EXPERIMENTAL AND NUMERICAL
INVESTIGATION OF STRAIN-RATE-
DEPENDENT BEHAVIOUR OF KANGAROO
SHOULDER CARTILAGE
Noyel Deegayu Namal Bandara Thibbotuwawa
B. Sc. Eng. (Hons), M. Phil.
Submitted in fulfilment of the requirements for the degree of
Doctor of Philosophy
School of Chemistry, Physics and Mechanical Engineering
Science and Engineering Faculty
Queensland University of Technology
2016
Keywords
Articular cartilage
Kangaroo
Shoulder cartilage
Articular cartilage biomechanics
Strain-rate-dependent behavior
Indentation testing
Porohyperelastic model
Hyperelastic coefficients
Permeability
Pore size
Strain-rate-dependent permeability
Superficial collagen
Proteoglycans
Tissue adaptation
Proteoglycan distribution
Collagen Structure
Experimental and Numerical Investigation of Strain-rate-dependent Behaviour of Kangaroo Shoulder Cartilage i
Abstract
The smooth functioning of the human shoulder joint is predominantly facilitated by a
thin articular layer called the shoulder cartilage placed between the humerus and
scapula bones. The responses of the shoulder joint to external forces are influenced
by the characteristics of shoulder cartilage, which facilitates the frictionless
movement of the shoulder joint and the distribution of load through a large contact
area to protect bone-ends from high contact stresses. Perhaps due to the relatively
low incidence of shoulder osteoarthritis in the past, only a handful of studies have
focused on identifying the characteristics and behaviour of human shoulder cartilage.
In particular, the dynamic characteristics of shoulder cartilage that are most likely
linked to osteoarthritis development [1, 2] have not been investigated. However, with
the reported increase in the incidence of shoulder osteoarthritis [3, 4], it is crucial to
investigate the behaviour of shoulder cartilage in order to identify the reasons behind
osteoarthritis, to develop better diagnostic strategies, and to engineer joint-specific
cartilage tissues.
Due to the unavailability of human shoulder cartilage samples and ethical
restrictions, kangaroo shoulder cartilage was chosen as the animal model for this
research, considering its anatomical and biomechanical similarities to that of a
human shoulder joint. In addition, the bipedal hopping locomotion of kangaroos
results in their shoulder joint being less loaded than their lower limbs, and this
provides an ideal source for investigating the effect of external loading on cartilage
tissue adaptation and its influence on the tissue’s functional behaviour. Therefore, to
investigate the strain-rate-dependent behaviour of shoulder cartilage tissues in the
present study, comprehensive indentation experiments (ranging from physiologically
ii Experimental and Numerical Investigation of Strain-rate-dependent Behaviour of Kangaroo Shoulder Cartilage
low (10-4/s) to high (10-2/s) strain-rates) were conducted on kangaroo shoulder
cartilage tissues. A porohyperelastic numerical model with a newly introduced strain-
rate-dependent permeability function was developed to understand the mechanisms
underlying strain-rate-dependent behaviour. Comparison of the porohyperelastic
model with constant, strain-dependent and strain-rate-dependent permeability models
indicated that the rate-dependent fluid flow significantly affects the behaviour of
shoulder cartilage, especially at high strain-rates. Based on the results of this
investigation it was concluded that, in addition to solid–interstitial fluid frictional-
interactions, pressure drag forces and inertia forces also begin to affect the tissue
behaviour at high strain-rates. This assists the tissue to retain fluid and act as a
protective mechanism that reduces excessive deformation of the cartilage at large
strain-rates. The results of sequential enzymatic degradation and indentation tests
indicated that proteoglycan and collagen degradation significantly compromise the
strain-rate-dependent behaviour and that superficial collagen plays a more significant
role than proteoglycans in facilitating the strain-rate-dependent behaviour of
kangaroo shoulder cartilage. Contrary to the results in studies on knee cartilage, the
superficial collagen was found to equally contribute to shoulder cartilage behaviour
at all strain-rates, thus affirming its significance in the mechanical behaviour of
shoulder cartilage.
The comparative compositional, microstructural and biomechanical
experiments showed that the proteoglycan distribution with depth in shoulder and
knee cartilage is different. A distinct, large deep zone was observed in knee cartilage,
while the size of the deep zone in shoulder cartilage was relatively small. The
superficial collagen was identified as the most significant feature in the collagen
network of shoulder cartilage tissue. Contrary to shoulder cartilage, proteoglycans
Experimental and Numerical Investigation of Strain-rate-dependent Behaviour of Kangaroo Shoulder Cartilage iii
dominated the behaviour of knee cartilage and had a considerably large effect on the
tissue behaviour at low strain-rates while the effect of superficial collagen increased
with an increase in strain-rate. The results indicated that the proteoglycan distribution
and the structural features of the collagen network adapt to external mechanical
stimuli, and hence depend on the local mechanical environment experienced by the
tissue.
Through systematic and in-depth investigation, this study has explored the
mechanisms underlying the strain-rate-dependent behaviour of kangaroo shoulder
cartilage. The findings of this study will help to bridge the existing gaps in
knowledge on the mechanical behaviours of shoulder cartilage tissues and on the
compositional and microstructural adaptation of cartilage tissues to external
mechanical loading. The findings will also help to inform the cartilage modelling
community and tissue engineers about the underlying mechanisms and extracellular
matrix features that need to be considered when modelling and engineering shoulder
cartilage tissues. The experimental strategies employed and the computational model
developed are useful for future studies on cartilage biomechanics, and will inspire
future research investigations to be carried out on shoulder cartilage tissue—an area
that has hitherto lacked attention.
iv Experimental and Numerical Investigation of Strain-rate-dependent Behaviour of Kangaroo Shoulder Cartilage
List of Publications
Journal articles
• Thibbotuwawa, N., Oloyede, A., Senadeera, W., Li, T., Gu, Y., 2015.
Investigation of the mechanical behavior of kangaroo humeral head cartilage
tissue by a porohyperelastic model based on the strain-rate-dependent
permeability. J Mech Behav Biomed Mater 51, 248-259.
• Thibbotuwawa, N., Oloyede, A., Li, T., Singh, S., Senadeera, W., Gu, Y.,
2015. Physical mechanisms underlying the strain-rate-dependent mechanical
behavior of kangaroo shoulder cartilage. Applied Physics Letters 107,
103701.
• Thibbotuwawa, N., Oloyede, A., Li, T., Singh, S., Senadeera, W., Gu, Y.,
2015. Compositional, Microstructural and Biomechanical differences
between kangaroo shoudler and knee cartilage: Implication on numerical
modelling and tissue engineering stratagies (In preparation).
Refereed conference proceedings and extended abstracts
• Thibbotuwawa, N., Gu, Y.T., Oloyede, A., Senadeera, W., Li, T., 2012.
Finite element shoulder models, In Proceedings of 4th International
Conference on Computational Methods (ICCM 2012).
• Thibbotuwawa, N., Li, T., Gu, Y.T., 2014. Porohyperelastic finite element
model for the kangaroo humeral head cartilage based on experimental study
and the consolidation theory, In Proceedings of 5th International Conference
on Computational Methods (ICCM 2014).
• Thibbotuwawa, N., Oloyede, A., Senadeera, W., Gu, Y.T., 2014.
Hyperelastic Constitutive Relationship for the Strain-Rate Dependent
Behavior of Shoulder and Other Joint Cartilages, The 15th International
Conference on Biomedical Engineering. Springer, pp. 255-258.
Experimental and Numerical Investigation of Strain-rate-dependent Behaviour of Kangaroo Shoulder Cartilage v
• Thibbotuwawa, N., Oloyede, A., Senadeera, W., Gu, Y.T.,2014. Exploration
of the biomechanical load bearing mechanisms of articular cartilage under
dynamic loading, Proceedings of the 9th Australasian Biomechanics
conference.
vi Experimental and Numerical Investigation of Strain-rate-dependent Behaviour of Kangaroo Shoulder Cartilage
Table of Contents
Keywords ................................................................................................................................................. i
Abstract .................................................................................................................................................. ii
List of Publications ................................................................................................................................. v
Table of Contents ................................................................................................................................. vii
List of Figures ......................................................................................................................................... x
List of Tables ....................................................................................................................................... xiv
Ethical Clearance for Tissue Use .......................................................................................................... xv
Statement of Original Authorship ........................................................................................................ xvi
Acknowledgement ............................................................................................................................. xviii
CHAPTER 1: INTRODUCTION ....................................................................................................... 1 1.1 Background .................................................................................................................................. 1
1.2 Research problem and questions .................................................................................................. 3
1.3 Research aims and objectives ...................................................................................................... 4
1.4 Research Significance, contribution and scope ............................................................................ 6
1.5 Thesis Outline .............................................................................................................................. 8
1.6 Research framework .................................................................................................................. 10
CHAPTER 2: LITERATURE REVIEW ......................................................................................... 13
2.1 Importance of shoulder joint and shoulder cartilage .................................................................. 13 2.1.1 Shoulder osteoarthritis and causes .................................................................................. 15 2.1.2 Adaptation of cartilage tissues in response to the mechanical environment ................... 18
2.2 Articular cartilage ...................................................................................................................... 20 2.2.1 Articular cartilage proteoglycan ..................................................................................... 21 2.2.2 Articular cartilage collagen ............................................................................................. 22 2.2.3 Articular cartilage structure ............................................................................................ 23 2.2.4 Articular cartilage collagen network architecture ........................................................... 26 2.2.5 Articular cartilage load-bearing unit: Proteoglycan and collagen entrapment ................ 27
2.3 Articular cartilage biomechanics: static and dynamic load-bearing mechanisms ...................... 29 2.3.1 Articular cartilage biomechanics: The structure–function relationship .......................... 34
2.4 Characteristics of shoulder cartilage .......................................................................................... 35
2.5 Biomechanical models of articular cartilage .............................................................................. 37
2.6 Summary and Implications ........................................................................................................ 40
CHAPTER 3: RESEARCH DESIGN AND METHODOLOGY .................................................... 43
3.1 experimental animal model for shoulder cartilage ..................................................................... 43
3.2 Experimental methodologies and materials ............................................................................... 45 3.2.1 Tissue harvesting and preparation .................................................................................. 45 3.2.2 Evaluation of potential thickness measurement methods ............................................... 46 3.2.3 Ultrasound speed in kangaroo shoulder cartilage tissues and thickness
measurement ................................................................................................................... 47 3.2.4 Biomechanical characterisation: Mechanical tests performed on articular
cartilage .......................................................................................................................... 50 3.2.5 Critical evaluation of confined, unconfined and indentation mechanical tests ............... 51 3.2.6 Mechanical testing protocol ............................................................................................ 53
Experimental and Numerical Investigation of Strain-rate-dependent Behaviour of Kangaroo Shoulder Cartilage vii
3.3 Numerical modelling methodology ........................................................................................... 56 3.3.1 Numerical modelling to investigate the physical mechanisms underlying the
mechanical behaviour of cartilage: Initial model development ...................................... 56 3.3.2 Assessing the suitability of the porohyperelastic model for investigating the solid
and fluid behaviour of cartilage tissues: The preliminary porohyperelastic FE model .............................................................................................................................. 64
3.3.3 Development of force–indentation relationship for the 2-term reduced hyperelastic model .......................................................................................................... 70
CHAPTER 4: EFFECT OF INTERSTITIAL FLUID ON THE STRAIN-RATE-DEPENDENT BEHAVIOUR OF KANGAROO SHOULDER CARTILAGE ...................................................... 75 4.1 Introduction ............................................................................................................................... 75
4.2 Aims and objectives ................................................................................................................... 76
4.3 Hypotheses ................................................................................................................................. 76
4.4 Strain-rate-dependent mechanical behavioUr of kangaroo shoulder cartilage ........................... 76 4.4.1 Tissue stiffness: Piecewise linear regression method ..................................................... 77 4.4.2 Stiffness variation with strain and strain-rate ................................................................. 78 4.4.3 Solid–fluid interaction and its effect on the strain-rate-dependent behaviour ................ 80
4.5 Porohyperelastic field theory for soft biological tissues ............................................................ 81 4.5.1 Porohyperelastic FE model development for indentation test ........................................ 84 4.5.2 Permeability variation with strain-rate ........................................................................... 87
4.6 Extension OF Porohyperelastic Field theory: Strain-rate-dependent permeability function ..... 88 4.6.1 Material parameter identification ................................................................................... 92
4.7 Results and Discussion .............................................................................................................. 92 4.7.1 Biomechanical parameters of kangaroo shoulder cartilage ............................................ 92 4.7.2 Comparison of constant, strain-dependent and strain-rate-dependent model
predictions ...................................................................................................................... 94 4.7.3 Effects of strain-dependent and strain-rate-dependent permeability .............................. 95 4.7.4 Mechanisms underlying the strain-rate-dependent tissue behaviour .............................. 97 4.7.5 Role of cartilage as a protective layer at large strain-rates ............................................. 99 4.7.6 Limitations of the strain-rate-dependent permeability model and possible
improvements to the FE porohyperelastic model ......................................................... 101
4.8 Conclusion and Remarks ......................................................................................................... 103
CHAPTER 5: EFFECT OF PROTEOGLYCAN AND SUPERFICIAL COLLAGEN ON THE STRAIN-RATE-DEPENDENT MECHANICAL BEHAVIOUR OF KANGAROO SHOULDER CARTILAGE ................................................................................................................................... 107 5.1 Introduction ............................................................................................................................. 107
5.2 Aims and Objectives ................................................................................................................ 109
5.3 Hypotheses ............................................................................................................................... 110
5.4 Experimental Methodology ..................................................................................................... 110 5.4.1 Assessment of tissue preservation methods: The PBS-solution at 4 °C vs the
multiple freeze–thaw method ....................................................................................... 111 5.4.2 Proteoglycan, superficial collagen degradation and surface delipidisation .................. 114 5.4.3 Statistical data analysis procedure ................................................................................ 122
5.5 Results and discussion ............................................................................................................. 122 5.5.1 Effect of proteoglycan degradation on strain-rate-dependent behaviour and
mechanical properties ................................................................................................... 124 5.5.2 Effect of superficial collagen degradation on strain-rate-dependent behaviour and
mechanical properties ................................................................................................... 127 5.5.3 Effect of surface phospholipid removal on strain-rate-dependent behaviour and
mechanical properties ................................................................................................... 131 5.5.4 Comparison of the effect of proteoglycan and superficial collagen on strain-rate-
dependent behaviour ..................................................................................................... 133
viiiExperimental and Numerical Investigation of Strain-rate-dependent Behaviour of Kangaroo Shoulder Cartilage
5.5.5 Effect of proteoglycan and superficial collagen degradation on long-term functional load-bearing ability of the tissue .................................................................. 137
5.6 Conclusion and remarks ........................................................................................................... 139
CHAPTER 6: COMPOSITIONAL, MICROSTRUCTURAL AND BIOMECHANICAL DIFFERENCES IN KANGAROO SHOULDER AND KNEE CARTILAGE ............................ 143
6.1 Introduction .............................................................................................................................. 143
6.2 Aims and objectives ................................................................................................................. 144
6.3 Hypothesis ............................................................................................................................... 145
6.4 Methods and materials ............................................................................................................. 145 6.4.1 Cryostat tissue sectioning: Tissue preparation for histological studies ......................... 145 6.4.2 Safranin-O staining protocol ......................................................................................... 146 6.4.3 Proteoglycan quantification: Optical absorbance measurements .................................. 147 6.4.4 Sample preparation for PLM measurements ................................................................. 148 6.4.5 Collagen quantification: PLM measurements ............................................................... 150
6.5 Results and discussion ............................................................................................................. 152 6.5.1 Differences in proteoglycan concentration with depth in knee and shoulder
cartilage ........................................................................................................................ 152 6.5.2 Differences in collagen network of knee and shoulder cartilage .................................. 154 6.5.3 Comparison of strain-rate-dependent mechanical behaviour and biomechanical
properties of knee and shoulder cartilage ..................................................................... 160 6.5.4 Significance and implications for numerical modelling and tissue engineering ........... 166
6.6 Conclusion and remarks ........................................................................................................... 168
CHAPTER 7: CONCLUSIONS ...................................................................................................... 171
7.1 Conclusions .............................................................................................................................. 171
7.2 Limitations ............................................................................................................................... 175
7.3 Future Research Directions ...................................................................................................... 175
BIBLIOGRAPHY ............................................................................................................................. 179
Experimental and Numerical Investigation of Strain-rate-dependent Behaviour of Kangaroo Shoulder Cartilage ix
List of Figures
Figure 1.1: Research framework ........................................................................................................... 11
Figure 2.1: Parts of the shoulder joint showing the humeral head, scapula, shoulder cartilage, and muscle and tendon structure of the rotator cuff [27]. .................................................... 14
Figure 2.2: Articular cartilage proteoglycan aggregate ......................................................................... 21
Figure 2.3: Collagen triple helix structure and assembly [110]. ........................................................... 23
Figure 2.4: Zonal variation of matrix components in articular cartilage [117] ..................................... 24
Figure 2.5: Complex fibre and proteoglycan molecule arrangement; the proteoglycans are trapped in three dimensional collagen fibre network [125] .................................................. 26
Figure 2.6: (a) Idealised fibre arrangement in articular cartilage – Benninghoff structure; (b) Typical segmental obliquity in collagen fibres observed; (c) Cross-linking and interlocking of collagen fibres facilitated by the segmental obliquity [129] ........................ 27
Figure 2.7 Three-dimensional network of fibres in articular cartilage generated from repeated cross-linking of string elements; (b) The balloons representing proteoglycans packed and constrained in the fibre network; (c) The complete balloon-string model showing upper tension diaphragm of the surface layer and fibre-proteoglycan arrangement [130]. ............................................................................................................... 28
Figure 2.8: The load-bearing functional unit of articular cartilage; the confining forces of the collagen network resist the osmotic swelling pressure of the proteoglycan and give rise to the intrinsic stiffness of the cartilage matrix [134] .................................................... 29
Figure 2.9: (a) Mechanical analogy illustrating the stress-sharing stress-transfer consolidation mechanism; (b) Pore pressure variation, when the articular cartilage is statically loaded, illustrating how the fluid supports the external load [135]. ..................................... 30
Figure 2.10: Extended rheological analogy of cartilage matrix featuring σa, applied stress; kv stiffness of instantaneous spring; di (i=0,1,2,…), length of ith dent; ki (i=1,2,…), stiffness of ith spring; D1, instantaneous damping coefficient for unbound fluid; D2, damping coefficient relating to bound fluid; Di, general damping coefficient of unbound fluid in the dashpot; σeff, solid’s effective stress; Q, permeability coefficient of the matrix; O, active osmotic component [151]. .............................................................. 32
Figure 3.1: (a) 8 mm diameter cartilage sample; (b) Specimen-harvested region (near the central area of the humeral head); (c) Bone was constrained using a stainless steel holder and submerged in physiological (0.15 M) saline solution; (d) Indentation testing on the sample ............................................................................................................ 46
Figure 3.2: (a) Needle probe measurement location in the sample; (b) A typical force–indentation curve during needle probe indentation – The curve is characterised by articular cartilage (AC) surface puncture due to piercing of the cartilage surface; the initial gradient of the curve significantly increases when the cartilage-calcified bone surface is reached ................................................................................................................. 49
Figure 3.3: Plot between ultrasound travel time and kangaroo shoulder cartilage thickness from needle probe measurements; the slope of the curve is the average ultrasound speed in the tissue ........................................................................................................................... 50
Figure 3.4: (a) The model geometry, mesh, boundary condition and loading configuration of the cartilage–bone FE model; (b) Simplified FE model (without the bone) with mesh and boundary conditions – In this model, the bone is replaced by a rigid constraint (indicated by the red line) which restricts the displacement of the bottom plane of the cartilage .......................................................................................................................... 61
Figure 3.5: (a) Stress distribution for cartilage on rigid bone indented to 10% strain; (b) Stress distribution for cartilage on rigid constraint indented to 10% strain; (c) Comparison
x Experimental and Numerical Investigation of Strain-rate-dependent Behaviour of Kangaroo Shoulder Cartilage
of the numerical model’s result (force on indenter) with the theoretical prediction of Eq. (3.2) and mesh sensitivity data; (d) Comparison of elastic cartilage on rigid bone laminate model results and elastic cartilage-rigid constraint model results with theoretical model results ....................................................................................................... 62
Figure 3.6: (a) Cartilage, indenter geometry (3 mm diameter with 0.1 mm fillet radius at the edge), the mesh, boundary condition and loading configuration based on mechanical testing carried on kangaroo shoulder cartilage samples; (b) Numerical result of elastic cartilage samples indented up to 30% engineering strain; (c) Variation of force on indenter based on mesh element number................................................................ 63
Figure 3.7: Boundary conditions employed in preliminary porohyperelastic FE model ....................... 66
Figure 3.8: Solid skeleton effective stress-strain curve fitted with a piecewise linear curve to extract the solid skeleton material parameters ...................................................................... 68
Figure 3.9: (a) Pore pressure measurements compared with FE model predictions; (b) Creep strain measurements compared with FE model predictions ................................................. 69
Figure 3.10: Mesh sensitivity analysis for (a) pore pressure predictions (b) creep strain prediction .............................................................................................................................. 70
Figure 3.11: Variation of correction factors k1, k2 with sample thickness ............................................ 74
Figure 4.1: Piecewise linear curve fit to nominal stress-nominal strain data ........................................ 77
Figure 4.2 (a) Strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage indicated by nominal stress-strain data; (b) Stiffness variation with strain and strain-rate – Stiffness was calculated by force divided by the indentation area and by displacement divided by the cartilage’s original thickness .................................................. 78
Figure 4.3: (a) Experimental data from 10-2/s of a representative sample fitted to neo-Hookean, Mooney–rivlin and 2-term reduced polynomial incompressible hyperelastic functions; (b) R-squared values indicating the goodness of fit of neo-Hookean, Mooney–rivlin and 2-term reduced polynomial incompressible hyperelastic functions to the experimental data ........................................................................................ 86
Figure 4.4: An exponential function (Eq. 4.19) fitted to Lai and Mow’s (1980) data; (b) Variation of coefficient M with coefficient a is approximated as a second-order polynomial function; (c) Variation of coefficient a with pressure difference (P) approximated as a power function ........................................................................................ 90
Figure 4.5: (a) Re-analysis of Oloyede and Broom’s [149] Variation of pressure difference (P) between the inside and outside of the tissue with strain-rate; (b) Variation of permeability with strain-rate as predicted by Eq. (4.23) ...................................................... 91
Figure 4.6: Comparison of constant, strain-dependent and strain-rate-dependent model prediction to average (n=10) experimental data of the samples tested – (a) Constant permeability; (b) Strain-dependent permeability; (c) Strain-rate-dependent permeability; (d) Model predictions in terms of R-squared (R2) values and the corresponding significant differences among constant, strain-dependent and strain-rate-dependent models at individual strain-rates .................................................................. 95
Figure 4.7: An ideal representation of part of a tissue with pores represented by circles (undeformed) and ellipse (deformed); (a) Constant permeability – Pore volume/effective fluid-flow area does not change; (b) Strain-dependent permeability – Pore volume/effective flow area is reduced due to application of strain (ε); (c) Strain-rate-dependent permeability – Large pressure differences due to suddenly applied load (Tt2<<<Tt1) result in larger drag forces; this will compact the tissue to reduce the pore size (indicated by the red hatched area), creating congestion for fluid particles to move through pores and, therefore, the fluid particles experience a reduction of pore size/effective flow area ............................................................................ 96
Figure 4.8 : Comparison of pore pressure and velocity profiles at 10-2/s – (a) Strain-dependent permeabilty; (b) Strain-rate-dependent permeability; (c) Fluid velocity at the bottom left (point P) of the cartilage matrix ..................................................................................... 98
Experimental and Numerical Investigation of Strain-rate-dependent Behaviour of Kangaroo Shoulder Cartilage xi
Figure 5.1: (a) Mechanical properties after 72 hrs in PBS-inhibitor solution at 4 °C; (b) Mechanical properties after 1 week in PBS-inhibitor solution at 4 °C ............................... 112
Figure 5.2 : Mechanical property change due to multiple freeze thaw cycles .................................... 113
Figure 5.3: Steps in sequential trypsin treatment (0.05 mg/ml) and mechanical testing on kangaroo shoulder cartilage samples .................................................................................. 115
Figure 5.4: Safranin-O staining of cryosectioned samples harvested from near the central load-bearing area of the humeral head – (a) Untreated sample; (b) 1 hr trypsin-treated sample; (c) 2 hr trypsin-treated sample; (c) 4 hr trypsin-treated sample ............................ 117
Figure 5.5: Steps carried out to investigate the effect of superficial collagen on the strain-rate-dependent behaviour of kangaroo shoulder cartilage ......................................................... 118
Figure 5.6: Safranin-O staining of (a) untreated samples; (b), (c), (d) collagenase-treated samples 1, 2, 3, respectively ............................................................................................... 119
Figure 5.7: (a) Alcian blue 0.1 ml mixed in 1 ml of distilled water (control test); (b) Alcian blue 0.1 ml mixed in the resulting solution after a cartilage sample being digested in 1 ml of 30 U/ml collagenase for 44 hrs; (c) Sample after digesting in collagenase was treated in 1 ml of trypsin–PBS solution (0.05 mg/ml) for 4 hrs and then mixed with 0.1 ml of alcian blue ........................................................................................................... 120
Figure 5.8: Steps carried out to investigate the effect of surface lipids on the mechanical behaviour of kangaroo shoulder cartilage .......................................................................... 121
Figure 5.9 : Effect of mechanical behaviour when (a) normal sample is (b) treated in trypsin (0.05 mg/ml) for 4 hrs, and when (c) normal sample is (d) treated with collagenase (30 U/ml) for 44 hrs, and when (e) normal sample is (f) treated with Folch reagent to remove surface lipids ......................................................................................................... 123
Figure 5.10: Effect of 1 hr, 2 hrs and 4 hrs of trypsin treatment (0.05 mg/ml) on (a) Young’s modulus and (b) nonlinear stiffness parameter of kangaroo shoulder cartilage for 10-
4/s, 5x10-4/s, 5x10-3/s and 10-2/s strain-rates ....................................................................... 125
Figure 5.11 : Effect of 44 hr collagenase treatment (30 U/ml) on (a) Young’s modulus and (b) nonlinear stiffness parameter of kangaroo shoulder cartilage for 10-4/s, 5x10-4/s, 5x10-3/s and 10-2/s strain-rates ............................................................................................ 128
Figure 5.12: Effect of surface lipid removal on (a) Young’s modulus; and (b) the nonlinear stiffness parameter of kangaroo shoulder cartilage for 10-4/s, 5x10-4/s, 5x10-3/s and 10-2/s strain-rates ................................................................................................................ 132
Figure 5.13: Effect on average normalised force–indentation curves due to (a) 4 hrs of trypsin (0.05mg/ml) treatment, i.e. proteoglycan completely removed (n=12); (b) 44 hrs of collagenase treatment, i.e. severe disruption to superficial collagen (n=10); and (c) surface phospholipid removal (n=9) .................................................................................. 135
Figure 5.14: Percentage decrease in (a) Young’s modulus and (b) nonlinear stiffness parameter of kangaroo shoulder cartilage due to complete removal of proteoglycans (4 hrs of treatment in 0.05 mg/ml trypsin) and severe disruption to superficial collagen (44 hrs of treatment in 30 U/ml collagenase) ........................................................................... 137
Figure 5.15: Variation of pore pressure with strain for 4 hr trypsin-treated and 44 hr collagenase-treated samples ............................................................................................... 138
Figure 6.1: Split line directions identified through the pin-prick test performed on (a) femur; (b) tibia; (c) humeral head; and (d) glenoid of kangaroo knee and shoulder joints ............ 150
Figure 6.2: (a) Variation in proteoglycan concentration (indicated by light absorbance by safranin-O) with depth for samples harvested from four locations of knee cartilage (i.e. lateral femur, medial femur, lateral tibia, medial tibia) and from central humeral head in shoulder joint; (b) Area under absorbance curve, i.e. proteoglycan content in LF, MF, LT, MT and H; (c) LF, MF, LT, MT and H cartilage stained by 0.1% safranin-O indicating the proteoglycan variation with depth ............................................. 153
Figure 6.3: Typical images of cartilage when exposed to polarised light – The dark image at 0° angle corresponds to the cross-polarised configuration; in a sequence of every 45°,
xii Experimental and Numerical Investigation of Strain-rate-dependent Behaviour of Kangaroo Shoulder Cartilage
bright images are visible indicating the polarised light has been transmitted through the samples; (b) For 0°,90°, 180° and 270° angles, the light transmittance through the samples is literally uniform; (c) For 45°,135°, 225° and 315° angles, the variation in light transmitted through depth corresponds to the zonal arrangement of cartilage fibres (These results are for typical samples harvested from a medial femur of kangaroo knee cartilage) ................................................................................................ 156
Figure 6.4: (a) PLM images obtained from four locations of kangaroo knee (LF, MF, LT and MT) and from the central humeral head; (b) Depth-dependent light transmittance profiles of the samples (i.e. LF, MF, LT, MT and H) ........................................................ 157
Figure 6.5: (a) Confocal image of a typical cartilage sample harvested from the kangaroo humeral head indicating the region (red arrow) near the calcified bone where fibres are perpendicular to the bone – in the same image, near to the top (black arrow), the transition from perpendicular fibres to a random fibre arrangement can be observed; (b) Confocal image enlarged from top region of image (a) indicating random fibre arrangement; (c) Transition from near perpendicular to random fibre arrangement is shown in this enlarged figure taken from the transitional area of image (a) ...................... 159
Figure 6.6: Normalised force vs indentation graphs for: (a) normal and trypsin-treated (in 0.1 mg/ml for 4 hrs) knee cartilage; (b) normal and trypsin-treated (in 0.05 mg/ml for 4 hrs) shoulder cartilage; (c) the effect of complete removal of proteoglycan (due to 0.1/mg/ml trypsin treatment for 4 hrs) on Young’s modulus of kangaroo knee cartilage; (d) Comparison of percentage decrease in Young’s modulus due to complete removal of proteoglycans for kangaroo knee and shoulder cartilage .................. 163
Figure 6.7: Percentage decrease in Young’s modulus after complete proteoglycan-removed samples were treated for 44 hrs in collagenase; (b) Contribution of superficial collagen to tissue behaviour at four strain-rates (10-4/s, 5x10-4/s, 5x10-3/s and 10-2/s) ....... 165
Experimental and Numerical Investigation of Strain-rate-dependent Behaviour of Kangaroo Shoulder Cartilagexiii
List of Tables
Table 1.1: Aims and objectives ............................................................................................................... 5
Table 2.1: Pathoanatomical changes in shoulder osteoarthritis [37, 38]. .............................................. 16
Table 3.1: Parameters of the model used for mesh, boundary and loading condition validations ........ 59
Table 3.2: Hyperelastic material parameters and permeability values used for the initial porohyperelastic FE model................................................................................................... 68
Table 5.1: Young’s moduli (MPa) of 4hr trypsin-treated and 44 hr collagenase-treated kangaroo shoulder cartilage at four strain-rates ................................................................. 134
Table 5.2: FE model parameters for normal, 4 hr trypsin-treated and 44 hr collagenase-treated samples ............................................................................................................................... 139
xivExperimental and Numerical Investigation of Strain-rate-dependent Behaviour of Kangaroo Shoulder Cartilage
Ethical Clearance for Tissue Use
The Research Ethics Unit of Queensland University of Technology approved
(approval number is 1200000376) the use of kangaroo shoulder cartilage tissues for
the present research study. The approval was later extended to use kangaroo knee
cartilage as well in the study. The shoulder and knee cartilage tissue samples were
obtained from the same source.
Experimental and Numerical Investigation of Strain-rate-dependent Behaviour of Kangaroo Shoulder Cartilage xv
QUT Verified Signature
To my beloved Father, Mother, Sister, Brother and Wife.
Experimental and Numerical Investigation of Strain-rate-dependent Behaviour of Kangaroo Shoulder Cartilagexvii
Acknowledgement
This thesis has been a one of the most challenging, yet an inspiring experience for
me. During this time, I have been fortunate to work with a number of people, whose
involvement has been instrumental for the success of this research study and hence
their sincere contribution truly needs to be acknowledged. It is my pleasure to
convey my heartfelt gratitude to all of them in this humble acknowledgment.
First and foremost I am greatly thankful to my supervisor, Professor YuanTong
Gu, for his constant support, inspiration and guidance provided during the course of
my PhD study in the School of Chemistry, Physics and Mechanical Engineering
(CPME) at Queensland University of Technology (QUT). I am amazed and
motivated by his affectionate nature, patience and courage and truly grateful for all
the productive discussions we had. His critical insight and academic experience have
been a constant inspiration to me. Without his immense contribution of time, support
and funding, I would not have been able to finish this research study successfully.
Words would not do justice to all the advice, guidance and assistance provided by
him and I am forever grateful to him.
I am thankful and appreciate the guidance provided by my associate
supervisors Professor Kunle Oloyede and Dr. Wijitha Senadeera. Professor
Oloyede’s original research in the field of biomechanics of articular cartilage and his
encouragement for philosophical thinking truly helped me to understand the depth
and value of a PhD study. He greatly helped me to mould my own research path
during the course of my PhD experience. His motivating and encouraging words will
always be remembered. I am grateful to Dr. Senadeera for introducing me to a PhD
opportunity in QUT and continuously encouraging me during my course of study.
xviiiExperimental and Numerical Investigation of Strain-rate-dependent Behaviour of Kangaroo Shoulder Cartilage
While envying his simplicity, affection and openness I will always remember him as
a truly remarkable human being. I am grateful for all his support too.
My heartfelt gratitude goes to QUT and its community as a whole for
providing me with generous funding in addition to space and a friendly environment.
The exposure, experience and relationships that I have accumulated during QUT has
not only helped my professional development but has enabled me to face the future
with a clear mindset and direction. Special thank also goes to the Higher Degree
Research (HDR) supporting team at Science and Engineering Faculty of QUT and
CPME school staff for their immense support in terms of taking care of all the
administrative work related to my PhD study and for been their always to guide
whenever there was an administrative problem.
I would also like to thank Dr. Sanjleena Singh, Mrs. Helen O’Conner, Ms.
