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Influence of stress rate on water loss, matrix deformation and chondrocyte viability in impacted articular cartilage
ARTICLE IN PRESS
Journal of Biomechanics 38 (2005) 493–502
Influence of stress rate on water loss, matrix deformation and chondrocyte viability in impacted articular cartilage Dejan Milentijevic*, Peter A. Torzilli Laboratory for Soft Tissue Research, Hospital for Special Surgery and Center for Biomedical Engineering, City University of New York, New York, NY 10021, USA Accepted 21 April 2004
Abstract The biomechanical response of articular cartilage to a wide range of impact loading rates was investigated for stress magnitudes that exist during joint trauma. Viable, intact bovine cartilage explants were impacted in confined compression with stress rates of 25, 50, 130 and 1000 MPa/s and stress magnitudes of 10, 20, 30 and 40 MPa. Water loss, cell viability, dynamic impact modulus (DIM) and matrix deformation were measured. Under all loading conditions the water loss was small (B15%); water loss increased linearly with increasing peak stress and decreased exponentially with increasing stress rate. Cell death was localized within the superficial zone (p 12% of total tissue thickness); the depth of cell death from the articular surface increased with peak stress and decreased with increasing stress rate. The DIM increased (200–700 MPa) and matrix deformation decreased with increasing stress rate. Initial water and proteoglycan (PG) content had a weak, yet significant influence on water loss, cell death and DIM. However, the significance of the inhomogeneous structure and composition of the cartilage matrix was accentuated when explants impacted on the deep zone had less water loss and matrix deformation, higher DIM, and no cell death compared to explants impacted on the articular surface. The mechano-biological response of articular cartilage depended on magnitude and rate of impact loading. r 2004 Elsevier Ltd. All rights reserved. Keywords: Cartilage; Impact; Stress rate; Stress magnitude; Water loss; Cell viability
1. Introduction Cartilage damage can occur from minor injuries such as in a sudden twist of an ankle (sport injury) to severe trauma such as in car accidents. During these different traumatic episodes the articular cartilage may experience different load and strain magnitudes as well as different load and strain rates. Articular cartilage resists joint loads by two distinctly different mechanisms that can be delineated by the rate (stress or strain) at which loading is applied. Under slow loading rates, cartilage behaves as a poroelastic material exhibiting fluid-flow dependent deformation, where the load is transiently resisted by the frictional drag between the fluid and the solid matrix while at equilibrium the load is resisted by *Corresponding author. Tel.: +1-212-774-2003; fax: +1-212-2492373. E-mail address: [email protected] (D. Milentijevic). 0021-9290/$ - see front matter r 2004 Elsevier Ltd. All rights reserved. doi:10.1016/j.jbiomech.2004.04.016
the solid phase only. On the other hand, under rapid and impact loading conditions cartilage exhibits fluid-flow independent deformation. Theoretical models predict that under impact loading conditions cartilage will respond as an incompressible, elastic material (fluid and solid assumed incompressible), where the interstitial fluid is incapable of escaping the dense matrix during the short time of loading (Armstrong et al., 1984; Mak et al., 1987; Ateshian and Wang, 1995; Soltz and Ateshian, 1998). How these different mechanical mechanisms effect the extent of damage to the cartilage and its post-injury survival is not well understood. Several studies have evaluated the affects of strain rate by loading cartilage explants in unconfined compression. Radin et al. (1970) loaded bovine cartilage (without underlying bone) at slow (0.1–16%/s) and fast (up to 670%/s) strain rates, and found that the compressive stiffness increased from 20 to 50 MPa with increasing strain rate (0.5–35%/s). Repo and Finlay
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(1977) impacted human articular cartilage-on-bone and reported that while matrix damage (fibrillation and cell death) was initiated at 20–30% strain (B25 MPa) it was independent of strain rate (500 vs. 1000 s1). They also attempted to quantify the amount of water loss but were unsuccessful because of large weighing errors. Later Oloyede et al. (1992) used a conventional materials testing machine to apple slow strain rates (5 105– 5 102 s1) and a free-swinging pendulum for impact strain rates (1000 s1). The compressive stiffness of the cartilage (modulus) increased with strain rate, however the modulus reached a maximum of B40 MPa for strain rates X5 102 s1. More recent studies have focused on the effects of strain rate ranging well below impact level. Quinn et al. (2001) loaded osteochondral specimens from young adult steer in unconfined compression at strain rates from 3 105 to 0.7 s1. At the slowest strain rate there was excessive chondrocyte death within the tissue (B100% loss of viability) with no visible matrix damage, while at the faster rates (0.3–0.7 s1) cell death occurred only in the superficial zone and adjacent to surface fissures. However, in another study using newborn calf cartilage, increasing strain rate (0.01–1 s1) resulted in a decrease in cell viability 3-days after compression (Kurz et al., 2001). Few studies have investigated how the impact loading (stress) rate effects cartilage damage. Ewers et al. (2001) impacted bovine cartilage samples (without underlying bone) in unconfined compression with a 40 MPa stress at two stress rates, 40 and 900 MPa/s, resulting in average strains of 41% and 48%, respectively. The impact caused superficial matrix damage (fissures), with cell death only close to the fissures at the high loading rate which propagated away from the fissures deeper into the matrix at the low loading rate. In a previous study, we impacted bovine cartilage in confined compression with peak stresses of 10–60 MPa at a single stress rate of 350 MPa/s (Milentijevic and Torzilli, 2003). Cell death was localized within the superficial region and increased linearly in depth from the surface with increasing peak stress. We also measured the water loss from the matrix immediately after impaction, which was small and also increased with increasing peak stress (1–10%). In previous studies, articular cartilage was loaded quasi-statically (slow rate) or with a single impact (fast rate). For the latter, drop-tower or pendulum type devices were predominantly used to generate high impact energy (fast rates). However, the loading rate and load magnitude are difficult to control independently, and thus it is difficult to assess their individual effects on cartilage response. To our knowledge, no study has evaluated how cartilage responds under a broad range of impact loading rates, nor have measurements been made of the resulting matrix compression
and concomitant water loss. Therefore, the objective of our study was to investigate the mechano-biological response of articular cartilage to varying loading rates and load magnitudes. We hypothesized that with increasing loading rate the cartilage matrix would respond mechanically by decreasing its matrix deformation and water loss, and thus reduce the amount and depth of cell death. To test this hypothesis, we impacted viable articular cartilage in confined compression with peak stresses of 10–40 MPa at stress rates of 25, 50, 130 and 1000 MPa/s, and measured water loss, cell viability, matrix deformation, and dynamic impact modulus (DIM). We found that cell death was localized within the superficial zone (p12% of total tissue thickness), and that the DIM increased (200–700 MPa) and matrix deformation and water loss decreased with increasing stress rate.
2. Materials and methods 2.1. Tissue acquisition Mature bovine knees with an intact capsule (young adult 18–24 months of age) were obtained from a local abattoir (Greenville, NJ) within twenty-four hours after death. A biopsy punch (diameter=8 mm) was used to cut circular explants from the femoral condyles and trochlear groove. Full-thickness disks were removed from the subchondral bone using a surgical blade, washed three times (15 min each) in Dulbecco’s Modified Eagle’s Media (DMEM, Gibco BRL, CA) containing 10% antibiotic/antimycotic (Ab/Am). The explants were then placed in DMEM with 10% Hepes buffer solution, 10% fetal bovine serum and 1% Ab/Am at 37 C and 95% humidity and incubated for least 24 h to allow the explants to reach an equilibrium water content and stabilize cell metabolism, as previously determined (Torzilli et al., 1997). Equilibrium matrix swelling was confirmed by independent measurements of water content after incubation for 20 min and 24 h (paired difference=0.170.4%, p>0.2, n=24). Initial water and proteoglycan (PG) contents and cell viability at the time of impact testing were determined as described below. 2.2. Impact test system A detailed description of the Cartilage Impact Test System (CITS) was reported earlier (Milentijevic and Torzilli, 2003). Briefly, the CITS uses a servo-controlled impactor (double-acting pneumatic cylinder) with a solid flat-ended cylindrical platen (6.35 mm) to apply a peak stress at a defined stress rate to the cartilage explant. A piezoelectric load transducer is used to control the peak stress while a flow control valve is used to set the loading rate. A high-speed, Linear Variable
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Differential Transformer (LVDT) is used to measure matrix deformation. At the time of testing, a cartilage disk (6.35 mm diameter) is impacted in confined compression by inserting it into a confining chamber (6.35 mm diameter) with the articular surface facing down onto a stainless-steel porous filter (39 mm pore size) (Fig. 1b). A data acquisition system is used to record the applied load, matrix deformation and time at a 5 kHz sampling rate per channel. The test system was calibrated for repeatability (precision) at each loading rate (stress rate, MPa/s) used in this study: 25 MPa/s (mean7SD=25.017 0.90 MPa/s, coefficient of variance (COV)=3.6%, n=5), 50 MPa/s (49.9670.83 MPa/s, COV=1.7%, n=5), 130 MPa/s (130.7871.99 MPa/s, COV=1.5%, n=5) and 1000 MPa/s (1003.9778.60 MPa/s, COV= 0.86%, n=5). In addition, the applied peak stresses (load/indenter’s cross-sectional area) of 10, 20, 30 and 40 MPa were also found to be repeatable at each stress rate (e.g., at 130 MPa/s and 30 MPa, peak stress= 30.0870.13 MPa, COV=0.44%, n=5). However, for the fastest stress rate of 1000 MPa/s the CITS was only able to apply peak stresses of 17, 23, 30 and 40 MPa due to the transient response of the test system. 2.3. Experimental design Impact tests were performed 1-to-3 days after cartilage harvest (post-stabilization). On the day of testing, cartilage explants were cored with a biopsy punch (d=6.35 mm) and the core (center) and remainder of the explant (ring) were returned to the media to stabilize (swell) for an additional 1 h (Fig. 1a). Independent tests were performed to confirm that the explants had reached a constant water content by measuring the wet weights of the cores and rings immediately after removal and 30 min and 1 h later (paired differences p0.1%, p>0.1, n=24). As described below, the cores
