**6. Biomechanics of implant biomaterials**

Orthopedic implants are commonly used for different types of surgical procedures to gain optimal function and provide stability to bone tissue. When inserting these implants, the characteristics of the material are important for surgical success, and the ideal implant must be biocompatible and nonallergenic from a biological point of view. However, when contoured an implant to the bone surface, its resistance can change significantly. Implants can be temporary or permanent in the body, and metal possesses properties that make it acceptable for bone repair. In orthopedic implants, metals and their alloys were the first materials used in their production, primarily due to their superior strength and biocompatibility. The metals used for implant production include nickel, iron, cobalt, titanium, vanadium, and aluminum. Metal alloys aim to achieve specific properties in the final mixture, such as ductility, strength, elasticity, and corrosion resistance [65]. Ductility is the ability of a material to absorb energy and plastically deform without fracturing. The term ductility is sometimes used to encompass both types of plasticity: tensile (ductility) and compressive (malleability). Current alloys used in orthopedic metal-based implants include stainless steels, cobalt-based alloys, and titanium-based alloys.

#### **6.1 Stainless steel**

Stainless steel 18-8 (18% chromium, 8% nickel) is the most common alloy. It has superior corrosion resistance obtained through compositional modifications by using additional metals, especially Cr [66]. The inclusion of Cr allows Cr2O3 promotes the formation of a strong and adherent layer that is beneficial for healing. Stainless steel is commonly used in removable orthopedic devices, such as plates, screws, and intramedullary pins, due to its affordability [50, 67]. Currently, the new stainless steel-based alloys contain Co-Cr, Mn, Ni, and a high nitrogen content. Stainless steel alloys have high resistance to corrosion due to their high chromium content (more than 12 wt%), which enables the formation of a strongly adherent, self-healing, and corrosion-resistant coating of Cr2O3 oxide. Different types of stainless steel are available for implant production, and the most widely used is austenitic stainless steel. Austenitic stainless steel, which contains austenite-stabilizing elements such as Ni or Mn, is the most commonly used type of stainless steel for implant manufacture. AISI 316L is the most widely used stainless steel in clinical applications, containing 0.03 wt % C, 17–20 wt% Cr, 12–14 wt% Ni, 2–3 wt% Mo, and minor amounts of nitrogen, manganese, phosphorus, silicon, and sulfur [68].

When compared to bone tissue, stainless steel alloys are significantly stiffer and have proven to be durable enough for osteosynthesis [69]. Additionally, stainless steel is relatively inexpensive and biologically well-tolerated, with a smooth surface from electropolishing. It is also ductile enough to allow for contouring of the plate without breaking [69].

#### **6.2 Titanium and titanium-based alloys**

Titanium (Ti) and its alloys were initially used in the field of aeronautics but later gained significant interest in the biomedical field due to their remarkable properties. These properties include a moderate elastic modulus of about 110GPa, good corrosion resistance, and low density (around 4700kgm�<sup>3</sup> ) [70].

For orthopedic devices, Ti may be used alone or in alloys with other metals, most commonly commercially pure (CP)-Ti and Ti-6Al-4V alloy; this designation refers to its chemical composition of almost 90% titanium, 6% aluminum, 4% vanadium, 0.25% iron (maximum content), and 0.2% oxygen (maximum content). They both provide stable fixation and a low risk of implant loosening [70].

The report of the osseointegration phenomenon for Ti implants by Branemark [71] led to the development of dental and surgical applications of Ti alloys. This property enables titanium and its alloys to tightly integrate with bone, resulting in the improved long-term behavior of the implanted devices, which in turn reduces the risks of loosening and failure.

CP Ti, grade 4 (ASTM F67) and Ti6Al4V (ASTM F136) are the titanium alloys most commonly used for orthopedic implants. For CP Ti-based implants, four grades are currently available varying their oxygen content. CP Ti grade 4 is the type having the highest amount of oxygen (up to 0.4%) and, consequently, the highest tensile and yield strengths [72].

The use of pure titanium has the following advantages: low weight and very good corrosion resistance, especially in saline solution. Ti and its alloys possess outstanding corrosion resistance, which can be attributed to the creation of a robust and adherent TiO2 oxide layer on their surface. About the surface properties, namely, wear, the performance is poor due to the low shear resistance of Ti and Ti alloys.

The ability to become tightly integrated into the bone greatly improves the longterm mechanical behavior of the implant as well as reduces the risk of loosening and failure of the device [73–75].

CP Ti, with a single-phase alpha microstructure, is currently used for dental implants production, while Ti6Al4V, with a biphasic alpha-beta microstructure, is mostly used in orthopedic implants and prostheses. The Al and V alloying elements stabilize the alpha-beta microstructure and improve the mechanical properties of CP Ti (typically twice the yield and ultimate strength values of CP Ti). Mechanical properties of CP Ti and their alloys can be altered by heat treatment and mechanical working. Although Ti and Ti alloys are characterized by an array of excellent properties (e.g., favorable mechanical characteristics, corrosion resistance, fatigue-corrosion

## *Biomechanical Basis of Bone Fracture and Fracture Osteosynthesis in Small Animals DOI: http://dx.doi.org/10.5772/intechopen.112777*

resistance, low density, and relatively low Young modulus), their processing is complex whether it is by machining, forging, or heat treating.

CP Ti and Ti alloys, on the other hand, more closely matches the modulus of elasticity of bone. This flexibility may be more conducive to fracture healing in points where more strain is required for a bone regeneration to develop. Titanium alloy is also more resistant to cyclic loading and notch sensitivity.

### **6.3 Cobalt-based alloys**

Cobalt-based alloys are superior to stainless steel in terms of strength [76]. However, cobalt alloys have better biocompatibility and are more corrosion-resistant. But these alloys are more expensive to produce. Cobalt-chromium-molybdenum alloy variants are specifically used for hip prosthesis implants due to their high abrasion resistance [77, 78].

### **6.4 Fatigue failure and cyclic loading of implants**

Fatigue failure and cyclic loading are two important concepts for guiding the choice of orthopedic implants to avoid construct failure. Clinically, acute deformation or catastrophic failure by a single applied load is a rare event. Several factors can influence the fatigue failure phenomenon such as the magnitude of the applied load (by consequence generates stress within the implant), the geometry of the implant, the material and how it was handled and manufactured, and the local environment of the fracture.

Experimental determination of the fatigue behavior of a material involves creating an S versus N curve, where S represents the applied stress and N represents the number of cycles required for failure (plotted logarithmically). If the applied stress is greater than the yield stress of the implant, the material fails in a few cycles, such as repeatedly bending a paper clip. The number of cycles to cause failure increases as the applied stress is reduced. The stress level at which a material can withstand an infinite number of cycles without failure is called the endurance limit, which is approximately 50% of the ultimate tensile stress for most metals [2]. A similar process can be used to characterize a fatigue behavior of a structure such as a bone plate, applying a load versus number curve. After determining the yield load, a series of progressively decreasing peak loads are established, and the number of cycles required to reach a defined failure point, such as breakage or reduced stiffness, is recorded. The number of cycles required to reach failure increases as the applied load is reduced. An implant's performance may be considered adequate if it survives a clinically relevant number of cycles, which is often set at 10<sup>6</sup> [2].

The response curve for implant construct may be more complex to interpret because geometry, material, and manufacturing factors may all interact. Factors, such as plate screw holes, may cause local stress concentrations that accelerate fatigue failure. The degree of cold working and even the purity of the production process may vary among different manufacturers of similar implants. Macroscopically visible small imperfections and cracks can trigger the implant failure cascade. Surgeons should also be aware that small notches on the surface of a structure can significantly decrease the endurance limit because it is a stress riser and should prompt intraoperative replacement.

In clinical practice, fatigue failure can be avoided by selecting implants of appropriate strength and dimension for the weight and bone size of the animal, minimizing notching, and, through good client compliance to discharge indications, reducing the magnitude and frequency of the applied loads.

