Skip to main content
NCBI home page
NIHPA Author Manuscripts logoLink to NIHPA Author Manuscripts
. Author manuscript; available in PMC: 2021 Jun 23.
Published in final edited form as: J Biomech. 2020 May 19;107:109852. doi: 10.1016/j.jbiomech.2020.109852

Immature Bovine Cartilage Wear by Fatigue Failure and Delamination

Krista M Durney 1, Courtney A Shaeffer 2, Brandon K Zimmerman 2, Robert J Nims 1, Sevan Oungoulian 2, Brian K Jones 2, James F Boorman-Padgett 1, Jason T Suh 2, Roshan P Shah 3, Clark T Hung 1, Gerard A Ateshian 1,2
PMCID: PMC7809915  NIHMSID: NIHMS1599037  PMID: 32517855

Abstract

This study investigated wear damage of immature bovine articular cartilage using reciprocal sliding of tibial cartilage strips against glass or cartilage. Experiments were conducted in physiological buffered saline (PBS) or mature bovine synovial fluid (SF). A total of 63 samples were tested, of which 47 exhibited wear damage due to delamination of the cartilage surface initiated in the middle zone, with no evidence of abrasive wear. There was no difference between the friction coefficient of damaged and undamaged samples, showing that delamination wear occurs even when friction remains low under a migrating contact area configuration. No difference was observed in the onset of damage or in the friction coefficient between samples tested in PBS or SF. The onset of damage occurred earlier when testing cartilage against glass versus cartilage against cartilage, supporting the hypothesis that delamination occurs due to fatigue failure of the collagen in the middle zone, since stiffer glass produces higher strains and tensile stresses under comparable loads. The findings of this study are novel because they establish that delamination of the articular surface, starting in the middle zone, may represent a primary mechanism of failure. Based on preliminary data, it is reasonable to hypothesize that delamination wear via subsurface fatigue failure is similarly the primary mechanism of human cartilage wear under normal loading conditions, albeit requiring far more cycles of loading than in immature bovine cartilage.

Keywords: Articular cartilage, friction and wear, fatigue failure, delamination

Introduction

Osteoarthritis (OA) affects millions of Americans: it is a progressive, complex, multi-tissue joint disease with degenerative changes in the articular cartilage and subchondral bone (Ashkavand et al., 2013), with a long asymptomatic early development and debilitating late stages. OA is viewed today as a disease of the joint as an organ, with inflammation, injury, and changes in bone, articular cartilage, and synovial fluid (SF) as potential driving forces. The degradation of the extracellular matrix (ECM) components of cartilage is key to the progression of the disease (Dijkgraaf et al., 1995; Loeser, 2013).

Regardless of the initiating factors of OA, cartilage stresses produced by sliding contact of the articular layers mediate the progression of tissue degeneration. The mechanisms by which stresses produce progressive tissue degeneration via mechanical pathways remain poorly understood. OA is often described as a natural process of wear and tear associated with aging, or an initiating traumatic event. The cartilage mechanics literature has mostly focused on examining the friction coefficient μ as a surrogate for understanding wear and tear in cartilage (Ateshian and Mow, 2005). The friction coefficient of articular cartilage is not constant (McCutchen, 1962). The lowest reported value of μ for cartilage against glass is typically μ≈0.002 (Krishnan et al., 2004), which is exceptionally low. However, μ may rise over time, depending on loading conditions, to achieve values as high as μ≈0.15 against glass (Krishnan et al., 2004), or even μ≈0.5 against stainless steel (Forster and Fisher, 1996; Forster and Fisher, 1999). These values are expected to be detrimental to the integrity of cartilage, though they are not normally achieved under physiologic loading conditions (Ateshian, 2009).

When cartilage slides against cartilage, it produces a migrating contact area (MCA) configuration that sustains elevated interstitial fluid load support (Caligaris and Ateshian, 2008). As a result, the friction coefficient remains low for sustained durations; for bovine and human cartilage in saline, it is typically μ≈0.025; for human cartilage in SF it is only slightly lower, μ≈0.020 (Caligaris and Ateshian, 2008; Caligaris et al., 2009). Implicitly, a low value of μ has been assumed to produce low wear while an elevated value could lead to significant wear. Despite the prominence of this hypothesized mechanism, only a few cartilage wear studies have been performed under controlled conditions, most notably by investigating PRG4 knockout mice (Jay et al., 2007), since PRG4/lubricin has been shown to reduce the friction coefficient of cartilage in vitro (Schmidt et al., 2007) and prevent degeneration in vivo (Flannery et al., 2009). However, it has also been shown that advancing osteoarthritic degeneration does not increase the friction coefficient of human cartilage (Caligaris et al., 2009), suggesting that OA wear may progress without a concomitant increase in μ.

