Functional Anatomy and Biomechanics

Carl W. Imhauser

The knee joint allows motion in all planes based on a complicated interaction between passive constraints including bones and soft tissues and the dynamic muscle stabilizers. The structure of these constraints affords the knee large rotations in the plane of progression (flexion/extension) while allowing more subtle motions in the remaining planes.¹

The knee joint must resist substantial multidirectional forces and moments during daily activities and sports. These loads arise from the contraction of muscles, the interaction of the foot with the ground, and the acceleration of the body. In fact, measurements from instrumented knee replacement indicate that compressive forces at the knee approach two and a half times body weight even during everyday activities such as walking.²

The shear forces generated between the foot and the ground are notable because they give rise to both forces and moments (i.e., torques) at the knee joint. The anterior-posterior force at the ground can be up to 20% of body weight.³,⁴ Because this anterior-posterior force acts at a distance from the knee, a moment is generated about the knee that acts to flex or extend it during functional activities. The anterior-posterior forces and flexion-extension moments can be effectively countered by the cruciates and the quadriceps and hamstring muscles.

The shear forces in the mediolateral direction are also substantial during daily activities and sports. For example, at the instant in walking where the foot strikes the ground, a lateral reaction force develops that can reach 5% of body weight.³,⁴ This force pushes the foot laterally, and because it acts at a distance from the knee, a moment is also generated about the knee that acts to rotate the tibia into valgus.⁵ Therefore, both a medial shear force and a varus moment must be generated within the knee to counteract the effect of the lateral component of the ground reaction force (Fig. 3.1).

The varus and valgus moments generated at the knee can be even more substantial during sporting activities such as running, side-stepping, and cutting⁶ and can serve as a predictor of injury risk. Elevated valgus moments during landing from a jump have been identified as a predictor of anterior cruciate ligament (ACL) injury risk in female athletes.⁷

MECHANISMS OF KNEE STABILITY

Frontal Plane Stability and the Role of Muscles

The bicondylar shape of the distal femur is well-suited to resisting the varus and valgus moments that act on the knee during daily activities and sports. When the knee compresses due to muscle contraction and/or the ground reaction force, an applied moment in the frontal plane will cause redistribution of the contact force acting on the medial and lateral condyles.⁸,⁹ The ability of the articular surfaces to resist the applied moment is proportional to the magnitude of the contact force and its distance from the joint center. This highlights the importance of the muscles in generating contact forces that increase the ability of the articular surfaces to resist frontal plane moments.

The articular surfaces can rapidly generate resistive frontal plane moments with minimal varus or valgus angulation (≤2°) in the presence of a compressive load (Fig. 3.2). The small amounts of angulation are due to compression of the cartilage and menisci. Condylar lift-off will occur when the applied moment has exceeded the capacity of the articulating surfaces to generate an internal counteracting moment. This instant is characterized by a sudden flattening in the slope of the moment-angular displacement response (see Fig. 3.2). As the knee continues to rotate into varus or valgus, the ligaments will engage and generate forces that cause the resistive moment to once again increase (see Fig. 3.2).

Stability Provided by the Articulating Surfaces

The shapes of the articulating surfaces of the knee provide an involuntary mechanism to provide joint stability.¹ The joint will unconsciously move to a contact position that can counter the direction of an applied force within an envelope defined by the size and shape of the articular surfaces (Fig. 3.3). The bounds of this envelope can be visualized by drawing lines that are perpendicular to the articulating surfaces at the extreme ends of the joint (Fig. 3.3A).¹ The directions of these bounding lines encompass all orientations of reaction forces that can be generated by the articulating surfaces. Thus, the articular surfaces may accommodate changes in the direction of the applied

loads through compression of the joint alone and by adjusting the relative position of the joint surfaces (Fig. 3.3 B,C).

Figure 3.1. A lateral shear force is generated between the foot and the ground during the heel strike phase of the gait cycle and more so during athletic activities such as running, cutting, and jump-landing. This lateral shear force creates a valgus moment at the knee joint. The externally applied valgus moment and lateral shear force must be balanced within the knee by a medial shear force and varus moment, respectively (shown with dashed lines).

Figure 3.2. Increasing the contact force via an applied axial load increases the ability of the bicondylar articular surfaces of the knee to resist frontal plane moments. The instant of condylar lift-off is identified by a sudden decrease in the slope of the moment-angular displacement curve in varus and valgus. The torque level at which this occurs increases with increasing axial load. (Redrawn from Markolf KL, Bargar WL, Shoemaker SC, et al. The role of joint load in knee stability. J Bone Joint Surg Am. 1981;63[4]:570-585, with permission.)

