4.1.2 Historical Perspective

The evolution of fracture repair has progressed greatly since it was first reported in 1886 by Carl Hansmann in conjunction with advances in anesthesia, antisepsis, and radiography [3]. Hansmanns’ original device consisted of using nickel coated sheet steel with part of the plate and shanks of screws protruding through the skin to allow for percutaneous removal four to eight weeks later (Figure 4.1). Although revolutionary at the time, the concept was limited as it failed to fully appreciate the engineering challenges associated with orthopedic repair such laws of stress, strain, and leverage. The plate functioned merely as a connector without the ability to allow for approximation or even compression of osseous fragments. The progress of osteosynthesis was further advanced in the 1960s by the Arbeitsgemeinschaft fur Osteosynthesefragen (AO) group and the development of the dynamic compression plate (DCP) in 1969 utilizing the laws of friction to allow for fragmentary compression [4, 5]. However, the extensive underside of the DCP interfered with the underlying periosteum and the blood supply to the cortex, which led to the development of the LC‐DCP in 1990 (Figure 4.2) [6]. The LC‐DCP has an approximately 50% reduced footprint to minimize this interference with biological healing, although the plate is still restricted in that it must be compressed to the bone and thus may cause a disturbance in the vascular supply [4, 7].

The successful use of the LC‐DCP is dependent on the frictional forces generated between the remaining, reduced footprint of the plate and the cortex of the bone. This frictional force is essentially limited by the degree of screw torque that can be placed. Using conventional screws, a plate can be compressed to the bone with a force of 2000–3000 N [8]. However, the resulting frictional force this generates in preventing plate sliding is dependent on the coefficient of friction between the two surfaces. To overcome these challenges and increase the coefficient of friction, soft tissue stripping and bone cements have been used. However, both of these have limitations and tissue stripping can further damage the periosteal blood supply. Early attempts to create a device less dependent on the bone‐plate frictional force or “internal fixator” included the point‐contact fixator (PC‐Fix) [9] and less invasive stabilization system (LISS plate) [10] with the resultant development of the LCP in 2001 [11].

4.1.2.1What Do We Mean by Better?

The term better can have many connotations. By “better,” we are implying we want to achieve optimal fracture healing. This suggest the restoration of the normal functional biomechanics of the bone, i.e. a return to pre‐fracture stiffness and strength of the osseous tissues [12]. From a clinical context, there are a plethora of factors that go into determining an optimal outcome. In a biomechanical context, we will assume that the interpretation of “better” is assigned to bone healing and biomechanical superiority of one implant versus the other in conjunction with what each technique may contribute to the overall biomechanical picture. In order to do this, one must understand the biological concepts of bone healing, and fracture repair. A variety of biomechanical concepts, definitions of which are provided in Table 4.1.

Biomechanical Term	Definition
Brittle Material	Material with a low or absent capacity for plastic deformation prior to ultimate failure.
Ductile Material	Material with a high capacity for plastic deformation prior to ultimate failure.
Elasticity	Reversible deformation of a material. When a material is unloaded after loading in the elastic range, it returns to its original shape and dimensions.
Elastic Modulus (E)	Constant of proportionality between stress and strain, also termed Young’s modulus., where E = 𝛿σ / 𝛿ε = tan φ.
Friction	The frictional force is directly proportional to the normal force (Amonton’s second law). The frictional force Ff is parallel to the surface and is in a directly opposite direction to the net applied external force, Fe the normal force Fn, is force exerted by each surface, directed perpendicular to the surface. Ff and Fn are connection via a proportional constant, μ (frictional coefficient), such that Ff = μ Fn.
Hysteresis	For time‐dependent elastic materials, during cyclic loading within the elastic region, hysteresis is the energy dissipated between loading and unloading cycles, usually in the form of heat energy, governed by the coefficient of restitution.
Plasticity	Plasticity describes the irreversible deformation of a material undergoing permanent morphological alterations in response to applied forces.
Stability	Degree of relative movement between structures. Relative stability in orthopedics indicates that motion exists between fracture gaps under loading, but return to the initial position during unloading. Healing is via callus formation (endochondral ossification). Absolute stability is where no motion occurs between fracture fragments during loading, and healing occurs through osteonal cutting cones (endosteal healing).
Stiffness	The resistance of a material to deform under load. Materials with high stiffness (rigid material) deform less under a given load (c.f flexible). The product of the cross‐sectional area (A) and the elastic modulus (E) expresses axial stiffness, such that Rax = A E. Bending stiffness is defined as the product of the axial area moment of inertia and the elastic modulus, such that Rbe = Iax E,
Strain, ε	Deformation of a material under a given load. It is expressed as elongation per unit of original length (l) and is dimensionless, although often given as a percentage, where ε = Δl/l0. For a material undergoing deformation, the ratio of transverse strain to axial strain is termed Poisson’s ration, or 𝛎.
Strength	Generally, strength indicates how much force a material can support before ultimate failure.
Stress, σ	Force per unit cross‐section area. Stress is directly proportional to strain with the elastic modulus as a constant of proportionality. Unit is Newton per m2 (Pascal, Pa). Can exist as normal stress, acting perpendicular to a surface, or shear stress when the force acts parallel to a surface.
Toughness	The energy absorbed by a structure during the loading process, determined by integrating the stress–strain curve, such that: where ϵ is strain ϵf is the strain upon failure σ is the stress Generally, toughness indicates how much energy a material can absorb before ultimate failure.
Fatigue	The relationship between stress magnitude and number of loading cycles, described by Wöhler’s curve. Fatigue stress is defined as an asymptotic line of Wöhler curve.

4.1.3 Fracture Stability

Stability is a crucial concept in orthopedics in order to permit successful fracture healing. As defined in Table 4.1, stability can be either relative or absolute depending on the method of repair [14]. Thus, the degree of implant and fragment stability determines the amount of strain at a fracture site; the level of stain present will determine the type of bone healing that occurs. Absolute stability occurs where fracture gap strain is less than 2% and results in primary bone healing via osteonal cutting cones if the fracture gap is <200 μm. Relative stability occurs where fracture gap strain is 2–10% and healing is via callus formation. At strains in excess of 10%, fracture healing cannot occur, and the site is destined to form a non‐union or mal‐union.

When fractures are loaded, and the fragments return to their original configuration when unloaded (i.e. elastic deformation), relative stability is present. If after loading, the fracture fragments do not return to their original configuration, plastic deformation has occurred and the fracture is considered unstable. The ability of different tissues to heal in the presence of strain varies. Lamellar bone has the lowest strain tolerance of 2% while granulation tissue can form while undergoing 100% strain [14, 15]. The advantage of callus and relative stability is that the process occurs much more rapidly than direct primary bone healing with absolute stability[14].

Compressing a fracture using a DCP can reduce gap lengths between unstable fragment sections to near zero values. As a result, any fracture site motion that occurs in this near zero gap will result in very high gap strains if the motion persists [16]. When presented with high fracture gap strains preventing healing, reduction in strain can occur via (i) increasing gap length, such as occurs with bone resorption at a fracture site, fracture comminution or imperfect reduction, or (ii) decreasing motion present at the fracture gap. Placing the LC‐DCP on the tension surface of the bone, where the metallic plate will be the strongest, gap motion and therefore fracture strain, will be limited as long as the LC‐DCP can maintain this function.

