PMCCPMCCPMCC

Search tips
Search criteria 

Advanced

 
Logo of nihpaAbout Author manuscriptsSubmit a manuscriptHHS Public Access; Author Manuscript; Accepted for publication in peer reviewed journal;
 
Osteoarthritis Cartilage. Author manuscript; available in PMC 2013 May 1.
Published in final edited form as:
PMCID: PMC3384701
NIHMSID: NIHMS360551

Response of cartilage and meniscus tissue explants to in vitro compressive overload

Abstract

Objective

To examine the relative susceptibility of cartilage and meniscus tissues to mechanical injury by applying a single, controlled overload and observing cellular, biochemical, and mechanical changes.

Design

Cartilage and meniscus tissue explants in radial confinement were subjected to a range of injury by indenting to 40% strain at three different strain rates: 0.5%/s (slow), 5%/s (medium), or 50%/s (fast). Following injury, samples were cultured for either 1 or 9 days. Explants were assayed for cell metabolic activity, water content, and sGAG content. Mechanical properties of explants were determined in torsional shear and unconfined compression. Conditioned medium was assayed for sGAG and LDH release.

Results

Peak injury force increased with strain rate but both tissues displayed little to no macroscopic damage. Cell metabolism was lowest in medium and fast groups on day 1. Cell lysis increased with peak injury force and loading rate in both tissues. In contrast, sGAG content and release did not significantly vary with loading rate. Additionally, mechanical properties did not significantly vary with loading rate in either tissue.

Conclusion

By use of a custom confinement chamber, large peak forces were obtained without macroscopic destruction of the explants. At the loads achieved in this studied, cell damage was induced without detectable physical or compositional changes. These results indicate that sub-failure injury can induce biologic damage that may not be readily detected and could be an early event in osteoarthritis genesis.

Keywords: Cartilage, Meniscus, Injury, In vitro, Overload

Introduction

Traumatic injury of the knee has been long associated with the development of osteoarthritis (OA)[1]. Prolonged exposure to overloading due to obesity[2, 3], occupational loading[3-5] or altered biomechanics[6-8] is associated with the early onset of OA, suggesting that non-traumatic overloading can also be detrimental to long-term joint health. Knee loading affects multiple tissues, but the post-traumatic osteoarthritis literature has predominantly focused on the responses of articular cartilage. The menisci are important in load transfer and joint stability of the knee joint[9], yet their response to sub-failure injury is relatively unknown. In this study, we explore the susceptibility of both cartilage and meniscus to a range of simulated, sub-catastrophic mechanical overloads.

In vivo experiments demonstrate that high impact loads to animal knee joints can induce osteoarthritic changes in cartilage such as proteoglycan loss, decreased cartilage integrity, and cell death. While in vivo studies highlight the role of mechanical trauma for OA development, it is difficult to gain insight into the injury response of specific tissues in the knee joint. A number of in vitro studies have investigated cartilage injury from drop-tower impacts[10, 11], controlled overload[12-16], and cyclic loading[17, 18]. These studies showed that greater loads impart greater cell death, proteoglycan release, collagen damage, surface fissuring, loss of mechanical properties, and decreased cell synthesis of matrix constituents. Investigators have sought to identify a threshold value of peak stress that cartilage can tolerate without significant damage. Depending on factors such as repetition, loading rate, presence of underlying bone, and radial confinement, this peak value can vary from 6-15 MPa[17, 19]. The precise stress level required to induce chondrocyte death has been related to the gel diffusion rate, the characteristic rate of load-induced fluid flow, which depends on both the loading configuration and the mechanical properties of the tissue[14]. In radially unconfined compression, no cell death was observed for stresses below 14 MPa at a low strain rate, whereas at higher strain rates cell death increased substantially. These studies demonstrate that peak strain, strain rate, and physical boundary conditions are all influential parameters in injury-induced cartilage damage.

In contrast to cartilage, meniscal injury has received comparatively little attention. Imaging studies reveal that meniscal lesions are seen in 70% of early OA patients[20], meniscal damage may precede that of articular cartilage in the lateral compartment of OA knees[21], and meniscal malposition increases the risk of cartilage loss[22]. A prospective case-control study found that the presence of meniscal damage yielded an odds ratio of 5.7 for developing radiographic OA within a 30-month period[23]. However, despite strong evidence that meniscal damage is implicated in OA development, relatively few studies have directly investigated injury to meniscus. ACL-transected animal models of OA showed meniscal tears as early as 12 weeks after injury in rabbits[24] and dogs[25], presumably due to overload and macroscopic tissue failure. In vitro studies show that meniscal explants exhibit increased proteoglycan release[26, 27], nitric oxide production[28, 29], and increased levels of gene expression for IL-1 and iNOS, and catabolic MMPs and aggrecanases[30] in response to dynamic mechanical compression, indicating that the meniscus may be a source of pro-inflammatory mediators[29] and catabolic activity[30] in the knee joint. However, in a direct comparison of the short-term response of cartilage and meniscus explants to in vitro overload in unconfined compression (50% at 100%/s), meniscus explants exhibited a downregulation of catabolic gene expression[31]. A loss of cell viability was found at the meniscus surface and throughout cartilage explants, but only cartilage explants displayed fissured surfaces.