Figure 3.3. Assume that the flexion angle remains fixed. A. The articular surfaces are capable of generating a reaction force within the range of orientations defined by perpendiculars to the joint surface at either side of the joint (dotted line).¹ The sweep of achievable orientations of reaction forces that can be generated by the curvature of the articular surface is encompassed in gray. B. The reaction force that joint contact is required to generate is shown. At its current position, contact between the articular surfaces alone is unable to generate the required reaction force at the articular surface because the perpendicular to the joint surface where contact occurs is not aligned with the required reaction force. C. To generate the required reaction force, the femur will slide from the position shown in B to the position shown in C. Here, the perpendicular to the articulating surface aligns with the required reaction force. Therefore, no other stabilizing mechanism besides compression of the joint is needed to keep the joint stable. (Adapted from Burstein AH, Wright TM. Fundamentals of Orthopaedic Biomechanics. Baltimore, MD: Williams and Wilkins; 1994.)

Figure 3.4. The conforming geometry of the knee joint in the frontal plane is well-suited to producing both medially and laterally directed contact forces. The required medially (A) and laterally (B) directed reaction forces are generated by a small lateral and medial shift of the tibia, respectively. The dotted lines are perpendicular to the joint surfaces. They identify the range of orientations of reaction forces that can be generated by the curvature of the articular surfaces.

Because the knee is conforming in the coronal plane, it takes only small shifts in the mediolateral position of the knee to produce mediolateral-directed contact forces (Fig. 3.4).

Similarly, the medial compartment of the knee will resist small changes in anterior-posterior-oriented forces in the sagittal plane due to its concave shape. The lateral compartment, however, based on its convex shape in the sagittal plane and lateral orientation in the frontal plane, cannot effectively resist anteriorly and medially directed external forces, respectively. Thus, it requires additional stabilizing mechanisms to accommodate these applied forces including the menisci and the ACL.

Figure 3.5. A. The load-displacement response of the ACL¹⁰ includes a toe region (circled in red). B. In the toe region, the ligament gradually generates force for the first few millimeters of the load-displacement response as shown between points 1 and 2. C. The toe region of the ligament allows for some relative movement of the articular surfaces between positions 1 and 2. The toe region combined with a ligament’s slack length enables motion of the contact point along the curvature of the joint. This permits the articular stabilizing mechanism to engage. (Adapted from Woo SL, Hollis JM, Adams DJ, et al. Tensile properties of the human femur-anterior cruciate ligament-tibia complex. The effects of specimen age and orientation. Am J Sports Med. 1991;19[3]:217-225; Burstein AH, Wright TM. Fundamentals of Orthopaedic Biomechanics. Baltimore, MD: Williams and Wilkins; 1994, with permission.)

Stability Provided by Ligaments

The forces generated by ligaments also contribute to joint stability. However, when acting to restrain the joint, they must not interfere with the normal range of motion of the knee. This is accomplished because their line of action passes close to the rotation axis of the joint.¹ This makes it impossible to generate large resistive moments against flexion and extension because the distance between rotation axis and the line of action of the force generated by the ligament remains small through flexion. When this principle is not adhered to during ligament reconstruction, for example, it can lead to clinical complications including flexion contracture and overcon-straint of the joint.

Ligaments must also allow some relative movement of the articular surfaces for the articular stabilizing mechanism discussed earlier to be effective.¹ Two characteristic features of ligaments facilitate this: their slack length and toe region. Slack length describes the length at which the ligament begins to carry load. The toe region refers to the initial portion of the load-displacement response of the ligament where load builds slowly as it is stretched. These two features allow the ligament to gradually generate force for the first few millimeters of relative movement of the joint surfaces. This enables the articular surfaces to move to a suitable contact point along their curvature to generate the required reaction force (Fig. 3.5).

If there is no suitable point along the curvature of the articular surfaces that can generate the required reaction force, the bones displace further apart. As the bones displace further, the ligaments will generate force more rapidly as they stiffen.¹⁰ Thus, they will make a larger contribution to resisting loads applied to the joint.

Finally, a ligament should be oriented as closely as possible in the direction of the applied force that it must resist.¹ Thus, the collateral ligaments are appropriately oriented to resist frontal plane moments based on their mediolateral distance from the joint center and proximal-distal orientation. Similarly,
the ACL has an oblique orientation with origin on the medial aspect of the lateral femoral condyle and insertion on the anterior aspect of the tibia. Thus, its orientation reveals its function in resisting forces that act to push the tibia anteriorly and medially.¹¹,¹²,¹³ Its small offset relative to the center of the joint also makes it a secondary stabilizer against varus and valgus.¹³

Knee Instability

Situations arise where the stabilizing mechanisms described earlier fail; that is, they cannot effectively counteract the magnitude and direction of the externally applied loads. This will lead to instances where the knee joint feels unstable. These instability events, or episodes where the knee gives way, compromise function and can lead to cartilage and meniscal damage.¹⁴,¹⁵ Relief from knee instability is a primary reason why patients seek treatment for ACL injury.

Instability has been defined as large changes in the position of articular contact with small changes in applied load.¹ A stable joint is one where the articular surfaces glide smoothly relative to each other without sudden jumps or abrupt changes in their relative position. Instability can occur when a stabilizing mechanism is absent (e.g., ACL insufficiency) or when existing stabilizing mechanisms are incapable of providing an appropriate resistive force. This might happen when a movement occurs too quickly for the muscles to generate the required force such as in a quick cutting maneuver while playing sports. Stabilizing mechanisms might also fail in combination such as muscle weakness combined with a ligament injury leading to a giving-way episode in an ACL-deficient patient.

