Biomechanics of the anterior cruciate ligament and implications for surgical reconstruction

Description of motion about the knee implies sagittal plane motion and rotational components of the femorotibial joint. It is best described by the “instantaneous centre of motion” theory, which suggests that motion occurs about any instant point that acts as the centre of rotation and therefore does not move [5]. Mapping of successive instant centres throughout the range of motion, however, does not generate a single motion axis. It rather allows determination of an “instant centre pathway”, which is shaped semicircularly and is located close to the joint surface in flexion and distant to the joint line in extension.

Relative surface motion between the tibia and the femur during flexion and extension occurring about the centre pathway appears to be that of gliding and rolling, the ratio of which is determined by the geometry of the articulating joint surfaces [6]. It changes from flexion to extension, with rolling of the femorotibial joint predominating near extension and gliding predominating as the knee is flexed [7]. The ratio of gliding and rolling, however, differs between the medial and the lateral condyle. While the medial femoral condyle sagittally is composed of two functional facets, the radius of which decreases from anterior to posterior, the lateral femoral condyle usually is composed of a single circular facet. This morphological principle implicates that rolling of the femur relative to the tibia would rather occur within the lateral compartment, whereas gliding would predominantly occur within the medial femorotibial compartment [8‐10].

The combined movement of the medial and the lateral compartment finally equates to external rotation of the tibia with extension and internal tibial rotation with flexion of the knee joint. The axis of external and internal rotation, which is primarily determined by the geometry of the articulating surfaces, generally passes through the medial intercondylar tubercle of the tibial plateau [11]. Due to the incongruency in the radius of the medial and lateral condyle, the total range of tibial rotation depends upon the degree of knee flexion and sharply decreases near extension. Although surface geometry has been identified as a major guide of femorotibial joint motion, the complex kinematics of the knee under physiological loading in vivo have yet to be completely investigated. It appears to be obvious that the constraints provided by the femorotibial joint surface are insufficient for functional knee stability, though a combination of resisting structures such as the menisci, the joint capsule, the muscles, and the intra- and extraarticular ligaments account for both functional range of motion and joint stability under loading conditions.

As the knee derives its stability from ligament structures rather than from its osteochondral surfaces, disruption of any of the supporting ligamentous structures will likely alter the overall motion characteristics of the knee. Moreover, damage to the central ligamentous column (i.e., the anterior and posterior cruciate ligaments), is found to physiologically stress other joint structures due to the pathologic shift in the sagittal and longitudinal axis of rotation [7]. The ligamentous structures of the knee provide stability as they compensate for tensile stresses acting in line with the axis of collagen fibres. As motion of the knee is mechanically complex, with several simultaneously changing axes of rotation, forces are not restricted by one ligament acting alone. Rather, they are restricted by a combination of ligament fibre bundles that are being recruited depending on both the flexion angle of the knee and the loading condition.

The amount that a specific structure contributes to the absorption of deforming forces has been described by the concept of primary and secondary stabilisers of the knee joint [12]. The cruciate ligaments, by guiding the motion of the femoral and tibial joint surfaces past one another, have been clearly identified as the primary stabilisers of anteroposterior translation when the knee is flexed. Studies have shown that the ACL provided more than 80% of anterior restraining force from 30° to 90° of knee flexion, while other ligamentous structures such as the medial joint capsule, the iliotibial tract, and the medial and lateral collateral ligaments provided no relevant secondary restraint to this motion [12]. Markolf et al. were able to demonstrate that anterior tibial translation was greatest between 20° and 45° of knee flexion, while beyond 90° of flexion both the superficial and the deep medial collateral ligaments secondarily contributed to anteroposterior stability [13].

Secondly, the ACL forces the tibia to internally rotate during anterior tibial translation, indicating that the ACL primarily restrains internal rotational moments during anteroposterior translation [14]. After sectioning the ACL, a significant increase of internal tibial rotation has been reported while subsequent sectioning of the collateral ligaments produced no progressive increase in internal tibial rotation near extension [15, 16]. With the knee flexed, anterolateral and posteromedial capsular structures are recruited during internal rotation as the ACL slackens and the posterior cruciate ligament tightens.

