LOADING OF THE FEMORAL NECK
The loads acting on the hip joint are among the largest loads occurring in the human body. These loads are generated by external and internal forces. The external forces at the hip joint are primarily created through the interaction of the foot with the ground during walking, running, or jumping, the so-called ground reaction forces (GRF). While standing on one leg, the GRF reflect the weight of the body and are measured in multiples of body weight (BW) to adjust for variable body mass. During locomotion, dynamic forces and mass inertia increase the GRF depending on the speed of locomotion. The largest GRF during walking are created when the heel strikes the ground and the body decelerates downward by transferring the load from the heel to the front foot. Although for level walking at normal speed, the GRF remain below 130% BW, the forces increase to above 300% BW during running and can reach up to 5 times BW for high-impact activities, like jumping.1
The direction of the external forces is mainly vertical to support the body in the gravitational field. Forces in the anterior–posterior direction are considerably smaller compared with the vertical forces. The anterior force peaks at about 20% BW when the forefoot is pushed off the ground and the body is propelled forward. Medial–lateral forces counteract the sway of the body and are responsible for the maintenance of balance. They constitute the smallest force component of GRF and typically remain below 10% BW.
The bony skeleton transfers the loads occurring at the foot to the hip joint. Peak loads occurring during impact activities are dampened through soft tissues in the foot and cartilage in the ankle and knee joints. The anatomy and bipedal gait of humans create lever arms between the point of ground contact of the foot and the point of reaction in the hip joint, resulting in bending and torsional moments along the shaft of the long bones and particularly in the femoral neck. The GRF and the moments create the external forces acting on the bony skeleton. However, these forces are amplified by internal forces created by the muscles. Muscle forces are required for maintaining body balance and accomplishing movement tasks. Typically, internal forces created by the muscles exceed external forces acting on the body. As an example, for the hip joint to be stable during single-leg stance, the gluteal muscles acting at the proximal femur are needed to counterbalance the BW. Due to the small lever arms between the muscular attachments at the greater trochanter and the center of the hip joint, as compared with the larger lever arm between the center of the body mass and the hip joint, muscle forces of more than 2 times BW are required to properly maintain balance. During 1-legged stance, the contact forces in the hip joint resulting from the vector sum of the muscle forces and gravitational BW amount to 250% BW.2 Forces during slow-level ground walking are of the same magnitude but increase with increasing walking speed or during walking on stairs.3 For more demanding activities, the hip joint contact forces can dramatically increase. Activities involving load lifting or load carrying have shown to result in hip joint contact forces between 400% and 600% of BW.4 The highest forces were observed during load transfers involving squatting, 1-legged weight bearing, or uncoordinated movement like stumbling.
The mechanical signal experienced by the bone under loading is tissue strain. Tissue strain can be both friend and foe to bone structures. Strain is the relevant signal to trigger and control bone remodeling and build up strong bone tissue where needed. However, strain is also the major determinant for the initiation of failure at the tissue level by the induction of microcracks or catastrophic breakage under tensile or compressive strain. Bone strains in the area of the femoral neck have been reported to be between 500 and 2000 microstrains for day-to-day activities, like walking, stair climbing, or jumping.5 According to Frost's6 Mechanostat paradigm, this is also the amount of strain for which bone is in homeostasis. Strains in the neck region during daily activities are typically largest at the inferior aspect of the neck, which coincides with a larger thickness of the neck inferiorly compared with the superior region. Only during activities with increased hip flexion and active hip abductors, such as in stair ascent, does the superior aspect of the femoral experience strains at or above the level of bone homeostasis.5 For healthy cortical bone to fracture, strain needs to exceed 1% (10,000 microstrain), so that there is sufficient safety margin at the femoral neck to survive day-to-day activities without spontaneous fracture.7,8 On the other hand, because excessive loading or stumbling can increase loads at the hip by 3- to 5-fold, and aging results in bone tissue deterioration and cortical thinning, strains may start to approach the yield strain limits at the femoral neck region.
FEMORAL NECK FRACTURES
Although the hip joint contact forces reported in the literature easily exceed 500% BW and can be as high as 4000–5000 N of load during daily activities, spontaneous fractures typically do not occur in healthy individuals. Age, osteoporosis, or metastatic lesions, however, may weaken bone tissue to such an extent that spontaneous fractures occur.9 In these fracture situations, the loads exceed the strength of the proximal femur. Not only is the amount of load important but also the direction of loading. Fractures of the femoral neck occur during vertical impact along the femoral shaft axis typically while the hip is flexed. They can also occur during sideways impact on the greater trochanter or during torsional moments acting on the leg, generally when the hip is abducted, when loads are transferred to the hip joint.