In this study, we employed an in vitro model to deliver controlled overloads to meniscus and cartilage tissue explants. Our goal was to compare responses of meniscal and cartilage tissues to varied levels of mechanical insult through quantifying resulting cell death and tissue damage. Based on prior reports, our hypothesis was that meniscal tissue may be more susceptible than cartilage to immediate mechanical damage and increased cell death and tissue degradation would lead to decreased tissue properties over time.

Methods

Materials

Immature bovine stifle joints were from Research 87 (Marlborough, MA). High glucose Dulbecco’s modified Eagle medium (DMEM), N-(2-hydroxyethyl)-piperazine-N’-2-ethanesulfonic acid (HEPES), non-essential amino acids (NEAA), antibiotic-antimycotic solution (100 U/mL penicillin, 100 mg/mL streptomycin, and 0.25 mg/mL amphotericin B), proteinase K, and phosphate buffered saline (PBS) were from Invitrogen (Carlsbad, CA). Fetal bovine serum was from HyClone (Logan, UT). L-ascorbic acid 2-phosphate, ammonium acetate, protease inhibitor cocktail, 1,9-dimethyl-methylene blue (DMMB) dye and shark chondroitin sulfate were from Sigma (St. Louis, MO). Protease inhibitor cocktail set I was from Calbiochem (San Diego, CA). Biopsy punches were from Miltex (York, PA). The Quick Cell Proliferation Assay Kit was from BioVision (Mountain View, CA) and the CytoTox96 Non-Radioactive Cytotoxicity Assay kit was from Promega (Madison, WI).

Tissue culture

Articular cartilage and meniscal fibrocartilage explants (n=10/tissue/condition/endpoint; total of 160 explants) were harvested from seven immature bovine stifles within one day of slaughter using a 6 mm biopsy punch (donor animals were not identified; stifles were from 4-7 animals). Cartilage samples were removed from both femoral condyles and meniscus samples were removed from tibial aspects of both menisci. Explants were trimmed to 2 mm thickness using a custom slicing block, leaving the articular surfaces intact, and heights measured at three locations with a contact-sensing micrometer were averaged to determine actual thickness of each sample. Prior to loading, explants were cultured for 6 days at 37oC, 5% CO2, and 95% relative humidity in high-glucose DMEM supplemented with 10% fetal bovine serum, 10 mM HEPES buffer, 0.1 mM NEAA, 50 g/mL L-ascorbic acid-2-PO4, and 1% antibiotic-antimycotic solution, with medium changes every other day. Explants were randomly allocated to loading and endpoint groups.

In vitro overload

A custom semi-confinement loading fixture was constructed for the loading protocol (Fig 1). The chamber consisted of an aluminum base and a 60 mm inner diameter polyvinyl chloride cylindrical wall. A porous, sintered stainless steel filter sat beneath the explants and a 5 mm thick silicon rubber ring with a 6 mm diameter hole limited radial bulging during compressive loading. An impermeable, stainless steel indenter with a 3.2 mm diameter hemispherical tip was used to injure explants at room temperature on a materials testing frame (Instron 5848, Norwood, MA).

Fig 1
Tissue explants (6mm diameter by 2mm thick) sat on a permeable base and were radially confined by a rubber ring. A 3.2mm impermeable, hemispherical indentor was used to compress explants by 40% of their measured thickness at variable strain rates. The ...

On the day of loading, sample heights measured at three locations were averaged to determine the thickness of each sample, which was used to define both peak displacement and displacement rate based on the targeted peak strain and strain rate. Explants were individually moved to the sanitized loading chamber containing PBS with 1% antibiotic-antimycotic. Indenter surfaces were sanitized with ethanol prior to loading of each explant. Explants were indented under displacement control to a peak 40% nominal compressive strain at three different nominal strain rates: Slow (S): 0.5%/s, Medium (M): 5%/s, and Fast (F): 50%/s. Peak strain was held for one minute and explants were then unloaded. Control (C) samples were placed in the loading chamber but not compressed for a duration equivalent to that of a slow ramp and hold. Explants were rinsed three times in PBS with 1% antibiotic-antimycotic and returned to culture in fresh medium. Force-displacement curves were recorded for each sample. Peak stresses for compressed explants were estimated as the peak force divided by the projected contact area predicted by a numerical analysis of large deformation, Hertzian contact between a rigid sphere and an incompressible, neo-Hookean layer of finite thickness[32].

Explants were cultured for 1 or 9 days post loading. Culture medium was changed and collected at 1, 2, 3, 5, 7 and 9 days post-loading, with medium volumes of 0.5 mL through 3 days and 1 mL subsequently. Conditioned medium was stored frozen for later biochemical analysis.

