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. 2019 Aug 14;477(9):2161–2174. doi: 10.1097/CORR.0000000000000884.

Biomechanical Function and Size of the Anteromedial and Posterolateral Bundles of the ACL Change Differently with Skeletal Growth in the Pig Model

Stephanie G Cone 1,2,3,4,5,6, Emily P Lambeth 1,2,3,4,5,6, Hongyu Ru 1,2,3,4,5,6, Lynn A Fordham 1,2,3,4,5,6, Jorge A Piedrahita 1,2,3,4,5,6, Jeffrey T Spang 1,2,3,4,5,6, Matthew B Fisher 1,2,3,4,5,6,✉
PMCID: PMC7000103  PMID: 31373947

Abstract

Background

ACL injuries are becoming increasingly common in children and adolescents, but little is known regarding age-specific ACL function in these patients. To improve our understanding of changes in musculoskeletal tissues during growth and given the limited availability of pediatric human cadaveric specimens, tissue structure and function can be assessed in large animal models, such as the pig.

Questions/purposes

Using cadaveric porcine specimens ranging throughout skeletal growth, we aimed to assess age-dependent changes in (1) joint kinematics under applied AP loads and varus-valgus moments, (2) biomechanical function of the ACL under the same loads, (3) the relative biomechanical function of the anteromedial and posterolateral bundles of the ACL; and (4) size and orientation of the anteromedial and posterolateral bundles.

Methods

Stifle joints (analogous to the human knee) were collected from female Yorkshire crossbreed pigs at five ages ranging from early youth to late adolescence (1.5, 3, 4.5, 6, and 18 months; n = 6 pigs per age group, 30 total), and MRIs were performed. A robotic testing system was used to determine joint kinematics (AP tibial translation and varus-valgus rotation) and in situ forces in the ACL and its bundles in response to applied anterior tibial loads and varus-valgus moments. To see if morphological changes to the ACL compared with biomechanical changes, ACL and bundle cross-sectional area, length, and orientation were calculated from MR images.

Results

Joint kinematics decreased with increasing age. Normalized AP tibial translation decreased by 44% from 1.5 months (0.34 ± 0.08) to 18 months (0.19 ± 0.02) at 60° of flexion (p < 0.001) and varus-valgus rotation decreased from 25° ± 2° at 1.5 months to 6° ± 2° at 18 months (p < 0.001). The ACL provided the majority of the resistance to anterior tibial loading at all age groups (75% to 111% of the applied anterior force; p = 0.630 between ages). Anteromedial and posterolateral bundle function in response to anterior loading and varus torque were similar in pigs of young ages. During adolescence (4.5 to 18 months), the in situ force carried by the anteromedial bundle increased relative to that carried by the posterolateral bundle, shifting from 59% ± 22% at 4.5 months to 92% ± 12% at 18 months (data for 60° of flexion, p < 0.001 between 4.5 and 18 months). The cross-sectional area of the anteromedial bundle increased by 30 mm2 throughout growth from 1.5 months (5 ± 2 mm2) through 18 months (35 ± 8 mm2; p < 0.001 between 1.5 and 18 months), while the cross-sectional area of the posterolateral bundle increased by 12 mm2 from 1.5 months (7 ± 2 mm2) to 4.5 months (19 ± 5 mm2; p = 0.004 between 1.5 and 4.5 months), with no further growth (17 ± 7 mm2 at 18 months; p = 0.999 between 4.5 and 18 months). However, changes in length and orientation were similar between the bundles.

Conclusion

We showed that the stifle joint (knee equivalent) in the pig has greater translational and rotational laxity in early youth (1.5 to 3 months) compared with adolescence (4.5 to 18 months), that the ACL functions as a primary stabilizer throughout growth, and that the relative biomechanical function and size of the anteromedial and posterolateral bundles change differently with growth.

Clinical Relevance

Given the large effects observed here, the age- and bundle-specific function, size, and orientation of the ACL may need to be considered regarding surgical timing, graft selection, and graft placement. In addition, the findings of this study will be used to motivate pre-clinical studies on the impact of partial and complete ACL injuries during skeletal growth.

Introduction

The ACL stabilizes the knee in multiple directions during activities of daily living and physically demanding athletic activities [16]. Its structure is well suited to the multidirectional loads it experiences; the ACL is commonly divided into two sub-bundles: the anteromedial and posterolateral bundles [4, 19, 26, 36, 39, 42]. The anteromedial bundle is more directly responsible for resisting anterior tibial loading and sees greater loads in deep flexion than the posterolateral bundle, while the posterolateral bundle contributes more to resisting rotational moments and in positions near full extension [16, 19, 34].