The ACL is a ligamentous structure composed of dense connective tissue containing parallel rows of fibroblasts and type I collagen, which has been shown to be the predominant structural component [17]. It originates from the posterior medial aspect of the lateral femoral condyle and inserts to the anterior and lateral aspect of the medial tibial spine [18]. The area of origin and insertion of the ACL is reported to average 113 and 136 mm², respectively. The cross-sectional area at midsubstance varies between 36 and 44 mm², while the length of the anterior and posterior aspect of the ligament is reported to vary between 22 and 41 mm [19‐22].

The ACL does not function as a simple tube of fibres with a constant tension, but rather consists of fibre groups that are subjected to episodes of lengthening and slackening throughout the range of motion (Fig. 1) [23]. This has advocated the functional subdivision of the ACL into an anteromedial and a posterolateral bundle, although histological investigations suggested a subdivision of fibre bundles to be rather arbitrary [24, 25]. According to functional observations, the fibres of the anteromedial bundle originate most anteriorly on the femoral side and insert anteriorly and medially at the tibial attachment. The fibres of the posterolateral bundle run from the posterior part of the femoral attachment to the posterior and lateral aspect of the tibial ACL footprint (Fig. 2). With the knee in full extension, the fibres of the smaller anteromedial and the larger posterolateral bundle run parallel, while during knee flexion the anteromedial fibres tighten and twist around the slackened posterolateral fibres, leaving the anteromedial fibres as the primary restraint against anterior tibial displacement at 90° of knee flexion. With both internal and external rotation, the ACL tightens so that it may operate as a major restraint against rotational moments acting about the knee joint [7].

Like any other ligamentous structure, the biomechanical properties of the ACL are determined by the geometry of the ligament as well as the tensile characteristics of both ligament midsubstance and the ligament-to-bone insertion site. Basically, they can be characterised by the relationship between ligament length and ligament tension, which can be determined when simultaneously measuring load and the corresponding elongation during experimental uniaxial tensile testing.

The resultant load-elongation curve can be divided into four distinct regions according to the structural properties of the ACL. A first nonlinear region, the so-called ‘toe region’, is described as collagen fibres, which are arranged in varying degrees of crimp, easily extend under low axial forces [26‐28]. The toe region is followed by a quasilinear region where collagen fibres reversibly deform. The slope of the linear region allows for reproducible determination of ligament stiffness (measured in Newtons per millimetre) and corresponds to the loads acting on the ACL during daily activities. In the intact knee, both the toe region and the linear region of the ACL loading curve will allow the tibia to translate anteriorly for 3–5 mm during knee motion as well as during an anterior drawer manoeuvre. With additional loading, the slope of the load-elongation curve decreases (yield-point) as plastic deformation of the collagen fibres occurs. Finally, the curve reaches the ultimate load, which is described as failure of the bone-ligamentbone complex. It may be derived from the load-elongation curve, that applying high loads to a ligament will increase the stiffness and may therefore sufficiently restrict excessive joint motion when high external loads are applied. Even more accurately, the biomechanical properties of a ligament are represented by the relation of stress and strain, where stress is defined as deformation per unit length (%) and where strain is defined as load per unit cross-sectional area (N/mm²) [26].

When a constant load is applied to a ligament, the increase in ligament length is called ‘creep’, whereas the decrease in load with the ligament constantly elongated is called ‘relaxation’. In vivo, cyclic loading of the ACL will cause gradual creep and relaxation, which results in increased knee laxity after physical activity. However, it will return to the original stiffness after a period of rest. The parameters derived during experimental ligament loading are considered to be essential to understanding ligament function and evaluating the appropriateness of different graft materials and fixation techniques commonly used in ACL reconstruction [26‐28]. In addition, visual or apperative control of the mode of failure during tensile testing allows identification of either graft slippage or deterioration of graft material under mono- or polycylic loading conditions, thus indicating what amount of graft tension loss should be expected during the postoperative rehabilitation period.

Estimations of ACL forces during activities of daily living calculated by Morrison revealed that ACL loads of 169 N may be expected during normal level walking, while descending stairs generated 445 N of in situ force due to the activation of the knee extensor apparatus [29‐31]. In contrast, ascending stairs as well as ascending or descending a ramp generated in situ forces below 100 N.

