1. Introduction

Figure 1. Diagram illustrating a synovial joint featuring an AC structure. The primary components determining the frictional and mechanical properties of AC are mainly collagens, particularly type 2 collagens (55-75%), proteoglycans (15-30%), and lipids, specifically phospholipids (10%). These components vary across the superficial (10%), middle (40-60%), and deep (30%) zones of the AC. Collagen fibers are densest in the superficial zone (1-200 µm thickness, 5-20 nm diameter), providing resistance against compressive, tensile, and shear forces. In the middle and deep zones, collagen fibers have a similar size with a diameter of 30-5000 nm. Proteoglycan density is lower in the superficial zone but increases in the middle and deep zones. The arrangement of extracellular matrix consisting of collagen fibers and proteoglycans in each zone contributes to the mechanical strength of the AC, with parallel arrangement in the superficial zone and more random arrangements in the middle and radial arrangement in the deep zones [21,22,23,24,25,26,27].

Figure 1. Diagram illustrating a synovial joint featuring an AC structure. The primary components determining the frictional and mechanical properties of AC are mainly collagens, particularly type 2 collagens (55-75%), proteoglycans (15-30%), and lipids, specifically phospholipids (10%). These components vary across the superficial (10%), middle (40-60%), and deep (30%) zones of the AC. Collagen fibers are densest in the superficial zone (1-200 µm thickness, 5-20 nm diameter), providing resistance against compressive, tensile, and shear forces. In the middle and deep zones, collagen fibers have a similar size with a diameter of 30-5000 nm. Proteoglycan density is lower in the superficial zone but increases in the middle and deep zones. The arrangement of extracellular matrix consisting of collagen fibers and proteoglycans in each zone contributes to the mechanical strength of the AC, with parallel arrangement in the superficial zone and more random arrangements in the middle and radial arrangement in the deep zones [21,22,23,24,25,26,27].

2. Chronology of cartilage lubrication

Figure 6. The human gait cycle is shown representing the possible lubrication mechanism at the cartilage interface to support the load (Recreated from reference) [97].

Figure 6. The human gait cycle is shown representing the possible lubrication mechanism at the cartilage interface to support the load (Recreated from reference) [97].

2.1. How does Lubrication models fail in the case of OA

Previous discussions regarding lubrication models for AC can be categorized by drawing from established tribological lubrication models, which include fluid-film, mixed, and boundary lubrication. These models have been adapted to account for the structure and function of AC, resulting in modifications to these lubrication models. When examining the concept of fluid film formation within AC, several models have been developed, including hydrodynamic, weeping, elastohydrodynamic, micro-elastohydrodynamic, and tribological rehydration models. Additionally, models have been proposed based on the formation of mixed film modes in AC, such as osmotic, squeeze film, boosted, biphasic, and triphasic models. Lastly, hydration model has been explained considering the lubrication effects of synovial constituents and their boundary lubrication properties. All these models focus on comprehending how healthy cartilage is lubricated with SF at the interface to ensure smooth functioning. However, when it comes to OA, the effectiveness of these models in explaining the progression of cartilage degeneration becomes questionable. In this section, we will dive into each model to assess its relevance in the context of OA.

2.1.1. Fluid-film lubrication

Fluid film lubrication theory was originally developed within the conventional context of tribological lubrication in journal bearings. The interface is separated by the pressure generated by the lubricant film at the contact point which has been adapted for understanding AC lubrication in synovial joints. This adaptation has led to the emergence of various theories, including hydrodynamic, hydrostatic (or weeping), elasto-hydrodynamic, micro-elasto-hydrodynamic, and tribological rehydration models. These lubrication models have been formulated while considering the function and physiological characteristics of AC.