Cell viability and lysis

Mitochondrial dehydrogenase activity, generally viewed as a measure of cell proliferation or viability, was measured using the BioVision Quick Cell Proliferation Assay Kit, which measures the cleavage of tetrazolium salt WST-1 into a red formazan product. At the end of each culture period, tissue explants were incubated in fresh medium with 1% WST-1 reagent (0.25 mL for 4 hours). Absorbance values at 440 nm of conditioned media were measured using a spectrophotometric plate reader. Explants were then stored at -20oC in PBS with protease inhibitors.

Cell lysis was measured by quantifying the release of the intracellular enzyme lactate dehydrogenase (LDH) using the Promega CytoTox96 Non-Radioactive Cytotoxicity Assay, which measures enzymatic cleavage of tetrazolium salt 2-(4-Iodophenyl)-3-(4-nitrophenyl)-5-phenyl-2H-tetrazolium chloride into a red formazan product. Absorbance values at 490 nm were measured using a spectrophotometric plate reader, with background media absorbance values subtracted.

Mechanical testing

Prior to mechanical testing, explants were visually examined for surface fissures and India ink was swabbed on half of the day 9 explants to highlight damage. The inner, directly loaded region of the explant was isolated from the surrounding tissue annulus using a 4 mm biopsy punch. Core thickness was measured using a contact sensing micrometer and used to determine strains for mechanical testing. All mechanical testing was conducted in PBS with protease inhibitors. Samples were first tested in oscillatory torsional shear at 37°C using an AR 2000ex torsional rheometer (TA Instruments, New Castle, DE). Samples were compressed by 10% of the measured thickness at 2 μm/s and allowed to stress relax for 20 minutes. A ±0.5% nominal shear strain was applied at 0.1Hz to determine dynamic shear modulus G*. Samples were unloaded, allowed to recover for 15 minutes at room temperature and then tested in unconfined compression at room temperature using an Instron 5848. Samples were sequentially compressed to 5, 10, 15, and 20% nominal compressive strain at 0.1%/s and allowed to stress relax at each offset for 20 minutes. The equilibrium compressive modulus Eeq was determined by linear regression of the relaxed stress against applied strain. After relaxation at 10% offset, a ±1.5% compressive strain was applied at 0.1 Hz to determine dynamic compressive modulus E*.

Biochemical analysis

Following mechanical testing, explants were weighed, lyophilized overnight (Freezone 4.5 Freeze Dry System, LabConco, Kansas City, MO), and re-weighed dry. Explants were digested in proteinase K (1mg/80mg of cartilage, 1mg/20mg of meniscus) buffered with 100mM ammonium acetate at 60°C overnight. Sulfated glycosaminoglycan (sGAG) contents of digested explants and conditioned media were assayed using the dimethymethylene blue (DMMB) assay[33] using chondroitin sulfate standards. Explant water and sGAG contents were expressed as fractions of wet mass.

Data analysis

Of the 160 samples, three were fully excluded from analysis because of artifacts during loading, two were excluded from mechanical testing and explant biochemical analysis due to errors/artifacts, and two were excluded from explant sGAG analysis due to handling errors. Data were analyzed using General Linear Models (GLMs) using Minitab (version 16, Minitab, Inc., State College, PA). Data for GLMs were processed with Box-Cox transformation using the exponent for each outcome measure that minimized the pooled standard deviation, with rounding when standard exponents (e.g., 1, 0.5, -0.5, 0) fell within the confidence intervals. Bonferroni’s test was used for pairwise planned comparisons for main or interaction effects with p<0.05. The peak loading force, estimated peak stress, thickness, and 24 hour LDH and sGAG release data were examined using three-way (tissue, loading rate, medial-lateral) GLMs with stifle as random factor. Mechanical, explant biochemical and WST-1 results were analyzed using four-way (tissue, loading rate, day, medial-lateral) GLMs with stifle as a random factor. For each tissue, relationships between outcomes and peak loading force were examined for log-transformed data via linear mixed models using SPSS Statistics (version 19, SPSS Inc., Chicago, IL), with day and medial-lateral source as fixed factors, log force as a covariate and stifle as a random factor. Best models were identified by minimizing the corrected Akaike information criterion through backwards selection. Data are presented as means with 95% confidence intervals.