While estimation of in vivo ACL forces during normal activity have been subject to computational analyses, the biomechanical properties of the native ACL have extensively been analysed in ex vivo studies. Measurements using a universal force-moment sensor revealed that the in situ force of the intact ACL was largest at 15° of knee flexion and continuously decreased until 90° of knee flexion [23]. Focusing on the ACL force and ligament deformation during anterior tibial translation, Sakane et al. demonstrated that near full extension, the in situ force of the anteromedial bundle significantly differed from that of the posterolateral bundle with 110 N of anterior tibial load applied [23]. While there was no significant difference in the in situ force of the anteromedial bundle throughout the range of knee motion, the in situ force of the posterolateral bundle was significantly lower at 90° of knee flexion when compared to full extension, similar to what was measured for the entire ACL. This suggests that the role of the posterolateral bundle in response to anterior tibial loads near extension may be of more importance than was previously thought (Fig. 3).

Performing tensile testing of the bone-ligament-bone complex, Woo et al. reported the ultimate failure load of the native femur-ACL-tibia complex in younger cadaveric specimens to average 2160±157 N, while mean ACL stiffness was 242±28 N/mm [32, 33]. They were also able to demonstrate that both the ultimate failure load and the linear stiffness significantly decreased with age and with the axis of loading. Ultimate failure loads in older specimens being loaded in an anterior drawer mechanism averaged 496±85 N with a mean stiffness of 124±16 N/mm. Given this data, Noyes et al. concluded that the initial fixation strength of an ACL graft required for sufficient knee stability during daily activities should exceed 450 N, although earlier studies performed by Shelbourne and Gray reported excellent clinical results using graft fixation techniques with a significantly lower initial failure strength of only 248 N [34‐36].

ACL injuries commonly occur during sport activities with sudden stresses applied to the knee joint while the tibia is in contact with the ground. Typically, the ACL is torn in a non-contact deceleration situation that produces valgus and internal rotational moments on the knee joint that begins to flex from near extension. This usually occurs in high-risk pivoting sports when the athlete lands on the leg and starts rotating to the opposite direction [37].

Active quadriceps pull is considered to play an important role in the pathomechanism of ACL injury. Reflective eccentric quadriceps contraction accompanied by apparent weakness of the hamstring muscles allows the extensor mechanism to strain the ACL during anterior tibial translation. This mechanism can be observed during jump stop landings. Less frequently, direct contact injuries occur as a result of extensive valgus stress or hyperextension of the knee joint.

Isolated disruption of the ACL is a rather rare condition, while complete or incomplete ruptures accompanied by traumatic lesions to the medial or lateral meniscus as well as to the medial collateral ligament are reported in 80% of all cases [3, 37‐39]. Those concomitant lesions are reported to significantly influence the long-term functional outcome. Additionally, in both isolated and combined ACL injuries, minor or major bruising of chondral and subchondral structures may be present. Histologic analyses of bruised bone performed by Johnson et al. and others demonstrated areas of necrotic osteocytes and chondrocytes, indicating severe damage to the local osteochondral tissue homeostasis after ligament injuries of the knee [38‐43].

Disruption of the ACL inevitably results in alterations in knee kinematics as a transfer of loads can be effective only if the joint is mechanically stable. ACL insufficiency causes deterioration of the physiologic roll-glide mechanism of the femorotibial joint and results in an increased anterior tibial translation as well as an increased internal tibial rotation. In the advent of muscular fatigue or deficient neuromuscular control, the patient will experience this combined anterior and rotatory instability as a subluxation of the femorotibial joint. According to the concept of primary and secondary restraints, failure of a primary restraint will cause recruitment of secondary structures in order to resist external forces and to stabilise joint motion. The increase in loads applied to secondary structures may render them more susceptible to degeneration or secondary failure [44‐46]. Studies investigating the long-term functional outcome after conservative treatment of ACL ruptures, though not characterised by well designed prospective cohort studies, reported on the prevalence of radiographic osteoarthritis in 60%-90% of subjects 10–15 years after injury [2, 3, 44, 47].

Current research supports the concept that under observance of several key factors, arthroscopically assisted ACL reconstruction done with a biologic autograft significantly improves the stability and function of the knee in most patients. The key factors significantly influencing the functional outcome are: