Contribution to the pathogenesis of osteochondrosis dissecans: Experimental investigations on the idea of a mechanical overload of the cancellous bone

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Contribution to the pathogenesis of osteochondrosis dissecans: Experimental investigations on the idea of a mechanical overload of the cancellous bone

B. Gasserl,z, H. Wagner3 and W. Bandi*

1 Dr. h.c. Robert Mathys Foundation, Bettlach, Switzerland

2 M.E.M. Institute for Biomechanics, University of Berne, Switzerland

3 Laboratory for Biomechanics, Orthopaedic Clinic, University of Mainz, Germany

4 AO-International, Beme, Switzerland

Summary

The pathogenesis of osteochondrosis dissecans (OD), i.e. the formation of an osteocartilaginous arthrolith, also called aseptic bone necrosis, has been discussed time and again. The formation of the osteochondral fragment which is typical of OD may be attributed to a bland ischemia or to a single or chronic trauma. However, neither of the theories explains the typical location of OD.

This study is based on the hypothesis that overload of the subchondral cancellous bone of the femoral condyles through deformation caused by mechanical factors with a slowly progressing fatigue fracture is a decisive factor in the pathogenesis of OD. Recordings made through in vitro measurements of quasi-axial and medio-lateral displacements of knee condyles allowed for consideration of anatomical differences of the individual knee specimens.

With a knee flexion of 0” to 120°, condylar measurements showed a maximum of 83 urn of quasi-axial condylar displacements and a maximum of 680 pm of medio-lateral condylar enlargements. Greatly varying deformations in both size and direction were found in the different knee models. This leads to depressions and protuberances during knee flexion, with corresponding alternations of

compressive and tensile load of the cancellous bone in the femoral condyles. Compared to the lateral condyle, larger deformations were found in the medial condyle in a quasi-axial direction according to predilected locations of OD. Medio-lateral condylar enlargements may lead to tensile elongations in the size range of the fracture elongation in cancellous bone. Even without considering load peaks, the results of this study point to alternating loads in the femoral condyle and thus support the hypothesis.

1. Introduction

1.1. Problem

Osteochondrosis dissecans (OD) was the name given in 1887 by F. Konig to the formation of an osteocartilaginous arthrolith from a convexly curved joint part (Konig, 1887). To date, questions regarding the pathogenesis of OD also described as aseptic osteonecrosis, cannot be fully answered. Different factors were discussed but until now none of the various theories prevailed. Different studies (Campbell et al., 1966; Bergmann, 1927; Enneking, 1977) attribute the formation of the loose body to

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local ischemia, i.e. to disturbances of local circulation. Representatives of the opposing theory (Burckhardt, 1924; Nagura, 1937; Aichroth, 1971/b) advocate a single or chronically effective trauma with very small, slowly progressing fatigue fractures in the subchondral cancellous bone. Because neither of the theories explains the typical location of OD, the underlying aspects of this study were recorded in the early eighties (Bandi, 1982 and 1984).

- OD affects almost exclusively joints lying between long skeletal levers.

- OD forms definitely on closely defined predilection sites, usually on a convex part of the joint.

- The osseous disintegration starts beneath a macroscopically intact joint cartilage.

- The loose joint fragment has invariably a more or less typical shape.

OD starts within the subchondral cancellous bone with a slowly progressing demarcation of a bean- shaped osteochondral body beneath an initially macroscopically normal joint cartilage. Eventually, the covering hyaline cartilage splits and the osteochondral fragment protrudes and migrates into the joint space. The nucleus of the fragment consists mostly of necrotic cancellous bone. The joint side surface of the fragment is covered by hyaline cartilage whereas the cancellous bone side surface as well as the fracture surface of the fragment bed are covered by fibrous cartilage (Fig. 1). OD occurs predominantly towards the end of the growth period in adolescent males with long limbs, who practise sports, and are therefore exposed to excessive intraarticular compressive forces due to joint movement. The fragmentation of a loose body occurs with less frequency on the lateral side of the medial knee joint condyle on the convex part of the trochlea tali and in the capitulum humeri.

Fig. 1: Osteochondral fragment in situ in a femoral condyle of the knee joint shown in a lateral X-ray view: a) A light demarcation zone develops around the fragment and delimits again by a denser area towards the condyle.

Fig. lb: The light demarcation zone contrasts clearly with the surroundings.

Fig. lc: The loose, bean-shaped fragment has a macroscopically normal hyaline cartilage on the joint side, however the fracture side is covered by irregular fibrous cartilage.

A clinical study by Aichroth (Aichroth, 1971/a) shows that 50% to 60% of the patients are between 10 and 20 years of age and that at least 75% of the cases occur at the classical predilection site. Occurrences of OD at the heelbone (calcaneum) or at the kneecap (patella) are rare.

1.2. Clinical importance

The manifestation of OD is different in each patient. Complaints are seldom typical and often negligible: Increased fatigue and/or dull not exactly localized pain have been recorded. However, the situation changes when the osteochondral fragment exits from its bed. Acute attacks of pain occur, sometimes accompanied by temporary joint locking; cartilage damage might cause gonarthrosis in the course of time.

Histologically speaking, the osteochondral fragment while still in its bed resembles an atrophic pseudarthrosis. Due to sudden outbreaks of pain

Gasser: Contribution to the pathogenesis 53

with or without blocking of the knee, the patient judges the impingement of the fragment between the joints to be an accident. However, based on the pathogenesis, a single exogenous trauma (accident) has to be rejected in most cases as the initiator of this disease.

If the formation of an osteochondral fragment is recognized in time, it can be stabilised by means of cortical pins or screws and will heal sufficiently (Fig. 2). -

Fig. 2: Patient (19 years) with osteochondrosis dissecans: a) The large bean-shaped loose fragment can clearly be recognized in the lateral view; b) status immediately after fixation with cortical bolts; c) situation after 12 years (frontal view); d) situation after 12 years (lateral view); e) and 0 tomography in different layers: complete reconstruction.

According to Jakob (1990), the healing process depends on the patient’s age. Whereas in children up to the age of 12 years only a slightly invasive treatment is necessary, i.e. usually careful treatment until healing of the defect is sufficient, surgery is often necessary in adolescent and adult patients. Different techniques such as the conventional method of fixation by wire or the new compression pinning system or, as shown in Fig. 2, fixations by means of bone bolts are used to reduce and fix the osteochondral fragments.

1.3. Hypothesis

The previous statements elucidate that neither an osteonecrosis caused by ischaemia nor a fragmentation caused by mechanical behaviour of the joints are sufficient to explain the pathogenesis of OD. Site-specific aspects such as geometry and material properties as well as local loading have to be included and considered as possible explanations. Predilection sites of OD have in common that they are located at convex joint surfaces and usually very close to transitions into concave surface parts. These specific joint surfaces show different and changing radii of curvature in the frontal and lateral view. It is known (Carter et al., 1987) that in curved surfaces under load, the kind of curvature substantially influences the kind of strain on the material beneath it. Whereas in convex surfaces hydrostatic pressure occurs, concave surfaces are marked by radial compression and tangential tensile strains and tensile elongations. When flexing the knee joint, high dynamic loads and deformations occur due to the changing load of high tibia-femoral and patello- femoral compressive forces and the unloading in the margin between the two compressive force components. At 0” and 30” and also between approximately 120’ and 130” of knee flexion, compression occurs in specific condylar segments due to body weight and patellar pressure respectively. On the other hand, a protuberance occurs between 40” and 110’ due to the tong-shaped influence of compressive patellar and tibia1 forces. The dynamics of the occurring loads in the mentioned predilection sites are given by geometry as well as by loads. Therefore, it may be assumed that under physiological load, specific joint parts are alternately subjected to depressions and protuberances which lead to compressive and tensile deformations (Fig. 3).

A further essential aspect seems to be the different behaviour of the viscoelastic material of cartilage and cancellous bone. Perren and Cordey (1977) reported at least a 4 to 5-fold fracture elongation of cartilage tissue compared to cancellous bone with a simultaneous 4 to IO-fold tensile strength. This implicates a discontinuity of material behaviour at the interface between cartilage and cancellous bone which, besides geometry and load, has a negative influence on the strain distribution in the endangered section.

The Knee 2994, Supplement 1

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patellar ligament

Fig. 3: Knee condyles are subjected to high compressive forces caused by body weight and by the patella and tibia which lead to minimal depressions in the condylar vault. A protuberance of the condylar contour occurs in the finite vault segment which, according to the angle of knee flexion, is neither covered nor compressed by the patella (dashed line = elevated deformation of the condyles). With each flexing and extending movement of the knee, depression and protuberance migrate over the condyles and the viscoelastic subchondral cancellous bone is exposed to dynamic loading so that fatigue fractures may occur.

Therefore, our deliberations point to the fact that the decisive factor in the pathogenesis of OD under given conditions of geometry, load distribution and material properties is a dynamic overload of the subchondral cancellous bone with ensuing slowly progressing fatigue fractures. Disturbances of local circulation, i.e. the circulation transfer at an epiphyseal closure can create a temporary disposition but they will never be the trigger of OD.

In using the method of finite elements it was possible to show (Nambu et al., 1991) that complex three- dimensional strain distributions and especially large deformations in the medial compared with the lateral condyle occur at the predilected site of OD.

1.4. Aim

In vitro measurements on knee models should be made leaning on the hypothesis of overload of the subchondral cancellous bone by mechanically condi- tioned joint deformations. The pathogenetically important deformations in the area of the femoral

condyles, the most frequent predilection site, should be established and quantified by means of simulated, varying loads under flexion and extension. Besides, the influence of anatomical differences should be evaluated by comparing the results of different knee joints.

2. Methods and materials

2.1. Methods

A specially designed device (Bodem et al., 1985) was used to measure the condylar deformations on a cadaver knee. This measuring device allows the investigation of selected biomechanical parameters on knee models during flexion and extending movements. The measuring device is designed for computerized data acquisition.

Knee models, consisting of the distal third of the femur and the proximal third of the tibia are fixed in the device to firmly lock the femoral stump and to anchor the tibia1 part in a hollow cylindrical mounting (Fig. 4).

Fig. 4: Schematic representation of a knee model in the measuring device with an indication of degrees of freedom and measured variables (aK=angle of knee flexion, aT=torsional angle of the tibia, a\r=angle of varus-valgus, dT=axial tibia1 displacement, F-I= tibia1 compressive force, FQ=tensile force of quadriceps).

This bracket allows for all tibia1 movements such as knee flexion, axial tibia1 displacement, tibia1 rotation and varus-valgus motion according to the degrees of freedom of the knee joint (Fig. 5). Moreover, the corresponding movements are potentiometrically analysed and computerized. Knee models investigated with an intact ligament apparatus and capsules can be either loaded with an axial tibia1 compressive load to simulate body weight or with a tensile load at the quadriceps to simulate the quadriceps tendon force. Simulating a constant preload, these loads are charged with weights and

Gusser: Contribution to the pathogenesis 55

cable-controlled onto the device and the models. Nonslip fixation is guaranteed by a special bracket to connect the quadriceps tendons with the wire. For control tibia1 compressive forces and quadriceps tensile forces are measured with a load-sensing device, based on strain gauge bridges, during data recording. The measuring system is equipped with a total of 15 measuring channels in addition to a cyclometer which allows the fitting of a whole series. Following the protocol of experimental procedure, additional measuring transducers can be mounted to record corresponding data (Fig. 6). The measuring frequency of the system is 300 Hz in order to take into account the dynamics during flexion and extension of the knee models. The flexing and extending movements of knee models are done by hand.

principle, and transmit the displacement to be recorded from the desired localisation to the measuring element through a sleeve-guided pin. The measuring element consists of a gauge-based bending spring whose deformation after calibration allows for displacement measurements. The two transducers for the measurements of axial or better quasi-axial condylar displacements were fixed at the proximal condylar insertion. Helped by a drill hole, the measuring pin, which in this case was a Kirschner wire, was able to transmit the displacements to the measuring element practically without friction. The tip of the Kirschner wire with short thread was used as an anchor at the predilection sites in the area of the subchondral bone. One of the transducers was placed in the medial condyle, the other in the lateral condyle (Fig. 8). One special bone screw each was placed medially and laterally to define measuring position and direction for measurements of lateral condylar displacements.