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were used for impact testing while the outer rings were used to determine the initial water and PG content of the core. Wet weights of the center (WWC, mg) and the ring (WWR, mg) were determined as follows. First the excess water was removed from the explant by gently blotting the specimen’s surfaces on a Kim Wipe (Kimberly Clark Corporation, Rowel, GA) and immediately weighing the explant using a micro-balance (resolution 71 mg, Cahn, Cerritos, CA). The wet weight was recorded exactly 15 s after placement on the balance to normalize for water evaporation and the explant returned to the fluid media. The center’s thickness was then measured using the LVDT by placing the explant in the CITS and manually lowering the load platen until a B60 kPa pressure was applied to the explant, at which point the thickness was recorded. After thickness measurement the center was returned into the fluid media. Immediately before impact loading, the center was blotted as described above and placed into the confining chamber with the articular surface resting on the dry porous filter. The center was then loaded with a stress rate of 25, 50, 130 or 1000 MPa/s and a peak stress of 10, 20 or 40 MPa. Eight specimens were impacted at each combination of stress rate and peak stress. Immediately after the impact, the center was quickly removed from the chamber, blotted, the wet weight (WWI, mg) measured, and the explant returned to the fluid media. After one hour of incubation the center was again blotted and weighed (WWS, mg). Control explants were treated identically but were not impacted. After determining the wet weights, the centers were evaluated for chondrocyte viability using the FDA (fluorescein diacetate, Molecular Probe, OR)/PI (propidium iodide, Sigma Chemicals, MO) live/dead cell essay. Thin cross-sections (B0.5 mm thick) were cut from each explant disk, incubated with FDA (10 mM)-PI (60 mM) stain for 5 min, rinsed in PBS (phosphate-buffered
Fig. 1. Schematic illustration of the impact test configurations used in the study. (a) The center of the cartilage explant (core, 6.35 mm diameter) was removed and impacted under different test configurations. The remainder of the explant (ring, 8.0 mm diameter) was used to determine the initial water and PG contents. (b) Articular surface loading (ASL): The core is positioned in the chamber with its articular surface facing the porous filter and the deep zone impacted with a nonporous load platen. The filter was placed on the bottom of the chamber to retain exuded water (due to gravity) when the explant was unloaded. (c) Deep zone loading (DZL): The cut-surface (deep zone) is placed against the porous filter and the articular surface impacted with the non-porous load platen. The ASL explants were impacted with 10, 20, 30 and 40 MPa at stress rates of 25, 50, 130 and 1000 MPa/s (n=9 explants per loading combination), while the DZL explants were impacted with a peak stress of 30 MPa at all stress rates (n=10 explants per loading combination).
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saline) and viewed using a fluorescent microscope (Nikon, NY). Digital images were recorded using an image capturing system (CCD camera, Optronics, Goleta, CA) attached to the microscope, and imaging software (SigmaScan, SPSS Science, Chicago, IL) was used to quantify the amount of chondrocyte death. Since cell death was initiated at the articular surface in all explants, the depth of cell death was calculated from the average of 20 measurements taken from the articular surface along the cross-section. The absolute depth (mm) of cell death from the surface was then calculated as the mean of two cross-sections per explant, and percent (%) depth calculated by normalizing this measurement to the initial cartilage thickness. The remaining rings were used to calculate the water and PG content of their paired cores as described below. The rings were freeze-dried overnight in a lyophilizer (Labconco, Kansas, MO) and their dry weights (DWR) determined. The rings were then digested in 1 ml Papain solution (Sigma-Aldrich Corp., St. Louis, MO) on a shaker water bath at 65C. Digestates were used to determine the PG content of each ring using the dimethylmethylene blue (DMMB) assay (Farndale et al., 1986). An independent impact test was performed to compare the amount of water lost from the deep zone (DZ) to that lost from articular surface (AS). Specimens were impacted with their cut surface (n=10) against the porous filter (Deep Zone Loading, DZL) (Fig. 1c) and with their AS (n=10) against the porous filter (Articular Surface Loading, ASL) (Fig. 1b), with a 30 MPa peak stress at a stress rate of 25, 50, 130 or 1000 MPa/s. Specimens not impacted served as control for each test configuration (n=5). 2.4. Initial water and pg content and water loss Preliminary experiments were performed to determine if there was a difference in initial water and PG contents between the center and outer ring from the same explant (Fig. 1a). Briefly, twelve explants (d=8 mm) were harvested from the bovine knee joint, cored (d=6.35 mm), and the paired wet weights, dry weights and PG content of the center (WWC, DWC and PGC) and ring (WWR, DWR and PGR) determined as described above. The water content of the center (WCC) and ring (WCR) were not statistically different (77.971.9% and 78.271.9%, respectively, p=0.7). A similar result was found for the PG content normalized per wet and dry weights (5.070.9% and 4.970.8%, p=0.89, and 23.775.9% and 24.374.5%, p=0.52, respectively). Since the water and PG contents of the center did not differ from the ring, we could calculate the initial water content in the center from the ring, and therefore utilize the center for the cell viability analysis. In the same fashion the PG content of the ring was used
to determine the initial PG content of the center. Finally, the change in water content (water loss) immediately after impact was calculated from Water Loss ¼ ðWWI WWC Þ=ðWWC WCR Þ;
ð1Þ
while the change in water content (water absorption) 1-h post impaction was calculated from Water Absorption ¼ ðWWS WWC Þ=ðWWC WCR Þ: ð2Þ 2.5. Data analysis The deformation of each explant was first corrected for apparatus compliance and individual stress–strain curves plotted to determine the maximum strain. Data graphing and statistical analysis were performed using commercially available software (SigmaPlot and SysStat,, SPSS Inc.). The DIM was calculated from the slope of the linear region of the loading phase for stresses X15 MPa. Linear regression analysis (with 795% confidence intervals) was performed for water loss and depth of cell death (propagation from the surface) as a function of peak stress, stress rate and corresponding maximum strains. A multiple regression analysis (forward, backwards, stepwise and all-possible procedures) was used to assess the effect of the independent parameters (peak stress, stress rate, PG content and water content) on water loss, depth of cell death, matrix deformation (strain) and DIM (Draper and Smith, 1966). The Student’s t-test (unpaired or paired, two-tailed, post-hoc for repeated measures) was used to test for differences between the dependent variables, while R2, t and F statistics were used to determine the best regression models. Statistical significance was set at a=0.05. Averaged data is given as mean7standard deviation unless otherwise indicated. For completeness, data from a previous study (Milentijevic and Torzilli, 2003) performed at a stress rate of 350 MPa/s was included in the analyses and figures.
3. Results The initial thickness and water and PG contents of the impacted explants was 1.2670.22 mm, 0.7770.02 mg/ mg (wet weight) and 0.2970.08 mg/mg (dry weight), respectively (n=105); for the non-impacted controls these values were 1.2570.18 mm, 0.7870.01 mg/mg and 0.2870.07 mg/mg (dry weight) (n=36). There was no significant difference in initial thickness and water and PG contents between impacted and non-impacted groups (p>0.7). The stress–strain response of the impacted samples exhibited a non-linear (toe) region (0–5 MPa) followed by a transition to a more linear region (>10 MPa)
ARTICLE IN PRESS D. Milentijevic, P.A. Torzilli / Journal of Biomechanics 38 (2005) 493–502 0.18
35 25 MPa/sec
1000 MPa/sec
40 MPa
30
0.15
Water Loss, mg/mg
Peak Stress, MPa
25
497
Peak Stress = 30 MPa
20 15 10 5
0.12
0.09
30 Mean + SEM
20 30
0.06
DZL 10
0.03
0
0.00 25 0.00
0.05
0.10
0.15
0.20
50
130
350
1000
Stress Rate, MPa/sec
(a)
Strain, mm/mm
(Fig. 2). The toe regions appeared nearly identical when the stress–strain responses were superimposed for all loading combinations, while the slopes of the linear regions (DIM) increased with increasing stress rate. The hysteresis area between the loading and unloading phases also decreased with increasing stress rate. Immediately after load removal (zero stress), the stress–strain response indicated a residual strain in the tissue (Fig. 2). The amount of residual strain increased with increasing peak stress but decreased with increasing stress rate (25–350 MPa/s). Explants impacted at the highest stress rate of 1000 MPa/s had minimal residual strain upon load removal, which also varied minimally with peak stress. There was no change in the water content of the nonimpacted explants (controls) during the time it took to perform the impact tests, that is, between weighing pre- and post-impact (water loss=0.0070.01 mg/mg). On the other hand, the water loss in impacted explants increased linearly with increasing peak stress and corresponding maximum strains (Fig. 3b), while decreasing non-linearly with increasing stress rate (Fig. 3a). Multiple linear regression analysis determined that the best statistical combination of the independent variables (predictors) that were associated with changes in water loss (WL) were the peak stress (s), stress rate ðsÞ ’ and initial water (WC) and PG contents (R2=0.768, po0.001), WLðmg=mgÞ ¼ 0:089ð70:922Þ þ 0:002ð70:001Þs 0:060ð70:139Þðln sÞ ’ þ 0:004ð70:011Þðln sÞ ’ 2 þ 0:044ð70:257ÞPG þ 0:335ð71:169ÞWC:
ð3Þ
0.18 25 MPa/s
0.15
Water Loss, mg/mg
Fig. 2. Typical stress–strain response of explants impacted with a peak stress of 30 MPa at the slowest (25 MPa/s) and highest (1000 MPa/s) stress rates. All impacted explants exhibited a non-linear toe-region followed by a linear region. The slope of the linear region was used to determine the DIM. The area between the loading and unloading phases (hysteresis region) decreased with increasing stress rate.