Wear is a generally complex phenomenon that may manifest itself in different ways. In the engineering tribology literature, a broad range of wear mechanisms are reported, many of which are mostly applicable to metals and other artificial surfaces (Moore, 1975). However, some of these mechanisms may also be candidates for wear of biological tissues. These mechanisms include abrasive wear, which removes particulates of matter from the bearing surfaces, third-body wear, where particulate matter causes further abrasion of the bearing surfaces, fatigue wear with delamination, where the load-bearing material fails below the surface due to fatigue and the failure propagates until a lamina shears off, and chemical wear, where breakdown of the bearing material is initiated by chemical reactions, such as proteolysis in biological tissues. Other phenomena such as adhesion or stick-slip friction have also been proposed as initiators of cartilage damage (Han and Eriten, 2018; Lee et al., 2013).

Biological tissues, such as articular cartilage and tendons have been reported to fail in fatigue under both tension and compression. In studying tensile fatigue of human articular cartilage, Weightman found that fatigue life is reduced with an increase in tensile stress and the tissue’s resistance to fatigue decreases with age (Weightman, 1976). In another study, Weightman et al. showed that fatigue failure of articular cartilage is a possibility in the span of an average lifetime (Weightman et al., 1978). Additionally, in a study of fatigue of human tendons, Schechtman and Bader found a highly significant relationship between stress and ln N, where N is the number of cycles to failure, suggesting that tendons fail in fatigue (Schechtman and Bader, 1997). Recent studies confirmed that fatigue is the failure mechanism of articular cartilage when under cyclic compression loading (Kaplan et al., 2017; Vazquez et al., 2019).

In our recent study on immature bovine cartilage (Oungoulian et al., 2015), wear tests were performed on cartilage plugs sliding against glass or various metals used in orthopaedic implants, producing delamination of the superficial zone with negligible abrasive wear. These results were consistent with the fact that delamination is a clinically recognized symptom of OA (Meachim, 1982; Pritzker et al., 2006). However, a potential limitation of that study was our adoption of a stationary contact area (SCA) testing configuration, which promoted loss of interstitial fluid pressurization over time. It could be argued that the elevated friction coefficient achieved under those conditions would not occur under more physiological loading conditions. Furthermore, with prolonged wear testing, complete delamination and removal of the top layer of these plugs was observed (Oungoulian et al., 2015), raising the possibility that the initiating failure resulted from edge effects between the flat counterface material and the circular edge of the plug surface.

Therefore, in this study, we performed experiments on immature bovine cartilage to test our primary hypothesis (H1) that delamination wear occurs even when the friction coefficient μ remains low under a migrating contact area configuration (MCA) (Caligaris and Ateshian, 2008; Caligaris et al., 2009; Northwood et al., 2007). We used large, rectangular cartilage strips harvested from the medial or lateral tibial plateau, loaded with a glass lens under low physiological contact stresses, such that the contact area remains well within the strip boundaries to avoid edge effects.

Based on prior literature findings regarding the role of SF boundary lubricants on the reduction of friction and wear (Flannery et al., 2009; Jay et al., 2007; Schmidt et al., 2007), we also tested the hypothesis (H2) that SF delays the onset of cartilage delamination when compared to physiological buffered saline (PBS).

Based on our previously reported model for the dependence of the frictional force on interstitial fluid load support and the solid-on-solid contact area fraction (Ateshian et al., 1998; Soltz et al., 2003), we tested the hypothesis (H3) that loading cartilage against cartilage delays the onset of delamination wear compared to testing glass on cartilage, since contacting porous cartilage layers exhibit a much smaller solid-on-solid contact area fraction than porous cartilage contacting impermeable glass.

Materials and Methods

Experimental Design

Hypotheses H1-H3 were investigated in a set of experiments using two identical friction testing devices that could apply contact loads up to 45 N (Study 1) and a third testing device that could apply a maximum contact load of 356 N (Study 2), producing more physiological levels of contact stress. Six test groups were included in Study 1, as summarized in Table 1, using either a semi-convex glass lens (G) or a femoral condylar cartilage counterface (C) with similar radius of curvature sliding against a tibial cartilage strip (Fig. 1). The average contact stress reported in Tables 1 and 2 was calculated from the contact area measured with pressure-sensitive film prior to the start of the wear test. For G-groups, little variability was observed in the contact area during preliminary measurements, thus only a mean value of 0.5 MPa is reported in Table 1. The choices of load magnitude (L1) and testing durations (8 h or 24 h) were motivated by observations from preliminary studies that wear damage could be induced under those conditions.