Vertical loading of the hip joint creates a bending moment at the femoral neck, resulting in tension at its superior aspect and compression at the inferior part of the neck. It is typically the tensional strain at the superior aspect, which causes the initiation of failure. Consequently, the most common site of fracture initiation during vertical overloading is the superior lateral aspect in the subcapital region of the femoral neck.10
In contrast to vertical impacts, sideways impacts on the greater trochanter result in compressive loading of the femoral neck. The forces required to induce a fracture through a sideways fall are considerably lowered compared with those for vertical impact loading (Table 1). The resulting fractures during sideways falls are more often basicervical or even trochanteric.11
Femoral neck fractures in younger individuals require relevant amounts of fracture energy resulting in considerable levels of impact load. In addition to the level of impact energy, several factors have been identified that are associated with increasing the risk for sustaining a femoral neck fracture. Most importantly, the bone mineral density at the femoral neck is a reliable predictor of femoral neck fracture risk and can be easily obtained by bone density assessment at the proximal femur with dual x-ray absorptiometry or with quantitative computed tomography.12 Other factors shown to contribute to the strength of the femoral neck include bone geometry, cortical thickness, and cortical bone density.13 The risk for femoral neck fractures increases with increased hip axis length, increased neck length, reduced cortical thickness, and reduced cross-sectional area of the neck.13,14 Adding these relevant factors together in numerical models of bone strength enables an accurate prediction of fracture loads for in vitro experiments11,15,16 and prediction of fracture risks in longitudinal in vivo studies. Finite element analysis (FEA) has been shown to provide accurate estimates for mechanical responses to loading in the elastic region (bone stiffness)17 and in the plastic region (bone strength).18,19 These numerical models can also be clinically used to provide an accurate risk assessment for the occurrence of fractures at the hip.20,21 Although FEA provides the most accurate estimate for failure loads, combining bone density and bone geometry information based on quantitative computed tomography and dual x-ray absorptiometry assessment has similar accuracy and might be easier to employ in clinical routine.19
In addition to the structural properties of bone, the material properties of the bone matrix and particularly the cortical bone matrix affect the overall mechanical properties of bone.22 For the proximal femur, it has been shown that the porosity of the cortex can be assessed by bone biopsy23 and that the amount of porosity is a strong predictor for mechanical strength.23,24 Finally, the composition of the bone matrix itself contributes to the mechanical properties of bone tissue and to its overall strength. Material-level properties identified to contribute to bone strength include the accumulation of advanced glycation end products,24 size of bone crystals,25 and collagen content.26 Resistance of bone to loading and fracture strength is hierarchical and includes structural and material properties at multiple scales, many of which are accessible to diagnosis27 and provide reliable biomarkers for the prediction of bone fracture risk.
FIXATION OF FEMORAL NECK FRACTURES
Fractures at the hip joint tend to be located in the region of the femoral neck region, and the likelihood of a fracture at the trochanteric region increases with increasing age.28 Thus, hip fractures in younger adults are much more likely to occur in the neck region as compared with other regions of the hip. Based on the Pauwels classification, young adults more frequently present vertically oriented, more unstable, type III fractures that require accurate reduction and stable internal fixation.29
Although a considerable armamentarium is available for the fixation of femoral neck fractures, complication rates are still unacceptably high, and there is no agreement on the most suitable technique to treat these fractures. The goal of fracture fixation at the femoral neck is to provide sufficient mechanical stability until the fracture has healed. In particular for comminuted fractures, the stability of the osteosynthesis construct heavily relies on the mechanical stability of the implant. The available techniques for fracture fixation at the femoral neck can be divided in 4 different types of fixation: (1) cannulated screws (CS), (2) hip screw systems, (3) proximal femur plates (PFP), and (4) cephallomedullary nails (CMN).