Melissa Johnston, Mr. Len Wilcox, Dr. Hayley Moody and Dr. Mark Wellard for the
immense support given, to the best of their capacity, especially for my experimental
studies. Also my special heartiest gratitude goes to each and every colleague in the
Laboratory for Advanced Modelling in Simulation in Engineering and Science
(LAMSES) group for their continuous support and encouraging words. Their
constructive advices have helped me remarkably in carrying out my research and
made my stay at QUT smooth. Special thanks go to Haifei Zhang, Chaminda
Karunasena, Tong Li, Trung Dung Nguyen, Hasitha Nayanajith, Yei Wei, Izzat
Thiyahuddin, Suchitra De Silva and Charith Rathnayaka who are some of my past
and present LAMSES colleague, for their friendship. My heartiest thanks go to Dilini
Galpaya for being a sister to me during my stay in QUT and helping a lot with my
experimental studies and I am also thankful to Helen Whittle for helping me to proof
read my thesis. Special thanks also goes to Sri Lankan student community in QUT
Experimental and Numerical Investigation of Strain-rate-dependent Behaviour of Kangaroo Shoulder Cartilagexix
and the people who I play cricket in every Saturday afternoon, who made my life in
QUT exciting.
Lastly, I am forever grateful to my mother, father, brother and sister for their
constant love, understanding and advices that have not only helped me to grow as a
person but have taught me to define ‘success’ in my own terms in life. Finally, I am
forever grateful to my wife, for all the continuous support given throughout my study
period and for filling my life with love and happiness.
xx Experimental and Numerical Investigation of Strain-rate-dependent Behaviour of Kangaroo Shoulder Cartilage
Chapter 1: Introduction
This chapter first provides an introduction to the research background (Section 1.1)
and research problems (Section 1.2). Then it presents the aims and objectives of this
study (Section 1.3), followed by an outline of the research significance, its
contribution to the field of biomechanics of cartilage tissue, and the research scope
(Section 1.4). Additionally, to provide a better understanding of the content covered
in this thesis, a brief outline of each chapter is included (Section 1.5). In the final
section (Section 1.6), the research framework is presented.
1.1 BACKGROUND
The shoulder joint is the most flexible of all major human joints [5-7]. From object
manipulating to typing and sporting activities, the shoulder joint plays a significant
role in performing a variety of physical activities that are essential for daily human
function. The shoulder joint’s smooth function and flexibility are vital for
individuals’ productivity and efficiency in performing daily activities, and thus are
significant in the human endeavour to lead a quality life. Smooth functioning of the
shoulder joint is predominantly facilitated by a thin articular layer called the shoulder
cartilage that is positioned between the humerus and scapula bones The response of
the shoulder joint to external forces is influenced by the characteristics of the
shoulder cartilage which facilitate the frictionless movement of the shoulder joint [6,
8] and the distribution of load through a large contact area so as to protect bone-ends
from high contact stresses [8, 9].
Flexibility of the shoulder joint comes at the expense of its vulnerability to
injuries and disease. Due to repetitive dislocation, overuse or trauma, the
Chapter 1: Introduction 1
functionality of the shoulder joint may be compromised and lead to degeneration of
its soft tissues [6]. The most prevalent among the potential degenerative changes is
known as arthritis—a complex family of degenerative joint conditions of which the
most common is osteoarthritis [10-12]. High contact stresses on shoulder cartilage
due to trauma and dislocation is known to trigger shoulder osteoarthritis [13], which
degrades the constituents of the tissue and leads to the structural and functional
breakdown of the cartilage. Athletes, wheelchair users and construction workers who
often utilise the shoulders for repetitive load-bearing tasks are the most at risk of
shoulder injuries and eventually of shoulder osteoarthritis development [13].
Osteoarthritis is the most common musculoskeletal joint disorder, affecting the
wellbeing of more than 15-20 million individuals worldwide. After the knee and hip,
the shoulder is the third most common osteoarthritis-affected joint, accounting for
3% of all types of osteoarthritis [14]. Total of 20,000–25,000 semi or total shoulder
arthroplasty procedures have been reported to the Australian Orthopaedic
Association National Joint Replacement Registry (AOANJRR) in the period of 2008
to 2013. This is approximately 5–6% of all joint replacements reported to AOANJRR
[15]. Probably due to the relatively low incidence of shoulder osteoarthritis in the
past, few studies have focused on characterising the behaviour of shoulder cartilage.
In particular, the dynamic characteristics of shoulder cartilage, which are most likely
linked to osteoarthritis development [1, 2], have not been investigated until now. Due
to the growing number of shoulder joint replacements and active participation in
sports among youth, there is now a growing interest in investigating the upper
extremity tissues such as the shoulder cartilage [13, 15].
Biological tissues such as shoulder cartilage are complicated materials which
are naturally designed and adapted to the kinematics and dynamics of the human
2 Chapter 1: Introduction
body. Due to the adaptability of biological structures to mechanical stimuli, it can be
hypothesised that shoulder cartilage is different in composition and structure to lower
limb joint cartilage such as knee cartilage. Until now, scarce research has been
carried out to specifically investigate the adaptation of shoulder cartilage tissues to
the mechanical environment and its biomechanical and functional implications for
cartilage behaviour. Further, it is essential that the compositional and structural
adaptation of cartilage tissues to mechanical loading is known in order to formulate
and regulate strategies to engineer cartilage tissues to specific joints such as the
shoulder. The present study therefore specifically focuses on investigating the
mechanical behaviour of shoulder cartilage and its structure–function relationship.
1.2 RESEARCH PROBLEM AND QUESTIONS
In physical activities, such as lifting and throwing, the shoulder cartilage is subjected
to physiologically different strain-rates [1]. Therefore, the shoulder cartilage should
have the ability to undergo controlled deformation in response to these different
external loading conditions, in order to reduce the risk of bone-to-bone contact in the
joint. Solid–fluid interaction is considered to play a significant role in facilitating this
behaviour of shoulder cartilage tissues. To date, however, there are limited
systematic studies investigating factors affecting the strain-rate-dependent
mechanical behaviour of shoulder cartilage tissues specifically. It is crucial to
understand the extent to which solid–fluid interaction facilitates the strain-rate-
dependent behaviour of shoulder cartilage tissues, in order to identify its implications
on the initiation of shoulder osteoarthritis and the development of artificial shoulder
cartilage tissues.
Differences in the contribution of cartilage constituents, namely, proteoglycans
and the collagen network, to the static and dynamic responses of knee cartilage
Chapter 1: Introduction 3
tissues have been reported [16, 17]. As mentioned earlier, since biological tissues
adapt to their mechanical environment, it is expected that proteoglycan composition,
proteoglycan distribution and the structural features of the collagen network in
shoulder cartilage and knee cartilage are different. Considering these differences, it is
plausible that the findings of the studies done on knee cartilages are not directly
applicable to shoulder cartilage. Based on the literature review (Chapter 2), the main
research questions that guided this study are as follows:
1) How does the behaviour of fluid under different loading-rates affect the
mechanical behaviour of shoulder cartilage?
2) How do the cartilage constituents (collagen structure and proteoglycans)
affect the mechanical behaviour of shoulder cartilage under different loading-
rates?
3) Are the composition and microstructure of shoulder cartilage distinctly
different from that of knee cartilage? And what are the implications of these
differences for the functional behaviours of shoulder and knee cartilage?
1.3 RESEARCH AIMS AND OBJECTIVES
The main aim of the present study is to comprehensively investigate the factors
affecting the strain-rate-dependent behaviour of shoulder cartilage. Further, it also
aims to establish the fact that, due to adaptation to different mechanical
environments, cartilage tissues (e.g. knee and shoulder cartilage) will demonstrate
functional, microstructural and composition differences. The aims and objectives of
this study are outlined in Table 1.1.
4 Chapter 1: Introduction
Table 1.1: Aims and objectives Aims Objectives
To investigate the effect of solid–fluid interaction on the strain-rate-dependent mechanical behaviour of shoulder cartilage
• To investigate the significance of the rate-dependent fluid flow on the strain-rate-dependent mechanical behaviour of shoulder cartilage tissue
To investiage the role of cartilage constituents (i.e. proteoglycan and collagen) on the strain-rate-dependent behaviour of shoulder cartilage
• To investigate the role of proteoglycans in the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage
• To investigate the role of superficial collagen and surface phospholipids in the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage
• To assess whether proteoglycans or superficial collagen dominate the mechanical behaviour of kangaroo shoulder cartilage
To investigate the compositional, microstructural and biomechanical differences between shoulder and knee cartilage
• To investigate the differences in proteoglycan concentration and distribution in knee and shoulder cartilage
• To investigate the differences in the collagen network in knee and shoulder cartilage
• To investigate the contribution of proteoglycans and superficial collagen to the strain-rate-dependent behaviour of knee cartilage, and compare it with shoulder cartilage
Chapter 1: Introduction 5
1.4 RESEARCH SIGNIFICANCE, CONTRIBUTION AND SCOPE
Few studies have been conducted to investigate the mechanical behaviours of
shoulder cartilage tissue. This research contributes to the current knowledge by
improving the understanding of the deformation mechanisms underlying the strain-
rate-dependent behaviour of shoulder cartilage tissues. Through systematic
investigation of the factors affecting the strain-rate-dependent behaviour of shoulder
cartilage tissues, the present study informs the cartilage biomechanics and modelling
community about the physical mechanisms that should be considered in the future
modelling of shoulder cartilage tissue. Further, by investigating the effect of the
mechanical environment on depth-dependent proteoglycan concentration and
features of the collagen structure, the present study identifies the factors to be
considered when engineering artificial cartilage for low and high compressive load-
bearing joints such as the shoulder and knee. This has been achieved by analysing
both the load-bearing abilities (at low and high physiological strain-rates) of shoulder
and knee cartilage tissues with respect to the tissue composition and microstructure.
In this study, due to the unavailability of human shoulder cartilage, and due to
anatomical and biomechanical similarities with the human shoulder, kangaroo was
chosen as a model to study shoulder cartilage (as discussed in detail in Chapter 3,
Section 3.1). Through the introduction of kangaroo as a model, this study provides
researchers with a natural source for investigating how the mechanical environment
affects the cartilage tissue composition and structure. Unlike most other animal
models, the kangaroo shoulder joint experiences relatively less force than its knee
joint, similar to the human shoulder joint. Therefore, it is a suitable animal model for
investigating how the mechanical environment and joint differences affect the
structure, composition and function of cartilage tissues. Future research into the
6 Chapter 1: Introduction
development and degradation of upper and lower extremity joint cartilages using
kangaroo as an animal model would reveal new insights into the some of the area
still unclear about development process [18] of cartilage tissues.
Through the results of this research it has been possible to show that the
contribution of the main cartilage constituents (i.e. proteoglycans and the collagen
network) to strain-rate-dependent behaviour is different in the case of shoulder
cartilage and knee cartilage. It was found that the collagen network was the main
contributor to the behaviour of shoulder cartilage, while the proteoglycans were the
main contributor to the behaviour of knee cartilage. The experimental protocols used
in this study were similar to those in the literature [19-22]; however, to the best of the
author’s knowledge, the experimental approach implemented in the present study has
not been carried out before. In previous studies, numerical models have been used to
investigate the contribution of cartilage constituents to strain-rate-dependent
behaviour of knee cartilage [16]. However, the present study has been able to
experimentally illustrate the effect of the constituents on the strain-rate-dependent
behaviour of shoulder cartilage tissues. The findings of the experiments confirm the
results in the literature [16] on high weight-bearing knee cartilage, and also extend
the knowledge on low weight-bearing shoulder cartilage.
In addition, by thoroughly investigating the effect of fluid behavior on the
strain-rate-dependent behaviour of cartilage tissues, this study demonstrates the
importance of rate-dependent fluid flow in the mechanical behaviour of shoulder
cartilage tissues. The proposed strain-rate-dependent permeability function and the
numerical model will enable future researchers to numerically investigate how rate-
dependent fluid behaviour affects the mechanical behaviours of cartilage tissues.
Chapter 1: Introduction 7
1.5 THESIS OUTLINE
Chapter 1 of this thesis provided an introduction to the research background
including the research problem and research questions, followed by the aims and
objectives. Subsequently, an overview of the significance of the study and its
contribution to the field of biomechanics of cartilage tissues was provided.
In order to elaborate the motivation, philosophy, rationale and research design
behind the present study, Chapter 2 presents a review of the literature. This includes
a discussion on the importance of studying the shoulder joint by emphasising its
significant role in carrying out daily activities as well as the social and economic
burdens of shoulder osteoarthritis. The chapter then discusses the limited research on
shoulder cartilage tissues and the potential adaptation of articular cartilage tissues to
the local mechanical environment. Next, the structure of the articular cartilage is
presented with an emphasis on the importance of individual constituents and the
structural integrity of constituents in the long-term performance of cartilage. The
stress-processing mechanisms of the articular cartilage are discussed next by
presenting its functional load-bearing unit. This is followed by an elaboration on the
structure–function relationship of articular cartilage and the available cartilage
biomechanical models. Lastly, the chapter summarize the research gaps identified
through literature review.
Chapter 3 discusses in detail the research design and methodology used in the
study to explore the research problem. In so doing, the rationale behind the selected
animal model, tissue harvesting and preparation method, ultrasound thickness
estimation procedures along with the ultrasound speed determination is explained in
detail. In addition, the rationale behind the selected mechanical testing method and
indentation testing protocol is elaborated in detail. The chapter then discusses the
8 Chapter 1: Introduction
experimental methodology, followed by the numerical methods that elaborate the
model development and validate its suitability for investigating cartilage
biomechanics.
In Chapter 4, the strain-rate-dependent mechanical behaviour of kangaroo
shoulder cartilage is investigated using indentation testing. This chapter also presents
the stiffness variation with strain and strain-rate, followed by a discussion on
possible reasons underlying the observed behaviour. Subsequently, porohyperelastic
field theory is presented and a suitable hyperelastic model for the solid skeleton of
shoulder cartilage tissue is evaluated. Then, the mechanical parameters of kangaroo
shoulder cartilage tissue are compared with the parameters of human shoulder
cartilage reported in the literature. Further, the predictions of the porohyperelastic
cartilage models which includes the existing constant and strain-dependent
permeability models are compared with the experimental results and a new strain-
rate-dependent permeability model is introduced. Lastly, based on the results of the
comparison with the experimental data, the effects of the strain-dependent fluid
behaviour on the strain-rate-dependent behaviour of shoulder cartilage tissues are
discussed in detail.
In Chapter 5, the mechanism underlying strain-rate-dependent behaviour of
shoulder cartilage is further investigated by studying how proteoglycan and
superficial collagen affect the strain-rate-dependent mechanical behaviour of the
tissue. Firstly, tissue preservation methods are evaluated in order to choose a
preservation method for this particular study and then details of the enzymatic
degradation experimental protocols are presented. The effect of proteoglycan and
superficial collagen on the strain-rate-dependent tissue behaviour is discussed next
by evaluating their relative contribution to the mechanical behaviour of the tissue. In
Chapter 1: Introduction 9
addition, the effect of proteoglycan and collagen degradation on the long-term
functional behaviour of the tissue is discussed using numerical modelling.
Chapter 6 discusses the compositional, microstructural and biomechanical
differences between the kangaroo knee and shoulder cartilage. In so doing, firstly,
the sample preparation procedures, staining protocols, and microscopy and image
processing techniques employed are explained in detail. Then, the differences in
proteoglycan distribution and collagen structure in knee and shoulder cartilage are
discussed, followed by a discussion of their strain-rate-dependent behaviours.
Differences in the contribution of cartilage constituents to the mechanical behaviour
of the two tissues are elaborated next. Implications of the results for the numerical
modelling of cartilage and tissue engineering strategies are further discussed at the
end of the chapter.
Finally, in concluding, Chapter 7 presents the main conclusions of the research
conducted in this study, in addition to the study limitations. This chapter also
discusses future research directions in order to further understand the dynamic load-
bearing mechanisms of cartilage tissues.
1.6 RESEARCH FRAMEWORK
The research framework is presented in Figure 1.1.
10 Chapter 1: Introduction
Figure 1.1: Research framework
Introduction
Background Research problem
Research aims & objectives
Significance & contribution
Research framework
Litrature review
Introduction Articular cartilage
Shoulder cartilage
Biomechanical models of
articular cartilage
Summary & implications
Research design and methodology
Animal model for shoulder cartilage
Experimental methodology & materials
Numerical modelling methodology
Research schedule
Effect of fluid behaviour on mechanical behaviour of kangaroo shoulder cartilage Development of strain-rate-dependent permeability model Comparison of porohyperelastic model with constant, strain-
dependent, strain-rate-dependent permeability
Research schedule
Effect of collagen structure and proteoglycans on strain-rate-dependent behaviour of kangaroo shoulder cartilage Sequential constituent degradation and mechanical testing Comparison of the effect of constituent degradation on
mechanical properties
Comparison of kangaroo knee and shoulder cartilage Compositional and microstructural differences Contribution of collagen structure and proteoglycans Implications of mechanical behavioural differences between
knee and shoulder cartilage
Conclusions
Conclusion
Future work
4-6
3
2
1
7
Limitations
Articular cartilage biomechanics
Chapter 1: Introduction 11
Chapter 2: Literature Review
This chapter first explains the importance of investigating the characteristics of
shoulder cartilage (Section 2.1) and then reviews the literature on the articular
cartilage structure, function and its constituents (Section 2.2). Subsequently, the
static and dynamic load-bearing mechanisms of cartilage are reviewed (Section 2.3)
and research studies conducted specifically on shoulder cartilage are summarised
(Section 2.4). The biomechanical models of articular cartilages are thoroughly
reviewed (Section 2.5). Finally, the literature review is summarised and the research
gaps are identified (Section 2.6).
2.1 IMPORTANCE OF SHOULDER JOINT AND SHOULDER CARTILAGE
The shoulder joint is the most mobile joint of all human joints [4, 6]. It has a large
range of motion, more than any other major diarthrodial joint in the human body and
plays a significant role in performing a variety of physical activities essential for
daily human function. Its flexibility, stability and strength are provided by a complex
integration of soft and hard tissues (Figure 2.1). The rotator cuff tendons and
muscles, glenohumeral ligaments, and glenohumeral capsule provide stability and
flexibility to the shoulder joint, while the biceps tendons and muscles are responsible
for assisting the shoulder joint movements [6]. Smooth functioning of the shoulder
joint is predominantly facilitated by a thin articular layer placed between the
humerus and scapula, which is known as shoulder cartilage. The very nature of the
shoulder joint’s response to external forces is influenced by the characteristics of this
cartilage, which helps to distribute contact stresses [23, 24]. This in turn helps to
Chapter 2: Literature Review 13
maintain stability of the shoulder joint by facilitating uniform stress distribution in
the shoulder capsule [6, 25, 26].
Figure 2.1: Parts of the shoulder joint showing the humeral head, scapula, shoulder cartilage, and muscle and tendon structure of the rotator cuff [27].
Healthy long-term function of the shoulder cartilage can be directly or indirectly
influenced by the amount and frequency of the forces it experiences and the proper
function of the shoulder joint’s elements such as the labrum and rotator cuff muscles
[7, 28, 29]. Alterations in the joint’s anatomy or deficiencies in ligamentous and/or
capsular components can cause motion abnormalities and may lead to the
development of high focal contact stresses on the cartilage. High contact stresses can
deteriorate the cartilage tissues and eventually they can be affected by shoulder
osteoarthritis [30-33].
14 Chapter 2: Literature Review
2.1.1 Shoulder osteoarthritis and causes
Osteoarthritis is a degenerative disease that degrades the functional quality of
articular cartilage and eventually results in bone-to-bone contact in joints, leading to
chronic pain and disability [10-12]. It is the most common musculoskeletal joint
disorder accounting for 50% of the musculoskeletal disease burden [10-12]. The
economic and social costs of osteoarthritis are unprecedented, with Australia
spending nearly 5-billion Australian dollars per year for arthritis related health costs
in 2007. In terms of total financial costs (when productivity costs and other indirect
costs are accounted) this value mounts to more than 20-billion [34]. Osteoarthritis is
affecting nearly 15–20 million individuals worldwide [10-12] and is projected to
increase from three million to as many as one in four Australians by 2040 [35, 36] .
Osteoarthritis can be classified as primary or secondary osteoarthritis. Primary
osteoarthritis occurs without any predisposed reason and is known to be triggered by
aging, while secondary osteoarthritis occurs due to injuries, obesity, inactivity or
other similar causes [1, 2]. Due to the limited capacity of cartilage to repair, and its
aneural, alymphatic and avascular nature, if cartilage damage is of unmanageable
size, there is a greater possibility of an onset of tissue degeneration and eventual
osteoarthritis development.
Osteoarthritis in the shoulder joint accounts for 3% of all osteoarthritis and it
is also the third most common osteoarthritis-affected joint after the knee and hip
[14]. Degenerative changes, irregularities, delamination and eventual cartilage loss
will affect movement in the shoulder cartilage (i.e. roughness and stiffness on
movement will increase) and may additionally result in joint instability [37]. The
most common anatomical degenerative changes due to shoulder osteoarthritis are
contracture of the anterior aspect of the capsule, posterior shoulder subluxation, and
Chapter 2: Literature Review 15
central and proximal wear of the articular cartilage of the humeral head [37, 38].
Details of the pathoanatomical degenerative changes in the shoulder joint due to
osteoarthritis are listed in Table 2.1.
Table 2.1: Pathoanatomical changes in shoulder osteoarthritis [37, 38]. Anatomical location or
position Pathoanatomical change
Humeral head
• Central cartilage erosion and flattening
• Central and superior sclerosis
• Peripheral osteophytes, especially inferiorly
• Subchondral bone cysts
Glenoid
• Cartilage loss, especially centrally and
posteriorly
• Sclerosis
• Peripheral osteophytes, especially along the
inferior border
• Subchondral bone cysts
Resting joint position • Posterior joint subluxation
Shoulder joint capsule • Anterior contracture
• Inferior enlargement
Shoulder osteoarthritis can occur due to initial traumatic events or associated
rotator cuff tears [39, 40]. It can also occur due to recurrent instability and
misalignments because of anterior dislocation and reduction, or repetitive
subluxation; all of which result in abnormally large compressions on the shoulder
cartilage [40, 41]. The population affected by shoulder osteoarthritis due to
instability in the shoulder joint is normally young and active [39]. Some patients with
shoulder instability have been diagnosed with post-operatively-developed shoulder
osteoarthritis [39, 42]. Studies also report that shoulder osteoarthritis occurs in
manual wheelchair users, athletes and construction workers [43-46]. Athletes such as
16 Chapter 2: Literature Review
weightlifters, baseball players and racquet sports athletes are more likely to develop
shoulder osteoarthritis [40, 47], mainly due to loads and load distribution patterns on
the shoulder cartilage during dynamic movements of their shoulder [48, 49]. For
instance, there is a reported incidence of osteoarthritis as high as 33% among older
tennis players, which is higher than age-matched control subjects [50]. Studies also
report that repetitive tasks such as lifting may increase the possibility of osteoarthritis
occurrence in the shoulder [51, 52].
Non-surgical interventions such as medication and low weight-bearing
exercises are the first steps to relieve the pain, discomfort and immobility caused by
shoulder osteoarthritis [53-55]. If these treatments are proven unsuccessful, and if the
disease has reached a severe stage, then joint arthroplasty can be considered as an
option [55, 56]. However, with time, arthroplasty is known to cause problems such
as the loosening and wear of the artificial implants [53, 57]. Additionally, longevity
of shoulder implants is reported to be low compared to arthoplasty in hip and knee
[58-60]. The component radiolucency at a 12 year follow up can be high as 89% with
44% been definitive loosening of the implant [54, 60]. Further, shoulder arthoplasty
in younger and active patients has not been highly successful [61, 62]. Nevertheless,
efforts to improve the arthroplasty implants and surgical procedures are ongoing.
Surgical procedures for osteoarthritis treatment such as joint resurfacing
(involving abrasion, drilling, debridement techniques, microfracture techniques or
arthroscopic shaving) and biological auto-grafts are found to be unsatisfactory in the
long run because the fibrocartilages formed using these techniques are mechanically
inadequate to support the joint function in the long-term [63-66]. Joint cartilage
tissues affected by injuries are treated by autologous chondrocyte transplantation
with a considerable success rate in the knee [67, 68] and specific treatment
Chapter 2: Literature Review 17
procedures for focal chondral defects in knee is well established [69, 70]. Shoulder
chondral defects are rarely reported and discovers several years after the original
incident when shoulder complications surface [13, 71]. Autologous chondrocyte
transplantation is relatively new for focal chondral defects in shoulder [71-73].
Nevertheless this treatment method has been reported ( 3 years and 9 years follow up
study) to be considerably successful with more than 90% of patients does not
requiring further treatment and surgery after the treatment [72, 74].
Considerable amount of research has also focused on bio-regenerative
approaches including the promotion of native tissue generation and biocompatible
artificial tissue generation [66]. Even though these efforts are successful to an extent,
matching the mechanical properties of native tissues and simultaneously ensuring
functionality in vivo remains a significant challenge [75-78]. Most research on
cartilage is conducted on knee and hip cartilages and, especially in the field of
cartilage biomechanics, research on shoulder cartilage is limited [79]. Most tissue
engineering efforts are also focused on repairing or engineering knee cartilage
tissues, probably due to the relatively lower occurrence of shoulder osteoarthritis
[14]. However, the facts mentioned earlier and the report by the Australian
orthopaedic association (AOA) that from 2008 to the beginning of 2014 there is
steady increase in the number of shoulder replacement imply the significant need to
carry out specific research on shoulder cartilage [15].
2.1.2 Adaptation of cartilage tissues in response to the mechanical environment
Kinematics and dynamics are important factors in regulating the composition,
microstructure and mechanical properties of biological tissues [80-84]. During
human development from childhood to maturity, the lower and upper extremity
joints undergo forces which are significantly different in magnitude and mode
18 Chapter 2: Literature Review
(section 6.1). In addition, the upper extremity joints have different functional
requirements compared to the lower extremity joints, and hence are different in terms
of the anatomy and biomechanics [85]. The mechanical properties and the potential
material behaviour of cartilage tissues are regulated by the magnitude and mode
(compressive, tensile and shear) of the force experienced by individual joints. There
is evidence to suggest that the properties of cartilage tissue vary significantly from
one joint to another [86]. In this regard, the shoulder joint is subject to forces which
are significantly different in magnitude and mode compared to the lower extremity
joints such as the knee. Hence, it is believed that the shoulder cartilage possesses
properties and characteristics which are unique to it.
The majority of investigations on cartilage tissues have focused on
investigating the mechanical behaviour, microstructure and composition of lower
extremity cartilage tissues such as the tissues in the knee and hip. There are only a
few investigations that have specifically focused on mechanical behavior of upper
extremity cartilage tissues such as the shoulder cartilage (In section 2.4 studies on
mechanical behavior of cartilage tissues are summarized). Although not primarily a
weight-bearing joint, in daily work and sports activities that involve grasping, lifting
or throwing, the shoulder undergoes a variety of static and dynamic loads. The
highest static reaction force experienced by the shoulder joint is 44–90% of the body
weight during arm elevation of 60°–100°, with shear forces being almost 50% of
body weight at 60° of abduction [87-90]. During 30°–90° abduction and 60°–90°
external rotations, the resulting articular contact pressure is reported to be in the
range of 2.2–5.1 MPa [31]. However, under dynamic loading, depending on the arm
position and velocity, the joints can experience 2 to 3 times loading than under static
loading conditions [91, 92]. For example, during a throwing movement in adults and
Chapter 2: Literature Review 19
professional baseball pitchers, compressive/distraction forces of 108% of body
weight and external rotational torques (causing shear stresses) in the range of 67–
92 Nm are generated [1, 93, 94]. For non-athletes and youth, these values can be 40–
50% lower [1]. The shoulder joint of a young baseball pitcher can experience
compressive/distraction forces of 49.8±8.3% (214.7±47.2N) of body weight and
external rotational torques up to 17.7±3.5 Nm [95]. Considering these facts and
based on the underlying philosophy that cartilage adapts based on external
mechanical stimuli, the present study investigates the mechanical properties and
material behaviour of shoulder cartilage under various loading-rates with the
objective of relating the mechanical behaviour of shoulder cartilage to its structure
and composition. The following sections elaborate on the general articular cartilage
structure, composition and function based on the findings reported in the literature.
2.2 ARTICULAR CARTILAGE
Diarthrodial joints experience a variety of mechanical loads under a range of
physiological loading-rates during daily activities. Articular cartilage, which is a thin
translucent tissue found at the bone ends of an diarthrodial joints, is predominantly a
mechanical biological tissue that has the ability to endure a lifetime of compression,
tension and shear forces at a variety of loading-rates, without any significant damage
to the tissue [96, 97]. Its superior mechanical properties and behaviour are due to the
structural organisation and properties of its constituents: water, a three-dimensional
collagen network and negatively-charged, highly-hydrated proteoglycan
macromolecules [98]. By composition, collagen and proteoglycan account for 30%
and 10% of the cartilage’s wet weight, while the water content ranges from 70–80%
of which 4–5% is bound water [99-101].
20 Chapter 2: Literature Review
2.2.1 Articular cartilage proteoglycan
Proteoglycans are glycosylated proteins which are synthesised by eukaryotic
cells. They are found in almost all mammalian tissues in diverse structures and
composition, depending on the specific biological and mechanical function of the
tissue [102, 103]. Aggrecan, the main proteoglycan of the cartilage, is also the
largest proteoglycan in all biological tissues [104] , consisting of a hyaluronic acid
core bonded to aggrecan monomers. Aggrecan monomers are brush-like structures
with glycosaminoglycans (chondroitin sulfate and keratan sulfate) side chains
attached to the hyaluronic acid core protein (Figure 2.2). Aggrecan has
approximately 150 glycosaminoglycan chains attached to the core protein, which is
the highest among all proteoglycans. [102, 104, 105].
Figure 2.2: Articular cartilage proteoglycan aggregate
The physical and biological characteristics of proteoglycans are governed by
the physiochemical nature and structure of glycosaminoglycan chains and the core
Chapter 2: Literature Review 21
protein. Glycosaminoglycans in cartilage, such as chondroitin sulfate, carry a high
negative charge and hence attract positively-charged ions and water molecules. This
accounts for unique swelling properties of cartilage, and hence its resilience to
compressive forces [100, 106]. Therefore, normal cartilage function is significantly
affected by the size of the proteoglycans, the number of glycosaminoglycan side
chains, and the concentration of proteoglycan aggregates [99, 107]. Osteoarthritis is
frequently characterised by the functional compromise of its aggrecan due to
cleavage of the glycosaminoglycan side chains or hyaluronic core protein. This
reduces the size of aggrecan, its water attraction ability, and its ability to aggregate
with other aggrecan [108]. Proteoglycan aggregates impede fluid flow when an
external load is applied on the tissue, resulting in the immediate rise of fluid pressure
and its gradual reduction with time when the water flows out. The rate of fluid flow
is controlled by the ability of the proteoglycan aggregates to impede flow, which
affects the compressive load-bearing ability of the tissue [109].
2.2.2 Articular cartilage collagen
Collagen is the building block of the supporting structure of the cartilage
extracellular matrix where the proteoglycan gel and other cartilage constituents
remain constrained. It accounts for two-thirds of the dry weight of adult cartilage
tissues. Collagen fibres are responsible for the tensile strength of the cartilage and,
together with proteoglycans, form an integrated tissue matrix [98]. Secreted by
articular cartilage cells, collagen-II accounts for more than 90% of all collagens
while the rest of the collagens are type XI (~3%) and IX (~1%). The typical triple-
helix structure of the collagen-II molecule is formed by three polypeptide α-chains
twisted in three directions. Collagen-II molecules bundle together in repeated
structures, 67 nm apart, to form collagen fibrils (Figure 2.3).
22 Chapter 2: Literature Review
Figure 2.3: Collagen triple helix structure and assembly [110].
Collagen fibres are formed by assembling parallel collagen-II fibrils
surrounding a collagen-XI fibril and an outer layer of collagen-IX fibrils. Collagen
fibres are 30–80nm in diameter and approximately 100 nm apart within the
extracellular matrix [111]. Intra and inter-molecular cross-links between the fibril
and collagen network considerably contribute to the mechanical properties of the
cartilage matrix [112, 113]. Therefore, cross-linking molecules such as families of
matrilins and small leucine-rich proteins play an important role in maintaining the
mechanical integrity of cartilage tissues. Cross-linking not only imparts cohesive
strength to the tissue matrix but also helps to constrain the swelling of proteoglycans
in the extracellular matrix. There is evidence that proteins such as decorin and
matrilin together also link the collagen fibre network and proteoglycans [106, 111,
113].
2.2.3 Articular cartilage structure
Health of the cartilage is maintained by equilibrium between the rate of synthesis and
rate of degradation of the extracellular matrix components by articular cartilage cells
Chapter 2: Literature Review 23
called “chondrocytes” [114-116]. Chondrocytes continuously synthesise
proteoglycans and collagens that are soaked in a physiological liquid environment
consisting mainly of water and mobile ions (Na+ and Cl- mostly). The concentration,
orientation and shape of the chondrocytes vary with the depth of the tissue, with the
highest concentration found in the surface layer. Chondrocytes found in the surface
layer are flat-horizontal in shape, while in the middle zone they are round shaped
(Figure 2.4).
Figure 2.4: Zonal variation of matrix components in articular cartilage [117]
In the deep zone, near to the tide mark, cells become vertically oriented, following
the direction of collagen fibres with the depth. Arguably, the depth-dependent
variation of chondrocytes and collagen structure is linked to the mechanical
environment, with high compressive load-bearing areas having a high cell volume to
synthesise more proteoglycans [109, 118]. The skeleton of the cartilage tissues,
which is arguably the collagen network, has a very low biological turnover. It is
responsible for the tensile strength of the cartilage, while cross-linking in the
24 Chapter 2: Literature Review
collagen structure provides cohesive strength to the tissue [119]. Based on the
arrangement of fibres in the surface and deep zone, the cartilage matrix is often
classified into different zones (Figure 2.4). In the superficial zone or surface zone,
the collagen fibres are predominantly parallel to the articular cartilage surface. The
parallel fibre arrangement transits to a random arrangement in the transition zone and
runs radially in the deep zone, anchoring perpendicular to the tidemark in the
calcified zone near the subchondral bone [120].
Proteoglycan concentration has been also found to vary with the depth of the
cartilage [121, 122]. The reported typical variation in the concentration of
proteoglycans with cartilage depth is a skewed bell shape distribution with an
increase in concentration from a low value in the superficial zone to a maximum
value around the deep zone, followed by a decrease nearer to the tidemark. The high
affinity of proteoglycans to water and the proteoglycan-collagen interaction create a
three-component gel-like structure wherein the collagen structure with the
interconnected network constrains the swelling of hydrated proteoglycans (Figure
2.5). Arguably, the zonal architecture and proteoglycan distribution with depth
protect the chondrocytes from being damaged by external loads. For example, in the
early stages of osteoarthritis when proteoglycan levels have been observed to
decrease, the chondrocytes respond by increasing the synthesis of proteoglycans to
repair the tissue [123]. However, if further degradation occurs, especially affecting
the collagen network, the chondrocytes may fail to repair the tissue further. Thereby,
the tissue may become functionally incapable of responding to external loads and
that may lead to chondrocyte damage and eventual osteoarthritis [124].