Table 1.

List of test groups used in Study 1. For all groups, sliding was performed using the specified ‘Counterface’ against a cartilage strip, using the specified ‘Lubricant’ and contact ‘Load’, over the given ‘Duration’. The number of specimens used in each group is given under ‘n’.

Group Counterface Lubricant Duration Load Contact Stress n
G-PBS-8-L1 Glass lens PBS 8 h 4.45 N 0.5 MPa 12
G-SF-8-L1 Glass lens SF 8 h 4.45 N 0.5 MPa 7
G-PBS-24-L1 Glass lens PBS 24 h 4.45 N 0.5 MPa 6
G-SF-24-L1 Glass lens SF 24 h 4.45 N 0.5 MPa 6
C-PBS-8-L1 Cartilage PBS 8 h 4.45 N 0.28±0.05 MPa 7
C-SF-8-L1 Cartilage SF 8 h 4.45 N 0.29±0.08 MPa 6
Open in a new tab

Fig. 1.

Fig. 1.

Open in a new tab

Representative photographs of wear testing configurations: (A) Cartilage femoral condylar plug against tibial cartilage strip (Study 1), (B) glass lens on tibial cartilage strip (Study 1), and (C) cartilage ankle condylar plug against tibial cartilage strip (Study 2).

Table 2.

List of test groups used in Study 2. For all groups, sliding was performed using the specified ‘Counterface’ against a cartilage strip, using the specified ‘Lubricant’ and contact ‘Load’, over the given ‘Duration’. The number of specimens used in each group is given under ‘n’.

Group Counterface Lubricant Duration Load Contact Stress n
C-PBS-24-L25 Cartilage PBS 24 h 111.2 N 2.2±0.2 MPa 11
C-SF-48-L25 Cartilage SF 48 h 111.2 N 2.1±0.2 MPa 8
Open in a new tab

Since gross delamination (by visual examination of the articular surface) was not observed in either of the cartilage-on-cartilage groups of Study 1 (C-PBS-8-L1, C-SF-8-L1), and since preliminary studies showed that increasing the testing duration to 24 h for these load conditions did not lead to gross delamination, Study 2 was performed on a different testing device that could apply a higher, more physiologic contact load. Two groups were included in Study 2, one using PBS and 24 h testing duration and the other using SF and 48 h testing duration (Table 2). In both groups the contact load was 111.2 N and a highly curved ankle condylar cartilage counterface (C) was used (Fig. 1), producing more than a fourfold increase in contact stress magnitudes.

All three hypotheses were tested by pooling the results from Studies 1 and 2, since none of the hypotheses depended on the magnitude of the contact load or stress. Hypothesis H1 was tested by examining whether the friction coefficient for samples that exhibited damage was statistically different from that measured in samples that did not get damaged, looking at the mean value μeff prior to the onset of damage. Hypothesis H2 was tested by statistically comparing the number of cycles to failure N for damaged samples tested in PBS versus SF. Hypothesis H3 was tested by statistically comparing the number of cycles to failure N in glass counterface groups versus cartilage counterface groups.

Experimental Measurements

Specimen preparation, wear testing devices, experimental protocols for measuring contact areas and stresses, friction coefficient, cartilage topography from laser scans, evidence of subsurface delamination and damage from polarized light microscopy (PLM) and histological staining, and detection of the onset of delamination from measurements of the frictional force versus tangential displacement data, are presented in the Supplementary Material section.

Results

Representative plots of the friction coefficient μ versus time for one undamaged and two damaged samples are presented in Fig. 2. Representative photographs, topographical scans, PLM, and histological images are presented in Figs. 3–5. As supported by the plots in Fig. 2, samples that remained undamaged at the completion of the test all exhibited a nearly constant μ throughout the testing duration (Fig. 2a). In some cases, samples that presented gross visual (Fig. 3) or occult PLM/histological (Fig. 4, Fig. 5) evidence of damage exhibited an increase in μ after the onset of damage (Fig. 2b), though this evidence of change in μ was not observed in all cases (Fig. 2c). These findings were observed regardless of counterface.

Fig. 2.

Fig. 2.

Open in a new tab

Representative plots of the friction coefficient μeff as a function of time for (a) undamaged sample, (b) damaged sample with evidence of increased friction, and (c) damaged sample with no evidence of increased friction. Arrows point to time of damage initiation. (Top: glass-on-cartilage; Bottom: cartilage-on-cartilage)

Fig. 3.

