## 3 Biomechanics of Articular Cartilage
### Lubrication Mechanisms in AC  
In normal synovial joints, the range of the friction coefficient is \(\upmu\approx\) 0.002-0.02 [185; 186; 187; 188]. Measurements of friction on both the AC and whole joint surfaces may depend on the energy dissipation mechanisms occurring during articulation [188]. Such a phenomenon is the result of the viscoelastic deformation of adjacent tissues, such as the synovial membrane, muscles, or ligaments, during joint rotation. Over the years, a number of studies have been conducted on the exact mechanism of lubrication, but to date, no answer has been provided as to why friction is so low in a healthy synovial joint [189; 187; 190; 191; 192; 193]. The low friction and wear of AC is probably related to the complex structure as well as the composition of the AC surface. Fluid film lubrication is based on the separation of sliding surfaces by a fluid film of much greater thickness than the surface roughness. Initially, the models describing the SF layer were based on the classical theory of hydrodynamic lubrication known from mechanical engineering, and in later work, these were extended to include the deformation properties of AC during movement [194; 195; 196]. The work of [PERSON], [PERSON], [PERSON] and others includes models that take into account the films between the articular surfaces formed by the outflow of interstitial fluid during AC compression [197; 198; 199; 200; 201; 202; 203]. Due to the fluid film, contact between the surfaces is reduced, friction and wear are minimised, and interstitial fluid pressure supports much of the cartilage load [188]. Consequently, regardless of the presence of a fluid layer between the sliding surfaces, the actual contact stresses between the AC surfaces may exhibit lower values than the high pressures that are recorded directly at their boundary [204]. This is because the hydrostatic pressure of the interstitial fluid plays a key role in load transfer [202]. Many tribological experiments using AC explants were conducted at physiological strain levels, but at relatively low pressures (\(\leq\)2 atm) [188]. This value is approximately 100 times lower than the maximum pressures observed in vivo on joint surfaces, regardless of their origin [205; 206]. The mechanism of load transfer is realised based on biphasic models, where the AC is treated as an incompressible, porous-permeable matrix composed of a collagen-proteoglycan network in which the interstitial fluid exhibits incompressible fluid behaviour [178; 207]. These models are used in mapping the biomechanical response of the AC subjected to loading, as well as in experimental studies using excised specimens [186; 208; 209; 210]. In their studies, [PERSON] et al. showed that during tests conducted at an applied pressure of \(\approx\)1.5 atm with a sample diameter of 6 mm, the friction coefficient increased from 0.04 to 0.18 within 15 min, with a simultaneous decrease in the interstitial pressure, which confirms the validity of the used model [186; 211; 212]. Many previous studies have assumed that interstitial pressure is responsible for transferring most of the load between opposing AC surfaces in vivo [213; 214; 215; 216; 217]. \begin{table}
\begin{tabular}{c c c c c c} \hline \hline
**Measurement Method** & **Measured Parameters** & **Test Characteristics** & **Advantages** & **Limitations** & **References** \\ \hline \multirow{3}{*}{Dynamic shear test} & Complex shear modulus, & Cyclic shear load application & Dynamic response & Complex apparatus, analysis, testing model & \multirow{3}{*}{[PERSON] et al. [175]} \\  & Shear loss angle & & of viscoelastic properties & dependent interpretation & \\ \hline \multirow{3}{*}{Equilibrium shear test} & \multirow{3}{*}{Equilibrium shear modulus} & Slow deformation & Measurement of elasticity after stabilisation of fluid flow & Long testing time, requires precision & \\  & & & & fluid flow & \\ \hline \hline \end{tabular}
\end{table}