As many as 250,000 ACL injuries occur in the United States each year [18], and the incidence of ACL injury is rising in the skeletally immature population [5]. Recent studies have found rapid increases in the number of ACL reconstruction procedures annually in pediatric and adolescent patients [5, 10, 17, 35], with one study reporting that the fastest growing number of ACL reconstruction procedures is in boys and girls younger than 14 years [5]. Current clinical treatments for this age group focus on restoring stability while minimizing interruption of the femoral and tibial physes for patients with considerable growth remaining [1, 11]; however, these treatments may not restore normal kinematics and contact stresses in pediatric patients [28]. This suggests a need for more knowledge regarding the normal function of the ACL during growth.

The size and orientation of the ACL also undergo substantive changes during growth in childhood. The ACL increases in steepness in the sagittal and coronal planes in humans during growth [24, 32]. The ACL cross-sectional area increases in children up to 10 years old, with more modest changes during adolescence [9, 40]. However, little is known regarding whether and how ACL function and anteromedial and posterolateral bundle function, size, and anatomic orientation change during growth.

Large-animal models are valuable when the ability to collect human data is limited, such as in pediatric cadaveric studies. Previously, the pig ACL model has been validated as a surrogate for the ACL in skeletally mature humans; the length and cross-sectional area of the healthy ACL were found to be comparable between humans and pigs [31]. Additionally, the pig model better replicates human anteromedial and posterolateral bundle function under anterior tibial loads than do sheep and goat models [41]. More recently, the skeletally immature pig model yielded similar changes in ACL orientation to the human ACL with growth [6, 24]. Another study of ACL reconstruction and repair also employed the pig model, using both young and mature animals [29]. Thus, the pig model can provide clinically useful insight into ACL function during growth.

The objectives were defined to study age-dependent changes in four parts: (1) to measure joint kinematics in response to applied AP loads and varus-valgus moments, (2) to assess the response of the ACL under the same applied loads, (3) to compare the relative functional contributions of the anteromedial and posterolateral bundles of the ACL, and (4) to study changes in the cross-sectional area, length, and angular orientation of the bundles of the ACL.

Materials and Methods

Study Design

The study design involved collecting hind limb specimens from pigs and assessing the specimens via MRI and biomechanical testing (Fig. 1). All pigs used in this study were obtained from a university-owned herd (Swine Education Unit, NC State University, Raleigh, NC, USA) and were healthy and of normal size. The animals were cared for according to the management practices outlined in the Guide for the Care and Use of Agricultural Animals in Teaching and Research [12], and experimental protocols were approved by our local institutional animal care and use committee. Hind limbs were collected from 30 female Yorkshire crossbreed pigs at ages ranging from 1.5 to 18 months. Age groups (n = 6 per group) and equivalencies to human growth were based on studies of porcine skeletal and sexual maturity [7, 33] and included: early youth (1.5 months), juvenile (3 months), early adolescent (4.5 months), adolescent (6 months), and late adolescent (18 months) groups. Ages of animals were within a range of +/- 3 days for the early youth and juvenile groups, and within a range of +/- 7 days for the early adolescent through late adolescent groups. Sample sizes were initially determined by a power analysis based on preliminary imaging data (effect size of 2), which found that n = 8 per age group would be sufficient to detect an effect size of 2 at a power of 0.8 and adjusting for multiple comparisons. After initial testing, we observed larger effect sizes and determined that n = 6 was sufficient to detect changes in major outcomes, so animals were limited to n = 6 per group. Specimens were wrapped in saline-soaked gauze and stored at -20 °C.

Fig. 1.

Fig. 1

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The overall methods for this manuscript are described here, including specimen collection, magnetic resonance imaging, and biomechanical testing.

MRI Analysis

MRI analysis was performed to study the anatomic orientation, length, and cross-sectional area of the ACL and its anteromedial and posterolateral bundles. Hind limbs were allowed to thaw at room temperature in preparation for MRI. Stifle joints were imaged at full extension (approximately 30° to 40° of flexion) at the Biomedical Research Imaging Center (University of North Carolina – Chapel Hill, Chapel Hill, NC, USA). MRI scans were performed in a 7.0-Tesla Siemens Magnetom scanner (Siemens Healthineers, Erlangen, Germany) using a double-echo steady-state sequence (flip angle: 25°; TR: 17 ms, TE: 6 ms; acquisition time: 24 minutes; FOV: 123 x 187 x 102 mm) with a 28-channel knee coil (Siemens Healthineers) and voxel size of 0.42 x 0.42 x 0.4 mm, with no gap between slices (see Fig. 1; Supplemental Digital Content, http://links.lww.com/CORR/A211). After imaging, each limb was wrapped in saline-soaked gauze and stored again at -20 °C.

