LIGAMENT FIBER LENGTH-TENSION PROPERTIES

In the early writings of Fick in 1911,²³ the description of ligaments having fibers grouped into bands was presented, along with the concepts that not all fibers of a ligament would be tense at one time and knee flexion angle played an important role in which fibers were under tension and which were slack. Since then, numerous publications have appeared that demonstrate how ligament fiber tension and function depend on the initial length and attachment site and on the specific joint position of both flexion-extension and tibial rotation.^28,31,32,72 These concepts are important because they form the basis for the placement of knee ligament grafts and provide insight into the concepts of single- and double-bundle anterior cruciate ligament (ACL) and posterior cruciate ligament (PCL) reconstruction. In addition, the ligament fiber length-tension properties form the basis for the tensioning rules regarding knee flexion and load applied for ligament grafts at surgery. Major disagreement still exists among investigators on these points, as is discussed in this chapter.

It is first important to understand the length-tension behavior of individual ligament fibers. The viscoelastic properties of ligament fibers allow them to initially undergo elongation under load in resisting joint displacements. The tension developed in the ligament fibers depends on its tibiofemoral attachment site, initial length, initial joint position, and subsequent joint displacement. In Figure 4-1, a force-versus-length curve is shown for human ACL fibers. The “Lo” point describes the initial loading of the ligament fiber; when the ligament bone attachment distance of the fiber is less than Lo, the fiber is slack. The fiber functions only when the distance between the fiber attachment sites reaches the length of the fiber to imitate loading. The fiber functions only within a narrow elongation or strain, which has been estimated at strains of 5% or less.¹⁹ Failure of the individual fiber begins in the upper range estimated at the 8% level of strain.

FIGURE 4-1 Force-length curve for human anterior cruciate ligament (ACL) fibers. Fiber force is shown as a function of the distance between attachments (length). In-vivo functional activities are thought to occur within the shaded region, which corresponds to fiber strain of only 5%.

(From Grood, E. S.: Placement of knee ligament grafts. In Finerman, G. A.; Noyes, F. R. [eds.]: Biology and Biomechanics of the Traumatized Synovial Joint: The Knee as a Model. Rosemont, IL: American Academy of Orthopaedic Surgeons [AAOS], 1992; pp. 393–417.)

Collagen fibers are not highly elastic and demonstrate failure at relatively low levels of elongation or strain. Because ligament fibers function only within a small range of lengths and the fiber lengths vary with knee rotations and translations, these fiber length patterns can be measured under in vitro conditions. Beynnon and coworkers^5–7 measured fiber length patterns in vivo, described further in Chapter 12, Scientific Basis of Rehabilitation after ACL Autogenous Reconstruction.

HOW LIGAMENT FIBER MICROGEOMETRY DETERMINES LIGAMENT FUNCTION AND FAILURE

The three-dimensional microgeometry of ligament fibers is only partially understood on a macroscopic and ultrastructural level. From a macroscopic level, Figures 4-2 to 4-5 show the concepts of how ligament fibers determine functional properties and the mechanisms of ligament failure. Figure 4-2 shows a stress-strain curve for a typical collagen fiber. The initial toe region involves straightening of the collagen fibers, followed by a linear region. At the end of the linear region, failure of individual collagen fibers occurs, which progresses to final failure of the whole structure. Note that the total fiber strain to failure is only 8% in this illustration. Figure 4-3 shows the effect on the structural properties of a ligament on increasing the cross-sectional area. There is an increase in stiffness and ultimate force to failure. Figure 4-4 shows the effect of different lengths of a collagen structure under load, both having the same cross-sectional area. In the illustration, there is a 50% decrease in stiffness in the ligament that is twice as long. This effect has been noted in ACL semitendinosus-gracilis tendon (STG) reconstructions where aperture fixation of the graft directly at the tunnel site may improve mechanical properties of the graft in addition to decreasing graft-bone tunnel motion. Figure 4-5 shows the effect of fiber microgeometry on structural properties. The finger-trap illustration shows a major decrease in stiffness and ultimate strength because the collagen fibers are not loaded in a symmetrical manner.

FIGURE 4-2 Example of a typical normalized stress-strain curve for collagen. The resulting curve provides mechanical (material) parameters that are independent of tissue dimensions.

(From Butler, D. L.; Grood, E. S.; Noyes, F. R.; Zernicke, R. F.: Biomechanics of ligaments and tendons. In Hutton, R. [ed.]: Exercise and Sports Science Review, Vol. 6. Philadelphia: Franklin Institute Press, 1978; pp. 125–182.)

FIGURE 4-3 The effects of increasing tissue cross-sectional area on the shape of the load-deformation curve.

(From Butler, D. L.; Grood, E. S.; Noyes, F. R.; Zernicke, R. F.: Biomechanics of ligaments and tendons. In Hutton, R. [ed.]: Exercise and Sports Science Review, Vol. 6. Philadelphia: Franklin Institute Press, 1978; pp. 125–182.)

FIGURE 4-4 The influence of original length of tissue fibers on the shape of the load-deformation curve.

(From Butler, D. L.; Grood, E. S.; Noyes, F. R.; Zernicke, R. F.: Biomechanics of ligaments and tendons. In Hutton, R. [ed.]: Exercise and Sports Science Review, Vol. 6. Philadelphia: Franklin Institute Press, 1978; pp. 125–182.)

FIGURE 4-5 The effect of altering the collagen fiber alignment is shown. Parallel collagen fibers are loading uniformly, resulting in structural properties of increased strength, decreased elongation to failure, and increased stiffness in contrast to the oblique fiber orientation representing joint capsular tissues.

The arrangement of ligament fibers in a parallel manner such as the medial collateral ligament (MCL), fibular collateral ligament (FCL), or a tendon provides a structure with a high maximum load to failure because the majority of the fibers are loaded in a symmetrical manner and share the loading profile. The structure will fail at a relatively small displacement, as is true for the MCL and FCL, which are discussed in Chapter 2, Lateral, Posterior, and Cruciate Knee Anatomy. Alternatively, the capsular structures are designed to provide a low stiffness, allowing elongation or compliance, which is built into the microgeometry similar to that shown in the illustration of the finger-trap in Figure 4-5. Certain portions of the capsule may have an arrangement in which some fibers exist in a more parallel array, such as the posterior oblique ligament of the posteromedial capsule. Failure of capsular structures can be difficult to detect because a microtearing process occurs throughout the entire capsule. The capsule may appear slack and elongated without an obvious failure or the capsule may be avulsed off of its femoral or tibial attachment and the failure site identified and repaired at surgery. Prior to ultimate failure of the capsule, the microfailure process produces a residual elongation or slackening that, in chronic knee injuries, may require plication procedures to restore tension and function. An example is plication of posteromedial or posterolateral capsular structures so that they will function in full extension and resist knee hyperextension. The length-tension behavior and failure pattern of the ACL and PCL are also dissimilar, with different fiber regions brought into the loading based on the position of the knee.

An example of the force elongation properties of the MCL and medial capsular structures is shown in Figure 4-6. This example is from a rhesus primate knee in which the medial structures were loaded in uniaxial tension. The high failure load of the MCL in comparison with the capsular structures is shown. The capsule undergoes a failure process that is somewhat continuous because different portions of the less stiff capsule are brought into the loading sequence and fail, without an abrupt failure.

FIGURE 4-6 Typical oscillograph record of force vs. time for failure tests on the medial collateral ligament (MCL) and two capsular preparations. A constant distraction rate was used, so the time axis is proportional to specimen elongation. The MCL supports the largest forces and fails abruptly. This is in contrast to the progressive failure of the ACL. Note the reduction in stiffness that occurs for both the MCL and the capsule prior to significant failure indicated by the sudden drop in resisting force. This is indicative of a progressive microfailure process, which occurs prior to macroscopic failure.

(From Butler, D. L.; Grood, E. S.; Noyes, F. R.; Zernicke, R. F.: Biomechanics of ligaments and tendons. In Hutton, R. [ed.]: Exercise and Sports Science Review, Vol. 6. Philadelphia: Franklin Institute Press, 1978; pp. 125–182.)

An example of a scanning electron microscopic analysis of failure of the ACL ligament is shown in Figure 4-7. In this series of experiments, the ACL of a rhesus primate knee was loaded to one half of its expected failure load. Scanning electron microscopic examination showed that some portions of the fiber bundles had separated and had lost some of their normal wavy attachment. This is not an artifact, because normal ACL fibers will show a wavy appearance. In certain portions of the ACL, there is complete failure of individual fibers that were brought into the loading sequence owing to their fiber length at the knee position selected for the loading test. In Figure 4-8, failure of a major fiber bundle is shown with failure of the individual fibers at different places within the fiber. Note that failure of a small vessel has also occurred at the site of fiber failure, which is unusual given the expected elongation that vessels undergo prior to ultimate failure.