95% CI
0.12
50 130
0.09
0.06 350
0.03
0.00 0.00 0.05
(b)
1000
0.10
0.15
0.20
0.25
Max. Strain, mm/mm
Fig. 3. Water loss from the cartilage matrix as a function of (a) impact stress rate (logarithmic scale) and (b) maximum strain. The water loss (mg) has been normalized to the total amount of water (mg) in the explant. (a) Water loss from the articular surface (ASL; solid lines) decreased exponentially with increasing stress rate. The DZL explants (dashed line), impacted with a 30 MPa peak stress, had significantly less water loss from the deep zone compared to the water loss in the ASL explants impacted with the same peak stress. (b) Water loss increased linearly with maximum strain and decreased with increasing stress rate. The 95% confidence intervals (dotted lines) are plotted only for the highest and lowest stress rates.
However, the stress and stress rate accounted for 68% of the WL, with the initial water and PG contents accounting for an additional 12% (change in R2; R2=0.683 when WC and PG were removed). One-hour after impact loading the explants regained their initial wet weight, that is, there was no measurable water gain or loss. In the impacted explants, chondrocyte death extended from the articular surface to the deeper zones and was uniformly distributed across the entire explant. The depth of cell death decreased non-linearly with increasing stress rate (Fig. 4a). However, the depth of cell death increased linearly with increasing peak stress and corresponding strain (Fig. 4b). Multiple linear regression analysis determined that the depth of the cell death (CD) was best accounted for by all of the independent
ARTICLE IN PRESS D. Milentijevic, P.A. Torzilli / Journal of Biomechanics 38 (2005) 493–502
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variables (s, s; ’ WC and PG, R2=0.677, po0.001), CDðmm=mmÞ ¼ 33:798ð798:778Þs þ 0:237ð70:198Þs 1:371ð71:574Þðln sÞ ’ þ 12:314ð727:967ÞPG þ 47:756ð7127:733ÞWC:
ð4Þ
Depth of Cell Death, % thickness
14 12
eðmm=mmÞ ¼ 0:093ð70:181Þ þ 0:003ð70:000Þs 0:017ð70:021Þðln sÞ ’
40 MPa
10
95% CI
30
þ 0:176ð70:446ÞPG:
8
ð5Þ
20
6 4
10
2 0 25
50
(a)
130
350
1000
Stress Rate, MPa/sec 16
Depth of Cell Death, % thickness
In this case, the depth of cell death was equally accounted for by each of the four independent parameters. The corresponding maximum strains (e) were found to decrease with increasing stress rate (Table 1). Statistically, the best combination of independent variables that accounted for the variation in matrix deformation were peak stress, stress rate and PG content (R2=0.556, po0.001),
25 MPa/s
14
1970:04ð75504:02ÞPG 2075:19ð75683:58ÞWC:
12 350
10
6 1000
4 2 0
(b)
0.10
0.15
0.20
ð6Þ
50
8
0.05
DIMðMPaÞ ¼ 1437:107ð74395:40Þ þ 156:319ð768:22Þðln sÞ ’
95% CI
130
The variation in the initial water content was not associated with the amount of explant deformation resulting from the impact load. The tissue stiffness or DIM increased exponentially with increasing rate of loading. Multiple linear regression analysis determined that explant stiffening was dependent on the stress rate and water and PG contents (R2=0.847, po0.001),
0.25
Max. Strain, mm/mm
Fig. 4. Depth of cell death from the articular surface, normalized to total explant thickness (%), as a function of (a) stress rate (logarithmic scale) and (b) maximum strain. (a) The depth of cell death increased with increasing peak stress and decreased with increasing stress rate. At the lowest peak stress of 10 MPa, cell viability (B2.5%) was independent of the stress rate. (b) The depth of cell death increased linearly with strain and decreased with increasing stress rate. The depth of cell death per unit strain (slope) was not significantly different between the 50, 130 and 350 MPa/s stress rates. The 95% confidence intervals (dotted lines) are plotted only for the highest and lowest stress rates.
The peak stress (>10 MPa) was not correlated with changes in the DIM, as expected. The stress rate was found to have the biggest influence on the DIM (R2=0.703 when PG and WC were removed), with the initial water and PG contents accounting for only 20% of the variation in DIM. In the experiments to compare the DZL to ASL explants (30 MPa at 25, 50, 130 and 1000 MPa/s), the DZL explants had less water loss (Fig. 3a), higher DIM and deformed less then the ASL explants at each stress rate (Table 1). Differences in all three parameters were greatest at the slowest stress rate (25 MPa/s). However, at the highest stress rate (1000 MPa/s) the DIM and maximum strain did not differ between the DZL and ASL configurations. A surprising finding was that the DZL explants had no cell death (similar to the unimpacted controls), compared to the ASL explants, which had cell death only in the superficial region.
Table 1 Effect of explant orientation (ASL vs. DZL) on maximum strain and DIM in explants impacted with 30 MPa peak stress at different stress rates (mean (SD), n=10) Stress rate
25 MPa/s
50 MPa/s
130 MPa/s
350 MPa/s
1000 MPa/s
Strain (mm/mm)
Asl Dzl
0.227 (0.046) 0.184 (0.041)a
0.187 (0.030) 0.159 (0.020)a
0.181 (0.035) 0.159 (0.034)
0.137 (0.024) 0.120 (0.020)
0.138 (0.031) 0.130 (0.020)
DIM (MPa)
Asl Dzl
224.7 (48.4) 300.0 (80.1)a
362.4 (62.5) 471.6 (66.9)a
382.4 (72.6) 525.0 (150.2)a
455.9 (111.9) 570.4 (137.0)a
849.9 (113.8) 1018.9 (340.3)
a
Significantly different from ASL group, po0.05.
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4. Discussion
0.20
0.16
Water Loss, mg/mg
The intention of the present study was to investigate the mechano-biological response of articular cartilage when subjected to similar impact stress rates and magnitudes (walking to joint injury) that may cause cartilage damage during joint trauma. Our findings showed that when impacted cartilage exudes a small amount of water relative to its initial water content, and that water loss decreases with increasing stress rate. Furthermore, chondrocyte death was initiated at the articular surface and propagated very little from the surface, remaining in the superficial tangential zone (STZ, 10–20% of total tissue thickness) for all combinations of peak stress and stress rate. In our study, we applied peak stress magnitudes of 10–40 MPa at stress rates of 25–1000 MPa/s. However, the exact in vivo physiological stress magnitudes and stress rates that articular cartilage experiences during normal dynamic and impact loading in the joint are not known. Peak stress magnitudes during normal activity are estimated to be in the range of 5–10 MPa (Afoke et al., 1987; Hodge et al., 1989), while much higher stress magnitudes (>40 MPa) are estimated to be necessary to produce intra-articular joint fractures (Borrelli et al., 1997; Haut, 1989). Stress rates can also be determined from these studies: 6–25 MPa/s during walking, 50–210 MPa/s during jogging, 140–250 MPa/s during jumping, B800 MPa/s for subchondral injury, and >1000 MPa/s for osteochondral fracture (Atkinson and Haut, 1995; Bergmann et al., 2001; Chang et al., 2001; Fukubayashi and Kurosawa, 1980; Haut, 1989; Hodge et al., 1989; King, 2001; Spagele et al., 1999; von Eisenhart et al., 1999). To our knowledge, this is the first study to specifically measure water loss (exudation) from articular cartilage when subjected to a wide range of impact loading rates. We found that when articular cartilage is subjected to impact loading, the majority of interstitial water is trapped within the matrix, probably due to the high frictional forces between the matrix and fluid as predicted by theoretical models (Armstrong et al., 1984; Mak et al., 1987). Impact loading times ranging from 18 msec (18 MPa at 1000 MPa/s) to 1600 msec (40 MPa at 25 MPa/s) caused water loss from 2.5% to 15% of the total water content, respectively (Fig. 5). At the highest stress rate (1000 MPa/s) the water loss was changed minimally (2–3%) with increasing stress magnitude (18–40 MPa) (Fig. 3b). Overall, water loss increased linearly with increasing loading time (Fig. 5). If our experimental results are extrapolated to an instantaneous response of the cartilage explant (i.e. loading duration of t=0+), approximately 2.24% of the water would still be exuded through the articular surface (Fig. 5). This indicates that independent of the loading time, a small amount of water loss would occur
499
Water Loss = Y0 + a*time, n = 183 Y0 = 0.0224 + 0.023 mg/mg, p<0.0001 a = 0.0882 + 0.044 (mg/mg)/sec, p<0.0001 r = 0.857, SEE = 0.0208 mg/mg
0.12 Mean + SEM
0.08
0.04
0.00 0.1
1
Loading Time, sec
Fig. 5. Water loss normalized to total water content in the explant as a function of loading time. Data is for all stress magnitudes (10, 20, 30 and 40 MPa) and stress rates (25, 50, 130, 350 and 1000 MPa/s). The amount of water loss varied linearly with loading time. For purposes of clarity, the loading time values were plotted on logarithmic scale.
