Patellar cartilage deformation in vivo after static versus dynamic loading

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Abstract

The objective of this study was to test the hypothesis that static loading (squatting at a 90° angle) and dynamic loading (30 deep knee bends) cause different extents and patterns of patellar cartilage deformation in vivo. The two activities were selected because they imply different types of joint loading and reflect a realistic and appropriate range of strenuous activity. Twelve healthy volunteers were examined and the volume and thickness of the patellar cartilage determined before and from 90 to 320 s after loading, using a water excitation gradient echo MR sequence and a three-dimensional (3D) distance transformation algorithm. Following knee bends, we observed a residual reduction of the patellar cartilage volume (−5.9±2.1%; p<0.01) and of the maximal cartilage thickness (−2.8±2.6%), the maximal deformation occurring in the superior lateral and the medial patellar facet. Following squatting, the change of patellar cartilage volume was −4.7±1.6% (p<0.01) and that of the maximal cartilage thickness −4.9±1.4% (p<0.01), the maximal deformation being recorded in the central aspect of the lateral patellar facet. The volume changes were significantly lower after squatting than after knee bends (p<0.05), but the maximal thickness changes higher (p<0.05). The results obtained in this study can serve to validate computer models of joint load transfer, to guide experiments on the mechanical regulation of chondrocyte biosynthesis, and to estimate the magnitude of deformation to be encountered by tissue-engineered cartilage within its target environment.

Introduction

Quantitative knowledge on the in vivo deformation of articular cartilage during or after normal joint loading is crucial in the context of several issues. Such data can, for instance, provide an estimate of the magnitude of deformation and strain to be encountered by potentially vulnerable, tissue-engineered articular cartilage once transplanted to a functioning joint. Secondly, accurate data on in vivo cartilage deformation are required for validating computer models that attempt to derive the interstitial fluid pressure and solid matrix stress in the cartilage (Mow et al., 1993; Ateshian et al., 1991; Suh et al., 1995; Goldsmith et al., 1996; Wu et al (1996), Wu et al (1997), Wu et al (1998); Mow and Ratcliffe, 1997; Donzelli and Spilker, 1998), particularly when these studies will be extended to the simulation of physiological loading patterns. Because it is the stress of the matrix, and not the interstitial hydrostatic pressure that is relevant in cartilage damage (Ateshian et al., 1991), these models can provide insight into the pathogenesis of cartilage injury caused by mechanical overuse. Finally, information on the normal deformation can guide in vitro experiments on the biosynthetic behavior of articular cartilage and thus allow to develop strategies for enhancing cartilage metabolism and repair by adequate physical stimulation. In these in vitro experiments, chondrocytes have been shown to yield a differential response to dynamic and static loading regimes as well as to different strain frequencies (Gray et al., 1988; Sah et al., 1989; Urban, 1994; Kim et al., 1995; Bachrach et al., 1995; Buschmann et al., 1996; Lee et al., 1998). Although there is a high interest in exploring mechano-transduction pathways enabling chondrocytes to sense mechanical signals and to process them into a biological response (e.g. Guilak et al., 1995; Guilak, 1995; Buschmann et al., 1996), currently there is a lack of data with regard to the actual magnitude of deformation and strain in articular cartilage in its in vivo environment.

Traditionally, the deformational behavior of articular cartilage has been examined in compression experiments of excised tissue samples, by indentation devices (Mow et al., 1984; Athanasiou et al., 1991; Goldsmith et al., 1996; Mow and Ratcliffe, 1997) and, in few instances, also in intact cadaver joints (Armstrong et al., 1979; Wayne et al., 1998; Herberhold et al (1998), Herberhold et al (1999)). However, the cartilage deformation occurring under in vivo conditions cannot be extrapolated from in vitro studies, because the magnitude and rate of the muscle and joint forces, which are exerted during various physiological activities, are unknown. The contact- and load-bearing areas will vary with changes in joint position, and due to the natural incongruity of the joint surfaces (e.g. Eckstein et al., 1994) they also depend on the particular load magnitude. At the joint surface, highly complex pressure gradients occur under such physiological loading conditions, which cannot be adequately simulated in an experiment.

In this context, dynamic and static loading regimens can be expected to yield different effects: During knee bends, for instance, the contact- and load-bearing areas travel rapidly across the entire patellar surface (Hehne, 1990), and thus a relatively homogeneous deformation pattern may be expected. Upon static loading at a 90° knee angle, on the other hand, the load transmission is confined to one relatively small area in the central aspect of the lateral patellar facet, the deformation potentially reaching higher values there, but smaller ones in the other parts of the joint surface. Moreover, the type of loading is different under these two conditions: During knee bends, the rate of loading is high, but its duration short, the tissue being able to recover in between repetitive loading cycles. During squatting, however, the loading rate is much lower, but the cartilage is compressed continuously at one particular site. Kääb et al. (1998) have been able to demonstrate different effects of static and dynamic loadings on the collagen structure of articular cartilage in a rabbit model. We believe that knee bends (dynamic) and squatting (static) reflect a realistic and appropriate range of joint loading, since most other strenuous activities will represent a combination of theses two loading conditions.

The objective of the present study was thus to investigate changes in cartilage volume and thickness immediately after static (squatting) and dynamic (knee bends) loading in an adaptable in vivo system, this information being of particular relevance in the context of the transplantation of tissue-engineered cartilage into its loaded target environment. The experiment proposed takes into account the unique, individual neuromuscular control functions, which cannot be simulated under in vitro conditions. Based upon the reasons given above, we hypothesized that static and dynamic loadings will cause different extents and patterns of patellar cartilage deformation in squatting causing a higher residual deformation at one particular site, and knee bends causing a more homogeneous deformation throughout the patellar joint surface.

Section snippets

Materials and methods

In 12 healthy volunteers without musculoskeletal disease or internal derangements of the knee (age 22–30 yr (mean 25.3 yr), 6 males, 6 females), MR imaging was performed with a commercial 1.5 T MR magnet (Magnetom Vision, Siemens, Erlangen, Germany) and a circularly polarized knee coil. We obtained serial, contiguous, transverse images of the right patellar cartilage. In contrast to previous studies (Tieschky et al., 1997; Eckstein et al., 1998a), we employed a spoiled three-dimensional (3D)

Results

After 30 knee bends, we observed a significant reduction of the patellar cartilage volume from an average of 3.67±0.60 to 3.46±0.61 ml, the mean pairwise differences amounting to −5.9 (±2.1)% (p<0.01) (Fig. 2), and the minimal and maximal changes being −2.9 and −9.6%, respectively (Fig. 3). The reduction of the maximal cartilage thickness was from an average of 5.52±0.99 to 5.36±0.95 mm, the mean difference being −2.8 (±2.6)% (not significant, minimum=+0.8%, maximum=−7.4%) (Fig. 2). The 3D

Discussion

Whereas reasonable estimates exist regarding the strains on the surfaces of bones during normal daily activity (Duncan and Turner, 1995), little is known about the deformation to be expected in articular cartilage during and after physiological loading. Bone tissue atrophies when not adequately stimulated, it is maintained when experiencing strains within a certain range (physiologic window: approx. 100–2000 μstrain; Duncan and Turner, 1995), it hypertrophies when critical threshold values are

Acknowledgements

We would like to thank the Deutsche Forschungsgemeinschaft for its support. These data presented in this study are part of the doctoral thesis (in preparation) of Bärbel Lemberger at the Ludwig-Maximilians-Universität München.

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