FIGURE 4-7 Scanning electron microphotographs of an ACL after it was loaded to one half its anticipated maximum load without evidence of failure. Microfailure occurs. A whole fiber bundle completely ruptured is shown at left (arrows). Separation of fibers within a bundle is shown at right.

(From Butler, D. L.; Grood, E. S.; Noyes, F. R.; Zernicke, R. F.: Biomechanics of ligaments and tendons. In Hutton, R. [ed.]: Exercise and Sports Science Review, Vol. 6. Philadelphia: Franklin Institute Press, 1978; pp. 125–182.)

FIGURE 4-8 Scanning electron photomicrograph of a collagen fiber bundle at the plane where failure occurred. The failure of collagen fibers at multiple levels within the fiber bundle demonstrates the pull-apart failure process. The arrows indicate a rich supply of vessels lying on the surface of the fiber bundle with perpendicular branches supplying the interior of the fiber bundle.

(From Noyes, F. R.; Grood, E. S.; Nussbaum, N. S.; Cooper, S. M.: Effect of intra-articular corticosteroids on ligament properties: a biomechanical and histological study in rhesus knees. Clin Orthop 123:197–209, 1977.)

IMPORTANCE OF FEMORAL ATTACHMENT ON LIGAMENT LENGTH-TENSION PATTERNS

Many factors determine ligament length patterns during knee motion; however, one of the most sensitive factors is the fiber’s femoral attachment (even more so than the tibial attachment). This concept is highly important for understanding ACL and PCL ligament fiber function and the location of ligament grafts for reconstruction.

The effect of the femoral attachment described by Grood and associates²⁸ is illustrated in Figure 4-9. In this model, the tibial attachment is assumed to be stationary, and flexion of the knee occurs about a fixed center of rotation (CR). The location of the CR is at the average rotation center of the human knee joint. A fiber that attaches anterior to the CR, such as fiber F_a, will lengthen with knee flexion up to 90° as it rises up to the top of the circle and thereafter shortens with knee flexion. The opposite effect is shown by fiber F_d, which shortens with knee flexion and, at 90° of flexion, would start to lengthen. The length patterns of the fibers are shown in the accompanying graph. This shows the complex length pattern of cruciate ligament fibers and is in marked contrast to more simplified descriptions that the ACL anteromedial fibers lengthen with knee flexion and the posterolateral fibers lengthen with knee extension, which is a gross oversimplification. The circle model is still only an approximation because the knee does not have a fixed CR, which provides an additional complexity in the analysis of ACL and PCL fiber function to be described.

FIGURE 4-9 Length patterns are determined by the femoral attachment location. A ligament fiber is shown with tibial attachment (T) and femoral attachment (Fa). Other femoral attachments are represented by points Fb through Fe. The femur is assumed to flex about a fixed center of rotation (CR). The graph shows how the distance between attachments (d) referenced to the distance between attachments at 0° (do) varies with flexion to 120°. The curves are solid in the region over which the ligament fiber is tense and dotted in the region over which it is slack. The functional region is obtained by assuming the fiber is strained 5% at its maximum length.

(From Grood, E. S.: Placement of knee ligament grafts. In Finerman, G. A.; Noyes, F. R. [eds.]: Biology and Biomechanics of the Traumatized Synovial Joint: The Knee as a Model. Rosemont, IL: American Academy of Orthopaedic Surgeons [AAOS], 1992; pp. 393–417.)

Contour plots are used to simplify the description of ligament length changes; a typical contour map is shown in Figure 4-10.³² A single contour line expresses all of the attachments that would produce the same maximum fiber strain. The region bounded by the contour line with the smallest strain (2% elongation) provides the most isometric femoral attachments. Note in Figure 4-10 that the tibiofemoral separation distance increases dramatically anterior to the contour line, and correspondingly, the length decreases posterior to the line with knee flexion. It should be noted that the concept of ligament isometricity is a misnomer because it is not expected that a ligament shows this behavior with a set minimal strain or elongation during knee flexion, as already discussed. A ligament functions with fibers of different lengths and their loading, which depends on attachment site characteristics. The disadvantage of the contour map is that the length of the individual native ligament fibers is not represented, and therefore, it is unknown at what point the ligament fibers in any region are tense or slack. The contour map is most useful to show the length changes that would occur for ligament grafts placed at certain positions within the ACL and PCL attachment sites at the maximum elongation or strain of the graft. In addition, note the knee position at which the graft is tensioned to not produce excessive graft elongation, which would either overconstrain the joint or result in graft failure.

FIGURE 4-10 Typical contour plot for the ACL. This figure was determined for knee flexion from 0° to 90° while an anterior force of 100 N was applied to the tibia. The tibial attachment site used in the analysis was located in the geometric center of the ACL’s tibial insertion.

(From Hefzy, M. S.; Grood, E. S.; Noyes, F. R.: Factors affecting the region of most isometric femoral attachments. Part II: the anterior cruciate ligament. Am J Sports Med 17:208–216, 1989.)

A number of studies show that there is an approximate region in the ACL or PCL at which one region lengthens with knee flexion, another region shortens, and fibers between these two regions undergo less elongation.^{8,28–30,32,61,72} However, there is a major discrepancy in the investigations regarding where the regions are located within the cruciate ligaments. The authors’ data and those of others⁷² have concluded that the line for the ACL is primarily oriented in a proximal-to-distal direction and that fibers anterior to the line become longer in flexion and posterior fibers become shorter (Fig. 4-11; see also Fig. 4-10). The region for the line for the PCL, in contrast, is oriented in an anterior-to-posterior direction, with distal fibers becoming longer with knee flexion and proximal fibers lengthening with knee extension. Thus, the line that divides the shortening and lengthening regions has a completely different orientation for the ACL than for the PCL. This is one of the most basic concepts of how ligament fibers function. Many current illustrations for ACL femoral attachment sites incorrectly show that the proximal fibers lengthen with knee flexion and the distal fibers lengthen with knee extension. In fact, there are fibers within the ACL proximal and distal femoral attachment regions, governed by the contour plot behavior (see Fig. 4-10), that do not show this reciprocal tension behavior with knee flexion and extension. The data for these studies are not based on true fiber attachment separation distances, and these differences have rather profound effects on where ACL and PCL grafts are placed and tensioned. The PCL fiber length-tension behaviors under loading are shown in Chapter 20, Function of the Posterior Cruciate Ligament and Posterolateral Ligament Structures.

FIGURE 4-11 The division of the femoral surface into shortening and lengthening regions. The line A-A, dividing the regions, has a different orientation for the two cruciate ligaments. Small errors in placement along the direction of the line have only small effects on length patterns. The approximate location of the line in reference to the posterior cruciate ligament (PCL) and ACL footprint is shown, with the qualification that the line changes its orientation with knee flexion. In contrast, placement errors perpendicular to the line cause large changes in length patterns.

(Redrawn from Grood, E. S.: Placement of knee ligament grafts. In Finerman, G. A.; Noyes, F. R. [eds.]: Biology and Biomechanics of the Traumatized Synovial Joint: The Knee as a Model. Rosemont, IL: American Academy of Orthopaedic Surgeons [AAOS], 1992; pp. 393–417.)

It should be noted that both Sidles and colleagues⁷² and Hefzy and coworkers³² reported that the contour maps describing fiber elongation patterns were strongly dependent on whether an anterior or a posterior displacement load was applied to the tibia and the contour maps provided were under loading conditions for the respective ACL or PCL. A more complete description of contour maps for the femoral and tibial attachments is provided for the reader,^28,32,72 and the effect of tibial attachment sites is discussed in Chapters 7, Anterior Cruciate Ligament Primary and Revision Reconstruction: Diagnosis, Operative Techniques, and Clinical Outcomes, and 20, Function of the Posterior Cruciate Ligament and Posterolateral Ligament Structures.