We analyzed MR images using commercially available software (Simpleware 7.0, Synopsys, Chantilly, VA, USA) to measure the orientation and size metrics. To calculate the orientation of each bundle, we used a multiangle measurement tool to determine the angle of each bundle relative to the AP axis of the tibial plateau in both the sagittal and coronal planes, as previously reported by us and others [6, 24]. We calculated tissue length by identifying coordinates of approximately 15 evenly distributed points at the tibial and femoral insertion sites for each tissue and calculating the 3-D centroid for each insertion using a custom Matlab code (Matlab, Mathworks, Natick, MA, USA). We then calculated length as the magnitude of the vector between the centroids. Images were segmented to create 3-D models of each tissue, specifically the entire ACL, the anteromedial bundle, and the posterolateral bundle (Fig. 2). Models were refined using “close” and “discrete Gaussian” filters before being exported as .stl files.

Fig. 2 A-D.

Fig. 2 A-D

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(A) Coronal and sagittal angles of the ACL bundles were measured on the most anterior (coronal, image shown) or the most medial (sagittal, image shown) slice depicting the full ACL relative to the AP or medial-lateral plane of the tibial plateau (determined from a different MRI slice and transferred for measurement). (B) Bundle length was calculated between insertions (marked with red stars for the posterolateral bundle) based on points collected near the femoral and tibial insertion sites as shown in sagittal planes. (C) MR images were translated into 3-D models by creating masks for the individual tissues throughout the MRI scan and compiling the image masks to create 3-D models. (D) The cross-sectional area was measured by isolating and rotating 3-D models of the ACL bundles, generating point clouds for each tissue, creating 2-D cross-sections of these point clouds (shown in blue inset), and measuring the cross-sectional area at slices in the midsubstance of the tissue.

Cross-sectional area values were calculated from MR images in a manner similar to the method of Fujimaki et al. [15]. The anteromedial and posterolateral bundles of the ACL were visualized, confirmed as inserting into the tibial plateau on the anterior and posterior side of the most medial aspect of the anterior insertion of the lateral meniscus (an anatomic feature of the porcine ACL), and transformed into individual masks (see Fig. 2; Supplemental Digital Content, http://links.lww.com/CORR/A211). The 3-D models were transformed into point clouds and imported into a custom Matlab code for further analysis. Point clouds for each tissue were rotated to align the line-of-best-fit (or long axis) to the Z-axis of a Cartesian coordinate system. The point cloud was then translated along the Z-axis to originate on the orthogonal X-Y plane, and points were projected onto planes distributed in 1-mm slices along the Z-axis. The area within the collapsed points was recorded for each slice, and the areas identified from the central 50% of the tissue were averaged to calculate the mean cross-sectional area value for the tissue.

Biomechanical Testing

We then performed biomechanical testing to study the AP tibial translation and varus-valgus rotation of the joints under applied loads and moments. Additionally, this testing was performed to study the in situ forces carried by the ACL and its anteromedial and posterolateral bundles under the applied forces and moments. Before biomechanical analysis, limbs were removed from storage and allowed to thaw at room temperature. The femur, tibia, and fibula were cut in the center of the diaphysis, and the soft tissue was removed up to the joint. The bones were fixed in molds using an epoxy compound (Everglass, Evercoat, Cincinnati, OH, USA). Joints were wrapped in saline-soaked gauze, and additional saline was applied as needed throughout testing.

Robotic testing systems have been widely used to analyze the in situ function of musculoskeletal joints and individual tissues [13, 14, 20, 27, 34]. Testing was performed using a 6 degree of freedom robotic system (KR300 R2500, Kuka, Shelby Charter Township, MI, USA) powered by a separate controller (KRC4, Kuka, Shelby Charter Township, MI, USA) combined with a 6 degree of freedom force/moment sensor (Omega160 IP65, ATI Industrial Automation, Apex, NC, USA) and integrated and controlled via the simVitro software package (Cleveland Clinic, Cleveland, OH, USA). This system is capable of operating under both kinematic and kinetic control, with kinematic repeatability of 0.1 mm and 0.1° and load cell sensitivity of 0.25 N.

Specimens were attached to the robotic system with custom clamps, and the anatomic coordinate system of the joint was determined relative to the coordinate system of the robotic manipulator using a point digitizer with an accuracy of 0.23 mm (G2X, Microscribe, Amherst, VA, USA) as previously described [27, 30, 34]. A passive path was established for each joint by increasing the flexion angle of the joint from full extension (approximately 40° of flexion, measured with a goniometer) to 90° by 1° increments while minimizing forces and moments in the other 5 degrees of freedom and recording the kinematics.