Fig. 5: Survey of the tibia1 assembly with mounted displacement and force transducers.

Fig. 7: Displacement transducer with strain gauge- based bending spring (position A) to measure a lateral deformation (horizontal direction in the Fig.).

Fig. 6: Survey of the devices for data acquisition.

To measure condylar deformations, 4 displacement transducers were mounted in order to record displacements at the measuring site in relation to their fixed sites on the femoral bone. Two transducers each served to measure the quasi-axial subchondral movements with regard to the femoral axis and the lateral (medial-lateral) movements. Displacement transducers (Fig. 7) are based on the bending spring

Fig. 8: Schematic view of positions and directions of the quasi-axial measuring transducer.


The displacements were transmitted by means of the measuring pin directly from the head of this screw to the transducer fixed to the femoral diaphysis (Fig. 9). The position of the reference point of the different transducers was chosen at the proximal condylar insertion and the femoral diaphysis as to be far enough from the measuring point in order to eliminate any influence on deformations within the measuring zone.

Fig. 9: Survey of the measuring device with the knee model: in the foreground measuring transducers for the lateral condylar deformation (position Al and the fixation device of the femoral diaphysis (position B).

Calibration of the displacement transducers was made by a micrometer and by a special device for measuring transducers. On each specimen, calibrations were carried out immediately prior to implantation of the transducers and were repeated after measuring. These calibrations permitted the determination of a minimal resolution of 8 urn for displacement transducers over the entire data acquisition system. Transducers were preloaded at implantation to allow displacement measurings in both directions. This preload was in the range of 0.5 to 0.8 mm. In preliminary tests, a minimum linearity range of 0.0 to 1.5 mm was determined for all transducers.

One measuring sequence of a given load always included a cycle of flexion, extension and again flexion over a knee flexion range of 0” up to at least 120’. The flexing movement was always conducted at a knee flexing rate of about 25”/sec. Loads for the axial tibia1 force were varied between 600 and 1000 N and for the quadriceps tendon tensile force between 700 - 2000 N. Based on former measuring experiences, these load values were chosen in order not to overload the cadaver specimens and to get the best possible physiological and measurable condylar deformations. After closing the test programme, some measuring sequences were executed a second time to test the reproducibility. During measuring, physiological movements of the tibia such as tibia1

rotation, varus-valgus motion and axial tibia1 displacements were determined besides the displacement values of quasi-axial and lateral

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transducers, the quadriceps tendon tensile force and axial tibia1 compressive force.

Special programmes were written and used to evaluate the measurements. These programmes comprised the conversion of the recorded data (bit values into HEX form) by means of calibration curves and the determination of measuring values. An evaluation frequency of a 3” knee flexion was selected for the graphic presentation of displacements. The corresponding value was a 10” knee flexion for loads and tibia1 movements. This required interpolations and average means of the 1st and 2nd flexion of a measuring sequence and, if possible, of the second load sequence.

2.2. Material

Measurements were carried out on 4 cadaver knees on whose donors unfortunately only limited data and especially no indications with regard to the functional state of the knee was available:

Before the investigation, knee models were preserved at -2OOC and prepared immediately prior to measurements. A completely intact ligament apparatus was found only in knee no. 1 whereas the condition of the collateral ligament apparatus in other specimens was unsatisfactory.

The positions of the quasi-axial Kirschner wires and of the lateral positioning screws were recorded radiologically (Fig. 10) before placing the specimens into the measuring system. Due to the fact that the measuring transducers could record displacements of the subchondral layer locally only, these positions were very important for the interpretation of the results.

3. Results

Data from the knee model no. 3 were deleted in the evaluation of measurements because bad bone quality of this model did not allow for a sufficient anchoring of the measuring transducers and moreover a fracture occurred after the 5th measuring sequence.

Gasser: Contribution to the pathogenesis 57

Fig. 10: Radiographic frontal and sagittal view of the knee model no. 1 with a representation of the position of measuring sites and directions.

Based on radiographies done prior to measuring, the following table shows the directions of the quasi-axial transducers in the medial and lateral condyles with regard to the femoral axis in the frontal and sagittal planes. For the 3 evaluated knee models, orientations of these transducers in the sagittal plane are represented at a flexion angle of 0”, 30”, 60” and 90” as in Fig. 11.

I dlllhfv /--am 0 deg 80 deg 60 deg 90 deg

Fig. 11: Directions of quasi-axial displacement transducers in the sagittal plane, represented for an angle of knee flexion at 0”, 30”, 60” and 90”: a) specimen no. 1 (above).

I I I

0 deg 30 deg 60 deg 90 deg

Fig 11 b: Specimen no. 2 (middle)

a dk I

m 0 deg 30 deg 60 deg 90 deg

Fig. llc: Specimen no. 4 (below).

Directions of quasi-axial transducers in the medial and lateral condyles: (frontal and sagittal view with regard to the axis of the femoral shaft)

Knee model: Condyle: Frontal Sag&l plane: plane:

No.- medial 3.Y 180 lateral 159 0”

No. 2 medial 25” 129 lateral 16” 20”

No. 4 medial 35” 10° lateral 10” 30”

The varying directions and positions of the measuring transducers are given by the limited implantation possibilities due to the specific anatomy of the different specimens and influence the displacement values, which depend on the flexion angle, accordingly.

Figure 12 shows a survey of the test arrangement during measurements and of the knee models with implanted transducers installed in the loading device. Evaluation of the measuring data und their comparison for each knee model ensued principally for two loads: only the variation of the quadriceps tendon tensile force was considered and investigated at 1100 N and 2000 N whereas the axial tibia1 compressive force was kept constant at 900 N.

Fig. 12: Testing of knee models with implanted measuring transducers: a) Frontal view of the knee model no. 1 at a flexion angle of approx. 80”.


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Fig 12b: Overview of the knee model no. 4 with the femoral fixation and tibia1 mounting at extension.

3.1. Loadings and knee flexing movements during flexion / extension

The course of the quadriceps tendon tensile force (Fig. 13) during a measuring sequence consisting of flexion, extension and repeated flexion resulted in a very similar characteristic at both loading level rates of 11OONand20OON.

Quadriceps Tendon Force [Nl

2400

2300

--t 1 st Fhion --t Extension - 2nd Flexion

2200 -

2100‘

2ooo-

1900-

1800 -

1700 -

1600- Knee Specimen No. 1

1500 I -20 0 20 40 60 80 100 120 140 160

Knee Flexion Angle [deg]

Fig. 13: Course of the quadriceps tendon tensile force of the knee model no. 1 at a loading rate of 2000 N.

Loads were subjected to dynamics by charging them with weights and guiding those weights over a cable control and deflection sheaves during knee flexion. A maximum loading amplitude of approximately +/- 80 N was recorded within one cycle of flexion or extension. In addition, the flexing load was always higher by approximately 200 N than the extension load which is due to the inertia and friction of the loading system.

In the evaluation, the axial tibia1 compressive force had to be compensated with the weight of the tibia1

mounting (total weight = 3.6 kg) depending on the flexion angle. The course of the axial tibia1 compressive force (Fig. 14) was subjected to even higher dynamics than the quadriceps tendon tensile force.

Axial Tibia Force [N]

._-

1400 I - 1stFlexion 1300 --t Extension

1200 - 2nd Flexion

1100 i

looo-

900-

800-

700-

500 1 7

-20 0 20 40 60 80 100 120 140 160

Knee Flexion Angle [dcg]

600 1 Knee Spccimcn No. 1

Fig. 14: Course of the axial tibia1 compressive force (load rate = 900 N) at a quadriceps tendon tensile force rate of 2000 N: a) knee model no. 1 without load

1

drop.

Axial Tibia Force [N]

1oao-

900-

800-

7OQ-

600 Knee Specimen No. 1

500 -20 0 20 40 60 80 loo 120 140 160

Knee Flexion Angle [dcg]

Fig. 14b: Knee model no. 4 with load drop at greater flexion angles.

In contrast to the quadriceps tendon tensile force, the tibia1 compressive force was smaller in flexion than in extension due to the inertia of the system. In addition, a distinct load drop and a renewed load build-up were recorded in a certain area of measurement at flexion angles ranging from 80” to 120”. This load drop could be allocated to no particular configuration of knee specimens or quadriceps tendon tensile force.

Gasser: Contributim to the pathogenesis s9

In all knee models, the course of tibia1 rotation and of varus-valgus motion was analogous qualitatively and practically identical quantitatively. Neither load dependences nor differences in flexion or extension could be determined. In tibia1 rotation (Fig. 15) with increased flexion, an internal rotation could be found which showed a stronger increase in the first 40” of flexion and reached maximum values of 20’ and 30° at full flexion. The value of the maximum internal rotation was dependent on the specimen and not on the quadriceps tendon tensile force. A similarity applies to the angle of varus-valgus (Fig. 16) which always decreased with increasing flexion from approx. +5” (varus) to -5” (valgus). This angle modi- fication was larger during the first 40” to 60’ of knee flexion than during the second half of the flexion cycle.

30

20

IO

Internal Tibia Rotation [deg]

- 1st Flexion ---t Extension -O- 2nd Flexion

Knee Specimen No. 4

-20 0 20 40 60 80 100 120 140 160


Fig. 15: Course of tibial rotation in the example of the knee model no. 4 with a load rate of 900 N and 2000 N for tibia1 compressive force and quadriceps tendon tensile force, respectively.

In comparing the different knee models as far as behaviour of loads and of tibia1 movements concerned, no marked differences were recorded.

the

the are

3.2. Condylar deformations in a quasi-axial direction

The determination of quasi-axial displacements of subchondral bone in the tip area of the measuring transducer was done in relation to the anchoring of the transducer at the condylar insertion. The re- spective transducer direction (see Fig. 11) has to be taken into consideration when evaluating the medial and lateral condyle for each knee model. Independent thereof, the characteristic of these quasi-axial displacements depending on the flexion angle is different for the various knee models and condyles. Under a larger quadriceps tendon tensile force, a

larger displacement was determined in all specimens. As a consequence of the increased load the increase of these displacements was again different in all three specimens.

10 -

Varus / Valgus Angle [deg]

-C 1st Flexion

Knee Specimen No. 4

10 1 I -20 0 20 40 60 80 100 120 140 160


Fig. 16: Course of varus-valgus motion in the example of the knee model no. 4 with a load rate of 900 N and 2000 N for the tibia1 compressive force and the quadriceps tendon tensile force respectively (positive angles correspond to a varus position).

The largest displacements in the quasi-axial direction of the knee model no. 1 (Fig. 17) were found for the flexion angle corresponding to the alignment of the measuring transducers in the direction of the tibia1 axis. This corresponds to a knee flexion angle of 30’ and 0” at the medial and lateral transducers, respectively. In this knee model, the increase of displacements with an increasing load can be clearly distinguished. Displacements are distinctly larger in the medial condyle than in the lateral condyle. Due to a somewhat more dynamic characteristic, load dependent differences are less noticeable in the knee model no. 2 (Fig. 18). Also in this knee model, maximum displacements were determined at the flexion angle where transducer directions corre- sponded to the tibia1 axis each at the beginning of a flexion. The characteristic for the knee model no. 4 (Fig. 19) where the lateral condyle is always under compression, looks different. Similar to the knee model no. 1, the displacement amplitude in the medial condyle is clearly larger than in the lateral condyle. The allocation of the transducers’ direction and of the largest displacements with regard to the angle of knee flexion is also here in accordance. At 5” to 10’ of knee flexion, the medial tranducer points in the direction of the tibia1 axis, the lateral transducer at approximately 30”. Considering the displacement amplitudes (defined as the difference between maximum and minimum values of displacements) of the individual measuring curves (Figs 17-19), the


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increase of the quadriceps tendon tensile force from 1100 N to 2000N results in the following enlargements:

Quasi-axial Displacement [mm] / Knee No. 1

0.00

-0.02

-0.04 - lat./l 100 N

- IatROOO N

I - med./llOON

__

- med.ROOON

-0.08 I 0 20 40 60 80 100 120 140

Knee Rexion Angle [deg]

Fig. 17: Quasi-axial displacements in the medial and lateral condyles of the knee model no. 1 for quadriceps tendon tensile load rates of 1100 N and 2000 N as a function of the angle of knee flexion (positive displacements correspond to a protuberance). The curves shown correspond to mean values of the 1 st and 2nd flexion phases.

Quasi-axial Displacement [mm] / Knee No. 2

0.02 I 0.01

0.00

-0.0 1

-0.02

-0.03 -I I 0 20 40 60 80 100 120 140

Knee Fiexiott Angle [deg]

Fig. 18: Quasi-axial displacements in the medial and lateral condyles of the knee model no. 2 for quadriceps tendon tensile load rates of 1100 N and 2000 N as a function of the angle of knee flexion (positive displacements correspond to a protuberance). The curves shown correspond to mean values of the 1st and 2nd flexion phases.

According to the knee model, enlargements of the displacement amplitudes between 7% and 100% can be determined for a load increase of 82%.

Outward M-L Displacement [mm] / Knee No. 4

0.02

o.Ou -__-

-0.02

-0.04 - lat./l 100 N

-0.06 - lat./2000 N , - med./l 100 N

-0.0x

-0. IO I 0 20

--t med./2000 N

I 40 60 80 100 120 I 40

Knee Flexion Angle [degl

Fig. 19: Quasi-axial displacements in the medial and lateral condyle of the knee model no. 4 for quadriceps tendon tensile load rates of 1100 N and 2000 N as a function of the angle of knee flexion (positive displacements correspond to a protuberance). The curves shown correspond to mean values of the 1st and 2nd flexion phases.

Increase of the quasi-axial maximum displacement amplitudes as a function of increased load [pm]:

Quadriceps tendon tensile force:

Knee model no.1 : 1100 N: 2000 N: Increase: medial 55 83 51% lateral 22 44 100%

1 medial I 38 1 59 I 55% I 1 lateral I 22 1 26 1 18% 1

The differences of quasi-axial displacements between the two flexion cycles are very small and amount to approximately 10 urn (Fig. 20). The corresponding differences between flexion and extension are about 20 pm.

Differences between the individual knee models and condyles can be well seen if displacements, i.e. protuberances and depressions of medial and lateral condyles at a selected angle of knee flexion are shown in a bar chart (Fig. 21). Compared to lateral condyles, these quasi-axial displacements in medial condyles are primarily larger in their variability regarding the angle of knee flexion than in their quantity.

Gasser: Contribution to the pathogenesis Sll

Quasi-axial Displacement [mm] I Knee No. 1

0.02

vertical to each other, the influence of the varying quadriceps tendon tensile force was less obvious.

0.00 __- __ Quasi-axial Protuberance I Depression [mm] Lateral Condyle at 2000 N

-0.02

- lat./l. loading - l&2. loading - lat./mean - med./l. loading

-0.06 - m&2. loading - medhnean

0.02

0.00

-0.02

-0.04

-0.06

-0.08

20 -I

40 60 80 100 120 140 Knee Rexion Angle [deg]

Fig. 20: Comparison of quasi-axial displacements for the 1st and 2nd flexion phases and their mean values in the example of the knee model no. 1 with a quadricepts tendon tensile load rate of 2000 N.

Quasi-axial Protuberance I Depression [mm] Medial Condyle at 2000 N

nn27 ._._-

0.00

-0.02

-0.04

-0.06

??Knee No. 1 ??KneeNo. 2 ??KneeNo.

1 -0.08 J

10 30 60 90 120


Fig. 21: Quasi-axial displacements of the three tested knee models at selected angles of knee flexion (lo’, 30”, 60”, 90” und 120’) and a quadriceps tendon tensile load rate of 2000 N (positive displacements correspond to a protuberance): a) medial condyle.

3.3. Condylar displacements in a medio-lateral direction

Measuring of lateral condylar displacements in medial and lateral directions was done with regard to the fixed femoral diaphysis. Similar to quasi-axial displacements, medial and lateral condylar displacements as a function of the flexion angle yielded very different values for each tested knee model. In this case, the direction of load and displacement being

??Knee No. 1 •l KneeNo. 2 •a Knee No.4

IO 30 60 90


120

Fig. 21 b: Lateral condyle.

In the knee model no. 1 (Fig. 22), the medial condyle moves medially whereas the lateral condyle displaces towards the interior knee, i.e. also medially. Condyles are displaced medially in relation to the femoral axis. The largest medio-lateral displacements are measured at a flexion angle of between 25’ and 45”. In contrast to the medial condyle, larger displacements (compressions) were measured in the lateral condyle with a larger quadriceps tendon tensile load. In the knee model no. 2 (Fig. 23), corresponding medio- lateral displacement values decrease with increasing flexion, especially in the medial condyle.

Maximum displacement values occur in extension and are smaller than in the knee model no. 1. With regard to the increase in the quadriceps tendon tensile force, a larger displacement in the medial and lateral condyles was found only at a flexion angle of over 50”. The knee model no. 4 (Fig. 24) shows a completely different course of displacement. Starting with a compression on both sides in extension, the outer sides of the medial and lateral condyles displace with a larger flexion angle outward, i.e. in a medial and lateral direction respectively. The maximum values of these displacements are much smaller than in the other knee models. The difference amounts almost to factor 10. In this knee model, the influence of the quadriceps tensile force on displacement values is negligible and only evident at flexions over 50°.


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Outward M-L .Displacement [mm] / Knee No. 1

0.6

- lat./l 100 N 0.4 I - lat.L?OOON

- med./llOON

0.2 - med.ROOON

0.0 _____________--- ____.

-0.2

-0.4

-0.6 4 I 0 20 40 60 80 100 120 140


Fig. 22: Medio-lateral displacements measured at the outer sides of medial and lateral condyles of the knee model no. 1 for quadriceps tendon tensile load rates of 1100 N and 2000 N as a function of the angle of knee flexion (positive displacements correspond to a protuberance). The curves shown correspond to mean values of the 1st and 2nd flexion phases.


0.6 I

-C lat./l 100 N

0 20 40 60 80 Knee Rexion Angle [deg]

100 120


Differences between displacement values of the 1st and 2nd flexion are larger than those of quasi-axial displacements and different in each knee. The largest differences were found in knee model no. 2 (Fig. 25) where they amounted to an approximate maximum of 0.15 mm at a mean range of knee flexion. In the knee model no. 1 the corresponding behaviour curves

practically coincide and also in knee no. 4 showing very small displacements, the differences never yielded more than 0.05 mm.

Outward M-L Displacement [mm] I Knee No. 4

0.02

0.00 __

-0.02

0.04

-0.06 - latJl100 N - latJ2000 N - med./llOON

-0.08 - medJ200ON

_-

20 -I

40 60 80 100 120 140 Knee Flexion Angle [degl



--)- lat./l. loading

0.6 --t IatR. loading

- lat./mean - med./l. loading

40 60 80 100 Knee Rexion Angle [deg]

Fig. 25: Comparison of medio-lateral displacements for the 1st and 2nd flexion phases as well as their mean values for the knee model no. 2 with a quadriceps tendon tensile load rate of 2000 N (positive displacements correspond to a protuberance).

A resulting lateral condylar enlargement can be calculated by adding medio-lateral displacements of the medial and lateral condyles (Fig. 26). The course of the condylar enlargement is again different in all

Gasser: Contribution to the pathogenesis

knee models. Only knee model no. 1 shows an effective load dependence on the quadriceps tendon tensile force. The largest condylar enlargement occurs at maximum flexion. A reduction of condylar enlargement occurs at the larger quadriceps tendon tensile force and a flexion angle of between 20” and 70”. This reduction fails to appear at the smaller tensile force of the quadriceps tendons. At this range of flexion angle, this knee model experiences a larger condylar enlargement at a smaller load. In the knee model no. 2, the condylar enlargement of approximately 0.25 mm in extension decreases with increasing flexion and becomes negative at full flexion. Therefore, the condyles are laterally compressed in the flexed position. Condylar enlargement of the knee no. 4 is quite the contrary. Starting from a compression in extension, it continuously increases with growing flexion and obtains a comparatively small enlargement of 0.06 mm in the flexed position. From mean values of the 1st and 2nd flexion, the following minimum and maximum values of condylar enlargements can be determined for individual knee models:

Condylar Enlargement [mm] / Knees No. 1, 2, 4

(izL - KneeNo.l/llOON

0.4 - - KneeNo.2/11OON - KneeNo.4/1lOON - KneeNo. 1/2OOON

0.2 - - KneeNo.212OOON

--o- KneeNo.4/2OOON

0.0 -------- --___.

-0.2 -I 0 20 40 60 80 100 120 140


Fig. 26: Lateral condylar enlargement of the tested knee models as a result of medio-lateral displacements in both condyles as a function of the angle of knee flexion and of both quadriceps tendon tensile load rates of 1100 N and 2000 N.

Minimum and maximum values lateral condylar enlargement as a load [pm]:

s13

of the medio- function of the

Knee model no.2

These extreme values show that depending on the knee model, large differences in displacement and also deformations with alternating elongation and compression changes may occur.

A bar chart of the connection between measuring sites of the medial and the lateral transducers can illustrate the movement of the condyles and the condylar mean plane with regard to the shaft axis. In Fig. 27, the connecting bars of all three knee models and of selected knee flexion angles with regard to the femoral axis are shown whereby displacements and deformations are strongly elevated compared to the effective connective distance between the two measuring points. Besides the differences of medio- lateral displacements in the individual specimens, it becomes evident that knee condyles also experience a solid displacement in a medial direction. This applies specially to the knee models no. 1 and no. 2 with maximum values of 0.61 mm (at 30” of knee flexion) and 0.28 mm (at 60” of knee flexion) respectively.

Due to very small lateral condylar displacements, knee model no. 4 moved almost imperceptibly medially at low flexion and at full flexion to a maximum of 0.04 mm laterally.


s 14

Medial-lateral Displacement and Deformation of the Condyles

Knee No. 1 Knee No. 2 Knee No. 4

loo

30”

60”

90”

120°

- lat.

Fig. 27: Displacement and deformation of a medio-lateral bony bar of the three knee models in selected flexed positions at a given quadriceps tendon load rate of 2000 N.

4. Discussion

The results of this study are based on the relatively small number of 3 tested knee models. This number is too small to allow for a conclusive survey of possible deformations in the knee condyles of different individuals. However, the varying results of the tested knees demonstrate that measurements of deformations on a large number of specimens would be necessary to deal with all anatomical differences such as age, height, weight, but also with bone geometry and quality. The great variety of measured deformation properties shows that some individuals might experience greater condylar deformations than others due to their bodily constitution. Yet in this study, no load differences between athletes and less active persons were taken into consideration.

The results of 3 knee models show that greater deformations might occur in younger individuals (knee specimen no. 1, 33 years) than in older ones (knee specimens no. 2 and 4, 40 and 63 years respectively). Height and body weight cannot be directly quoted to explain knee-based different results as all knee models were tested under the same conditions with the same experimental set-up. In vivo, a greater height would call for an increase in condylar deformations due to the longer lever arms

of the limbs. A greater age or older knee models could lead to a change, especially to an erroneous increase of measured displacements due to incorrect loading displacements as a result of a slackening of the capsule ligament apparatus as was possibly the case with knee models no. 3 and no. 4. Both change and erroneous increase of measured displacements do not apply to our measurements of condylar deformations with maximum values for knee specimen no. 1 so that the different results have to be attributed mainly to individual properties of bone geometry and structure. The only reservation in a possible comparison of the results of the three knee models is the fact that displacement transducers had a somewhat different position and direction in each knee. Even though the transducer positions are taken into consideration in each evaluation, they render a direct comparison more difficult.

Of course, it would be interesting to investigate knee specimens of even younger individuals showing symptoms of OD or who had OD earlier on. Only the comparisons with healthy knee specimens of the same age could give detailed information regarding our deformation hypothesis. Despite the rather small number of the investigation material, the individually

Gasser: Contribution to the pathogenesis s 15

measured differences with regard to kind and extent of displacements and deformations of the knee condyles absolutely support the hypothesis regarding local mechanical overload of the cancellous bone with subsequent fatigue fracture.

When evaluating the kinematics of the investigated knee models, it has to be taken into consideration that only knee no. 1 had a fully intact ligament apparatus. Tibia1 rotations and varus-valgus alterations mainly depend on the ligament apparatus. Nevertheless, one could proceed from practically the same behaviour in all knee models.

During the knee flexion phase, an internal tibia1 rotation of 20” to 30” was measured and a lateral movement of approximately 10” whereby a transition from varus to valgus took place at the movement from extension to flexion. There is some uncertainty as to the localisation of the zero point of intemal- external rotation and varus-valgus position in these knee movements due to limited evaluation possibilities at implantation of the specimens into the testing device. A knee loaded in vivo has a somewhat larger movement range with a mean freedom of tibia1 rotation of approximately 35” (Ruetsch et al., 1977). The comparison with an intact knee yields an accordance in that with unimpaired knee movement starting from a varus position in extension an internal rotation ensues with increasing flexion, a fact which was found in all knee models. The measured medial displacement of the whole condylar body which occurs with increasing flexion can be explained physically by the increasing valgus position. The extent of this solid displacement could be somewhat elevated due to possible ligament slackening or the kind of loading.

Measurements were difficult in that small displacements had to be measured in viscoelastic material, i.e. in cancellous bone as well as in places with difficult access. External calibrations of the measuring transducers immediately prior and after specimen measurements partly resulted in small differences at evaluation, which were considered by taking the mean of the two values. Another uncertainty is the preloading of transducers: due to the viscoelastic behaviour of the cancellous bone, a relatively quick partial loss of this preload has to be assumed without the possibility of measuring it. The data system, originally designed for a different sort of measurement, was equipped with a relatively small test amplifying sensibility for the acquisition of these data. This disadvantage was partly compensated by the high measuring frequency of the system by means of the same angle of flexion and means of the first and second loading cycle. In addition, linear interpolations with both neighbouring values were

executed to calculate the measuring values of 3O flexion angle steps. Further inaccuracies which could not be determined may also have been caused by ligament extensions or slippings which can hardly be excluded in older knee models. This makes it obvious that the accuracy of absolute values of condylar deformations is very difficult to assess. Therefore, measurement results are only to be considered as relative alterations over the range of flexion which should be sufficient for the evaluation of our hypothesis.

The forces exerted on the knee models of this study correspond approximately to those of an above- average static load due to body weight. The tibia1 compressive force of 900 N represents a considerable body weight. Quadriceps tendon tensile forces of 1100 N and 2000 N can occur in daily activities such as getting up from a chair (1150 N) or going upstairs (1600 N) (Ahmed et al., 1983). It is well known that especially in athletic activities, much higher peak loads can occur so that much greater displacements and deformations are possible compared to the present results. It has therefore to be taken into consideration that according to the measured displacements, there is no linear relationship between condylar deformations and charged loads.

Comparing the measuring values of the 1st and 2nd flexion within the loading cycle, it was possible to prove the reproducibility of quasi-axial and medio- lateral displacement measurements. When repeating individual loading cycles after the execution of the scheduled measuring programme, larger quantitative but no qualitative discrepancies were established due to the disadvantages described such as viscoelastic material behaviour, possible slackening of the capsule apparatus or other non measurable influences.

In spite of different positions and directions of measuring transducers, the results of quasi-axial displacements of the tested knee models allow the conclusion of a varying deformation behaviour regarding extent and dynamics. The youngest knee model has considerably larger displacements than the other two models. Dynamic displacement behaviour depending on the flexion angle with alternating load, i.e. elongation and compression alternations, was established in all specimens. Maximum displacement amplitudes also show that the medial condyle and with it the predilected site of OD experiences larger displacements than the lateral condyle. In our investigation series, this did not apply to the knee model no. 2 with the smaller quadriceps tendon tensile load.