instantaneously, most probably from the superficial matrix region at the time of load application as was predicted by the biphasic model (Mow et al., 1980). Our data further suggests that complete water entrapment within the cartilage matrix probably never (realistically) occurs. The amount of water loss measured in our study was solely caused by the impact loading. Possible experimental artifacts, such as blotting, water evaporation or handling of the explants during the impact tests did not cause any change in the water content of the nonimpacted controls (water loss=0.0070.01 mg/mg), which were processed identical to the impacted explants. In addition, the impacted explants did not appear to be structurally damaged, as they did not swell 1-h after impaction, an indicator for collagen network integrity (Maroudas, 1976). However, this later outcome was expected since the explants were impacted in confined compression, where the confinement prevented lateral expansion and therefore collagen failure in tension (Basser et al., 1998). While the cartilage explants underwent a large volumetric compression (confined compression), the amount of matrix deformation was always greater then the amount of water loss; this discrepancy increased with increasing stress rate. There are two possibilities for why the compressive strains were higher than the amount of water loss. First is that the water may have been lost only from the superficial tangential zone of the matrix. This zone is more likely to compact more (higher strain) then the deeper zones since it has a higher water content (Brocklehurst et al., 1984; Torzilli et al., 1990) and lower compressive modulus (Chen et al., 2001; O’Connor et al., 1988; Schinagl et al., 1997). The second possibility is that the impacted explant absorbed
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some of its exuded water when it recovered elastically upon unloading. This is supported by our data since the amount of residual strain (at zero stress) was not statistically different compared to the water loss (p=0.47), at any combination of peak stress and stress rate (Fig. 6). Since the amount water loss is volumetrically proportional to the matrix axial deformation at the end of loading, then the discrepancy between the amount of water loss and maximum strain could have occurred due to water absorption by the matrix during the unloading phase confirming our earlier assumption (Milentijevic and Torzilli, 2003). Even though we cannot conclude that the difference in water loss and matrix strain is a result of STZ compaction or water absorption during unloading, we believe that both of these mechanisms contributed to this discrepancy. This was also evident from our statistical analysis. The best predictors for water loss were peak stress and stress rate, while the initial water and PG contents were weaker predictors. This may be because the mechanical parameters (stress magnitude and rate) were well controlled and could be varied over a wider range than the initial water and PG contents, which did not vary as much between the specimens (COV=2.1% and 28.9%, respectively). As important, these later parameters are measures of total matrix content and do not describe their spatial variation with depth from the articular surface. We believe that the amount of water exuded through the articular surface is more affected by the spatial variation in the matrix composition and structure then by the total water and PG contents. This appears validated from the results of the explants loaded on the deep zone (DZL), as these 0.18
Water Loss, mg/mg
0.15
25 MPa/sec 50 MPa/sec 130 MPa/sec 1000 MPa/sec
0.12 95% CI
0.09
0.06 Slope = 0.91 + − 0.35 r = 0.933
0.03
0.00 0.00
0.03
0.06
0.09
0.12
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0.18
Residual Strain, mm/mm
Fig. 6. Water loss normalized to total water content in the explant as a function of the residual strain in the cartilage explants for all loading combinations (ASL only). The residual strain was measured at the end of the unloading phase when the stress was equal to zero. The measured water loss was equal to the residual strain and was dependent on the magnitude of the applied stress but not on the stress rate. The best-fit linear regression (solid line), 95% confidence intervals (dotted lines), and the statistical parameters for the regression line are shown in the figure.
explants lost less water compared to the explants loaded on the articular surface (ASL), thus indicating the significance of the inhomogeneous content and matrix structure in the ASL configuration to water loss. This supports our previous hypothesis that the superficial tangential zone will immediately collapse when loaded unless the articular surface is pressurized against the opposing cartilage surface, where the STZ would then resist compaction and result in no cell death (Milentijevic and Torzilli, 2003). A similar matrix compositional-structural depth dependency was probably responsible for the localization of the load-induced cell death to the superficial region of the matrix. The maximum depth of cell death from the surface was only 12% of the total tissue thickness (Fig. 4a). Even at the lowest stress magnitude of 10 MPa, cell viability was compromised. Statistical analysis found that not only the mechanical stimulus but also the composition of the cartilage matrix affected the depth of chondrocyte death. Furthermore, when water was restricted from exuding through the articular surface, as in the DZL explants, there was no detectable cell death in the superficial region, unlike that seen in the ASL explants. The superficial tangential zone in the DZL explants was able to resist compaction due to water pressurization when facing the impervious base of the confining chamber, and thus protected the chondrocytes in this region, as we previously observed in surface-to-surface explants loaded in confined compression (Milentijevic and Torzilli, 2003). Furthermore, the cells in the deeper regions were protected by the hydrostatic pressurization of the interstitial fluid within this region, which restricted matrix and cell deformation (water is considered incompressible and becomes immovable). Earlier studies have also shown that chondrocyte death depends on the loading magnitude and rate of loading (Ewers et al., 2001; Kurz et al., 2001; Quinn et al., 2001; Repo and Finlay, 1977; Torzilli et al., 1999). However, in these studies the explants were loaded in unconfined compression, resulting in large compressive strains (50–95%) due to the lateral expansion of the cartilage, often with failure of the matrix in tension (Armstrong et al., 1984). Thus, we believe there are two possible mechanisms responsible for chondrocyte damage (death) in the superficial region during impact loading. The first is due to STZ compaction caused by the exudation of water from the articular surface, as we found in this study. The second mechanism is due to STZ tensile failure, where cell death is found adjacent to surface fissures (Ewers et al., 2001; Quinn et al., 2001; Repo et al., 1977). As expected, the dynamic stiffness of the cartilage matrix (DIM) increased with increasing stress rate. We also found that the ASL explants had a lower DIM than the DZL explants (Table 1), which we attribute to the
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loss of water and compaction of the STZ. It should be noted that at the highest stress rate, where there is a negligible water movement (loss), the DIM is a good estimate of the intrinsic modulus (lower limit) for the solid components of the cartilage matrix. Hard tissues, such as bone (cortical and trabecular), also exhibit a rate-dependent compressive modulus. The compressive modulus of cortical bone is about 10 GPa, while for trabecular bone it ranges up to 500 MPa (Carter and Hayes, 1977). In our study, the DIM of cartilage was at or above the modulus of trabecular bone at the highest stress rates. In a special case such as in a hip where ‘‘ball-and-socket’’ confinement exists, the articulating cartilage layers may significantly increase in stiffness (due to pressurized fluid) during dynamic and impact loading. Our data suggest that when rapidly loaded the cartilage may exceed the stiffness of the trabecular bone, such that the impact energy could be transferred to and might cause damage to the bone. This hypothesis is supported from sports trauma (anterior cruciate ligament injuries) and car accidents (dash-board injuries) which commonly result in the ‘‘bone bruises’’ (hemorrhage or edema within the bone due to trabecular microfracture) seen on magnetic resonance images of the cartilage (Bealle and Johnson, 2000; Rosen et al., 1991). These lesions in the underlying trabecular bone indicate that significant trauma has occurred in a joint and is probably due to energy absorption in the bone as described above. In this study, we impacted viable articular cartilage with a broad range of impact stress rates and stress magnitudes that may exist during traumatic joint injury. We found that the amount of water loss and the depth of cell death were small relative to the total fluid volume and thickness of cartilage, respectively. Even though small, these findings were statistically significant. However, one must use caution when interpreting our findings as they may not represent in vivo joint loading. Obviously in joints such as the hip, shoulder and elbow, where good surface congruity and confinement exist, physiological dynamic and impact loading do not always harm the articular cartilage. A better understanding of the association between blunt trauma factors, such as stress magnitude and stress rate, and articular cartilage injury will be essential for the prevention and management of post-traumatic arthritis due to joint injury.
Acknowledgements This study was supported by a grant from the National Institutes of Health (AR465748). The authors gratefully acknowledge the assistance of Victoria K. Potterton for biochemical assays, Ali Sadegh, Ph.D. for helpful suggestions on the design of the impact tests, and
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David L. Helfet, M.D. and Hollis Potter, M.D. for clinical interpretation.
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Hodge, W.A., Carlson, K.L., Fijan, R.S., Burgess, R.G., Riley, P.O., Harris, W.H., Mann, R.W., 1989. Contact pressures from an instrumented hip endoprosthesis. Journal of Bone and Joint Surgery (American) 71 (9), 1378–1386. King, A.I., 2001. Fundamentals of Impact Biomechanics: Part—2 Biomechanics of the Abdomen, Pelvis, and Lower Extremities. Annual Review of Biomedical Engineering 3, 27–55. Kurz, B., Jin, M., Patwari, P., Cheng, D.M., Lark, M.W., Grodzinsky, A.J., 2001. Biosynthetic response and mechanical properties of articular cartilage after injurious compression. Journal of Orthopaedic Research 19 (6), 1140–1146. Mak, A.F., Lai, W.M., Mow, V.C., 1987. Biphasic indentation of articular cartilage—I. Theoretical analysis. Journal of Biomechanics 20 (7), 703–714. Maroudas, A.I., 1976. Balance between swelling pressure and collagen tension in normal and degenerate cartilage. Nature 260 (5554), 808–809. Milentijevic, D., Torzilli, P.A., 2003. Influence of stress magnitude on water loss and chondrocyte viability in impacted articular cartilage. Journal of Biomechanical Engineering 125 (5), 594–601. Mow, V.C., Kuei, S.C., Lai, W.M., Armstrong, C.G., 1980. Biphasic creep and stress relaxation of articular cartilage in compression? Theory and experiments. Journal of Biomechanical Engineering 102 (1), 73–84. O’Connor, P., Orford, C.R., Gardner, D.L., 1988. Differential response to compressive loads of zones of canine hyaline articular cartilage: micromechanical, light and electron microscopic studies. Annals of the Rheumatic Diseases 47 (5), 414–420. Oloyede, A., Flachsmann, R., Broom, N.D., 1992. The dramatic influence of loading velocity on the compressive response of articular cartilage. Connective Tissue Research 27 (4), 211–224. Quinn, T.M., Allen, R.G., Schalet, B.J., Perumbuli, P., Hunziker, E.B., 2001. Matrix and cell injury due to sub-impact loading of adult bovine articular cartilage explants: effects of strain rate and peak stress. Journal of Orthopaedic Research 19 (2), 242–249. Radin, E.L., Paul, I.L., Lowy, M., 1970. A comparison of the dynamic force transmitting properties of subchondral bone and articular
cartilage. Journal of Bone and Joint Surgery (American) 52 (3), 444–456. Repo, R.U., Finlay, J.B., 1977. Survival of articular cartilage after controlled impact. Journal of Bone and Joint Surgery (Amercian) 59 (8), 1068–1076. Rosen, M.A., Jackson, D.W., Berger, P.E., 1991. Occult osseous lesions documented by magnetic resonance imaging associated with anterior cruciate ligament ruptures. Arthroscopy 7 (1), 45–51. Schinagl, R.M., Gurskis, D., Chen, A.C., Sah, R.L., 1997. Depthdependent confined compression modulus of full-thickness bovine articular cartilage. Journal of Orthopaedic Research 15 (4), 499–506. Soltz, M.A., Ateshian, G.A., 1998. Experimental verification and theoretical prediction of cartilage interstitial fluid pressurization at an impermeable contact interface in confined compression. Journal of Biomechanics 31 (10), 927–934. Spagele, T., Kistner, A., Gollhofer, A., 1999. Modelling, simulation and optimisation of a human vertical jump. Journal of Biomechanics 32 (5), 521–530. Torzilli, P.A., Askari, E., Jenkins, J., 1990. Water content and solute diffusion properties of articular cartilage. In: Mow, V.C., Ratcliffe, A., Woo, S.L-Y. (Eds.), Biomechanics of Diarthroidial Joints. Springer, New York, pp. 363–390. Torzilli, P.A., Grigiene, R., Huang, C., Friedman, S.M., Doty, S.B., Boskey, A.L., Lust, G., 1997. Characterization of cartilage metabolic response to static and dynamic stress using a mechanical explant test system. Journal of Biomechanics 30 (1), 1–9. Torzilli, P.A., Grigiene, R., Borrelli Jr., J., Helfet, D.L., 1999. Effect of impact load on articular cartilage: cell metabolism and viability, and matrix water content. Journal of Biomechanical Engineering 121 (5), 433–441. von Eisenhart, R., Adam, C., Steinlechner, M., Muller-Gerbl, M., Eckstein, F., 1999. Quantitative determination of joint incongruity and pressure distribution during simulated gait and cartilage thickness in the human hip joint. Journal of Orthopaedic Research 17 (4), 532–539.
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Influence of stress rate on water loss, matrix deformation and chondrocyte viability in impacted articular cartilage Dejan Milentijevic*, Peter A. Torzilli Laboratory for Soft Tissue Research, Hospital for Special Surgery and Center for Biomedical Engineering, City University of New York, New York, NY 10021, USA Accepted 21 April 2004
Abstract The biomechanical response of articular cartilage to a wide range of impact loading rates was investigated for stress magnitudes that exist during joint trauma. Viable, intact bovine cartilage explants were impacted in confined compression with stress rates of 25, 50, 130 and 1000 MPa/s and stress magnitudes of 10, 20, 30 and 40 MPa. Water loss, cell viability, dynamic impact modulus (DIM) and matrix deformation were measured. Under all loading conditions the water loss was small (B15%); water loss increased linearly with increasing peak stress and decreased exponentially with increasing stress rate. Cell death was localized within the superficial zone (p 12% of total tissue thickness); the depth of cell death from the articular surface increased with peak stress and decreased with increasing stress rate. The DIM increased (200–700 MPa) and matrix deformation decreased with increasing stress rate. Initial water and proteoglycan (PG) content had a weak, yet significant influence on water loss, cell death and DIM. However, the significance of the inhomogeneous structure and composition of the cartilage matrix was accentuated when explants impacted on the deep zone had less water loss and matrix deformation, higher DIM, and no cell death compared to explants impacted on the articular surface. The mechano-biological response of articular cartilage depended on magnitude and rate of impact loading. r 2004 Elsevier Ltd. All rights reserved. Keywords: Cartilage; Impact; Stress rate; Stress magnitude; Water loss; Cell viability
1. Introduction Cartilage damage can occur from minor injuries such as in a sudden twist of an ankle (sport injury) to severe trauma such as in car accidents. During these different traumatic episodes the articular cartilage may experience different load and strain magnitudes as well as different load and strain rates. Articular cartilage resists joint loads by two distinctly different mechanisms that can be delineated by the rate (stress or strain) at which loading is applied. Under slow loading rates, cartilage behaves as a poroelastic material exhibiting fluid-flow dependent deformation, where the load is transiently resisted by the frictional drag between the fluid and the solid matrix while at equilibrium the load is resisted by *Corresponding author. Tel.: +1-212-774-2003; fax: +1-212-2492373. E-mail address: [email protected] (D. Milentijevic). 0021-9290/$ - see front matter r 2004 Elsevier Ltd. All rights reserved. doi:10.1016/j.jbiomech.2004.04.016
the solid phase only. On the other hand, under rapid and impact loading conditions cartilage exhibits fluid-flow independent deformation. Theoretical models predict that under impact loading conditions cartilage will respond as an incompressible, elastic material (fluid and solid assumed incompressible), where the interstitial fluid is incapable of escaping the dense matrix during the short time of loading (Armstrong et al., 1984; Mak et al., 1987; Ateshian and Wang, 1995; Soltz and Ateshian, 1998). How these different mechanical mechanisms effect the extent of damage to the cartilage and its post-injury survival is not well understood. Several studies have evaluated the affects of strain rate by loading cartilage explants in unconfined compression. Radin et al. (1970) loaded bovine cartilage (without underlying bone) at slow (0.1–16%/s) and fast (up to 670%/s) strain rates, and found that the compressive stiffness increased from 20 to 50 MPa with increasing strain rate (0.5–35%/s). Repo and Finlay
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(1977) impacted human articular cartilage-on-bone and reported that while matrix damage (fibrillation and cell death) was initiated at 20–30% strain (B25 MPa) it was independent of strain rate (500 vs. 1000 s1). They also attempted to quantify the amount of water loss but were unsuccessful because of large weighing errors. Later Oloyede et al. (1992) used a conventional materials testing machine to apple slow strain rates (5 105– 5 102 s1) and a free-swinging pendulum for impact strain rates (1000 s1). The compressive stiffness of the cartilage (modulus) increased with strain rate, however the modulus reached a maximum of B40 MPa for strain rates X5 102 s1. More recent studies have focused on the effects of strain rate ranging well below impact level. Quinn et al. (2001) loaded osteochondral specimens from young adult steer in unconfined compression at strain rates from 3 105 to 0.7 s1. At the slowest strain rate there was excessive chondrocyte death within the tissue (B100% loss of viability) with no visible matrix damage, while at the faster rates (0.3–0.7 s1) cell death occurred only in the superficial zone and adjacent to surface fissures. However, in another study using newborn calf cartilage, increasing strain rate (0.01–1 s1) resulted in a decrease in cell viability 3-days after compression (Kurz et al., 2001). Few studies have investigated how the impact loading (stress) rate effects cartilage damage. Ewers et al. (2001) impacted bovine cartilage samples (without underlying bone) in unconfined compression with a 40 MPa stress at two stress rates, 40 and 900 MPa/s, resulting in average strains of 41% and 48%, respectively. The impact caused superficial matrix damage (fissures), with cell death only close to the fissures at the high loading rate which propagated away from the fissures deeper into the matrix at the low loading rate. In a previous study, we impacted bovine cartilage in confined compression with peak stresses of 10–60 MPa at a single stress rate of 350 MPa/s (Milentijevic and Torzilli, 2003). Cell death was localized within the superficial region and increased linearly in depth from the surface with increasing peak stress. We also measured the water loss from the matrix immediately after impaction, which was small and also increased with increasing peak stress (1–10%). In previous studies, articular cartilage was loaded quasi-statically (slow rate) or with a single impact (fast rate). For the latter, drop-tower or pendulum type devices were predominantly used to generate high impact energy (fast rates). However, the loading rate and load magnitude are difficult to control independently, and thus it is difficult to assess their individual effects on cartilage response. To our knowledge, no study has evaluated how cartilage responds under a broad range of impact loading rates, nor have measurements been made of the resulting matrix compression
and concomitant water loss. Therefore, the objective of our study was to investigate the mechano-biological response of articular cartilage to varying loading rates and load magnitudes. We hypothesized that with increasing loading rate the cartilage matrix would respond mechanically by decreasing its matrix deformation and water loss, and thus reduce the amount and depth of cell death. To test this hypothesis, we impacted viable articular cartilage in confined compression with peak stresses of 10–40 MPa at stress rates of 25, 50, 130 and 1000 MPa/s, and measured water loss, cell viability, matrix deformation, and dynamic impact modulus (DIM). We found that cell death was localized within the superficial zone (p12% of total tissue thickness), and that the DIM increased (200–700 MPa) and matrix deformation and water loss decreased with increasing stress rate.
2. Materials and methods 2.1. Tissue acquisition Mature bovine knees with an intact capsule (young adult 18–24 months of age) were obtained from a local abattoir (Greenville, NJ) within twenty-four hours after death. A biopsy punch (diameter=8 mm) was used to cut circular explants from the femoral condyles and trochlear groove. Full-thickness disks were removed from the subchondral bone using a surgical blade, washed three times (15 min each) in Dulbecco’s Modified Eagle’s Media (DMEM, Gibco BRL, CA) containing 10% antibiotic/antimycotic (Ab/Am). The explants were then placed in DMEM with 10% Hepes buffer solution, 10% fetal bovine serum and 1% Ab/Am at 37 C and 95% humidity and incubated for least 24 h to allow the explants to reach an equilibrium water content and stabilize cell metabolism, as previously determined (Torzilli et al., 1997). Equilibrium matrix swelling was confirmed by independent measurements of water content after incubation for 20 min and 24 h (paired difference=0.170.4%, p>0.2, n=24). Initial water and proteoglycan (PG) contents and cell viability at the time of impact testing were determined as described below. 2.2. Impact test system A detailed description of the Cartilage Impact Test System (CITS) was reported earlier (Milentijevic and Torzilli, 2003). Briefly, the CITS uses a servo-controlled impactor (double-acting pneumatic cylinder) with a solid flat-ended cylindrical platen (6.35 mm) to apply a peak stress at a defined stress rate to the cartilage explant. A piezoelectric load transducer is used to control the peak stress while a flow control valve is used to set the loading rate. A high-speed, Linear Variable
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Differential Transformer (LVDT) is used to measure matrix deformation. At the time of testing, a cartilage disk (6.35 mm diameter) is impacted in confined compression by inserting it into a confining chamber (6.35 mm diameter) with the articular surface facing down onto a stainless-steel porous filter (39 mm pore size) (Fig. 1b). A data acquisition system is used to record the applied load, matrix deformation and time at a 5 kHz sampling rate per channel. The test system was calibrated for repeatability (precision) at each loading rate (stress rate, MPa/s) used in this study: 25 MPa/s (mean7SD=25.017 0.90 MPa/s, coefficient of variance (COV)=3.6%, n=5), 50 MPa/s (49.9670.83 MPa/s, COV=1.7%, n=5), 130 MPa/s (130.7871.99 MPa/s, COV=1.5%, n=5) and 1000 MPa/s (1003.9778.60 MPa/s, COV= 0.86%, n=5). In addition, the applied peak stresses (load/indenter’s cross-sectional area) of 10, 20, 30 and 40 MPa were also found to be repeatable at each stress rate (e.g., at 130 MPa/s and 30 MPa, peak stress= 30.0870.13 MPa, COV=0.44%, n=5). However, for the fastest stress rate of 1000 MPa/s the CITS was only able to apply peak stresses of 17, 23, 30 and 40 MPa due to the transient response of the test system. 2.3. Experimental design Impact tests were performed 1-to-3 days after cartilage harvest (post-stabilization). On the day of testing, cartilage explants were cored with a biopsy punch (d=6.35 mm) and the core (center) and remainder of the explant (ring) were returned to the media to stabilize (swell) for an additional 1 h (Fig. 1a). Independent tests were performed to confirm that the explants had reached a constant water content by measuring the wet weights of the cores and rings immediately after removal and 30 min and 1 h later (paired differences p0.1%, p>0.1, n=24). As described below, the cores
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were used for impact testing while the outer rings were used to determine the initial water and PG content of the core. Wet weights of the center (WWC, mg) and the ring (WWR, mg) were determined as follows. First the excess water was removed from the explant by gently blotting the specimen’s surfaces on a Kim Wipe (Kimberly Clark Corporation, Rowel, GA) and immediately weighing the explant using a micro-balance (resolution 71 mg, Cahn, Cerritos, CA). The wet weight was recorded exactly 15 s after placement on the balance to normalize for water evaporation and the explant returned to the fluid media. The center’s thickness was then measured using the LVDT by placing the explant in the CITS and manually lowering the load platen until a B60 kPa pressure was applied to the explant, at which point the thickness was recorded. After thickness measurement the center was returned into the fluid media. Immediately before impact loading, the center was blotted as described above and placed into the confining chamber with the articular surface resting on the dry porous filter. The center was then loaded with a stress rate of 25, 50, 130 or 1000 MPa/s and a peak stress of 10, 20 or 40 MPa. Eight specimens were impacted at each combination of stress rate and peak stress. Immediately after the impact, the center was quickly removed from the chamber, blotted, the wet weight (WWI, mg) measured, and the explant returned to the fluid media. After one hour of incubation the center was again blotted and weighed (WWS, mg). Control explants were treated identically but were not impacted. After determining the wet weights, the centers were evaluated for chondrocyte viability using the FDA (fluorescein diacetate, Molecular Probe, OR)/PI (propidium iodide, Sigma Chemicals, MO) live/dead cell essay. Thin cross-sections (B0.5 mm thick) were cut from each explant disk, incubated with FDA (10 mM)-PI (60 mM) stain for 5 min, rinsed in PBS (phosphate-buffered
Fig. 1. Schematic illustration of the impact test configurations used in the study. (a) The center of the cartilage explant (core, 6.35 mm diameter) was removed and impacted under different test configurations. The remainder of the explant (ring, 8.0 mm diameter) was used to determine the initial water and PG contents. (b) Articular surface loading (ASL): The core is positioned in the chamber with its articular surface facing the porous filter and the deep zone impacted with a nonporous load platen. The filter was placed on the bottom of the chamber to retain exuded water (due to gravity) when the explant was unloaded. (c) Deep zone loading (DZL): The cut-surface (deep zone) is placed against the porous filter and the articular surface impacted with the non-porous load platen. The ASL explants were impacted with 10, 20, 30 and 40 MPa at stress rates of 25, 50, 130 and 1000 MPa/s (n=9 explants per loading combination), while the DZL explants were impacted with a peak stress of 30 MPa at all stress rates (n=10 explants per loading combination).
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saline) and viewed using a fluorescent microscope (Nikon, NY). Digital images were recorded using an image capturing system (CCD camera, Optronics, Goleta, CA) attached to the microscope, and imaging software (SigmaScan, SPSS Science, Chicago, IL) was used to quantify the amount of chondrocyte death. Since cell death was initiated at the articular surface in all explants, the depth of cell death was calculated from the average of 20 measurements taken from the articular surface along the cross-section. The absolute depth (mm) of cell death from the surface was then calculated as the mean of two cross-sections per explant, and percent (%) depth calculated by normalizing this measurement to the initial cartilage thickness. The remaining rings were used to calculate the water and PG content of their paired cores as described below. The rings were freeze-dried overnight in a lyophilizer (Labconco, Kansas, MO) and their dry weights (DWR) determined. The rings were then digested in 1 ml Papain solution (Sigma-Aldrich Corp., St. Louis, MO) on a shaker water bath at 65C. Digestates were used to determine the PG content of each ring using the dimethylmethylene blue (DMMB) assay (Farndale et al., 1986). An independent impact test was performed to compare the amount of water lost from the deep zone (DZ) to that lost from articular surface (AS). Specimens were impacted with their cut surface (n=10) against the porous filter (Deep Zone Loading, DZL) (Fig. 1c) and with their AS (n=10) against the porous filter (Articular Surface Loading, ASL) (Fig. 1b), with a 30 MPa peak stress at a stress rate of 25, 50, 130 or 1000 MPa/s. Specimens not impacted served as control for each test configuration (n=5). 2.4. Initial water and pg content and water loss Preliminary experiments were performed to determine if there was a difference in initial water and PG contents between the center and outer ring from the same explant (Fig. 1a). Briefly, twelve explants (d=8 mm) were harvested from the bovine knee joint, cored (d=6.35 mm), and the paired wet weights, dry weights and PG content of the center (WWC, DWC and PGC) and ring (WWR, DWR and PGR) determined as described above. The water content of the center (WCC) and ring (WCR) were not statistically different (77.971.9% and 78.271.9%, respectively, p=0.7). A similar result was found for the PG content normalized per wet and dry weights (5.070.9% and 4.970.8%, p=0.89, and 23.775.9% and 24.374.5%, p=0.52, respectively). Since the water and PG contents of the center did not differ from the ring, we could calculate the initial water content in the center from the ring, and therefore utilize the center for the cell viability analysis. In the same fashion the PG content of the ring was used
to determine the initial PG content of the center. Finally, the change in water content (water loss) immediately after impact was calculated from Water Loss ¼ ðWWI WWC Þ=ðWWC WCR Þ;
ð1Þ
while the change in water content (water absorption) 1-h post impaction was calculated from Water Absorption ¼ ðWWS WWC Þ=ðWWC WCR Þ: ð2Þ 2.5. Data analysis The deformation of each explant was first corrected for apparatus compliance and individual stress–strain curves plotted to determine the maximum strain. Data graphing and statistical analysis were performed using commercially available software (SigmaPlot and SysStat,, SPSS Inc.). The DIM was calculated from the slope of the linear region of the loading phase for stresses X15 MPa. Linear regression analysis (with 795% confidence intervals) was performed for water loss and depth of cell death (propagation from the surface) as a function of peak stress, stress rate and corresponding maximum strains. A multiple regression analysis (forward, backwards, stepwise and all-possible procedures) was used to assess the effect of the independent parameters (peak stress, stress rate, PG content and water content) on water loss, depth of cell death, matrix deformation (strain) and DIM (Draper and Smith, 1966). The Student’s t-test (unpaired or paired, two-tailed, post-hoc for repeated measures) was used to test for differences between the dependent variables, while R2, t and F statistics were used to determine the best regression models. Statistical significance was set at a=0.05. Averaged data is given as mean7standard deviation unless otherwise indicated. For completeness, data from a previous study (Milentijevic and Torzilli, 2003) performed at a stress rate of 350 MPa/s was included in the analyses and figures.