Fig. 3.

Open in a new tab

Tibial cartilage strip with visual evidence of gross damage (arrows), away from sample edges: (A) light photography, (B) 3D laser scan.

Fig. 5.

Fig. 5.

Open in a new tab

Histological sections of tibial cartilage strip sliced and stained with Picrosirius Red for collagen. Section (A) shows an undamaged sample with no subsurface delamination. Section (B) shows a sample with clear delamination of the superficial zone from the deep zone resulting in a blister. Section (C) shows a sample with occult delamination of the superficial zone from the middle and deep zones, but no blistering. White arrows point to gaps in collagen indicating delamination. No evidence of abrasive wear is apparent for any of the articular surfaces. *cartilage canals.

Fig. 4.

Fig. 4.

Open in a new tab

Polarized light microscopy depicting damage onset and damage progression in cross-sections taken through the travel paths of six strips halted mid-test at spaced intervals (N+M): (A) no damage (4.45 N, SF, 2800 cycles) (B-F) damage (4.45 N in PBS). Bright regions indicate collagen fiber alignment (SZ, surface zone; DZ, deep zone); dark regions indicate random collagen organization (MZ, middle zone).

PLM sections were obtained from a separate set of six specimens used for validating the identification of the cycle N of damage onset and the severity of damage subsequent to that onset (see Supplementary Material section). Images of these sections (Fig. 4) show that delamination occurred between the superficial and middle zones, where cartilage fibril orientation transitions from a tangential orientation in the superficial zone (bright contrast on PLM), to a more disordered orientation in the middle zone (dark contrast on PLM) and back to an ordered radial orientation in the deep zone (bright contrast).

Visual assessment of gross surface damage (Fig. 3), polarized light microscopy (Fig. 4) and/or histological staining (Fig. 5) for each tested sample confirmed that damage occurred in the form of subsurface delamination in all cases, with little evidence of abrasive wear from examination of the integrity of the articular surface. In 15 samples for which there was no gross visual evidence of damage, our algorithm for detecting onset of damage, which examined raw frictional force versus tangential displacement data (see Supplementary Material), indicated that damage could have occurred. In these cases, we performed histology to verify evidence of failure (true positives, n=10), or lack thereof (false positives, n=5). For true positives, delamination similarly occurred at the transition of the superficial and middle zones, as seen in the representative histological sections of Fig. 5.

Subsurface delamination damage occurred in a total of 47 of the 63 samples in Studies 1 and 2 (32 of 44 damaged in Study 1; 15 of 19 damaged in Study 2). The mean value of the friction coefficient of undamaged samples over all loading cycles was μeff = 0.021 ± 0.017 whereas that of damaged samples (calculated prior to damage onset) was μeff = 0.024 ± 0.017. A Shapiro-Wilk test of normality showed that neither of these two groups of μeff values exhibited normal distributions (p<0.01), though the undamaged group followed a log-normal distribution (p=0.15). We performed a two-tailed Student t-test with equal variance (as verified from a Fisher ratio test, p=0.42) on ln μeff and found no statistical difference between the undamaged and damaged samples (p=0.46), in support of hypothesis H1 that delamination occurred even when the friction coefficient remained low.

For all 47 damaged samples in Studies 1 and 2, the number of cycles to failure N varied from 140 to 12560. A Shapiro-Wilk test of normality demonstrated that N exhibited a log-normal distribution (p=0.62). A one-tailed Student t-test with equal variance (validated with a Fisher ratio test, p=0.48) performed on ln N between damaged samples tested in PBS (n=28) versus SF (n=19) showed that there was no significant difference between ln N in SF or PBS (ln N = 6.70 ± 0.99 in PBS and ln N = 7.11 ± 1.15 in SF; p=0.10). This result does not support hypothesis H2 that SF delayed the onset of cartilage delamination when compared to PBS. There was no statistical difference in ln μeff between these two groups (Student t-test, two-tailed, equal variance, p=0.48).

Among the 47 damaged tibial cartilage strips in Studies 1 and 2, there were n=26 tested against glass versus n=21 tested against cartilage. A Fisher ratio test (p=0.18) was used to verify that the log-normal distributions of N for these two groups had equal variances. A one-tailed Student t-test with equal variance demonstrated that ln N = 6.43 ± 0.83 for cartilage-on-glass was significantly lower than ln N = 7.40 ± 1.10 for cartilage-on-cartilage (p<0.001), in support of hypothesis H3 that loading cartilage-on-cartilage in a MCA configuration delayed the onset of cartilage delamination when compared to testing cartilage against glass in the same configuration. The friction coefficient for these cartilage-on-glass samples was μeff = 0.014 ± 0.003, whereas that for cartilage-on-cartilage was μeff = 0.035 ± 0.020. Both of these measures exhibited a normal distribution (Shapiro-Wilk test, p=0.10 for glass, p=0.28 for cartilage). A two-tailed Student t-test with unequal variances (Fisher ratio test p<0.001) showed that μeff for cartilage-on-glass was significantly lower than for cartilage-on-cartilage (p<0.0001).