The effects of changing the tibial attachment of an ACL graft in cadaveric specimens is shown in Figure 4-12. The six ACL fiber locations on the tibia all originate at the approximate center of the ACL femoral attachment. Moving the fiber attachment on the tibia in a medial-to-lateral direction had only a small effect on fiber length and normalized length patterns. In contrast, moving the fiber in an anterior-to-posterior direction had a strong effect on fiber length. An anterior shift of the tibial attachment of 7.5 mm increased the fiber length by 5 to 8 mm. An anterior tibial attachment resulted in the fiber length behavior to become more taut in flexion. In Figure 4-13, the strong effect of moving the femoral attachment location in an anterior-to-posterior location on changing the length with knee flexion is shown. The location of the tibial insertion site is unchanged. The effect of moving the femoral attachment in a proximal-to-distal direction is shown in Figure 4-14. The more distal attached fibers were markedly shortened and had a concave length pattern, whereas the longer proximal fibers had a slightly convex pattern. In summary, these studies show the complex behavior of the cruciate ligaments in which individual fibers and fiber regions are loaded to resist coupled motions of tibial translation and rotation. For the surgeon, the placement of a graft for the ACL in an anterior-to-posterior location within its footprint has significant effects in determining the tension and loading conditions of the graft, as demonstrated. For the PCL, as is discussed in Chapter 20, Function of the Posterior Cruciate Ligament and Posterolateral Ligament Structures, placing the PCL graft in a proximal-to-distal direction has similar profound effects on its length-tension behaviors. The surgeon does have the ability to approximately select the range of knee flexion (low flexion, high flexion positions) under which the graft will be at its longest length and, therefore, should be tensioned at this position. Tensioning a graft at a flexion range at which the graft is in its shortest position would be expected to produce failure of portions of the graft as it elongates at a different knee flexion position. Within each graft, there will be asymmetrical loading of the graft fibers and the contour plots show the different fiber elongations of fibers in the center of the graft compared with the fibers at the periphery of the graft. The concept of placing more than one graft is designed to allow adjustments to more closely simulate ligament fiber behaviors, although it is probable that the surgeon is adding collagen fibers to fill the ACL or PCL footprint and that native cruciate fiber behavior is never achieved.

FIGURE 4-12 Sensitivity of fiber-length patterns to tibial attachment location for specimen 901.

(From Hefzy, M. S.; Grood, E. S.: Sensitivity of insertion locations on length patterns of anterior cruciate ligament fibers. J Biomech Eng 108:73–82, 1986.)

FIGURE 4-13 Sensitivity of fiber-length patterns to femoral attachment location for specimen 901.

(From Hefzy, M. S.; Grood, E. S.: Sensitivity of insertion locations on length patterns of anterior cruciate ligament fibers. J Biomech Eng 108:73–82, 1986.)

FIGURE 4-14 Sensitivity of fiber-length patterns to femoral attachment for specimen 898. Shown are the femoral attachments, which are near the anteroposterior (AP) midline of the anterior cruciate insertion.

(From Hefzy, M. S.; Grood, E. S.: Sensitivity of insertion locations on length patterns of anterior cruciate ligament fibers. J Biomech Eng 108:73–82, 1986.)

In agreement with Grood and associates’ studies,²⁸ Markolf and colleagues⁴⁷ reported in a cadaveric in vitro biomechanical study that positioning the femoral ACL tunnel in an anterior-to-posterior is more critical in terms of graft forces and function than is a proximal-to-distal direction.

LIGAMENT FIBER LENGTH-TENSION PATTERNS AND BURMESTER CURVES

As already discussed, one of the primary scientific bases for ligament fiber function and graft placement is the measurement or computation of fiber length patterns. An alternative approach for determining ligament function proposed by Menschik^50,51 is frequently referenced in the literature. This approach is based on the four-bar model of ACL and PCL ligament function shown in Figure 4-15 first proposed by Strasser⁷⁵ and on the mathematical theory for knee motions developed by Burmester.⁹ This mathematical theory was developed to determine all pairs of femoral and tibial attachments that remain the same distance apart (isometric) as the knee is moved through small (infinitesimal) motions. The pairs of attachment points are represented by the two curves shown in Figure 4-16, one for all possible femoral attachments and one for all possible tibial attachments, which are described as a cubic curve because of their mathematical form. A single tibial and femoral attachment of a ligament is obtained by drawing a straight line through the instant center of knee flexion controlled by the four-bar mechanism of where the cruciate links cross each other. The MCL tibial and femoral isometric attachments are the points at which a line crosses the tibial and femoral cubics.

FIGURE 4-15 Burmester curves for the four-bar mechanism. There are two separate curves, a femoral cubic (CF) and a tibial cubic (CT). Pairs of nearly isometric attachments can be obtained by passing a line through instant center (P) and extending the line until it crosses both cubics. The points CF and CT where the line crosses the cubics are nearly isometric. Other pairs of nearly isometric attachments can be found by rotating the line about the instant center. The lines shown here represent theoretical fibers of the MCL.

FIGURE 4-16 Burmester curves computed for flexion angles of 33°, 43°, and 53°. The curves for 43° of flexion are nearly identical to the curves published by Menschik^50,51 and Mueller.⁵⁴ Note that most of the tibial cubics for 33° and 53° of flexion are not on the tibia, and, thus they are not possible ligament attachments.

(From Grood, E. S.: Placement of knee ligament grafts. In Finerman, G. A.; Noyes, F. R. [eds.]: Biology and Biomechanics of the Traumatized Synovial Joint: The Knee as a Model. Rosemont, IL: American Academy of Orthopaedic Surgeons [AAOS], 1992; pp. 393–417.)

The Burmester curves have been used by many authors to explain the attachment locations of the collateral ligaments and capsular structures, and references are available proposing these curves can explain ligament function.⁵⁴ However, there are important scientific objections to using this model to define ligament function and possible graft attachment locations. The shape of the curves depends on the knee flexion angle chosen, and different curves can be computed for different flexion angles (see Fig. 4-16). Menschik^50,51 published the curves for 43° of flexion without providing the specific criteria used for computing the curves at this particular position of knee flexion. The authors of this chapter computed the curves for other flexion angles that differed markedly in shape and had different sites for the attachment of the collateral ligaments that may not be close to their anatomic attachment (see Fig. 4-16). In addition, the Burmester curves are based on small knee motions, and it has been determined that there are tibial attachment locations that are as isometric that are not on the Burmester curve when a wider range of knee motion is considered.

MECHANICAL PROPERTIES OF LIGAMENTS

The ligaments of the knee have unique anatomic and mechanical characteristics that cannot be generalized from one ligament to another. The mechanical behavior of ligaments depends on the material properties of the fibers, the geometric arrangement of collagen fibrils^18,21,65 and fiber bundles (as already described), the proportion of different types of fibrous constituents, and the effect of the surrounding ground substance.^{18,27,35,53,62} The ligament insertion site and underlying bone are additional portions of the ligament unit that must be considered when evaluating overall mechanical properties.

One approach for investigating each component of a ligament unit is to use a bone-ligament-bone specimen, such as a femur-ACL-tibia complex (FATC) preparation. This allows the mechanical properties of all of the components of the ligament unit to be studied together. The combined interaction of the components determines the properties of the entire structure in vivo. The problem still exists to load the bone-ligament-bone unit in a manner that allows in vitro data to be extrapolated to in vivo loading conditions. This does represent a problem because there are so many different in vivo loading conditions at the time of ACL failure that it is simply not possible to reproduce these conditions in the laboratory. There is less of a problem in studying the effect of one factor or variable on the mechanical properties of a ligament-bone unit because the same loading conditions are used throughout the experiment.

Effect of Strain Rate on Mechanical Properties

The strain rate describes the rate of deformation used in mechanical studies (length change/initial length per unit of time). The authors conducted a series of fast and slow strain-rate tests on ACL bone-ligament-bone specimens in rhesus monkeys.⁵⁶ A slow rate of deformation was chosen to represent the strain rates used in many prior in vitro experiments. A fast rate of deformation was chosen to represent the physiologic in vivo loading conditions to which ligaments may be subjected. The ligament units were tested as right-left pairs at the two strain rates to exclude animal variability. The femur and tibia were at an angle that corresponded to 45° of knee flexion to allow as nearly uniform loading of the ligament as possible. In this experiment, a distraction of the joint was produced. An alternative loading sequence would be to induce anterior translation. Both of these loading profiles have been used experimentally.

Dividing the extension rate by the initial length of the ligament produced approximate strain rates of 0.662 sec^–1 at the fast rate and of 0.00662 sec^–1 at the slow rate, a 100-fold difference. At the fast rate, the ligament underwent the initial 15% to 20% elongation in approximately 0.25 second.