3. Results The initial thickness and water and PG contents of the impacted explants was 1.2670.22 mm, 0.7770.02 mg/ mg (wet weight) and 0.2970.08 mg/mg (dry weight), respectively (n=105); for the non-impacted controls these values were 1.2570.18 mm, 0.7870.01 mg/mg and 0.2870.07 mg/mg (dry weight) (n=36). There was no significant difference in initial thickness and water and PG contents between impacted and non-impacted groups (p>0.7). The stress–strain response of the impacted samples exhibited a non-linear (toe) region (0–5 MPa) followed by a transition to a more linear region (>10 MPa)
ARTICLE IN PRESS D. Milentijevic, P.A. Torzilli / Journal of Biomechanics 38 (2005) 493–502 0.18
35 25 MPa/sec
1000 MPa/sec
40 MPa
30
0.15
Water Loss, mg/mg
Peak Stress, MPa
25
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Peak Stress = 30 MPa
20 15 10 5
0.12
0.09
30 Mean + SEM
20 30
0.06
DZL 10
0.03
0
0.00 25 0.00
0.05
0.10
0.15
0.20
50
130
350
1000
Stress Rate, MPa/sec
(a)
Strain, mm/mm
(Fig. 2). The toe regions appeared nearly identical when the stress–strain responses were superimposed for all loading combinations, while the slopes of the linear regions (DIM) increased with increasing stress rate. The hysteresis area between the loading and unloading phases also decreased with increasing stress rate. Immediately after load removal (zero stress), the stress–strain response indicated a residual strain in the tissue (Fig. 2). The amount of residual strain increased with increasing peak stress but decreased with increasing stress rate (25–350 MPa/s). Explants impacted at the highest stress rate of 1000 MPa/s had minimal residual strain upon load removal, which also varied minimally with peak stress. There was no change in the water content of the nonimpacted explants (controls) during the time it took to perform the impact tests, that is, between weighing pre- and post-impact (water loss=0.0070.01 mg/mg). On the other hand, the water loss in impacted explants increased linearly with increasing peak stress and corresponding maximum strains (Fig. 3b), while decreasing non-linearly with increasing stress rate (Fig. 3a). Multiple linear regression analysis determined that the best statistical combination of the independent variables (predictors) that were associated with changes in water loss (WL) were the peak stress (s), stress rate ðsÞ ’ and initial water (WC) and PG contents (R2=0.768, po0.001), WLðmg=mgÞ ¼ 0:089ð70:922Þ þ 0:002ð70:001Þs 0:060ð70:139Þðln sÞ ’ þ 0:004ð70:011Þðln sÞ ’ 2 þ 0:044ð70:257ÞPG þ 0:335ð71:169ÞWC:
ð3Þ
0.18 25 MPa/s
0.15
Water Loss, mg/mg
Fig. 2. Typical stress–strain response of explants impacted with a peak stress of 30 MPa at the slowest (25 MPa/s) and highest (1000 MPa/s) stress rates. All impacted explants exhibited a non-linear toe-region followed by a linear region. The slope of the linear region was used to determine the DIM. The area between the loading and unloading phases (hysteresis region) decreased with increasing stress rate.
95% CI
0.12
50 130
0.09
0.06 350
0.03
0.00 0.00 0.05
(b)
1000
0.10
0.15
0.20
0.25
Max. Strain, mm/mm
Fig. 3. Water loss from the cartilage matrix as a function of (a) impact stress rate (logarithmic scale) and (b) maximum strain. The water loss (mg) has been normalized to the total amount of water (mg) in the explant. (a) Water loss from the articular surface (ASL; solid lines) decreased exponentially with increasing stress rate. The DZL explants (dashed line), impacted with a 30 MPa peak stress, had significantly less water loss from the deep zone compared to the water loss in the ASL explants impacted with the same peak stress. (b) Water loss increased linearly with maximum strain and decreased with increasing stress rate. The 95% confidence intervals (dotted lines) are plotted only for the highest and lowest stress rates.
However, the stress and stress rate accounted for 68% of the WL, with the initial water and PG contents accounting for an additional 12% (change in R2; R2=0.683 when WC and PG were removed). One-hour after impact loading the explants regained their initial wet weight, that is, there was no measurable water gain or loss. In the impacted explants, chondrocyte death extended from the articular surface to the deeper zones and was uniformly distributed across the entire explant. The depth of cell death decreased non-linearly with increasing stress rate (Fig. 4a). However, the depth of cell death increased linearly with increasing peak stress and corresponding strain (Fig. 4b). Multiple linear regression analysis determined that the depth of the cell death (CD) was best accounted for by all of the independent
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variables (s, s; ’ WC and PG, R2=0.677, po0.001), CDðmm=mmÞ ¼ 33:798ð798:778Þs þ 0:237ð70:198Þs 1:371ð71:574Þðln sÞ ’ þ 12:314ð727:967ÞPG þ 47:756ð7127:733ÞWC:
ð4Þ
Depth of Cell Death, % thickness
14 12
eðmm=mmÞ ¼ 0:093ð70:181Þ þ 0:003ð70:000Þs 0:017ð70:021Þðln sÞ ’
40 MPa
10
95% CI
30
þ 0:176ð70:446ÞPG:
8
ð5Þ
20
6 4
10
2 0 25
50
(a)
130
350
1000
Stress Rate, MPa/sec 16
Depth of Cell Death, % thickness
In this case, the depth of cell death was equally accounted for by each of the four independent parameters. The corresponding maximum strains (e) were found to decrease with increasing stress rate (Table 1). Statistically, the best combination of independent variables that accounted for the variation in matrix deformation were peak stress, stress rate and PG content (R2=0.556, po0.001),
25 MPa/s
14
1970:04ð75504:02ÞPG 2075:19ð75683:58ÞWC:
12 350
10
6 1000
4 2 0
(b)
0.10
0.15
0.20
ð6Þ
50
8
0.05
DIMðMPaÞ ¼ 1437:107ð74395:40Þ þ 156:319ð768:22Þðln sÞ ’
95% CI
130
The variation in the initial water content was not associated with the amount of explant deformation resulting from the impact load. The tissue stiffness or DIM increased exponentially with increasing rate of loading. Multiple linear regression analysis determined that explant stiffening was dependent on the stress rate and water and PG contents (R2=0.847, po0.001),
0.25
Max. Strain, mm/mm
Fig. 4. Depth of cell death from the articular surface, normalized to total explant thickness (%), as a function of (a) stress rate (logarithmic scale) and (b) maximum strain. (a) The depth of cell death increased with increasing peak stress and decreased with increasing stress rate. At the lowest peak stress of 10 MPa, cell viability (B2.5%) was independent of the stress rate. (b) The depth of cell death increased linearly with strain and decreased with increasing stress rate. The depth of cell death per unit strain (slope) was not significantly different between the 50, 130 and 350 MPa/s stress rates. The 95% confidence intervals (dotted lines) are plotted only for the highest and lowest stress rates.
The peak stress (>10 MPa) was not correlated with changes in the DIM, as expected. The stress rate was found to have the biggest influence on the DIM (R2=0.703 when PG and WC were removed), with the initial water and PG contents accounting for only 20% of the variation in DIM. In the experiments to compare the DZL to ASL explants (30 MPa at 25, 50, 130 and 1000 MPa/s), the DZL explants had less water loss (Fig. 3a), higher DIM and deformed less then the ASL explants at each stress rate (Table 1). Differences in all three parameters were greatest at the slowest stress rate (25 MPa/s). However, at the highest stress rate (1000 MPa/s) the DIM and maximum strain did not differ between the DZL and ASL configurations. A surprising finding was that the DZL explants had no cell death (similar to the unimpacted controls), compared to the ASL explants, which had cell death only in the superficial region.
Table 1 Effect of explant orientation (ASL vs. DZL) on maximum strain and DIM in explants impacted with 30 MPa peak stress at different stress rates (mean (SD), n=10) Stress rate
25 MPa/s
50 MPa/s
130 MPa/s
350 MPa/s
1000 MPa/s
Strain (mm/mm)
Asl Dzl
0.227 (0.046) 0.184 (0.041)a
0.187 (0.030) 0.159 (0.020)a
0.181 (0.035) 0.159 (0.034)
0.137 (0.024) 0.120 (0.020)
0.138 (0.031) 0.130 (0.020)
DIM (MPa)
Asl Dzl
224.7 (48.4) 300.0 (80.1)a
362.4 (62.5) 471.6 (66.9)a
382.4 (72.6) 525.0 (150.2)a
455.9 (111.9) 570.4 (137.0)a
849.9 (113.8) 1018.9 (340.3)
a
Significantly different from ASL group, po0.05.
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4. Discussion
0.20
0.16
Water Loss, mg/mg
The intention of the present study was to investigate the mechano-biological response of articular cartilage when subjected to similar impact stress rates and magnitudes (walking to joint injury) that may cause cartilage damage during joint trauma. Our findings showed that when impacted cartilage exudes a small amount of water relative to its initial water content, and that water loss decreases with increasing stress rate. Furthermore, chondrocyte death was initiated at the articular surface and propagated very little from the surface, remaining in the superficial tangential zone (STZ, 10–20% of total tissue thickness) for all combinations of peak stress and stress rate. In our study, we applied peak stress magnitudes of 10–40 MPa at stress rates of 25–1000 MPa/s. However, the exact in vivo physiological stress magnitudes and stress rates that articular cartilage experiences during normal dynamic and impact loading in the joint are not known. Peak stress magnitudes during normal activity are estimated to be in the range of 5–10 MPa (Afoke et al., 1987; Hodge et al., 1989), while much higher stress magnitudes (>40 MPa) are estimated to be necessary to produce intra-articular joint fractures (Borrelli et al., 1997; Haut, 1989). Stress rates can also be determined from these studies: 6–25 MPa/s during walking, 50–210 MPa/s during jogging, 140–250 MPa/s during jumping, B800 MPa/s for subchondral injury, and >1000 MPa/s for osteochondral fracture (Atkinson and Haut, 1995; Bergmann et al., 2001; Chang et al., 2001; Fukubayashi and Kurosawa, 1980; Haut, 1989; Hodge et al., 1989; King, 2001; Spagele et al., 1999; von Eisenhart et al., 1999). To our knowledge, this is the first study to specifically measure water loss (exudation) from articular cartilage when subjected to a wide range of impact loading rates. We found that when articular cartilage is subjected to impact loading, the majority of interstitial water is trapped within the matrix, probably due to the high frictional forces between the matrix and fluid as predicted by theoretical models (Armstrong et al., 1984; Mak et al., 1987). Impact loading times ranging from 18 msec (18 MPa at 1000 MPa/s) to 1600 msec (40 MPa at 25 MPa/s) caused water loss from 2.5% to 15% of the total water content, respectively (Fig. 5). At the highest stress rate (1000 MPa/s) the water loss was changed minimally (2–3%) with increasing stress magnitude (18–40 MPa) (Fig. 3b). Overall, water loss increased linearly with increasing loading time (Fig. 5). If our experimental results are extrapolated to an instantaneous response of the cartilage explant (i.e. loading duration of t=0+), approximately 2.24% of the water would still be exuded through the articular surface (Fig. 5). This indicates that independent of the loading time, a small amount of water loss would occur
499
Water Loss = Y0 + a*time, n = 183 Y0 = 0.0224 + 0.023 mg/mg, p<0.0001 a = 0.0882 + 0.044 (mg/mg)/sec, p<0.0001 r = 0.857, SEE = 0.0208 mg/mg
0.12 Mean + SEM
0.08
0.04
0.00 0.1
1
Loading Time, sec
Fig. 5. Water loss normalized to total water content in the explant as a function of loading time. Data is for all stress magnitudes (10, 20, 30 and 40 MPa) and stress rates (25, 50, 130, 350 and 1000 MPa/s). The amount of water loss varied linearly with loading time. For purposes of clarity, the loading time values were plotted on logarithmic scale.