Once the passive path positions were established, joint kinematics under applied anterior tibial loads at 40°, 60°, and 90° and varus-valgus moments at 60° were obtained (Table 1). The robotic system was operated under force control to apply selected loads at 40°, 60°, and 90° of flexion. Age-specific forces and moments (see Table 1; Supplemental Digital Content, http://links.lww.com/CORR/A211) were selected based on preliminary experiments to engage the connective tissues of the knee and reach the linear region of the load-displacement curve for the ACL under anterior loads. These load changes (sevenfold increase) were scaled with bone size increases (sevenfold increase). Kinematics were recorded under these applied loads and repeated in position control while force and moment data were collected. The anteromedial bundle of the ACL was isolated and transected, and the kinematics of the intact joint were repeated to obtain force and moment data remaining in the anteromedial-deficient state. The posterolateral bundle was then transected. This order was maintained throughout testing because of our inability to access the posterolateral bundle from the anterior aspect of the joint before transecting the anteromedial bundle; however, in preliminary validation studies, blunt separation of the bundles resulted in minimal change in recorded resultant force under applied load (mean of 1.7 N in the 3 month old juvenile group [n = 5] and 3.8 N in the 18 month old late adolescent group [n = 5]). This suggests minimal interaction between the anteromedial and posterolateral bundles. The kinematics of the intact joint were again repeated to measure the loads and moments resisted in the ACL-deficient state.

Table 1.

Experimental protocol for robotic testing of the stifle joints of pigs

graphic file with name abjs-477-2161-g003.jpg

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Biomechanical Data Processing

Kinematic measurements included AP tibial translation, calculated as the AP distance between the point of maximum applied anterior force and posterior force, and varus-valgus rotation, measured as the rotation in degrees between the maximum applied varus and valgus torques. AP tibial translation was normalized to the length of the tibial plateau in the sagittal plane, measured from previously collected MR images, to correct for size differences because of growth between age groups. Forces were recorded in the AP, medial-lateral, and proximal-distal directions at the peak force or moment of each applied condition. The principle of superposition was applied to calculate the in situ forces for each force component in the anteromedial bundle, posterolateral bundle, and ACL under the applied loads in all three conditions [34, 41]. Resultant forces were calculated as the result of the force vector for each tissue under each applied load. The load carried in a tissue can exceed 100% of the net force of another tissue simultaneously applies an opposing force (that is, posterior force to offset anterior force). These effects are fairly minimal, but can result in tissue force values greater than the applied target force. Normalized forces were calculated as a tissue-specific percentage of the overall force in the joint, measured at the peak load.

Primary and Secondary Outcomes of Interest

Our primary study outcomes were age-dependent changes in AP tibial translation and varus-valgus rotation, the biomechanical contribution of the ACL under applied anterior tibial loads, and relative function of the anteromedial and posterolateral bundles of the ACL under anterior tibial loads. We tested this by applying an anterior tibial load to joints of various ages and recording the resulting deformations and sensed forces.

Our secondary study outcomes were changes in the size and angular orientation of the ACL bundles during growth. We tested these by comparing changes in tissue cross-sectional area, length, and both sagittal and coronal tissue orientation throughout skeletal growth in the porcine model.

Statistical Analysis

The statistical analysis was performed using commercial software (JMP Pro 13.0, SAS Institute, Cary, NC, USA). For AP tibial translation, normalized AP tibial translation, and ACL force contribution, we performed a two-way ANOVA with age as an independent variable and flexion angle as a repeated measure. Varus-valgus rotation, ACL angle, bundle angle, ACL length, and cross-sectional area of the ACL were analyzed with a one-way ANOVA using age as the independent variable. We analyzed bundle contributions under varus and valgus torque, bundle angle, bundle length, and bundle cross-sectional area using a two-way ANOVA with age as an independent variable and bundle as a repeated measure. Normalized bundle contributions to anterior tibial force were analyzed individually by bundle using a two-way ANOVA with age as the independent variable and flexion angle and bundle as repeated measures; paired t-tests were used to compare values between bundles. For each ANOVA, Tukey’s honestly significant difference post-hoc analyses were performed. For all tests, the overall alpha value was set at 0.05. Data are reported in tables and supplementary material as the mean ± SD (95% CI). Figures showing individual data points and means with 95% CIs are reported in the main text, with minimal statistical markings for clarity.

Results

Age-Dependent Joint Kinematics

No differences in AP tibial translation were found in response to applied loads because of age, despite increases in specimen size (see Table 2; Supplemental Digital Content, http://links.lww.com/CORR/A211). Once it was normalized to the length of the tibial plateau in the sagittal plane, AP tibial translation decreased with increasing age (Fig. 3). The mean values for normalized AP tibial translation in the late adolescent group (18 months) were only 50%, 56%, and 63% of that of the early youth group (1.5 months) at 40°, 60°, and 90° of flexion, respectively (see Table 3; Supplemental Digital Content, http://links.lww.com/CORR/A211). These decreases occurred primarily before the onset of adolescence, with no changes during adolescence (between 4.5 and 18 months old) across all flexion angles. Varus-valgus rotation decreased with increasing age (Fig. 3). The mean values for varus-valgus rotation of the late adolescent (18 months) were 25% of that of the early youth group (1.5 months) (see Table 4; Supplemental Digital Content, http://links.lww.com/CORR/A211).

Fig. 3 A-B.

