August 2017

Biomechanics of femoral neck fractures in runners

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Factors related to joint loading during running contribute to the development of stress fractures in the femoral neck and other lower extremity structures. Research suggests these loading parameters can be reduced by altering running foot strike pattern, stride length, speed, and step rate.

By Mark Riebel, PT, DSc, OCS, SCS

Running is currently one of the most popular competitive and recreational activities in the world. Due to its accessibility, relatively low cost, minimal equipment requirements, and a myriad of health- and fitness-related benefits, millions of adults worldwide regularly engage annually in some degree of running.1,2 Running, however, can result in injury, particularly to the lower extremities. Lower extremity injury prevalence associated with running has been reported at 68%, with incidence between 19.4% and 79.3%.1,3

Due to the atraumatic nature of running, the majority of lower extremity injuries reported are overuse injuries. In bone, these are termed bone stress injuries. As part of the normal loading and unloading of bone that occurs during running, osteoclastic cells are stimulated in the presence of stress, and bone is resorbed at the site of the stress. Once the stress is removed, osteoblastic cells then rebuild and reinforce the areas to which stress was applied. Because the upwardly directed ground reaction force (GRF) during running may be at least 150% greater than during walking, an imbalance in the breakdown and buildup of bone material may develop.4 When the osteoclastic actions due to stress outweigh the osteoblastic actions, a stress injury can occur and structural weakness may be present.

Bone stress pathology progresses from stress reaction to stress fracture and finally to complete bone fracture.5 The majority of lower extremity stress injuries occur in the metatarsals and tibia, and are deemed low risk because they are associated with a low incidence of complications and favorable natural history.

Though only .6% to 12% of lower extremity stress injuries involve the proximal femur, such injuries can be more serious than those involving the metatarsals or tibia.

Though representing only .6% to 12% of stress injuries in the lower extremities, a stress injury of the proximal femur is of greater concern.6,7 A proximal femur stress injury, particularly to the femoral neck, is high risk in nature due to the propensity for delayed union, nonunion, the need for surgical fixation, or combinations of these.8 Femoral neck stress fractures are of particular importance in military populations, as incidence can lead to significant cost, convalescence, and morbidity.9 Proper understanding of the development of femoral neck stress fractures in runners is of great importance for clinicians.

Hip joint anatomy

The coxofemoral joint, or hip joint, consists of the articulation between the acetabulum of the pelvis and the head of the femur. These bones form a diarthrodial, multiaxial joint with 3º of freedom of motion. These motions consist of internal and external rotation in the transverse plane about a longitudinal axis, flexion and extension in the sagittal plane about a horizontal axis, and abduction and adduction about an anterior to posterior axis. The ball-and-socket joint of the hip functions primarily to support the weight of the head, trunk, and upper extremities during weightbearing activities.

Figure 1. Trabecular systems of the femoral head and neck. Two major systems (the medial compressive and lateral tensile) show the primary transmission of forces. Three minor systems attenuate the remaining forces.

The acetabulum is comprised of the three innominate bones—the pubis, the ilium, and the ischium. These bones are typically fully ossified when individuals are aged between 20 and 25 years.10 The superior portion of the acetabulum, the lunate surface, is covered with hyaline cartilage and is the only portion of the acetabulum that articulates with the femur. The acetabulum is oriented laterally with some degree of anteversion, a combination of inferior and anterior angulation. Angles of acetabular anteversion have been reported from 18.5° to 40° and vary by sex, with women typically having a greater degree of anteversion than men.11 The fibrocartilaginous acetabular labrum surrounds the periphery of the acetabulum, functionally deepening it and improving joint hip stability. Evidence suggests the acetabular labrum has no loadbearing function at the hip joint but rather serves a proprioceptive role at the hip and helps to protect the rim of the acetabulum.11

The femur is the longest and heaviest bone in the body, and its head forms the distal articulation of the hip joint. The head of the femur may range in shape from a hemisphere to two-thirds of a hemisphere and is covered almost entirely in hyaline cartilage. At the most medial point of the head lies the fovea capitis, the attachment site for the ligament of the head of the femur, which attaches to the peripheral edge of the acetabular notch. This ligament offers little in actual hip stability, primarily functioning as a synovial fold to conduct a small branch of the obturator artery.10 The femoral head is joined to the shaft of the femur by the femoral neck, a trapezoidal portion of the bone approximately 5 cm in length whose broader base is continuous with the shaft.

Two primary angulations affect the orientation of the femur relative to the acetabulum: the angle of inclination and the angle of torsion. Abnormalities in these angulations can substantially alter hip joint stability and weightbearing biomechanics, as well as muscle biomechanics.11

The angle of inclination of the femur occurs in the frontal plane between an axis that runs through the femoral head and neck and the longitudinal axis of the shaft. The typical range for the angle of inclination is between 115° to 140° for an unimpaired adult. The angle averages around 135° in infancy and childhood and gradually decreases throughout the lifespan, averaging around 120° in a normal older adult.10 A pathologic increase in this angle is termed coxa vara, while a pathologic decrease is called coxa valga.

The angle of torsion of the femur occurs in the transverse plane between an axis through the femoral head and neck and an axis through the femoral condyles. The angle of torsion for an unimpaired adult normally lies between 10° and 20°, though this has been shown in some populations to lie around 35°.11 The angle of torsion typically decreases across the lifespan from an estimated angle of 40° in infancy. A pathologic increase in the angle of torsion is termed anteversion, while a pathologic decrease is called retroversion.

Abnormalities of all the aforementioned angulations may also have an effect on the joints, both proximally and distally. The hip joint is reinforced by its dense fibrous capsule and three reinforcing capsular ligaments—the iliofemoral and pubofemoral ligaments anteriorly, and the ischiofemoral ligament posteriorly. These ligaments provide significant stability to the hip joint, though there is evidence that the anterior ligaments are stronger than the ischiofemoral ligament.12 The combination of the capsule and ligaments allow for little or no joint distraction at the hip under traction. The close-packed position for the hip is in extension with slight abduction and internal rotation. The open-packed position is in slight flexion, abduction, and relatively neutral rotation.11

Weightbearing biomechanics

During upright weightbearing tasks such as walking and running, the weight of the head, trunk, and upper extremities is transmitted through the pelvis to the head of the femur, while the GRF travels up the femoral shaft. Because of the offset of the head of the femur from the shaft, these two opposing forces create a force couple with a moment arm equal to the distance between the force of body weight through the head and the GRF through the shaft. These forces cause a bending moment across the femoral neck,13 which creates a compressive force on the inferior aspect of the femoral neck and a tensile force on the superior aspect. Structural resistance to these shear forces, which can be substantial, comes in the form of cortical bone and trabecular systems. The two major (medial compressive and lateral tensile) and three minor trabecular systems comprise the internal resistance to shear forces (Figure 1). These trabecular areas are particularly strong in regions where they overlap.

However, a small area of the femoral neck has no overlapping trabeculae and is therefore more susceptible to failure. This area, known as the zone of weakness, may be compromised with exposure to excessive force or when decreased bony integrity makes the femoral neck more vulnerable to typical forces.

The trabecular systems are reinforced by the cortical bone, which may bear more than 50% of the load placed on the proximal femur.14 Ground reaction force and the weight of the head, trunk, and upper extremities are both also attenuated through various muscle actions around the hip joint. Due to the structure of the hip, actions of the muscles that directly influence the joint change in
accordance with the joint’s position and motion.

Running biomechanics

Normal walking gait consists of alternating cycles of single-limb stance and double-limb stance. The transition from walking to running occurs when the double-limb stance phase of walking gait gives way to a swing phase in which neither foot is in contact with the ground. As such, the running cycle is divided into two phases: stance and swing. Stance is subdivided into foot strike, midsupport, and takeoff; swing is subdivided into follow-through, forward swing, and foot descent. Kinematically, the hip joint moves through a total range of approximately 85° in the sagittal plane during this cycle.15 Typical sagittal plane range of motion values for the hip during each phase of running are shown in Table 1.

With respect to kinetics in this plane, moments are generated by action of the hip flexor and hip extensor muscles. The hip extensors increase in activity during forward swing as they begin to contract, while the hip decreases in flexion with a simultaneously extending knee. They are the dominant muscles just prior to foot strike through the first half of the stance, as they contract concentrically to propel the body forward and move the hip into extension. The hip flexors are dominant beginning in the second half of the stance, as they help to decelerate the extending thigh in preparation for the swing phase. They remain dominant through the first half of swing.16

Some hip motion also takes place in the frontal plane. Because the body’s center of mass must be directly over the single-support foot to maintain balance, the pelvis shifts laterally during stance phase and the GRF is directed medially to the hip joint. This creates adduction at the hip that is partially countered by the hip abductor muscles.17 Due to gravitational and acceleration loads that are greater than the abduction moment produced by the hip abductors, functional limb varus of 10° greater than walking may occur.18 Functional limb varus is defined as the angle between the bisection of the leg and the floor, and originates at the hip with adduction of the joint.15 Though hip range of motion in the frontal plane is relatively small compared with the sagittal plane, medially directed shear forces of 3.75 times body weight can occur at the femoral neck during stance phase.19 During terminal stance and swing, the hip abductors contribute to hip extension and abduction, respectively.