Individual differences of the 3 knee models are especially distinct in measurements of medio-lateral displacements. Each specimen shows a different dis-


S 16

placement characteristic as medio-lateral displacements at the flexion angle range either remaining almost unchanged, or decreasing or increasing. In addition, absolute displacement values are vary distinctly with differences up to a factor of 10. The results of the condylar enlargement are accordingly specimen-specific. The medial solid body displacement of the condyles seems to be the only common aspect in medio-lateral displacements whereby, due to very small displacements, this medial shift is not obvious in knee model no. 4. Maximum values of condylar enlargements show clearly that in a medio-lateral direction, states of elongation as well as of compression occur. Assuming a condylar width of 75 mm, the calculation of the mean condylar strain shows that the measured maximum condylar enlargements can reach values in the range of the fracture elongation of cancellous bone under tensile load of 0.6% (Perren et al., 1977).

The results of these experimental investigations correspond with those of the finite element analysis (Nambu et al., 1991) in that both studies show strains of an alternating amount with tensile and compressive alternations in the area of OD predilection sites. Whereas the advantages of analytical investigations are in the precise localisation of result determination, the experimental investigation can take into consideration anatomical differences of the individual specimens, especially with regard to the properties of cancellous bone, of joint cartilage and of geometry. Measured results show essential specimen-specific differences and thereby explain the importance of an experimental study. Compared to the lateral condyle, both proceedings can establish larger displacements and deformations in the medial condyle. Depressions and protuberances were measured in experimental proceedings whereas concentrations of stresses and strains occurred at corresponding joint surfaces in the analytical studies. Consequently, the possibility of local overload is supported by the results of both studies.

Further and even more detailed investigations of younger knee models, perhaps even of specimens with the beginning of OD, could bring additional findings in connection with the hypothesis serving as a base here. A variation of the tibia1 force in the frontal plane could explain a possible influence of the varus-valgus position. A preciser introduction of the quadriceps tendon tensile force by separation into different muscle groups such as M. rectus femoris, M. vastus medialis and M. vastus lateralis with corresponding tensile directions and amounts as well as repeated charges of extreme loads could contribute to the corroboration of the established results.

5. Conclusions

The investigation of only three knee models already shows the importance of experimental procedures and with it the consideration of anatomical differences to clarify the pathogenesis of osteochondrosis dissecans (OD). In spite of reservations regarding nonmeasurable modifications during the test process such as viscoelastic material behaviour or slackening of the capsule ligament apparatus, the selected procedure could determine with sufficient accuracy the condylar movements.

The measured quasi-axial and medio-lateral condylar displacements are small and specific to the different knee models. They act dynamically over an angle range of 0” to 120’ of knee flexion while changing the amount as well as the direction which leads to depressions and protuberances with corresponding compressive and tensile loads in the cancellous bone of the femoral condyles. Besides this alternating load and based on the measured displacements, the condyles were additionally subjected to a medial shift simulating a solid body displacement. Corresponding to the predilected sites of OD, larger displacement amplitudes were established in the medial condyle compared to the lateral one, especially in a quasi- axial direction. The amount of the average medio- lateral condylar enlargement can easily lead to tensile expansions in the area of the fracture elongation of cancellous bone.

Therefore, the measured condylar displacements and deformations can support our hypothesis. Accordingly, local overload of the subchondral cancellous bone due to mechanically induced joint deformation and subsequent fatigue fractures can lead to the formation of the osteochondral fragment which is typical of OD. Anatomical differences based on condylar geometry and bone properties can contribute to the mechanical overload.

Acknowledgements: The authors would like to gratefully acknowledge the assisstance of Dr. F. Bodem and Dr. A Wackerhagen (Biomech. Laboratory, Univ. of Mainz) during the experiments and the contributions of Mr. H. Wyder and Mr. D. Wyder (Miiller - Institute for Biomechanics, Univ. of Beme) for developing and manufacturing the measuring transducers and calibration devices. Many thanks to Mrs. H. Dobler (Robert Mathys Foundation, Bettlach) for translation of the manuscript.

Gasser: Contribution to the pathogenesis s17

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Ahmed A.M., Burke D.L. and Yu A. (1983) In Vitro Measurement of Static Pressure Distribution in Synovial Joints, Part II. J. Biomech. Engineering 105, 226 - 236.

Bandi W. (1982) Die retropatellaren Kniegelenks- schtiden. 2nd edition, Hans Huber, Berne.

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Burckhardt H. (1924) Uber die Entstehung der freien Gelenkkijrper und Mechanik des Kniegelenks. Bruns Beitr. Klin. Chir. 130,205.

Campell C.J. and Ranavat Chitranjan S. (1966) Osteochondritis dissecans: The Question of Etiology. J. Trauma, 6 (21,201.

Carter D.R., Rapperport D.J., Fyhrie D.P. et al. (1987) Relation of coxarthrosis to stresses and morphogenesis. A finite element analysis. Acta Orthop. Scud., 58 (61,611 - 619.

Enneking W.F. (1977) Clinical Musculoskeletal Pathology. Shorter, Gainsville.

Jakob R.P. (1990) Die Behandlung der Osteochondrosis dissecans am Kniegelenk mit einem neuen Kompressionsdrahtsystem. Z. Unfallchir. Vers. Med., 83 (21 104 -110.

Nagura S. (1937) Das Wesen und die Entstehung der Osteochondrosis dissecans am wachsenden Knochen. Zbl. Chir. 35,2049.

Nambu T., Gasser B., Schneider E. et al. (1991) Deformation of the Distal Femur: A Contribution towards the Pathogenesis of Ostechondritis Dissecans in the Knee Joint. 1. Biomech., 24 (61,421 - 433.

Perren S.M. and Cordey J. (1977) Die Gewebedifferenzienmg in der Frakturheilung. UnfalZheiZkunde 80,161 - 164.

Ruetsch H. and Morscher E. (1977) Measurement of the rotatory stability of the knee joint. In: Chapchal G. (Ed.) Injuries of the ligaments and their repair. Thieme, Stuttgart, pp. 116 - 122.

Koenig F. (1887) ijber freie K&per in den Gelenken. Deutsche Z. Chir. 27,90 - 109.


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Contribution to the pathogenesis of osteochondrosis dissecans: Experimental investigations on the idea of a mechanical overload of the cancellous bone