Fig. 3 A-B

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(A) In response to applied tibial loads, AP tibial translation decreased with age (p < 0.001) when normalized to the length of the tibial plateau in the sagittal plane (data shown at 60° of flexion). (B) Additionally, varus-valgus rotation decreased with age (p < 0.001). The bars represent the mean and 95% CI and the points represent data from individual specimens.

ACL Function Under Applied Anterior Loads

The ACL served as the primary restraint to anterior tibial load in the joint throughout skeletal growth (Fig. 4). The functional contribution of the ACL was measured as the percentage of the overall joint anterior force at the peak applied anterior tibial load. Specifically, the mean contribution of the ACL ranged from 75% to 104% at 40° of flexion, 97% to 110% at 60° of flexion, and 98% to 111% at 90° of flexion across all age groups relative to the intended target load (see Table 5; Supplemental Digital Content, http://links.lww.com/CORR/A211).

Fig. 4.

Fig. 4

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The ACL was a dominant soft-tissue restraint, reported here as a percentage of the target load, to applied anterior tibial loading, with no effect detected because of age (p = 0.63). Percentages can exceed 100% of the target value due to opposing contributions of other soft tissues in the joint. Data are presented with a representative flexion angle (60°); the bars represent the mean and 95% CI.

ACL Bundle Biomechanical Function Under Applied Loads

Although the biomechanical function of the entire ACL was consistent across ages, the functional contributions of the anteromedial and posterolateral bundles shifted from both bundles having substantial but variable contributions in the younger age groups towards dominance of the anteromedial bundle under anterior loading after the onset of adolescence (Fig. 5). In early youth (1.5 months), the mean contribution of the anteromedial bundle was 44%, 50%, and 49% at 40°, 60°, and 90° of flexion, respectively (see Table 6; Supplemental Digital Content, http://links.lww.com/CORR/A211). The functional contribution of the anteromedial bundle did not increase substantially from youth to the onset of adolescence (4.5 months). By late adolescence (18 months), the contributions of the anteromedial bundle had a mean of 89%, 92%, and 86% of the anterior force of the ACL at 40°, 60°, and 90° of flexion, respectively (see Table 6; Supplemental Digital Content, http://links.lww.com/CORR/A211). This divergence of functional contributions was such that there was a difference in contributions between the bundles by 18 months of age for all flexion angles (see Table 6; Supplemental Digital Content, http://links.lww.com/CORR/A211).

Fig. 5.

Fig. 5

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The in situ force of the anteromedial and posterolateral bundles of the ACL under applied anterior loads is shown at (A) 40°, (B) 60°, and (C) 90° of flexion (reported as a percentage of the total force in the ACL). Data varied in younger groups, with substantial forces carried by both bundles (p values ranging from 0.04 to 0.98 between bundles for the 1.5- and 3-month age groups across all flexion angles); however, after the onset of adolescence, the contribution of the anteromedial bundle became dominant under an anterior load (p < 0.001 for all flexion angles at 18-month age group). The bars represent the mean and 95% CI; *p < 0.05 between bundles. AM = anteromedial; PL = posterolateral.

Under applied varus and valgus torques at 60° of flexion, the behavior of the anteromedial and posterolateral bundles also varied as a function of age (Fig. 6). In the young age groups (1.5 to 3 months), biomechanical functional contributions from the anteromedial and posterolateral bundles were similar under varus and valgus loading. Beginning in early adolescence (4.5 months), the anteromedial bundle began to carry a greater portion of the resultant load under an applied varus moment. By adolescence (6 months), the anteromedial bundle carried a mean of 86% of the resultant force. In the late adolescent group (18 months), the anteromedial bundle carried 80% of the resultant force. Dissimilarly, under valgus loading, both bundles continued to have a substantial role with increased age. Highly variable behavior under valgus loads resulted in mean contributions from the anteromedial bundle that ranged between 44% and 72%, with no effect because of age (see Table 7; Supplemental Digital Content, http://links.lww.com/CORR/A211).

Fig. 6 A-B.

Fig. 6 A-B

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The in situ force of the anteromedial and posterolateral bundles in the ACL under applied varus and valgus moments is shown (normalized to the total in situ force of the ACL). (A) Under varus moment, data varied in younger age groups, with both bundles carrying a portion of the forces reported as a percentage of the total force in the ACL (p = 0.99 at 1.5 months and p = 0.94 at 3 months between bundles); the anteromedial bundle playing a greater role at later ages (p = 0.002 at 4.5 months, p < 0.001 at 6 months, and p < 0.001 at 18 months between bundles). (B) Dissimilarly, the data were highly variable throughout all age groups under an applied valgus moment. The anteromedial and posterolateral bundles did not exhibit bundle-specific behavior at any age under the valgus moment (p = 0.15 for the main effect of bundle type). The bars represent the mean and 95% CI; *p < 0.05 between bundles. AM = anteromedial; PL = posterolateral.

