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Bone geometry markedly influences structural mechanical properties. Axial stiffness, which is the resistance of bone to deformation during loading in tension or compression, is proportional to the cross‐sectional area, while bending and torsional stiffness depend on how the bone material is distributed around the axis of bending or torque. Material most distant from the neutral axis has the largest effect on resisting bending and torque loads. Two geometric properties, the area moment of inertia and the polar moment of inertia, quantify the contribution of geometry to a particular bone’s resistance to bending and torsion, respectively.

Application of a load at a distance from the centre of a bone induces a bending moment. The bending moment is a product of the magnitude of the force applied and the length of the moment arm about which the force is applied. Moment arm length is the perpendicular distance from the line of action of the force to an axis of rotation. A longer moment arm increases the bending effect of the force applied.

Figure 3.11 Example of a stress–strain curve of a bone sample.

Source: Modified from Lopez [43].

Figure 3.12 Compressive stress–strain behaviour of one compact and two trabecular bone samples demonstrating the influence of apparent density (P) on the material properties of bone. Compact bone is more dense than trabecular bone and exhibits greater stiffness and strength under compressive loading, but tolerates minimal strain before failure. Trabecular bone is porous with a low resistance to compressive stress but is capable of enduring much higher strains than compact bone.

Source: Modified from Keaveny and Hayes [72].

The amount a bone will deform under a bending force is related to the magnitude of the bending moment, the elastic modulus of the bone material and the area moment of inertia (I) about the neutral axis (Figure 3.13). The area moment of inertia is defined as the capacity of a cross‐section to resist bending. For any given bending moment, bone deformation can be reduced by decreasing moment arm length, increasing stiffness of the bone (i.e. larger elastic modulus) or increasing the area moment of inertia.

Figure 3.13 Factors influencing the deformation of bone subjected to a bending force include the magnitude of the bending moment (a product of the force applied (F) and the length of the moment arm), the elastic modulus of the bone material, and the area moment of inertia about a neutral axis.

Factors that affect bone strength and stiffness in torsion are similar to those that operate in bending: the applied torque (force applied to induce a rotation), the length of the bone, the shear modulus and the polar moment of inertia (J) about the torsional axis. The polar moment of inertia is defined as the moment of inertia with respect to an axis perpendicular to the plane of the area. The shear stress created in a bone when loaded under torsion is inversely related to the polar moment of inertia. Thus, in a bone with a high polar moment of inertia, the same torque will result in smaller shear stress than in a bone with a lower polar moment of inertia.

The area moment of inertia and the polar moment of inertia are proportional to the fourth power of the radius for circular cross‐sections (Figure 3.14). For example, if the diaphysis of the bone is considered a hollow cylinder [73–75], small increases in bone diameter will result in exponentially greater bending and torsional strengths. Periosteal callus will contribute substantially more to the bending and torsional stiffness of a bone than endosteal callus, as the new bone material is located further away from the diaphyseal (neutral) axis. Similarly, a small amount of compact bone loss near the marrow cavity may have a relatively small effect on overall bending and torsional rigidity.

Although helpful in fostering a conceptual understanding, assumptions of cylindrical or elliptical geometry underestimate the complexity of bone structure [76]. Experimentally, finite‐element (FE) modelling, wherein geometry and material properties are obtained from quantitative computed tomography (QCT), is used to generate 3D models that more accurately predict the structural response of bones with irregular and variable cross‐sectional characteristics to different loading conditions [77, 78].