Results

Overload system

After six days of culture prior to loading, cartilage explants swelled by 8.49% (6.69%, 10.29%) whereas meniscus did not exhibit swelling (-1.47%, 1.10%) as defined by thickness change. Cartilage explants were significantly thicker than meniscus samples at time of loading (p=0.010), but there were no thickness differences among groups for either tissue prior to loading (p=0.85). For each sample, the thickness at time of loading was used as a basis for target strains and strain rates. Peak forces (Fig 2A) increased with strain rate in both tissues. Cartilage explants experienced higher peak forces than meniscus explants at each strain rate and peak force increased significantly with each increase in strain rate for both tissues (p≤0.0001). Forces were higher for lateral explants than medial explants for meniscus (p≤0.0001) but not cartilage (p=1.00). Patterns of estimated peak stress (Fig 2B) were similar to those of peak force. For the fast loading groups, estimated peak stresses for cartilage and meniscus were 11.0 MPa (9.37 MPa, 12.7 MPa) and 4.63 MPa (3.17 MPa, 6.09 MPa), respectively. Despite high peak stresses, relatively little macroscopic tissue damage was evident. No visible surface damage was observed following overload, and India ink staining revealed little surface damage. In cartilage, one medium strain rate explant had light staining and one fast rate explant had intense staining indicative of surface fissures. In meniscus, no load-induced surface damage was observed. Some loaded cartilage explants exhibited bulging of the annular ring and distinct indentation at the loading site after 9 days of culture, but no similar changes were noted in meniscus explants.

Fig 2Fig 2
Peak applied forces (A) and estimated stresses (B) for Control, Slow, Medium and Fast rate groups. Peak forces and stresses increased with increasing strain rate, were higher for cartilage than meniscus, and higher for lateral menisci than medial menisci. ...

Cell viability and lysis

WST-1 conversion (Fig 3), which is expected to reflect the number of viable cells, was greater for cartilage on day 1 than for meniscus on day 1 or either tissue on day 9 (p<0.0001). Additionally, WST-1 conversion on day 1 was significantly lower for medium and fast groups than for slow or control groups (p≤0.031), but across groups did not vary significantly with peak force for either tissue (Table I). WST-1 conversion did not vary significantly between tissues or among loading rate groups at day 9 (p≥0.20) or with medial-lateral source (p=0.35).

Fig 3
Cell viability as measured by the WST-1 colorimetric assay for cartilage (A) and meniscus (B) explants on day 1 (black bars) and day 9 (grey bars) after loading. Asterisks (*) indicate p≤0.05 vs. other rate groups on that day. Precise significance ...
Table I
Estimated relationships between log(outcome) and log(peak force)

Cell lysis measured by LDH release was greatest in the first day post-loading. Twenty-four hour LDH release (Fig 4A) was greater in meniscus than in cartilage for slow and control groups (p≤0.0001). Cartilage LDH release increased with strain rate among loaded explants, and was higher for medium and fast groups than control or slow groups (p=0.0019), but did not vary with medial-lateral source (p=1.00). Meniscus LDH release was higher in the fast group than in the slow group (p=0.0132) and greater for lateral samples (p=0.001). LDH release increased with peak loading force in both tissues (Table I).

Fig 4Fig 4
LDH release (A) and sGAG release (B) to the media for cartilage (filled circles) and meniscus (open circles) explants during the first 24 hours after loading. LDH and sGAG release were higher for lateral menisci than medial menisci. Coefficients for model ...

sGAG release

Sulfated glycosaminoglycan (sGAG) release was quantified by measuring sGAG content in the media throughout the culture period. Twenty-four hour sGAG release (Fig 4B) was greater for cartilage explants than for meniscus explants (p<0.0001), but did not vary among loading rate groups (p=0.24). For both tissues, sGAG release increased with loading force and was greater for lateral than medial samples (Table I).

Explant composition

The water mass fraction was greater for cartilage than for meniscus (p<0.0001). Water fraction was slightly higher in the fast group than in medium group (p=0.048). The cartilage water content increased significantly from day 1 to day 9 (p<0.0001) but did not differ for meniscus (p=0.48). Cartilage explants had substantially higher sGAG contents (Fig. 5) than meniscus explants (p<0.0001) representative of baseline differences between tissues. sGAG content was greater at day 1 than at day 9 across groups (p=0.0006). For meniscus explants, the sGAG content was positively related to peak force at day 1 (Table I) and higher for lateral samples (p=0.0009).

Fig 5
Explant core sGAG content normalized to wet mass for cartilage (A) and meniscus (B) explants at day 1 (black bars) and day 9 (grey bars) after loading. sGAG content was higher for lateral menisci than medial menisci. Precise significance values are reported ...

Mechanical testing

All measured mechanical properties were higher for cartilage explants than for meniscus explants (p<0.0001), representative of baseline differences between tissues. The equilibrium modulus was lower at day 9 than at day 1 across groups for cartilage explants (p<0.0001) but not for meniscus explants (p=0.23). Both the dynamic compressive modulus and dynamic shear modulus were lower at day 9 than at day 1 across groups for both tissues (p≤0.0404). For cartilage explants, the equilibrium modulus for lateral samples was positively related to peak force, but no other properties varied strongly with peak force. For meniscus explants, equilibrium and dynamic compressive moduli at day 1 were positively related to peak force, and all moduli were greater for lateral samples (Table I).