Chapter 2: Literature Review 25
Figure 2.5: Complex fibre and proteoglycan molecule arrangement; the proteoglycans are trapped in three dimensional collagen fibre network [125]
2.2.4 Articular cartilage collagen network architecture
The specialised collagen architecture of the articular cartilage significantly affects
the mechanical behaviour of cartilage tissues and is responsible for proteoglycan
entrapment. The current structural models of cartilage’s collagen architecture depict
an arcade-like structure (Figure 2.6(a)), where the collagens are parallel to the
cartilage surface [126]. The parallel collagen fibres change to a radial arrangement
in the deep layers of the cartilage and anchor perpendicular to the subchondral bone
[127, 128]. At the ultrastructural level, fibrils deviate from the radial direction
repeatedly in short ranges to form segmental obliquity (Figure 2.6(b)). Segmental
obliquity facilitates collagen cross-linking to form an interlocking collagen structure
that gives rise to a highly efficient entrapment system that enables the cartilage
matrix to contain swelling proteoglycans (Figure 2.6(c)). The degree of obliquity and
26 Chapter 2: Literature Review
the number of collagen fibre interactions determine the amount of constraint on
proteoglycans and their swelling potential, which in turn is responsible for the load-
carriage properties of the cartilage [98, 119].
Figure 2.6: (a) Idealised fibre arrangement in articular cartilage – Benninghoff structure; (b) Typical segmental obliquity in collagen fibres observed; (c) Cross-linking and interlocking of collagen fibres facilitated by the segmental obliquity [129]
2.2.5 Articular cartilage load-bearing unit: Proteoglycan and collagen entrapment
The structure of the articular cartilage can be physically represented by the balloon-
string model of Broom and Marra [130] (Figure 2.7). Highly-packed, hydrated
proteoglycan aggregate macromolecules (Figure 2.7(b)) experience repulsive forces
due to proteoglycan–proteoglycan interaction and are constrained by the cross-linked
collagen meshwork (Figure 2.7(a)).
Chapter 2: Literature Review 27
Figure 2.7 Three-dimensional network of fibres in articular cartilage generated from repeated cross-linking of string elements; (b) The balloons representing proteoglycans packed and constrained in the fibre network; (c) The complete balloon-string model showing upper tension diaphragm of the surface layer and fibre-proteoglycan arrangement [130].
The articular cartilage’s load-bearing functional unit is this system of entrapped
fluid-swollen proteoglycan molecules and the cross-linked collagen meshwork
(Figure 2.8). The negative charge in proteoglycan aggregates is neutralised by mobile
ions in the solution; and, due to the imbalance of mobile ion distribution, an osmotic
pressure is generated [98]. Due to the osmotic pressure, the cartilage swells and
enters a pre-stress state (Figure 2.8), enhancing its compressive load-bearing capacity
[98, 131]. The collagen fibres are predominantly tension-resisting structures, while
proteoglycans are highly deformable under direct compression. However, due to the
structural arrangement, these constituents together form a functional load-bearing
unit that can withstand a wide range of loading (Figures 2.7(c) and 2.8). Although
cartilage is mostly known as a compressive load-bearing structure, it cannot function
without the intrinsic strength of the matrix [98, 106]. Experiments have shown that a
reduction in fibre connections and removal of glycosaminoglycan tend to reduce the
compressive load-bearing ability [132, 133]. Therefore, this integrated system of
fibres and hydrated macromolecules is essential to create an efficient compression-
resisting system [98] that can enable cartilage to bear external loads.
28 Chapter 2: Literature Review
Figure 2.8: The load-bearing functional unit of articular cartilage; the confining forces of the collagen network resist the osmotic swelling pressure of the proteoglycan and give rise to the intrinsic stiffness of the cartilage matrix [134]
2.3 ARTICULAR CARTILAGE BIOMECHANICS: STATIC AND DYNAMIC LOAD-BEARING MECHANISMS
Under static loading, the fluid inside the cartilage initially bears the stress (Figure
2.9(a)) which develops a hydrostatic excess pore pressure over and above the
osmotic pressure. Then, the pore pressure reaches a maximum value (Figure 2.9(b)).
However, as deformation progresses, the pores in the proteoglycan-entrapped
collagen structure close and this reduces the tissue’s permeability [135, 136]. The
increased resistance to flow, caused by pore closure, not only reduces the fluid
pressure but also gradually transfers the stress to the collagen meshwork (Figures
2.9(a) and (b)). The tissue’s deformation eventually stops when the solid skeleton is
able to fully resist the externally applied load. This stress-sharing and stress-transfer
mechanism is called “consolidation” and is also observed in porous materials such as
clay and soil [135].
Chapter 2: Literature Review 29
Figure 2.9: (a) Mechanical analogy illustrating the stress-sharing stress-transfer consolidation mechanism; (b) Pore pressure variation, when the articular cartilage is statically loaded, illustrating how the fluid supports the external load, data adapted from [135].
Although cartilage can be considered as a compressive load-bearing tissue, in
vivo, the compression is localised; hence, the surrounding tissue is additionally
subjected to tension. Therefore, arguably, mechanical indentation tests are
physiologically close to the loading experienced by tissue in vivo, and hence were
used for experimentations in the present study. Acknowledging the stress-sharing and
stress-transfer mechanism, the tension-resisting superficial collagen fibres also
influence the tissue behaviour during indentation [137, 138]. However, in confined
compressions, arguably, superficial collagen does not significantly contribute to the
load-bearing function as there are no significant tensile stresses in the superficial
cartilage layer. Nonetheless, depth-dependent proteoglycan variation is known to
30 Chapter 2: Literature Review
contribute to the load-bearing ability of tissue under both confined and unconfined
compression [139, 140].
Although the mechanisms underlying the static load-bearing of cartilage are
well understood, deformation mechanisms under different loading-rates and dynamic
loading are still under investigation [16, 79]. This is due to the difficulty of
monitoring time-dependent internal tissue behaviour under varying external loads.
Evidence in the literature indicates that the mechanical behaviour of articular
cartilage is strain-rate-dependent [16, 141-149]. According to experimental findings,
with increasing strain-rate, the stiffness of cartilage increases at the beginning and
then it approaches an asymptotic value [141, 142, 150] . The interplay between solid
and interstitial fluid contributes significantly to this behaviour, with 70%–80% of the
load being supported by the matrix at low strain-rates (10-4/s) [149], while the fluid
contributes to a similar percentage at moderately large strain-rates (10-2/s) [144, 149].
In order to explain the strain-rate-dependent behaviour of cartilage tissue, a
rheological model (Figure 2.10) has been proposed, with springs demonstrating the
progressive stiffening behaviour of the tissue with deformation [142, 151]. During
deformation, the articular cartilage exudes water out of the matrix. The rate of water
exudation is controlled by the tissue permeability. This role of water in the
deformation process is represented by a special fluid dashpot, Q, in Figure 2.10.
Immediately after applying the load, the dashpot is set in motion and the fluid is
exuded out of the matrix due to the increase in pressure difference. Dashpots D1 and
D2 model the frictional drag due to interaction between the fluid and solid
components during deformation. It is believed that when the strain-rate is increased,
the fluid exudation from the matrix becomes increasingly difficult due to the drag
forces [142].
Chapter 2: Literature Review 31
Figure 2.10: Extended rheological analogy of cartilage matrix featuring σa, applied stress; kv stiffness of instantaneous spring; di (i=0,1,2,…), length of ith dent; ki (i=1,2,…), stiffness of ith spring; D1, instantaneous damping coefficient for unbound fluid; D2, damping coefficient relating to bound fluid; Di, general damping coefficient of unbound fluid in the dashpot; σeff, solid’s effective stress; Q, permeability coefficient of the matrix; O, active osmotic component [151].
Although it is known that solid–interstitial fluid interaction affects the cartilage
behaviour significantly, detailed investigations about its effect on the strain-rate-
dependent behaviour of shoulder cartilage are limited [79]. In particular, there are no
reported studies in the literature on the significant or insignificant nature of the
strain-rate-dependent fluid flow in regard to cartilage deformation behaviour.
Although studies [145, 149] suggest that the loading velocity affects the fluid
behaviour inside the tissue, scarce research has investigated the effect of loading-
32 Chapter 2: Literature Review
rate dependent fluid behavior mechanical behavior of shoulder cartilage tissues.
Further, Oloyede and Broom [149] reported that the relationship between the
effective-matrix-stress and the pore pressure changes considerably when the strain-
rate is increased from 10-3/s to 10-2/s. They found a fundamental change in the
deformation mechanism when the strain-rate increased and claimed that this could be
due to fluid being locked/contained inside the tissue with the increase of loading
velocity. Based on these facts, it is anticipated that drag forces introduced by the
reduction of permeability and solid–interstitial fluid frictional interactions largely
contribute to the strain-rate-dependent behaviour and that the locking effect is mainly
due to the pressure drag forces and possibly inertia forces affecting the tissue
behaviour.
In addition, it has been stated that flow-independent matrix viscoelasticity
affects the strain-rate-dependent behaviour of cartilage tissue. For instance,
DiSilvestro and Zhu [146], investigated tissue behaviour from 10-2/s to 10-4/s and
stated that viscoelasticity governs the tissue behaviour at high strain-rates, while
fluid flow affects the tissue behaviour at low strain-rates. Although there is
experimental evidence for flow-independent viscoelasticity [152-154], the question
of whether the cartilage matrix truly possesses viscoelasticity is subject to ongoing
investigation [155, 156]. The effect of superficial collagen and flow-independent and
flow-dependent viscoelastic mechanisms in combination is also postulated to affect
the indentation tissue behaviour at large strain-rates [156]. On the other hand, direct
fluid pressure measurements inside the cartilage have shown that flow-dependent
drag forces mainly govern the apparent viscoelastic response of cartilage under
dynamic loading [157].
Chapter 2: Literature Review 33
2.3.1 Articular cartilage biomechanics: The structure–function relationship
Acknowledging that the collagen network’s structure and depth-dependent
proteoglycan distribution are the result of the mechanical environment to which the
tissue is subject, we believe that tissue composition and structure have a significant
effect on the tissue behaviour. For example, the dynamic properties of cartilage
(extracted at high strain-rates) are considered to be governed by the collagen network
structure [20, 158]. Taking into consideration the cartilage structure and its
composition, Julkunen et al.’s [16] finite element (FE) model has shown that the
superficial collagen can considerably affect the tissue behaviour at high strain-rates.
Further, during fast loading, the collagen meshwork is shown to limit the shape
changes of the cartilage in unconfined compression [159]. In contrast, the
equilibrium properties of cartilage (extracted at very low strain-rates) are mainly
affected by proteoglycans [17, 20]. Similarly, proteoglycan concentration has been
found to correlate to the equilibrium properties of articular cartilage [160-162].
Chondrocytes dynamically synthesise the extracellular matrix (i.e.
proteoglycans and collagen) based on the external loading stimuli they receive [163-
165]. Therefore, proteoglycan composition and the structural features of the collagen
network may adapt to external mechanical stimuli, and hence may depend on the
local mechanical environment of the tissue [86, 166-171]. The proteoglycan content
of knee cartilage that bears high compressive loads is anticipated to be higher than
the proteoglycan content of shoulder cartilage that experiences low compressive
loads [168, 172, 173]. In addition, there may be differences in the distribution of
depth-dependent proteoglycan concentration and the features of the collagen
network. However, the studies that have investigated the structure–function
34 Chapter 2: Literature Review
relationship of cartilage have predominantly focused on knee cartilage [16, 17, 20,
158].
Given the potential differences in composition and microstructure, it is highly
possible that the conclusions drawn for the structure-function relationship of knee
cartilage tissues may not be directly applicable to shoulder cartilage tissues. While
systematically investigating the mechanisms underlying the strain-rate-dependent
mechanical behaviour of shoulder cartilage, this study also investigates how the
structural features of the shoulder cartilage affect its strain-rate-dependent
mechanical behaviour. This will help to comprehensively understand the underlying
factors affecting the strain-rate-dependent behaviour of shoulder cartilage. Before
proceeding to investigate the shoulder cartilage, it was crucial to review the previous
studies conducted specifically on shoulder cartilage. These findings are summarised
in the following section.
2.4 CHARACTERISTICS OF SHOULDER CARTILAGE
Using stereophotogrammetric studies, the thicknesses of the humeral head cartilage
and glenoid cartilage are found to be 1.5–2 mm and 2–4 mm, respectively [8, 174].
Magnetic resonance imaging (MRI) and ultrasound studies have also shown that the
shoulder cartilage is thicker in the central part of the humeral head while it is thickest
in the periphery of the glenoid [175]
Cohen and Lai [176], Mow and Bigliani [28] and Huang and Stankiewicz [177]
are the only researchers who have investigated the mechanical behavior and
properties of shoulder cartilage tissue. The compressive properties such as the
aggregate modulus obtained in these studies have not shown any statistically
significant difference within or between the glenoid and humeral head cartilages [28,
177]: for humeral head cartilage, the Young’s modulus reported by Cohen and Lai
Chapter 2: Literature Review 35
[176] and the aggregate modulus reported by Mow and Bigliani [28] and Huang and
Stankiewicz [177] are 0.46 MPa, 0.63-0.77MPa and 0.15MPa, respectively. The
tensile modulus of humeral head cartilage is reported to be significantly higher than
its compressive modulus, and is in the range of 4.23–5.81 MPa [178]. This is one of
the characteristics, known as tension–compression nonlinearity [179, 180], which
significantly affect the mechanical behaviour of cartilage tissues. Tension–
compression nonlinearity stems from the tension-resistive superficial collagen fibres
which impart a higher tensile modulus on the tissue compared to its compressive
properties.
There are no reported significant differences in permeability between humeral
head cartilage and glenoid cartilage. However, Huang and Stankiewicz [177]
reported significant differences in permeability values within different locations of
the humeral head cartilage, while Mow and Bigliani [28] reported no significant
differences. The average permeability values reported in these studies for the central
region of the humeral head cartilage range from 1.82x10-14m4/Ns [178] to 5.1x10-
15m4/Ns [28]. Although the aforementioned values are different from each other, the
average age of the subjects from which the specimens were harvested was similar
(p<0.05). The findings from these studies indicate that there are no significant
differences in the mechanical properties in humeral head cartilage and glenoid
cartilage. Therefore, the present study focused on the central load-bearing area of the
humeral head cartilage which is also where osteoarthritis-related changes mainly
manifest.
36 Chapter 2: Literature Review
2.5 BIOMECHANICAL MODELS OF ARTICULAR CARTILAGE
Due to the remarkable load-bearing ability of articular cartilage, its susceptibility to
osteoarthritis, and the difficulty in experimentally investigating the internal tissue
behaviour, cartilage has been of interest to biomechanical modellers during the past
half-a-century. Biomechanical models have investigated the mechanisms underlying
the mechanical behaviour responses of normal and osteoarthritis-affected cartilages;
the role of cartilage components on tissue behaviour; and how the cartilage
composition and structure affect the tissues’ mechanical behaviour [181]. The main
challenge in modelling articular cartilage is identifying a suitable way to
conceptualise the tissue and finding ways to include structural features into the
models [113]. Biomechanical models of articular cartilage can be broadly divided
into three types: mixture models, Biot’s [182] theory-based models, and fibril-
reinforced composite models.
Mixture models for cartilages were first developed by Mow et al. [183] who
conceptualised the cartilage as a mixture of its constituents. These models consider
that the cartilage consists of distinctly different phases, namely, solid and fluid
phases. While these earliest mixture models explain the macroscopically-observed
confined compression data satisfactorily [183]. However, some limitations in model
predictions have been observed [184, 185] . In unconfined compression, the earlier
mixture models were unable to capture the typical, large stress relaxation behaviours
(ratio of peak force to equilibrium force can be high as 10) observed for cartilage
tissues, where they have a theoretical limit of 1.5 for the peak force to equilibrium
force ratio [186]. Accounting for the friction between cartilage–indenter interfaces
was not enough to improve the theoretical limits [187, 188]. Deviation of mixture
models in creep experiments was observed during the early stages of creep
Chapter 2: Literature Review 37
deformation [160, 189]. To date, there have been numerous extensions to the original
mixture model to account for tissue anisotropy [176] , large deformations [190, 191],
viscoelasticity [192], swelling [193] and tension-compression nonlinearity [180].
These extended models have been able to improve the limitation of mixture models
and is able to explain the tissue behaviour under various loading conditions.
There were discussions in the late 1970s to early 1990s about whether the
conceptualisation of cartilage as a mixture is representative of the functional roles of
cartilage constituents [135, 194]. For example, collagen (which mainly resists
tension) and hydrated proteoglycans (which deform excessively under direct
compressive loads) are unlikely to function separately to form a remarkable
compressive load-bearing unit in the cartilage [98]. Based on these arguments, there
was debate about whether the conceptualisation of cartilage as a mixture was
acceptable [194]. Therefore, some researchers prefer to use to use Biot’s theory
[182] rather than mixture theory for cartilage. This is because, conceptually, Biot’s
theory does not consider cartilage to consist of two distinctly different phases.
However, it should be noted that mixture theory was reviewed by Simon [195], who
concluded that, mathematically, the mixture models are similar to the models based
on Biot’s theory.
Soulhat et al. [179] and Li et al. [185] introduced fibril-reinforced cartilage
models by conceptualising cartilage as fibril-reinforced composite materials that
distinctly distinguishing the role of fibrils and the non-fibrillar matrix (proteoglycan
and water). The main objective of the fibril-reinforced composite models is to
explain the large stress relaxations observed in the unconfined compression of
cartilage tissues and to understand the role of fibrils and non-fibrillar components.
Fibril-reinforced models have been extended to non-homogeneous models [196] with
38 Chapter 2: Literature Review
depth dependent material parameters [197] and further extended to a 3D cartilage
model to include the realistic 3D collagen fibre orientation [198]. These models are
advantageous in explaining the large stress relaxations observed in cartilage and in
understanding the contribution of matrix constituents to the mechanical behaviour of
the tissue.
Acknowledging the advantages of the fibril-reinforced models, it is important
to mention some of the limitations and concerns about these models. One of the main
limitations of the fibril-reinforced models is the inability to represent collagen cross-
linkers and the dense collagen meshwork [113, 155]. . Furthermore, fibril-reinforced
models rely on seemingly non-physical parameters such as the Poisson’s ratio of the
proteoglycan gel [199]. The limitations of fibril-reinforced models have been
partially addressed by Wilson, van Donkelaar [200] and Wilson, van Donkelaar
[201]. Their models included a greater number of springs and paid more attention to
the interactions between the collagen and proteoglycans [199].
Another concern about the conceptualisation of cartilage as a fibril-reinforced
model arises from the experimental observation of Broom and Silyn‐Roberts [106].
They observed that proteoglycan itself has little influence on the cohesive strength of
the fibril matrix. Further, there are experimental findings supporting the argument
that there are no significant mechanical interactions between the proteoglycan gel
and the fibril matrix [202]. Considering these observations, it is unlikely that there is
a significant shear stress transfer between the fibril network and proteoglycans,
unlike in fibril-reinforced composite materials [98].
In the ABAQUS 6.12 commercial FE software (Abaqus 6.12, SIMULIA,
Rhode Island, USA), Biot theory based porous media models can be easily
implemented using built in software capabilities. Therefore, the present study
Chapter 2: Literature Review 39
modelled cartilage as a porous media saturated with fluid, based on Biot’s theory
[182]. This theory was originally developed by Terzaghi [203] for the one-
dimensional consolidation of soils and was later generalised to three dimensions by
Biot [182]. Oloyede and Broom [135], using a series of experiments [131, 142, 204],
demonstrated that the theory is applicable to cartilage. The theory considers cartilage
to comprise an incompressible solid skeleton saturated with incompressible mobile
fluid. Under load, the fluid is able to exude; hence, the solid skeleton is deformed
due to volume loss.
2.6 SUMMARY AND IMPLICATIONS
Based on the above literature review, the following main points and research gaps
have been identified:
• A large number of studies have been conducted to investigate the behaviours
of lower limb cartilage tissues such as knee cartilages. However, scarce
research is available on one of the most important joint cartilages, namely,
shoulder cartilage. While the lack of studies on shoulder cartilage—the third-
most osteoarthritis-affected joint—is surprising, the growing incidence of
shoulder osteoarthritis and shoulder arthroplasty, as well as the increasing
numbers of people participating in sporting activities, suggest that it is crucial
to investigate the characteristics of shoulder cartilage.
• Evidence reported in the literature shows that biological tissues such as
cartilage adapt according to the mechanical loading environment. Although
not mainly a load-bearing joint, shoulder joint undergoes a variety of forces
during daily activities. These forces are both static and dynamic. Hence,
unique compositional, microstructural and biomechanical properties and
behaviours are expected in shoulder cartilage tissues, although these have
40 Chapter 2: Literature Review
hitherto lacked attention. Furthermore, given that the shoulder cartilage
undergoes both static and dynamic loading, and that the individuals who are
most prone to shoulder osteoarthritis are those who perform dynamic
activities using the shoulder, it remains important to investigate the dynamic
loading behaviours of shoulder cartilage.
• To best of author’s knowledge, an investigation to identify the mechanisms
facilitating the strain-rate-dependent mechanical behaviour of shoulder
cartilage tissues has not been conducted previously. Therefore, it remains
important to understand the factors affecting the dynamic load bearing ability
of shoulder cartilage.
• The findings and conclusions on the structure–function relationship in
cartilage have mostly been based on investigations carried out on knee
cartilage. Considering the potential differences in composition and
microstructure in the knee and shoulder cartilage, it is highly likely that the
conclusions that have been drawn may not be directly applicable to shoulder
cartilage. This remains a significant gap in knowledge which could have
important implications for both the numerical modelling and tissue
engineering strategies catering for specific joint cartilages. Therefore, it is
important to investigate the structure–function relationship of shoulder
cartilage specifically.
• The existing literature on shoulder cartilage points out that there are no
significant differences between the properties of the cartilage layer in the
humeral head and glenoid. Therefore, in the current study, only the behaviour
of humeral head cartilage tissues was investigated.
Chapter 2: Literature Review 41
• Cartilage can be modelled as a porous media saturated with fluid, based on
Biot’s theory. The theory considers cartilage to consist of an incompressible
solid skeleton saturated with incompressible mobile fluid which exudes under
the application of load.
• The main hypothesis of this thesis was: Due to the adaptability of biological
structures to mechanical loading, compositional and structural differences in
shoulder cartilage is anticipated in comparison to high weight bearing
cartilages and these differences is reflected in the factors affecting the
mechanical behavior of the shoulder cartilage. This main hypothesis is broken
down in to several hypotheses in the three studies conducted in this thesis.
Based on the identified research gaps, it can be stated that investigation into the
behaviour of shoulder cartilage remains an important research area. Hence, the
following chapters report this study’s investigation of the strain-rate-dependent
mechanical behaviour of shoulder cartilage tissues with the objective of
understanding the underlying mechanism and its relationship with the shoulder
cartilage structure.
42 Chapter 2: Literature Review
Chapter 3: Research Design and Methodology
This study is based on the underlying philosophy that the structure of biological
tissues such as cartilage adapts to the external mechanical stimuli and that their
mechanical behaviour depends on the structure and composition of the tissue.
Therefore, we hypothesized that microstructural features and composition of
shoulder cartilage are different from cartilage tissues that are more frequently loaded.
Further, these differences should be reflected in the mechanical properties and
behaviour of shoulder cartilage.
This chapter presents the rationale for the research methodologies employed in
this study and describes the research design and methods adopted to achieve the aims
and objectives (as stated in Chapter 1, Section 1.3). Section 3.1 discusses the animal
model chosen for experimentation purposes. Section 3.2 elaborates on the
experimental methodology and explains the stages of the implemented research
methodology including the preliminary results that were required for the subsequent
stages of the study. Lastly, Section 3.3 presents the computational modelling
methodology used to investigate the research problems identified.
3.1 EXPERIMENTAL ANIMAL MODEL FOR SHOULDER CARTILAGE
Cartilage tissue obtained from the human shoulder joint would have been the most
suitable sample for the present investigation, since joint kinematics and dynamics are
regarded as important factors that regulate the mechanical properties of biological
tissues [81, 205]. However, human specimens which are usually diverse (in terms of
weight, age etc.) are not only difficult to obtain but their usage involves ethical and
Chapter 3: Research Design and Methodology 43
legal restrictions. On the other hand, animal models are more readily available and
involve fewer ethical and financial constraints [206].
In selecting a suitable animal model for shoulder cartilage research, the
shoulder joint of the animal model should be anatomically and biomechanically
similar to that of a human shoulder joint. Ovine, bovine, steer, canine and rat have
been previously used to investigate the mechanical properties of shoulder cartilage
tissue [207-209] . However, except for rats, the anatomy and biomechanics of the
shoulder joint in other animal models are different to that of a human shoulder [210].
This is because the quadruped animals use forelimbs for weight-bearing during
locomotion with minimal overhead activity. Their movements are largely restricted
to the sagittal plane. Humans are bipedal and do not use the shoulder much for
weight-bearing activities. Further, humans can additionally rotate and move in the
coronal plane, giving the shoulder more mobility [211]. These differences have
significant implications for the adaptation and architecture of shoulder cartilages
[210], and potentially for its mechanical properties and behaviour too.
Apart from non-human primates, macropods, rats and certain types of mice
(kangaroo mice, hopping mice etc.) have shoulder joints similar to that of humans.
Rat is one of the most commonly used animal models for shoulder research because
it is considered to have similar bone anatomy and overhead activity to that of a
human shoulder [210]. However, the small tissue thickness of its articular cartilage is
a disadvantage in carrying out macroscale mechanical testing. On the other hand,
there are ethical and economic concerns that limit the use of non-human primate
tissues for experimentations [210]. The shoulder joint of rare species such as tree
kangaroo also has a very similar anatomy and biomechanics to that of a human
shoulder [210, 211]. Recently, kangaroo has been postulated as a potential animal
44 Chapter 3: Research Design and Methodology
model to study upper limb joint cartilages [79, 168, 212-214] due to the significantly
lower amounts of loading experienced by the upper limb joints compared to the
lower limb joints, making it similar to humans. Considering these factors, kangaroo
was selected as the most suitable animal model for the shoulder cartilage research
conducted in this study.
3.2 EXPERIMENTAL METHODOLOGIES AND MATERIALS
3.2.1 Tissue harvesting and preparation
Visually normal [ICRS [215] macroscopic score=0] cartilage samples of 8 mm
diameter with 2–3 mm of subchondral bone intact (Figure 3.1(a)) were harvested
using a specially designed stainless steel puncher. The samples were obtained from
the central load-bearing area of the humeral head (Figure 3.1(b)), from adult red
kangaroos (approximately 5 years old), bought from an abattoir within 24 hrs of
slaughter. After harvesting, the samples were preserved in phosphate-buffered saline
(PBS)-inhibitor solution containing inhibitors of proteolytic enzymes (5 mM
benzamide-HCL and 5 mM EDTA) and antibiotics (200 mM L-glutamine 10000
units of penicillin and 10 mg/mL of streptomycin; Sigma-Aldrich, Castle Hill, NSW)
and were stored at -20 °C [216]. Before subsequent biomechanical testings, the
samples were thawed for approximately 30 minutes in PBS at room temperature (i.e.
approximately 27 °C) [217]. The samples went through a single freeze–thaw cycle so
as to ensure that their composition and structure were not affected by multiple
freeze–thaw cycles [216, 217].
Chapter 3: Research Design and Methodology 45
Figure 3.1: (a) 8 mm diameter cartilage sample; (b) Specimen-harvested region (near the central area of the humeral head); (c) Bone was constrained using a stainless steel holder and submerged in physiological (0.15 M) saline solution; (d) Indentation testing on the sample
3.2.2 Evaluation of potential thickness measurement methods
Several methods are available to measure the thickness of cartilage tissues. Notably,
these methods can be broadly divided into destructive and non-destructive methods.
Destructive methods of thickness measurement include anatomical sectioning or
punch probes/biopsy [218], needle probe [219], optical [148], and
stereophotogrammetry [220]. In anatomical sectioning, a biopsy is obtained from the
place of interest before thickness is measured using a precision caliper or through
microscopic observation. In the needle probe method, a sample with subchondral
bone intact is fixed to a mechanical testing machine attached with a needle indenter
and is pierced at a point where the thickness measurements are of interest. In this
method, the cartilage surface and calcified bone interface are identified using the
force–displacement profile and the cartilage thickness is obtained by finding the
difference in measurements. In the optical method, the osteochondral cartilage
samples are imaged using a microscope and the subsequent thickness measurements
are analysed using these images. In the stereophotogrammetric method, the joint
cartilage surface and its underlying bone are imaged after dissolving the cartilage in a
46 Chapter 3: Research Design and Methodology
5.25% sodium hypochlorite solution and the difference is considered to be the
cartilage thickness [220].
MRI, near-infrared resonance (NIR) and ultrasound are the main non-
destructive methods used to measure the thickness of cartilage. MRI is especially
used for in vivo measurements [221] and has not been frequently used to measure the
cartilage thickness of osteochondral plugs. Recently, NIR has been identified as a
potential method to measure cartilage thickness [222]. However, in this method, in
addition to the complex post-processing of NIR spectral data, extensive calibration
(requiring simultaneous NIR and needle probe measurements of a considerably large
number of samples (90–100)) of the specific tissue of interest is also required [109].
On the other hand, the ultrasound technique measures the thickness of the cartilage
based on the difference in the time taken for echoes to reach the sensor after
reflection from the cartilage surface and cartilage–bone interface; hence, it does not
require complex post-processing [223]. Therefore, in the present study where
thickness measurements were required before indentation testing, ultrasound was
chosen as the most suitable technique. However, for the thickness calculation using
ultrasound measurements it is necessary to know the speed of the sound in the
respective cartilage. The procedure employed for ultrasound speed measurements in
the kangaroo shoulder cartilage is elaborated in the next section.
3.2.3 Ultrasound speed in kangaroo shoulder cartilage tissues and thickness measurement
Ultrasound thickness measurement may incur errors due to the assumed value of
ultrasound speed in cartilage [224], which depends on tissue microstructure and
composition. In the present study, in order to calculate the speed of the sound in
kangaroo shoulder cartilage, the travel time of sound (based on the difference
between echoes from the saline–cartilage interface and osteochondral junction) was
Chapter 3: Research Design and Methodology 47
recorded for the samples harvested from the central load-bearing area of the humeral
head. A 10 MHz, Ø3 mm plane-ended contact transducer (V129 Panametrics Inc.,
Massachusetts, USA) was used for the ultrasound measurements and a 3 mm
distance was set between the cartilage surface and transducer during these
measurements based on previously published results in our laboratory [225]. The
transducer, excited by a pulser-receiver (Model 5072PR) was connected to an
oscilloscope (Model PC 5204) that converts the analogue signal to digital. The
sampling frequency of the pulser-receiver was 50 MHz. The echoes from the surface
and subchondral junction reflections were recorded using PicoScope software (Pico
Technology Limited, Cambridgeshire, UK).
After ultrasound measurement, the samples were placed in a sample holder to
reduce the possible movement of the sample and were fixed to a high-resolution
mechanical testing machine called an Instron (Model 5944, Instron, Canton, MA,
USA). A custom-made needle indenter (needle probe) was attached to the load cell
of the machine for needle probe measurements. Then the needle probe was gently
lowered until it just touched the tissue surface and the sample was indented at 10
mm/min [219, 222] until the cartilage–bone interface was reached, as identified by
the load displacement curve (Figure 3.2(b)). Five needle probe measurements were
obtained within an approximate 3 mm diameter circular area (Figure 3.2(a)). The
measurements were averaged and plotted against the respective ultrasound travel
times for all the samples tested (n=43). Based on a previous study, 40–50 samples
were chosen for this testing [219]. The previous study observed a good linear
correlation between the needle probe and ultrasound measurements [219], whereby
the gradient of the plot gave the average speed of sound in the cartilage tissues.
48 Chapter 3: Research Design and Methodology
Figure 3.2: (a) Needle probe measurement location in the sample; (b) A typical force–indentation curve during needle probe indentation – The curve is characterised by articular cartilage (AC) surface puncture due to piercing of the cartilage surface; the initial gradient of the curve significantly increases when the cartilage-calcified bone surface is reached
The ultrasound travel times plotted against the needle probe thicknesses for the
tested samples are shown in Figure 3.3. The two measurements showed a good
correlation for the tested samples (R2=0.9829, p<0.005). The slope of the curve
indicated that the average speed of the sound inside the kangaroo shoulder cartilage
is 1658.27 ms-1 (Figure 3.3). This value was taken for all the subsequent thickness
calculations for kangaroo shoulder cartilage in this research.
Ultrasound Measurement area
Needle probe measurement points
Sample surface
AC surface
AC surface puncture Tide mark
Estimated uncalcified cartilage thickness
(a) (b)
Chapter 3: Research Design and Methodology 49
2x10-7 4x10-7 6x10-7
5x10-4
1x10-3
2x10-3N
eedl
e pr
obe
thic
knes
s(m
)
Ultrasound travel time(s)
Figure 3.3: Plot between ultrasound travel time and kangaroo shoulder cartilage thickness from needle probe measurements; the slope of the curve is the average ultrasound speed in the tissue
3.2.4 Biomechanical characterisation: Mechanical tests performed on articular cartilage
Indentation, confined compression and unconfined compression tests are commonly
used to assess the mechanical properties of articular cartilages [226]. In confined
compression, a cartilage sample without subchondral bone is tested with a porous
disc on top, after confining the perimeters of the sample in a confining chamber.
Using the confining configuration, the most commonly obtained measurements
include the consolidation measurements and equilibrium stress-strain measurements
of the tissue [135, 226]. In the consolidation measurement, the pore pressure
underneath the samples and the strain measurement are first obtained. Afterwards,
based on the effective stress principle, the effective stress-strain curve is obtained
and the solid skeleton properties of the cartilage are evaluated based on the curve
[135]. In the equilibrium measurement, the samples are compressed stepwise and
Needle probe thickness = 1658.27 x Ultrasound travel time R2=0.9829
50 Chapter 3: Research Design and Methodology
allowed to relax after each compression in order to obtain the equilibrium stress-
strain curve. The aggregate modulus ( AH ), which is an indicator of solid matrix
stiffness, is obtained from the gradient in the linear region of the equilibrium stress-
strain curve. On the other hand, in unconfined compression, stepwise ramp
compression is employed on samples which are not confined in perimeter and
Young’s modulus of solid matrix ( sE ) is extracted from the linear part of the
equilibrium stress-strain curve [226]. Young’s modulus of solid matrix is related to
the aggregate modulus using Eq. (3.1), where sν is the solid matrix Poisson’s ratio:
As
sss HE
ννν
−−+
=1
)21)(1( (3.1)
Indentation tests, using porous and nonporous indenters, are commonly
performed on cartilage samples with the bone intact [189, 227]. Samples are
compressed stepwise and allowed to relax after each indentation and the subsequent
equilibrium stress-strain curves are used to obtain the mechanical parameters when
the porous indenter is used [146].
3.2.5 Critical evaluation of confined, unconfined and indentation mechanical tests
In the case of confined compression, it is difficult to fit the samples into the
confining chamber perfectly. If the sample cannot be perfectly fitted to the confining
chamber, frictional forces may act between the sample and the confining chamber
during compression [184]. This may affect the mechanical parameters measured,
especially when extracting dynamic properties [146, 226]. Additionally, the
mechanical parameter obtained may be affected by the interdigitation of the cartilage
surface and the pores of the disc during testing [184]. This may possibly damage the
Chapter 3: Research Design and Methodology 51
tissue, given that interdigitation depends on the pore size and pore distribution,
which may not be uniform unless specifically controlled [160, 184, 228].
In both confined and unconfined compression tests, removing the sample from
the subchondral bone affects the integrity of the natural cartilage–bone system.
Further, if precision cutters are not employed during sample-shaving, the cartilage
may possibly be damaged, especially if the cartilage samples are thin. The natural
collagen architecture of cartilage is such that fibres are anchored perpendicular to the
underlying subchondral bone in the calcified zone. Disruption to the integrity of this
structure may affect the mechanical properties of the cartilage. Studies have reported
that Young’s modulus obtained from indentation tests is significantly higher than that
obtained from confined and unconfined compression tests, and this has been partially
attributed to the cartilage samples being removed from the bone [226]. Unconfined
compression tests on bone-intact samples can also be used to obtain the mechanical
properties of cartilage tissue. However, the geometrically complex behaviour of the
tissue at the periphery of the cartilage–bone interface makes it difficult to obtain an
analytical solution to this type of compression problem [179, 229].
Due to the above reasons, indentation testing can be considered a better
alternative to unconfined and confined compression tests. Indentation moduli have
also been identified to correlate well with Young’s modulus obtained from
unconfined compression or the aggregate modulus obtained from confined
compression [226, 230] . Although the size of the indenter affects the force–
displacement results, instant and equilibrium shear moduli are found to be
independent of the indenter radius [231, 232]. However, when using the porous
indenter, not only can it result in interdigitation but there are additional difficulties in
precisely implementing the indenter–cartilage interface boundary conditions in a
52 Chapter 3: Research Design and Methodology
numerical model [184]. Moreover, Jurvelin and Kiviranta [233] stated that the
instant and equilibrium response should be the same for both porous and non-porous
indenters, since there is no water flow through the indenter in both cases [234].
Indentation using a non-porous indenter is shown to resemble the physiological
cartilage-to-cartilage contact during joint function [204]. However, during
indentation, friction between the cartilage and indenter interface may affect the tissue
behaviour, especially when a flat indenter is used. In addition, during repeated tests it
is also possible for the tissue to be damaged. In order to address these concerns,
synovial fluid can be employed on the indenter. By using an indenter with rounded
edges, both friction and potential damage to cartilage can also be reduced. Moreover,
indentation tests on cartilage have been widely used to obtain the mechanical
properties of tissues for the clinical diagnosis of osteoarthritis both in vitro and in
vivo, and therefore serve as an important method of tissue characterisation [230,
235-238]. Hence, in the present study, the indentation test was chosen to characterise
the kangaroo shoulder cartilage tissue. Details of the indentation test conducted in the
present study are provided in the following section.
3.2.6 Mechanical testing protocol
3.2.6.1 Physiological strain-rates and strains experienced by joints
A number of earlier studies have tested cartilage under different strain-rates. Radin
and Paul [141] tested cartilage from 2.7×10-3/s to 3.5×10-2/s and Lai and Mow [239]
loaded the cartilage from 3.3×10-5/s to 3.3×10-4/s. Investigating the cartilage
response for a wide range of strain-rates, Oloyede and Flachsmann [142] indented
cartilage from 10-5/s to 103/s. More recently, Langelier and Buschmann [148],
DiSilvestro and Zhu [146] and Li and Buschmann [145] studied cartilage in the
range of 10-4/s to 5×10-2/s strain-rates. To the best of the author’s knowledge, there
Chapter 3: Research Design and Methodology 53
are no existing studies that have measured the physiological strain-rates experienced
by cartilage in vivo. However, the maximum physiological strain-rate measured by
Rubin and Lanyon [240] in a treadmill study in which strain gages were attached
directly to the radial and tibia mid shaft of dogs and horses was 8×10-2/s (when the
animals were galloping). In [241], load cells were implanted on the tibia of a rabbit
approximately 1 cm below the knee joint and strain-rates of 3×10-2/s were reported
when impulse loads were applied on limbs. Therefore, a strain-rate in the order of 10-
2/s can be considered as at the higher end of the physiological strain-rates. Also
strain-rates in the range of 10-5/s to 10-4/s are generally considered to be at the low
end of the physiological loading-rates [142].
The peak load of an impact on joints, which potentially causes tissue damage,
occurs in much less than a second after the initial application of the load [242, 243].
For example, motor vehicle impact accidents, which occur in milliseconds, generally
happen at around 103/s strain-rate [244]. Sub-impact loads occur in several seconds
(in the order of 1/s strain-rates), and may induce surface cracks and chondrocyte
death [242]. Further, from reported in vivo deformation data on the tibiofemoral
joint, physiological strains on average can go up to 30%, ranging from 10% to 40%
during daily activities [245, 246]. Based on the above information, in the current
study, strain-rates ranging from 10-4/s to 10-2/s were chosen to cover the
physiological low and high ends of strain-rates and cartilage was loaded up to 30%
strain to represent the average loading in joint cartilages.
3.2.6.2 Indentation testing
Before conducting mechanical testing, the subchondral bone underneath the cartilage
sample was properly fixed using a stainless steel holder (Figure 3.1(c)) to ensure that
the deformation data obtained in the testing were only related to the cartilage
54 Chapter 3: Research Design and Methodology
deformation. The indentation testing was carried out at 10-4/s, 5x10-4/s, 5x10-3/s and
10-2/s strain-rates (Figure 3.1(d)). Depending on the thickness of the samples, the
speed of the indentation was adjusted in order to obtain the required strain-rate. The
samples were indented up to 30% engineering strain and a further limit of 3.5 MPa
was imposed on the amount of stress that the samples were subjected to, in order to
minimise potential damage to the tissues. A safe limit of 3.5 MPa for strain-rates
between 3×10-5/s and 7×10-1/s has been suggested to prevent damage to the cartilage
matrix [242, 247]. In addition, before and after every test, the sample surfaces were
microscopically examined (Leica MZ6, Leica Microsystems, Heerbrugg,
Switzerland) to check whether testing had induced any damage to the cartilage.
Although we rarely found damage to the tissue, testing on the specific sample was
terminated in the cases where damage was found.
The testing was done on a high-resolution Instron testing machine
(Model 5944, Instron, Canton, MA, USA) using a plane-ended polished indenter of 3
mm diameter with 0.1 mm radius rounded edge. An indenter with rounded edge was
chosen in order to reduce possible local damage to the cartilage due to stress
concentration at the indenter edges. Spherical indentation could also be used;
however, the theoretical analysis, especially the solid–fluid interaction analysis,
would become complicated since the contact area changes with the deformation of
the tissue. Additionally, it has been experimentally demonstrated that the contact of
the cartilage–solid indenter resembles cartilage–cartilage contact [204]. After each
test, the cartilage was unloaded and allowed to recover for 1 hr in PBS-inhibitor
solution at 4 °C prior to the next test.
Chapter 3: Research Design and Methodology 55
3.2.6.3 Surface detection, zero-point determination
Identifying the point where the sample surface first touches the indenter (zero-point)
is important in order to accurately extract the mechanical properties from the force–
displacement results. This is especially relevant for soft material such as cartilage
where a small pre-load can have a significant deformation in the tissue, leading to
overestimation of the material’s properties [248, 249]. It is customary in cartilage
testing to apply a pre-load of 0.01N-0.05N to make sure that the sample is in contact
with the indenter [177, 250]. However, in the present study, the surface detection
method developed by Cao and Yang [251] as described by Kaufman and Klapperich
[248] was used to identify the cartilage surface before the indentation tests. This
involves carefully lowering the indenter using the ‘fine position controller’ in the
Instron machine until it just touches the cartilage surface (indicated by the positive
readings of the load cell) and then retracting a step backwards to make sure that the
indenter is just above the sample surface.
3.3 NUMERICAL MODELLING METHODOLOGY
3.3.1 Numerical modelling to investigate the physical mechanisms underlying the mechanical behaviour of cartilage: Initial model development
FE modelling has been widely used to investigate cartilage behaviour due to
experimental difficulties in probing the internal tissue behaviour. For instance,
inserting sensors inside the tissue will affect its structure and hence also affect the
natural behaviour of the tissue. Additionally, under dynamic loading conditions, the
measurement of fluid velocity, fluid pressure and collagen network stresses requires
sophisticated experimental set-ups, especially for small tissues such as shoulder
cartilage. In an earlier attempt to carry out an experimental investigation using MRI
in this study, to measure fluid velocity inside the tissue, it was found that facilities
are extremely expensive to acquire. Therefore, the most pragmatic method to
56 Chapter 3: Research Design and Methodology
investigate the research problems in this study was to carry out a combination of
experimental and numerical investigations that complemented each other. The
following sections summarise the procedures employed to develop the numerical
model of cartilage, and the initial numerical studies that were carried out to confirm
the appropriateness of the modelling technique. The modelling in this study was
conducted in ABAQUS 6.12 commercial FE modelling software (Abaqus 6.12,
Simulia, Rhode Island, USA), which is widely used in developing biomechanical
models of cartilage tissues.
Cartilages undergo nonlinear lager deformations [177]. Hyperelastic material
models can be used to represent large deformation behaviour of cartilage tissues
[252]. However, there are no analytical equations available for indentation of
hyperelastic solids that could be used for the present study’s indenter geometry. In
section 3.3 an analytical equation for the indenter geometry used in this study is
developed. Before that, in order to check the appropriateness of boundary conditions
and mesh density of the initial numerical model an analytical relationship reported
[253] for the spherical indentation of isotropic linear elastic material is used in this
section..
The analytical relationship, which relates indentation force )(F to indentation
depth )(δ , for a linear elastic solid with thickness h , indented with a spherical
indenter (radius R ) is [253]
++
+−+−
−= 4
0
2304
030
2303
22
200
2
23
21
5316
154842
1)1(3
4χβ
πα
πα
χβπ
απ
χπα
χπαδ
vERF (3.2)
where hRδχ = , the constant 0α and 0β are functions (Eq. (3.3) and Eq. (3.4)) of
Poisson’s ratio )(ν and E is the Young’s modulus.
Chapter 3: Research Design and Methodology 57
νννα
−+−
−=1
3442.14678.12876.1 2
0 (3.3)
νννβ
−+−
=1
5164.10227.16387.0 2
0 (3.4)
Eq. (3.2) above has been developed based on the assumption that, in order to
maintain the material lineally, the strain should not exceed 10%. Further, a constraint
of Rh 1.0≥ is suggested for the validity of Eq. (3.2). In the present study, we first
developed a numerical model for spherical indentation by representing the cartilage
as a linear elastic material. Then, the numerical model prediction was compared with
the prediction of Eq. (3.2) to confirm the appropriateness of the boundary conditions
in the model. The effect of mesh density on the model predictions was also
evaluated. All the details related to the model development are set out in the
following sections.
3.3.1.1 Model geometry
The geometry of numerical models were based on the dimensions of the individual
samples tested (Table 3.1). The maximum thickness of kangaroo cartilage measured
in this study was less than 1.5 mm. Therefore, the thickness value of the initial
cartilage model, that was used to obtain the mesh density and to check model
validity, was taken as 1.5 mm. Further, considering that the radius of the indenter
used in the present study was 1.5 mm, this value was taken as the radius of the
spherical indenter. The bone and indenter were assumed to be rigid throughout the
study due to the fact that the stiffness of the bone and indenter is several magnitudes
larger than the stiffness of cartilage tissues. A test simulation was performed and
indicated that representing both bone and indenter as rigid bodies did not affect the
simulation results. Young’s modulus of cartilage was taken as 0.5 MPa, which is a
58 Chapter 3: Research Design and Methodology
typical average value for cartilage tissues [254]. The material was assumed to be
nearly incompressible for simplicity, and Poisson’s ratio was taken as 0.499 in
accordance with the ABAQUS 6.12 user manual [255]. It is noted that these
assumptions did not affect the subsequent conclusions made.
Table 3.1: Parameters of the model used for mesh, boundary and loading condition validations
Diameter of sample (mm) 8.0
h (mm) 1.5
Thickness of bone (mm) 3.0
R (mm) 1.5
E (MPa) 0.5
ν 0.499
3.3.1.2 Model mesh
Considering the geometry of the cartilage samples, the assumption of homogeneity
and isotropy, and the loading conditions used during mechanical testing an
axisymmetric model were employed. Cartilage was meshed with 4-node bilinear
axisymmetric quadrilateral (CAX4). The cartilage mesh and model are shown in
Figure 3.4(a). Model simulations for 800, 8000 and 22,400 elements were also
evaluated.
3.3.1.3 Boundary conditions of the model
The boundary conditions of the model were set according to the boundary conditions
to which the specimens were subjected during mechanical testing. The enforced
boundary conditions are summarised as follows:
Chapter 3: Research Design and Methodology 59
• The interaction boundary condition between the indenter and cartilage was
specified as ‘frictionless’ using ABAQUS software options. This simulated
frictionless contact between the cartilage surface and indenter.
• The boundary between the cartilage and bone (highlighted in a red line in
Figure 3.4(a)) was simulated using the ‘Tie constraint’ boundary condition in
ABAQUS 6.12.
• During the experiment, the bone was constrained using a stainless steel
holder to ensure that the deformation was only caused by deformation of the
cartilage samples. Therefore, the displacement and rotation of the bone in all
directions were set to zero by fixing the reference point (RP) of the bone
(Figure 3.4(a)).
• Deformation in the ‘Z’ direction was allowed on the left symmetric plane of
the cartilage model.
• The indenter was prescribed with displacement at its RP in order to simulate
the indentation process. The amount of displacement was specified based on
the deformation to which the samples were subjected during the experiments.
In order to assess whether the model in Figure 3.4(a) could be further
simplified, the bone was replaced by a rigid constraint. This restricted the
displacement and rotation of the lower cartilage boundary in all directions
(highlighted in a red line in Figure 3.4(a)), while keeping all other boundary
conditions the same. The two model results were compared with each other as well
as with the theoretical prediction of Eq. (3.2).
60 Chapter 3: Research Design and Methodology
Figure 3.4: (a) The model geometry, mesh, boundary condition and loading configuration of the cartilage–bone FE model; (b) Simplified FE model (without the bone) with mesh and boundary conditions – In this model, the bone is replaced by a rigid constraint (indicated by the red line) which restricts the displacement of the bottom plane of the cartilage
3.3.1.4 Solution methodology
The static stress analysis procedure in ABAQUS/Standard was used for the above
indentation problem. The procedure uses the Newton method to solve nonlinear
equilibrium equations. The initial time increment was set at 10-5/s, and the maximum
time increment was set at 0.1 s.
3.3.1.5 Comparison of the model results and the theoretical predictions
Figure 3.5(a) and Figure 3.5(c) show the results of the model simulation along with
its comparison with the theoretical prediction for 800, 8000 and 24,400 elements. It
was observed that the numerical results and the theoretical model conformed well,
with only a slight deviation at large deformation. Further, the two models (i.e. the
model with the rigid bone and the model with the rigid constraint) showed identical
Chapter 3: Research Design and Methodology 61
force–indentation graphs (Figure 3.5(d)), implying that the proposed model
simplification could be carried out. Stress distributions were also observed to be
identical (compare Figure 3.5(a) and Figure 3.5 (b)).
Figure 3.5: (a) Stress distribution for cartilage on rigid bone indented to 10% strain; (b) Stress distribution for cartilage on rigid constraint indented to 10% strain; (c) Comparison of the numerical model’s result (force on indenter) with the theoretical prediction of Eq. (3.2) and mesh sensitivity data; (d) Comparison of elastic cartilage on rigid bone laminate model results and elastic cartilage-rigid constraint model results with theoretical model results
After confirming that the FE mesh and the model including the boundary
conditions are adequate, the spherical indenter was replaced with the geometry used
in the present study and the number of elements was evaluated to obtain model
62 Chapter 3: Research Design and Methodology
accuracy of 10-2N for force measurements. The developed model mesh is shown in
Figure 3.6(a). Based on the results it was noted that, in order to obtain 10-2N level of
accuracy, 8000 to 12,000 elements were required (Figure 3.6(c)). Therefore, 9600
elements were considered as appropriate and were used for all simulations in the
study.
Figure 3.6: (a) Cartilage, indenter geometry (3 mm diameter with 0.1 mm fillet radius at the edge), the mesh, boundary condition and loading configuration based on mechanical testing carried on kangaroo shoulder cartilage samples; (b) Numerical result of elastic cartilage samples indented up to 30% engineering strain; (c) Variation of force on indenter based on mesh element number
Chapter 3: Research Design and Methodology 63
3.3.2 Assessing the suitability of the porohyperelastic model for investigating the solid and fluid behaviour of cartilage tissues: The preliminary porohyperelastic FE model
In Chapter 2 it was explained that the poroelastic model based on Biot’s theory [182]
is an appropriate starting point for investigating the mechanisms underlying the
mechanical behaviour of shoulder cartilage tissues. Further, it was hypothesised that
the solid–fluid interplay governs the mechanical behaviour of the tissue. Since
cartilage generally undergoes large deformations, it was necessary to impart a
theoretical framework that extends Biot’s theory [182] to include large deformations.
Therefore, porohyperelastic field theory, an extension to Biot’s theory [182] (using
the hyperelastic solid skeleton to account for large deformation) proposed by Simon
[195] was chosen as the theoretical basis for the present study. This theory is
discussed in Chapter 4 in detail. The advantage of using Biot’s theory [182] is that it
can be directly implemented in ABAQUS 6.12 commercial software with a
hyperelastic solid skeleton to simulate large deformation behaviour. However, before
applying the model based on porohyperelastic field theory to investigate the research
problem identified in the literature review, it was necessary to check the
appropriateness of the model for investigating the solid and fluid behaviour of the
tissue. Therefore, a preliminary porohyperelastic model was first developed based on
experiments carried out by Oloyede and Broom [135] and then model
appropriateness was evaluated. Details of this preliminary model development and
comparison with the experimental data are set out in the following sections.
3.3.2.1 Model geometry
The static confined consolidation experiments of Oloyede and Broom [135] were
used for the development and validation of the porohyperelastic FE cartilage model
due to the availability of reported direct internal pore pressure measurements. The
64 Chapter 3: Research Design and Methodology
thickness and diameter of cartilage samples based on the sample geometry in
Oloyede and Broom [135] were taken as 1.6 mm and 10 mm, respectively.
3.3.2.2 Model mesh and solution methodology
Due to the symmetric nature of the sample and the type of loading scenario, an
axisymmetric model was developed in order to reduce computational time. The
model was meshed with 9,600 4-node bilinear displacement and bilinear pore
pressure elements (CAX4P). Transient consolidation analysis with the ‘Full Newton’
solution procedure was used to solve the equilibrium equations. The transient
coupled pore pressure/effective stress analysis uses a backward difference operator to
integrate the continuity equation, and therefore it provides unconditional stability.
However, the only concern regarding time integration was accuracy. The minimal
time increment )( t∆ for saturated flows, such as in cartilage, to avoid any evident
oscillation is: [255];
2
2
)(16
)1(l
KE
Ekt
g
ww ∆∆
−
+>
βυg (3.5)
where wg is the specific weight of the wetting fluid, E is the Young’s modulus, k is
the permeability, wυ is the velocity of the pore fluid, β is the velocity coefficient
( 0=β for Darcy flow), gK is the bulk modulus and l∆ is a typical element
dimension. The initial time increment for transient consolidation analysis was set at
10-5/s and, considering Eq. (3.5), the maximum time increment was restricted by the
time scale of the problem solved.
3.3.2.3 Boundary conditions
The static confined consolidation experiments by Oloyede and Broom [135] were
simulated by applying a constant load (1.35 MPa) to the upper boundary in the
Chapter 3: Research Design and Methodology 65
negative Z direction (Figure 3.7). Based on the experiment’s boundary conditions,
the ‘pore pressure (p)’ (p=0) boundary condition was enforced on the upper surface
to enable free fluid flow out of the cartilage. The displacement (U) boundary
condition on the right side of the model was set to allow deformation in the Z
direction, while sample deformation in the X direction was constrained. The
displacement at the bottom of the sample was set to zero to mimic the sample glued
to the base of the experimental rig.
Figure 3.7: Boundary conditions employed in preliminary porohyperelastic FE model
3.3.2.4 Initial conditions
The cartilage was considered to be fully saturated, and 80% of the tissue was
assumed to be filled with fluid that fully occupied the pores inside the tissue.
Therefore, the initial value (0e ) as the ratio of the volume of pores to the volume of
solid was set at 4.0. The void ratio was calculated by n
ne−
=1
, where n is the
porosity. Further, the initial pore pressure value was set at zero given that the
osmotic pressure of the tissue was not considered during the study [183].
3.3.2.5 Solid skeleton material model
To account for large deformation, the solid skeleton was modelled as an isotropic
hyperelastic material [256, 257]. The neo-Hookean model, which is the simplest
hyperelastic material model, was used for this initial model. However, it was later
found that the nonlinearity of the solid skeleton of kangaroo shoulder cartilage was
66 Chapter 3: Research Design and Methodology
more accurately represented by the 2-term reduced polynomial hyperelastic model
(Chapter 4, Section 4.5.1.1). The isotropic elastic strain energy potential W of the
neo-Hookean model used in the present model is given by:
2
1110 )1(1)3( −+−= J
DICW (3.6)
where )(1 CtrI = is the first invariant of the distortional part FFC T= of the right
Cauchy deformation tensor FFC T= where FF 31−= J is the distortional part of
the deformation gradient F , and Fdet=J is the volume ratio.
Furthermore, 102C=µ , where µ is the shear modulus of the linear elasticity and
κ21 =D , where κ is the bulk modulus of the linear elasticity [79].
3.3.2.6 Strain-dependent permeability
The fluid flow was modelled based on Darcy’s law. The permeability ( k ) decreases
under strain application and can be represented by Eq. (3.7) [191, 258, 259].
Parameter 0k is the permeability of the tissue in an undeformed configuration and M
and m are dimensionless material parameters which were taken as 4.638 and 0.0848,
respectively [191]. Eq. (3.7) is expressed as follows:
−
++
= 1
e1e1
2Mexp
eekk
2
0
m
00 (3.7)
3.3.2.7 Solid skeleton material parameter identification
Young’s modulus was extracted from the initial gradient of the solid skeleton
effective stress-strain data reported in [135], by fitting it to a piecewise linear graph
as shown in Figure 3.8. For bovine cartilage, the Poisson’s ratio was taken as 0.2
Chapter 3: Research Design and Methodology 67
according to a reported study [260]. The parameters 10C and 1D were calculated
from Eq. (3.8). The permeability at an undeformed state of the tissue was taken as
2.58x10-16 Ns/m4, in accordance with values reported in the literature [261]. The
material parameters that were used in the model are summarised in Table 3.2.
)1(410 ν+=
EC , E
D )21(61
ν−= (3.8)
Figure 3.8: Solid skeleton effective stress-strain curve fitted with a piecewise linear curve to
extract the solid skeleton material parameters
Table 3.2: Hyperelastic material parameters and permeability values used for the initial porohyperelastic FE model
10C (MPa) 1D (1/MPa) k (Ns/m4)
0.158 4.738 2.58x10-16
3.3.2.8 Model prediction
The pore pressure values obtained from the bottom ‘P’ point (Figure 3.7) and the
displacement values obtained from the compression platen of the model are shown in
Figure 3.9(a) and Figure 3.9(b), respectively. The model prediction indicated that
pore pressure and creep strain, although not perfect, followed experimentally
68 Chapter 3: Research Design and Methodology
observed trends acceptably. The model using 400, 6400 and 14,400 elements did not
show any considerable differences in model prediction (Figure 3.10(a) and Figure
3.10(b)). These results demonstrated that the porohyperelastic model can be
considered as an initial model to investigate the solid and fluid behaviour of cartilage
tissues. Therefore, the model was used to investigate the mechanism underlying the
mechanical behaviour observed during the indentation experiments in this study.
The mechanical properties of cartilage tissues can be extracted using
indentation tests where experimental force–indentation data are fitted to a specific
analytical solution. As mentioned above, there are, however, limited analytical
solutions available for the indentation of hyperelastic materials. Especially for the
geometry of the indenter used in the present study, no analytical equations have been
reported. Therefore, for the 2-term reduced polynomial hyperelastic function, we
derived a force–displacement relationship for the indentation problem and used it to
extract the hyperelastic mechanical properties of kangaroo shoulder cartilage. The
methodology and background of the derivation are set out below.
Figure 3.9: (a) Pore pressure measurements compared with FE model predictions; (b) Creep strain measurements compared with FE model predictions
Chapter 3: Research Design and Methodology 69
Figure 3.10: Mesh sensitivity analysis for (a) pore pressure predictions (b) creep strain prediction
3.3.3 Development of force–indentation relationship for the 2-term reduced hyperelastic model
The Hertz relationship [262] is the most widely known analytical solution for the
indentation problem of elastic solids which arises from indentation of infinitely thick
samples. Hayes and Keer [263] introduced a correction factor to account for the
finite thickness of samples, for infinitesimal deformation. Jurvelin and Kiviranta
[233] extended the values of the correction factors for large aspect ratios. Further,
Zhang and Zheng [264] considered the correction factors for large deformations and
friction between the indenter and the sample. The bond between the sample and
substrate has also been identified as an important factor when extracting mechanical
properties [265]. Considering the bond between the sample and the substrate, a
force–indentation relationship for spherical indentation was derived by Dimitriadis
and Horkay [253]. The main characteristic of these formulations is that the material
was considered to be linear elastic; hence, the typical nonlinear large deformation
responses of biological tissue cannot be adequately represented.
Based on the hyperelastic models that describe the nonlinear material
behaviour, analytical equations for force–indentation problems have been recently
70 Chapter 3: Research Design and Methodology
reported [266, 267]. However, the available equations are only for spherical
indentation and do not consider the effect of samples bonded to substrates and hence
the effect of finite thickness on force–indentation results. Given this limitation, using
finite element analysis (FEA) and theoretical analysis, we introduced correction
factors to adjust the hyperelastic material parameters for the thicknesses and indenter
geometry encountered in this study. In doing so, given the inability of the neo-
Hookean and Mooney–Rivlin models to predict certain large strain behaviours [267],
the 2-term reduced polynomial hyperelastic function was considered as the starting
point. Chapter 4 (Section 4.5.1.1) presents the evidence that this 2-term reduced
polynomial hyperelastic model is the most appropriate model for kangaroo shoulder
cartilage as compared to neo-Hookean and Mooney–Rivlin models.
The strain energy potential (W ) of the reduced polynomial model is:
iN
i i ICW )3(1 10 −=∑ =
(3.9)
Here, the first strain invariant is 2221 zyxI λλλ ++= , where yx λλ , and zλ are stretch
ratios in the X, Y and Z direction. Assuming material incompressibility and the
uniaxial loading is in the X direction, then 2/1, λλλλλ === zyx . Therefore,
1
12
02 )32()1(2
−
=∑ −+−=iN
i iiC λλλλσ (3.10)
In the case of the 2-term reduced polynomial hyperelastic function 2=N .
Therefore, nominal stress ( λσ ∂∂= W ) is:
−+
−+
−= 321412 2
220210 λλ
λλ
λλσ CC (3.11)
Chapter 3: Research Design and Methodology 71
where µ=102C and µ is the linear shear modulus; 20C is a nonlinear stiffness
parameter.
The force–indentation relationship obeying the constitutive relationship
(Eq. (3.10)) can be derived by firstly accounting for the sign conversion for the
nominal stress, that is, by replacing σ with *σ− for nominal stress in indentation.
This is because indentation exercises a compressive strain on the material while the
positive convention in Eq. (3.10) is for the tension. Additionally, the relationship
between indentation strain and stretch ratio is: *1 ελ −= . Incorporating these
relationships into Eq. (3.11) gives Eq. (3.12):
−−
+−+−
+
+−+−
= *
3*2*
*2*
*2*3*
20*2*
*2*3*
10*
13
12334
12332
εεε
εεεεε
εεεεεσ CC (3.12)
Assuming that the indenter–sample adhesive force is negligible, the average
nominal stress is equal to the mean contact pressure: 2* aF πσ = where F is the
indentation force and a is the contact radius. In the current formulation, the
indentation strain ( *ε ) is defined as rδε =* where δ is the indentation depth and
r is the indenter radius. Therefore, the following relationship between indentation
force and depth can be obtained:
−−
+−+−
+
+−+−
= 23
32
22
3223
2022
3223
103
2334
2332
rrr
rrrrrC
rrrrrCF
δδδ
δδδδδ
πδδ
δδδπ (3.13)
Here, 10C is related to Young’s modulus ( E ) and Poisson’s ratio )(ν
through)1(3 210 νπ −
=EC . The force indentation Eq. (3.13) accounts for the material
nonlinearity; however, it does not account for the finite thickness of samples.
72 Chapter 3: Research Design and Methodology
Therefore, two correction factors, namely 1k and 2k , are introduced into Eq. (3.13) to
obtain the correct values for coefficients 10C and 20C . The modified equation is as
follows:
−−
+−+−
+
+−+−
= 23
32
22
3223
20222
3223
1013
2334
2332
rrr
rrrrrCk
rrrrrCkF
δδδ
δδδδδ
πδδ
δδδπ (3.14)
3.3.3.1 Relationship between correction factors and sample thickness
The above correction factors were obtained by calibrating Eq. (3.13) to a prediction
of the FE model which simulated the indentation of a flat cartilage sample using
indenter geometry in the present study (as discussed in Section 3.3.1 in relation to the
model development). In the FE model, the 2-term reduced hyperelastic polynomial
function was chosen as the material, with the values of the material parameters, 10C
and 20C , being 0.1 MPa and 0.1 MPa, respectively. The model was set up with
different thicknesses of 0.5, 0.55, 0.65, 0.7, 0.75, 1.0, 1.25 and 1.5 mm. The diameter
of the model was set to be 8 mm. In addition, it used an indenter with 3 mm diameter
and a rounded edge of 0.1 mm radius.
The force–indentation curves for 30% strain obtained from the FE model were
curve fitted to Eq. (3.14), using a custom-made nonlinear curve fitting code in
MATLAB R2014a (The MathWorks Australia Pty. Limited, NSW, Australia), in
order to obtain the 1k and 2k values for the respective thickness and indentation
depth. The dependency of the correction factors on the sample thicknesses is shown
in Figure 3.11, and can be represented by the following equations:
568.1
1 306.2−
=
rhk (3.15)
Chapter 3: Research Design and Methodology 73
435
2 9580.
.−
=
rhk (3.16)
This variation of correction factors with was incorporated in Eq. (3.14) to
obtain the correct values of 10C and 20C based on the thickness of the individual
samples. Since Eq. (3.9) has been developed assuming material incompressibility,
Young’s modulus was calculated using )1(3 210 νπ −
=EC , assuming the Poisson’s
ratio to be 0.5.
Figure 3.11: Variation of correction factors k1, k2 with sample thickness
This chapter elaborated on the research methodology including the animal
model selection, thickness measurement methodology, mechanical testing procedure,
numerical model development, material model validation and hyperelastic
mechanical property extraction. In the following chapters, the results of the
mechanical experiments are reported and discussed along with the studies conducted
to understand the mechanisms underlying the strain-rate-dependent mechanical
behaviour of kangaroo shoulder cartilage.
74 Chapter 3: Research Design and Methodology
Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage
4.1 INTRODUCTION
This chapter aims to characterise the strain-rate-dependent behaviour of kangaroo
shoulder cartilage (Section 4.2) using indentation testing at different strain-rates
(Section 4.4). After analysing the force–indentation curves and observing the
characteristic stiffness variation of the tissue with strain and strain-rate, it was
anticipated that the interaction between the solid and fluid components would be able
to fully explain the behaviour of the tissue (Section 4.3). Since the porohyperelastic
model developed in Chapter 3 (Section 3.3) is able to acceptably explain the solid
and fluid behaviour of cartilage tissues, we then evaluated the ability of the existing
constant and strain-dependent permeability models to capture the strain-rate-
dependent behaviour of kangaroo shoulder cartilage (Section 4.5). It was revealed
that the strain-rate-dependent tissue behaviour cannot be fully explained by these two
models. It is postulated that this inability is due to the rate-dependent fluid behaviour
that might be prevalent during rate-dependent loading. Therefore, the existing
porohyperelastic model was extended to incorporate strain-rate-dependent
permeability (Section 4.6) in order to comprehensively analyse the effect of fluid
behaviour on the solid–interstitial fluid interaction of the tissue. The results of the
investigation are presented and discussed (Section 4.7). The study in this chapter
resulted in a journal article named “Investigation of the mechanical behavior of
kangaroo humeral head cartilage tissue by a porohyperelastic model based on strain-
Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage 75
rate-dependent permeability” in the Journal of Mechanical Behavior of Biomedical
Materials [79].
4.2 AIMS AND OBJECTIVES
The research reported in this chapter aimed to:
1. Investigate the behaviour of kangaroo shoulder cartilage under different
strain-rates.
2. Compare the mechanical behaviour and biomechanical parameters of
kangaroo shoulder cartilage with the mechanical behaviour and
biomechanical parameters of human shoulder cartilage reported in the
literature.
3. Fully investigate the effect of interstitial fluid on mechanical behaviour of
kangaroo shoulder cartilage under different strain-rates.
4.3 HYPOTHESES
This part of the study explored two main hypotheses:
1) Solid–interstitial fluid interaction governs the strain-rate-dependent
mechanical behaviour of shoulder cartilage tissues.
2) The strain-rate-dependent fluid behaviour significantly affects the
mechanical behaviour of kangaroo shoulder cartilage.
4.4 STRAIN-RATE-DEPENDENT MECHANICAL BEHAVIOUR OF KANGAROO SHOULDER CARTILAGE
For experimentation, as previously mentioned in Chapter 3, firstly, 8 mm diameter
osteochondral samples were harvested from the central load-bearing area of the
humeral head cartilage of ten adult red kangaroos (approximately 5 years old).
Afterwards, the thickness of each sample was calculated based on ultrasound
76 Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage
measurements. Mechanical testing was conducted on osteochondral plugs where the
subchondral bone was constrained using the specially-designed apparatus. A 3 mm
diameter indenter with 0.1 mm fillet at the edge was used for indentation. Testing
was conducted by loading samples up to 30% engineering strain at four strain-rates:
10-4/s, 5x10-4/s, 5x10-3/s and 10-2/s. After each test, the samples were unloaded and
allowed to recover for at least an hour in PBS-inhibitor solution at 4 °C. The details
of this experimental procedure were presented in Chapter 3 (Section 3.2.6).
4.4.1 Tissue stiffness: Piecewise linear regression method
The obtained force–indentation data were processed and nominal stress )(σ was
plotted against nominal strain )(ε (Figure 4.1). The stiffness at different strains was
extracted from the nominal stress-strain curves by fitting a piecewise linear curve to
the data points using the shape language modelling technique [268]. Available
MATLAB source codes [268] to implement shape language modelling can be readily
specified to extract the gradients of the stress-strain curves at a given strain value.
The piecewise linear regression fitted well to the stress-strain curves (R-
squared>0.9900; Figure 4.1). The obtained variation of stiffness with strain and
strain-rate is illustrated in Figure 4.2(b).
Figure 4.1: Piecewise linear curve fit to nominal stress-nominal strain data
Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage 77
4.4.2 Stiffness variation with strain and strain-rate
The characteristics of kangaroo shoulder cartilage are not reported in the literature.
Therefore, a suitable initial step was to compare the mechanical behaviour of
kangaroo shoulder tissues with the reported values and behaviour of human shoulder
cartilage tissues. To the best of the author’s knowledge, there are no reported data on
the strain-rate-dependent nature of human shoulder cartilage tissues. However, the
experimental trends observed in the present study (Figure 4.2(a) and Figure 4.2(b))
are consistent with the data reported for bovine patellar cartilages [142, 145, 146,
148]. According to the literature, with increasing strain-rate, the stiffness of cartilage
increases and then approaches an asymptotic value at large strain-rates [142,
150].Similar to the results in the literature, the experimental results (Figure 4.2(a)) of
the current study indicated that the mechanical behaviour of kangaroo shoulder
cartilage is strain-rate-dependent, with tissue being increasingly resistive to
deformation and loading-rates as implied by the increase in tissue stiffness with
strain and strain-rate (Figure 4.2(b)).
Figure 4.2 (a) Strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage indicated by nominal stress-strain data; (b) Stiffness variation with strain and strain-rate – Stiffness was calculated by force divided by the indentation area and by displacement divided by the cartilage’s original thickness
78 Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage
Repeated measures analysis of variance (ANOVA) showed that the amount of
strain and strain-rate applied on the cartilage significantly affected the tissue
response (p<0.005). Further, Tukey’s pairwise comparison test showed that the
stiffness values at all levels of strain and strain-rate were significantly different
(p<0.005). The stiffness values reported by Langelier and Buschmann [148] for
bovine patella cartilages under different strain-rates (5x10-4/s, 5x10-3/s and 5x10-2/s)
are higher than the stiffness values in the present study (p<0.05). Nonetheless, these
differences are reasonable due to the fact that patellar cartilage bears high
compressive loading compared to shoulder cartilage in vivo. Apart from that,
differences in species may also contribute to the differences. It is noted that, despite
tight control in sample selection and experimental set-up, a relatively large standard
deviation was observed in the experimental data. However, this does not seem
unusual when looking at the standard deviation in the reported experimental data for
human shoulder cartilage [177]. Therefore, the large standard deviation is most likely
due to inherent biological variations in the samples. Throughout the thesis, the
average data of the samples with the corresponding positive standard deviation is
illustrated in figures where necessary.
As mentioned in Chapter 2, upper limb cartilage tissue such as shoulder
cartilage has noticeably low modulus in compression than in tension, differing even
up to two orders of magnitude [178]. The reasons for and implications of these
differences are still under investigation [156]. Considering similar observations in
other tissues such as tendons and ligaments, which are primarily loaded in tension, it
is postulated that the lower compressive loading experienced by these tissues is the
primary reason for this difference [156]. This disparity between tensile and
compressive properties (tension–compression nonlinearity) may have significant
Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage 79
implications for the mechanical behaviour [179, 269] and the initiation and
progression of osteoarthritis [270].
Taking into account the considerable differences in compressive loading
experienced by knee and shoulder cartilage, in future, kangaroo may provide a
potentially suitable animal model to investigate the origins of the disparity between
the tensile and compressive properties and its effect on cartilage behaviour and
health.
4.4.3 Solid–fluid interaction and its effect on the strain-rate-dependent behaviour
As mentioned in Chapter 2, articular cartilage is a fluid-saturated tissue with water-
swollen proteoglycans constrained by a three-dimensional collagen meshwork. The
interplay between solid and interstitial fluid is known to contribute significantly to
the mechanical behaviour of cartilage tissues [149]. For example, investigations on
the strain-rate-dependent behaviour of bovine patellar suggest that 70–80% of the
load is supported by the collagen meshwork at low strain-rates (10-4/s) [149] while
the fluid contributes a similar percentage at moderately large strain-rates (10-2/s)
[144, 149]. There have been claims that the drag forces introduced by permeability
reduction and solid–interstitial fluid frictional interactions contribute largely to
strain-rate-dependent behaviour [142, 149]. Therefore, looking at the characteristic
stiffness variation with strain and strain-rate in the present study, it was anticipated
that the interaction between the solid and fluid components could fully explain the
behaviour of the kangaroo shoulder cartilage tissue.
The solid–fluid interaction of cartilage tissue is often investigated using FE
models due to experimental difficulties in investigating the tissue’s internal
behaviour [144-146]. The most commonly used FE models, as discussed in the
literature review (Chapter 2), are poroelastic models based on Mixture theory [183],
80 Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage
Biot’s theory [182] and fibril-reinforced models [179] or their extensions for large
deformation. In this study, we used the extension to Biot’s theory proposed by Simon
[195] , which accounts for the large deformation behaviour of the solid skeleton. This
theory is summarised in the following section.
4.5 POROHYPERELASTIC FIELD THEORY FOR SOFT BIOLOGICAL TISSUES
The governing equations of the porohyperelastic field theory proposed by Simon
[195], Kaufmann [271] and Ayyalasomayajula and Vande Geest [256] are
summarised in this section. Porohyperelastic field theory assumes that biological
tissues can be represented as a porous incompressible solid skeleton (s) statured with
an incompressible fluid (f). The solid skeleton is hyperelastic in nature. The fluid is
free to move relative to the solid depending on the solid deformation behaviour and
friction [271]. Further, the pores are assumed to be small. Therefore, the material can
be viewed as a continuum.
Further, the pores are assumed to be small. Therefore, the material can be viewed as
a continuum.
According to porohyperelastic field theory:
fs dVdVdV += (4.1)
where V is the volume of individual solid and fluid components. The porosity ( n )
and void ratio ( e ) is defined as:
)n(JdVdVn
f
0
1 11 −−== − ;0n is the initial porosity (4.2)
nn
dVdVe
s
f
−==
1 (4.3)
Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage 81
The deformation gradient ( F ) and the volume ratio ( J ) is given by:
I
iiI X
F∂∂
=χ (4.4)
)det(F=J (4.5)
where χ is the configuration map, which maps, at each time, t , points
),,( 321 XXXX = in the reference configuration to a point ),,( 321 xxxx = in space,
that is, ),( tXx χ= .
The governing equations when expressed using Lagrangian description are as
follows:
• Conservation of linear momentum
0=∂∂
I
iI
XT ( IiiI TT ≠ ) (4.6)
Here, T is the first Piola-Kirchhoff stress.
• Conservation of fluid mass (Darcy’s law)
−=
j
f
ijfs
i dxdkv π~ ; )(0 efkkij = (4.7)
Here, the velocity of the fluid relative to the solid sffs ννν −= and the
filtration velocity )(~ sffs n ννν −= . In Eq. (4.7), fπ and k are excess pore
fluid pressure and hydraulic permeability, respectively. In addition, given that
permeability is a function of deformation, it depends on the void ratio (the
relationship between k and e is mentioned in Section 4.3.1.2). The parameter
0k is the tissue permeability at an undeformed state.
82 Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage
• Conservation of total mass
Neglecting the inertial terms, the equation of the balance of mass is:
0])1[( =+− fs nndiv νν (4.8)
Hence,
0)]([)( =−+ sfs ndivdiv ννν (4.9)
Using the definition of filtration velocity:
0)~()( =+ fss divdiv νν (4.10)
• Effective stress principle
ijfeff
ijtotalij δσσ π−= ;
=01
ijδjiji
≠= (4.11)
ijfeff
ijtotalij HJSS π−= (4.12)
where total
ijσ , eff
ijσ are the total Cauchy stress and effective stress of the solid
skeleton, respectively. The corresponding components of the second Piola-
Kirchhoff stress are totalijS and eff
ijS . Here, H is the Finger strain which is
given by 11 −−= JkIkIJ FFH . The eff
ijσ is calculated from the drained effective
strain energy density function ( eW ) as follows:
TJj
effIJiI
effij FSFJ 1−=σ , T
JjIJiIeffij FJFS −−= σ1 (4.13)
ij
e
ij
eeffij E
WCWS
∂∂
=∂∂
= 2 (4.14)
Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage 83
The Green strain tensor )(21 ICE −= and FFC T= is the right Cauchy–
Green deformation tensor. The drained effective strain energy potential for
Isotropic hyperelastic material can be written as:
),,( 21 JIIWW ee = (4.15)
where 1
3/2
1 IJI −= , 2
3/2
2 IJI −= and )(1 CtrI = , )(2 CCtrI = .
According to the above theoretical framework, the FE model was initially
developed to explore the suitability of porohyperelastic modelling for investigating
the solid and fluid behaviour of cartilage tissue as set out in Chapter 3 (Section
3.3.2). The results indicated that the developed FE model is acceptable for studying
the solid and fluid behaviour inside the tissue.
4.5.1 Porohyperelastic FE model development for indentation test
After confirming the suitability of the modelling, the porohyperelastic FE model for
indentation tests was developed in ABAQUS 6.12, similar to the model developed in
Chapter 3. Axisymmetric elements were adopted to reduce the computational cost
based on the characteristics of the test sample and loadings. The FE mesh consisted
of 9600 4-node bilinear displacement and bilinear pore pressure elements. The
number of elements was decided based on the numerical tests performed earlier
(Chapter 3, Section 3.3.15). Large deformations and geometric nonlinearity were
considered in the calculation using the ‘NLGEOM’ option in ABAQUS 6.12.
Transient consolidation analysis with the ‘Full Newton’ solution procedure was used
to solve the equilibrium equations. The pore pressure (p) (p=0) boundary condition
was enforced on the upper surface of the portion—where the indenter does not touch
the cartilage surface—and on the right side of the cartilage so as to enable flow of
84 Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage
fluid through these boundaries. The surface-to-surface contact between the cartilage
and the indenter was modelled as frictionless.
The impermeable boundary between the indenter and cartilage does not require
special boundary conditions since it is a default boundary condition in ABAQUS
6.12 and has been previously used to model the contact between cartilage and rigid
indenters [185, 252]. Given that the stiffness of the indenter and bone is several
orders of magnitude higher than that of the cartilage, both were modelled as rigid
bodies. Preliminary investigation indicated that representing the bone and indenter as
rigid bodies did not affect the results of the simulations.
4.5.1.1 Solid skeleton material model
To account for the nonlinear large deformation, the solid skeleton was modelled as
an isotropic hyperelastic material. For isotropic hyperelastic materials, the decoupled
potential with linear bulk modulus (κ ) small or of the same order of magnitude to
that of the linear shear modulus (µ ) would impose compressibility on the material.
In fact, Simon and Kaufmann [257] and Ayyalasomayajula and Vande Geest [256]
also used similar formulations in their studies. Hence, this study used the decoupled
formulation by treating cartilage as an isotropic material. However, it is important to
mention that, in an anisotropic model, the use of the decoupled formulation would
generally yield inaccurate results [272, 273].
The lower-order material models such as the neo-Hookean or Mooney–Rivlin
were identified as incapable of representing the highly nonlinear stress-strain
behaviour observed during this study (Figure 4.3(a)). Higher-order models such as
the Yeoh model are considered more suitable for explaining the nonlinearity of
cartilage tissues [274]. However, the 2-term reduced polynomial hyperelastic model
gave an accurate description of the material behaviour for the cartilage samples
Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage 85
tested in this study (Figure 4.3(b)). Due to having few parameters, unlike models
such as the Yeoh model, the 2-term reduced polynomial model can be used to obtain
unique sets of material parameters. The hyperelastic material parameters were
obtained based on Eq. (3.14) developed in Chapter 3. The form of the 2-term reduced
polynomial model used in the present study is:
2
1
2
120110 )1(1)3()3( −+−+−= JD
ICICW e (4.16)
As mentioned earlier, here, eW is the isotropic elastic strain energy potential, 1I is
the first invariant of the distortional part C of the right Cauchy deformation
tensor C , and Fdet=J is the volume ratio. Furthermore, µ210 =C , where µ is the
shear modulus of linear elasticity, and 20C is a nonlinear stiffness parameter. In
addition, κ21 =D , where κ is the bulk modulus of linear elasticity.
Figure 4.3: (a) Experimental data from 10-2/s of a representative sample fitted to neo-Hookean, Mooney–rivlin and 2-term reduced polynomial incompressible hyperelastic functions; (b) R-squared values indicating the goodness of fit of neo-Hookean, Mooney–rivlin and 2-term reduced polynomial incompressible hyperelastic functions to the experimental data
86 Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage
4.5.1.2 Permeability variation with strain
Under compressive strain, the permeability ( k ) of cartilage decreases exponentially
(Eq. (4.17)) as mentioned earlier [191, 258, 259]. In Eq. (4.17), e is the void ratio
(i.e. the ratio of the volume of pores to the volume of solid), which is a quantity
representing dilatation, while 0e is the initial void ratio. The parameter 0k is the
initial undeformed state permeability and M and m are dimensionless material
parameters. The parameters 0e , M and m were chosen to be 4.0, 4.638 and 0.0848
[191], respectively, due to the extensive use of these values in the cartilage
literature. Eq. (4.17) is expressed as follows:
−
++
= 1
e1e1
2Mexp
eekk
2
0
m
00 (4.17)
During the study it was noted that the porohyperelastic model with strain-
dependent permeability was insufficient to capture the strain-rate-dependent tissue
behaviour (discussed in Section 4.7.2). It was postulated that the fluid behaviour
might significantly affect the tissue behaviour depending on the strain-rate/loading
velocity. Therefore, the permeability function that takes into consideration the effect
of the strain-rate was introduced, as detailed in the next section.
4.5.2 Permeability variation with strain-rate
The permeability of cartilage is not only dependent on strain, but is also a function of
applied pressure difference [259]. A higher pressure difference would result in
smaller permeability. This is due to the pressure drag forces that are developed when
the pressure difference is increased, which restricts the fluid movement [259].
According to Oloyede and Broom [149], as strain-rate increases, the pore pressure
inside the cartilage also increases. Therefore, an increase in pressure difference
Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage 87
inside and outside of the tissue is expected with an increase in strain-rate. This results
in a reduction of permeability. To the best of the author’s knowledge, a permeability
function that considers this phenomenon has not been reported to date. Nevertheless,
the following section re-analyses the data presented by Lai and Mow [259] and
Oloyede and Broom [149] and introduces a new mathematical relationship (Section
4.6) that represents the permeability variation with strain and strain-rate. In addition
to applied strain, the mathematical relationship illustrates that permeability decreases
with an increase in strain-rate.
The obtained permeability function, that is, the strain-rate-dependent
permeability, was included in the porohyperelastic framework and was compared
with the porohyperelastic models, which include constant permeability and strain-
dependent permeability, in order to comprehensively investigate the effect of fluid
behaviour on the sold-fluid interaction of kangaroo shoulder cartilage tissues.
Throughout the rest of this thesis, the porohyperelastic model with strain-dependent
permeability and the porohyperelastic model with strain-rate-dependent permeability
are referred to as the strain-dependent model and strain-rate-dependent model,
respectively.
4.6 EXTENSION OF POROHYPERELASTIC FIELD THEORY: STRAIN-RATE-DEPENDENT PERMEABILITY FUNCTION
Data from studies by Lai and Mow [259] and Oloyede and Broom [149] were used to
obtain the relationship along with certain approximations which are specified below.
Through experiments, Mow and Lai [96] found that permeability is a function of
strain and applied pressure difference ( P ). An exponential relationship for
permeability has been reported for infinitesimal strain [259]. This equation has been
further extended for large strains [191, 258].
The isotropic permeability tensor (k ) can be expressed as:
88 Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage
Ik k= (4.18)
where I is the unit tensor and k is the scalar valued permeability function.
The k can be represented by the following equation, where permeability is an
exponential function of J , the volume ratio [191, 275].
( )2/)1(exp 2 −= JaMk (4.19)
The data from Mow and Lai [96] can be fitted to Eq. (4.19) for each pressure
difference ( P ) (Figure 4.4(a)) to obtain the corresponding a and M values
depending on the applied strain and pressure difference. At a high value of P , due to
large pressure drag forces, fluid velocities become considerably small and hence the
fluid will be contained inside the tissue. This will reduce the tissue permeability to
almost zero. Therefore, at this pressure value, a and M will be almost zero. The
resulting variation of a vs M (Figure 4.4 (b)) can be approximated by a second-
order polynomial function in the form of Eq. (4.20), where α and β are empirical
constants:
2aaM βα += (4.20)
The P value at which coefficient a becomes almost zero is not available in
the literature. Therefore, we assumed this P to be a high physiological joint contact
pressure (e.g. contact pressure during high-speed running). Contact pressure inside
the knee during high-speed running (5-10.5 m/s) is approximately 3.5 times [91, 92]
t the static contact pressure. This assumes that the ground reaction forces are
proportional to the joint contact pressures. The reported static mean contact pressure
values in the knee are 2.75–3.79 MPa [276-278].
Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage 89
Therefore, P at which a becomes zero was taken as 13.625 MPa. The
variation of P vs a was approximated by a power relationship (Eq. (4.21) and
Figure 4.4(c)), whereg , δ and λ are empirical constants. In this derivation, we
have used simple functional forms that can represent the relationship between
different variables. Some researchers may prefer to use other functional forms which
are also correct. Eq. (4.21) is expressed as follows:
λg δ += Pa (4.21)
Figure 4.4: An exponential function (Eq. 4.19) fitted to Lai and Mow’s (1980) data; (b) Variation of coefficient M with coefficient a is approximated as a second-order polynomial function; (c) Variation of coefficient a with pressure difference (P) approximated as a power function
Re-analysis of Oloyede and Broom’s [149] data suggests that the internal pore
pressure of cartilage increases—under compression—with increasing strain and
strain-rate (Figure 4.5(a)). This increase in pore pressure values results in increased
90 Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage
pressure differences between the inside and outside of the loaded area of the tissue.
Assuming that the internal pore pressure is proportional to P , an approximate
relationship between P , the strain-rate and strain can be obtained (Eq. (4.22)):
2/1)1( DJP −= φ (4.22)
Here, φ is an empirical constant and D is the rate of deformation. The deformation
rate tensor, )(νD grad= (i.e. jiij xD ∂∂= υ ).
Combining Eq. (4.19) and Eq. (4.22), a mathematical relationship can be
obtained for permeability (Eq. (4.23)), where permeability is a function of strain and
strain-rate:
( )2/)1(exp)( 2 −+= Jaaak βα where λg δ += Pa (4.23)
The values of the empirical parameters obtained using the data of Lai and Mow [259]
and Oloyede and Broom [149] are: α =0.0827, β =0.1071, φ =350.27,λ =-6.399,
g =7.942, and δ =-0.0791. Based on the obtained parameters, we predicted the
variation of permeability in terms of strain and strain-rate as shown in Figure 4.5(b).
The permeability was found to be decreasing with both strain and strain-rate.
Figure 4.5: (a) Re-analysis of Oloyede and Broom’s [149] Variation of pressure difference (P) between the inside and outside of the tissue with strain-rate; (b) Variation of permeability with strain-rate as predicted by Eq. (4.23)
Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage 91
4.6.1 Material parameter identification
There were four material parameters, namely,10C ,
20C , 1D and0k , to be estimated
(assuming all of the above parameters are constant) in the current porohyperelastic
model. In order to obtain them, the following procedure was employed:
• Following the approach developed by Simon et al. [279], considering the
material as incompressible at the highest strain-rate (10-2/s), 10C and
20C were
obtained by fitting the force–indentation experimental data to Eq. (3.14)
derived in Chapter 3.
• Using 10C and
20C determined in the previous step, parameters 1D and
0k were obtained by fitting the experimental data on force–indentation at the
lowest strain-rate (10-4/s) to a porohyperelastic FE model, considering the
material to be compressible.
The obtained parameters were used to predict the strain-rate-dependent
behaviour of the cartilage tissues and were compared with the experimental results.
Based on the performance of the three porohyperelastic models—constant, strain-
dependent, and strain-rate-dependent—the effect of solid–interstitial fluid interaction
on strain-rate-dependent behaviour was evaluated.
4.7 RESULTS AND DISCUSSION
4.7.1 Biomechanical parameters of kangaroo shoulder cartilage
The 2-term reduced polynomial hyperelastic function fitted well to high strain-rate
data (R2=0.9890±0.0044, p<0.000). The obtained stiffness parameters (i.e. 10C
and20C ) were 0.0988±0.0622 MPa and 0.1482±0.061 MPa, respectively. The FE
porohyperelastic model also fitted well to the low strain-rate data
(R2=0.9855±0.0098, p<0.000). The obtained compressibility parameter ( 1/1 D ) and
92 Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage
the permeability (0k ) in an undeformed configuration were 0.0782±0.055 MPa and
1.32±0.98 x 10-14m4/Ns, respectively.
The hyperelastic material parameters for human shoulder cartilage tissues have
not been reported elsewhere. Nevertheless, considering the reported Poisson’s ratio
for shoulder cartilage tissue, which is approximately 0.15 [226, 280], Young’s
modulus (E) for kangaroo shoulder cartilage was estimated to be 0.454±0.286 MPa.
The calculated Young’s modulus (Eq. (4.24)) for human shoulder cartilage using
Huang et al.’s [177] and Mow et al.’s [28] modulus in uniaxial strain (HA) were 0.214
MPa and 0.624 MPa, respectively. Eq. (4.24) is expressed as follows:
υ)υ)((υ)E(H A 211
1−+
−=
(4.24)
The Young’s modulus of kangaroo shoulder cartilage was not significantly
different (p=0.109) from the value calculated in Mow et al.’s [28] study; however, it
was significantly different (p<0.05) from the value calculated in Huang et al.’s [177]
work. Nevertheless, the average value falls within the values calculated for human
shoulder cartilage. The average thickness of the kangaroo shoulder cartilage samples
obtained through ultrasound measurements was 0.72±0.10 mm. The reported average
thickness value for human shoulder cartilage is 1.44 mm [174], which is higher than
that of kangaroo cartilage (p<0.005). Considering the aforementioned differences in
thickness and potential differences in the tissue composition of different species we
would consider the average E value obtained in this study to be acceptable. The
permeability value of the kangaroo humeral head cartilage was relatively low, but
was not significantly different (p=0.145) from the value reported for the central
region of human humeral head cartilage, which is 1.82±1.27x10-14 m4/Ns [177].
Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage 93
The biomechanical properties and behaviour of kangaroo shoulder cartilage
were in general agreement with that of human shoulder cartilage, indicating that
kangaroo can be considered as a potential animal model for shoulder cartilage
research in the future. However, extensive experimentation (histological,
biochemical assessment etc.) to investigate the development and degradation of this
tissue is required in order to confidently use kangaroo as an animal model to
investigate the pathologies related to shoulder cartilage.
4.7.2 Comparison of constant, strain-dependent and strain-rate-dependent model predictions
The average experimental stress-strain data were compared with the porohyperelastic
FE prediction, for constant, strain-dependent and strain-rate-dependent models
(Figures 4.6(a), 4.6(b) and 4.6(c), respectively). In general, all the models were
sensitive to the effect of strain-rate, indicating the ability of the poromechanics
framework [182, 195, 281] to capture the strain-rate dependency. The models with
strain-rate-dependent and strain-dependent permeability outperformed (p<0.05) the
model with constant permeability at all strain-rates (Figure 4.6(d)). At intermediate
strain-rates, statistically significant differences (p<0.05) were identified between the
constant-permeability-model and the predictions of the other two models. This
difference was even more significant when a similar comparison was made at the
highest strain-rate (p<0.005), that is, at 10-2/s.
The predictions of the strain-rate-dependent model were better than the
predictions of the strain-dependent model at all strain-rates (Figure 4.6(d)). However,
statistical differences were not identified between the predictions of the strain-
dependent and strain-rate-dependent models at 5x10-3/s (p=0.179), although at 5x10-
4/s they were significantly different (p<0.05). Nevertheless, compared to the strain-
dependent model (R2=0.7937±0.1478), the strain-rate-dependent model
94 Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage
(R2=0.8915±0.0662) was significantly better at predicting the stress-strain variation
at the highest strain-rate (p<0.005).
Figure 4.6: Comparison of constant, strain-dependent and strain-rate-dependent model prediction to average (n=10) experimental data of the samples tested – (a) Constant permeability; (b) Strain-dependent permeability; (c) Strain-rate-dependent permeability; (d) Model predictions in terms of R-squared (R2) values and the corresponding significant differences among constant, strain-dependent and strain-rate-dependent models at individual strain-rates
4.7.3 Effects of strain-dependent and strain-rate-dependent permeability
In comparison with the model with constant permeability, strain-dependent
permeability takes into account the shrinkage of pores (Figures 4.7(a) and 4.7(b)) and
its concomitant effect on permeability during tissue deformation. Due to the
reduction in pore size (i.e. reduction of effective flow area) with deformation, it
becomes difficult for fluid to move out of the tissue. The reduction of tissue
Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage 95
permeability leads to an increase in solid–interstitial fluid frictional interactions. This
could be the reason why both the strain-dependent and strain-rate-dependent models
significantly outperformed the FE model with constant permeability (p<0.05) at all
strain-rates.
Figure 4.7: An ideal representation of part of a tissue with pores represented by circles (undeformed) and ellipse (deformed); (a) Constant permeability – Pore volume/effective fluid-flow area does not change; (b) Strain-dependent permeability – Pore volume/effective flow area is reduced due to application of strain (ε); (c) Strain-rate-dependent permeability – Large pressure differences due to suddenly applied load (Tt2<<<Tt1) result in larger drag forces; this will compact the tissue to reduce the pore size (indicated by the red hatched area), creating congestion for fluid particles to move through pores and, therefore, the fluid particles experience a reduction of pore size/effective flow area
As noted above, when the effect of strain-rate was considered, the effect of
solid–interstitial fluid interaction on strain-rate-dependent behaviour was more
significant. In two cases, namely, (b) (Figure 4.7(b)) and (c) (Figure 4.7(c)), the
tissue was deformed to a strain ε at a time Tt1 and Tt2 (<<<Tt1), respectively. In case
(c), due to the sudden application of strain, fluid may experience large pressure
differences in comparison to case (b). Due to this, in case (c), there will be a rush of
fluid to move away from the deformed areas to other parts of the tissue. Meanwhile,
due to high pressure difference there will be higher pressure drag forces inside the
tissue, leading to compaction of the tissue matrix [183, 239, 259]. The compaction of
the cartilage tissue matrix was theoretically predicted and explained by Lai and Mow
96 Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage
[259]. Due to tissue compaction, in addition to the effect of strain, the pore size also
further decreases; hence, fluid particles will experience a reduced effective flow area
(Figure 4.7(c)). The compaction coupled with the rush of the fluid to move out of the
deformed area creates more congestion for the fluid particles moving through the
pores in case (c), reducing the permeability as predicted by the strain-rate-dependent
permeability. Therefore, when the strain-rate is increased from low to high, at a
certain strain-rate, the pressure drag force may start to affect the tissue behaviour
significantly. This could be the reason why the strain-rate-dependent model
performed significantly better in capturing the experimental results at the highest
strain-rate tested (p<0.005) compared to the strain-dependent model. Therefore, in
addition to the effect of strain on permeability, in order to better explain the tissue
behaviour at high strain-rates—where pressure drag forces become significant—the
effect of strain-rate on permeability can be taken into account in future FE models.
4.7.4 Mechanisms underlying the strain-rate-dependent tissue behaviour
Based on the above observations, this section summarises the solid–fluid frictional
interactions and the pressure drag forces as some of the possible mechanisms
underlying the strain-rate-dependent behaviour of kangaroo shoulder cartilage
tissues.
Solid–fluid frictional interactions: In the present study, fluid was considered to be
inviscid, that is, its viscosity manifests itself only in the fact that there is a non-zero
resistance to fluid flow, which suggests that the permeability is not infinite.
However, since there is no other viscous effect, the fluid can only bear hydrostatic
Cauchy stresses (i.e. no viscous shear stress in the fluid). Therefore, the solid–fluid
frictional drag forces depend on the tissue permeability. The model with constant
permeability had a finite permeability value and was sensitive to strain-rate—thus,
Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage 97
indicating the presence of frictional interaction between the solid and fluid. The
reduction of permeability with strain will increase the magnitude of frictional
interactions. This was reflected by the increased sensitivity of the model with strain-
dependent permeability to strain-rates. Therefore, it was concluded that the frictional
interactions between solid and fluid can be stated as one of the mechanisms
facilitating the strain-rate-dependent behaviour.
Figure 4.8 : Comparison of pore pressure and velocity profiles at 10-2/s – (a) Strain-dependent permeabilty; (b) Strain-rate-dependent permeability; (c) Fluid velocity at the bottom left (point P) of the cartilage matrix
Pressure drag forces: The strain-rate-dependent model showed higher pore pressure
values than the strain-dependent model (Figures 4.8(a) and 4.8(b)). The smaller fluid
velocities observed in the former model reflected the effect of the higher drag forces
and generated due to the higher pressure difference (Figure 4.8(b)). Hence it can be
stated that that, at high strain-rates, the strain-rate-dependent permeability enhances
the fluid pressurisation. This enables the tissue to respond to large strain-rates more
98 Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage
effectively, such that excessive deformation of the tissue is minimised. Therefore, the
strain-rate-dependent fluid behaviour can be stated as one of the mechanisms that
support the load-bearing of cartilage tissue and is significantly prevalent at large
strain-rates.
4.7.5 Role of cartilage as a protective layer at large strain-rates
As mentioned previously (Section 4.7.3), at 5x10-3/s strain-rate, the strain-dependent
model prediction was not significantly different from the strain-rate-dependent
model, but when the strain-rate was increased to 10-2/s it became significantly
different. Similarly, Oloyede and Broom [149] observed a significant increase in the
effective stress of the cartilage solid skeleton in comparison to the pore pressure
increase when the strain-rate was increased from 10-3/s to 10-2/s. This is believed to
be due to the same phenomenon of the increase in pressure drag, which reduces the
fluid movement at high strain-rates. Also there is a possibility that inertia forces
might begin to affect the fluid behavior at very large strain-rates which could further
impede fluid movement inside the tissue. The ability of cartilage tissue to contain the
fluid inside can be attributed to its small pores that are in the range of 20–65 Å in
their undeformed state [100, 202, 282]. The pore size of the tissue in the undeformed
state was calculated based on the formulation described in the next section.
4.7.5.1 Pore size calculation based on permeability
The network of pores in the cartilage tissue determines the fluid behaviour and
permeability. The pore size can be considered as a microstructural parameter of the
tissue. In order to understand the fluid behaviour inside the tissue, the pore sizes
were calculated and matched with the tissue behaviour. The calculation was based on
Maroudas’s [100] study as summarised next.
Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage 99
Consider cartilage as a porous material, comprising of network of pores with an
average pore radius (r) and tortuosity (τ ). The path of the fluid movement due to
pressure difference ( dP ) in cartilage tissue is likely to be tortuous. The path of the
fluid movement in a porous structure is also likely to be tortuous. Therefore, the true
pressure gradient, accounting for tortuosity, can be written as:
e
e
dLdP
LL
dLdP
= , where P is the pressure (4.25)
Here, LLe is the ratio between the true path length and the Darcian length. Applying
the Poiseuille equation to the pore capillaries, the velocity of flow ( fv ):
e
f LL
ddPr
ddPrv
L 8L 8
2
e
2
−=
−=
ηη (4.26)
Here, η is the viscosity of the fluid. The velocity of the flow is related to the fluid
flux ( q ) and porosity or the fractional water content ( H ) by:
LL
Hqv e
f = (4.27)
Combining Eq. (4.26) and Eq. (4.27), the fluid flux is:
dLdP
LLHrq
e
22
8
−=
η (4.28)
By comparing Eq. (4.28) with Darcy’s law, the following equation for
permeability can be obtained:
100 Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage
ητ 81
2
2
=
Hrk ;LLe=τ (4.29)
In this study, the volume fraction of the water )(H , tortuosity of the flow path(τ ) and
fluid viscosity )(η were taken as 0.8, 1.4 and 10-3 Ns/m2, respectively, considering
the values reported in the literature [100].
Based on Eq. (4.28) in the above formulation, the pore size of the tissue in an
undeformed state was 154.55±46.1 Å. The pores which are 10–50 times the size of a
water molecule (approximately 3 Å) reduces further under deformation. Hence,
depending on the pressure gradient at different strain-rates, the movement of the fluid
inside will be reduced. This containment of fluid inside the tissue at large strain-rates
will provide protection to the cartilage and underlying bone by reducing the
excessive deformation of tissues. Therefore, the reduction of permeability with the
strain-rate plays an important role at high strain-rates. However, in the case of
osteoarthritis, for which the proteoglycan content is reported to be low, the pore size
of the tissue can be relatively high and thus can interrupt this phenomenon. This will
result in excessive deformation of the tissue in comparison to a healthy tissue, and
thus increase the risk of bone-to-bone contact and tissue damage.
4.7.6 Limitations of the strain-rate-dependent permeability model and possible improvements to the FE porohyperelastic model
It is important to assess the limitations and assumptions under which the
mathematical expression of strain-rate-dependent permeability has been formulated
and the possible implications of the assumptions. Currently, there are no practical
methods available to directly measure the fluid flow rates inside cartilage tissue
under different loading-rates. The indirect method, as employed in the present study,
is to use fluid pressure measurements inside the tissue under different strain-rates and
Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage 101
relate them to available static permeability measurements under different pressure
gradients. In the static permeability measurement, a pressure gradient is imposed on
already compressed cartilage and the amount of fluid outflow is measured at
equilibrium. When the tissue is compressed at a certain loading-rate, the fluid
pressure in the loaded area will increase and there will be a pressure gradient
between the loaded area and outside which happens on a timescale smaller than that
of static measurements. Additionally, during compression, fluid pressure in the
loaded area continuously changes and is non-uniform. Therefore, static permeability
measurements extracted at different pressure gradients do not represent the actual
conditions ideally and hence may affect the permeability values predicted by the
strain-rate-dependent permeability model. However, the extent to which the
assumptions affect the strain-rate-dependent model predictions requires more
investigations which could be potentially carried out in the future if a methodology
can be devised to measure the fluid velocities inside cartilage under different
loading-rates..
The strain-rate-dependent model did not fit well to the stress-strain response at
the highest strain-rate (10-2/s) and some low strain-rates such as 5x10-4/s. Although
the main focus of the present study was not to present a comprehensive FE model to
predict the experimental data, this can be stated as one of the limitations of the strain-
rate-dependent model. Given that in the present study we have fully evaluated the
isotropic fluid and solid behaviour, the model could be possibly improved by
incorporating anisotropic fluid behaviour, anisotropy of the collagen network and/or
the viscoelasticity of the tissue. Earlier studies have shown that fluid pressurisation is
enhanced by anisotropy of the elastic properties of the tissue [156, 283].
Furthermore, the anisotropy of cartilage permeability due to the glycosaminoglycan
102 Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage
network deformation [275, 284, 285] and the collagen fibre orientation [272, 286-
290] affects the fluid pressurisation at both small and large strains. Additionally, as
mentioned earlier, there is evidence that the flow-independent viscoelasticity of the
cartilage matrix affects the strain-rate-dependent behaviour of the tissue at small and
large strain-rates [146, 150, 156]. Hence, the lower stresses predicted by the strain-
rate-dependent model at low strain-rates, for example at 5x10-4/s, might be improved
by the inclusion of matrix viscoelasticity in the model.
4.8 CONCLUSION AND REMARKS
In this study, by introducing kangaroo as an animal model we have explored the
strain-rate-dependent mechanical behaviour and the underlying mechanisms of
shoulder cartilage. By introducing the strain-rate-dependent permeability model and
comparing the model’s predictions with constant and strain-dependent models, the
present study explored how the solid–interstitial fluid interaction facilitates the
strain-rate-dependent behaviour of shoulder cartilage tissues and its physiological
relevance. The following conclusions were made based on the results of the current
study:
• Kangaroo can be considered as potentially suitable animal model for future
biomechanical research on shoulder cartilage. This is because the
biomechanical properties and behaviour of kangaroo cartilage tissues are in
general agreement with that of human shoulder cartilage tissues.
• Further, the different loadings encountered by the upper and lower limb
cartilage of kangaroo provide a natural source for investigating how
mechanical forces affect the development, composition (e.g. proteoglycan
distribution) and structure (e.g. collagen architecture) of cartilage, and the
progression of osteoarthritis. In addition, experimentations on this animal
Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage 103
model may also have the potential to give insights into how tissue-
engineering strategies must be adjusted to develop joint-specific tissues.
• The mechanical behaviour of kangaroo shoulder cartilage is strain-rate-
dependent and nonlinear. Its solid skeleton behaviour can be adequately
represented by the 2-term reduced polynomial hyperelastic model. However,
lower-order material models such as the neo-Hookean and Mooney–Rivlin
models were found to be inadequate to explain the nonlinear solid skeleton
behaviour of the tissue.
• In addition to strain, permeability has a dependency on strain-rate: it
decreases when the strain-rate is increased.
• Both constant and strain-dependent models are sensitive to strain-rates.
Further, strain-dependent models were found to outperform the constant
permeability model. Therefore it can be said that solid–fluid frictional
interaction is one of the main reasons for strain-rate-dependency.
• This study found that both strain-dependent and strain-rate-dependent models
significantly affect the tissue behaviour. At high strain-rates, the latter model
becomes more significant than the former. Therefore, it can be concluded that
at high strain-rates in addition to solid–interstitial frictional fluid interaction,
pressure drag forces and possibly inertia forces begin to play a significant
role in the tissue behaviour.
• Based on an earlier study, where a transition of tissue behaviour was
observed at 10-2/s [149], we postulate that this phenomenon could be due to
the small pore size of the cartilage (in the order of 10-9m) and its size
reduction under deformation. The pores facilitate the ability of cartilage
tissues to contain the fluid within the matrix at large strain-rates, and thereby
104 Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage
to effectively reduce excessive deformation. This assists the tissues to
function as a protective layer for bone-ends during injurious loads at high
strain-rates.
• Based on the findings of this study, it can be concluded that FE models will
also benefit from the inclusion of strain-rate-dependent permeability to better
predict the cartilage response.
• Since all aspects of the isotropic fluid and solid behaviour were evaluated in
the present study, the model deviations at the highest strain-rate can be
attributed to the anisotropic solid, fluid properties and viscoelasticity of the
matrix.
The introduction of kangaroo as a model for shoulder cartilage investigation, the
mathematical expression for strain-rate-dependent permeability and the strain-rate-
dependent FE model employed in the present study provide insights and open new
avenues for investigating the mechanisms underlying the strain-rate-dependent
mechanical behaviour of cartilage tissues. However, since the FE model developed in
this study was unable to fully explain the tissue behaviour, the factors affecting the
strain-rate-dependent behaviour of kangaroo shoulder cartilage tissue were further
analysed as reported in the following chapter.
Chapter 4: Effect of interstitial fluid on the strain-rate-dependent behaviour of kangaroo shoulder cartilage 105
Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage
This chapter begins with a brief introduction to the investigations reported in the
literature on the role of cartilage tissue constituents on mechanical behaviour
(Section 5.1). The aims and objectives of this part of the study and the hypotheses are
then discussed (Sections 5.2 and 5.3, respectively). The experimental methodology is
set out (Section 5.4), followed by the results and discussion (Section 5.5). Lastly, the
conclusions and recommendations for follow-up studies are provided (Section 5.5).
The study in this chapter resulted in a journal article named “Physical mechanisms
underlying the strain-rate-dependent mechanical behavior of kangaroo shoulder
cartilage” in the Journal of Applied Physics letters [214].
5.1 INTRODUCTION
In Chapter 4 it was noted that the porohyperelastic model with strain-rate-dependent
permeability was unable to capture the tissue behaviour at the highest strain-rate
tested. Further, we concluded that the differences identified between the
experimental data and FE model prediction could, to an extent, be due to anisotropy
of the solid skeleton. Meanwhile, the literature has indicated that superficial collagen
considerably affects the tissue behaviour at large strain-rates [16]. Therefore, in order
to comprehensively understand the factors affecting its mechanical behavior of
kangaroo shoulder cartilage, it is important investigate how the cartilage extracellular
matrix components affect the strain-rate dependent behaviour. Hence, the
Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage 107
investigations in this chapter focused on the effect of cartilage constituents (i.e.
proteoglycans and collagen network) on the strain-rate-dependent mechanical
behaviour of kangaroo shoulder cartilage tissue.
The dynamic properties of cartilage (extracted at high strain-rates) are known
to be governed by the structure of the collagen network [20, 158]. Based on an FE
model that considered the cartilage structure and composition, Julkunen and Jurvelin
[16] showed that superficial collagen can considerably affect the tissue behaviour at
physiologically high strain-rates (10-1/s in their study). In contrast, the equilibrium
properties of cartilage (extracted at very low strain-rates) are mainly affected by
proteoglycans [17, 20]. It is also well accepted that the compressive properties of
cartilage are governed by the hydrated proteoglycans constrained by the collagen
network [98]. However, the extent to which the superficial layer and proteoglycans
affect the strain-rate-dependent behaviour of shoulder cartilage tissues has not been
investigated thoroughly.
While conclusions from the above studies have been made for knee cartilage in
particular, we believe they cannot be generalised to all joint cartilages due to possible
differences in the composition and microstructure of tissues which are regulated by
the different mechanical environments experienced by the tissues. Chondrocytes
dynamically synthesise the extracellular matrix (i.e. proteoglycans and collagen)
based on the external loading stimuli they receive [163-165]. Therefore,
proteoglycan composition and structural features of the collagen network potentially
adapt to external mechanical stimuli, and hence depend on the local mechanical
environment of the tissue [86, 166-171]. The conclusions of reported studies [16, 17,
20, 158] which are predominantly for knee cartilage should therefore be evaluated in
the context of the specific tissue being studied. As shoulder cartilage experiences
108 Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage
considerably less compressive loading, it was postulated in this study that the
collagen network (including the superficial layer) plays a more significant role in
facilitating the strain-rate-dependent behaviour of the shoulder cartilage than
proteoglycans. In order to test this belief, we carried out a series of experiments
which involved simultaneous artificial degradation of cartilage constituents
(proteoglycan and superficial collagen) and mechanical testing.
Artificial degradation through enzyme treatment is commonly used to model
proteoglycan loss and superficial collagen damage [20, 22]. The main advantage of
artificial degradation is that the level of damage to the tissue can be controlled by
enzyme concentration, enzyme type and the duration of the exposure [22]. Hence,
gradual degradation of the constituents and simultaneous assessment of the
mechanical properties can give direct insight into the contribution of proteoglycans
and superficial collagen to the strain-rate-dependent behaviour. Since enzymatic
treatments have also been used to mimic some of the characteristics of osteoarthritis
(e.g. superficial collagen degradation and proteoglycan depletion) [19, 291], the
results of this study can also assist to evaluate the effect of osteoarthritis on the time-
dependent behaviour of shoulder cartilage tissues. Because the surface lipid layer of
cartilage is also known to be affected during the early stages of osteoarthritis [292] ,
we also performed an experiment to assess the effect of surface layer on time-
dependent tissue behaviour. Since damage to the superficial collagen may also affect
the surface lipid layer, this experiment helped to further confirm the contribution of
the superficial collagen to tissue behaviour.
5.2 AIMS AND OBJECTIVES
This part of the study was mainly focused on achieving the following objectives:
Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage 109
1) Investigate the role of proteoglycan in the strain-rate-dependent mechanical
behaviour of kangaroo shoulder cartilage.
2) Investigate the role of superficial collagen and surface lipid layer in the
strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage.
3) Assess whether the proteoglycan or the superficial collagen dominates the
strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage.
5.3 HYPOTHESES
As mentioned in the introduction to this chapter, this part of the study aimed to test
the following two main hypotheses:
1) Superficial collagen plays a more significant role in facilitating the strain-
rate-dependent behaviour of the shoulder cartilage than proteoglycans.
2) The contribution of superficial collagen to tissue behaviour at high strain-
rates is significantly larger than at small strain-rates.
5.4 EXPERIMENTAL METHODOLOGY
Indentation on a cartilage sample at four strain-rates (10-4/s, 5x10-4/s, 5x10-3/s and
10-2/s) requires one day for completing the testing procedure. However, during the
experimental design stage of this study, it was noted that each sample testing can
take more than a day to complete. Therefore, it was essential to design the
experimental procedure in such a way that ensured the least possible effect on the
tissues due to sample preservation. There were two methods for preserving the
samples. The first method would be to preserve the samples in a PBS-inhibitor
solution at 4 °C until the experimentations were completed. The second method
would be to immediately freeze the samples after harvesting in a PBS-inhibitor
solution and then to thaw the samples in PBS at room temperature (24–27 °C) for 30
minutes before mechanical testing. However, these methods may affect the
110 Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage
mechanical properties of the tissue. Therefore, in order to assess the effect of the
preservation method on the mechanical properties of the tissue, preliminary studies
were conducted as mentioned in the next section.
5.4.1 Assessment of tissue preservation methods: The PBS-solution at 4 °C vs the multiple freeze–thaw method
5.4.1.1 Effect due to preservation of sample in PBS-inhibitor solution at 4 °C
Ten visually normal (ICRS [215] macroscopic score=0) cylindrical kangaroo
cartilage plugs of 8 mm diameter with 2–3 mm subchondral bone were harvested
from the near central load-bearing area of the humeral head. The thicknesses of the
samples were estimated using ultrasound measurements. Then, indentation testing
was carried out on individual samples using a plane-ended, polished indenter of 3
mm diameter with a rounded edge of 0.1 mm radius at four different strain-rates: 10-
4/s, 5x10-4/s, 5x10-3/s and 10-2/s. The samples were indented up to 25% engineering
strain throughout the study and the same experimental procedures described in
Chapter 3 (Section 3.2.6) were used. After mechanical testing, the samples were
divided into two groups. The initial experimental design indicated that three days
were required to complete the study on the proteoglycans (Section 5.4.2.1 presents
the details on the proteoglycan study). Therefore, samples from the first group were
kept for three days in a PBS-inhibitor solution at 4 °C prior to subsequent mechanical
testing, while samples from the other group were kept for one week. Based on the
force–indentation results, the mechanical properties were extracted by fitting to Eq.
(3.14) in Chapter 3 and the effect of the preservation at 4 °C on mechanical
properties was assessed (Figure 5.1).
Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage 111
Figure 5.1: (a) Mechanical properties after 72 hrs in PBS-inhibitor solution at 4 °C; (b) Mechanical properties after 1 week in PBS-inhibitor solution at 4 °C
5.4.1.2 Effect of sample preservation methods: Multiple freeze–thaw cycle
In order to assess the effect of freeze–thaw cycles on the mechanical properties and
behaviour of kangaroo shoulder cartilage, mechanical tests were conducted on five
samples immediately after harvesting from the central load-bearing area of the
humeral head. Then, the samples were put in containers filled with PBS-inhibitor
solution and frozen at -20 °C overnight. The following day, the frozen samples were
thawed for 30 minutes in the PBS solution at room temperature (24–27 °C), after
which the mechanical tests were conducted. This entire procedure was repeated
thrice in the next three days. Samples were tested at one strain-rate (10-2/s); based on
the literature, this was considered enough to ascertain the effect of the freeze–thaw
cycles on the mechanical properties of the tissue [217]. After testing, Young’s
modulus was extracted from the force–indentation results and the effect of the
multiple freeze–thaw cycles on the tissue properties was assessed (Figure 5.2).
112 Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage
Figure 5.2 : Mechanical property change due to multiple freeze thaw cycles
The results (Figure 5.1(a)) of the above experiments indicated that the
mechanical properties were not significantly affected by keeping the samples at 4 °C
in PBS-inhibitor solution for three days (p>0.1). However, the samples that were
kept in the PBS-inhibitor solution for one week (Figure 5.1(b)) showed a significant
decrease in mechanical stiffness at small strain-rates (i.e. at 10-4/s and 5x10-4/s)
(p<0.05). It is noteworthy that these results are similar to findings in previously
reported studies on other cartilage tissues [217]. The samples that went through the
multiple freeze–thaw cycle (i.e. four times) also did not show (Figure 5.2) any
significant change in mechanical properties (p>0.2). On the one hand, some studies
have reported that multiple freeze–thaw cycles may degrade the biomechanical
properties and composition of cartilage [293-295]. On the other hand, some studies
have stated that the decrease in mechanical properties is due to the thawing
procedure employed [217]. It has also been suggested that the sample size and
species may influence the subsequent effect of freeze–thaw cycles [217].
Considering the lack of consensus on this issue in the literature, the multiple freeze–
thaw cycle method was not used in this study to preserve the tissues. Instead, the first
Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage 113
preservation method which kept the cartilage sample in PBS-inhibitor solution at 4
°C was chosen.
5.4.2 Proteoglycan, superficial collagen degradation and surface delipidisation
Visually normal cartilage plugs (ICRS macroscopic score=0) of 8 mm diameter were
harvested from the central load-bearing area of the humeral head. After obtaining the
thickness values using ultrasound measurements, before and after the proteoglycan
and collagenase degradation, the samples were subjected to mechanical testing under
four strain-rates (10-4/s, 5x10-4/s, 5x10-3/s and 10-2/s) to check how the progressive
degradation of the cartilage components affects the strain-rate-dependent mechanical
behaviour and the properties of the tissue. The results of the five samples tested
earlier (i.e. the samples tested after being kept in PBS-inhibitor solution for 72 hrs at
4 °C) were considered as the control test for these studies. Hence, additional control
studies were not conducted.
5.4.2.1 Proteoglycan degradation: Trypsin-PBS (phosphate buffered saline) solution
Several enzymes such as trypsin, chondroitinase ABC, cathepsin D and elastase have
been used to remove proteoglycans from cartilage tissues. Chondroitinase ABC [19,
296] and trypsin [291, 297-300] are the most commonly used enzymes for this
purpose. Chondroitinase ABC acts on proteoglycans to degrade chondroitin 4-
sulfate, chondroitin 6-sulfate and dermatan sulfate, and hyaluronate slowly. Trypsin
can cleave peptides on the C-terminal side of lysine and arginine residues of
proteoglycans. For the present study, trypsin was chosen to degrade the
proteoglycans due to its availability and extensive use in cartilage-related research.
Firstly, 100 ml of 0.01M PBS (P4417-100TAB, Sigma-Aldrich, Castle Hill,
NSW, Australia) solution (pH 7.4) which contained 1 ml of L-glutamine–penicillin–
114 Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage
streptomycin solution was mixed with 0.1 mg trypsin (from bovine pancreas, T1426,
Sigma-Aldrich, Castle Hill, NSW, Australia). The obtained 0.1 mg/mL trypsin–PBS
solution was then diluted to make 0.05 mg/ml. After the first cycle of mechanical
testing on a cartilage sample, it was treated with 0.05 mg/ml of trypsin under 37 °C
for 1 hr in an incubator. Afterwards, the sample was removed from the solution for
further mechanical tests (Figure 5.3). The cartilage sample was again treated for 1 hr
in trypsin before indentation testing and then again for 2 hrs before testing in order to
gradually degrade the proteoglycans, followed by mechanical testing to assess the
effects of 2 hrs and 4 hrs of trypsin–PBS treatment on the mechanical properties and
behaviour of the tissue. All samples (n=10) were tested as above and were always
preserved in PBS-inhibitor solution at 4 °C throughout the experiment.
Figure 5.3: Steps in sequential trypsin treatment (0.05 mg/ml) and mechanical testing on kangaroo shoulder cartilage samples
Treatments of cartilage samples for 1 hr, 2 hrs and 4 hrs in trypsin have shown
the need to remove the proteoglycans gradually in a wavefront manner through the
depth of cartilage [22, 301]. Noticeable variation in the amount depleted has also
been observed depending on the initial proteoglycan concentration [22]. In order to
investigate the manner in which the above-mentioned trypsin treatment degrades the
proteoglycan in kangaroo shoulder cartilage, a preliminary histological study was
conducted. Three cartilage samples were harvested from near the central load-
bearing area of the humeral head of adult red kangaroos. The samples were
cryosectioned to 10 μm and were stained for proteoglycans using 0.1% safranin-O
Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage 115
(details of the cryosectioning procedure and staining protocol are set out in Chapter
6, Sections 6.4.1 and 6.4.2). Then, the samples were subjected to 1 hr, 2 hrs and 4 hrs
of trypsin treatment according to the above-mentioned protocol. After each trypsin
treatment, the samples were cryosectioned and stained in safranin-O in order to
assess the proteoglycan depletion. The safranin-O stains proteoglycan in red/orange
colour (Figure 5.4(a)) and the stain intensity is proportional to the concentration of
proteoglycans.
The results of the histology study indicated that treatment of the samples for
one hour in trypsin removed almost all the proteoglycans in the superficial zone and
in part of the middle zone (Figure 5.4(c)). Further treatment for one hour removed a
substantial amount of proteoglycans in the deep zone (Figure 5.4(c)) while 4 hrs of
trypsin treatment removed almost all the proteoglycans (Figure 5.4(d)). The gradual
removal of proteoglycans in the form of a wavefront was also observed (as can be
seen in a comparison of Figures 5.4(a), (b), (c) and (d)). These results are consistent
with the reported findings in previous studies [22, 301].
The selective removal of primary proteoglycans through trypsin treatment for a
smaller period of time (<8 hrs) has been reported not to affect the structural
appearance of the collagen network [302]. There is limited evidence to suggest that
selective proteoglycan removal significantly affects the collagen network [21, 112,
302]. However, there is a possibility that trypsin may attack collagen molecules
which are already cleaved. This is most likely to be minimal due to the small
exposure time (≤ 4 hrs), but the inability to control the effect of trypsin specifically
should be noted as a possible limitation of this study.
116 Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage
Figure 5.4: Safranin-O staining of cryosectioned samples harvested from near the central load-bearing area of the humeral head – (a) Untreated sample; (b) 1 hr trypsin-treated sample; (c) 2 hr trypsin-treated sample; (c) 4 hr trypsin-treated sample
5.4.2.2 Superficial collagen degradation: Collagenase solution
Collagenase isolated from clostridium histolyticum is commonly used to
degrade the collagen of cartilage tissues. Studies commonly use 30 U/ml
concentration of collagenase solution for 24 hrs [237] to partially disrupt the
superficial collagen or for 40–44 hrs [19-21] to significantly damage it. It has been
identified that this collagenase treatment results in minor proteoglycan loss in the
superficial zone and middle zone of cartilage, possibly due to the diffusion of the
proteoglycans through the damage to the collagen network [19, 20]. However, some
studies have reported that proteoglycan depletion is histologically not visible when
treated with collagenase for 40 hrs [303].
Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage 117
A 3KU bottle of collagenase (from Clostridium histolyticum, C0773, Sigma-
Aldrich, Castle Hill, NSW, Australia), 100 ml of 0.01M PBS solution (pH 7.4), and 1
ml of L-glutamine–penicillin–streptomycin solution were mixed together to make 30
U/ml collagenase solution in the current study [19-21]. After the first cycle of the
mechanical tests, samples (n=12) were immersed in this collagenase solution and
placed in an incubator at 37°C for 44hrs. After the first cycle of mechanical tests, the
samples (n=12) were immersed in this collagenase solution and placed in an
incubator at 37 °C for 44 hrs. After this treatment, further mechanical tests were
carried out to assess the effect of the collagenase on the mechanical properties and
behaviour of the tissue (Figure 5.5). Since testing required more than one day, all the
samples were preserved overnight in PBS-inhibitor solution at 4 °C.
Figure 5.5: Steps carried out to investigate the effect of superficial collagen on the strain-rate-dependent behaviour of kangaroo shoulder cartilage
In order to check the effect of collagenase treatment on proteoglycans, a histological
study was also conducted on the cryosectioned samples harvested from three
kangaroo humeral heads. Safranin-O staining was conducted before (Figure 5.6(a))
and after collagenase treatment (Figures 5.6(b), (c) and (d)) and did not show a
noticeable effect on the proteoglycans. However, in order to further confirm these
results, an alcian blue test was conducted to detect any proteoglycans leaching out of
the matrix due to the collagenase treatment. Details of the alcian blue test are
presented in the next section.
118 Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage
Figure 5.6: Safranin-O staining of (a) untreated samples; (b), (c), (d) collagenase-treated samples 1, 2, 3, respectively
5.4.2.3 Alcian blue test: Effect of collagenase treatment on proteoglycans
Five kangaroo cartilage samples harvested from the central load-bearing area of the
humeral head were incubated at 37 °C for 44 hrs in 1 ml of 30 U/ml collagenase
solution. Afterwards, 0.1 ml of alcian blue (1% alcian blue in 3% acetic acid, PH 2.5)
was added to that solution and was kept for 24 hrs to settle the precipitate. At low pH
and in the presence of glycosaminoglycans, alcian blue forms a compound and
precipitate [304]. The resulting samples were then treated with 1 ml of 0.05 mg/ml
trypsin–PBS solution for 4 hrs. Afterwards, 0.1 ml of alcian blue was added to the
solution and the solution was kept for 24 hrs for the precipitate to settle so as to
check the amount of proteoglycans left after collagenase treatment. In addition, 1 ml
distilled water mixed with 0.1 ml of alcian blue served as the control for the
experiment. Results of the experiment are shown in figure 5.7. Based on the
experiment, it was confirmed that only a small amount of proteoglycans was
Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage 119
removed by collagenase treatment while a majority of the proteoglycans were still
intact in the tissue matrix (as can be seen by a comparison of Figures 5.7(a), 5.7(b)
and 5.7(c)).
Figure 5.7: (a) Alcian blue 0.1 ml mixed in 1 ml of distilled water (control test); (b) Alcian blue 0.1 ml mixed in the resulting solution after a cartilage sample being digested in 1 ml of 30 U/ml collagenase for 44 hrs; (c) Sample after digesting in collagenase was treated in 1 ml of trypsin–PBS solution (0.05 mg/ml) for 4 hrs and then mixed with 0.1 ml of alcian blue
5.4.2.4 Removal of surface lipids: Chloroform, methanol mixture
There are several methods to remove lipids from the surface of the cartilage tissue.
They include mechanical, enzymatic and chemical delipidisation. The mechanical
delipidisation involves carefully wiping the cartilage surface with emery cloth,
sandpaper or glasspaper [292]. The uncertainty of grit size and the rate of removal of
the lipid layer make it difficult to control the amount that is delipidised in the
process. Enzymatic delipidisation uses enzymes to remove the surface lipids by
specifically hydrolysing the phospholipid chains [305, 306]. Although this method
selectively removes the lipids from surfaces, it is not commonly used to delipidise
the surface of cartilages.
Chemical delipidisation is the most common method of removing lipids from
the cartilage matrix and surface. Reagents such as ethanol [307], propylene glycol
120 Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage
[307], chloroform [307, 308] and Folch [309-311] have been used to remove lipids
from both the matrix and on the surface. However, the Folch reagent, which is a
mixture of chloroform and ethanol in 2:1 ratio has been proposed to remove lipids
quickly. Therefore, it reduces the tissue exposure time to the chemical, resulting in
fewer risks associated with compromising the integrity of the cartilage matrix [292,
306, 311]. Given this advantage, the Folch reagent was used in the current study to
remove the surface lipids from the cartilage surface and to subsequently assess its
effect on the strain-rate-dependent mechanical behaviour of cartilage tissues.
It has been established that carefully wiping the cartilage surface (for 20–30
minutes) using a Folch reagent will significantly remove the surface lipid layer [312].
Therefore, the surfaces of the kangaroo cartilage samples (n=9) harvested from near
the central load-bearing area of the humeral head were wiped continuously for
approximately 30 minutes using Kimwipes soaked in Folch reagent. Before and after
delipidisation, the samples were subjected to mechanical testing (Figure 5.8)
following the same protocols as before. The mechanical properties were extracted
from the force–indention curves and the effect of surface delipidisation on the strain-
rate-dependent mechanical behaviour was assessed.
Figure 5.8: Steps carried out to investigate the effect of surface lipids on the mechanical behaviour of kangaroo shoulder cartilage
Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage 121
5.4.3 Statistical data analysis procedure
The mechanical behaviour and properties of normal, proteoglycan-degraded and
collagenase-degraded tissues were statistically compared. The repeated measure
ANOVA was used to identify the statistical significance of the treatments, while
Tukey’s pairwise comparison test was employed to compare the individual levels of
treatments. The random effect model shown in Eq. (5.1) was used for the statistical
analysis with the variable ‘sample’ as the random factor. Minitab Version 16.1.1
(2010 Minitab Inc.) was used for statistical analysis. The statistical significance is
reported at both 95% (p<0.05) and 99.5% (p<0.005) confidence intervals. Eq. (5.1) is
expressed as follows:
Y= Sample + Strain + Strain-rate + Exposure-time + Sample*Strain +Sample*Strain-
rate +Sample*Exposure-time + Strain*Strain-rate +Strain*Exposure-time + Strain-
rate*Exposure-time (5.1)
In the above model, ‘Exposure-time’ is the amount of time the samples were
subjected to the chemical (i.e. trypsin–PBS for 1 hr, 2 hrs and 4 hrs, collagenase for
44 hrs, and Folch reagent for 0 hr). ‘Strain-rate’ had four levels (10-4/s, 5x10-4/s,
5x10-3/s and 10-2/s). Six levels (0%, 5%, 10%, 15%, 20%, 25%) were chosen for
‘Strain’.
5.5 RESULTS AND DISCUSSION
The responses of typical cartilage samples before and after 4 hrs of trypsin treatment
(Figures 5.9(a) and 5.9(b)), 44 hrs of collagenase treatment (Figures 5.9(c) and
5.9(d)) and chloroform/methanol treatment (Figures 5.9(e) and 5.9(f)) are shown
122 Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage
with the corresponding hyperelastic coefficients extracted using Eq. (3.14) (Chapter
3).
Figure 5.9 : Effect of mechanical behaviour when (a) normal sample is (b) treated in trypsin (0.05 mg/ml) for 4 hrs, and when (c) normal sample is (d) treated with collagenase (30 U/ml) for 44 hrs, and when (e) normal sample is (f) treated with Folch reagent to remove surface lipids
Samples from post-trypsin and post-collagenase treatments showed a reduction
in stiffness while no such difference was revealed in samples from post-
Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage 123
chloroform/methanol treatment. The effects of these treatments are detailed in the
following sections. The sample thicknesses calculated using ultrasound
measurements were 0.76±0.16 mm (n=10), 0.78±0.10 mm (n=12) and 0.63±0.01 mm
(n=9) for the trypsin treated, collagenase treated and surface lipid-removed cartilage
sample groups, respectively.
5.5.1 Effect of proteoglycan degradation on strain-rate-dependent behaviour and mechanical properties
5.5.1.1 Effect on tissue stiffness and nonlinearity
As shown above in Figure 5.9, for a typical sample, trypsin treatment reduced the
ability of the tissue to withstand external compressive loads. This was observed in all
strain-rates and was further confirmed by the variation of Young’s modulus with
exposure time to trypsin (Figure 5.10(a)). As shown, the Young’s modulus, which is
an indicator of matrix stiffness, reduced gradually with the progressive removal of
proteoglycans. This was observed at all strain-rates.
The decrease in stiffness was statistically significant for 1 hr, 2 hrs and 4 hrs of
trypsin treatment when compared with normal tissue (p<0.05). Except at 10-4/s
strain-rate, the stiffness of the 1 hr- and 2 hr-treated samples was not statistically
different (p>0.05). Nonetheless, the stiffness of the 1 hr-treated samples was
considerably lower (p<0.05) than the normal tissue samples. This indicated that a
considerable amount of proteoglycans might have been removed from the samples
during the first hour of trypsin treatment. The stiffness reduction from 2 hrs to 4 hrs
compared with 1 hr to 2 hrs was high but not statistically significant (p>0.05).
The nonlinear stiffness parameter 20C , which is an indicator of the nonlinearity
of the stress-strain behaviour of cartilage, reduced gradually with the progressive
degradation of proteoglycans (Figure 5.10(b)). However, the decrease in 20C was not
statistically significant at the smallest (10-4/s) and largest strain-rates (10-2/s) tested
124 Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage
when compared with untreated samples. Nevertheless, significant differences were
observed at intermediate strain-rates (p<0.05). The results indicated that the
nonlinearity of the tissue response may not have been affected by the proteoglycan
degradation at the smallest and highest strain-rates. The reason for this behaviour has
not been reported in the literature and remained unclear in this study; therefore, it
requires further investigation. However, the observations might be related to small
fluid velocities at the smallest strain-rate and the fluid containment effect at large
strain-rates. Overall, the gradual reduction of stiffness with removal of proteoglycans
confirms the already-established belief that proteoglycans directly affect the
compressive load-bearing function of cartilage tissues.
Figure 5.10: Effect of 1 hr, 2 hrs and 4 hrs of trypsin treatment (0.05 mg/ml) on (a) Young’s modulus and (b) nonlinear stiffness parameter of kangaroo shoulder cartilage for 10-4/s, 5x10-4/s, 5x10-3/s and 10-2/s strain-rates
5.5.1.2 Effect on tissue permeability and pore size
In order to check the increase in pore size after removal of proteoglycans,
permeability was extracted for normal and 4 hr trypsin-treated tissues by curve fitting
a porohyperelastic model to experimental force–indention data at the lowest strain-
rate. The results showed that the permeability increased from 1.38±0.83×10-14 m4/Ns
to 3.03±1.43×10-14 m4/Ns due to proteoglycan degradation (p<0.005), similar to the
Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage 125
results in previously reported studies [20]. Based on Eq. (4.28) in Chapter 4, these
permeability values corresponded to pore sizes of 160.38±37.49 Å and 238.96±48.00
Å, respectively, which represents an increase of 1.48 times.
It was presumed that complete removal of proteoglycans would considerably
reduce the ability of cartilage tissues to respond to varying rates of loads, and hence
would completely remove the strain-rate-dependent nature of cartilage tissues. This
was based on the anticipation that the complete removal of proteoglycans will
increase the pore size of the tissue to an extent where solid–interstitial fluid frictional
interactions are considerably reduced. However, even after 4 hrs of trypsin treatment,
which removed almost all the proteoglycans, the strain-rate-dependent behaviour was
observed in the tested samples. This may be because the dense collagen meshwork is
still able to sustain the pore sizes in cartilage to such an extent that the solid–
interstitial fluid frictional interaction is able to facilitate the tissue’s ability to respond
to varying rates of loads, or it could also be due to the viscoelasticity of the collagen
network.
5.5.1.3 Role of proteoglycans in strain-rate-dependent behaviour
One of the objectives of the current study was to investigate the contribution of
proteoglycans to the tissue behaviour at individual strain-rates. The result of the
histological study described in Section 5.4.2.1 indicated that 4 hrs of trypsin
treatment resulted in almost complete removal of proteoglycans from kangaroo
shoulder cartilage. Therefore, the percentage reduction of Young’s modulus and the
nonlinear stiffness parameter at individual strain-rates due to proteoglycan
degradation (after 4 hrs of trypsin treatment) were calculated and compared. The
percentage decreases in Young’s modulus were 36.5±22.8%, 35.0±21.3%,
39.4±17.0% and 33.1±14.0% for 10-4/s, 5x10-4/s, 5x10-3/s and 10-2/s strain-rates,
126 Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage
respectively. These values indicated that complete removal of proteoglycans reduced
the Young’s modulus of the tissues more or less to a similar extent at the different
strain-rates tested. The statistical analysis indicated that the percentage decreases in
Young’s modulus at each strain-rate were not significantly different from each other
(p>0.05). For the tested strain-rates of 10-4/s, 5x10-4/s, 5x10-3/s and 10-2/s, the
reduction of the nonlinear stiffness parameter was 33.5±30.8%, 48.2±24.9%,
46.8±28.2% and 23.2±35.7%, respectively. Reductions of 20C at the lowest and
highest strain-rates were smaller than at intermediate strain-rates. However,
statistically significant differences were only identified between the highest and
intermediate strain-rates (p<0.05).
Previous studies have stated that the proteoglycans of cartilage tissue are
responsible for the equilibrium compressive properties (extracted at very low strain-
rates) of the tissue [20]. Further, compositional-based FE models specifically
focusing on the strain-rate-dependent behaviour of cartilage have shown that, at
small strain-rates (10-3/s), proteoglycans contribute more to the mechanical
behaviour of the tissue [16]. However, in the present study, irrespective of the strain-
rate, proteoglycans seem to have similar contribution to the compressive load-
bearing properties of the cartilage tissues, thus confirming the important role of
proteoglycans in tissue behaviour at both low and high strain-rates.
5.5.2 Effect of superficial collagen degradation on strain-rate-dependent behaviour and mechanical properties
5.5.2.1 Effect on tissue stiffness and nonlinearity
Similar to the trypsin treatment, treating kangaroo shoulder cartilage samples in
collagenase for 44 hrs also reduced the ability of the tissue to withstand external
loading at all strain-rates (Figures 5.9(c) and 5.9(d)). Figure 5.11(a) presents the
reduction of tissue stiffness due to collagenase treatment indicated by the decrease in
Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage 127
Young’s modulus. In general, stiffness of the tissues decreased at all strain-rates
tested, with the difference being statistically significant compared to untreated tissues
(p<0.005). Unlike in the case of trypsin treatment, where significant reductions in the
nonlinear stiffness parameter were only evident at intermediate strain-rates (5x10-4/s
and 5x10-3/s), 20C showed a statistically significant reduction at all strain-rates
(p<0.05) (Figure 5.11(b)).
Figure 5.11 : Effect of 44 hr collagenase treatment (30 U/ml) on (a) Young’s modulus and (b) nonlinear stiffness parameter of kangaroo shoulder cartilage for 10-4/s, 5x10-4/s, 5x10-3/s and 10-2/s strain-rates
Collagenase treatment for 40–44 hrs is known to significantly degrade the
superficial collagen [19-21, 237]. Longer exposure time may affect deep zone
collagen. However, the difficulty in exactly controlling the effect of collagenase on
the collagen network can be considered as a possible limitation of this study.
Nevertheless, the results of the histology (Section 5.4.2.2) and alcian blue (Section
5.4.2.3) experiments confirmed that only a small amount of proteoglycans was
removed by collagenase treatment, while the majority of the proteoglycans were still
intact in the tissue matrix, which is similar to the results of previously reported
studies [19, 303]. Therefore, it can be said that the reduction of the tissue’s stiffness
to a large extent is indeed due to the degraded superficial collagen. Overall, the
128 Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage
results indicated that the stiffness and nonlinearity of the kangaroo shoulder cartilage
were compromised due to superficial collagen degradation.
5.5.2.2 Effect on tissue permeability and pore size
The results of the present study indicated the tendency of the strain-rate-dependency
to reduce with collagen disruption. The inverse FEA method was used to fit the
force–indentation curves at the smallest strain-rate to a porohyperelastic model to
find out the permeability. It was found that, due to collagenase treatment, the
permeability increased from 1.36±0.41×10-14m4/Ns to 4.19±2.79×10-14m4/Ns.
These permeability values corresponded to pore sizes of 162.16±22.32 Å and
270.22±94.96 Å, respectively—an increase of 1.69 times. The increase of
permeability due to collagenase treatment has also been reported previously [20]. It
is highly possible that the collagen disruption increases the interspaces between
collagen fibrils and hence the pore size of the tissue. Therefore, the collagen
disruption and the increase in pore size compromise the tissue’s ability to respond to
varying strain-rates, as observed in this study.
5.5.2.3 Role of superficial collagen in strain-rate-dependent behaviour
The decrease in stiffness due to collagenase treatment has been reported in numerous
studies [19-21, 137]. However, there are limited studies on the effect of superficial
collagen on the strain-rate-dependent behaviour of shoulder cartilage. There are
reported numerical studies stating that the superficial collagen affects cartilage tissue
behaviour at large strain-rates much more than at small strain-rates [16]. However,
most of these studies have been conducted on knee cartilage. Due to the differences
in the mechanical environments experienced by knee cartilage (high magnitude and
frequent compressive loads) and shoulder cartilage (low magnitude compressive
loads), it is likely that the collagen network (including the superficial layer) plays a
Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage 129
more significant role than proteoglycans in shoulder cartilage. Further, considering
that the porohyperelastic model with strain-rate-dependent permeability was unable
to predict the tissue behaviour at the highest strain-rate tested (10-2/s), it was thought
that the collagen network might be playing a larger role in determining tissue
behaviour at high strain-rates. Further, for knee cartilage, there is evidence that the
effect of superficial collagen on tissue behaviour increases with strain-rate
[16].Therefore, one of the objectives of the present study was to check the
contribution of superficial collagen to the mechanical behaviour and properties of the
kangaroo shoulder cartilage at individual strain-rates.
To do so, the percentage decrease in Young’s modulus for 44 hr collagenase
treatment, which potentially would have greatly degraded the superficial collagen,
was assessed. The percentage decreases in Young’s modulus were 53.0±29.6%,
55.4±28.3%, 55.4±25.4% and 55.7±16.3% for strain-rates of 10-4/s, 5x10-4/s, 5x10-3/s
and 10-2/s, respectively. The differences in the percentage decreases of Young’s
modulus between the strain-rates were statistically insignificant (p>0.1). This
indicated that the superficial collagen more or less had an equal effect on the tissue
behaviour at all strain-rates. The decreases in the nonlinear stiffness parameter were
0.72±13.9%, 65.8±26.3%, 58.3±31.9% and 70.2±24.0% for the above strain-rates.
Although there were differences in the percentage decreases in the Young’s modulus
and nonlinear stiffness parameter for different strain-rates, these were not identified
to be statistically different (p>0.05). This indicated that the superficial collagen
affected the kangaroo shoulder cartilage behaviour more or less to the same extent
for all the strain-rates tested.
However, the aforementioned result differs from the findings reported in the
literature. As mentioned before, it has been reported that, at high strain-rates, the
130 Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage
superficial collagen content and collagen network architecture influence the tissue
behaviour considerably more than at low strain-rates [16]. These different findings
could be mainly due to the differences of the specific joint cartilages studied:
previous studies have used bovine knee cartilage, while this study used kangaroo
shoulder cartilage. In terms of microstructure and composition, there are noticeable
differences between shoulder and knee cartilage tissues (as discussed in detail in the
next chapter). Therefore, the FE model parameters (water volume fraction, fixed
charge density, collagen volume content, superficial and deep zone thickness) used in
the literature can be different from the parameters necessary for shoulder cartilage.
Therefore, it is our belief that the conclusions made through numerical models
should be evaluated according to the relevant tissue investigated (i.e. whether it is
knee or shoulder cartilage).
Superficial collagen degradation, in addition to collagen, may remove the
surface lipids of the tissue as well. Therefore, it was necessary to evaluate the effect
of surface lipids on tissue behaviour in order to confirm that the effects mentioned
above were mainly due to the superficial collagen degradation. The findings of the
succeeding study can also inform the research community about the role of surface
lipids in the strain-rate-dependent behaviour of the tissue.
5.5.3 Effect of surface phospholipid removal on strain-rate-dependent behaviour and mechanical properties
A comparison of Figures 5.9(e) and 5.9(f) above, for typical cartilage samples, seems
to suggest that the removal of surface phospholipids has a minimal effect on the
tissue’s ability to withstand external loading. This can be further identified by
comparing the average Young’s modulus of normal and surface lipid-removed
samples (Figure 5.12(a)). In general, stiffness of the cartilage matrix was reduced by
a small amount due to surface lipid removal. However, these differences were
Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage 131
statistically insignificant (p>0.05). Similarly, the nonlinear stiffness parameter was
also reduced due to the removal of the surface lipids. However, in this case,
statistically significant differences were only identified at 5x10-4/s and 10-2/s strain-
rates (p<0.05) (Figure 5.12(b)). The statistically insignificant differences imply that
the removal of surface lipids does not adversely affect the strain-rate-dependent
behaviour of kangaroo shoulder cartilage.
Figure 5.12: Effect of surface lipid removal on (a) Young’s modulus; and (b) the nonlinear stiffness parameter of kangaroo shoulder cartilage for 10-4/s, 5x10-4/s, 5x10-3/s and 10-2/s strain-rates
Previous studies [307, 308] have reported that the stiffness of cartilage could
reduce at low strain-rates (1.3x10-4/s) and increase at high strain-rates (10/s) when
lipids are removed from the tissue. These studies used a slightly different
methodology where cartilage was delipidised by immersing the specimen in a
reagent solution followed by a vacuum-drying process to evaporate the reagent.
During this method there is a high possibility for lipids inside the cartilage matrix as
well as on the surface to be removed. In addition, there are possible changes in the
tissue matrix due to the dehydration process. To a large extent, this could be the main
reason for the differences observed in the present study and other reported studies.
132 Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage
5.5.4 Comparison of the effect of proteoglycan and superficial collagen on strain-rate-dependent behaviour
Acknowledging the fact that each cartilage constituent plays an important role in the
cartilage behaviour, we specifically analysed their effect on the strain-rate-dependent
indentation behaviour of kangaroo shoulder cartilage tissues. In order to compare the
effects of complete removal of proteoglycans (Figure 5.13(a)), severe disruption to
the superficial collagen (Figure 5.13(b)) and removal of surface lipids (Figure
5.13(c)), normalised force vs indentation plots were considered.
The force–indentation graphs indicated that the ability of the cartilage to
withstand external loading was reduced more when the superficial collagen was
significantly degraded than in the case of complete removal of proteoglycans
(compare Figure 5.13(a) and Figure 5.13(b)). The effect of surface lipid removal on
the tissue behaviour was small (Figure 5.13(c)). Interestingly, when the effects of 4
hr trypsin treatment and 44 hr collagenase treatment were compared, it was noted
that the latter treatment reduced tissue stiffness more at all strain-rates (p<0.05)
(Figure 5.14(a) and Table 5.1). The percentage decrease in the nonlinear stiffness
parameter was also significantly higher in response to superficial collagen disruption
rather than proteoglycan degradation (p<0.05) (Figure 5.14(b)). Therefore, it can be
concluded that the contribution of superficial collagen to tissue behaviour in
kangaroo shoulder cartilage is higher than the contribution of proteoglycans. This is
understandable considering that chondrocytes are able to synthesise the extracellular
matrix according to the mechanical inputs the tissue receives. Hence, the larger and
more frequent the compressive forces, the higher the proteoglycan composition and
the greater its role in the tissue behaviour. Given that the shoulder joint experiences
low magnitude compressive loads (as discussed by comparison to knee joint in
Chapter 6, Section 6.1), the stimulation of chondrocytes by compressive forces is
Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage 133
also rather low. Thus, compared to knee cartilage, shoulder cartilage may comprise a
small amount of proteoglycans. This implies that the collagen architecture may play
a significantly larger role in determining shoulder cartilage tissue behaviour, which is
confirmed by the results of the present study.
Table 5.1: Young’s moduli (MPa) of 4hr trypsin-treated and 44 hr collagenase-treated kangaroo shoulder cartilage at four strain-rates
Strain-rates 10-4/s 5x10-4/s 5x10-3/s 10-2/s
0 hrs in trypsin (n=12) 0.040 ± 0.016
0.078 ± 0.445
0.360 ± 0.261
0.523 ± 0.299
4 hrs in trypsin 0.026 ± 0.014
0.048 ± 0.026
0.194 ± 0.076
0.328 ± 0.131
0 hrs in collagenase (n=10)
0.10 ± 0.045
0.217 ± 0.125
1.023 ± 0.580
1.412 ± 0.485
44 hrs in collagenase 0.043 ± 0.025
0.084 ± 0.049
0.377 ± 0.214
0.633 ± 0.341
These findings were further reinforced by the observation that 44 hrs of
collagen degradation more or less had an equal (p>0.1) effect on the tissue behaviour
at all strain-rates tested, while similar observations were also made for 4 hrs of
proteoglycan degradation (Figure 5.14(a)). These findings are contrary to reported
findings on knee cartilage. In investigating knee cartilage behaviour from 10-3/s to
10-1/s, Julkunen et al. [16] reported that, compared to low strain-rates, superficial
collagen substantially contributed to the tissue behaviour at the higher strain-rates
(10-1/s). However, in their study, proteoglycan contribution (approximately 37.2%) to
the tissue behaviour at 10-1/s was still much higher than the superficial collagen
contribution (14.7%).
134 Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage
Figure 5.13: Effect on average normalised force–indentation curves due to (a) 4 hrs of trypsin (0.05mg/ml) treatment, i.e. proteoglycan completely removed (n=12); (b) 44 hrs of collagenase treatment, i.e. severe disruption to superficial collagen (n=10); and (c) surface phospholipid removal (n=9)
On the contrary, by calculating the average percentage decrease in Young’s
modulus (Figure 5.14(a)) at the strain-rates tested, the results of the present study
showed that the contribution of the superficial collagen and proteoglycans to the
tissue behaviour of shoulder cartilage was 54.88±1.11% and 35.99±2.3%,
respectively. This difference in the observations in the current study and in the
literature is reasonable because the previous studies focused on knee cartilage that is
most likely structurally and compositionally different from shoulder cartilage. When
the proteoglycan content of kangaroo shoulder and knee cartilage was analysed (as
discussed in the next chapter), noticeable differences were identified. As mentioned
Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage 135
before, the collagen network plays a dominant role in the mechanical behaviour of
shoulder cartilage, to an even larger extent than proteoglycans. Therefore, a finding
that collagen has an equally dominant effect on the mechanical behaviour of shoulder
cartilage at all strain-rates is justifiable.
In cartilage, water-swollen proteoglycans are constrained by the three-
dimensional collagen network to form the functional load-bearing unit of cartilage.
Any disruption of the collagen network reduces its ability to constrain the
proteoglycans, compromising the matrix integrity and in turn its ability to act as an
effective load-bearing unit. In addition to the reduction in tissue stiffness, superficial
collagen disruption and proteoglycan degradation also increased the tissue
permeability (p<0.005). It was noted that the collagen disruption affected the
permeability more than the proteoglycan degradation (p<0.005). In the case of
proteoglycan degradation, the increase was 2.19 times; in the case of collagen
network disruption, it was 3.04 times. Therefore, the significant reduction in strain-
rate-dependency observed when the collagen network was disrupted—even more
than in the case of proteoglycan degradation—confirmed the importance of the
collagen network in facilitating the strain-rate-dependent behaviour of cartilage.
Assuming that the effect of collagenase and trypsin treatments on the tissue are
independent, the collagen disruption and proteoglycan degradation in total
contributed to approximately 89–95% (Figure 5.14(a)) reduction in tissue stiffness.
This implies that the total removal of proteoglycans and significant disruption of
superficial collagen would render the kangaroo shoulder cartilage almost incapable
of responding to varying rates of external loads. Although dominated by the collagen
network, this shows the important functional interdependency of collagen and
136 Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage
proteoglycans in facilitating the strain-rate-dependent behaviour of shoulder
cartilage.
As mentioned in Chapter 2, it is known that in cartilage tissues the value of the
modulus in tension is several times higher than in compression. This is more
pronounced in upper limb cartilage tissues compared to lower limb cartilage tissues.
The results of the current study, through the indication of the considerably high
contribution of superficial collagen in determining the cartilage behavior, indicated
the importance taking this tension-compression nonlinearity into consideration when
modelling shoulder cartilage tissues.
Figure 5.14: Percentage decrease in (a) Young’s modulus and (b) nonlinear stiffness parameter of kangaroo shoulder cartilage due to complete removal of proteoglycans (4 hrs of treatment in 0.05 mg/ml trypsin) and severe disruption to superficial collagen (44 hrs of treatment in 30 U/ml collagenase)
5.5.5 Effect of proteoglycan and superficial collagen degradation on long-term functional load-bearing ability of the tissue
In order to understand the changes in the internal tissue behaviour when
proteoglycans and superficial collagens are degraded, the porohyperelastic FE model
which was validated for internal pore pressure measurements was employed (Chapter
3, Section 3.3.2). Based on the FE model predictions, as shown in Figure 5.15, for
Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage 137
the strain-rate of 10-4/s, the hydrostatic excess pore pressure noticeably decreased
due to degradation of the proteoglycans and superficial collagen. The reasons for
these results are the decrease in elastic properties and increase in permeability when
the tissue is degraded. The fluid is less capable of contributing to the load-bearing
function in the case of superficial collagen degradation as compared to proteoglycan
degradation. Therefore, there will be more burden on the collagen network when the
superficial collagen is degraded, resulting in the collagen network being further
damaged and eventually dysfunctional.
Figure 5.15: Variation of pore pressure with strain for 4 hr trypsin-treated and 44 hr collagenase-treated samples
The model parameters used for simulating the behaviour of tissue after proteoglycan
and collagen degradation is shown in Table 5.2. The stiffness parameter 10C was
calculated based on the percentage reduction of Young’s modulus due to 4 hrs of
trypsin treatment and 44 hrs of collagenase treatment (i.e. 35.99% and 54.88%,
respectively). Similarly, the permeability values were calculated based on the
increase in permeability due to trypsin treatment for 4 hrs (2.19 times) and due to
collagenase treatment for 44 hrs (3.06 times). The nonlinear stiffness parameter 20C
used for the simulation was assumed to be small (i.e. 0.01 MPa), considering the
138 Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage
linearity of the stress-strain response data [135] used to validate the FE model and
was maintained constant throughout the simulations. It is noted that these
assumptions did not affect the results and conclusion of this study.
Table 5.2: FE model parameters for normal, 4 hr trypsin-treated and 44 hr collagenase-treated samples
Normal tissue
Proteoglycan-
degraded tissue (4
hrs in trypsin)
Collagen-degraded
tissue (44 hrs in
collagenase)
10C (MPa) 0.158 0.101 0.0554
20C (MPa) 0.01 0.01 0.01
1D (1/MPa) 4.738 7.403 13.536
k (Ns/m4) 2.58x10-16 5.64x10-16 7.73x10-16
5.6 CONCLUSION AND REMARKS
Considering the results reported in Chapter 4, we carried out a systematic and
comprehensive study to investigate the effect of the cartilage constituents on the
strain-rate-dependent behaviour of kangaroo shoulder cartilage. In doing so,
enzymatic degradations of proteoglycans and superficial collagen along with
simultaneous indentation tests at different strain-rates were conducted. The removal
of proteoglycans increased the pore size of the tissue. The exposure of the tissue to
trypsin was limited to 4 hrs in order to reduce the possible effects on the collagen
network and was identified as enough to remove almost all the proteoglycans. The
collagenase treatment was limited to 44 hrs in order to significantly degrade the
superficial collagen and this was found to remove only a small amount of
proteoglycans from the tissue. The following remarks and conclusions can be made
based on the results of the current study:
Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage 139
• Neither of the tissue preservation methods (i.e. preservation of the tissue at 4
°C in PBS-inhibitor solution and freeze–thaw methods) affected the
mechanical properties and behaviour of the tissue significantly. However,
considering reported concerns about the freeze–thaw method, preservation of
the tissue at 4 °C in PBS-inhibitor solution was chosen as the preservation
method for the current study.
• Proteoglycan depletion and superficial collagen disruption substantially
compromised the tissues’ ability to respond to and withstand external
compressive loading at different strain-rates. This would increase the risk of
bone-to-bone contact.
• Proving the direct role of proteoglycans in facilitating the compressive load
bearing ability of the tissue, the removal of proteoglycan using 1 hr, 2 hr and
4 hr trypsin treatments gradually reduced the tissue stiffness at all strain-rates.
• Total proteoglycan degradation increased the permeability of kangaroo
shoulder cartilage by 2.19 times and the pore size by 1.48 times. Even after
most proteoglycans have been removed (4 hrs of trypsin treatment), the
kangaroo shoulder cartilage demonstrated strain-rate-dependent behaviour.
This could be due to dense collagen meshwork is been able to keep pore sizes
such that solid-interstitial fluid frictional interactions are able to facilitate the
strain-rate-dependent behavior.
• Similarly, even after severe disruption to the superficial collagen network (44
hrs of collagenase treatment), the strain-rate dependency was still a
characteristic of the tissue. In this case, the permeability increased by 3.04
times and the pore size by 1.69 times.
140 Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage
• The effect of surface lipids on the strain-rate-dependent mechanical
behaviour of kangaroo shoulder cartilage was found to be small and
statistically insignificant. Additionally, superficial collagen disruption was
found to remove only a small amount of proteoglycans while keeping the
majority intact. These two results together confirmed that the main reason for
the reduction of tissue stiffness due to collagenase treatment is the disruption
to superficial collagen.
• Superficial collagen degradation had the largest effect on strain-rate-
dependency compared with complete removal of proteoglycans. Therefore, it
was concluded that the contribution of superficial collagen is higher than the
proteoglycans in facilitating the strain-rate-dependent behaviour.
• Superficial collagen was found to contribute evenly to tissue behaviour at all
strain-rates, thus confirming the importance of superficial collagen in
governing the mechanical properties and behaviour of kangaroo shoulder
cartilage at both low and high strain-rates.
• The above two findings are different to the conclusions reported in the
literature on knee cartilage. In previous studies, the superficial collagen is
reported to contribute less than proteoglycans to the mechanical behaviour.
Furthermore, the contribution of superficial collagen is reported to be
considerably large at high strain-rates when compared to its contribution at
low strain-rates. These differences can be attributed to the potential
compositional and microstructural differences in the two types of tissues (i.e.
knee cartilage and shoulder cartilage). An investigation into the differences in
these two types of tissues is discussed in the next chapter.
Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage 141
• Based on porohyperelastic modelling, it was revealed that collagen disruption
leads to faster shoulder cartilage damage than when proteoglycans were
depleted. This is due to interstitial fluid being less capable of supporting
external loads.
The investigations reported in this chapter gave insights into the role of
proteoglycans and the collagen network in the strain-rate-dependent mechanical
behaviour of kangaroo shoulder cartilage. Further, the results of the experiments
pointed to the differences in knee and shoulder cartilage as the reason for the
different results in the present study and in the literature. Nevertheless, to further
investigate and confirm this conclusion, the compositional and macrostructural
differences in knee and shoulder cartilage were studied as discussed in the next
chapter.
142 Chapter 5: Effect of proteoglycan and superficial collagen on the strain-rate-dependent mechanical behaviour of kangaroo shoulder cartilage
Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage
This chapter begins with a brief introduction to the forces experienced by shoulder
and knee cartilage tissues (Section 6.1). The aims and objectives of this part of the
study as well as the hypothesis are presented (Sections 6.2 and 6.3, respectively). The
methods and materials used are set out in detail (Section 6.4), followed by the results
and discussion (Section 6.5). Lastly, the chapter presents the conclusions and
remarks based on the results obtained in the study (Section 6.6).
6.1 INTRODUCTION
Anatomically different knee and shoulder joints undergo forces which are
significantly different in magnitude and mode due to the bipedal nature of humans.
For example, the shoulder joint experiences compressive forces of 44–90% body
weight during arm elevation of 60°–100°, with shear forces being almost 50% body
weight at 60° of abduction [87-90]. In comparison, the knee joint experiences
compressive forces of 50% to 600% body weight during flexion angles of 0°–90° in
static joint positions [313]. These differences are more prominent when dynamic
movements are considered where joint can experience 2-3 times loading than in the
static loading conditions [91, 92]. A professional baseball pitcher’s shoulder joint
experiences at minimum compressive/distractive forces of 108% BW and generates
external rotational torques (causing shear stresses) in the range of 67-92 Nm [1, 93,
94]. On the other hand, during normal walking and jogging, the knee joint
Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage143
experiences compressive forces of 300% body weight to 700% body weight,
respectively [313].
Arguably, the mechanical properties and mechanical behaviour of cartilage
tissues are regulated by the magnitude and mode of forces experienced by the tissue.
Given the considerable differences in forces experienced by the two joints, it is
expected that the shoulder cartilage is compositionally and microstructurally
different from the knee cartilage. This could perhaps be the main reason why the
conclusions made in the earlier part of this study (Chapter 5) are different from the
findings of studies in the literature that focused mainly on knee cartilage. The animal
model used throughout this thesis, namely, kangaroo, experiences considerably
different magnitudes of load in the lower limb joints compared to the upper limb
joints due to its bipedal-hopping locomotion. Therefore, kangaroo as an animal
model provides a unique avenue for investigating the effect of joint loading on the
composition and microstructure of cartilage. Following on from the investigation
reported in Chapter 5, the objectives of this part of the study are presented in the next
section.
6.2 AIMS AND OBJECTIVES
The main aims and objectives of this part of the study were:
1. To investigate the differences in proteoglycan concentration and distribution
between knee and shoulder cartilage.
2. To investigate the differences in features of the collagen network of knee and
shoulder cartilage.
3. To investigate the contribution of proteoglycan and superficial collagen to the
strain-rate-dependent behaviour of knee cartilage and compare the findings
with shoulder cartilage.
144 Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage
4. To relate the mechanical behavioural differences of knee and shoulder
cartilage to compositional and microstructural differences.
6.3 HYPOTHESIS
The main hypothesis in this part of the study was: Compositionally and
microstructurally, the shoulder cartilage and knee cartilage are distinctly different
from each other and these differences are reflected in the biomechanical behaviour of
the tissues.
6.4 METHODS AND MATERIALS
In the present investigation, histological and polarised light microscopy (PLM)
studies on kangaroo knee and shoulder cartilage were conducted. Further, knee
cartilage was subjected to sequential proteoglycan and collagenase degradation along
with mechanical tests under different strain-rates. The results of these mechanical
tests on knee cartilage were compared with the results on shoulder cartilage
presented in Chapter 5. The mechanical tests and histological studies were carried
out on separate samples. This is because, in between consecutive mechanical
testings, the sample size and geometry should be maintained, while histological
studies require a biopsy to be taken. The preservation of the sample geometry and
dimensions is especially important for the meaningful extraction of mechanical
properties from force–indentation data. The following section describes the sample
preparation methods along with details of the histological and PLM studies.
6.4.1 Cryostat tissue sectioning: Tissue preparation for histological studies
Biopsy samples were obtained from three fresh, adult kangaroo knee and shoulder
joints. For shoulder joints, the biopsy sample was taken from near the central load-
bearing area of the humeral head, while for knee joints, the samples were harvested
Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage145
from the medial and lateral sides of the femoral condyle and tibial plateau (2 samples
from each location). Then, a cryostat was used to reduce the sample thickness to a
suitable size (10 μm) for histological and PLM studies. Firstly, the temperature of the
cryostat with mounting metal chucks inside was set to -20 °C to -22 °C. Then, the
samples were placed inside the metal chuck and immediately embedded and frozen
in an optimal cutting temperature medium. The samples were embedded such that the
cryostat blade was parallel to the articular surface. The embedded samples were
cryosectioned and 10 μm specimens were then collected carefully on microscope
slides and left to dry in an oven at 30 °C for one week before histological staining.
This procedure made the samples stick to the slide and reduced the possible loss of
the sample during staining.
6.4.2 Safranin-O staining protocol
Safranin-O, which is a cationic dye, binds to positively-charged glycosaminoglycan
in a one-to-one ratio. The negatively-charged carboxyl or sulphate group in
glycosaminoglycan helps to form the bonding with the positively-charged dye
molecules [314, 315]. Therefore, safranin-O is considered to be a suitable dye for
the histochemical quantification of proteoglycans in cartilage tissues [314]. Hence,
in accordance with the standard safranin-O staining protocol, the prepared slides
were fixed in 95% alcohol for 30 seconds and left to dry in air. Then, the specimens
were hydrated in distilled water and rinsed in 1% acetic acid for 15 seconds. Staining
was conducted by immersing the slides for 5 minutes in a 0.1% safranin-O solution.
Following this, the slides were dehydrated with 95% alcohol for 6 dips and 100%
alcohol for 8 dips. During the dehydration, the metachromatic dye-dye bonding is
destroyed to form an orthochromatic bonding. Therefore, stain intensity can be
directly related to proteoglycan concentration [314, 315].
146 Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage
6.4.3 Proteoglycan quantification: Optical absorbance measurements
In accordance with the protocol developed by Moody and Brown [22] and validated
by Afara and Singh [301], the Beer–Lambert law of optical absorbance was used to
quantify the proteoglycans in the present study. As mentioned, since one safranin-O
molecule binds to one negatively-charged chondroitin 6-sulphate or keratan sulphate
[316], the stain intensity directly corresponds to the proteoglycan concentration;
hence, digital densitometry techniques can be applied to quantify the proteoglycan
concentration.
In the present study, by comparing the digital images of unstained and stained
tissue samples, proteoglycan concentration was quantified based on safranin-O stain
intensity. The unstained samples were imaged using a Nikon Labophot-POL
microscope illuminated by a 6V, 20W halogen lamp. The images were captured
using a Nikon DS-Fi1 5-megapixel CCD camera (Nikon Instruments, Sound Vision
Inc.) at 2560x1960 resolution. After safranin-O staining, images were again taken
while maintaining the same sample position and orientation as for the unstained
samples so that the obtained images could be matched with the images of the
unstained samples. All the settings including the shutter speed and light intensity
were kept unchanged during the image capture.
6.4.3.1 Image processing
The Beer–Lambert law of optical absorbance which relates the stain intensity to the
absorbance is used as the principle for image processing. This law states that the
amount of visible light absorbed by a substance, which is the absorbance (A), is
linear proportional to the concentration (C) of the substance and the length of the
path on which the light has travelled. The light absorbance by a sample is related to
the light transmittance (T) through the sample by the following equation:
Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage147
logTA −= where 0IIT = (6.1)
Here, 0I is the initial light intensity and I is the light intensity after
transmittance of light through the sample. In this study, absorbance values were
taken as a measure of proteoglycan concentration by implementing Eq. (6.1) above
in ImageJ 1.48v software (National Institute of Health, USA). Firstly, both the
stained and unstained images were opened in ImageJ and converted to 32-bit images.
The unstained converted image was copied and pasted on the stained image using the
“paste control-blend” function. Afterwards, the images were matched manually and
the pixel values were divided using the “paste control-divide” function. Then, the
logarithm to the base-10 was taken at each pixel using the “log” function in ImageJ
and was multiplied by -1 to obtain the absorbance of safranin-O. The image was
adjusted for brightness/contrast afterwards by setting the minimum and maximum
displayed values to 0 and 3, respectively (image-adjust-brightness/contrast-set) in
accordance with the methodology proposed by Moody and Brown [22]. Next, the
scale measurements were adjusted based on the size (in pixels) of the picture and
corresponding distance in millimetres using the “analyse-set scale” function.
Thereafter, absorbance profiles were obtained from the cartilage surface to bone
which correlates to proteoglycan variation with depth in cartilage.
6.4.4 Sample preparation for PLM measurements
Superficial collagen fibres are parallel to the articular surface and their direction
depends on the joint location. The pin-prick test is a standard test that is widely used
to identify the fibre directions in the surface layer of cartilage tissues. The linear-
markers being formed during the pin-prick test are known as split line directions. The
tensile properties are the highest parallel to the split line direction and are
148 Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage
significantly lower in the perpendicular direction of the split lines [317]. Hence,
arguably, the split line direction indicates the predominant fibre direction in the
superficial cartilage layer. The sample thickness, location and orientation of
harvested samples are known to affect the PLM measurement [318, 319].
Birefringence values obtained through PLM measurements are an indication of tissue
anisotropy and affected by the orientation of harvested samples (i.e. whether or not
the sample’s direction is parallel to the split line direction). Therefore, in the present
study, before harvesting the samples for PLM measurements, the split line directions
were identified in three separate shoulder and knee joints using the pin-prick test.
Pin-prick tests were performed by piercing the cartilage surface
perpendicularly using a sharp needle tip dipped in Indian ink. Then, excess ink was
washed away using PBS to observe the predominant fibre directions on the surface of
the cartilage tissues (Figure 6.1). Considering the split line directions, the samples
were harvested such that the longitudinal directions of the samples were parallel to
the split line directions. The samples were harvested from near the central load-
bearing area of the humeral head and from the medial and lateral sides of the femoral
condyle and tibial plateau, and then cryosectioned to 10 μm specimens as described
above in Section 6.4.1.
Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage149
Figure 6.1: Split line directions identified through the pin-prick test performed on (a) femur; (b) tibia; (c) humeral head; and (d) glenoid of kangaroo knee and shoulder joints
6.4.5 Collagen quantification: PLM measurements
The PLM is widely used to investigate the anisotropic nature of biological tissues
[320, 321]. In principle, when polarised light travels through an anisotropic material
the light is refracted into two polarised wave components, namely, the ordinary wave
and the extraordinary wave. This phenomenon is called double refraction, and the
optically anisotropic materials that demonstrate this characteristic are said to be
birefringent. Depending on the direction of propagation inside the material of
interest, the velocities of these wavefronts would vary. When these out-of-phase
wavefronts are re-combined, the features of the resulting wavefront are considered to
be indicators of material anisotropy.
150 Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage
Electrical fields, vibrating in planes perpendicular to the transmitted direction
of natural light, can be restricted to one plane using filters. The resulting wave that
occurs due to the restriction is referred to as polarised light. In PLM, a filter called
the “polariser” is responsible for converting the natural light to polarised light. The
ordinary and extraordinary wavefronts, created after the polarised light is projected
onto the material, are combined together using a second polariser called the
“analyser”. When these two filters are perpendicular to each other (i.e. cross-
polarised), there would be no resulting wavefront. This is indicated by a dark view
through the microscope eyepiece. When PLM is in cross-polarised configuration, and
when supposedly anisotropic material placed on the circular microscopy stage is
rotated, the sample will be illuminated depending on the anisotropic characteristics
of the material. For cartilage specimens, the maximum illumination or contrast can
be seen when the axis perpendicular to the cartilage surface makes 45° with the axis
of the microscope.
In this study, a Nikon Labophot-POL microscope was used for the PLM
measurements. While the polariser, analyser and λ/4 wave plate in place, the
exposure time, gain, focus and contrast were first adjusted to obtain a clear image.
Then, the sample was placed on the rotatable stage with the articular surface facing
the microscope. The rotatable stage was positioned to 0° and the cross-polarised
configuration was checked before proceeding further. The sample was then rotated in
intervals of 45° from its original principal direction and images were taken in order
to have a comparative idea of the anisotropy of the tissue. The obtained images were
analysed using ImageJ software to obtain a quantitative value for the transmittance of
light through the samples and hence an indicator of the fibre directions with depth.
Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage151
6.5 RESULTS AND DISCUSSION
6.5.1 Differences in proteoglycan concentration with depth in knee and shoulder cartilage
For the samples harvested from knee and shoulder joints, the variations of
proteoglycan concentration with cartilage depth and location based on the optical
absorbance measurements are shown in Figure 6.2. For the lateral femur (LF), medial
femur (MF) and lateral tibia (LT) cartilages, with depth, proteoglycan concentration
increased to a maximum value of approximately 20–30% of tissue depth and
decreased afterwards. In contrast, for the medial tibia (MT), proteoglycan
concentration increased with depth, while for the humeral head (H) cartilage
apparent increase or decrease in proteoglycan concentration were not identified
(Figure 6.2(a)), therefore can be considered as homogeneous compared to
proteoglycan variation in knee cartilage. The findings on proteoglycan distribution
with depth for MT cartilage is similar to the findings on canine MT cartilage reported
in a previous study [322]. To the best of the authors’ knowledge, depth-dependent
proteoglycan variations for other locations in the knee and humeral head cartilage
have not been specifically reported in the literature. Nonetheless, we believe that the
proteoglycan variation with depth must be related to the patterns and magnitudes of
loads experienced by the specific tissue. Therefore, specific investigations need to be
carried out in future research in order to identify the relationship between load
magnitude, pattern and depth-dependent proteoglycan distribution. This could be
invaluable for fine-tuning tissue engineering strategies for cartilage tissues.
The areas under absorbance curves (which can be taken as an indicator of the
amount of proteoglycans in tissue [301]), showed that the LF, MF and LT cartilages
had a higher amount of proteoglycans, while in H and MT cartilages the amount was
low (Figure 6.2(b)). Proteoglycan content in the knee cartilage, at all four locations,
152 Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage
were significantly higher than in the shoulder cartilage (p<0.005). Statistical power
analysis indicated that sample size, n=3, is even sufficient (Power=0.9928) to
differentiate concentration differences of 0.05 absorbance units. The safranin-O
staining for LF, MF, LT, MT and H cartilages, shown in Figure 6.2(c), also reflected
these results, in addition to the proteoglycan distribution shown in Figure 6.2(a).
Figure 6.2: (a) Variation in proteoglycan concentration (indicated by light absorbance by safranin-O) with depth for samples harvested from four locations of knee cartilage (i.e. lateral femur, medial femur, lateral tibia, medial tibia) and from central humeral head in shoulder joint; (b) Area under absorbance curve, i.e. proteoglycan content in LF, MF, LT, MT and H; (c) LF, MF, LT, MT and H cartilage stained by 0.1% safranin-O indicating the proteoglycan variation with depth
Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage153
These results together explain the anticipated reasons for the differences
observed in the findings on kangaroo shoulder cartilage in Chapter 5 and the reported
studies in the literature for knee cartilage. In other words, chondrocytes synthesise
the cartilage matrix based on the mechanical stimuli the tissue receives: the higher
the magnitude and frequency of compressive forces, the higher the amount of
proteoglycans in the tissue. Due to the higher amount of proteoglycans in knee
cartilage, it can also be postulated that its contribution to the tissue behaviour is
considerably higher than in shoulder cartilage. This, in fact, could be the case for
the results discussed later in this chapter (Section 6.5.3). In the following sections,
we further prove this by showing the collagen structural differences and their
possible implications for the mechanical behaviour differences in knee and shoulder
cartilage.
6.5.2 Differences in collagen network of knee and shoulder cartilage
The PLM images of samples harvested from the MF of a kangaroo knee joint are
shown in Figure 6.3(a). At 0°, when the articular surface of the samples was facing
the microscope, confirming the cross-polarised configuration, the light did not reach
the eyepiece and thus appeared completely dark. When the sample was rotated in 45°
increments, sequential bright and dark images were observed (Figure 6.3(a)). This
was similar for samples harvested from the LF, LT and MT. However, there were
differences between the samples harvested from the knee and humeral head which
are discussed later.
In images that appeared bright (e.g. the image at 45° in Figure 6.3(a)), the
superficial zone was brightly visible, followed by a dark zone which was the
transition or middle zone of the cartilage. The large bright zone that appeared next to
154 Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage
the middle zone was the deep zone or radial zone of the cartilage. These
characteristic images are typical for cartilage samples exposed to a polarised light
and have been reported frequently in the literature [320, 321]. The dark appearance
in the transitional zone was due to the random orientation of the fibres. However, in
the superficial zone and radial zone there were definitive predominant directions of
fibre arrangements parallel to the cartilage surface and perpendicular to the
subchondral bone, respectively.
Figures 6.3(b) and 6.3(c) show the gray value of the polarised light transmitted
through the samples after converting the PLM images to 32-bit images in the ImageJ
software. The image analysis gave a quantitative indication of the birefringence or
anisotropy of the cartilage samples. As observed in these graphs, to the upper end of
the graph there is an initial dip in transmitted light values, which is the top of the
cartilage surface. As shown (Figures 6.3(b) and 6.3(c)), for the ease of image
comparison, relative distance has been normalised so that its value at the start of the
subchondral bone is set to unity. A comparison of Figure 6.3(b) and Figure 6.3(c)
clearly indicates that the first set of samples (0°, 90°, 180° and 270° angles) did not
transmit light, while the second set (45°,135°,225° and 315° angles) showed a
definite variation in transmitted light from surface to bone.
Similar to reported studies, based on the graphs in Figure 6.3(c) it can be
observed that the maximum contrast or brightness was for images taken at 45°. In
these images, the initial peak is related to the brightness observed in the superficial
collagen layer. This peak drops to a relatively low value and gradually increases to a
higher value, as seen in the plateau region of the graphs. This variation corresponds
to the change from random fibre configuration (dark appearance) in the transition
zone to a radial configuration in the deep zone, which is the brightest under polarised
Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage155
light. Similar fibre orientation variations were observed for samples from the rest of
the locations of the knee, while differences were observed for shoulder cartilage as
discussed next.
Figure 6.3: Typical images of cartilage when exposed to polarised light – The dark image at 0° angle corresponds to the cross-polarised configuration; in a sequence of every 45°, bright images are visible indicating the polarised light has been transmitted through the samples; (b) For 0°,90°, 180° and 270° angles, the light transmittance through the samples is literally uniform; (c) For 45°,135°, 225° and 315° angles, the variation in light transmitted through depth corresponds to the zonal arrangement of cartilage fibres (These results are for typical samples harvested from a medial femur of kangaroo knee cartilage)
Figures 6.4(a) and 6.4(b) show the PLM images for LF, MF, LT, MT and H
cartilage and their respective birefringence profiles with depth. Although there seem
to be apparent differences such as the size of the transitional zone and the brightness
of the superficial layer, the birefringence profiles of the knee cartilage samples, in
general, were similar (Figure 6.4(b)). In contrast, the PLM images of the shoulder
cartilage samples (H in Figure 6.4(b)) were characterised by a bright superficial layer
followed by a dark region up to the subchondral bone. Therefore, it seemed that the
156 Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage
fibres in the region, after the superficial layer of the H cartilage, were randomly
oriented and there seemed to be no apparent radial zone. In order to further confirm
this finding, 10 μm samples were stained with standard picrosirius-red staining
protocol and imaged using a confocal laser microscope (Nikon A1R confocal, Nikon,
Japan) with a 40x Nikon oil-immersion objective lens. For picrosirius-red staining,
samples were stained in 0.1% picrosirius red (Sirius red F3B and saturated picric
acid) for 90 minutes.
Figure 6.4: (a) PLM images obtained from four locations of kangaroo knee (LF, MF, LT and MT) and from the central humeral head; (b) Depth-dependent light transmittance profiles of the samples (i.e. LF, MF, LT, MT and H)
Figure 6.5(a) shows the confocal images taken near the subchondral bone of a
typical sample harvested from the humeral head of a kangaroo shoulder. Although
Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage157
not definitive, the fibre direction was observed to be perpendicular near the bone
(indicated by the red arrow in Figure 6.5(a)). However, careful observation showed
that the fibre arrangement in the top portion of Figure 6.5(a) seemed to have already
changed to a random arrangement (the black arrow in Figure 6.5(a)). This random
fibre arrangement and the transition of fibres from a perpendicular arrangement to a
random arrangement are clearly visible in Figure 6.5(b) and Figure 6.5(c),
respectively. Therefore, it seems that only a small deep zone of perpendicular fibres
is present in shoulder cartilage, while the transitional zone is noticeably high. This is
arguably due to the lower compressive forces experienced by the shoulder cartilage.
The PLM images in Figure 6.4(a) show that the deep zone was clearly visible
in all the samples harvested from four locations in the knee joint. In contrast, the
deep zone of the shoulder cartilage was not visible in the PLM images, while the
confocal images confirmed that the deep zone of the shoulder cartilage was
considerably small. The size of the deep zone was largest in the MF and tended to
vary with location. Therefore, the results altogether pointed to the conclusion that the
fibre arrangement might have adapted to the magnitude, frequency and mode of
forces experienced by the tissue. Given this, it is highly possible that MF cartilage is
the largest compressible load-bearing cartilage in kangaroo knee, while the
magnitude and frequency of the compressive loads on shoulder cartilage might be
considerably small. By considering and comparing the PLM and confocal images
obtained for both shoulder and knee cartilage it can be confidently concluded that the
superficial layer is the most prominent feature of the shoulder collagen network and
hence may considerably affect the mechanical behaviour.
158 Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage
Figure 6.5: (a) Confocal image of a typical cartilage sample harvested from the kangaroo humeral head indicating the region (red arrow) near the calcified bone where fibres are perpendicular to the bone – in the same image, near to the top (black arrow), the transition from perpendicular fibres to a random fibre arrangement can be observed; (b) Confocal image enlarged from top region of image (a) indicating random fibre arrangement; (c) Transition from near perpendicular to random fibre arrangement is shown in this enlarged figure taken from the transitional area of image (a)
Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage159
6.5.3 Comparison of strain-rate-dependent mechanical behaviour and biomechanical properties of knee and shoulder cartilage
In order to compare the mechanical behaviour and properties of knee and shoulder
cartilage and to investigate the proteoglycan and superficial collagen contributions to
the strain-rate-dependent behaviour of knee cartilage, a similar mechanical testing
procedure employed for shoulder cartilage (Section 5.4) was conducted on knee
cartilage. In other words, 8 mm diameter osteochondral cartilage samples (2–3 mm
subchondral bone) harvested from the MF were tested at 10-4/s, 5x10-4/s, 5x10-3/s and
10-2/s strain-rates. The MF was chosen for this study based on the above results of
the PLM and histological studies. The PLM study indicated that the collagen
architecture of MF cartilage is structurally different from the shoulder cartilage and
that proteoglycan concentration profiles with depth in the MF cartilage is
inhomogeneous. It is believed that the conclusions made in the follow-up study based
on MF cartilage would not change had the tests been performed on samples taken
from the LF, LT and MT of the knee. This is because, in the previous sections, it was
shown that collagen network of samples from all four locations had similar
anisotropic and inhomogeneous nature.
The harvested MF samples were subjected to 25% engineering strain where
thickness was estimated based on the average ultrasound speed reported for knee
cartilage in the literature, which is 1627 m/s [323]. Next, the samples were incubated
in 0.1 mg/ml trypsin–PBS solution at 37 oC for 4 hrs to completely remove the
proteoglycans and were tested again at the aforementioned four strain-rates.
Subsequently, the proteoglycan-removed samples were also treated with 30 U/ml
collagenase for 44 hrs to severely degrade the superficial collagen. Afterwards, the
mechanical testings were conducted again. The objective was to quantify the
contribution of proteoglycans and superficial collagen to the mechanical behaviour
160 Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage
of kangaroo knee cartilage and to compare the findings with the conclusions made
for shoulder cartilage. The same experimental protocols described in Chapters 4 and
5 were followed in these testings too.
6.5.3.1 Contribution of proteoglycans to the strain-rate-dependent behaviour of kangaroo knee cartilage
The results indicated that kangaroo knee cartilage was strain-rate-dependent (Figure
6.6(a)). The average thickness (1.04±0.23 mm) of the knee cartilage samples (n=9)
was significantly larger than the average thickness (0.72±0.13 mm) of the tested
shoulder cartilage samples (n=51) (p<0.005). The reasons for this difference in
thicknesses can be attributed to the different forces experienced and differences in
joint congruence in knee and shoulder joints [324]. The average stiffness of the knee
cartilage (n=9), indicated by the Young’s modulus, was also higher than that of the
shoulder cartilage (n=51) at all four strain-rates tested (p<0.05). Statistical power
analysis indicated that sample size, n=10, is sufficient (Power=0.8049) to
differentiate between mean stiffness of shoulder and knee cartilage.
Trypsin treatment reduced the stiffness of the knee cartilage significantly at all
strain-rates as indicated by the decreases in Young’s modulus (p<0.005) (Figure
6.6(c)). Although the treatment degraded the ability of the tissue to respond to strain-
rates, it still demonstrated the strain-rate-dependency (Figure 6.6(a)), similar to
observations made on the shoulder cartilage in Chapter 5. For comparison purposes,
variations of normalised force with indentation are shown for both knee and shoulder
cartilage when the proteoglycans were completely removed (Figure 6.6(a) and Figure
6.6(b)). The results clearly indicated that the removal of all proteoglycans had much
more effect on the knee cartilage at all strain-rates compared to the shoulder
cartilage. This was further demonstrated when the percentage decreases in Young’s
modulus due to trypsin treatment for knee and shoulder cartilage were compared
Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage161
(Figure 6.6(d)). In other words, the percentage decrease in Young’s modulus for the
knee cartilage was 83.7±5.9%, 85.0±3.6%, 64.7±15.4% and 53.6±19.5% for 10-4/s,
5x10-4/s, 5x10-3/s and 10-2/s strain-rates, respectively. For the shoulder cartilage,
however, it was approximately in the range of 35–40%, as reported in Chapter 5 and
demonstrated in Figure 6.6(d). Further, this percentage decrease in Young’s modulus
of the knee cartilage was significantly higher than for the shoulder cartilage
(p<0.005). These results suggest that proteoglycans in the knee cartilage play a
relatively larger role in the mechanical behaviour of tissue than the proteoglycans in
the shoulder cartilage. These results are reasonable and not surprising considering
that knee cartilage bears more compressive load than shoulder cartilage. Hence, the
amount and concentration of proteoglycans are higher in the knee cartilage, and
hence the contribution of the proteoglycans to the mechanical behaviour of the tissue
is higher.
It was intriguing to note that the percentage decreases in Young’s modulus at
the lowest two strain-rates (i.e. 10-4/s and 5x10-4/s) were considerably larger than at
the highest two strain-rates (i.e. 5x10-3/s and 10-2/s), with a difference of
approximately 25%. Further, the percentage decrease in Young’s modulus at 10-4/s
was significantly different to the percentage decrease in Young’s modulus at 5x10-3/s
and 10-2/s (p<0.05). However, no significant differences in the percentage decrease
in Young’s modulus were identified at the two lowest strain-rates (p=0.512).
Moreover, the percentage decrease in Young’s modulus at 5x10-4/s was significantly
different to the decreases at 5x10-3/s and 10-2/s (p<0.005). Additionally, the
percentage decreases in Young’s modulus at the two highest strain-rates were not
significantly different either (p=0.403). Together, these results confirmed that
162 Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage
proteoglycans play a more significant role in tissue behaviour at low strain-rates (10-
4/s) than at large-strain-rates (10-2/s) in knee cartilage tissues.
Figure 6.6: Normalised force vs indentation graphs for: (a) normal and trypsin-treated (in 0.1 mg/ml for 4 hrs) knee cartilage; (b) normal and trypsin-treated (in 0.05 mg/ml for 4 hrs) shoulder cartilage; (c) the effect of complete removal of proteoglycan (due to 0.1/mg/ml trypsin treatment for 4 hrs) on Young’s modulus of kangaroo knee cartilage; (d) Comparison of percentage decrease in Young’s modulus due to complete removal of proteoglycans for kangaroo knee and shoulder cartilage
These results on kangaroo knee cartilage are remarkably consistent with the
conclusions found in previous studies in the literature [16, 19]. As mentioned in
Chapter 5, it has been noted in the literature that the equilibrium properties (extracted
at very low strain-rates) of cartilage are governed by proteoglycans, while the
dynamic properties (extracted at high strain-rates) are governed by the collagen
network. Hence, with an increase in strain-rate, the proteoglycan contribution
Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage163
decreases and the collagen network begins to considerably affect the tissue
behaviour. This is, however, different for shoulder cartilage where the proteoglycans
equally contribute to the tissue behaviour at small and large strain-rates. Further,
based on the findings in the literature [16] and the above-mentioned findings (Figure
6.6) it was speculated that mechanical testing on knee cartilage tissues in which the
proteoglycan had been completely removed and which had been treated in
collagenase for 44 hrs to degrade the superficial collagen would show that superficial
collagen contributions at low and high strain-rates are different. The results of this
testing are discussed in the following section.
6.5.3.2 Contribution of superficial collagen to the strain-rate-dependent behaviour of kangaroo knee cartilage
Collagenase treatment for 44 hrs reduced the tissue stiffness at all strain-rates and the
differences were statistically significant compared with the stiffness after trypsin
treatment (p<0.005) [6.7(a)]. The results further indicated that the contribution of
superficial collagen at low (10-4/s and 5x10-4/s) and high strain-rates (5x10-3/s and
10-2/s) were statistically different and that the contribution increased with the strain-
rate (Figure 6.7(b)). The superficial collagen contributions were 6.5±6.2%,
4.8±2.9%, 13.8±12.4% and 14.6±7.2% at 10-4/s, 5x10-4/s, 5x10-3/s and 10-2/s strain-
rates, respectively. The superficial collagen contribution values reported in the
literature for bovine cartilage are 3.6%, 6.6% and 14.8% for 5x10-4/s, 5x10-3/s and
5x10-2/s strain-rates, respectively [16]. Although these values are smaller than the
values in the present study (perhaps due to the different tissue used), the finding on
the increased superficial collagen contribution with strain-rate is similar to the
findings in the literature.
The superficial collagen contributions at the two lowest strain-rates (10-4/s and
5x10-4/s) were not significantly different (p=0.538). Likewise, the contributions were
164 Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage
not significantly different at the two highest strain-rates (5x10-3/s and 10-2/s)
(p=0.449). However, the contribution of superficial collagen to tissue behaviour was
significantly low at 10-4/s when compared with the contribution at 10-2/s strain-rate.
Probably due to high standard deviation at 5x10-3/s, the statistical differences were
not identified when the superficial collagen contributions at 10-4/s and 5x10-3/s were
compared (p=0.058). Nevertheless, the superficial collagen contribution at 5x10-4/s
was significantly different from the contribution at the two highest strain-rates
(p<0.005). Therefore, as expected, it can be stated that with an increase in strain-rate
the superficial collagen begins to considerably affect the knee cartilage behaviour,
thus confirming the results reported in both the previous section (Section 6.5.3.1) of
this study and the literature.
Figure 6.7: Percentage decrease in Young’s modulus after complete proteoglycan-removed samples were treated for 44 hrs in collagenase; (b) Contribution of superficial collagen to tissue behaviour at four strain-rates (10-4/s, 5x10-4/s, 5x10-3/s and 10-2/s)
The mechanical test results confirmed that there were definite differences in
the mechanical properties of the knee and shoulder cartilage. Even though the
responses of both cartilage types were more or less similar, especially in terms of
strain-rate-dependency, the contributions of the underlying components (in
Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage165
particular, collagen and proteoglycans) were different. The results of the present
study not only further cemented the conclusions drawn in Chapter 5 but also
confirmed the importance of superficial collagen and collagen architecture in the
behaviour of kangaroo shoulder cartilage.
6.5.4 Significance and implications for numerical modelling and tissue engineering
6.5.4.1 Implications for tissue engineering strategies
The composition and structural features of the collagen network and hence the
mechanical properties depend on the magnitude and pattern of loading experienced
by the cartilage tissue. The results of the present study demonstrate that the
proteoglycan distribution with depth and the features of the collagen network in knee
and shoulder cartilage are noticeably different. The magnitude and pattern of loading
seem to affect the size of the superficial, transitional and radial zones. Hence, it is
plausible that the mechanical and biological characteristics of these zones are linked
to the stress, strain, and fluid flow in these zones as also postulated by earlier studies
[18, 325].
Therefore, joint-specific cartilage tissue generation would require approaches
that are catered to the specific tissue of interest so as to achieve the compositional
and structural features that are appropriate to it. Nonetheless, replicating a tissue
similar to the characteristics of the native tissue (e.g. the distribution of proteoglycan
concentration with depth and the arcade-like collagen network) remains a significant
challenge in the tissue engineering field. The findings of the present study suggest
that it is important to include procedures that signal to the (new) growing tissue the
mechanical forces it will experience in the future. As a result, the tissue will be able
to facilitate its growth to a status (or to an extent) where it is able to withstand
external forces specific to the joint.
166 Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage
6.5.4.2 Joint-specific differences in the prevalence of osteoarthritis
Osteoarthritis does not affect all joints equally. It is more prevalent in the hip, knee,
spine and metatarsophalangeal and interphalangeal joints [169]. Although
degenerative changes happen in ankle, wrist, elbow and shoulder joint cartilage, they
do not progress to a disease state as frequently as in the hip and knee. The reasons
why some joints are more affected by osteoarthritis than others still remain elusive
[169, 326]. The most common example for this is the difference in progression of
osteoarthritis in the knee and ankle [169, 327]. Although loaded more heavily than
the knee, the initiation and progression of osteoarthritis in the ankle has been
identified to be less frequent than in the knee. Anatomical differences such as joint
congruence and stability are suggested as the reasons; however, these do not fully
explain the differences [169]. Extensive studies [326-328] have concluded that ankle
chondrocytes respond quickly to damage with higher proteoglycan synthesis and are
also metabolically more active in terms of degradation compared to knee
chondrocytes. This is an indication that ankle cartilage has a higher capacity to repair
than knee cartilage. Ruling out the genetic difference between knee and ankle
chondrocytes, through a series of experiments it has been concluded that the
difference in native extracellular matrix of the two tissues is the main factor affecting
the chondrocyte metabolism [169]. Hence, it is important to investigate the
differences in the native matrix of joints that have different susceptibilities to
osteoarthritis. Therefore, this study, which compared the shoulder and knee
extracellular matrix, can be considered as a necessary step towards identifying
plausible reasons for differences in osteoarthritis initiation in different joints.
6.5.4.3 Implications for numerical modelling
Modelling cartilage tissue requires an understanding about the relationship between
its structure and the respective biomechanical functions. Most numerical models
Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage167
available in the literature are for knee cartilage tissues. Only limited investigations
have been carried out to explore the structure–function relationships of shoulder
cartilage. As indicated by the results of the present study, the conclusions drawn for
knee cartilage cannot be directly generalised to other cartilage tissues, such as
shoulder cartilage, which are loaded differently. Therefore, in modelling joint-
specific cartilage tissues, differences in the collagen structural feature and
composition should be acknowledged. For example, the results in the current study
indicated differences in the proteoglycan distribution in kangaroo knee and shoulder
cartilage. Additionally, differences in the anisotropy of collagen network in knee and
shoulder cartilage in terms of the size of the superficial, transitional and deep zones
as well as differences in fibre directions in individual zones were also observed.
These differences are likely to be important in developing numerical models and
could significantly affect model predictions.
6.6 CONCLUSION AND REMARKS
In this chapter, we mainly investigated the compositional, microstructural and
biomechanical differences between knee and shoulder cartilage with the objective of
explaining the different results observed in Chapter 5 and in reported studies in the
literature. Additionally, we investigated how the microstructure and composition of
the kangaroo shoulder and knee cartilage affect the mechanical behaviour of the
tissue. The implications of the results of these investigations for the numerical
modelling of cartilage and tissue engineering strategies were also discussed in this
chapter. The following conclusions can be made from the results of the study:
• High and frequent weight-bearing knee cartilage showed a clear depth-
dependent proteoglycan distribution as opposed to the shoulder cartilage (a
low weight-bearing cartilage) where there were no apparent trends in
168 Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage
proteoglycan distribution with cartilage depth. In addition, the proteoglycan
content in the knee cartilage was higher than in the shoulder cartilage.
• Proving that the magnitude and frequency of loading affects the structural
feature of the collagen network, knee cartilage showed a distinctly different
collagen structure than shoulder cartilage. A prominent deep zone, where
collagen is anchored perpendicular to the subchondral bone, was observed in
the knee cartilage. Differences in the size of the superficial, transitional and
deep zones were also identified in the samples taken from different locations
of the knee.
• Superficial collagen appeared to be the most prominent feature of the
collagen network of shoulder cartilage. The confocal images confirmed that
the size of the deep zone in shoulder cartilage was small and the transitional
zone was considerably large.
• Proteoglycan contribution to knee cartilage behaviour was significantly larger
than collagen contribution at all strain-rates tested. However, as mentioned in
chapter 5, results for shoulder cartilage were different where collagen
contribution on the tissue behavior was larger than proteoglycans.
• In congruence with the findings in the literature, this study was able to
conclude that proteoglycans significantly affect the knee cartilage behaviour
at low strain-rates; hence, the equilibrium properties of knee cartilage tissues
would be governed by proteoglycans.
• The effect of superficial collagen on the knee cartilage behaviour increased
with increasing strain-rate, reaching a considerably high value (~15%) at the
highest strain-rate tested. However, unlike in the knee cartilage, no trend was
evident in the superficial collagen contribution with an increase in strain-rate.
Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage169
• Together, these results confirmed the findings in Chapter 5 and the reported
studies in the literature [16], demonstrating the significant contribution of
superficial collagen on the strain-rate-dependent behaviour of shoulder
cartilage.
• Based on the results, it is important to control tissue engineering strategies in
order to generate joint-specific tissues. The findings are useful for the
numerical modelling of shoulder cartilage tissue.
170 Chapter 6: Compositional, microstructural and biomechanical differences in kangaroo shoulder and knee cartilage
Chapter 7: Conclusions
7.1 CONCLUSIONS
This research explored the factors affecting the strain-rate-dependent behaviour of
kangaroo shoulder cartilage by systematic experimental and numerical studies. New
knowledge on the mechanical behaviour of kangaroo shoulder cartilage was
developed by introducing: 1) a porohyperelastic model with a strain-rate-dependent
permeability function; and 2) an experimental approach to investigating the effects of
cartilage constituents on cartilage mechanical behaviour. In addition, by introducing
kangaroo as a model, the present study has comprehensively compared shoulder and
knee cartilage with the objective of investigating the relationship between cartilage
structure and mechanical behaviour in the context of the strain-rate-dependent
behaviour of cartilage tissues. The findings of this research inform the cartilage
modelling community about extracellular matrix features and the underlying factors
that need to be considered when modelling low and high weight-bearing cartilage
tissues such as the shoulder and knee, respectively. In addition, this research informs
tissue engineers about what might be the most important features to consider when
engineering a cartilage tissue for low and high compressive load-bearing joints such
as the shoulder and knee, respectively.
Based on this research, it can be concluded that kangaroo is a suitable model
for future biomechanical research on shoulder cartilage. The biomechanical
properties and behaviour of the kangaroo shoulder cartilage tissues were in general
agreement with that of human shoulder cartilage tissues. Further, the different
loadings encountered in the kangaroo’s upper and lower limb cartilage provide a
natural source for investigating how mechanical forces affect the development,
Chapter 7: Conclusions 171
composition (e.g. proteoglycan distribution) and structure (e.g. collagen architecture)
of the low and high load bearing cartilages and the progression of osteoarthritis in
low and high load bearing joints. Future experiments using kangaroo as a model are
promising for gaining insights into tissue engineering strategies that are necessary to
develop joint-specific cartilage tissues.
The findings of the present study indicate that the mechanical behaviour of
kangaroo shoulder cartilage is strain-rate-dependent. Its solid skeleton behaviour was
adequately represented by the 2-term reduced polynomial hyperelastic model in this
study’s experiments, while lower-order material models such as the neo-Hookean
and Mooney–Rivlin models were not adequate. Based on the formulation of the
strain-rate-dependent permeability model, it was found that the permeability and
strain-rate are negatively correlated (i.e. permeability is reduced when the strain-rate
is increased). The strain-dependent and strain-rate-dependent models affected the
tissue behaviour considerably at all strain-rates tested, while at high strain-rates the
latter became more significant. Therefore, in addition to solid–interstitial frictional
fluid interaction, the pressure drag forces and possibly the inertia forces begin to play
a significant role in tissue behaviour at high strain-rates. Therefore, it was concluded
that strain-rate-dependent permeability is one of the mechanisms governing the
strain-rate-dependent behaviour of cartilage tissues. Computational models have the
potential to benefit from the inclusion of strain-rate-dependent permeability in terms
of better prediction of cartilage response to external loads, especially at high strain-
rates. Since all aspects of isotropic fluid and solid behaviour were evaluated in the
strain-rate-dependent model, the model deviations at the highest strain-rate can be
concluded to be a result of the anisotropic solid properties and fluid behavior of the
tissue.
172 Chapter 7: Conclusions
Proving the direct role of proteoglycan in facilitating the compressive loads,
the progressive removal of proteoglycans reduced the tissue stiffness gradually at all
strain-rates. Proteoglycan depletion and superficial collagen disruption substantially
compromised the ability of the tissues to resist external compressive loading at
different strain-rates, increasing the risk of bone-to-bone contact. Total proteoglycan
removal and severe degradation of superficial collagen increased the permeability of
cartilage tissue. Even after complete removal of proteoglycans or severe degradation
of superficial collagen, the tissues still exhibited strain-rate-dependency. Therefore, it
was concluded that the dense collagen meshwork has the ability to sustain the pore
size of cartilage so that solid–interstitial fluid frictional interaction is able to facilitate
the strain-rate-dependent behaviour. It was also concluded that the effect of surface
lipids on the strain-rate-dependent mechanical behaviour of cartilage tissues is
minimal.
Superficial collagen degradation had the largest effect on tissue stiffness and
strain-rate-dependency when compared with proteoglycan degradation. Hence, it was
concluded that superficial collagen plays a more significant role than proteoglycans
in facilitating the strain-rate-dependent behaviour of kangaroo shoulder cartilage. In
addition, superficial collagen contributed evenly to tissue behaviour at all strain-
rates, which confirmed its significance in governing the mechanical behaviour of
shoulder cartilage tissues. However, these findings were different from the
observations made on kangaroo knee cartilage where superficial collagen contributed
less than proteoglycans to the mechanical behaviour, although its role became
substantially more significant at the higher strain-rates tested. Based on
computational modelling, it was found that collagen disruption would lead to
shoulder cartilage being damaged faster than when proteoglycans were depleted.
Chapter 6: Conclusions 173
Therefore, for shoulder cartilage, collagen disruption can be regarded as more
damaging than proteoglycan depletion.
Adaptation of the cartilage tissues to the mechanical environment was
confirmed by the conclusions drawn from the study of the composition and
microstructure of both shoulder and knee cartilage. It was concluded that knee
cartilage (high and more frequent weight-bearing) has an inhomogeneous
proteoglycan distribution compared to shoulder cartilage (low weight-bearing). In
addition, the proteoglycan content in knee cartilage was considerably higher than in
shoulder cartilage, indicating that the magnitude and frequency of loading affect the
proteoglycan distribution and content. The collagen network of knee cartilage
showed a distinctly different collagen structure than the collagen network of shoulder
cartilage. Although structurally similar, differences in the size of the superficial,
transitional and deep zones were also identified in different locations of the knee.
Therefore, it can be concluded that the structural features of the collagen network
also adapt to external mechanical loading. Superficial collagen was the most
prominent feature of the shoulder collagen network and the microstructure from the
confocal study confirmed that the size of the deep zone in kangaroo shoulder
cartilage is small (i.e. fibres perpendicular to the calcified zone were rarely visible in
the shoulder cartilage).
In summary, the research conducted in this project explored the factors
affecting the strain-rate-dependent behaviour of shoulder cartilage tissues. The main
findings of this study have significant implications for the computational modelling
of shoulder cartilage tissues and tissue engineering approaches to engineering joint-
specific cartilages.
174 Chapter 7: Conclusions
7.2 LIMITATIONS
In terms of experimental limitations, ideally, the present study required human
cartilage tissues. However, obtaining human tissues, especially in the required
quantities, was not feasible due to ethical considerations and difficulties in obtaining
young to mature tissues. Therefore, kangaroo was chosen as the animal model which
the findings of the present study indicated as a suitable for biomechanical
investigation of the shoulder cartilage. However, there were additional limitations
such as the inability to obtain shoulder and knee joints from the same animal because
animals were not specifically euthanised for the present study.
The scope of the present study had to be limited to indentation tests. Although
tension tests are required to develop a comprehensive FE model for the tissue tested,
there were no appropriate resources available to carry out tension tests. An
investigation into the effect of osmotic pressure on the strain-rate-dependent
behaviour of the kangaroo shoulder cartilage was not conducted. However, it would
be important to carry out that study in order to comprehensively understand the
response of the shoulder cartilage to external environmental changes (i.e. changes in
saline concentration).
7.3 FUTURE RESEARCH DIRECTIONS
Despite vast knowledge on cartilage tissues there still exists a significant need for
more research to be carried out in order to achieve the ultimate objectives such as
facilitating the early diagnosis of osteoarthritis, understanding the reasons for
osteoarthritis and engineering cartilage tissues that can function appropriately in vivo
under static and dynamic loading. The extension of this research can mainly be
divided into two parts: improvement of the porohyperelastic cartilage model, and
experimental investigation of the dynamic fluid and solid-skeleton behaviour.
Chapter 6: Conclusions 175
In terms of improving the porohyperelastic FE model with strain-rate-
dependent permeability, which was developed in the present study, tension
experiments on cartilage tissue should be performed to obtain the transverse material
parameters. The inclusion of transverse material parameters in the model could
improve the model results, as implied by the results in Chapter 5. Additionally,
follow-up studies investigating the effect of cross-linking on strain-rate-dependent
cartilage tissue behaviour may provide information on how cross-linking, which is
one of the main features of the load-bearing unit of cartilage, facilitates the dynamic
load-bearing ability of cartilage tissues.
The potential techniques that can be utilised to investigate the fluid behaviour
of cartilage tissue under dynamic loading are high-resolution MRI coupled with
microscopy–micro/nano indentation techniques. High-resolution MRI coupled with
simultaneous mechanical testing can provide potentially valuable information about
fluid behaviour in normal and osteoarthritis-affected cartilage. Moreover, this
methodology would also be valuable in assessing the performance of engineered
cartilage tissues so as to identify whether fluid support is appropriate for long-term
performance.
There are a few plausible ways to investigate collagen behaviour during
dynamic loading. One way is to couple confocal microscopy with micro/nano
manipulation devices or tracking cells during deformation in order to infer
information regarding the potential deformation of collagen fibre as a bulk. Injecting
particles, in particular fluorescence particles, and tracking them using imaging
techniques can also provide information regarding collagen behaviour under dynamic
loading. The results from such studies, if successful, would provide important
information regarding differences in the structural response of normal and
176 Chapter 7: Conclusions
osteoarthritis-affected cartilages. This information could also be used for the
validation of existing and future multiscale numerical models.
Future research that compares the development processes of cartilage tissues
from joints that have different susceptibilities to osteoarthritis (i.e. the knee and
shoulder) will provide insights into the underlying reasons for these different
susceptibilities and thereby lead to the identification of possible preventive strategies
for osteoarthritis. In this context, kangaroo can serve as a valuable animal model for
future research studies in this area.
If conducted, these recommended studies and methodologies will hopefully
contribute to better understanding the internal behaviour of cartilage tissues in
addition to the physiological factors affecting the health and development of the
tissue. Thereby, it will also provide valuable insights into the development of
functionally viable cartilage tissues and the factors affecting osteoarthritis
development.
Chapter 6: Conclusions 177
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