Discussion

The motivation for this study was to verify that cartilage wear occurs by surface delamination, when adopting a physiologically more realistic testing configuration than our prior study of glass or metal sliding against cylindrical cartilage plugs in PBS (Oungoulian et al., 2015). In this study, either a semi-convex glass lens or an ellipsoidal condylar cartilage counterface was slid against a cartilage strip under average contact stresses ranging from 0.2 to 2.5 MPa, using immature bovine cartilage, with a bath consisting of mature bovine SF or PBS. The contact area did not extend to any of the edges of the articular surface(s), thus precluding any edge effects. Cartilage damage was observed in 47 of 63 samples, showing clear evidence of delamination between the superficial and middle zones (Fig. 4, Fig. 5), with no PLM or histological evidence of abrasive wear on the articular surface. Based on these results, along with our earlier finding of minimal evidence of abrasive wear under harsher testing conditions (Oungoulian et al., 2013; Oungoulian et al., 2015), we can conclude with confidence that the primary mode of wear in immature bovine cartilage is delamination, initiated approximately below the interface between the superficial and middle zones.

In a series of studies on bovine and human articular cartilage, Buckley and co-workers have shown that the cartilage shear modulus varies through the depth of the articular surface, achieving its lowest value at “50–250 μm below the articular surface in a region just below the superficial zone” (Buckley et al., 2010; Buckley et al., 2008). This location coincides with the initiating site of delamination in the current study. They also showed that this location exhibits the greatest amount of energy dissipation during reciprocal tangential loading of the articular surface (Buckley et al., 2013). This combined evidence strongly supports the hypothesis that the delamination observed in the current study occurs due to fatigue failure (a dissipative process) in this transition zone.

In all cases, results of this study showed that the friction coefficient remained low (ranging from 0.005 to 0.083, with a mean value of 0.023), consistent with prior findings in the cartilage lubrication literature under sustained interstitial fluid pressurization (Ateshian, 2009). Results also showed that the friction coefficient was not statistically different in samples that became damaged, compared to those which remained undamaged, implying that wear was not induced by a higher friction coefficient (hypothesis H1). This outcome was consistent with the prior finding that osteoarthritic human cartilage has a comparable friction coefficient to healthier cartilage (Caligaris et al., 2009).

SF was not found to delay the onset of damage compared to PBS (hypothesis H2), as measured by the number of cycles to failure. Similarly, the friction coefficient of samples tested in SF was not statistically different than those tested in PBS. Prior studies that have examined the friction coefficient of cartilage against cartilage, or another material such as glass or steel, in saline versus SF, have usually reported a lower friction coefficient in SF (Schmidt et al., 2007; Schmidt and Sah, 2007), though this effect has not always been strong (Caligaris and Ateshian, 2008; Caligaris et al., 2009; Forster and Fisher, 1996; Forster and Fisher, 1999), nor consistently observed over the range of variation of the friction coefficient (Caligaris and Ateshian, 2008; Gleghorn and Bonassar, 2008). An examination of this prior literature suggests that differences in μeff between SF and PBS are more evident at lower contact stresses (e.g., ≤ 0.5 MPa, (Schmidt and Sah, 2007)) and tend to diminish with increasing values (e.g., up to 4 MPa, (Forster and Fisher, 1996)). The same trend was observed in the current study when examining μeff in Study 1 versus Study 2 (results not shown), though this trend was no longer apparent upon combining all samples for the statistical analyses.

Sliding glass against cartilage led to premature onset of delamination when compared to sliding cartilage against cartilage (hypothesis H3), even though the friction coefficient of glass-on-cartilage was lower than cartilage-on-cartilage, further supporting the hypothesis that the friction coefficient does not mediate the onset of damage. The most likely explanation for earlier onset of damage against glass is that, for a given contact load, glass-on-cartilage produced higher deformations and stresses in the cartilage collagen matrix than cartilage-on-cartilage, due to the much higher stiffness of glass. Based on the experimental tensile fatigue failure studies of Weightman and co-workers on human cartilage (Weightman et al., 1973), the number of cycles to failure decreased with increasing tensile stress. This proposed explanation can be tested using theoretical or computational models of sliding contact of glass versus cartilage, using frictional or idealized frictionless conditions. These types of investigations are currently in progress.

Perhaps the most intriguing result of this study was the finding that the great majority of cartilage samples got damaged in less than 48 hours of continuous frictional loading, under low-to-moderate contact stresses, with one sample failing after only 140 cycles. We believe that this result was particular to our use of immature bovine articular cartilage from young calves, which is known to contain a significant percentage of degraded collagen (Nims et al., 2016; Temple et al., 2006) and is similarly expected to have a significant fraction of immature collagen crosslinks (Gineyts et al., 2010). These conditions may explain the vulnerability of immature bovine cartilage to damage under the testing conditions of this study. Indeed, in a set of preliminary studies using adult human cartilage (mildly osteoarthritic), we were not able to observe any damage using testing conditions employed in our earlier study of sliding glass against a cartilage plug (Oungoulian et al., 2015), not even after continuous loading for one week. This preliminary assessment of the resiliency of adult human cartilage was not surprising, since this avascular tissue, which has limited repair capacity, is typically able to sustain frictional loading for several decades of life prior to the onset of osteoarthritic degeneration. Nevertheless, based on the results of the current study, it is reasonable to hypothesize that delamination wear via subsurface fatigue failure is similarly the primary mechanism of human cartilage wear under normal loading conditions, albeit requiring far more cycles of loading than in immature bovine cartilage. Histological sections of cartilage from osteoarthritic human knee joints provide evidence of delamination in close analogy to results with immature bovine cartilage (Fig. 6), in support of this hypothesis.

Fig. 6.

Fig. 6.

Open in a new tab

Picrosirius Red histology slides for sections through (A) damaged immature bovine tibial cartilage strip with complete removal of delaminated superficial zone; (B) osteoarthritic human distal femoral cartilage layer with evidence of subsurface delamination damage, and (C) osteoarthritic human distal femoral cartilage layer with complete removal of delaminated superficial zone (arrows). Bottom row shows location of sections for all cases. Note the analogy between immature bovine and adult human cartilage in (A) and (C) and between (B) and Fig. 5C.

In summary, this study has demonstrated that wear of immature bovine articular cartilage under low-to-moderate contact stresses occurs in the form of delamination of the extracellular matrix at the interface between the superficial and middle zones. This wear is initiated independently of the magnitude of the friction coefficient or the lubricant. The most likely explanation for wear progression via delamination is fatigue failure of the collagen matrix at its most vulnerable location across the thickness of the articular layer, due to the cyclical tensile stresses produced there by the reciprocal sliding contact motion.

Supplementary Material

2

Acknowledgements

This study was supported with funds from the National Institute of General Medical Sciences of U.S. National Institutes of Health (Award No. R01GM083925), the National Science Foundation Graduate Research Fellowship Program (DGE-11-44155, BKZ), and the Office of the U.S. Assistant Secretary of Defense for Health Affairs and the Defense Health Agency J9, Research and Development Directorate, through the Peer Reviewed Medical Research Program Investigator-Initiated Research Award (Award No. W81XWH-18-10361). Opinions, interpretations, conclusions and recommendations are those of the author and are not necessarily endorsed by the National Institutes of Health, the National Science Foundation, or the Department of Defense. The authors would also like to thank Prof. X. Edward Guo (Columbia University, New York) for providing support for polarized light microscopy measurements.

Footnotes

Conflict of interest statement

We do not have any conflicts of interest with regard to this study and the materials contained herein.