Discussion

Acute injury or chronic overload of the knee joint can involve elevated loading of both cartilage and meniscus tissues, but there has been little comparison between tissues regarding susceptibility to overload. Furthermore, many studies have focused on macroscopic damage to the cartilage matrix though high overloads[10, 34, 35]. In this study, we investigated damage to cartilage and meniscal explants in response to a range of loads in a system intended to limit macroscopic damage. The diameter of the spherical indenter was smaller than the explant to allow surrounding tissue to constrain deformations and minimize involvement of the disrupted tissue at the cut surfaces of the explant. Likewise, retention of the tissue surface layers avoided loading the tissue through a cut (and therefore damaged) surface. The confinement ring limited lateral expansion and permitted production of large forces that would typically cause gross destruction of explants in unconfined compression. This is particularly true for meniscus tissue, which is highly compliant in unconfined compression[36-38].

This in vitro system allows reproducible loading under conditions allowing direct comparisons of cartilage and meniscus tissues. As with other in vitro overload models, this system does not fully replicate the complex loading patterns involved in clinically relevant injuries, and in particular does not reproduce the circumferential tensile stress induced in the meniscus under physiologic loading. However, compressive overloading at low rates is relevant to situations such as occupational overload and obesity[3] in which joints are overloaded for long durations of time, as well as clinical scenarios such as femoroacetabular impingement, in which both labral fibrocartilage and acetabular cartilage may see chronic, abnormal compression[39]. Like other similar in vitro single overload models employing low strain rates[13, 14], high peak strains achieved at low strain rates produced neither substantial cell death nor matrix failure. Conversely, large stress induced at higher rates of loading is a strong predictor for injury, perhaps due to rapid interstitial fluid pressurization[40, 41].

Immature bovine tissues have been widely used to study cartilage mechanics[42, 43], mechanobiology[44, 45] and degradation[46, 47] due in part to their ready availability. While appropriate for initial comparisons of tissue overload responses, this model has limitations. Tissue composition and cell metabolism change with maturation, and chondrocyte viability following in vitro injury was found to be lower in mature tissues[18, 48]. Future investigations would benefit by investigating overload response in mature tissue to identify age-related changes in meniscal tissues, particularly in tissues exhibiting early degenerative changes. Additionally, while LDH release (an indirect indicator of cell death) increased with loading severity for both tissues, we did not extensively examine the spatial distribution of cell death or matrix damage. Chondrocyte death following in vitro injury was found to be greatest at the surface[17], where it appears to initiate[18, 41], and progress deeper into the tissue with lower strain rate and higher peak stress. Early chondrocyte death was reported to occur primarily through necrosis[49], which is followed by cellular apoptosis that may spread via intercellular signaling[50]. We observed in preliminary live-dead stains of cartilage explants at one and nine days after loading (data not shown) that medium and fast loaded groups had extensive regions of cell death through the tissue depth whereas the slow rate had cell death confined to the surfaces of explants. A more thorough examination of the spatial distribution and mechanisms of meniscus cell death may provide valuable information on regional susceptibility to injury and the functional roles of the meniscal surface layer.

In this study, explants were precultured for six days prior to loading. While a shorter preculture might reflect native tissue conditions more accurately, both sGAG and LDH release rates during this period were low, mitigating concerns about influences of the preculture period on patterns of results. Typical of in vitro culture, cartilage explants swelled significantly during the preculture period, likely due to additional sGAG accumulation and some disruption of the collagen network due to explant isolation. In contrast, meniscus explants did not swell, likely due to the lower proteoglycan content and lower proteoglycan synthesis rates of meniscus. Because of tissue swelling, target strains and strain rates were determined using the measured thickness at the time of loading, reflecting the true state of the tissue.

All measured mechanical properties were much lower for meniscus explants than for cartilage explants, reflecting in part baseline difference in sGAG content between tissues. We observed positive dependence of mechanical properties, sGAG release, and sGAG content on peak force. While it is possible that these patterns were responses to applied loading, it should be noted that the load levels incurred in this study, particularly for meniscus explants, were below levels required to initiate physical tissue damage. Thus, these patterns may in part reflect the induction of greater peak forces through compression of samples with higher initial sGAG contents and therefore higher moduli, as demonstrated by higher forces, sGAG content, and moduli of lateral explants.

LDH release for both tissues was highest in the fast strain rate group one day after loading. On the same day, we also observed lower WST-1 conversion in the medium and fast strain rate groups. By nine days post-loading, both measures of cell viability were stable among groups. While WST-1 conversion is generally viewed as reflecting the number of viable cells, it may also be temporarily decreased by treatments that do not kill cells. Additionally, the WST-1 assay likely disproportionately reflects activity of cell populations near the surface of the explant, due to the greater diffusion times between the explant center and the medium. As implemented, this may increase the assay’s sensitivity to cell death near the explant surface. As employed, the WST-1 assay measures a reduction in a signal that may be dominated by the activity of unaffected cells in the unloaded region. In contrast, the LDH assay is dominated by an increased signal from lysed cells, and is thus more sensitive to localized cell death.

