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In situ measurement of articular cartilage deformation in intact femoropatellar joints under static loading
Journal of Biomechanics 32 (1999) 1287}1295
In situ measurement of articular cartilage deformation in intact femoropatellar joints under static loading C. Herberhold , S. Faber, T. Stammberger, M. Steinlechner, R. Putz , K.H. Englmeier, M. Reiser, F. Eckstein * Musculoskeletal Research Group, Institute of Anatomy, Ludwig-Maximilians-Universita( t, Pettenkoferstr. 11, D-80336 Munich, Germany Institut fu( r Radiologische Diagnostik, Klinikum Gro}hadern, Ludwig-Maximilians-Universita( t, D-81377 Munich, Germany Institut fu( r Medizinische Informatik und Systemforschung, GSF-Forschungszentrum fu( r Umwelt und Gesundheit Neuherberg, D-85764 Oberschlei}heim, Germany Institut fu( r Gerichtliche Medizin der Universita( t Innsbruck, Mu( llerstrasse 44/III, A-6020 Innsbruck, Austria Accepted 14 June 1999
Abstract The deformational behavior of articular cartilage has been investigated in con"ned and uncon"ned compression experiments and indentation tests, but to date there exist no reliable data on the in situ deformation of the cartilage during static loading. The objective of the current study was to perform a systematic study into cartilage compression of intact human femoro-patellar joints under shortand long-term static loading with MR imaging. A non-metallic pneumatic pressure device was used to apply loads of 150% body weight to six joints within the extremity coil of an MRI scanner. The cartilage was delineated during the compression experiment with previously validated 2D and 3D fat-suppressed gradient echo sequences. We observed a mean (maximal) in situ deformation of 44% (57%) in patellar cartilage after 3 h of loading (mean contact pressure 3.6 MPa), the femoral cartilage showing a smaller amount of deformation than the patella. However, only around 7% of the "nal deformation (3% absolute deformation) occurred during the "rst minute of loading. A 43% #uid loss from the interstitial patellar matrix was recorded, the initial #uid #ux being 0.217$0.083 lm/s, and a high inter-individual variability of the deformational behavior (coe$cients of variation 11}38%). In conjunction with "nite-element analyses, these data may be used to compute the load partitioning between the solid matrix and #uid phase, and to elucidate the etiologic factors relevant in mechanically induced osteoarthritis. They can also provide direct estimates of the mechanical strain to be encountered by cartilage transplants. 1999 Elsevier Science Ltd. All rights reserved. Keywords: Articular cartilage; Cartilage mechanics; MR imaging; Joint loading
1. Introduction Magnetic resonance (MR) imaging has recently evolved as a technique for accurate measurement of cartilage volume and thickness in intact joints in vitro and in vivo (Peterfy et al., 1994; Eckstein et al., 1996,1997,1998a,b; Piplani et al., 1996; Tieschky et al., 1997; Cohen et al., 1999; Stammberger et al., 1999a). This implies the possibility of measuring cartilage deformation and volume changes in intact joints, for instance after physical activity (Eckstein et al., 1998c). Up to now, the deformational behavior of articular cartilage usually has been examined
* Corresponding author. Tel.:#49-89-5160-4847; fax:#49-89-51604802. E-mail address: [email protected] (F. Eckstein)
in explants or exposed articular surfaces (e.g. in con"ned or uncon"ned compression of excised tissue samples or with metallic indentation devices * Hayes et al. (1972), Mow et al. (1984,1989), Athanasiou et al. (1991), Suh and Spilker (1994), Mow and Ratcli!e (1997) and Froimson et al. (1997)). Although these experiments have provided important insight into the function of articular cartilage and have been a valuable basis for the theoretical characterization of its mechanical behavior, there exist considerable di!erences to the situation in the intact joint. The anchoring of the collagen "brils in the surrounding matrix, for example, and the resistance to #uid redistribution is altered in con"ned and uncon"ned compression, in which the explants have been cut at their edges. In indentation experiments, the deformational behavior depends on the form and permeability of the arti"cial indenter (Hayes et al., 1972) and represent data for
0021-9290/99/$ - see front matter 1999 Elsevier Science Ltd. All rights reserved. PII: S 0 0 2 1 - 9 2 9 0 ( 9 9 ) 0 0 1 3 0 - X
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cartilage}indenter, but not for cartilage}cartilage contact conditions. In the intact joint, however, the contact areas and pressure distribution are a function of both the magnitude and time of the load application (due to the natural incongruity of the joint surfaces), and these conditions are not adequately simulated in these experiments. Furthermore, opening of the joint capsule alters the natural biophysical boundary conditions and may in#uence the results of mechanical testing. Only sparse data are available on the in situ cartilage deformation in intact joints. Armstrong et al. (1979) used X-ray arthrography and Wayne et al. (1998) radiographs to determine cartilage deformation in human joints upon mechanical loading. However, the deformation throughout the joint surface or volume changes cannot be adequately characterized with projectional techniques. Moreover, Armstrong et al. (1979) had to partially open the joint capsule, in order to place metallic markers into the cartilage, and a contrast medium was injected into the joint space, to visualize the articular surface. The in situ deformation of articular cartilage has also been determined in intact joints after (but not during) mechanical loading in an animal experiment (in vitro loading simulation and subsequent histologic analysis * KaK aK b et al. (1998)) and in humans after knee bends (MR imaging and 3D reconstruction * Eckstein et al. (1998c,1999)). These experiments, however, have not allowed characterization of the time-dependent deformation of cartilage under a de"ned static load. Precise knowledge of the actual deformation of the cartilage during static loading is, nevertheless, one important prerequisite for calculating the hydrostatic interstitial #uid pressure, #uid #ow, and solid matrix stress in the cartilage tissue based on theoretical models (Spilker et al., 1992; Ateshian et al., 1994; Ateshian and Wang, 1995; Suh et al., 1995; Wu et al., 1998; Donzelli and Spilker, 1998). The elastic solid (proteoglycan}collagen) matrix stress has been suggested to play a crucial role in the pathogenesis of mechanically induced cartilage degradation (osteoarthritis), but the elastic matrix stress cannot be measured experimentally, and must be derived from theoretical models (Mow et al., 1993; Ateshian et al., 1994). In order to be able to provide estimates for the situation of a normal, intact joint, relevant input data are required for these calculations (e.g. Donzelli and Spilker, 1998). Moreover, data on the in situ deformation can provide estimates of the mechanical stress to be encountered by transplanted cartilage tissue (autologous or tissue-engineered), within its target environment. Herberhold et al. (1998) have recently presented a non-metallic apparatus for loading intact femoro-patellar joints with 150% body weight within the extremity coil of a clinical MR scanner. The objective of the current study was to re"ne this technique and to perform a systematic study into the short- and long-term in situ deformation in human femoro-patellar joints under static
loading with MR imaging. In this context we focused on the following speci"c questions: 1. What amount of deformation occurs in the joint during long-term static loading? 2. Do the patellar and femoral cartilage show di!erences in deformational behavior? 3. What part of the "nal deformation occurs during the "rst few minutes of compression? 4. How large are the volume changes and the rate of the interstitial #uid loss during loading? 5. Are there large di!erences in the deformational behavior between individuals?