In general, the mechanical stability of the osteosynthesis of a fractured bone is composed of the stability of the implant and the stability of the bone; hence, we talk about the stability of an osteosynthesis construct or bone–implant construct. The overall strength of an osteosynthesis construct at the femoral head is less than the strength of intact bone, and even after completion of bone healing, the healed osteosynthesis construct has not been completely restored to the mechanical strength of intact bone.30,31
Implant systems for femoral neck fracture fixation have been tested in biomechanical experiments on human cadaver specimens, various plastic analogue bones, and numerical models (FEA). Bone analogues provide a very reproducible and geometrically accurate template for creating fractures and placing osteosynthesis implants.32 However, biomechanical test results do not adequately represent real-bone behavior in terms of fracture patterns, construct stiffness, and strength.33 Numerical models can also not generally be adapted from one fracture model to another and need careful validation to make their findings trustworthy.
When treating femoral neck fractures with CS systems, the stability of the fixation construct has been shown to be affected by several factors related to the CS: screw type, screw number, screw thickness, and position, direction, and configuration of screws. Most biomechanical studies would agree that 3 screws should be used with diameters of larger than 6 mm.34,35 Adding a fourth screw has not consistently shown to provide biomechanical benefits, most likely due to the weakening of the lateral wall by the screw insertion holes.36 Screws should be placed in a triangular arrangement with an inverted triangle32 and spread apart as far as achievable under the anatomic circumstances.37,38 This configuration will provide the best axial and torsional stiffness and result in improved failure strength. The angulation of the screws seems to be of minor importance for the mechanical stability.34 The screws should apply compression of the fracture gap (compression screws) to minimize the remaining fracture gap and accelerate the healing process. Parallel placement of the screws enables sliding of the neck and impaction of the fracture gap in cases of bone resorption and bone remodeling during the healing process.39 Headless, fully threaded, self-compressing screw designs demonstrated some biomechanical advantages compared with conventional compression screws for the initial phase of stabilization.40 Similarly, the use of washers allows for higher compressional forces and thus increases the initial stability of the osteosynthesis construct.41 Washers further prevent the screw heads from penetrating the lateral cortex, which has shown to be beneficial in particular in osteoporotic bone.42
HIP SCREW SYSTEMS
Hip screw systems (sliding hip screw, dynamic hip screw, and compression hip screw) are extramedullary plate devices that penetrate the lateral cortex with a barrel that receives a lag screw. The fixation of the plate to the shaft is secured by at least 2 cortical compression screws. Longer plates are able to integrate more screws, which is not necessary in young healthy cortical bone. The barrel provides an angular stable support for the lag screw, securing the neck shaft angle. Finally, the lag screw enables retention and compaction of the fracture gap and also allows for sliding of the femoral neck along the lag screw axis during the healing process.
The lag screw should be placed with the tip of the screw in the center of the femoral head (center/center position) with a superior position within the head being more disadvantageous as compared with a slightly inferior position of the tip of the screw.43,44 Furthermore, the screw tip should be close to the subcortical bone of the femoral head, with an overall distance from the apex (tip–apex distance) of less than 25 mm.45 To secure the rotation of the femoral head, a cannulated antirotation (sometimes also called derotation) screw can be inserted before the placement of the lag screw. This antirotation screw secures the femoral head and avoids rotation of the head during insertion of the lag screw. Left in place, the antirotation screw provides additional mechanical stability for osteosynthesis and improves axial and rotational stability.46 Using a blade design in femoral neck fracture fixation, as compared with a screw design for the lag screw, has not shown biomechanical advantages.47,48
The angulation of the lag screw with respect to the shaft also has shown to affect the stability of the osteosynthesis construct. In Pauwels type III fractures, increasing the angulation of the lag screw from 135 to 150 degrees has shown to increase failure load and also reduce fracture displacement.49
PROXIMAL FEMORAL PLATES
Locking compression plates contoured to the proximal femur have been suggested for the fixation of femoral neck fractures. Compared with hip screw systems and CS, PFP constructs have shown biomechanical advantages regarding the initial stability of fracture fixation (ie, construct stiffness),50 additionally in plates designed with young patients in mind.51 However, clinically, PFPs are associated with considerable failure rates related to screw breakage or screw loosening. A biomechanical study identified inaccurate placement of the locking screws within the locking hole of the plate as a potential source of these failures.52 In response to the shortcomings of traditional PFP methods as treatments for femoral neck fractures, novel methods of plating these types of fractures have surfaced in recent years.