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 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. Ashkavand Z, Malekinejad H, and Vishwanath BS, 2013. The pathophysiology of osteoarthritis. Journal of Pharmacy Research 7, 132–138 [Google Scholar]
  2. Ateshian GA, 2009. The role of interstitial fluid pressurization in articular cartilage lubrication. J Biomech 42, 1163–1176. [DOI] [PMC free article] [PubMed] [Google Scholar]
  3. Ateshian GA, and Mow VC, 2005. Lubrication and wear of diarthrodial joints In: Basic orthopaedic biomechanics & mechano-biology. Mow VC, and Huiskes R, (Eds.). Lippincott Williams & Wilkins, Philadelphia, Pa.; London. [Google Scholar]
  4. Ateshian GA, Wang HQ, and Lai WM, 1998. The role of interstitial fluid pressurization and surface porosities on the boundary friction of articular cartilage. Journal of Tribology-Transactions of the Asme 120, 241–248 [Google Scholar]
  5. Buckley MR, Bergou AJ, Fouchard J, Bonassar LJ, and Cohen I, 2010. High-resolution spatial mapping of shear properties in cartilage. J Biomech 43, 796–800. [DOI] [PMC free article] [PubMed] [Google Scholar]
  6. Buckley MR, Bonassar LJ, and Cohen I, 2013. Localization of viscous behavior and shear energy dissipation in articular cartilage under dynamic shear loading. J Biomech Eng 135, 31002. [DOI] [PMC free article] [PubMed] [Google Scholar]
  7. Buckley MR, Gleghorn JP, Bonassar LJ, and Cohen I, 2008. Mapping the depth dependence of shear properties in articular cartilage. J Biomech 41, 2430–2437 [DOI] [PubMed] [Google Scholar]
  8. Caligaris M, and Ateshian GA, 2008. Effects of sustained interstitial fluid pressurization under migrating contact area, and boundary lubrication by synovial fluid, on cartilage friction. Osteoarthritis Cartilage 16, 1220–1227. [DOI] [PMC free article] [PubMed] [Google Scholar]
  9. Caligaris M, Canal CE, Ahmad CS, Gardner TR, and Ateshian GA, 2009. Investigation of the frictional response of osteoarthritic human tibiofemoral joints and the potential beneficial tribological effect of healthy synovial fluid. Osteoarthritis Cartilage 17, 1327–1332. [DOI] [PMC free article] [PubMed] [Google Scholar]
  10. Dijkgraaf LC, de Bont LG, Boering G, and Liem RS, 1995. The structure, biochemistry, and metabolism of osteoarthritic cartilage: a review of the literature. Journal of oral and maxillofacial surgery 53, 1182–1192 [DOI] [PubMed] [Google Scholar]
  11. Flannery CR, Zollner R, Corcoran C, Jones AR, Root A, Rivera-Bermudez MA, Blanchet T, Gleghorn JP, Bonassar LJ, Bendele AM, Morris EA, and Glasson SS, 2009. Prevention of cartilage degeneration in a rat model of osteoarthritis by intraarticular treatment with recombinant lubricin. Arthritis Rheum 60, 840–847 [DOI] [PubMed] [Google Scholar]
  12. Forster H, and Fisher J, 1996. The influence of loading time and lubricant on the friction of articular cartilage. Proc Inst Mech Eng H 210, 109–119 [DOI] [PubMed] [Google Scholar]
  13. Forster H, and Fisher J, 1999. The influence of continuous sliding and subsequent surface wear on the friction of articular cartilage. Proc Inst Mech Eng H 213, 329–345 [DOI] [PubMed] [Google Scholar]
  14. Gineyts E, Borel O, Chapurlat R, and Garnero P, 2010. Quantification of immature and mature collagen crosslinks by liquid chromatography-electrospray ionization mass spectrometry in connective tissues. J Chromatogr B Analyt Technol Biomed Life Sci 878, 1449–1454 [DOI] [PubMed] [Google Scholar]
  15. Gleghorn JP, and Bonassar LJ, 2008. Lubrication mode analysis of articular cartilage using Stribeck surfaces. J Biomech 41, 1910–1918 [DOI] [PubMed] [Google Scholar]
  16. Han G, and Eriten M, 2018. Effect of relaxation-dependent adhesion on pre-sliding response of cartilage. R Soc Open Sci 5, 172051. [DOI] [PMC free article] [PubMed] [Google Scholar]
  17. Jay GD, Torres JR, Rhee DK, Helminen HJ, Hytinnen MM, Cha CJ, Elsaid K, Kim KS, Cui Y, and Warman ML, 2007. Association between friction and wear in diarthrodial joints lacking lubricin. Arthritis Rheum 56, 3662–3669. [DOI] [PMC free article] [PubMed] [Google Scholar]
  18. Kaplan JT, Neu CP, Drissi H, Emery NC, and Pierce DM, 2017. Cyclic loading of human articular cartilage: The transition from compaction to fatigue. J Mech Behav Biomed Mater 65, 734–742 [DOI] [PubMed] [Google Scholar]