LDH release was positively correlated to peak force in both tissues, indicating that cell death occurred in both tissues at similar levels of loads. This indicates that the overload configuration used in this study induces biological damage at load levels well below those required to produce detectable physical damage or changes in tissue composition. This behavior was observed in low-level impacted cartilage explants where chondrocyte death preceded matrix damage[51]. As found in a similar study[31], the levels of cell death in meniscus explants were comparable to those in cartilage despite the lack of visible signs of damage in meniscal samples. In a preliminary study, explants (n=5/tissue/condition) were loaded with greater peak forces that produced estimated peak stresses for cartilage and meniscus of 17.9 MPa (8.83 MPa, 26.9 MPa) and 25.8 MPa (19.4 MPa, 32.2 MPa), respectively, for the highest loading group. After 9 days of culture, cartilage compressive moduli were reduced in the highest loading group, but material properties did not significantly change for meniscus explants. However, cell lysis measured by LDH release was greater in meniscus than in cartilage (p≤0.0001) and increased with peak load, consistent with observations in this report.

Taken together, results of these studies indicate that biologic damage of both cartilage and meniscus can occur at levels well below that required to produce detectable physical damage. This suggests that even moderate overload may be capable of inducing early biological damage in both tissues. Cell death was an early event in tissue trauma and may be undetectable beyond 24 hours after injury using methods presented. Thus, biological damage to cartilage and meniscus may be difficult to clinically detect. Additionally, it appears that meniscus may be more physically robust to macroscopic damage than cartilage, but may experience similar levels of cellular trauma. This raises the possibility of “hidden” biological damage in an apparently intact meniscus as a potential early event in the initiation or progression of meniscal degeneration. Further studies of the biological response to sub-failure injury may lead to a greater understanding of meniscal involvement in the early stages of knee OA.

Acknowledgements

This study was supported by a grant from the National Institutes of Health (NIAMS R01AR052861) and a National Science Foundation Graduate Research Fellowship.

Footnotes

Author Contributions: JN and ML participated in study conception and design, data analysis and interpretation, and drafting and revision of the manuscript. ML supervised the study and JN performed the experiments and acquired data.

Conflicts of Interest: The authors have no conflicts of interest to declare.

Publisher's Disclaimer: This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting proof before it is published in its final citable form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.