An example of a load-versus-time record from a typical test at the fast strain rate is shown in Figure 4-17. The curves demonstrate an initial concave toe region, followed by a fairly linear region up to the first significant failure. At this point, the load is termed the linear load because it denotes the approximate point at which the ligament preparation enters into the major failure region. Minor failures occur first, with a drop in load. Decreases in load after the linear load indicate successive failures until complete failure, at which time the load falls to zero. The maximum load occurs in the major failure region. The time axis is proportional to the relative displacement of the grips and represents elongation of the ligament after the compliance of the testing device and the grips are computed. Because the extension rate is known, specimen elongation at the linear load, maximum load, and zero load (complete failure) is determined. The area under the load-deformation curve indicates the energy absorbed by the specimen up to complete failure. The slope of the curve in the linear region provides the approximate stiffness of the preparation. It was not possible in this experiment to calculate the cross-sectional area of the ACL with any degree of reproducible accuracy. Force values are reported as load rather than as stress (force per unit area). Alternative experiments, discussed later, employ other testing configurations in which ligament strain is measured by sensitive optic techniques.

FIGURE 4-17 Oscillograph record of force vs. time for a tension test to failure of a rhesus femur-ACL-tibia preparation. A constant distraction rate was used so that the time axis was proportional to specimen elongation. The photographs, obtained from high-speed movies taken during the test, show the preparation at four stages of the test.

(From Butler, D. L.; Grood, E. S.; Noyes, F. R.; Zernicke, R. F.: Biomechanics of ligaments and tendons. In Hutton, R. [ed.]: Exercise and Sports Science Review, Vol. 6. Philadelphia: Franklin Institute Press, 1978; pp. 125–182.)

The anterior cruciate bone-ligament-bone specimens showed a complex and varied mode of failure. These included avulsion of minor to major bone fragments from insertion sites, failure at the bone-ligament interface without bone avulsion, and ligament failure that occurred initially by tensile pull-apart failure of fiber bundles and then by shear failure between the disrupted fibers. The predominant mode of failure of each specimen was classified as one of the following: (1) ligamentous failure, (2) tibial avulsion failure, (3) combined ligament failure–tibial avulsion fracture, (4) femoral avulsion fracture, or (5) combined ligament failure–femoral avulsion fracture.

A statistically significant difference was found in the predominant mode of failure of the specimens tested at the fast strain rate compared with those tested at the slow rate (P < .05; Fig. 4-18). At the fast rate, two thirds of the specimens demonstrated a ligamentous failure and 28% failed by tibial avulsion fracture. At the slow rate, 57% of the specimens failed by tibial avulsion fracture and 29% failed by ligamentous failure.

FIGURE 4-18 The major mechanism of specimen failure is shown for the 32 knees tested at the fast strain rate and the 28 knees tested at the slow strain rate. The difference in specimen failure at the two strain rates is statistically significant (P < .05).

(From Noyes, F. R.; DeLucas, J. L.; Torvik, P. J.: Biomechanics of anterior cruciate ligament failure: an analysis of strain-rate sensitivity and mechanisms of failure in primates. J Bone Joint Surg Am 56:236–253, 1974.)

The results of the strain-rate properties according to the modes of failure are shown in Table 4-1. Under slow strain-rate conditions, the bone-insertion area of the tibia was the weakest component, demonstrated by the reduced maximum load and energy absorbed to failure. At the fast strain rate, representing the more physiologic loading condition, the ligament and bone components failed at a similar maximum load and energy. The overall difference in strength properties was due in part to increased strength of the bone component at the fast rate. At the slow rate, failure by tibial avulsion fracture was interpreted as premature failure, indicating a decrease in the specimen’s potential strength. The size of the animals was not a factor contributing to the predominant mode of failure. The different types of failure were distributed throughout each strain-rate group regardless of the weight of the animal. In addition, juvenile animals were not used, which could have biased the results toward a higher frequency of tibial avulsion fractures.

TABLE 4-1 Strain-Rate Results by Failure Mode

Analysis of Failure-Mode Mechanisms

Macroscopic Analysis

High-speed films of specimen loadings demonstrated that the mechanism of failure usually involved both the bony and the ligamentous components in a progressive manner until complete failure occurred. Sequential failure of ligament fibers commonly occurred first, prior to major tibial avulsion fractures. In other specimens, minor avulsion fractures were associated with major ligamentous failure (Fig. 4-19).

FIGURE 4-19 Photomicrograph after tibial avulsion mode of failure shows that cleavage occurred through cancellous bone below the more dense cortical bone at the site of ligament insertion. The zonal arrangement at the bone-ligament interface is shown. Ligament, zone 1; fibrocartilage, zone 2; mineralized fibrocartilage, zone 3; and bone, zone 4 (hematoxylin and eosin, x90).

(From Noyes, F. R.; DeLucas, J. L.; Torvik, P. J.: Biomechanics of anterior cruciate ligament failure: an analysis of strain-rate sensitivity and mechanisms of failure in primates. J Bone Joint Surg Am 56:236–253, 1974.)

Microscopic Analysis

Microscopic analyses were conducted after the loading tests to define the mechanisms of specimen failure. Resorption changes at the bone-ligament interface, indicative of premature failure, were not visualized in any of the femoral or tibial insertion sites. Polarized light microscopy demonstrated that the ACL inserted at femoral and tibial sites through four well-defined zones of fibrocartilage (see Fig. 4-19).¹⁷ The zonal arrangement at the insertion site represents a change in composition of the surrounding medium as the ligament collagen fibers (zone 1) pass through fibrocartilage (zone 2), across a blue line to mineralized fibrocartilage (zone 3), and into the bone (zone 4).

The specimens in this investigation demonstrated three histologic modes of failure, which were consistent with the macroscopic observations. First, ligamentous failure occurred by an initial tensile failure of collage fibers, followed by a pulling-apart shear failure of the disrupted fibers (Fig. 4-20). Rupture of collagen fiber bundles was also observed throughout the ligament. The ligament fiber microgeometry under the loading conditions in the experiment determined which fibers were subjected to the greatest deformation leading to the serial rupture of fibers and accounted for rupture of fibers in areas adjacent to intact fibers. Although branching between individual collagen fiber bundles was observed, the pulling-apart type of failure suggests that little in the way of cohesive properties exists between the major fiber bundles. A more explicit definition of the mechanism of failure of collagen tissues also involves the ultrastructural changes of fibrils and microfibrils under failure conditions.⁵⁶

FIGURE 4-20 Photomicrograph of a portion of the ACL after failure. The mechanism of failure appeared to be a tensile, pulling-apart type of failure. The collagen fibers are shown to have ruptured at different portions of the ligament, giving an uneven cleavage line corresponding to the common “mop-end” gross appearance (trichrome, x45).

(From Noyes, F. R.; DeLucas, J. L.; Torvik, P. J.: Biomechanics of anterior cruciate ligament failure: an analysis of strain-rate sensitivity and mechanisms of failure in primates. J Bone Joint Surg Am 56:236–253, 1974.)

Second, avulsion fracture at the attachment site occurred most commonly through the cancellous bone immediately beneath the dense cortical bone (Fig. 4-21). The minute gritty material sometimes just barely palpable on the end of the ligament after failure represented fragments of bone beneath the cartilaginous zone of ligament insertion.⁵⁶

FIGURE 4-21 Photomicrograph shows predilection for cleavage at the ligament-bone interface to occur through the mineralized fibrocartilage zone, in this case just distal to the blue line. Inset, Columnar arrangement of chondrocytes (periodic acid–Schiff, x80).

(From Noyes, F. R.; DeLucas, J. L.; Torvik, P. J.: Biomechanics of anterior cruciate ligament failure: an analysis of strain-rate sensitivity and mechanisms of failure in primates. J Bone Joint Surg Am 56:236–253, 1974.)

