is necessary to assist in cardiac, muscular, nervous, bone and renal function [14]. Its ion (Mg2+) that is released during degradation is used in the regular metabolism having no critical toxic limits or side effects reported [3, 13]. It has great ability to be absorb and excreted by the body without causing any harm (for example, the corrosion product MgO is totally removed in the urine [11]). During the first half of the 20th century, magnesium was used in surgeries to connect blood vessels and, in some cases, portions of the intestine in animals and in humans. The most advanced clinical applications are biodegradable cardiovascular magnesium stents; however, because it is bioabsorbable, research has been carried out aiming orthopedic applications as non-permanent implants.

The biggest issue that hinders the use of Mg as a biomedical material is its too fast corrosion in body fluid. This characteristic can be related to three main factors: Mg alloys are very reactive, the corrosion products formed in aqueous solutions containing chloride is not protective, and the micro-galvanic interaction between the constituent phases causes the corrosion of the Mg alloys to be significantly greater than that of the Mg of high purity [15].

Besides an early loss their structural function, a fast corrosion releases a big quantity of hydrogen, what may induce some local pH changes, affecting some physiological reactions and leading to an alkaline poisoning effect. Moreover, gas bubbles can accumulate in the implants surrounding surface delay healing and causing necroses due tissue detachments in some cases [3, 16].

#### **3.1 Mg corrosion**

Two types of corrosion occur preferentially in Mg and its alloys, galvanic corrosion and localized corrosion (by pitting or filiform). Galvanic corrosion occurs because magnesium generally behaves anodically in contact with other metals, and it is often used as a sacrificial anode. The standard reduction potential of Mg to form Mg2+ is E ° = −2.37 VSHE and one of its main alloy elements is aluminum (Al), which has a reduction potential E ° = −1.66 VSHE in the reaction reduction for Al3+. Galvanic pairs are generally formed when the concentration of the alloying element exceeds the maximum solubility limit. The extent of the galvanic effect depends on several factors, such as the crystalline orientation of the magnesium matrix, the type of secondary phases, particles of impurities, the size of the grain and the medium.

The concentration and distribution of the secondary phases are also important. A fine and continuous distribution of the secondary phases normally increases the corrosion resistance of the metal in various media. Alloy elements and other intermetallic phases formed, with electrochemical potential close to that of magnesium, can increase corrosion resistance, reducing galvanic internal corrosion [16].

Magnesium, when immersed in an aqueous solution, presents a double oxide film, an internal layer of MgO, which is adhered to the metal surface and an external and porous layer of Mg (OH)2 [17]. MgO oxide is a n-type semiconductor, with ΔG formation of −136 kcal/mol at 25°C, has a PB (Pilling Bedworth) ratio of less than 1, and is therefore only partially protective, since the volume of the oxide formed is less than the volume of the reagent metal. The Mg (OH)2 film has a lamellar structure that facilitates basal cleavage, and has a PB ratio of 1.77, which generates internal compressive stress, leading to cracks in the Mg(OH)2 film.

*Mg-Based Composites for Biomedical Applications DOI: http://dx.doi.org/10.5772/intechopen.95079*

In the case of magnesium, unlike aluminum or steel, the oxide layer is crystalline. There is no epitaxy between the oxide layer and the matrix with the compact hexagonal mesh (HCP), leading to a high compressive stress of the layer [18]. One of the ways to reduce discontinuity and have less disorder between the oxide layer and the metal surface is to introduce a large fraction of grain outlines per unit area [19].

Cathodic reactions in neutral or alkaline aqueous media are oxygen reduction to hydrogen evolution, expressed in Eq. 1:

$$\text{H}\_2\text{O}\_{(l)} + 2\text{ e}^- \rightarrow \text{H}\_{2(g)} + 2\text{ OH}^-\_{(aq)}\tag{1}$$

In order to have a complete understanding of the Mg corrosion mechanism, it is necessary to consider a phenomenon called Negative Difference Effect (NDE). This phenomenon is experimentally characterized by an unexpected increase in the hydrogen evolution reaction when the anodic overvoltage is increased. NDE is the main cathodic reaction in the corrosion of Mg and occurs spontaneously in Mg under open circuit conditions. When Mg and its alloys are anodically polarized, hydrogen evolution also occurs, which is normally restricted to the cathodic branch. Anodic polarization causes higher rates of dissolution, and higher rates of NDE are also observed [20]. To explain the corrosion of magnesium and more particularly the NDE phenomenon, several mechanisms have been proposed including the formation of magnesium hydrides, metastable monovalent ions, hydroxides and magnesium oxides [17].

The corrosion rate of magnesium-based biomaterials, when implanted in the human body, depends on the composition of the material and the environmental conditions in which the implant is found, such as temperature, pH and concentration of other ions. A study [21] indicates that, the effects of Mg corrosion products on cellular activity is concentration-dependent, on bone marrow derived stem cells and on osteoclastogenesis *in vitro*. This will vary depending on the state of differentiation of cells and length of exposure. Also, the presence of the corrosion products significantly altered the cells' metabolic and proliferative activities, which further affected cell fusion/differentiation. Maradze and coauthors [21] summarizes the cellular response in the presence of a corroding Mg biomaterial *in vivo* in **Figure 1**.

The corrosion rate of Mg can be reduced by increasing the intensity of the basal planes parallel to the surface [22]. Other strategies to diminishing the degradation rate are grain refinement, alloying, surface modification, conversion coating, and incorporation of other materials to create composites. It is worth noting that, in the case of composites, some of these strategies can be combined. For example, a processing technique that enables a refined microstructure can be selected. Grain refinement has a recognized importance for improving mechanical resistance and also for diminishing the corrosion rate with time due to the development of a more homogeneous protective layer on composites surface [23–25]. Also, the matrix can either be pure Mg or an alloy. Although many Mg-alloys are been developed for improved corrosion resistance and superior mechanical strength [4, 11, 26], care must be taken to choose a composition free of any toxic elements. The insertion of a biocompatible second phase in a magnesium matrix could bring a better mechanical resistance and less pitch corrosion. Furthermore, some bioactive materials that could promote a better bonding between implant and the tissue can be incorporated and accelerate body's healing response in some cases. Also, the addition of a hard phase, combined with mechanical processing, enables a better grain refinement and strengthening. Thus, magnesium matrix composites are potential candidates for structural orthopedic implants.

**Figure 1.**

*Representation of cellular response in the presence of a corroding Mg biomaterial in vivo [21]. Scientific reports – Open access.*