References

1. Gelber A, Hochberg M, Mead L, Wang N-Y, Wigley F, Klag M. Joint injury in young adults and risk for subsequent knee and hip osteoarthritis. Annals of Internal Medicine. 2000;133:321–328. [PubMed]
2. Davis MA, Ettinger WH, Neuhaus JM, Cho SA, Hauck WW. The association of knee injury and obesity with unilateral and bilateral osteoarthritis of the knee. American journal of epidemiology. 1989;130:278–288. [PubMed]
3. Anderson JJ, Felson DT. Factors associated with osteoarthritis of the knee in the first national Health and Nutrition Examination Survey (HANES I). Evidence for an association with overweight, race, and physical demands of work. Am J Epidemiol. 1988;128:179–189. [PubMed]
4. Rytter S, Egund N, Jensen LK, Bonde JP. Occupational kneeling and radiographic tibiofemoral and patellofemoral osteoarthritis. J Occup Med Toxicol. 2009;4:19. [PMC free article] [PubMed]
5. Coggon D, Croft P, Kellingray S, Barrett D, McLaren M, Cooper C. Occupational physical activities and osteoarthritis of the knee. Arthritis Rheum. 2000;43:1443–1449. [PubMed]
6. Brandt KD. Insights into the natural history of osteoarthritis provided by the cruciate-deficient dog. An animal model of osteoarthritis. Annals of the New York Academy of Sciences. 1994;732:199–205. [PubMed]
7. Englund M. The role of the meniscus in osteoarthritis genesis. Rheumatic diseases clinics of North America. 2008;34:573–579. [PubMed]
8. Setton LA, Elliott DM, Mow VC. Altered mechanics of cartilage with osteoarthritis: human osteoarthritis and an experimental model of joint degeneration. Osteoarthritis and Cartilage. 1999;7:2–14. [PubMed]
9. Fithian DC, Kelly MA, Mow VC. Material properties and structure-function relationships in the menisci. Clinical orthopaedics and related research. 1990;(252):19–31. [PubMed]
10. Jeffrey JE, Thomson LA, Aspden RM. Matrix loss and synthesis following a single impact load on articular cartilage in vitro. Biochimica et biophysica acta. 1997;1334:223–232. [PubMed]
11. Natoli RM, Scott CC, Athanasiou KA. Temporal effects of impact on articular cartilage cell death, gene expression, matrix biochemistry, and biomechanics. Annals of Biomedical Engineering. 2008;36:780–792. [PubMed]
12. DiMicco MA, Patwari P, Siparsky PN, Kumar S, Pratta MA, Lark MW, et al. Mechanisms and kinetics of glycosaminoglycan release following in vitro cartilage injury. Arthritis and rheumatism. 2004;50:840–848. [PubMed]
13. Kurz B, Jin M, Patwari P, Cheng DM, Lark MW, Grodzinsky AJ. Biosynthetic response and mechanical properties of articular cartilage after injurious compression. Journal of Orthopaedic Research. 2001;19:1140–1146. [PubMed]
14. Morel V, Quinn TM. Cartilage injury by ramp compression near the gel diffusion rate. Journal of Orthopaedic Research. 2004;22:145–151. [PubMed]
15. Morel V, Quinn TM. Short-term changes in cell and matrix damage following mechanical injury of articular cartilage explants and modelling of microphysical mediators. Biorheology. 2004;41:509–519. [PubMed]
16. Quinn TM, Allen RG, Schalet BJ, Perumbuli P, Hunziker EB. 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. 2001;19:242–249. [PubMed]
17. Clements KM, Bee ZC, Crossingham GV, Adams MA, Sharif M. How severe must repetitive loading be to kill chondrocytes in articular cartilage? Osteoarthritis and Cartilage. 2001;9:499–507. [PubMed]
18. Levin AS, Chen C-TC, Torzilli PA. Effect of tissue maturity on cell viability in load-injured articular cartilage explants. Osteoarthritis and Cartilage. 2005;13:488–496. [PubMed]
19. Torzilli PA, Grigiene R, Borrelli J, Helfet DL. Effect of impact load on articular cartilage: cell metabolism and viability, and matrix water content. Journal of Biomechanical Engineering. 1999;121:433–441. [PubMed]
20. Karachalios T, Zibis A, Papanagiotou P, Karantanas AH, Malizos KN, Roidis N. MR imaging findings in early osteoarthritis of the knee. European Journal of Radiology. 2004;50:225–230. [PubMed]
21. Bennett LD, Buckland-Wright JC. Meniscal and articular cartilage changes in knee osteoarthritis: a cross-sectional double-contrast macroradiographic study. Rheumatology. 2002;41:917–923. [PubMed]
22. Hunter DJ, Zhang YQ, Niu JB, Tu X, Amin S, Clancy M, et al. The association of meniscal pathologic changes with cartilage loss in symptomatic knee osteoarthritis. Arthritis and rheumatism. 2006;54:795–801. [PubMed]
23. Englund M, Guermazi A, Roemer FW, Aliabadi P, Yang M, Lewis CE, et al. Meniscal tear in knees without surgery and the development of radiographic osteoarthritis among middle-aged and elderly persons: The Multicenter Osteoarthritis Study. Arthritis and rheumatism. 2009;60:831–839. [PMC free article] [PubMed]
24. Killian ML, Isaac DI, Haut RC, Déjardin LM, Leetun D, Donahue TL Haut. Traumatic anterior cruciate ligament tear and its implications on meniscal degradation: A preliminary novel lapine osteoarthritis model. The Journal of surgical research. 2009;8:1–8. [PubMed]
25. Smith GN, Mickler EA, Albrecht ME, Myers SL, Brandt KD. Severity of medial meniscus damage in the canine knee after anterior cruciate ligament transection. Osteoarthritis and Cartilage. 2002;10:321–326. [PubMed]
26. McHenry JA, Zielinska B, Donahue TLH. Proteoglycan breakdown of meniscal explants following dynamic compression using a novel bioreactor. Annals of Biomedical Engineering. 2006;34:1758–1766. [PMC free article] [PubMed]