The proposal of a hydrodynamic lubrication mechanism is based on the notion that SF in synovial joints creates a wedge-shaped fluid layer at the articulating junction as shown in Figure 7a [83,108]. This fluid layer prevents direct contact between the two surfaces by acting as a thin lubricant and generating a hydrodynamic lift (about 10 µm [128]). However, unlike typical hydrodynamic bearings that require high-speed transverse motion to establish this lift (500 to 4000 rpm for 40mm radius bearings [129]). The pin on plate experiments performed by Gleghorn et.al., using bovine AC failed to obtain the hydrodynamic lubrication even at high speed 50mm/s and low strain 5% [69]. The non-conformal nature of biological joint surfaces and absence of consistent high-speed and lighter loads, make the contacting geometries unfavorable for this mechanism. In the case of OA, the properties of SF can vary due to aging or damage at the articulating surfaces. The reduction of constituents such as hyaluronic acids, lubricin in diseased joints can affect the viscosity of the SF which could not support interstitial fluid pressurization (viscosity reduction was reduced by the factor 20 at shear rate of 10^-1s^-1 and a factor of 3 at shear rate of 10³s^-1) [130-132]. Hence, hydrodynamic lubrication is insufficient to explain in degradation mechanism of OA.

Hydrostatic / Weeping lubrication was introduced by McCutchen in 1959, drawing inspiration from hydrostatic bearings [109,110]. According to this theory, when two articulating surfaces are compressed, the interstitial fluid is forced out of the AC pores, resulting in a pressurized fluid film that supports the solid matrix as shown in Figure 7b [107,133]. Consequently, the frictional forces primarily act on the portion of the load that is transferred across the solid matrix, which signifies the load sharing mechanism [85]. These findings were substantiated through experimental observations of cartilage on glass over a period. Further they demonstrated that the frictional force did not decrease when the load was removed for a brief period of 1s, but rather increased upon reapplication of the load [133]. This mechanism provided an explanation for the time-dependent frictional behavior of AC, which could not be accounted by hydrodynamic or boundary lubrication theories. However, this proposed mechanism sparked controversies regarding the principle of conservation of linear momentum [134]. The concept of hydrostatic lubrication cannot be applied in the case of OA due to the changes in the permeability of the cartilage surface. Researchers have identified that the permeability of the cartilage changes in the extracellular matrix structure allowing rapid fluid flow causing uneven deformation [21]. This permeability is elevated from 4.19 ± 3.78×10⁻¹⁷ m⁴/N s to 10.2 ± 9.38×10⁻¹⁷ m⁴/N s influencing the biomechanical properties of the chondrocytes in the AC tissue [135]. Thus, the fluid load support as mentioned in the weeping model fails to lubricate the degraded cartilage interface.

Figure 7. (a) Hydrodynamic lubrication mechanism shown at the AC interface along with the schematic representation of the pressure generated by the interstitial fluid flow (b) The hydrostatic/weeping lubrication at the AC interface is shown in the figure with the schematic representation of the flow of interstitial fluid load support during the loading and unloading cycles (c) The interaction of the asperities at the AC surface with the surface is shown in the figure and the pressure is being generated at these interfaces due to the hydrodynamic lubrication.

Figure 7. (a) Hydrodynamic lubrication mechanism shown at the AC interface along with the schematic representation of the pressure generated by the interstitial fluid flow (b) The hydrostatic/weeping lubrication at the AC interface is shown in the figure with the schematic representation of the flow of interstitial fluid load support during the loading and unloading cycles (c) The interaction of the asperities at the AC surface with the surface is shown in the figure and the pressure is being generated at these interfaces due to the hydrodynamic lubrication.