An emerging trend in locking PFP development involves combining fixed-angle technology with parallel telescoping screws, which allows controlled compression via a dynamic sliding mechanism. Constructs that incorporate these features have the potential to utilize the respective benefits of both CS and hip screw methods, by allowing compression of the fracture site while preserving cancellous bone and providing a rotationally stable healing environment.53,54 The Targon FN system (Aesculap B. Braun, Tuttlingen, Germany), developed in 2007, is a short femoral locking side plate that has the potential for up to 4 telescopic sliding screws that are placed in the femoral head and bridge the fracture site.55 A biomechanical study of the Targon FN demonstrated higher loads to failure and enhanced mechanical strength when compared with the sliding hip screw and CS methods.56 Clinical results for the Targon system are encouraging, with reduced nonunion rates,53 reduced risk for surgical revision,57 and less fracture subsidence.58 When examined specifically in young female athletes, the Targon FN implant offered a safe alternative for displaced stress fractures of the femoral neck.59
A retrospective study by Xiao et al60 presents an additional method of plating femoral neck fractures with a dynamic locking compression system that combines 3 parallel screws with a lateral pressure locking plate and interlocking tail caps. All 36 patients treated with this novel system achieved bony union within 4 months, and only 1 experienced mild screw back out.60 New developments in the field of fracture plating that encourage constrained motion in biomechanically advantageous directions, sometimes referred to as dynamic technology, may offer improvements in orthopedic care.61 Further research on these types of innovations for the femoral neck, such as novel plating constructs or telescoping screws combined with conventional locking plates, should be pursued.
CMNs are able to transfer the loads encountered at the femoral neck to the shaft of the femur, where the nail can have a long intramedullary support providing enhanced mechanical stability. In particular, for very unstable basicervical fractures, CMNs have demonstrated adequate mechanical stability. Biomechanical experiments have demonstrated that position and alignment of the femoral head can be preserved, physiological weight-bearing activities can be withstood, and supraphysiological loads can be endured in the case of adequate bone mineral density.31 At least in elderly patients, this biomechanical advantage can be transferred to the clinical application of CMNs for unstable basicervical fractures.62 For the rotational stability of the fixation, devices with 2 lag screws or with additional CS demonstrate significant advantages.63,64
Clinical recommendations for implant usage must often rely on findings from biomechanical testing. Clinical studies comparing outcome for different implant designs typically lack statistical power to identify differences in clinical outcome measured by functional outcome scores or by patient satisfaction. Thus, if reassuring findings from multiple biomechanical studies can be identified, this can lead to strong recommendations for or against the use of specific implant configurations. Biomechanical studies comparing different implants for femoral neck fracture fixation in nonelderly adults mainly focused on the 4 types of implants discussed above. However, it is not only the type of implant but also the type and remaining stability of the fracture that strongly affect the mechanical performance of an osteosynthesis construct.32,65 The remaining stability of the fracture determines the amount of load that the implant can share with the bone. For stable fractures, the implant can be more load sharing, whereas for unstable fractures, the implant needs to be load bearing.
When comparing CS with hip screw systems, hip screws with additional antirotation screw have consistently performed better biomechanically.66–68 CS systems consistently were less stable with respect to shear displacement and varus collapse of the head fragment; CS systems also failed significantly earlier. This observation of biomechanical advantages of hip screw systems compared with cancellous screw systems (Pauwels configuration) has been confirmed in clinical observations on displaced femoral neck fractures in young adults.69 Also, compared with PFP systems, hip screw systems have shown biomechanical advantages with respect to stabilization of interfragmentary shear movement and much more so with respect to load and energy to failure.70,71
Cephallomedullary implants have been used to stabilize unstable or comminuted basicervical fractures, or Pauwels type III fractures, and have shown similar mechanical strength compared with hip screw systems and superior strength compared with cancellous screw systems.72
In general, basicervical fractures or Pauwels type III fractures and comminuted fractures can be considered as mechanically unstable, requiring a load-bearing implant, such as hip screws, with antirotational screws or intramedullary nails. Subcapital or transcervical fracture patterns and noncomminuted fractures enable load sharing and can be securely fixed with CS or solitary hip screw systems without compromising fixation stability. However, despite all the biomechanical evidence that is available, the choice of implants for femoral neck fractures in young adults remains to be controversial.73
The role of PFP will have to be critically explored, in particular with the newer plate system combining the mechanical advantages of angular stability and multiple telescoping cancellous screws.
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