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Biomechanics and fracture risk

Presumably due to the low incidence of injury, there is a dearth of information on risk factors specifically for femoral neck stress fractures. In Air Force basic trainees, Kupferer et al found that prevalence of femoral neck stress fracture was significantly higher in women than men.9 When assessing trainees’ physical fitness levels, both men and women with slower 1.5-mile run times demonstrated increased risk of femoral neck stress fracture (odds ratios of 3.24 and 1.32, respectively) than trainees with faster run times. Greater numbers of push-ups and sit-ups performed on physical fitness assessment demonstrated a protective effect for stress injury in female trainees only (odds ratios of .55 and .62, respectively).

The authors hypothesized these factors affect risk levels for stress injury potentially because they indicate individuals who may lack the physiologic reserve to recover from repetitive impact activity or who have poor overall physical conditioning and stamina.These findings may not be as applicable to individuals outside the military due to the unique demands on military trainees, who cannot self-select modes of exercise to avoid more serious injury.

When considering all lower extremity stress fractures associated with running, stress fractures tend to occur more commonly in those increasing volume or intensity without appropriate levels of rest and recovery.5 Increased risk has also been associated with amenorrhea for at least six months, poor physical fitness, low bone mineral density, low dietary calcium and vitamin D intake, and smoking.5,20-23 Cosman et al found the incidence of stress fracture in US Military Academy cadets was elevated in those with a smaller tibial cortical diameter and smaller femoral neck diameter.24

With regard to biomechanical factors associated with running, athletes demonstrating higher vertical GRF loading rates and peak acceleration at foot strike have a higher incidence of tibial stress fractures.25 Foot strike during running may influence this.

Two primary categories are used to classify the type of foot strike pattern at the beginning of stance during running: rearfoot strike (RFS) and forefoot strike (FFS).26 The majority of habitually shod runners employ a RFS pattern.27

Several studies have shown that RFS runners demonstrate a double peak in the vertical component of the GRF, while FFS runners demonstrate a single peak.28,29 The first spike in vertical GRF seen with RFS running occurs when the heel contacts the ground, and is referred to as the impact peak.30 This occurs within the first 20 milliseconds of stance.

Evidence is equivocal as to which strike pattern results in a greater overall magnitude of force imparted on the lower extremities.26,31 However, runners employing a RFS pattern consistently show significantly greater average and instantaneous rates of loading than those with a FFS pattern, likely due to the presence of the impact peak.32

These loading rates may be of greater importance than the magnitude of peak force for the development of stress fracture. Zadpoor et al conducted a systematic review and meta-analysis of studies exploring the relationship of lower extremity stress fractures to both magnitude and rate of loading. They found that only rate of loading was associated with lower extremity stress fracture.32 It may therefore be possible that reducing the rate of loading during running, by altering technique or training parameters, would reduce the incidence of lower extremity stress fracture.

Although it is impractical to obtain in vivo measurements of forces on the hip joint during running without an instrumented hip prosthesis, Rooney used a biomechanical model to show that rate of loading and peak force on the hip are reduced when habitual RFS runners adopt a FFS pattern.26 Biomechanical modeling from Boyer and Derrick also showed habitual FFS runners experienced lesser hip joint loads compared with habitual RFS runners.33 It has also been shown that habitual RFS runners who are instructed in FFS running demonstrate similar biomech­anics as runners with a habitual FFS.34 This may be an important factor in reducing the potential for femoral neck stress fractures.

Although not specifically directed at stress fractures of the femoral neck, some research suggests alterations in running mechanics can help decrease the potential for stress fracture. Edwards et al investigated the effect of altering stride length on forces through the lower extremity using a probabilistic model of tibial stress fractures in 10 male runners.35 If stride length is shortened, there is a subsequent decrease in strain experienced by the lower extremities, but the number of impact cycles, or cumulative loading, may be increased due to the shorter distance traveled per stride.