ACL Size and Orientation

The cross-sectional area of the ACL increased steadily with age (444% mean increase from early youth to late adolescence) (see Table 8; Supplemental Digital Content, http://links.lww.com/CORR/A211). These changes were rapid between 3 and 4.5 months old (Fig. 7). Simultaneously, the ACL had a 250% increase in length from 1.5 to 18 months of age in the pig model (see Fig. 3; Supplemental Digital Content, http://links.lww.com/CORR/A211). These increases occurred gradually throughout skeletal growth.

Fig. 7 A-B.

Fig. 7 A-B

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(A) The cross-sectional area of the ACL increased with increasing age (p < 0.001). (B) The cross-sectional area values of the anteromedial and posterolateral bundles were similar through 4.5 months old (p = 0.99 at 1.5 months, p = 0.98 at 3 months, and p = 0.99 at 4.5 months between bundles), with the anteromedial bundle becoming larger in the 6- and 18-month age groups (p = 0.003 at 6 months and p < 0.001 at 18 months between bundles). The bars represent the mean and 95% CI. *p < 0.05 between bundles. AM = anteromedial; PL = posterolateral.

Both bundles had an increased cross-sectional area during youth, but the anteromedial bundle had continued increases in the cross-sectional area throughout adolescence while the posterolateral bundle plateaued at the onset of adolescence. The midsubstance cross-sectional area of the anteromedial bundle increased by sevenfold from 1.5 to 18 months old (Fig. 7). Simultaneously, the cross-sectional area of the posterolateral increased by only 2.3 times (see Table 8; Supplemental Digital Content, http://links.lww.com/CORR/A211). Between bundles, the cross-sectional area values only differed in the 6- and 18-month age groups, with the anteromedial bundle reaching 203% of the cross-sectional area of the posterolateral bundle by 18 months old (Fig. 7). From birth through skeletal maturity, there were similar increases in the anteromedial and posterolateral bundle lengths (262% and 284%, respectively) (see Fig. 3; Supplemental Digital Content, http://links.lww.com/CORR/A211). The length of the anteromedial bundle was greater than that of the posterolateral bundle in all groups between 1.5 and 18 months old (see Table 9; Supplemental Digital Content, http://links.lww.com/CORR/A211).

The angular orientation of the anteromedial and posterolateral bundles similarly increased with increasing age in both the sagittal and coronal planes. In the sagittal plane, these changes occurred in similar manners across bundles, with a mean increase of 29° in the posterolateral bundle (early youth: 36° ± 3° [95% CI 32 to 39]; late adolescence: 66° ± 4° [95% CI 62 to 71]; p < 0.001) and 23° in the anteromedial bundle (early youth: 32° ± 3° [95% CI 29 to 35]; late adolescence: 55° ± 5° [95% CI 50 to 61]; p < 0.001) from early youth (1.5 months) to late adolescence (18 months) (Fig. 8). Increases occurred between consecutive age groups only at the early adolescent stage (3 to 4.5 months) (see Table 10; Supplemental Digital Content, http://links.lww.com/CORR/A211). The mean coronal angle of the anteromedial bundle increased by 28°, while that of the posterolateral bundle increased by 31° from youth through late adolescence (Fig. 8). These increases occurred earlier in skeletal growth, with increases in both bundles between early youth (1.5 months) and the juvenile (3 months) stage, and increases in the anteromedial bundle in the juvenile (3 months) and adolescent (4.5 months) stages (see Table 11; Supplemental Digital Content, http://links.lww.com/CORR/A211).

Fig. 8 A-B.

Fig. 8 A-B

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The (A) sagittal and (B) coronal angles of the anteromedial and posterolateral bundles of the ACL increased similarly with increasing age (p < 0.001). The bars represent the mean and 95% CI. AM = anteromedial; PL = posterolateral.

Discussion

Recent increases in the demand for ACL reconstructions in skeletally immature patients, coupled with reports of suboptimal outcomes from traditional ACL reconstruction procedures in this population, have led to a need for age-specific studies of the native function and structure of the ACL during growth. As such, in this work we studied the joint stability, the biomechanical function of the ACL and its anteromedial and posterolateral bundles, and the structure and orientation of these bundles in a skeletally immature preclinical model of the knee, the porcine stifle joint. We found that increasing age resulted in lower joint laxity under both applied AP loads and varus-valgus moments. Additionally, we showed that the ACL carried the majority of force under an applied anterior load across ages and contributed to rotatory stability and the relative function of the anteromedial and posterolateral bundles of the ACL varied with age under anterior loading and varus torque. Near early adolescence, a shift in both relative size (cross-sectional area) and function occurred, where the anteromedial bundle had an increase in cross-sectional area and functional contribution, whereas the posterolateral bundle reached a plateau in cross-sectional area and decreased functional contribution. The angular orientation of both bundles also increased with increasing age; however, the change in the orientation of the anteromedial and posterolateral bundles was similar throughout adolescence. These findings are important because the in situ functional behavior of a tissue depends on the tissue size, orientation, and intrinsic material properties, so by identifying growth stages where major structural and functional shifts occur in the ACL, we may be able to improve clinical planning to adapt for age-specific needs.