27. Shin S-J, Fermor B, Weinberg JB, Pisetsky DS, Guilak F. Regulation of matrix turnover in meniscal explants: role of mechanical stress, interleukin-1, and nitric oxide. Journal of Applied Physiology. 2003;95:308–313. [PubMed]
28. Gupta T, Zielinska B, McHenry J, Kadmiel M, Donahue TL Haut. IL-1 and iNOS gene expression and NO synthesis in the superior region of meniscal explants are dependent on the magnitude of compressive strains. Osteoarthritis and Cartilage. 2008;16:1213–1219. [PubMed]
29. Hennerbichler A, Fermor B, Hennerbichler D, Weinberg JB, Guilak F. Regional differences in prostaglandin E2 and nitric oxide production in the knee meniscus in response to dynamic compression. Biochemical and biophysical research communications. 2007;358:1047–1053. [PMC free article] [PubMed]
30. Zielinska B, Killian M, Kadmiel M, Nelsen M, Donahue TL Haut. Meniscal tissue explants response depends on level of dynamic compressive strain. Osteoarthritis and Cartilage. 2009;17:754–760. [PubMed]
31. Kisiday JD, Vanderploeg EJ, McIlwraith CW, Grodzinsky AJ, Frisbie DD. Mechanical injury of explants from the articulating surface of the inner meniscus. Archives of biochemistry and biophysics. 2010;494:138–144. [PubMed]
32. Lin Y, Chang C, Lee W. Effects of thickness on the largely-deformed JKR (Johnson–Kendall–Roberts) test of soft elastic layers. International Journal of Solids and Structures. 2008;45:2220–2232.
33. Farndale RW, Buttle DJ, Barrett AJ. Improved quantitation and discrimination of sulphated glycosaminoglycans by use of dimethylmethylene blue. Biochimica et biophysica acta. 1986;883:173–177. [PubMed]
34. Ewers BJ, Dvoracek-Driksna D, Orth MW, Haut RC. The extent of matrix damage and chondrocyte death in mechanically traumatized articular cartilage explants depends on rate of loading. J Orthop Res. 2001;19:779–784. [PubMed]
35. Lewis JL, Deloria LB, Oyen-Tiesma M, Thompson RC, Jr., Ericson M, Oegema TR., Jr Cell death after cartilage impact occurs around matrix cracks. J Orthop Res. 2003;21:881–887. [PubMed]
36. Chia HN, Hull ML. Compressive moduli of the human medial meniscus in the axial and radial directions at equilibrium and at a physiological strain rate. Journal of Orthopaedic Research. 2008;26:951–956. [PubMed]
37. Bursac P, Arnoczky S, York A. Dynamic compressive behavior of human meniscus correlates with its extra-cellular matrix composition. Biorheology. 2009;46:227–237. [PubMed]
38. Lai JH, Levenston ME. Meniscus and cartilage exhibit distinct intra-tissue strain distributions under unconfined compression. Osteoarthritis and Cartilage. 2010;18:1291–1299. [PMC free article] [PubMed]
39. Beck M, Kalhor M, Leunig M, Ganz R. Hip morphology influences the pattern of damage to the acetabular cartilage: femoroacetabular impingement as a cause of early osteoarthritis of the hip. J Bone Joint Surg Br. 2005;87:1012–1018. [PubMed]
40. Chen CT, Burton-Wurster N, Lust G, Bank RA, Tekoppele JM. Compositional and metabolic changes in damaged cartilage are peak-stress, stress-rate, and loading-duration dependent. J Orthop Res. 1999;17:870–879. [PubMed]
41. Milentijevic D, Torzilli PA. Influence of stress rate on water loss, matrix deformation and chondrocyte viability in impacted articular cartilage. Journal of Biomechanics. 2005;38:493–502. [PubMed]
42. Bursac PM, Obitz TW, Eisenberg SR, Stamenovic D. Confined and unconfined stress relaxation of cartilage: appropriateness of a transversely isotropic analysis. J Biomech. 1999;32:1125–1130. [PubMed]
43. Chen AC, Bae WC, Schinagl RM, Sah RL. Depth- and strain-dependent mechanical and electromechanical properties of full-thickness bovine articular cartilage in confined compression. J Biomech. 2001;34:1–12. [PubMed]
44. Bonassar LJ, Grodzinsky AJ, Frank EH, Davila SG, Bhaktav NR, Trippel SB. The effect of dynamic compression on the response of articular cartilage to insulin-like growth factor-I. J Orthop Res. 2001;19:11–17. [PubMed]
45. Jin M, Frank EH, Quinn TM, Hunziker EB, Grodzinsky AJ. Tissue shear deformation stimulates proteoglycan and protein biosynthesis in bovine cartilage explants. Arch Biochem Biophys. 2001;395:41–48. [PubMed]
46. Patwari P, Gao G, Lee JH, Grodzinsky AJ, Sandy JD. Analysis of ADAMTS4 and MT4-MMP indicates that both are involved in aggrecanolysis in interleukin-1-treated bovine cartilage. Osteoarthritis Cartilage. 2005;13:269–277. [PMC free article] [PubMed]
47. Quinn TM, Grodzinsky AJ, Hunziker EB, Sandy JD. Effects of injurious compression on matrix turnover around individual cells in calf articular cartilage explants. J Orthop Res. 1998;16:490–499. [PubMed]
48. Kurz B, Lemke A, Kehn M, Domm C, Patwari P, Frank EH, et al. Influence of tissue maturation and antioxidants on the apoptotic response of articular cartilage after injurious compression. Arthritis and rheumatism. 2004;50:123–130. [PubMed]
49. Chen CT, Burton-Wurster N, Borden C, Hueffer K, Bloom SE, Lust G. Chondrocyte necrosis and apoptosis in impact damaged articular cartilage. J Orthop Res. 2001;19:703–711. [PubMed]
50. Levin A, Burton-Wurster N, Chen CT, Lust G. Intercellular signaling as a cause of cell death in cyclically impacted cartilage explants. Osteoarthritis Cartilage. 2001;9:702–711. [PubMed]
51. Duda GN, Eilers M, Loh L, Hoffman JE, Kaab M, Schaser K. Chondrocyte death precedes structural damage in blunt impact trauma. Clin Orthop Relat Res. 2001:302–309. [PubMed]