  19. Krishnan R, Caligaris M, Mauck RL, Hung CT, Costa KD, and Ateshian GA, 2004. Removal of the superficial zone of bovine articular cartilage does not increase its frictional coefficient. Osteoarthritis Cartilage 12, 947–955. [DOI] [PMC free article] [PubMed] [Google Scholar]
  20. Lee DW, Banquy X, and Israelachvili JN, 2013. Stick-slip friction and wear of articular joints. Proc Natl Acad Sci U S A 110, E567–574. [DOI] [PMC free article] [PubMed] [Google Scholar]
  21. Loeser RF, 2013. Osteoarthritis year in review 2013: biology. Osteoarthritis and Cartilage 21, 1436–1442 [DOI] [PMC free article] [PubMed] [Google Scholar]
  22. McCutchen CW, 1962. The frictional properties of animal joints. Wear 5, 1–17 [Google Scholar]
  23. Meachim G, 1982. Age-related degeneration of patellar articular cartilage. J Anat 134, 365–371. [PMC free article] [PubMed] [Google Scholar]
  24. Moore DF, 1975. Principles and applications of tribology. Pergamon Press, Oxford, New York,. [Google Scholar]
  25. Nims RJ, Durney KM, Cigan AD, Dusseaux A, Hung CT, and Ateshian GA, 2016. Continuum theory of fibrous tissue damage mechanics using bond kinetics: application to cartilage tissue engineering. Interface Focus 6, 20150063. [DOI] [PMC free article] [PubMed] [Google Scholar]
  26. Northwood E, Fisher J, and Kowalski R, 2007. Investigation of the friction and surface degradation of innovative chondroplasty materials against articular cartilage. Proc Inst Mech Eng H 221, 263–279 [DOI] [PubMed] [Google Scholar]
  27. Oungoulian SR, Chang S, Bortz O, Hehir KE, Zhu K, Willis CE, Hung CT, and Ateshian GA, 2013. Articular cartilage wear characterization with a particle sizing and counting analyzer. J Biomech Eng 135, 024501. [DOI] [PMC free article] [PubMed] [Google Scholar]
  28. Oungoulian SR, Durney KM, Jones BK, Ahmad CS, Hung CT, and Ateshian GA, 2015. Wear and damage of articular cartilage with friction against orthopedic implant materials. J Biomech 48, 1957–1964. [DOI] [PMC free article] [PubMed] [Google Scholar]
  29. Pritzker KP, Gay S, Jimenez SA, Ostergaard K, Pelletier JP, Revell PA, Salter D, and van den Berg WB, 2006. Osteoarthritis cartilage histopathology: grading and staging. Osteoarthritis Cartilage 14, 13–29 [DOI] [PubMed] [Google Scholar]
  30. Schechtman H, and Bader DL, 1997. In vitro fatigue of human tendons. J Biomech 30, 829–835 [DOI] [PubMed] [Google Scholar]
  31. Schmidt TA, Gastelum NS, Nguyen QT, Schumacher BL, and Sah RL, 2007. Boundary lubrication of articular cartilage: role of synovial fluid constituents. Arthritis Rheum 56, 882–891 [DOI] [PubMed] [Google Scholar]
  32. Schmidt TA, and Sah RL, 2007. Effect of synovial fluid on boundary lubrication of articular cartilage. Osteoarthritis Cartilage 15, 35–47 [DOI] [PubMed] [Google Scholar]
  33. Soltz MA, Basalo IM, and Ateshian GA, 2003. Hydrostatic pressurization and depletion of trapped lubricant pool during creep contact of a rippled indenter against a biphasic articular cartilage layer. J Biomech Eng 125, 585–593. [DOI] [PMC free article] [PubMed] [Google Scholar]
  34. Temple MM, Xue Y, Chen MQ, and Sah RL, 2006. Interleukin-1alpha induction of tensile weakening associated with collagen degradation in bovine articular cartilage. Arthritis Rheum 54, 3267–3276 [DOI] [PubMed] [Google Scholar]
  35. Vazquez KJ, Andreae JT, and Henak CR, 2019. Cartilage-on-cartilage cyclic loading induces mechanical and structural damage. J Mech Behav Biomed Mater 98, 262–267. [DOI] [PMC free article] [PubMed] [Google Scholar]
  36. Weightman B, 1976. Tensile fatigue of human articular cartilage. J Biomech 9, 193–200 [DOI] [PubMed] [Google Scholar]
  37. Weightman B, Chappell DJ, and Jenkins EA, 1978. A second study of tensile fatigue properties of human articular cartilage. Ann Rheum Dis 37, 58–63. [DOI] [PMC free article] [PubMed] [Google Scholar]
  38. Weightman BO, Freeman MA, and Swanson SA, 1973. Fatigue of articular cartilage. Nature 244, 303–304 [DOI] [PubMed] [Google Scholar]

Associated Data

This section collects any data citations, data availability statements, or supplementary materials included in this article.

Supplementary Materials

2
NIHMS1599037-supplement-2.pdf (2.4MB, pdf)

RESOURCES