2. Materials and methods 2.1. Specimens and loading apparatus Six normal knee joint specimens (age 18}67 yr [mean 42.5 yr], 4 male, 2 female; body weight 55}84 kg; Table 1) were harvested within 48 h of death, frozen at !203C, and gradually thawed at room temperature before the investigation. Exposure of the joint surfaces (after the compression experiments) did not reveal any signs of joint disease or cartilage degradation. The joints were mounted in the adapted compression apparatus (Herberhold et al., 1998), which "ts into the extremity coil of a clinical MRI scanner (1.5 T, Magnetom Vision, Siemens, Erlangen, Germany) (Fig. 1A). The femora were "rmly "xed (no degree of freedom) to the basis of the apparatus at a 403 angle by guiding a rope through a medio-lateral drill hole (situated at the level of the epicondyles), the joint capsule being left intact (Fig. 1B). The tibia was positioned on the 203 slant at the opposite side, resulting in a 603 #exion of the knee joint. The patella was "rmly attached to a mobile sleigh with bone cement (Durus Arthroplasty Research and Development Ltd., South Africa), the sleigh (and the patella) being able to move freely within the transverse image plane of the MR-scanner translation, but no tilting or shifting was possible relative to this plane (3 degrees of freedom "xed and 3 left free) (Herberhold et al., 1998). This allows one to use faster 2D pulse sequence for the thickness measurements during the "rst minutes of the compression experiment, since projectional artifacts due to relative motion of the patella are avoided. For improving the cited technique, a new pneumatic pressure device was constructed, consisting of a pressure piston in a cylindrical frame, the magnitude of load being independent of the distance that the piston has to travel until it meets mechanical resistance (Fig. 1B). The pressure device was attached to the top of the sleigh and the reaction force transferred to the basis of the apparatus through a "rm nylon band. The calibration of the pressure device was performed with a uniaxial material-testing machine
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Table 1 Characteristics of the investigated specimens. Age, sex, body weight, maximal and mean cartilage thickness, cartilage volume and articular surface of the patellae (determined from 3D reconstruction of the unloaded patellar cartilage) Specimen
Age (yr)
Sex
Body Weight (kg)
Maximal thickness Mean thickness (mm) (mm)
Cartilage volume (mm)
Articular surface area (cm)
1 2 3 4 5 6
31 67 18 51 65 23
m m m f f m
58 55 84 65 65 77
3.8 5.2 5.3 5.2 4.7 5.2
1.9 2.3 2.7 2.2 2.1 2.8
3680 3640 4620 2745 2560 5150
14.8 14.5 16.9 11.6 11.6 17.8
Mean SD
42.5 20
70 10
4.9 0.54
2.3 0.31
3730 930
14.5 2.35
Fig. 1. Compression apparatus for applying a load of 150% body weight to an intact femoro-patellar joint within the extremity coil of a clinical MRI scanner. (A) Apparatus in the extremity coil of the scanner, in front of the main magnet, (B) apparatus with a femoro-patellar specimen. Please note the pneumatic pressure device on top of the apparatus.
(Zwick 1445, Ulm, Germany). As a control, sagittal MR images were recorded before and at the end of load application, to assess whether the femur or the patella had tilted or shifted relative to the transverse image plane.
2.2. MR imaging Before load application, two sets of transverse images of the femoro-patellar joint were recorded with a
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previously validated fat-suppressed 3D gradient echo sequence (FLASH) at a resolution of 2;0.31;0.31 mm (FOV"16cm, matrix"512 pixels, TR"45 ms, TE"11 ms, FA"303, N. acq."1, imaging time"8 min). The mean values obtained from these data sets were used as a reference (for the volume and 3D thickness changes during compression), and as a basis for the selection of one transverse slice within the center of the contact area. This 2D slice was chosen to monitor the thickness changes with a fast 2D MR sequence during the "rst minutes of loading (N. acq."4, imaging time"70 s, imaging parameters otherwise identical to the 3D sequence described above) and with the 3D sequence up to 214 min after load application. Three transverse MR-images were recorded with this 2D sequence before compression. The patella was then loaded with 150% body weight (Table 1) within approx. 10 s, and six 2D images were recorded at intervals of 90 s up to 9 min after load application, the load being kept constant throughout the entire experiment. 3D data sets of the entire patellar cartilage (approx. 25 sections, including the selected central slice) were registered at 10 min intervals from 10 to 210 min after load application ("rst acquisition from 10 to 18 min). The middle of the acquisition time was used to characterize the time-dependant deformation of the cartilage (e.g. 14 min for the "rst 3D acquisition). In one specimen the compression experiment was extended by a recovery experiment, 3D data sets of the patellar cartilage being acquired from the moment of load release at 10 min intervals over 4 h.
#ux in lm/s) from the volume di!erences at various time steps. To be able to track the location- and time-dependant cartilage thickness changes, it was mandatory to reidentify corresponding points in the consecutive MR data sets. In the 2D slices, a scalar function was de"ned (representing the distance between corresponding cartilage}bone interfaces in consecutive images) and minimized within the parameter space of allowed patellar movements (translation in two axes, rotation around the center of mass * Herberhold et al. (1998), and Stammberger et al. (1998)). For evaluating the cartilage deformation throughout the joint surface, a new 3D matching algorithm was applied, which adjusts the corresponding bone}cartilage interfaces in consecutive 3D image data sets, and allows one to directly visualize the spatial di!erence between the thickness distribution patterns (Stammberger et al., unpublished work). 2.4. Joint pressure analysis with FUJI xlm Following imaging, the joints were left in an unloaded position for 8 h, opened, and two pieces of pressuresensitive "lm (FUJI low sensitivity: 2.5}10 MPa) positioned within the joint, one at the lateral and one at the medial patellar facet. The pressure device was then in#ated for 30 s to the same value as during the compression experiment. The "lms were "nally calibrated with a 0.5 cm stamp in a material-testing machine, applying a set of de"ned loads, and the staining of the "lm converted into pressure intervals (MPa) with image analysis.
2.3. Digital image analysis 3. Results Following imaging, all data sets were transferred digitally to a symmetrical multiprocessing computer (Octane Duo, Silicon Graphics Inc., Mountain View, CA, USA), the segmentation, 3D reconstruction, and digital image analysis being performed with the software designed in our laboratory (detailed descriptions in Herberhold et al., 1998; Eckstein et al., 1998a,b; Stammberger et al., 1998,1999b). Based on a Euclidean distance transformation algorithm (Stammberger et al., 1999a), these techniques make it possible to compute the total cartilage volume, the articular surface area, and the 3D mean and maximal cartilage thickness, independent of the original slice orientation. Whereas the 3D analyses were con"ned to the patella, the mean and maximal cartilage thickness within the selected 2D slice were determined for both the patellar and femoral cartilage layers. Based on the assumptions that the solid matrix is incompressible (Bachrach et al., 1998) (and therefore all volume changes are due to #uid #ow), and that during compression the #uid #ow will occur through the articular surface into the joint cavity, we estimated a #ow rate per square centimetre of the articular surface (the #uid
The maximal 3D cartilage thickness (before compression) in the six patellae ranged from 3.8 to 5.3 mm, and the mean thickness from 1.9 to 2.8 mm (Table 1). The cartilage volumes amounted to between 2560 and 5150 mm and the surface areas to between 11.6 and 17.8 cm (Table 1). During the compression experiment, the patella showed a shift of 3.8 ($1.1) mm in the direction of the applied force and of 0 ($0.4) mm perpendicular to it within the selected 2D image plane. The average rotation was !0.5 ($2)3. Movements perpendicular to the image plane did not occur, as demonstrated by comparison of sagittal images before and after the experiment. The analysis of the cartilage deformation in the selected central 2D slice showed a mean reduction of 44$15% of the initial thickness in the patella, and of 30$10% in the femur after 214 min of static loading with 150% body weight (Fig. 2A, Table 2), the deformation of the patella generally exceeding that of the femur. The maximal deformation was 57$15% in the patella and 44$7 in the femur (Fig. 2B, Table 2). The cartilage
C. Herberhold et al. / Journal of Biomechanics 32 (1999) 1287}1295
thickness decreased in an approximately exponential manner (Figs. 2A and B). Only 7$10% of the "nal (214 min) deformation was reached during the "rst minute of static loading of the
Fig. 2. In situ thickness changes during loading with 150% body weight. The error bars show one standard deviation of the variability between individuals. (A) Mean cartilage deformation in the selected central 2D slice within the contact area, (B) Maximal cartilage deformation in the selected central 2D slice. A mean (maximal) in situ deformation of 44% (57%) was observed in the patellar cartilage after 3 h of static loading (mean contact pressure"3.6 MPa). The femoral cartilage generally showed a smaller amount of in situ deformation than the patella.
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patellar cartilage (3% absolute deformation) and 25$10% after approx. 8 min (11% absolute deformation), the values for the femur being 4$12% and 30$12% (absolute deformation 1.3 and 9%), respectively (Fig. 3). The patellar cartilage volume was reduced by 8$5% after 14 min and by 29$3.2% after 214 min of loading (Fig. 4A, Table 2). The mean 3D patellar thickness decreased by 10$4% (14 min) and 29$5% (214 min), respectively, the time-dependant 3D thickness changes being very similar to those of the volume (Fig. 4A). The maximal and mean deformation in the selected 2D slice showed a high linear correlation with the volume changes observed in the 3D data (r"0.96 and 0.82, respectively). The rate of interstitial #uid loss from the matrix was 1.3 ($0.5) mm/min per square centimeter surface area (#uid #ux"0.217$0.083 lm/s] for the "rst 14 min of loading, and 0.22 ($0.04) mm/min cm (0.037$0.007 lm/s) in the terminal phase of the experiment ('120 min). There was a high interindividual variability of the deformational behavior, both in the 2D analysis (mean and maximal thickness) and in the 3D analysis (volume change, #uid #ux), the coe$cients of variation ranging from 11% (volume) to 38% (#ux) (Table 2). The thickness changes were highly inhomogeneous throughout the joint surface, the site of maximal deformation (the lateral patellar facet) being identical with that of the maximal cartilage thickness (Fig. 5). The FUJI "lm experiment conducted after imaging showed pressure maxima of 3.6$1.25 MPa, in the lateral patellar facet (Table 2), the pressure distribution being very similar to the pattern of cartilage deformation (Fig. 5). The one specimen used for the recovery experiment displayed almost full recovery (98%) of the patellar cartilage after approx. 4 h, the rate of #uid re-uptake by the cartilage matrix being 0.9 mm/min cm (0.15 lm/s) within the "rst 14 min of recovery, and 0.14 mm/min cm (0.023 lm/s) in the terminal phase.
Table 2 Deformational behavior of the di!erent specimens during loading with 150% body weight (2D and 3D analysis, patellar cartilage, maximal (mean) deformation and volume changes after 3.5 h of static loading) (1 mm/min cm is equivalent to 0.167 lm/s) Specimen
Mean deformation 2D slice (%)
Maximal deformation 2D slice (%)
Volume change Peak initial #ux (%) (mm/min cm))
1 2 3 4 5 6
30 27 34 70 52 48
60 41 40 84 65 53
29 26 25 33 32 28
Mean SD CV%
44 15 34
57 15 26
29 3.2 11
1.4 0.5 0.75 2.1 1.4 1.4 1.3 0.5 38
Terminal #ux (mm/min cm))
Articular pressure (MPa)
0.25 0.18 0.25 0.21 0.16 0.24
3.0 1.5 3.75 5.25 5.25 3.0
0.22 0.04 18
3.6 1.25 35
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Fig. 3. Cartilage deformation in the "rst minutes of loading (150% body weight) measured in the selected central 2D slice within the contact area. The error bars show one standard deviation of the variability between individuals. Only approx. 7% of the "nal deformation of the patellar cartilage (3% absolute deformation) occurred after the "rst minute, and 25% of the "nal deformation (11% absolute deformation) after 8 min of compression. The values for the femur were 4% (1.3% absolute deformation) and 30% (9% absolute deformation), respectively.
4. Discussion The objective of the current study was to investigate the in situ deformation of articular cartilage in the intact femoro-patellar joint under static loading with MR imaging. To our knowledge, this is the "rst systematic study to report the in situ deformation and volume changes of articular cartilage in an intact joint, under normal, nonlinear contact conditions and with the joint capsule being intact. Because the compressive behavior of the cartilage over the "rst few minutes of loading is physiologically most relevant, we analyzed the initial deformation with 2D MR sequences. These allow faster acquisition than 3D protocols (imaging time around 1 min) and have been shown to produce results identical to those of previously validated 3D fat-suppressed gradient echo sequences (Herberhold et al., 1998). To exclude movement relative to the image plane (which would interfere with accurate 2D measurement) the patellar motion was e!ectively restricted by the apparatus to the transverse image plane. The results obtained for the selected central 2D slice are shown to be consistent with those obtained for the entire patella (3D data) after long-term loading, and thus appear representative for comparative measurements. We "nd a mean (maximal) in situ deformation of 44% (57%) in the patellar cartilage after 3 hrs of static loading with 150% body weight (mean contact pressure"3.6 MPa). Only approximately 7% of the "nal deformation (3% absolute deformation) occurred during the "rst minute, and 25% (11% absolute deformation) during 8 min of loading. The data show that, under in
Fig. 4. In situ volume changes and #uid #ux during loading with 150% body weight. The error bars show one standard deviation of the variability between individuals. (A) Volume change of the patellar cartilage (3D analysis), (B) rate of #uid loss from the interstitial patellar cartilage matrix per square centimeter surface area (and #uid #ux in lm/s), based on the assumption that during compression the #uid #ow will occur through the articular surface into the joint cavity. A 43% #uid loss was observed from the interstitial matrix after 3 h of static loading, the #uid #ow rate per square centimeter articular surface being initially 1.3 ($0.5) mm/min cm (0.217$0.083 lm/s) for the "rst 14 min after loading, and 0.22 ($0.04) mm/min cm (0.037$ 0.007 lm/s) in the terminal phase of the experiment ('120 min).
situ conditions, only a small fraction of the "nal deformation (steady state) is reached for physiologically relevant loading periods, which typically last several seconds or minutes. These results are consistent with previous theoretical analyses (Mak et al., 1987; Ateshian et al., 1994), showing that there is little time for #uid to escape initially and therefore little change in the cartilage volume, due to the low permeability of the cartilage. These "ndings support the idea that the initial response of articular cartilage is nearly incompressible elastic, and that the #uid supports a substantial portion of the applied load
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Fig. 5. Cartilage thickness changes throughout the patellar cartilage surface. 3D matching of corresponding bone}cartilage interfaces in the consecutive 3D image data sets. (A) Cartilage thickness distribution before loading (The 8 gray values indicate thickness intervals of 0.6 mm (black"0}0.6 mm; white"'4.2 mm), (B) cartilage thickness distribution after loading [same gray value legend as for (A)], (C) direct visualization of thickness di!erences after 3D matching (dark"areas without deformation; white"areas of maximal deformation (gray values indicate intervals of 0.4 mm). There is a highly inhomogeneous deformation throughout the joint surface, the maximal deformation (2.4 mm) occurring in the lateral patellar facet, at the site of maximal cartilage thickness [see (A)].
(Ateshian et al., 1994; Ateshian and Wang, 1995), the apparent sti!ness of the tissue being much greater in the instantaneous than in the equilibrium response. The mechanism of hydrostatic pressurization of the interstitial #uid, which is assumed to protect the solid matrix from elastic deformation for several hundred seconds after the onset of loading (Ateshian et al., 1994; Ateshian and Wang, 1995) is assumed to be a crucial factor in protecting the matrix from undue stress and tissue damage during normal loading. The femoral cartilage generally showed a smaller amount of in situ deformation than the patella, this being consistent with the results of Froimson et al. (1997) who found a lower sti!ness and higher permeability of the patellar cartilage in relation to the femur in indentation studies. This demonstrates that the technique presented is able to e!ectively reveal di!erences in the mechanical behavior (and material properties) of di!erent cartilage plates. The observed di!erences may also be relevant in explaining the empirical observation that the patella generally shows earlier and more severe signs of osteoarthritic changes than the femur (Ficat and Hungerford,
1977; Froimson et al., 1997). Considering that the cartilage consists of 70% #uid (Mow and Ratcli!e, 1997), a 43% #uid loss from the interstitial matrix was observed in this study after 3 h of static loading. Assuming that during compression the #uid #ow occurs through the articular surface into the joint cavity, the #uid #ux was 0.216 lm/s initially, and 0.037 lm/s at the terminal phase of the experiment. The rate of #uid re-uptake observed during recovery (0.023 lm/s) in the terminal phase * one specimen) is in the same order of magnitude as that observed in an in vivo recovery experiment after 100 knee bends (0.027 lm/s) after a 5% volume change (Eckstein et al., 1999). These results suggest that the data obtained with the compression apparatus can, within some limits, be transferred to the in vivo situation. The #uid #ow through the joint surface can apparently be increased by a factor of approximately 10 over that during free recovery, when statically loading the cartilage with 150% body weight and with peak pressures of around 3.6 MPa. The high inter-individual variability observed in this study is also in agreement with previous indentation
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experiments (Athanasiou et al., 1991; Froimson et al., 1997) and in vivo results (Eckstein et al., 1998c,1999), both the deformation and the recovery of the cartilage yielding considerable di!erences between individuals. It is up to speculation whether these observed di!erences relate to the observation that some individuals show early signs of cartilage damage, whereas in others the cartilage maintains normal morphology and function over several decades (Mow and Ratcli!e, 1997). The data presented here may serve as a basis for "nite-element simulations, to validate the non-linear contact algorithms developed for biphasic cartilage layers (Donzelli and Spilker, 1998). By matching the experimental and computational results, it should become feasible to analyze the load partitioning between the solid (proteoglycan}collagen) matrix and #uid phase, to determine the #uid #ux within the cartilage matrix, and to elucidate etiologic factors relevant in mechanically induced osteoarthritis. The data can also provide a direct estimate of the mechanical strain to be encountered by autologous or tissue-engineered cartilage transplants, once they are located in their target environment.
Acknowledgements Parts of this study were performed within the framework of the doctoral thesis (in preparation) of Christoph Herberhold at the Medical Faculty of the LudwigsMaximilians-UniversitaK t, MuK nchen, Germany. We would like to thank Professor Gerard Ateshian, Columbia University (New York) for his very helpful discussion and the Deutsche Forschungsgemeinschaft (DFG) for supporting this work.
References Armstrong, C.G., Bahrani, A.S., Gardner, D.L., 1979. In vitro measurement of articular cartilage deformations in the intact human hip joint under load. Journal of Bone and Joint Surgery 61A, 744}755. Ateshian, G.A., Lai, W.M., Zhu, W.B., Mow, V.C., 1994. An asymptotic solution for the contact of two biphasic cartilage layers. Journal of Biomechanics 27, 1347}1360. Ateshian, G.A., Wang, H., 1995. A theoretical solution for the frictionless rolling contact of cylindrical biphasic articular cartilage layers. Journal of Biomechanics 28, 1341}1355. Athanasiou, K.A., Rosenwasser, M.P., Buckwalter, J.A., Malinin, T.I., Mow, V.C., 1991. Interspecies comparison of in situ intrinsic mechanical properties of distal femoral cartilage. Journal of Orthopedic Research 9, 330}340. Bachrach, N.M., Mow, V.C., Guilak, F., 1998. Incompressibility of the solid matrix of articular cartilage under high hydrostatic pressures. Journal of Biomechanics 31, 445}451. Cohen, Z.A., McCarthy, D.M., Kwak, S.D., Legrand, P., Fogarasi, F., Ciaccio, E.J., Ateshian, G.A., 1999. Knee cartilage topography, thickness, and contact areas from MRI: in vitro calibration and in vivo measurements. Osteoarthritis and Cartilage 7, 95}109.
Donzelli, P., Spilker, R.L., 1998. A contact "nite element formulation for biological soft hydrated tissues. Computer Methods in Applied Mechanics and Engineering 153, 68}79. Eckstein, F., Adam, C., Sittek, H., Becker, C., Milz, S., Schulte, E., Reiser, M., Putz, R., 1997. Non-invasive determination of cartilage thickness throughout joint surfaces using magnetic resonance imaging. Journal of Biomechanics 30, 285}289. Eckstein, F., Gavazzeni, A., Sittek, H., Haubner, M., LoK sch, A., Milz, S., Englmeier, K.H., Schulte, E., Putz, R., Reiser, M., 1996. Determination of knee joint cartilage thickness using three-dimensional magnetic resonance chondro-crassometry (3D MR-CCM). Magnetic Resonance in Medicine 36, 256}265. Eckstein, F., Schnier, M., Haubner, M., Priebsch, J., Glaser, C., Englmeier, K.-H., Reiser, M., 1998a. Accuracy of three-dimensional knee joint cartilage volume and thickness measurements with MRI. Clinical Orthopedics and Related Reserach 352, 137}148. Eckstein, F., Westho!, J., Sittek, H., Maag, K.P., Haubner, M., Faber, S., Englmeier, K.H., Reiser, M., 1998b. In vivo reproducibility of three-dimensional cartilage volume and thickness measurements with MR imaging. American Journal of Roentgenology 170, 593}597. Eckstein, F., Tieschky, M., Faber, S., Haubner, M., Kolem, H., Englmeier, K.-H., Reiser, M., 1998c. E!ects of physical exercise on cartilage volume and thickness in vivo * an MR imaging study. Radiology 207, 243}248. Eckstein, F., Tieschky, M., Faber, S., Englmeier, K.-H., Reiser, M., 1999. Functional analysis of articular cartilage deformation, recovery, and #uid #ow following dynamic exercise in vivo. Anatomy and Embryology, in press Ficat, P., Hungerford, D.S., 1977. Disorders of the Patellofemoral Joint. Masson, Paris. Froimson, M.I., Ratcli!e, A., Gardner, T.R., Mow, V.C., 1997. Di!erences in patellofemoral joint cartilage material properties and their signi"cance to the etiology of cartilage surface "brillation. Osteoarthritis and Cartilage 5, 377}386. Hayes, W.C., Keer, L.M., Herrmann, G., Mockros, L.F., 1972. A mathematical analysis for indentation tests of articular cartilage. Journal of Biomechanics 5, 541}551. Herberhold, C., Stammberger, T., Faber, S., Putz, R., Englmeier, K.-H., Reiser, M., Eckstein, F., 1998. An MR-based technique for quantifying the deformation of articular cartilage during mechanical loading in an intact cadaver joint. Magnetic Resonance in Medicine 39, 843}850. KaK aK b, M.J., Ito, K., Clark, J.M., NoK tzli, H.P., 1998. Deformation of articular cartilage collagen structure under static and cyclic loading. Journal of Orthopedic Research 16, 743}751. Mak, A.F., Lai, W.M., Mow, V.C., 1987. Biphasic indentation of articular cartilage * I. Theoretical analysis. Journal of Biomechanics 20, 703}714. Mow, V.C., Ateshian, G.A., Spilker, R.L., 1993. Biomechanics of diarthrodial joints: a review of twenty years of progress. Journal of Biomechanic Engineering 115, 460}467. Mow, V.C., Gibbs, M.C., Lai, W.M., Zhu, W.B., Athanasiou, K.A., 1989. Biphasic indentation of articular cartilage. II. A numerical algorithm and an experimental study. Journal of Biomechanics 22, 853}861. Mow, V.C., Holmes, M.H., Lai, W.M., 1984. Fluid transport and mechanical properties of articular cartilage: a review. Journal of Biomechanics 17, 377}394. Mow, V.C., Ratcli!e, A., 1997. Structure and function of articular cartilage and meniscus. In: Mow, V.C., Hayes, W.C. (Eds.), Basic Orthopaedic Biomechanics. Lippincott-Raven, New York, pp. 113}177. Peterfy, C.G., van Dijke, C.F., Janzen, D.L., Gluer, C.C., Namba, R., Majumdar, S., Lang, P., Genant, H.K., 1994. Quanti"cation of articular cartilage in the knee with pulsed saturation transfer subtraction and fat-suppressed MR imaging: optimization and validation. Radiology 192, 485}491.
C. Herberhold et al. / Journal of Biomechanics 32 (1999) 1287}1295 Piplani, M.A., Disler, D.G., McCauley, T.R., Holmes, T.J., Cousins, J.P., 1996. Articular cartilage volume in the knee: semiautomated determination from three-dimensional reformations of MR images. Radiology 198, 855}859. Spilker, R.L., de Almeida, E.S., Donzelli, P.S., 1992. Finite element methods for the biomechanics of soft hydrated tissues: nonlinear analysis and adaptive control of meshes. Critical Revue of Biomedical Engineering 20, 279}313. Stammberger, T., Herberhold, C., Faber, S., Englmeier, K.H., Reiser, M., Eckstein, F., 1998. A method for quantifying time dependant changes in MR signal intensity of articular cartilage as a function of tissue deformation in intact joints. Medical Engineering and Physics 20, 741}749. Stammberger, T., Eckstein, F., Michaelis, M., Englmeier, K.-H., Reiser, M., 1999b. Interobserver reproducibility of quantitative cartilage measurements: comparsion between B-spline snakes and manual segmentation. Magnetic Resonance Imaging, in press. Stammberger, T., Eckstein, F., Englmeier, K.-H., Reiser, M., 1999a. Determination of 3D cartilage thickness data from MR imaging
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* computational method and reproducibility in the living. Magnetic Resonance in Medicine 41, 529}536. Suh, J.K., Li, Z., Woo, S.L., 1995. Dynamic behavior of a biphasic cartilage model under cyclic compressive loading. Journal of Biomechanics 28, 357}364. Suh, J.K., Spilker, R.L., 1994. Indentation analysis of biphasic articular cartilage: nonlinear phenomena under "nite deformation. Journal of Biomechanical Engineering 116, 1}9. Tieschky, M., Faber, S., Haubner, M., Kolem, H., Schulte, E., Englmeier, K.-H., Reiser, M., Eckstein, F., 1997. Repeatability of patellar cartilage thickness patterns in the living, using a fat-suppressed MRI sequence with short acquisition time and 3D data processing. Journal of Orthopedic Research 15, 808}813. Wayne, J.S., Brodrick, C.W., Mukherjee, N., 1998. Measurement of cartilage thickness in the articulated knee. Annals of Biomedical Engineering 26, 96}102. Wu, J.Z., Herzog, W., Epstein, M., 1998. Articular joint mechanics with biphasic cartilage layers under dynamic loading. Journal of Biomechanical Engineering 120, 77}84.
In situ measurement of articular cartilage deformation in intact femoropatellar joints under static loading C. Herberhold , S. Faber, T. Stammberger, M. Steinlechner, R. Putz , K.H. Englmeier, M. Reiser, F. Eckstein * Musculoskeletal Research Group, Institute of Anatomy, Ludwig-Maximilians-Universita( t, Pettenkoferstr. 11, D-80336 Munich, Germany Institut fu( r Radiologische Diagnostik, Klinikum Gro}hadern, Ludwig-Maximilians-Universita( t, D-81377 Munich, Germany Institut fu( r Medizinische Informatik und Systemforschung, GSF-Forschungszentrum fu( r Umwelt und Gesundheit Neuherberg, D-85764 Oberschlei}heim, Germany Institut fu( r Gerichtliche Medizin der Universita( t Innsbruck, Mu( llerstrasse 44/III, A-6020 Innsbruck, Austria Accepted 14 June 1999
Abstract The deformational behavior of articular cartilage has been investigated in con"ned and uncon"ned compression experiments and indentation tests, but to date there exist no reliable data on the in situ deformation of the cartilage during static loading. The objective of the current study was to perform a systematic study into cartilage compression of intact human femoro-patellar joints under shortand long-term static loading with MR imaging. A non-metallic pneumatic pressure device was used to apply loads of 150% body weight to six joints within the extremity coil of an MRI scanner. The cartilage was delineated during the compression experiment with previously validated 2D and 3D fat-suppressed gradient echo sequences. We observed a mean (maximal) in situ deformation of 44% (57%) in patellar cartilage after 3 h of loading (mean contact pressure 3.6 MPa), the femoral cartilage showing a smaller amount of deformation than the patella. However, only around 7% of the "nal deformation (3% absolute deformation) occurred during the "rst minute of loading. A 43% #uid loss from the interstitial patellar matrix was recorded, the initial #uid #ux being 0.217$0.083 lm/s, and a high inter-individual variability of the deformational behavior (coe$cients of variation 11}38%). In conjunction with "nite-element analyses, these data may be used to compute the load partitioning between the solid matrix and #uid phase, and to elucidate the etiologic factors relevant in mechanically induced osteoarthritis. They can also provide direct estimates of the mechanical strain to be encountered by cartilage transplants. 1999 Elsevier Science Ltd. All rights reserved. Keywords: Articular cartilage; Cartilage mechanics; MR imaging; Joint loading
1. Introduction Magnetic resonance (MR) imaging has recently evolved as a technique for accurate measurement of cartilage volume and thickness in intact joints in vitro and in vivo (Peterfy et al., 1994; Eckstein et al., 1996,1997,1998a,b; Piplani et al., 1996; Tieschky et al., 1997; Cohen et al., 1999; Stammberger et al., 1999a). This implies the possibility of measuring cartilage deformation and volume changes in intact joints, for instance after physical activity (Eckstein et al., 1998c). Up to now, the deformational behavior of articular cartilage usually has been examined
* Corresponding author. Tel.:#49-89-5160-4847; fax:#49-89-51604802. E-mail address: [email protected] (F. Eckstein)
in explants or exposed articular surfaces (e.g. in con"ned or uncon"ned compression of excised tissue samples or with metallic indentation devices * Hayes et al. (1972), Mow et al. (1984,1989), Athanasiou et al. (1991), Suh and Spilker (1994), Mow and Ratcli!e (1997) and Froimson et al. (1997)). Although these experiments have provided important insight into the function of articular cartilage and have been a valuable basis for the theoretical characterization of its mechanical behavior, there exist considerable di!erences to the situation in the intact joint. The anchoring of the collagen "brils in the surrounding matrix, for example, and the resistance to #uid redistribution is altered in con"ned and uncon"ned compression, in which the explants have been cut at their edges. In indentation experiments, the deformational behavior depends on the form and permeability of the arti"cial indenter (Hayes et al., 1972) and represent data for
0021-9290/99/$ - see front matter 1999 Elsevier Science Ltd. All rights reserved. PII: S 0 0 2 1 - 9 2 9 0 ( 9 9 ) 0 0 1 3 0 - X
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cartilage}indenter, but not for cartilage}cartilage contact conditions. In the intact joint, however, the contact areas and pressure distribution are a function of both the magnitude and time of the load application (due to the natural incongruity of the joint surfaces), and these conditions are not adequately simulated in these experiments. Furthermore, opening of the joint capsule alters the natural biophysical boundary conditions and may in#uence the results of mechanical testing. Only sparse data are available on the in situ cartilage deformation in intact joints. Armstrong et al. (1979) used X-ray arthrography and Wayne et al. (1998) radiographs to determine cartilage deformation in human joints upon mechanical loading. However, the deformation throughout the joint surface or volume changes cannot be adequately characterized with projectional techniques. Moreover, Armstrong et al. (1979) had to partially open the joint capsule, in order to place metallic markers into the cartilage, and a contrast medium was injected into the joint space, to visualize the articular surface. The in situ deformation of articular cartilage has also been determined in intact joints after (but not during) mechanical loading in an animal experiment (in vitro loading simulation and subsequent histologic analysis * KaK aK b et al. (1998)) and in humans after knee bends (MR imaging and 3D reconstruction * Eckstein et al. (1998c,1999)). These experiments, however, have not allowed characterization of the time-dependent deformation of cartilage under a de"ned static load. Precise knowledge of the actual deformation of the cartilage during static loading is, nevertheless, one important prerequisite for calculating the hydrostatic interstitial #uid pressure, #uid #ow, and solid matrix stress in the cartilage tissue based on theoretical models (Spilker et al., 1992; Ateshian et al., 1994; Ateshian and Wang, 1995; Suh et al., 1995; Wu et al., 1998; Donzelli and Spilker, 1998). The elastic solid (proteoglycan}collagen) matrix stress has been suggested to play a crucial role in the pathogenesis of mechanically induced cartilage degradation (osteoarthritis), but the elastic matrix stress cannot be measured experimentally, and must be derived from theoretical models (Mow et al., 1993; Ateshian et al., 1994). In order to be able to provide estimates for the situation of a normal, intact joint, relevant input data are required for these calculations (e.g. Donzelli and Spilker, 1998). Moreover, data on the in situ deformation can provide estimates of the mechanical stress to be encountered by transplanted cartilage tissue (autologous or tissue-engineered), within its target environment. Herberhold et al. (1998) have recently presented a non-metallic apparatus for loading intact femoro-patellar joints with 150% body weight within the extremity coil of a clinical MR scanner. The objective of the current study was to re"ne this technique and to perform a systematic study into the short- and long-term in situ deformation in human femoro-patellar joints under static
loading with MR imaging. In this context we focused on the following speci"c questions: 1. What amount of deformation occurs in the joint during long-term static loading? 2. Do the patellar and femoral cartilage show di!erences in deformational behavior? 3. What part of the "nal deformation occurs during the "rst few minutes of compression? 4. How large are the volume changes and the rate of the interstitial #uid loss during loading? 5. Are there large di!erences in the deformational behavior between individuals?
2. Materials and methods 2.1. Specimens and loading apparatus Six normal knee joint specimens (age 18}67 yr [mean 42.5 yr], 4 male, 2 female; body weight 55}84 kg; Table 1) were harvested within 48 h of death, frozen at !203C, and gradually thawed at room temperature before the investigation. Exposure of the joint surfaces (after the compression experiments) did not reveal any signs of joint disease or cartilage degradation. The joints were mounted in the adapted compression apparatus (Herberhold et al., 1998), which "ts into the extremity coil of a clinical MRI scanner (1.5 T, Magnetom Vision, Siemens, Erlangen, Germany) (Fig. 1A). The femora were "rmly "xed (no degree of freedom) to the basis of the apparatus at a 403 angle by guiding a rope through a medio-lateral drill hole (situated at the level of the epicondyles), the joint capsule being left intact (Fig. 1B). The tibia was positioned on the 203 slant at the opposite side, resulting in a 603 #exion of the knee joint. The patella was "rmly attached to a mobile sleigh with bone cement (Durus Arthroplasty Research and Development Ltd., South Africa), the sleigh (and the patella) being able to move freely within the transverse image plane of the MR-scanner translation, but no tilting or shifting was possible relative to this plane (3 degrees of freedom "xed and 3 left free) (Herberhold et al., 1998). This allows one to use faster 2D pulse sequence for the thickness measurements during the "rst minutes of the compression experiment, since projectional artifacts due to relative motion of the patella are avoided. For improving the cited technique, a new pneumatic pressure device was constructed, consisting of a pressure piston in a cylindrical frame, the magnitude of load being independent of the distance that the piston has to travel until it meets mechanical resistance (Fig. 1B). The pressure device was attached to the top of the sleigh and the reaction force transferred to the basis of the apparatus through a "rm nylon band. The calibration of the pressure device was performed with a uniaxial material-testing machine
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Table 1 Characteristics of the investigated specimens. Age, sex, body weight, maximal and mean cartilage thickness, cartilage volume and articular surface of the patellae (determined from 3D reconstruction of the unloaded patellar cartilage) Specimen
Age (yr)
Sex
Body Weight (kg)
Maximal thickness Mean thickness (mm) (mm)
Cartilage volume (mm)
Articular surface area (cm)
1 2 3 4 5 6
31 67 18 51 65 23
m m m f f m
58 55 84 65 65 77
3.8 5.2 5.3 5.2 4.7 5.2
1.9 2.3 2.7 2.2 2.1 2.8
3680 3640 4620 2745 2560 5150
14.8 14.5 16.9 11.6 11.6 17.8
Mean SD
42.5 20
70 10
4.9 0.54
2.3 0.31
3730 930
14.5 2.35
Fig. 1. Compression apparatus for applying a load of 150% body weight to an intact femoro-patellar joint within the extremity coil of a clinical MRI scanner. (A) Apparatus in the extremity coil of the scanner, in front of the main magnet, (B) apparatus with a femoro-patellar specimen. Please note the pneumatic pressure device on top of the apparatus.
(Zwick 1445, Ulm, Germany). As a control, sagittal MR images were recorded before and at the end of load application, to assess whether the femur or the patella had tilted or shifted relative to the transverse image plane.
2.2. MR imaging Before load application, two sets of transverse images of the femoro-patellar joint were recorded with a
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previously validated fat-suppressed 3D gradient echo sequence (FLASH) at a resolution of 2;0.31;0.31 mm (FOV"16cm, matrix"512 pixels, TR"45 ms, TE"11 ms, FA"303, N. acq."1, imaging time"8 min). The mean values obtained from these data sets were used as a reference (for the volume and 3D thickness changes during compression), and as a basis for the selection of one transverse slice within the center of the contact area. This 2D slice was chosen to monitor the thickness changes with a fast 2D MR sequence during the "rst minutes of loading (N. acq."4, imaging time"70 s, imaging parameters otherwise identical to the 3D sequence described above) and with the 3D sequence up to 214 min after load application. Three transverse MR-images were recorded with this 2D sequence before compression. The patella was then loaded with 150% body weight (Table 1) within approx. 10 s, and six 2D images were recorded at intervals of 90 s up to 9 min after load application, the load being kept constant throughout the entire experiment. 3D data sets of the entire patellar cartilage (approx. 25 sections, including the selected central slice) were registered at 10 min intervals from 10 to 210 min after load application ("rst acquisition from 10 to 18 min). The middle of the acquisition time was used to characterize the time-dependant deformation of the cartilage (e.g. 14 min for the "rst 3D acquisition). In one specimen the compression experiment was extended by a recovery experiment, 3D data sets of the patellar cartilage being acquired from the moment of load release at 10 min intervals over 4 h.
#ux in lm/s) from the volume di!erences at various time steps. To be able to track the location- and time-dependant cartilage thickness changes, it was mandatory to reidentify corresponding points in the consecutive MR data sets. In the 2D slices, a scalar function was de"ned (representing the distance between corresponding cartilage}bone interfaces in consecutive images) and minimized within the parameter space of allowed patellar movements (translation in two axes, rotation around the center of mass * Herberhold et al. (1998), and Stammberger et al. (1998)). For evaluating the cartilage deformation throughout the joint surface, a new 3D matching algorithm was applied, which adjusts the corresponding bone}cartilage interfaces in consecutive 3D image data sets, and allows one to directly visualize the spatial di!erence between the thickness distribution patterns (Stammberger et al., unpublished work). 2.4. Joint pressure analysis with FUJI xlm Following imaging, the joints were left in an unloaded position for 8 h, opened, and two pieces of pressuresensitive "lm (FUJI low sensitivity: 2.5}10 MPa) positioned within the joint, one at the lateral and one at the medial patellar facet. The pressure device was then in#ated for 30 s to the same value as during the compression experiment. The "lms were "nally calibrated with a 0.5 cm stamp in a material-testing machine, applying a set of de"ned loads, and the staining of the "lm converted into pressure intervals (MPa) with image analysis.
2.3. Digital image analysis 3. Results Following imaging, all data sets were transferred digitally to a symmetrical multiprocessing computer (Octane Duo, Silicon Graphics Inc., Mountain View, CA, USA), the segmentation, 3D reconstruction, and digital image analysis being performed with the software designed in our laboratory (detailed descriptions in Herberhold et al., 1998; Eckstein et al., 1998a,b; Stammberger et al., 1998,1999b). Based on a Euclidean distance transformation algorithm (Stammberger et al., 1999a), these techniques make it possible to compute the total cartilage volume, the articular surface area, and the 3D mean and maximal cartilage thickness, independent of the original slice orientation. Whereas the 3D analyses were con"ned to the patella, the mean and maximal cartilage thickness within the selected 2D slice were determined for both the patellar and femoral cartilage layers. Based on the assumptions that the solid matrix is incompressible (Bachrach et al., 1998) (and therefore all volume changes are due to #uid #ow), and that during compression the #uid #ow will occur through the articular surface into the joint cavity, we estimated a #ow rate per square centimetre of the articular surface (the #uid
The maximal 3D cartilage thickness (before compression) in the six patellae ranged from 3.8 to 5.3 mm, and the mean thickness from 1.9 to 2.8 mm (Table 1). The cartilage volumes amounted to between 2560 and 5150 mm and the surface areas to between 11.6 and 17.8 cm (Table 1). During the compression experiment, the patella showed a shift of 3.8 ($1.1) mm in the direction of the applied force and of 0 ($0.4) mm perpendicular to it within the selected 2D image plane. The average rotation was !0.5 ($2)3. Movements perpendicular to the image plane did not occur, as demonstrated by comparison of sagittal images before and after the experiment. The analysis of the cartilage deformation in the selected central 2D slice showed a mean reduction of 44$15% of the initial thickness in the patella, and of 30$10% in the femur after 214 min of static loading with 150% body weight (Fig. 2A, Table 2), the deformation of the patella generally exceeding that of the femur. The maximal deformation was 57$15% in the patella and 44$7 in the femur (Fig. 2B, Table 2). The cartilage
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thickness decreased in an approximately exponential manner (Figs. 2A and B). Only 7$10% of the "nal (214 min) deformation was reached during the "rst minute of static loading of the
Fig. 2. In situ thickness changes during loading with 150% body weight. The error bars show one standard deviation of the variability between individuals. (A) Mean cartilage deformation in the selected central 2D slice within the contact area, (B) Maximal cartilage deformation in the selected central 2D slice. A mean (maximal) in situ deformation of 44% (57%) was observed in the patellar cartilage after 3 h of static loading (mean contact pressure"3.6 MPa). The femoral cartilage generally showed a smaller amount of in situ deformation than the patella.
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patellar cartilage (3% absolute deformation) and 25$10% after approx. 8 min (11% absolute deformation), the values for the femur being 4$12% and 30$12% (absolute deformation 1.3 and 9%), respectively (Fig. 3). The patellar cartilage volume was reduced by 8$5% after 14 min and by 29$3.2% after 214 min of loading (Fig. 4A, Table 2). The mean 3D patellar thickness decreased by 10$4% (14 min) and 29$5% (214 min), respectively, the time-dependant 3D thickness changes being very similar to those of the volume (Fig. 4A). The maximal and mean deformation in the selected 2D slice showed a high linear correlation with the volume changes observed in the 3D data (r"0.96 and 0.82, respectively). The rate of interstitial #uid loss from the matrix was 1.3 ($0.5) mm/min per square centimeter surface area (#uid #ux"0.217$0.083 lm/s] for the "rst 14 min of loading, and 0.22 ($0.04) mm/min cm (0.037$0.007 lm/s) in the terminal phase of the experiment ('120 min). There was a high interindividual variability of the deformational behavior, both in the 2D analysis (mean and maximal thickness) and in the 3D analysis (volume change, #uid #ux), the coe$cients of variation ranging from 11% (volume) to 38% (#ux) (Table 2). The thickness changes were highly inhomogeneous throughout the joint surface, the site of maximal deformation (the lateral patellar facet) being identical with that of the maximal cartilage thickness (Fig. 5). The FUJI "lm experiment conducted after imaging showed pressure maxima of 3.6$1.25 MPa, in the lateral patellar facet (Table 2), the pressure distribution being very similar to the pattern of cartilage deformation (Fig. 5). The one specimen used for the recovery experiment displayed almost full recovery (98%) of the patellar cartilage after approx. 4 h, the rate of #uid re-uptake by the cartilage matrix being 0.9 mm/min cm (0.15 lm/s) within the "rst 14 min of recovery, and 0.14 mm/min cm (0.023 lm/s) in the terminal phase.
Table 2 Deformational behavior of the di!erent specimens during loading with 150% body weight (2D and 3D analysis, patellar cartilage, maximal (mean) deformation and volume changes after 3.5 h of static loading) (1 mm/min cm is equivalent to 0.167 lm/s) Specimen
Mean deformation 2D slice (%)
Maximal deformation 2D slice (%)
Volume change Peak initial #ux (%) (mm/min cm))
1 2 3 4 5 6
30 27 34 70 52 48
60 41 40 84 65 53
29 26 25 33 32 28
Mean SD CV%
44 15 34
57 15 26
29 3.2 11
1.4 0.5 0.75 2.1 1.4 1.4 1.3 0.5 38
Terminal #ux (mm/min cm))
Articular pressure (MPa)
0.25 0.18 0.25 0.21 0.16 0.24
3.0 1.5 3.75 5.25 5.25 3.0
0.22 0.04 18
3.6 1.25 35
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Fig. 3. Cartilage deformation in the "rst minutes of loading (150% body weight) measured in the selected central 2D slice within the contact area. The error bars show one standard deviation of the variability between individuals. Only approx. 7% of the "nal deformation of the patellar cartilage (3% absolute deformation) occurred after the "rst minute, and 25% of the "nal deformation (11% absolute deformation) after 8 min of compression. The values for the femur were 4% (1.3% absolute deformation) and 30% (9% absolute deformation), respectively.
4. Discussion The objective of the current study was to investigate the in situ deformation of articular cartilage in the intact femoro-patellar joint under static loading with MR imaging. To our knowledge, this is the "rst systematic study to report the in situ deformation and volume changes of articular cartilage in an intact joint, under normal, nonlinear contact conditions and with the joint capsule being intact. Because the compressive behavior of the cartilage over the "rst few minutes of loading is physiologically most relevant, we analyzed the initial deformation with 2D MR sequences. These allow faster acquisition than 3D protocols (imaging time around 1 min) and have been shown to produce results identical to those of previously validated 3D fat-suppressed gradient echo sequences (Herberhold et al., 1998). To exclude movement relative to the image plane (which would interfere with accurate 2D measurement) the patellar motion was e!ectively restricted by the apparatus to the transverse image plane. The results obtained for the selected central 2D slice are shown to be consistent with those obtained for the entire patella (3D data) after long-term loading, and thus appear representative for comparative measurements. We "nd a mean (maximal) in situ deformation of 44% (57%) in the patellar cartilage after 3 hrs of static loading with 150% body weight (mean contact pressure"3.6 MPa). Only approximately 7% of the "nal deformation (3% absolute deformation) occurred during the "rst minute, and 25% (11% absolute deformation) during 8 min of loading. The data show that, under in
Fig. 4. In situ volume changes and #uid #ux during loading with 150% body weight. The error bars show one standard deviation of the variability between individuals. (A) Volume change of the patellar cartilage (3D analysis), (B) rate of #uid loss from the interstitial patellar cartilage matrix per square centimeter surface area (and #uid #ux in lm/s), based on the assumption that during compression the #uid #ow will occur through the articular surface into the joint cavity. A 43% #uid loss was observed from the interstitial matrix after 3 h of static loading, the #uid #ow rate per square centimeter articular surface being initially 1.3 ($0.5) mm/min cm (0.217$0.083 lm/s) for the "rst 14 min after loading, and 0.22 ($0.04) mm/min cm (0.037$ 0.007 lm/s) in the terminal phase of the experiment ('120 min).
situ conditions, only a small fraction of the "nal deformation (steady state) is reached for physiologically relevant loading periods, which typically last several seconds or minutes. These results are consistent with previous theoretical analyses (Mak et al., 1987; Ateshian et al., 1994), showing that there is little time for #uid to escape initially and therefore little change in the cartilage volume, due to the low permeability of the cartilage. These "ndings support the idea that the initial response of articular cartilage is nearly incompressible elastic, and that the #uid supports a substantial portion of the applied load
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Fig. 5. Cartilage thickness changes throughout the patellar cartilage surface. 3D matching of corresponding bone}cartilage interfaces in the consecutive 3D image data sets. (A) Cartilage thickness distribution before loading (The 8 gray values indicate thickness intervals of 0.6 mm (black"0}0.6 mm; white"'4.2 mm), (B) cartilage thickness distribution after loading [same gray value legend as for (A)], (C) direct visualization of thickness di!erences after 3D matching (dark"areas without deformation; white"areas of maximal deformation (gray values indicate intervals of 0.4 mm). There is a highly inhomogeneous deformation throughout the joint surface, the maximal deformation (2.4 mm) occurring in the lateral patellar facet, at the site of maximal cartilage thickness [see (A)].
(Ateshian et al., 1994; Ateshian and Wang, 1995), the apparent sti!ness of the tissue being much greater in the instantaneous than in the equilibrium response. The mechanism of hydrostatic pressurization of the interstitial #uid, which is assumed to protect the solid matrix from elastic deformation for several hundred seconds after the onset of loading (Ateshian et al., 1994; Ateshian and Wang, 1995) is assumed to be a crucial factor in protecting the matrix from undue stress and tissue damage during normal loading. The femoral cartilage generally showed a smaller amount of in situ deformation than the patella, this being consistent with the results of Froimson et al. (1997) who found a lower sti!ness and higher permeability of the patellar cartilage in relation to the femur in indentation studies. This demonstrates that the technique presented is able to e!ectively reveal di!erences in the mechanical behavior (and material properties) of di!erent cartilage plates. The observed di!erences may also be relevant in explaining the empirical observation that the patella generally shows earlier and more severe signs of osteoarthritic changes than the femur (Ficat and Hungerford,
1977; Froimson et al., 1997). Considering that the cartilage consists of 70% #uid (Mow and Ratcli!e, 1997), a 43% #uid loss from the interstitial matrix was observed in this study after 3 h of static loading. Assuming that during compression the #uid #ow occurs through the articular surface into the joint cavity, the #uid #ux was 0.216 lm/s initially, and 0.037 lm/s at the terminal phase of the experiment. The rate of #uid re-uptake observed during recovery (0.023 lm/s) in the terminal phase * one specimen) is in the same order of magnitude as that observed in an in vivo recovery experiment after 100 knee bends (0.027 lm/s) after a 5% volume change (Eckstein et al., 1999). These results suggest that the data obtained with the compression apparatus can, within some limits, be transferred to the in vivo situation. The #uid #ow through the joint surface can apparently be increased by a factor of approximately 10 over that during free recovery, when statically loading the cartilage with 150% body weight and with peak pressures of around 3.6 MPa. The high inter-individual variability observed in this study is also in agreement with previous indentation
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experiments (Athanasiou et al., 1991; Froimson et al., 1997) and in vivo results (Eckstein et al., 1998c,1999), both the deformation and the recovery of the cartilage yielding considerable di!erences between individuals. It is up to speculation whether these observed di!erences relate to the observation that some individuals show early signs of cartilage damage, whereas in others the cartilage maintains normal morphology and function over several decades (Mow and Ratcli!e, 1997). The data presented here may serve as a basis for "nite-element simulations, to validate the non-linear contact algorithms developed for biphasic cartilage layers (Donzelli and Spilker, 1998). By matching the experimental and computational results, it should become feasible to analyze the load partitioning between the solid (proteoglycan}collagen) matrix and #uid phase, to determine the #uid #ux within the cartilage matrix, and to elucidate etiologic factors relevant in mechanically induced osteoarthritis. The data can also provide a direct estimate of the mechanical strain to be encountered by autologous or tissue-engineered cartilage transplants, once they are located in their target environment.
Acknowledgements Parts of this study were performed within the framework of the doctoral thesis (in preparation) of Christoph Herberhold at the Medical Faculty of the LudwigsMaximilians-UniversitaK t, MuK nchen, Germany. We would like to thank Professor Gerard Ateshian, Columbia University (New York) for his very helpful discussion and the Deutsche Forschungsgemeinschaft (DFG) for supporting this work.
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