4.2 The LC‐DCP

4.2.1 Construct Basics

Plate‐bone‐screw constructs can function as either load‐sharing or load‐bearing devices. The goal of these conventional plating techniques is to provide absolute stability. Occurrences when absolute stability and primary bone healing are essential include articular fractures where joint congruity is crucial; it has been shown that it is important to limit any joint step off to be less than a critical value of 2 mm of incongruity [17–19]. When compression plates are axially loaded in either tension or compression, they serve to convert these forces to shear stress (parallel to the surface) at the bone‐plate interface. This shear stress is countered by a frictional force. The normal force, which acts perpendicular to the plate/bone surface, is equal to the axial force generated by the torque applied to the screws fixing the plate to the bone, which is approximately 3–5 Nm for 3.5 mm cortical screws placed into human femora [20]. However, not all screws are tightened with the same degree of torque; therefore, the screw with the greatest insertional torque bears the greatest amount of load in this system. Currently, during screw insertion, surgeons use subjective feel when inserting these screws and stop tightening the screw (stopping torque) when they feel it is “tight,” as further torque will strip the threads (stripping torque). It has been demonstrated in several studies that human surgeons achieve a stopping/stripping torque ratio of 66–92%, resulting in under‐ and overtightening of screws on a regular basis [21–24]. If screw insertion torque exceeds the shear resistance of the bone, then screws will strip and there may be subsequent loss of fixation. Thus, when an external force, F_e, is applied which exceeds the frictional force, F_f, the plate will slip on the bone (Figure 4.3).

4.2.2 Axial Loads

One of the major benefits of locking screws is that they can be used in bone of poor quality, such as osteoporotic bone. With conventional plating, osteoporotic bone can only withstand a maximum screw insertion torque of approximately 3 Nm, with values often less than this [14]. Experimentally, it has been shown that 3 Nm of screw torque permits motion between the plate and the bone at loads as low as 500 N [9, 20]. This screw torque is insufficient to generate a sufficient normal force, F_n to prevent plate and fracture motion, which can lead to excessive gap strains that exceed the 10% maximum for secondary bone healing. This 3 Nm value is important in veterinary medicine in relation to neonatal surgeries. Although the process of osteoporosis that occurs in humans is less of a concern in animals, it has been shown that a significant number of 4.5 mm cortical screws placed in neonatal calf femora will strip prior to achieving 3 Nm of torque [25].

The success of the LC‐DCP is therefore highly dependent on the level of frictional forces that can be generated. As such, increasing the coefficient of friction, μ, or increasing the screw torque to increase the normal force F_n, will lead to increases in the frictional force, F_f, that the plate can generate in order to resist an applied load, F_e. Once the frictional force between the plate and the bone is overcome, the resistance to axial loading is transferred to the single screw furthest from the fracture site in the direction of loading. For a 3.5 mm cortical screw, the maximum load that can be withstood once motion has occurred at the plate–bone interface is 1200 N [20]. As the screwhead is not locked in conventional plating, the cis‐cortex for that screw experiences high stresses, which can result in bone resorption and screw loosening if they exceed the strength of the bone. Conventional screws therefore fail by this form of toggling within the bone and, as each screw effectively functions in series, all stresses are concentrated at a single individual screw at any given time [14]. With the resultant screw loosening, there is increased strain at the fracture sites and increased motion, which prevents healing via callus formation once it exceeds 10% strain. It should be noted, therefore, that the LC‐DCP construct is strongest immediately after application; with the progress of time, the axial screw forces exerted by the screws is diminished due to remodeling around the screw threads, which leads to a corresponding decrease in normal force F_n and the fictional force F_f at the bone–plate interface.

4.2.3 Bending Loads

For conventional plating methods, resistance to bending loads is equal to the bending stiffness of the plate for gap lengths greater than zero. When bone fragments are compressed and the gap length is equal to zero, the resistance to bending is determined by the resistance of the bone threads of a single screw to shear stress. The location of which screw experiences these stresses is dependent on the biomechanics of the fracture and the plate. When the plate is placed on the tension surface of a bone during bending, the highest shear stresses associated with the screw threads will occur at the screws at the end of the plate. If the plate is placed on the compression side of the fracture, the threads and screws closest to the fracture site will have the highest shear stresses. The most critical screws placed are those in close proximity to the fracture, thus favoring the placement of the LC‐DCP on the tension surface of the bone. High screw shear stresses eventually culminate in screw pullout (Figure 4.4c). The required pullout force depends on bone quality, cortical thickness, and outer diameters of the screw.

4.2.4 The Bone–Screw Interface

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