Stress and Strain
Stress, like pressure, is a local force expressed in units of force per unit area. Strain is a local deformation expressed as units of length per length.
Strength and Stiffness
Strength denotes the ultimate load a material can withstand before catastrophic failure. Stiffness refers to the rate at which a material deforms when a load is applied.
The ultimate goal of any fracture repair is to achieve a functional result that is indistinguishable from the pre-injury level.
Bone has the ability to optimize itself by forming in areas of high stress and not in areas of little stress according to Wolff's law. Its strength is dependent on the rate at which it is loaded. The severity of a fracture is directly related to the rate of loading. For example, after a rapidly loaded bone reaches the failure point, it releases more energy, which results in more complex fractures and greater soft tissue damage. On the other hand, fractures that occur at low loading rates tend to be simple fractures with less soft tissue damage because less energy was stored during loading and released at failure.
Fracture patterns are largely determined by the orientation of the forces that caused the fracture and the relative strength of the bone in each loading orientation. Compressive forces result in oblique fractures: when loaded in compression, bone fails along the lines of highest shear stress rather than compressive stress, which tends to be at an angle of 30 to 45 degrees to the direction of the compressive force. Tensile forces tend to result in transverse fractures perpendicular to the direction of loading. Most tensile fractures actually also have a bending component and are influenced by bone shape or the presence of open physes. Pure bending forces also result in transverse fractures: the convex side is under tension, and the concave side is under compression.
When considering repair options, it must be kept in mind that small fracture gaps may actually delay bone healing because they concentrate strain. Because bone formation requires a low strain environment (> 2%), the treatment of fractures should not just decrease the fracture gap, but also decrease interfragmentary strain. The body primarily decreases interfragmentary strain in two ways: fracture resorption and callus formation. Fracture resorption appears radiographically during the first few weeks of healing and is the result of osteoclastic removal of dead bone at the fracture margins. This, in turn, increases interfragmentary distance and decreases strain. Periosteal callus formation functions primarily to provide stability. A large periosteal callus allows stiffness to increase and secondarily decreases strain.
Most implant failure occurs in bending. Area moment of inertia (AMI) is a structural parameter, therefore not material dependent, and determines the ability of that structure (implant) to resist bending. It depends on the cross-sectional area of the implant in question. Different formulae are used for cylindrical (IM Pin) and rectangular (plate) structures. In general, the thicker the implant, the stiffer it is in bending. Therefore, a bone plate is stiffer in bending along its width than its depth. An example of this is the biomechanical superiority of applying a medial instead of a dorsal bone plate to distal radial fractures.
In pure compression or tension, the cross-sectional area alone determines a structure's strength, but when bending is involved, one side of structure experiences tension, one compression, while the plane along the center of the structure that experiences no force is termed the neutral axis.
Cerclage wire is probably one of the most inappropriately used orthopedic implants. Its only function is to provide tensile strength across an intact cylinder of bone. If greater than ~30 N of tension are applied, the wire will in general not slip; greater tension applied while the wire is being secured results in its ability to withstand greater load prior to loosening. Collapse of only 1% of the cylinder results in tension dropping to < 30 N. Cerclage wires should only be applied to fractures 2.5–3 times the diameter of the bone in length; at least two should be placed, but more is better, and they need to be spaced ~ ½ a bone diameter apart and away from the fracture ends.
The tensile strength an appropriately placed cerclage wire can exert is dependent on its cross-sectional area. For dogs > 20 kg, it is recommended to use 18-ga wire, while 20-ga wire can be used for those smaller. The weakest point in a cerclage wire is the method of securing.
The standard method of twisting cerclage wires is probably most commonly used and the one that has the lowest tension at time of completion: prior to cutting the mean tension is 116.6 N, but after cutting it decreases to 82 ± 46 N. More tension is lost if the wire is wiggled during cutting or if the twist is formed and then folded. Twisted wires do resist a greater peak load prior to deformation, but they are easily deformed, and once they are, function is lost. Twisted cerclage wires fail by untwisting.
Single-loop cerclage wires have a mean tension of 166.6 ± 42.2 N at time of completion, while double-loop cerclage wires achieve a mean tension of 392.0 ± 116.6 N. Appropriate technique, especially maintenance of tension during folding, is particularly important when placing these. Both single- and double-loop cerclage wires fail by unbending of the ends.
The strength and stiffness of intramedullary (IM) pins is determined by the AMI. If they are used alone, they should fill approximately 70% of the medullary cavity, but when used with plates and screws, filling of 20–40% of the medullary cavity optimizes efficacy. In conjunction with plates (plate-rod construct), the presence of an IM pin decreases plate strain and increases stiffness of the entire construct significantly.
Plates and Screws
There are many different plate manufacturers and options available today. Plates can generally be divided into non-locking, locking, and specialty plates. Examples of non-locking plates are dynamic compression (DCP) and limited-contact dynamic compression (LC-DCP) plates. Locking plates include locking compression plates (LCP), string of pearl (SOP), and the advanced locking plate system (ALPS). Specialty plates include cuttable, lengthening, and reconstruction plates.
Non-locking plates function by maintaining friction between the bone and the implants requiring accurate contouring. Locking plates, on the other hand, are an 'internal fixator' and result in a fixed-angle construct. Because the screws lock into the plate, as well as engage the bone, accurate contouring is not required. This also results in decreased damage to the periosteal blood supply.
For any of the plates listed above, appropriate screw placement is important. In general, purchase of 6 cortices above and below the fracture results in optimal biomechanical stability. The advent of self-tapping screws has decreased surgical time, but if the cutting flutes of these are not inserted completely through the trans-cortex, the pull-out strength is significantly decreased. However, when appropriately placed, the pull-out strength of self-tapping screws is comparable to regular screws.
Not only is the appropriate selection of implants important, but also the surgical technique employed. It is generally accepted that perfect reconstruction of most fractures is not in the animal's best interest and results in longer healing times due to destruction of blood supply and loss of the fracture hematoma. The very important exception to this is the articular fracture for which perfect reconstruction, ideally with compression, is required to minimize long-term osteoarthritis formation. Open-but-do-not-touch (OBDNT), minimally invasive plate osteosynthesis (MIPO), and closed external fixator placement are all recognized methods for minimizing iatrogenic trauma while achieving rapid healing.