This study had several limitations. Care should be taken when trying to directly extrapolate results from a translational large-animal model to humans. However, the stifle joint in pigs has commonly been accepted as a surrogate for human knees, with congruence between the structural and functional properties of the ACL in mature populations [31, 41]. Yet, some anatomic differences exist between the pig stifle joint and human knees. For example, testing in this study was limited to maximum extension at 40°. This resulted in a limitation because the posterolateral bundle plays a greater functional role near full extension (0°) in humans, and we were unable to study age-related changes for this flexion angle [34]. Additionally, potential limitations may stem from our protocol, in which we transected the anteromedial bundle before the posterolateral bundle. While transecting the bundles in this order was necessary for subsequent applied loading conditions not reported here, if the anteromedial and posterolateral bundles have a considerable physical interaction with one another, this may have affected the data presented in this study. Specifically, transection of the anteromedial bundle may have resulted in an underestimation of the functional contribution of the posterolateral bundle, because the principle of superposition assumes the anteromedial and posterolateral bundles are independent [34, 41]. However, preliminary studies in our laboratory have confirmed that blunt separation of the bundles does not alter the forces carried by the ACL under these applied loads across a range of age groups.

Additionally, we limited our comparison to the commonly defined anteromedial and posterolateral bundles, although recent studies have highlighted the presence of three major bundles in the ACL in both humans and pigs: the anteromedial, posterolateral, and intermediate bundles [21, 22]. Although the anteromedial and posterolateral bundles were easily discerned across all ages, the intermediate bundle could only be separated from the posterolateral bundle at older ages during dissection, and it was not consistently discernible on MR images. For consistency across ages, the posterolateral and intermediate bundles were considered a single bundle. Reported sex-dependent differences in the incidence of ACL injury and the structural and biomechanical properties of the knee have motivated a need to investigate the effect of sex on the function of the ACL in both healthy and injured people during skeletal growth [2, 37, 38]. As such, future studies should replicate this study within a male pig population to identify any sex-dependent differences in ACL function throughout growth. Additionally, noninvasive parameters, such as those collected using MRI, should be explored to enable a longitudinal approach to studying functional properties in lieu of the current cross-sectional approach [3].

Our imaging studies were limited by the resolution of our scanner and the repeatability of our analysis methods. The scanner used in this study had a resolution of 0.42 x 0.42 x 0.44 mm, which is well below the magnitude of our reported values, and well below our between-group differences in length and cross-sectional area. Inter-viewer repeatability studies have shown that our sagittal angle measurements were repeatable within 2.6°, and our coronal angle measurements were repeatable within 2.7° across separate viewers. Again, these differences are well below the reported between-group differences in this study.

Analysis of the kinematics resulting from applied AP loads and varus-valgus moments in this study revealed that the general joint laxity decreases considerably with increasing age in the pig model. Other studies have addressed the kinematic response of porcine joints to these loads in adolescent and skeletally mature animals, and while our study did not match the specific porcine breed, age, and loading conditions, our findings for the adolescent groups generally fit with the magnitude of AP tibial translation and varus-valgus rotation reported previously [8, 25].

In this work we found that the ACL provided the vast majority of physical restraint to applied anterior tibial loads, not only in mature cases but throughout skeletally immature growth. These findings agreed with and built off of previous work showing that the ACL is the primary restraint to anterior tibial loads in both pig models and humans. Specifically, Xerogeanes et al. [41] found that under 100 N applied anterior tibial load, the adolescent porcine ACL restrained 95.6 N ± 10.6 N while the human ACL restrained 94.0 N ± 4.0 N.

Our findings regarding the functional contributions of the anteromedial and posterolateral bundles build from previous work studying the relative function of these bundles in skeletally mature human and animal models. Specifically, the relative force contributions of the anteromedial and posterolateral bundles under applied loads in adolescent (7-month-old) pigs were found to be more similar to those of humans than those of goats and sheep [41]. In the aforementioned study [41], the anteromedial bundle carried 65% ± 21% of the in situ force of the ACL under a 100 N applied anterior load for 6-month-old pigs, which was similar to our findings. Furthermore, it has been shown that the anteromedial bundle is a more considerable contributor under anterior tibial loads in the joints of skeletally mature pigs [27]. However, we showed that this is not true in younger age groups. Additionally, in younger age groups, we found that the posterolateral bundle plays a functional role under anterior tibial loads at all flexion angles in addition to applied rotational torque.