Episodes of giving way in ACL-deficient individuals have been serendipitously recorded in the gait laboratory.¹⁶,¹⁷ This provides a unique opportunity to quantify the time sequence of loads and motions experienced by the knee over the course of this event. A moderate giving-way episode suffered by an individual while descending stairs revealed that an abnormally high applied valgus moment on the tibia preceded the instability event.¹⁶ The instability event was associated with large and rapid anterior translation of the tibia relative to the femur and increased knee flexion compared to successfully performed trials. Another giving-way episode recorded during a cutting activity was preceded by reductions in the applied varus moment.¹⁷ This was followed by a transition from an external rotation moment to an internal rotation moment acting on the tibia. These abnormal moments were followed by increased internal rotation of the tibia relative to the femur and increased flexion of the knee compared to successful cutting trials.

These unique data bear the hallmark signs of the clinical pivot shift, which is also generated by a complex combination of applied valgus and internal rotation moments combined with knee flexion.¹⁴,¹⁵ This reveals the major function of the ACL as a restraint to anterolateral subluxation of the tibia against moments that tend to abduct and internally rotate the tibia.¹⁸,¹⁹

Interestingly, these giving-way episodes were rapid. Therefore, the restraint provided by muscle forces was likely low because the instability occurred too quickly to be resisted by a voluntary muscle response.²⁰ This indicates that the remaining articular and ligamentous passive stabilizers in the ACL-deficient knee were unable to maintain the position of the knee under these multiplanar loading conditions.

A closer inspection of the interaction of the articular and ligamentous stabilizers against the applied valgus and internal rotation moments provides insight into a potential cause of the instability event. This example ties together our discussion of stabilizing mechanisms of the knee focusing on the role of the ACL. First, the valgus moment acting on the tibia gives rise to a force on the convex lateral compartment that has both anterior and compressive components. Then, an internal rotation moment acts to drive the lateral compartment of the tibia anteriorly. The ACL is oriented to resist these anterior forces. However, in the ACL-deficient knee, this stabilizing mechanism is not available, and the lateral compartment of the tibia will be subject to an abrupt anterior translation relative to the femur. Only when secondary stabilizers such as the lateral meniscus engage to provide the needed posterior reaction force on the tibia will the lateral compartment cease translating anteriorly.

ANATOMY AND BIOMECHANICAL FUNCTION OF THE ANTERIOR CRUCIATE LIGAMENT

The ACL has classically been divided into two functional bundles: an anteromedial and posterolateral portion. This terminology stems from the relative location of its tibial insertion and was described initially by Abbot and colleagues²¹ and later by Girgis et al.²² This description has proven convenient in grossly demarcating function of major portions of the ACL; however, these bundles may not be easily identifiable in all individuals.²³,²⁴ Assessment of the material properties of these bundles indicated that the elastic modulus and the maximum stress of the anteromedial bundle is about 45% greater and at least twice as high as that of the posterolateral bundle, respectively.²⁵ Taken to its logical extreme, the ACL could be further subdivided.²⁶,²⁷ In fact, the ACL acts as a continuum of fibers²⁸ and has been divided into as many as 20 smaller bundles each about 1 mm in diameter.²⁷

The ACL appears flat in extension and fans out about 10 to 12 mm distal to the femoral attachment,²² thus forming a flat hourglass shape along its length.²⁹ This specialized shape enables the ACL to achieve an anterior and medial orientation to allow it to effectively resist loads in these directions while also accommodating the femoral notch.

Early observations by Groves³⁰ indicated that the bulk of the ACL was taut in extension, and this was later corroborated by Girgis et al.²² The ACL twists on itself as the orientation of the femoral attachment changes from more vertical in extension to more horizontal in flexion.²² The tissue inserting most proximal and anterior into the femur remains tight on visual inspection while the tissue inserting more distally into the femur slackens²² (Fig. 3.6).

Pioneering work by Artmann and Wirth³¹ quantified the relative changes in distance between the tibial insertion and various points on the femoral insertion through flexion by monitoring the change in length of cables spanning the tibial and femoral insertions. They reported that anteroproximal locations on the femur were isometric. This approach has been reproduced by others with similar results.³² However, such studies rely on measurement of the straight-line distance between tibia and femur when in fact the individual ACL fibers wrap around one another and do not follow a linear path.²⁷ This was recognized
and elegantly addressed by Amis and Dawkins²³ by traversing wires through the tibial insertion along fiber bundles comprising the anteromedial bundle, posterolateral bundle, and an intermediate bundle between them. The wires were fixed in the femoral insertion of each bundle. The length changes of these wires were then measured while the knee was passively flexed from full extension to 130 degrees flexion. None of the three fiber bundles was isometric through flexion²³ (Fig. 3.7). However, the anteromedial bundle did experience smaller changes in length (˜3.5 mm) than the posterolateral bundle (˜5 mm) from full extension to 130 degrees flexion.

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