Third, cleavage at the ligament-bone interface showed a predilection for the fracture cleavage line to occur through the zone of mineralized fibrocartilage (zone 3) at or just distal to the blue line. Failure at this interface was the least common mode of failure. It was identified microscopically in 6 of 22 specimens; however, it represented only a minor part of the failure process, the major failure having occurred by either a ligamentous or a bone avulsion mode. Failure through the fibrocartilaginous zone was, therefore, never the major mode of specimen failure, but not infrequently accompanied the other two modes of failure.⁵⁶

In summary, it was possible to classify the major mechanism of failure histologically, although minor modes of failure involved to different degrees all portions of the bone-ligament-bone unit. Thus, it was not unusual to find microscopic evidence of minor to moderate ligament disruption associated with major bone avulsion failures or vice versa.⁵⁶

This study showed for the first time that an ACL-ligament-bone preparation failed at a higher load and at greater elongation, and absorbed significantly more energy, at a fast rate of deformation than at a slow rate. These findings show the time-dependent behavior of bone-ligament-bone preparations and questioned the results of prior ligament failure studies that often showed premature bone failure at less than expected failure loads. At the slow deformation rate, the bone insertion area of the tibia was the weakest component. At the fast deformation rate, the ligament and tibial osseous component were balanced as to strength properties with an increased frequency of ligamentous failure.⁵⁶ The specimen under the fast-deformation rate withstood greater loads without premature osseous failure, allowing the ligament to participate in the failure process.⁶⁷ Even with expected viscoelastic effects in the ligament as a result of the higher strain rate, the strength of the bone increased faster than did an expected increase in strength of the ligament.⁵⁶

Many cadaveric and experimental animal studies have reported the difficulty of producing ligamentous injuries and have traditionally stated that the bone is the weakest component of the bone-ligament-bone system.^{34,52,73,78,7984,85,93} Other than the difficulty in reproducing in the laboratory the complex loading and displacement conditions that occur in traumatic injuries, the deformation rate of the experimental tests and the effect of disuse-induced bone changes due to animal caging or aged cadavers are believed to be the major explanations for premature bone avulsion of the ligament-bone unit.⁵⁶

In agreement with these results, Welsh and associates⁸⁹ demonstrated that the failure mode of a calcaneal–Achilles tendon unit changed from bone avulsion at low deformation rates to failure at the tendon-grip apparatus under high deformation rates. The ultimate load of the specimens in that investigation may not have been obtained because of the difficulty of clamping the tendons; however, the bone insertion site did not fail at the high strain rate despite the specimen reaching a higher load.

In cases of tibial avulsion fracture, the ligament failure component could not be appreciated on gross examination after the test. In view of the demonstrated elongation on the ligament in the region of failure, it is reasonable to expect damage to collagen fibers and blood vessels within the ligament, even though continuity remains. It is well appreciated that the visual determination of continuity of a ligament at surgery provides an inadequate determination of the extent of ligament damage that has actually occurred and its functional capacity. Such factors, among others, may contribute to the occasional unsatisfactory result in cases of partial ligament disruption with continuity and in cases of ligament reinsertion after bone avulsion.⁵⁶

The fibrocartilaginous zone at the ligament insertion site is believed to be advantageous in producing a gradual change in mechanical properties, decreasing the stress-concentration effect of the ligament’s insertion into the stiffer bone structure.¹⁷ This zone protects against fatigue and shear failure at the bone-ligament junction, providing a composite structure with the addition of cartilage and mineralized fibrocartilage to the collagen network. The fibrocartilaginous ground substance may provide greater cohesion between fiber bundles and provide a mechanism for diffusion of load over the entire insertion site, avoiding the deleterious effects of stress concentration. Based on the microscopic analysis of specimen failure and the infrequent occurrence of failure through the fibrocartilaginous zone, the zonal insertion arrangement appears to be mechanically advantageous.⁵⁶

EFFECTS OF IMMOBILIZATION AND DISUSE ON LIGAMENT BIOMECHANICAL PROPERTIES

It is well appreciated that immobility may lead to deleterious changes in bone, joint, and soft tissue.* The authors conducted a study to determine the effect of altered activity levels on the mechanical properties of ligaments in primates.⁶⁰ Whereas exercise is accepted as beneficial in preventing muscle atrophy and maintaining joint motion, its effect in preventing deterioration or strengthening a bone-ligament-bone unit had not been established. The study simulated certain clinical states of immobility and determined: (1) the effect of immobilization on the mechanical properties of a bone-ligament-bone unit, (2) the effect of an exercise program in preventing changes in these properties, and (3) the extent to which the changes remained after a 20-week reconditioning period. The reader is referred to the publication for the detailed explanation of the study.⁶⁰

The rhesus animals were placed into four groups: (1) group I, control group, 30 knees; (2) group II, 9 animals immobilized for 8 weeks in total body plaster casts that included both lower limbs; (3) group III, 11 animals in which one lower limb was immobilized for 8 weeks and the other lower limb was exercised daily; and (4) group IV, 11 animals immobilized as in group II, reconditioned in room-sized gang cages for 5 months that were wide and deep and provided ample room for jumping and other activities.

The animals in exercise group (group III) performed an active pushing movement against resistance with one free lower extremity in response to a food-reward system. The pushing movement was from 90° of hip and knee flexion to 20° of hip and knee flexion. The foot and ankle were secured by a boot to an exercise apparatus, and with each successful stroke, the animals received a pellet of food. The animals performed the exercise diligently, and on occasion, food had to be restricted. The force required to perform the exercise was minimal initially and then gradually increased over the first 2 weeks so that the animals performed the task at least 600 times daily. The resistance, length of push, and number of repetitions were recorded daily. The animals usually depressed the apparatus 600 to 900 times daily against a resistance of 1.6 to 2.4 kg (about one third body weight).

The average daily work performed by each primate was 11,000 ± 4000 kg force-cm (range, 6600–20,200). The average total work performed by each animal during the 8 weeks of immobilization was 6.19 x 10⁵ ± 2.4 x 10⁵ kg force-cm (range, 3.69 x 10⁵–1.11 x 10⁶ kg force-cm). There was no significant correlation between the total amount of work each animal performed and the strength of the ligament preparation.

The results showed that 8 weeks of immobilization produced a marked decrease of nearly one half of the maximum failure load (P < .001; Table 4-2). No significant differences were found in ultimate failure load between the immobilized or the exercise group or between the right and the left ligament units from the exercised group. The failure load for the reconditioned animals indicates partial but incomplete recovery at 20 weeks (Fig. 4-22).

FIGURE 4-22 Calculated load-deformation curves show ligament behavior under loading for all specimens in each of the animal groups. A significant decrease in stiffness (slope of the load-deformation curve) of the ligament preparation is seen in specimens from the exercised and immobilized group. Partial recovery occurred at 20 weeks.

(From Noyes, F. R.; Torvik, P. J.; Hyde, W. B.; DeLucas, J. L.: Biomechanics of ligament failure. II. An analysis of immobilization, exercise, and reconditioning effects in primates. J Bone Joint Surg Am 56:1406–1418, 1974.)

The data demonstrate that immobilization leads to significant weakening of a functional ligament unit. The exercise program had little effect in preventing the decline in the strength properties of the ligament unit during immobilization, and only partial recovery in strength properties occurred after 20 weeks of resumed activity. The decrease in strength properties was associated with resorption of haversian bone and weakening of the cortex beneath the ligament’s insertion site at both the tibia and the femur.

The mechanism of failure of a ligament-bone unit after disuse-induced changes depends on the anatomic characteristics at the ligament-bone insertion site. Most knee ligaments insert into bone through the well-defined zones of fibrocartilage previously described.^17,46,56,68 These zones appear to be protective, because subsynovial and subperiosteal (femoral site) resorption of bone was observed at the peripheral margins of the ligament insertion site but not at the ligament-bone junction. Ultimate failure of the ligament unit occurred either through the underlying cortical bone at the insertion site or through the body of the ligament (or at both of these locations), but not through the ligament-bone junction. This protective effect of the fibrocartilage probably also applies to tendons that have a similar type of attachment.^17,68,74 The zone also imposes a barrier to a vascular supply of the ligament from the bone beneath the insertion site.⁶⁸ The tibial attachment of the long fibers of the MCL is into the periosteum and Sharpy fibers without a fibrocartilaginous interface. Immobility has a pronounced effect on producing subperiosteal bone reabsorption that decreases the tibial MCL attachment site strength, resulting in premature failure prior to involving the ligament in the failure process.^41,43,46 Because of the premature attachment failure, the use of a MCL bone-ligament unit is not ideal and disuse-induced cage confinement of experimental animals induces this as a variable affecting the results. The authors agree with Laros and coworkers⁴⁶ that the reported increase in ligament strength above control values resulting from increased activity in animals may represent the prevention or reduction of the disuse effect of caging.

The effect of immobility on the ligament itself is of importance to its mechanical function. The change in the relationship between force and ligament elongation (stiffness) after immobilization correlated with the inactivity of the primates. The exercised ligament units were somewhat less affected than the fully immobilized specimens. Both groups were significantly different compared with the control group. The stiffness of the ligament unit in the reconditioned specimens had nearly returned to normal at 20 weeks, although the recovery in strength was incomplete. Other studies have shown a similar correlation between the general level of activity during confinement and the strength of a ligament unit.^{1,46,78,80,85,93}

The clinical relevance of this study relates to the suggestion that after immobility, an extended period of time may be required before the functional capacity of a ligament unit returns to normal. Disuse-induced changes after fracture or ligament reconstruction may extend well into the time period after which normal activity has been resumed by the patient.

STRENGTH OF THE ACL: AGE-RELATED AND SPECIES-RELATED CHANGES

Important changes occur in the functional properties of bone-ligament units with age and, as well, differences from one animal species to another. The authors⁵⁷ conducted a study in which the mechanical properties of ACL specimens from humans and rhesus monkeys were determined in tension to failure under high strain-rate conditions. The age of the cadaver specimens ranged from 16 to 86 years. One of the purposes of the study was to determine whether discrepancies in strength and mechanisms of ligament failure between human and animal specimens were due to size differences, to variables in the human specimens such as age or disuse atrophy, or to certain experimental testing variables. In addition, human data are required to define the strength properties required for ACL graft replacements.

Comparison of Structural Parameters

A typical force/elongation curve for one young and one older adult human specimen is shown in Figure 4-23. The mean values for the stiffness and strength parameters of the human and rhesus ligament specimens are shown in Table 4-4. A significant difference was found between the older and the younger human specimens. The preparations from the younger humans failed at a maximum force that was an average of 2.4 times that of the specimens from the older humans. The ligaments from the animals failed at force values higher than those of the older human specimens, despite an approximate fivefold difference in the respective cross-sectional areas of the ligaments.

FIGURE 4-23 Typical oscillograph force-vs.-time records for a younger and an older human specimen show behavior of the ACL preparation in the mechanical test. The specimen from the older human demonstrates a decrease in stiffness (slope of the curve) and failure at lower force and strain values. Failure in the specimen from the older human occurred by femoral avulsion fracture just beneath the ligament attachment site in contrast to the ligamentous mode of failure in the preparation from the younger human.

(From Noyes, F. R.; Grood, E. S.: The strength of the anterior cruciate ligament in humans and rhesus monkeys. Age-related and species-related changes. J Bone Joint Surg Am 58:1074–1082, 1976.)

TABLE 4-4 Anterior Cruciate Ligament Strength: Comparison of Structural Properties

Comparison of Material Properties

The data shown in Table 4-5 represent measurements of the mechanical properties of the ligament collagen as a material. They are normalizations of the structural parameters, adjusted for variations in the cross-sectional areas and lengths of the specimens. There were large differences in all parameters shown between ligaments from younger and older humans. The mean value for the elastic modulus of the ligaments from young adult humans was 1.7 times that of the ligaments from the older humans, but was significantly lower than the mean value for the ligaments from the animals.

TABLE 4-5 Anterior Cruciate Ligament Strength: Comparison of Material Properties*

The maximum stress and the strain energy to failure in the specimens from younger humans were 2.8 and 3.3 times the respective values for the older human group. The maximum stress and the strain energy to failure in the ligaments from the animals were 1.8 and 1.9 times the respective values for the specimens from younger humans.

Critical Points STRENGTH OF THE ACL: AGE-RELATED AND SPECIES-RELATED CHANGES

The authors conducted a study in which the mechanical properties of ACL specimens from humans and rhesus monkeys were determined in tension to failure under high strain-rate conditions.

The purposes of the study were to determine whether discrepancies in strength and mechanisms of ligament failure between human and animal specimens were due to size differences, to variables in the human specimens such as age or disuse atrophy, or to certain experimental testing variables.

The ligament lengths and areas of the animals were approximately one half and one fourth those in the human donors, respectively.

Significant differences were found between the older and the younger human specimens in stiffness, strength, and mechanical properties. Preparations from the younger humans failed at a maximum force that was an average of 2.4 times that of the specimens from the older humans.

Ligaments from the rhesus animals failed at normalized force values (to body mass) over three times higher than those of the younger human specimens.

The specimens from older humans showed a decrease in cortical thickness and trabecular bone at the insertion of the ligaments. Failure occurred by fracture through the cortical and underlying trabecular bone, owing to both age-related and probably disuse-induced changes.

Age-related Changes

Statistically significant decreases were found with age in the elastic modulus, maximum stress, and strain energy for the specimens that failed by a ligamentous mode, in which prior antemortem effects were excluded, but not for the specimens from older humans that failed by premature bone–avulsion fracture, in which significant antemortem effects were suspected.

The regression equation (y) and correlation coefficient (r) for the specimens that failed by a ligamentous mode are shown in Figure 4-24. Maximum stress showed the highest correlation with age (R = 0.863; P < .005). A low correction was found with strain energy to failure (0.75) and elastic modulus (0.712), but all correlations were statistically significant.

FIGURE 4-24 Correlation of ligament strength with age. The solid line represents the statistically significant correlation with age found in the trauma specimens (orange triangle) and younger cadaver preparations (yellow triangle), all of which failed through the body of the ligament. The interrupted line represents the correlation (not statistically significant) found in the amputation (blue circle) and older cadaver (red circle) preparations, which failed by avulsion fracture of the bone underneath the ligament insertion site.

(From Noyes, F. R.; Grood, E. S.: The strength of the anterior cruciate ligament in humans and rhesus monkeys. Age-related and species-related changes. J Bone Joint Surg Am 58:1074–1082, 1976.)

Comparison of Ligament Structural Properties

Alm and colleagues² conducted a study on the tensile strength of the canine cruciate ligament in which the data was normalized by calculating the ratio between ligament separation force and animal mass. Ratios were reported ranging from 80 to 95 N/kg body mass. Similar ratios ranging from 80 to 118 N/kg body mass may be calculated from data reported by Viidik^84,85 in the rabbit model. Smith⁷³ reported a ratio that ranged from 40 to 80 N/kg of body mass in rabbits. A ratio of 119 N/kg body mass was found in the rhesus monkey preparations in the authors’ study, and the specimens from older and younger humans had ratios of 10 and 33 N/kg of body mass, respectively. The human maximum force ratios were markedly low in comparison with those of the rhesus animals as well as other animal models.

It may be invalid, however, to conduct such comparisons based on body mass. The functional demands placed on ligaments vary according to the mode of locomotion, and one cannot compare the stresses of the bipedal gait of humans with the gaits of rabbits, monkeys, or dogs. In addition, the normalization of ligament strength as a function of body mass does not show parallel findings with respect to the variations in size of the ligaments, regardless of which scaling rule is applied. Differences in experimental protocol and testing procedures, such as strain rates, angles of knee flexion, and methods used in gripping the specimens, also makes comparisons between studies difficult.

Age-related Changes

Whereas it is reasonable to expect an alteration in ligament strength properties with age, our data found greater than expected declines in strength from the younger human specimens compared with the older human preparations (ultimate failure means, 1730 N and 734 N, respectively). The mean body mass of the younger donors was 53 kg; even higher values would be expected for individuals of greater body mass. In the failed specimens from the older humans, there was a decrease in thickness of the cortical and trabecular bone beneath the fibrocartilaginous ligament-bone attachment site. The underlying trabecular bone had wider spaces between trabeculae, and the individual trabeculae appeared thinner than in the specimens from the younger humans. Failure of the ligament-bone unit occurred by an osseous failure owing to both age-related and probably disuse-induced changes.

Kennedy and colleagues⁴⁴ reported an ultimate tensile strength of 626 N for human specimens obtained from donors with a mean age of 62 years. Trent and coworkers⁸¹ reported a range of 285 to 1718 N in five cadaver specimens with a donor age of 29 to 55 years. Woo and associates⁹¹ evaluated the structural properties of 27 pairs of human cadaver knees to determine effects of specimen age on the tensile properties of the FATC. Tensile tests of the bone-ligament unit were performed at 30° of knee flexion with the ACL aligned vertically along the direction of applied tension load. One knee from each pair was oriented anatomically, and the second knee was oriented with the tibial aligned vertically. Specimen age did not affect anteroposterior (AP) displacement tests in intact knees at 30° and 90° of flexion. Structural properties of linear stiffness, ultimate load, and energy absorbed significantly decreased with specimen age (Fig. 4-25). Specimen orientation affected these properties, because higher values were found in specimens tested in the anatomic orientation (Table 4-6).

FIGURE 4-25 The effect of specimen age on femur-ACL-tibia complex (FATC) ultimate load. Data on ultimate load as a function of specimen age and orientation using a least squares curve fit demonstrated that the strength of the FATC decreases in an exponential manner.

(From Woo, S. L.-Y.; Hollis, J. M.; Adams, D. J.; et al.: Tensile properties of the human femur–anterior cruciate ligament–tibia complex. The effects of specimen age and orientation. Am J Sports Med 19:217–225, 1991.)

TABLE 4-6 Effects of Specimen Age and Orientation on the Structural Properties of the Femur–Anterior Cruciate Ligament–Tibia Complex*

It is well known that humans begin to lose bone mass beginning in the 4th decade of life. This phenomenon has important implications in the strength of the osseous components of ligament units for individuals in whom the process is advanced or when it occurs in association with other disuse or disease states in the presence of osteopenia.

Other studies of age-related changes in the mechanical properties of collagenous soft tissues have shown increases in tensile strength and stiffness and decreases in elongation properties with aging.^86,87 However, these studies involved young animals, and the results reflect a maturation process rather than aging, which were probably related to changes in insoluble collagen, increased intermolecular and intramolecular cross-linking, and increases in the collagen-to-glycosaminoglycan and collagen-to-water ratios.⁷³ The physiochemical and mechanical changes that occur in collagen after completion of maturation, which can be ascribed to an aging process in the absence of adverse environmental or disease factors, have not been defined.

EFFECT OF INTRA-ARTICULAR CORTICOSTEROIDS ON LIGAMENT PROPERTIES

The effect of an intra-articular, slightly soluble corticosteroid (methylprednisolone acetate) was investigated on the mechanical properties of a ligament unit.^58,59 ACL preparations from wild rhesus animals were used to determine the effect of corticosteroid dosage and time duration after administration.

Intra-articular Steroid Injection

In the second investigation, two animal groups were studied.⁵⁹ The first group (high corticosteroid dosage) comprised 10 animals that received a total of three intra-articular injections of methylprednisolone acetate (6.0 mg/kg) into one knee. The injections were spaced 1 week apart. A sham procedure was performed on the opposite knee that consisted of an intra-articular saline injection in equal volume to that of the drug. These animals were sacrificed 6 weeks after the first injection. The second group (low corticosteroid dosage) comprised 12 animals that received two injections of a dosage one tenth that used in the other animal group (methylprednisolone acetate 0.6 mg/kg). The two injections were given 2 weeks apart, with a sham saline injection administered into the opposite knee. The animals were sacrificed 15 weeks after the first injection.

A load-versus-time record from a typical right and left specimen pair in an animal from the high corticosteroid dosage group is shown in Figure 4-26. The strain parameters are approximated owing to inherent errors in measuring ligament length and in using the tibia-femur separation distance as a measure of ligament elongation during testing.

FIGURE 4-26 Typical load-vs.-time records for a right-left pair of specimens. The sham specimen demonstrates normal behavior under loading. The steroid side demonstrates a decrease in stiffness (initial slope of the curve), failure at lower loads, and an increase in ligament elongation at complete failure.

(From Noyes, F. R.; Grood, E. S.; Nussbaum, N. S.; Cooper, S. M.: Effect of intra-articular corticosteroids on ligament properties: a biomechanical and histological study in rhesus knees. Clin Orthop 123:197–209, 1977.)

The histologic analysis of the failed ligament units showed no detectable corticosteroid effect in that the fibrocartilage zones at the ligament insertion sites were intact. There were no bone resorptive changes at the ligament attachment sites. On scanning electron microscopy, no alterations were found in the surface appearance of collagen fibers and fibrils in the ligaments that received the corticosteroid.⁵⁹

The alterations in ligament behavior were dependent on the dosage of the drug and the time that has elapsed after injection. In the higher-dosage group, the maximum failure load of the ligament unit decreased by 11% after 6 weeks and 20% after 15 weeks. The stiffness of the ligament preparation showed a minimal decrease 6 weeks after injection; however, at 15 weeks, a significant decline had occurred. In the lower-dosage group, the maximum failure load of the ligament unit decreased by 9% and the energy absorbed prior to failure decreased by 8%. The decreases were found when the right-left specimen pairs were compared and were statistically significant. No significant change occurred in the stiffness of the ligament unit. These small changes suggest only slight alterations in the projected functional capacity of the ligament unit.

The specific ultrastructural mechanisms responsible for the alterations in ligament mechanical properties after administration of corticosteroid are unknown. There are known inhibitory effects of corticosteroids on the synthesis of glycosaminoglycans, proteins, and collagen. Whether decreased collagen synthesis occurs owing to the drugs’ effect on specific collagen precursors and enzymes or from a general inhibitory effect on overall protein synthesis is unknown. Corticosteroid preparations are known to alter fibroblast proliferation and metabolism. The cellular and subcellular effects are highly complex and beyond the scope of this chapter. The end result of these drugs is generally antianabolic and responsible for the adverse effect on wound and soft tissue healing.

The clinical implications of this study are suggestive that intra-articular, slightly soluble corticosteroids in high and frequent doses have the potential to alter ligament strength and function. The risk associated with such alterations may be minimal if infrequent, low-dose injections are administered. Alternatively, a single direct injection of a slightly soluble steroid preparation has a marked effect on decreasing ligament functional properties, which remained up to 1 year.

Wiggins and colleagues⁹⁰ and Walsh and coworkers⁸⁸ reported extensive studies on the effect of a steroid injection in human equivalent doses on the healing of a rabbit MCL model. Their data show statistically significant reductions in maximum failure loads and alteration in histologic properties compared with appropriate controls. The authors concluded that a single steroid injection at the site of a healing ligament would decrease or retard healing properties; however, they suggested that the effect was probably reversible. The authors reviewed a number of mechanisms for the deleterious effects. It should be noted that the studies were performed on a rabbit model that exhibits a known hypersensitivity to steroids that is more pronounced than that expected in other animals and humans.

EFFECT OF VASCULARIZED AND NONVASCULARIZED ACL GRAFTS ON BIOMECHANICAL PROPERTIES

After ACL graft reconstruction, there is marked cellular invasion that removes necrotic collagenous tissues with replacement collagenous fibers that are a ligament in a parallel fashion, but do not resemble the native microgeometry of the ACL. The question that has been posed is whether maintaining a vascular supply (and any concomitant neural innervation) of an ACL graft would limit the so-called necrotic phase of ligament remodeling and retain more normal ligament functional properties. An investigation was conducted to measure the mechanical properties of patellar tendon autografts used to replace the ACL in the cynomolgus monkey at four time periods up to 1 year.¹¹ The ACL was replaced with the medial half of the patellar tendon as a vascularized graft in one knee and as a nonvascularized (or free) graft in the contralateral knee. The 5-mm-wide medial patellar tendon free grafts were prepared through a standard medial parapatellar exposure. The vascularized grafts were prepared in a similar manner, except the medial retinacular vascular supply and a portion of the fat pad were preserved. The tibial attachment sites were prepared by producing a thin bony trough in the anteromedial aspect of the tibia to the site of the ACL attachment, and the ACL grafts with the vascularized pedicle were transposed into a bony trough. An identical tibial and femoral placement was used for the free grafts. Standard graft fixation methods were used. The knees were immobilized for 4 weeks at 30° of flexion followed by unrestricted activity in large cages.

Both the vascularized and the nonvascularized grafts underwent significant reductions in structural mechanical and material properties as early as 7 weeks after surgery. Stiffness values were as low as 24% to 28% of control ACL and medial patellar tendon values. Maximum force at 7 weeks was 16% of control ACL values, increasing to 39% of control maximum force by 1 year (Fig. 4-27A). Modulus and maximum stress showed even greater reductions and were only 34% and 26%, respectively, of control ACL values at 1 year (see Fig. 4-27B). The rate of return was slow up to 1 year for all variables studied. The changes in maximum force over time were generally consistent with other studies that used free grafts, showing on a mechanical basis that an extended period of time is required for ACL graft remodeling and partial return of strength. The study concluded that retention of a portion of the blood supply to the ligament is not sufficient to avoid a major reduction in ligament mechanical properties that occur in vivo after implantation.

FIGURE 4-27 A, Normalized graft stiffness and maximum force (1 standard error of the mean [SEM]) are plotted against weeks postoperation. Note the more rapid increase in the stiffness values over time. By 1 year, stiffness and maximum force were 57% and 39% of control, respectively. B, Normalized graft modulus and maximum stress are plotted against weeks postoperation. No significant differences are present between the parameters at any time period. Modulus and maximum stress are lower than corresponding stiffness and maximum force in A, suggesting that the collagen material is both very compliant and weak.

(From Butler, D. L.; Grood, E. S.; Noyes, F. R.; et al.: Mechanical properties of primate vascularized vs. nonvascularized patellar tendon grafts; changes over time. J Orthop Res 7:68–79, 1989. Reprinted with permission of Wiley-Liss, Inc., a subsidiary of John Wiley & Sons, Inc.)

EFFECT OF CELLULAR NECROSIS AND STRESS-SHIELDING ON ACL HEALING AND RETENTION OF NATIVE ACL FIBER MICROARCHITECTURE

Jackson and associates^37–39 performed a series of landmark studies in a goat model in which the ACL was frozen in situ, producing cellular necrosis while retaining the normal fiber microarchitecture. Mechanical loading tests of the ACL bone-ligament-bone preparation at 6 months showed no difference in the maximum failure load, stiffness, or AP translation compared with those in untreated controls. The authors theorized that the large decreases in ACL mechanical properties and strength in experimental animals after ACL reconstruction could not be explained by cell death and devascularization.

To provide additional experimental information, a study was conducted in the authors’ laboratory in which the previously described devitalized freeze model in the goat ACL was used and two conditions were added.¹⁰ In one group of knees, the ACL tibial attachment was elevated by osteotomy and then replaced in its normal anatomic position. In a second group of knees, the ACL tibial attachment was osteotomized and replaced on the tibia 5 mm posterior to its normal attachment to study the effects of an altered loading state (stress-shielding) on ligament remodeling properties. The results of the ACL mechanical failure tests are summarized in Table 4-7. The stiffness, maximum force, and stress and modulus were greater in the anatomic placement group than in the posterior ACL placement group; however, they were still less than historical controls. Increased failures were noted at the tibial attachment site in the posterior ACL placement group.

The results of all of the studies* are summarized in Figure 4-28, which compares historical maximum failure loads in experimental animals with the frozen devitalized specimens in Jackson and associates’ study³⁸ and the anatomic- and posterior-placed ACL preparations.¹⁰ Although the anatomic-placed ACL (after tibial osteotomy) had a reduction in strength compared with the native frozen ACL, the strength was still greater than ACL graft reconstructions. The presumed loss of the normal stress and fiber loading in the posterior-placed ACL, in addition to tibial healing effects, produced marked reductions in all mechanical properties.

FIGURE 4-28 Maximum load to failure versus postoperative time. The range (shaded region) of means for allograft reconstructions (open circles) are from previously published studies. The mean (±SEM) for ligaments frozen and their tibial insertion moved 5 mm posterior (Posterior) lie within this range. Ligaments frozen and replaced in the anatomic position (Anatomic) have greater maximum loads but still less than for in situ freezing alone as reported by Jackson et al (Frozen).

(From Bush-Joseph, C. A.; Cummings, J. F.; Buseck, M.; et al.: Effect of tibial attachment location on the healing of the anterior cruciate ligament freeze model. J Orthop Res 14:534–541, 1996. Reprinted with permission of Wiley-Liss, Inc., a subsidiary of John Wiley & Sons, Inc.)

The results of these experimental studies strongly suggest that retaining the complex native microgeometry of the ACL provides a necessary stimulus for ligament remodeling. In addition, the use of ACL ligament grafts, which have no similarity to the native ACL fiber organization and in vivo microstrains, results in replacement of the ACL graft with a scarlike disorganized collagen framework with marked loss of mechanical strength and stiffness. Insofar as the results apply to clinical ACL reconstructions, the suggestion is that ACL grafts are limited in their ability to truly restore native ACL fiber microgeometry and, therefore, never function in a manner similar to that of a normal ligament. The authors hypothesize that ACL grafts function to provide a gross checkrein to abnormal knee displacements in contrast to the so-called fine-tuning of joint motions by a ligament fiber guiding microarchitecture mechanism. It is probable that even two-bundle graft constructs behave in a similar manner, replacing the collagen fiber mass of the ligament at tibial and femoral attachments. However, with remodeling, even two-bundle constructs have altered mechanical properties and function that do not approximate native ligament function.

ALLOGRAFTS AND AUTOGRAFTS: BIOMECHANICAL PROPERTIES AFTER IMPLANTATION AND EFFECT OF IRRADIATION

Multiple studies have been published on the biomechanical properties of autografts and allografts in experimental animal models.^{3,10,20,37–40,70,71,77,83} The majority demonstrate that allografts have inferior results compared with those of autografts (and contralateral controls) in regard to mechanical strength properties. In 1989, Thorson and colleagues⁷⁷ studied the 4-month postoperative mechanical properties of adult canines that received either bone-tendon-bone allografts or autografts. The allograft group showed a mean load to failure of only 17% of the contralateral control ligament, compared with 41% in the autograft group. Shino and coworkers,^70,71 in two studies using the canine model, reported that allografts had a mean maximum tensile load of approximately 30% that of contralateral controls 1 year postimplantation. Jackson and associates³⁹ reported in 1993 in a goat model that the strength of fresh frozen patellar tendon allografts was only 27% of controls at 6 months postimplantation. In contrast, the strength of patellar tendon autografts was 62% of control ACLs. The authors concluded that allograft remodeling is much slower than autograft and suggested that patients who receive allografts should probably be protected against maximum activities for a longer period of time than those who receive autografts.

In another study in which freeze-dried bone-ACL-bone allografts were implanted, Jackson and associates³⁷ reported that the maximum load of the allografts was only 25% of the contralateral ACL controls at 1 year postimplantation. Figure 4-29¹⁰ summarizes a number of allograft ACL reconstruction studies and demonstrates that all had mechanical properties in low ranges from 4 to 52 weeks postoperative. The healing effects of autografts and allografts are covered in greater detail in Chapter 5, Biology of ACL Graft Healing.

FIGURE 4-29 Average dose-dependent response curves for all four mechanical properties (solid lines), expressed as percentages of values for contralateral frozen controls. Normalized data for 2 and 3 Mrad (broken lines) from earlier studies are also shown. Those studies used different testing conditions. Note the nearly linear declines in all curves between 2 and 6 Mrad.

(From Salehpour, A.; Butler, D. L.; Proch, F. S.; et al.: Dose-dependent response of gamma irradiation on mechanical properties and related biochemical composition of goat bone–patellar tendon–bone allografts. J Orthop Res 13:898–906, 1995. Reprinted with permission of Wiley-Liss, Inc., a subsidiary of John Wiley & Sons, Inc.)

The authors’ laboratory has studied the effects of gamma irradiation on the mechanical and material properties of allografts in the goat model and in human cadavers. In the first study, Gibbons and colleagues²⁵ reported on the effects of 2 and 3 Mrads of gamma irradiation on goat ACL bone–patellar tendon–bone in vitro properties. The study reported that the maximum stress, maximum strain, and strain energy were significantly reduced after 3 Mrads; however, there were no signification reductions after 2 Mrads of irradiation (Table 4-8).

In a second study, the effects of a higher level of irradiation (4 Mrad) were studied on human cadaver donor (aged 18–59 yr) frozen patellar tendon-bone allografts and were compared with a frozen control graft (0 Mrad).⁶⁴ Irradiation produced a small, but significant, decrease in graft length (0.6 mm; P < .01; Table 4-9). The irradiated grafts showed significant reductions in stiffness (P < .025) and maximum force (P < .001).

In a third study, Salehpour and associates⁶⁶ studied the effects of 4, 6, and 8 Mrad (40,000, 60,000, or 80,000 Gys) of gamma irradiation on the in vitro properties of a bone–patellar tendon–bone allograft unit retrieved from mature female goats. On average, stiffness decreased by 18%, 40%, and 42% at 4, 6, and 8 Mrad, respectively (P < .05 for all comparisons). The data as a whole showed the overall dose-dependent effects of irradiation on ligament mechanical properties. Doses less than 2 Mrad have minimal effects. Recently, Schwartz and coworkers⁶⁹ examined the in vivo effect of 4 Mrad of gamma irradiation on bone–patellar tendon–bone units in adult goats at time 0 and 6 months postimplantation. The irradiation significantly altered structural but not material properties. Stiffness was reduced by 30% and maximum force by 21%, resulting in these parameters averaging 12% to 20% of normal ACL values (Table 4-10).

The authors concluded that 4 Mrad of gamma irradiation affect ACL allograft subfailure viscoelastic and structural properties but not material or biochemical properties over time. The ACL allografts of the 0 Mrad failed at 497 N (see Table 4-9), which in comparison with the first study of Gibbons and associates²⁵ showed a control graft of 1400 N maximum force. This suggests that the effects of the remodeling process, resulting in a weakened graft with altered mechanical properties as previously described, produced even more profound deleterious effects than the irradiation treatment. At present, there are no experimental data to show that the low-dose irradiation as used to secondarily sterilize allografts has a deleterious effect on graft mechanical properties. One clinical study⁶³ reported inferior outcomes in knees that received irradiated Achilles tendon allografts compared with knees that received frozen Achilles tendon allografts, and the authors recommended that other secondary sterilization methods be pursued. These concepts are further discussed in Chapter 7, ACL Primary and Revision Reconstruction: Diagnosis, Operative Techniques, and Clinical Outcomes. The authors prefer to obtain allografts for patients from tissue banks that use secondary chemical sterilization for bacterial contamination.