Along with our analysis of functional changes in the ACL, this study compared the cross-sectional area, length, and angular orientation of the ACL bundles across stages of growth. While there is a paucity of functional data on the human ACL during childhood and adolescence, some comparisons have been made between anatomic changes in the ACL of pigs and humans throughout growth. These studies have shown that the angular orientation of the ACL becomes steeper relative to the tibial plateau in both the sagittal and coronal planes during skeletal growth in both species [6, 24]. Additional studies in humans have found a plateau in the growth of the cross-sectional area of the ACL before the end of overall body growth [9]. In our pig model, we found similar results; the cross-sectional growth in the ACL plateaued during adolescence. In the future, data regarding the growth of the cross-sectional area of the anteromedial and posterolateral bundles in humans should be pursued to study similarities or differences in human and pig bundle-specific changes.

These findings are clinically relevant because of the rapidly increasing numbers of ACL injuries in children and adolescents [5, 10] and the relatively high proportion of osteoarthritis and secondary injuries in this patient population [1]. Although double-bundle surgeries may be difficult or impossible because of limited joint space, the information presented here suggests directions that can guide future clinical research toward age-specific approaches. In particular, factors worthy of future consideration include single-bundle graft size and placement with respect to the current ACL size, orientation, and function, as well as anticipated changes in these parameters with future growth. Before implementing our findings into clinical practice, further work is needed to confirm the occurrence and specific timing of changes in ACL function and structure in humans. Previously, modifications have been made in the surgical technique used for these populations relative to adults, primarily to avoid disturbing the tibial and femoral growth plates in patients with considerable growth remaining [11]. Data on the age-specific function of the bundles of the ACL could be useful in developing age-specific procedures for the treatment of childhood ACL injuries, and in this study we have determined that the relative function of the anteromedial and posterolateral bundles changes throughout skeletal growth, with both bundles playing an important role under anterior tibial loads in youth and early adolescence. This expands on the previous understanding of the field, which held that the anteromedial bundle was dominant under applied anterior tibial loads in mature cases, while the posterolateral bundle primarily functioned under rotatory movements and near full extension [41]. Current adaptations for ACL reconstruction in young patients (such as all-epiphyseal and over-the-top methods) may not match the orientation and functional behavior of the ACL, and any changes in size, orientation, and function that would occur normally after surgery could further complicate effective surgical treatment [23]. Further data from human studies, both clinical and cadaveric, may confirm the importance of matching age-specific function via graft selection and placement in children. The findings presented here can aid in directing future clinical studies to focus on aspects such as the timing of changes in the cross-sectional area of the ACL bundles, and direct biomechanical studies to anticipate general changes in laxity throughout growth and to focus on a redistribution of mechanical function between the bundles of the ACL.

In summary, while the overall function of the ACL as a primary stabilizer against anterior tibial load is maintained throughout skeletal growth, the individual roles of the anteromedial and posterolateral bundles undergo a major shift from shared biomechanical function in youth and early adolescence to anteromedial bundle dominant behavior after the onset of adolescence in pigs. These changes in function were coupled with changes in ACL size, with continued increase in the cross-sectional area of the anteromedial bundle and a plateau in posterolateral bundle cross-sectional area, but no differences between bundles in orientation as the bundles increased similarly throughout growth. These findings relating to a shift in ACL bundle function suggest that age-specific surgical treatments with a focus on replicating the shared function of the anteromedial and posterolateral bundles during youth, and the anteromedial bundle dominance in adolescence, may result in improved functional outcomes in children; however, clinical studies are needed to confirm the relevance of these findings in humans. Moving forward, this work will be expanded to compare male and female pigs during growth, and the findings presented here will be used to motivate preclinical studies on the effect of partial and complete ACL injuries during skeletal growth. Additionally, the structural findings in this work can be compared with retrospective imaging databases of human growth to determine the similarities or differences in the timing of changes in the bundles of the ACL across species.

Acknowledgments

We thank Paul Warren MS, and Stephanie Teeter MA, for their assistance during mechanical experiments. We thank Hope Piercy BS, for her assistance with image processing, and we thank Sean Simpson BS, for his assistance with animal work. Additionally, we thank the Swine Educational Unit at North Carolina State University for their contributions to this work. We thank the Small Animal Imaging Facility at the UNC Biomedical Imaging Research Center for providing the 7T magnetic resonance imaging service.

Footnotes

The institution or one or more of the authors (SGC, MBF) has received funding from the National Institutes of Health (R03 AR068112, R01 AR071985) and the National Science Foundation (DGE-1252376). Partial support was provided by a National Cancer Institute core grant, P30-CA016086-40. Each author certifies that he or she has no commercial associations (consultancies, stock ownership, equity interest, patent/licensing arrangements, etc.) that might pose a conflict of interest in connection with the submitted article.

All ICMJE Conflict of Interest Forms for authors and Clinical Orthopaedics and Related Research® editors and board members are on file with the publication and can be viewed on request.

Each author certifies that his or her institution approved the animal protocol of this investigation and that all investigations were conducted in conformity with ethical principles of research.

This work was performed at North Carolina State University (functional testing), Raleigh, NC, USA, and the University of North Carolina (MRI), Chapel Hill, NC, USA.

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