instantaneously, most probably from the superficial matrix region at the time of load application as was predicted by the biphasic model (Mow et al., 1980). Our data further suggests that complete water entrapment within the cartilage matrix probably never (realistically) occurs. The amount of water loss measured in our study was solely caused by the impact loading. Possible experimental artifacts, such as blotting, water evaporation or handling of the explants during the impact tests did not cause any change in the water content of the nonimpacted controls (water loss=0.0070.01 mg/mg), which were processed identical to the impacted explants. In addition, the impacted explants did not appear to be structurally damaged, as they did not swell 1-h after impaction, an indicator for collagen network integrity (Maroudas, 1976). However, this later outcome was expected since the explants were impacted in confined compression, where the confinement prevented lateral expansion and therefore collagen failure in tension (Basser et al., 1998). While the cartilage explants underwent a large volumetric compression (confined compression), the amount of matrix deformation was always greater then the amount of water loss; this discrepancy increased with increasing stress rate. There are two possibilities for why the compressive strains were higher than the amount of water loss. First is that the water may have been lost only from the superficial tangential zone of the matrix. This zone is more likely to compact more (higher strain) then the deeper zones since it has a higher water content (Brocklehurst et al., 1984; Torzilli et al., 1990) and lower compressive modulus (Chen et al., 2001; O’Connor et al., 1988; Schinagl et al., 1997). The second possibility is that the impacted explant absorbed
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some of its exuded water when it recovered elastically upon unloading. This is supported by our data since the amount of residual strain (at zero stress) was not statistically different compared to the water loss (p=0.47), at any combination of peak stress and stress rate (Fig. 6). Since the amount water loss is volumetrically proportional to the matrix axial deformation at the end of loading, then the discrepancy between the amount of water loss and maximum strain could have occurred due to water absorption by the matrix during the unloading phase confirming our earlier assumption (Milentijevic and Torzilli, 2003). Even though we cannot conclude that the difference in water loss and matrix strain is a result of STZ compaction or water absorption during unloading, we believe that both of these mechanisms contributed to this discrepancy. This was also evident from our statistical analysis. The best predictors for water loss were peak stress and stress rate, while the initial water and PG contents were weaker predictors. This may be because the mechanical parameters (stress magnitude and rate) were well controlled and could be varied over a wider range than the initial water and PG contents, which did not vary as much between the specimens (COV=2.1% and 28.9%, respectively). As important, these later parameters are measures of total matrix content and do not describe their spatial variation with depth from the articular surface. We believe that the amount of water exuded through the articular surface is more affected by the spatial variation in the matrix composition and structure then by the total water and PG contents. This appears validated from the results of the explants loaded on the deep zone (DZL), as these 0.18
Water Loss, mg/mg
0.15
25 MPa/sec 50 MPa/sec 130 MPa/sec 1000 MPa/sec
0.12 95% CI
0.09
0.06 Slope = 0.91 + − 0.35 r = 0.933
0.03
0.00 0.00
0.03
0.06
0.09
0.12
0.15
0.18
Residual Strain, mm/mm
Fig. 6. Water loss normalized to total water content in the explant as a function of the residual strain in the cartilage explants for all loading combinations (ASL only). The residual strain was measured at the end of the unloading phase when the stress was equal to zero. The measured water loss was equal to the residual strain and was dependent on the magnitude of the applied stress but not on the stress rate. The best-fit linear regression (solid line), 95% confidence intervals (dotted lines), and the statistical parameters for the regression line are shown in the figure.
explants lost less water compared to the explants loaded on the articular surface (ASL), thus indicating the significance of the inhomogeneous content and matrix structure in the ASL configuration to water loss. This supports our previous hypothesis that the superficial tangential zone will immediately collapse when loaded unless the articular surface is pressurized against the opposing cartilage surface, where the STZ would then resist compaction and result in no cell death (Milentijevic and Torzilli, 2003). A similar matrix compositional-structural depth dependency was probably responsible for the localization of the load-induced cell death to the superficial region of the matrix. The maximum depth of cell death from the surface was only 12% of the total tissue thickness (Fig. 4a). Even at the lowest stress magnitude of 10 MPa, cell viability was compromised. Statistical analysis found that not only the mechanical stimulus but also the composition of the cartilage matrix affected the depth of chondrocyte death. Furthermore, when water was restricted from exuding through the articular surface, as in the DZL explants, there was no detectable cell death in the superficial region, unlike that seen in the ASL explants. The superficial tangential zone in the DZL explants was able to resist compaction due to water pressurization when facing the impervious base of the confining chamber, and thus protected the chondrocytes in this region, as we previously observed in surface-to-surface explants loaded in confined compression (Milentijevic and Torzilli, 2003). Furthermore, the cells in the deeper regions were protected by the hydrostatic pressurization of the interstitial fluid within this region, which restricted matrix and cell deformation (water is considered incompressible and becomes immovable). Earlier studies have also shown that chondrocyte death depends on the loading magnitude and rate of loading (Ewers et al., 2001; Kurz et al., 2001; Quinn et al., 2001; Repo and Finlay, 1977; Torzilli et al., 1999). However, in these studies the explants were loaded in unconfined compression, resulting in large compressive strains (50–95%) due to the lateral expansion of the cartilage, often with failure of the matrix in tension (Armstrong et al., 1984). Thus, we believe there are two possible mechanisms responsible for chondrocyte damage (death) in the superficial region during impact loading. The first is due to STZ compaction caused by the exudation of water from the articular surface, as we found in this study. The second mechanism is due to STZ tensile failure, where cell death is found adjacent to surface fissures (Ewers et al., 2001; Quinn et al., 2001; Repo et al., 1977). As expected, the dynamic stiffness of the cartilage matrix (DIM) increased with increasing stress rate. We also found that the ASL explants had a lower DIM than the DZL explants (Table 1), which we attribute to the
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loss of water and compaction of the STZ. It should be noted that at the highest stress rate, where there is a negligible water movement (loss), the DIM is a good estimate of the intrinsic modulus (lower limit) for the solid components of the cartilage matrix. Hard tissues, such as bone (cortical and trabecular), also exhibit a rate-dependent compressive modulus. The compressive modulus of cortical bone is about 10 GPa, while for trabecular bone it ranges up to 500 MPa (Carter and Hayes, 1977). In our study, the DIM of cartilage was at or above the modulus of trabecular bone at the highest stress rates. In a special case such as in a hip where ‘‘ball-and-socket’’ confinement exists, the articulating cartilage layers may significantly increase in stiffness (due to pressurized fluid) during dynamic and impact loading. Our data suggest that when rapidly loaded the cartilage may exceed the stiffness of the trabecular bone, such that the impact energy could be transferred to and might cause damage to the bone. This hypothesis is supported from sports trauma (anterior cruciate ligament injuries) and car accidents (dash-board injuries) which commonly result in the ‘‘bone bruises’’ (hemorrhage or edema within the bone due to trabecular microfracture) seen on magnetic resonance images of the cartilage (Bealle and Johnson, 2000; Rosen et al., 1991). These lesions in the underlying trabecular bone indicate that significant trauma has occurred in a joint and is probably due to energy absorption in the bone as described above. In this study, we impacted viable articular cartilage with a broad range of impact stress rates and stress magnitudes that may exist during traumatic joint injury. We found that the amount of water loss and the depth of cell death were small relative to the total fluid volume and thickness of cartilage, respectively. Even though small, these findings were statistically significant. However, one must use caution when interpreting our findings as they may not represent in vivo joint loading. Obviously in joints such as the hip, shoulder and elbow, where good surface congruity and confinement exist, physiological dynamic and impact loading do not always harm the articular cartilage. A better understanding of the association between blunt trauma factors, such as stress magnitude and stress rate, and articular cartilage injury will be essential for the prevention and management of post-traumatic arthritis due to joint injury.
Acknowledgements This study was supported by a grant from the National Institutes of Health (AR465748). The authors gratefully acknowledge the assistance of Victoria K. Potterton for biochemical assays, Ali Sadegh, Ph.D. for helpful suggestions on the design of the impact tests, and
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David L. Helfet, M.D. and Hollis Potter, M.D. for clinical interpretation.
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