Optimizing the rate at which the one regenerates at the fracture gap consolidates also helps avoid fatigue failure because stress in the plate decreases during the regenerative phase of bone due to progressive load sharing.

The local environmental factors related to fractures also can influence fatigue failure of implants. Factors like load sharing between plate and bone, when anatomical reconstruction of fracture is possible (e.g., simple fracture of long bones), highly reduce the cyclic loading magnitude and early failure of the implants [79]. Technical factors related to the correct application of DCP or LC-DCP plates are also crucial to fatigue failure. An illustrative example is the use of DCP plates without pre-bending or overbending of at line fracture. In this case scenario, there will be compression under the plate and distraction on the opposite cortex, causing failure of load sharing and altering strain distribution over the fracture and increasing the magnitude of cyclic loading and fatigue failure of implant more probable [80]. The correct magnitude of pre-bending of the plate is 2 mm prior to a fixation on a convex side of long bones and provides the most compression at the far cortex and consequently the load sharing between bones and plates [80].

LCP has over the years progressively replaced the use of DCP and LC-DCP plates. One of the main advantages of applying this type of implant is the possibility of applying those without adequately contoured and affixed directly to the bone for stable internal fixation of the fracture. For this reason, it has been used in minimally invasive osteosynthesis modalities such as in MIPO and supports biological osteosynthesis by functioning as an internal fixator, rather than as a full (DCP) or limited contact bone plate (LC-DCP) [18, 81]. Additionally, it was reported that LCPs were more resistant to cyclic loading in different force vectors than DCP and LC-DCP [82]. However, to maintain biomechanical advantages, it is advisable that LCP must not be more than 2 mm away from the surface of the bone [81–83].

Bone regeneration in high-strain fractures occurs only if the interfragmentary strain is less than 2%. According to Claes et al., transverse line osteotomies can tolerate up to 2 mm of micromotion without causing harmful damage to bone regeneration [24]. In this type of fracture, anatomical reconstruction is necessary and the strain at the fracture site caused by different force vectors must be neutralized by the implants during the reparative phase of bone regeneration to avoid complications such as delayed or nonunion [84]. With high strain rates, the magnitude and frequency of loading cycles are also greater because the animal will use the limb very early, and the implants will endure a greater number of loading cycles predisposing to fatigue failure. On the other hand, the reparative phase of bone regeneration develops over time, alleviating the magnitude of load cycles due to load sharing.

In low-strain fractures, the interfragmentary movement is not very harmful to the repair process woven bone can tolerate 2–10% of interfragmentary strain [85]. The main objective in this type of fracture is the indirect reduction of bone fragments with bridge plating or external fixation, aiming to re-establish the mechanical axis and bone length and promote secondary bone healing by relative stability [86]. The great advantage of this method is the possibility of a minimally invasive application, and therefore, it is appropriately used in MIPO, where the preservation of the fracture environment is maximized, and bone healing is optimized and even faster than in open reduction and internal fixation (ORIF) [13, 56, 61, 87]. On the other side, from a mechanical standpoint, plates experienced a greater magnitude of strain, increasing the risk of fatigue failure. However, there are surgical options for a sparing effect on

*Biomechanical Basis of Bone Fracture and Fracture Osteosynthesis in Small Animals DOI: http://dx.doi.org/10.5772/intechopen.112777*

**Figure 13.** *Image of biological healing plate with a central section without screw holes.*

plate site that bridges fracture site and reduce strain. The use of an intramedullary rod (IMR) also helps the restoration of alignment, a substantial challenge in MIPO inherent to the lack of fragment observation and biological healing plates [53, 88]. These former implants are designed to support high strain values for bending and torsional force vectors by possessing a central section without screw holes (**Figure 13**). The screw holes located at the plate's outer section allow the implant to be fixed to the intact proximal and distal fragments, which avoids the need for anatomical reduction of the diaphysis. Additionally, the use of a locking version of this type of plate can improve the performance of the implant by decreasing the pull-out of screws and the need to exactly contour the plate to the bone surface. By applying a MIPO approach, the soft tissue disruption can be minimized, improving biological factors at comminuted fracture sites and hastening bone regeneration.