As discussed previously, the cyclical loads experienced by the lower extremities during running will cause some natural breakdown of osseous structures, and this subsequent increase in cumulative loading could lead to too great of a total stress on the lower extremities and negate any potential benefits in decreased strain. To examine this tradeoff in decreased strain versus an increased number of impact cycles, the researchers analyzed the runners at constant velocity with both a preferred stride length and at a stride length 10% shorter than preferred. Using a tibial stress fracture model that incorporated bone fatigue, bone repair, and bone adaptation, while taking into account lower extremity strain and loading cycles, the authors calculated that reducing a runner’s stride length by 10% may reduce the chance of tibial stress fracture by 3% to 6%, indicating the magnitude of strain is potentially more influential than the number of impact cycles.35

Running speed may also be of interest with regard to stress fracture risk. A positive relationship between GRF and running speed has been demonstrated, indicating that running at faster speeds will lead to greater GRF experienced by the body.36 If a positive relationship between running speed and stride length is assumed, these faster speeds will lead to longer strides and fewer foot strikes for a given distance.36  One may then infer that running at slower speeds would lead to lower GRFs with an increase in the number of foot strikes for a given distance. This results in a similar tradeoff in injury potential, with lower peak stresses at slower speeds but a greater amount of cumulative stress.

Edwards et al examined this negative association between magnitude of stress and number of impact cycles by recording 10 individuals running at three speeds (2.5 m/s, 3.5 m/s, and 4.5 m/s) in random order while wearing anatomical markers.36 Force plate and video data were then input into a computer musculoskeletal model that calculated joint force, bone strain, and tibial stress fracture risk. Their data indicated that decreasing running speed from 4.5 m/s to 3.5 m/s reduced absolute peak principle strain and decreased the probability of stress fracture by 7%. Further reducing speed to 2.5 m/s reduced stress fracture risk by another 10%. These data corroborate the researchers’ earlier findings that the magnitude of strain may be more influential on stress fracture develop­ment than the number of impact cycles.36

To investigate the variable of step rate on lower extremity loading parameters, Heiderscheit et al assessed 45 healthy recreational runners as they ran at a self-selected moderate pace and step rate.37 Visual and GRF data from an instrumented treadmill were then collected at self-selected, ± 5%, and ± 10% step rates while keeping the self-selected speed constant. Data were entered into a computerized musculoskeletal model along with each participant’s mass, height, and segmental lengths.

The results showed that as step rate increased, step length, braking impulse, and force absorbed at the knee all significantly decreased relative to the preferred step rate. Peak GRF and force absorbed at the hip only decreased at plus 10% of preferred step rate. Although the authors acknowledged the tradeoff between decreased magnitude of stress and increased loading cycles at faster step rates, they concluded the decreased joint loading may outweigh any detriments of increased loading cycles and be beneficial in the treatment and prevention of lower extremity running injuries.37

Although these data may seem to conflict with the findings of the Zadpoor et al systematic review and meta-analysis that concluded the magnitude of GRF was not significant in the development of stress fractures,32 two aspects of these three studies must be considered. First, these three studies used a predictive model rather than assessing runners in a prospective manner that quantified actual injury. Additionally, GRF does not represent all force exerted on the lower extremity, but rather is one of the most easily measured variables, and therefore one of the most frequently collected. As such, it is possible that individual differences in bone strength and geometry combined with varying aspects of GRF more completely explain stress fracture development than GRF alone.32

Conclusion

Because of the low incidence of femoral neck stress fracture and the relative difficulty of obtaining biomechanical data specifically at the hip joint and femoral neck, there are sparse data directly pertaining to this condition. However, due to the potential for morbidity and high medical costs often associated with a femoral neck stress fracture, clinicians must be familiar with the etiology of the disorder and ways to modify certain parameters that may influence its development in runners.

Among other factors, higher rates of loading in the lower extremities while running, and potentially the magnitude of this loading, affect the development of femoral neck and other lower extremity stress fractures. Some studies have shown that these loading parameters can be reduced by altering running foot strike pattern, stride length, speed, and step rate. These may be important factors when designing rehabilitation or injury prevention programs for runners. Areas for future research include more studies directed at the hip joint and femoral neck, and more prospective studies focused on femoral neck stress fractures in runners.

Mark Riebel, PT, DSc, OCS, SCS, is a physical therapist in the US Navy currently stationed at Naval Health Clinic Quantico in Virginia. 

Disclosure: The opinions or assertions contained herein are the private views of the author and are not to be construed as official or as reflecting the views of the US Navy or the Department of Defense.

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