Previous lubrication models did not take into account the deformation of AC as a factor in explaining the frictionless motion of synovial joints [86]. Dintenfass developed a AC lubrication model based on elastohydrodynamic theory, considering the viscous resistance of the SF (viscosity of SF is 10 Ns/m² at 0.1 s) and the elastic deformation (10.78 MPa when compressed in water) of the AC surface as shown in Figure 7c [136]. This model was supported by the observation that the viscosity of SF is negligible under pressure and the elasticity modulus of the AC is very small [136,137]. Therefore, when a high load is applied, the film thickness formed at the interface is minimal due to the flattening of the AC surface and the increased contact zone. This flattening phenomenon of the AC surface under physiological load was later referred to as the micro-elastohydrodynamic effect [88,138]. However, for the elastohydrodynamic lubrication the film thickness at the AC interface for an healthy joint was in the range of 0.1 to 1 µm with surface roughness of 2-5 µm thus this theory was substantiated by various theoretical calculations and experimental results [139]. Tanner also calculated the film thickness of hip joints during normal walking speeds, yielding a thickness of 10^-5cm or greater at 20cm² area with friction coefficient of 0.003 [128]. These experimental results demonstrated the formation of a film thickness during relative motion. But in terms of the degraded AC the surface roughness increases due to superficial fronding at the upper layer. The surface roughness with arithmetic average absolute values, maximum peak heights and the mean spacing between local peaks of the OA cartilage surface was 71%, 80% and 51% respectively to the healthy cartilage [140]. These studies shows that the load bearing ability of the AC cannot be explained using the micro-elasto or elasto hydrodynamic lubrication models.

2.1.2. Boundary lubrication

When the hydrodynamic lubrication mechanism was proposed, it faced skepticism from researchers for not considering the contribution of synovial compositions. In 1959, Charnley postulated the presence of boundary lubrication on the AC surface which was due to the adherence of the SF at the AC interface [116,141-143]. The adherence, or boundary action, is governed by the glycoprotein fraction of SF, which forms a surface film with a thickness ranging from 1 to 100 nm [144]. Swann et al. demonstrated that polypeptide chains within SF functioning as boundary lubricants when isolated from the SF [117,118]. Subsequently, researchers began investigating the individual components of SF to determine their contributions to the boundary lubrication nature [119]. The main constituents identified were hyaluronic acid, aggrecans, lubricin, and phospholipids. Several reviews have outlined the significant contributions of these constituents and the techniques employed to study their boundary lubricant properties [145,146]. The complete explanation of the boundary lubrication mechanism at the articulating surfaces of biological joints cannot be solely attributed to the presence of hyaluronic acids, lubricin and phospholipids constituted in SF [147]. Through in-depth studies on friction conducted using surface force balance measurements, it has been observed that hydrated ions are trapped within the phospholipid solutions [148]. These hydrated ions act as effective lubricants by creating a strong repulsion during compression, facilitated by the repulsive hydration shells surrounding the trapped counterions. This phenomenon is referred to as the hydration lubrication mechanism (Figure 8) [120,149]. The underlying reason for this low friction behavior is the formation of a slip plane between closely packed phosphocholine liposomes, wherein the outer surfaces of liposomes oppose each other within a group [150]. Studies on degraded cartilage indicate that an increase in surface roughness and permeability of the cartilage leads to an extension of the boundary mode plateau, resulting in higher friction at the interface. This is primarily attributed to the loss of boundary lubricants present in the SF at the AC interface. To counteract this, boundary modes can be achieved by either increasing the viscosity of the lubricants (for instance, if the lubricant viscosity for healthy cartilage is 150 mPas, degraded cartilage might require 1500 mPas to achieve healthy friction values) or by elevating the sliding speeds at the interface [151]. By understanding and harnessing this mechanism, it becomes possible to develop intra-articular injections containing suitable lubricants that effectively reduce friction. Furthermore, this knowledge can be applied in the field of tissue engineering scaffolds for diseased cartilage [147].

2.1.3. Mixed-mode lubrication

The film thickness formation in the cases of hydrodynamic lubrication and the boundary effect of the synovial constituents cannot completely answer the complex lubrication mechanism of the synovial joints. There must be different modes of lubrications that comes together into the play for this complex structure that helps in explaining its low friction and wear properties [98]. This mixed lubrication regimes can be called adaptive multi-mode lubrication that can explain each individual phase of gait cycles. Based on this knowledge various lubrication models such osmotic, squeeze film, boosted, biphasic, triphasic lubrication models were proposed.

Squeeze-film lubrication model was proposed by Dowson in 1966 suggests that during joint motion the SF between the AC surfaces in synovial joints experiences compression (around 18MPa in physiological loading condition [152]), resulting in the generation of pressure [153]. This pressure is attributed to the resistance offered by the lubricant as it is squeezed out by the contacting surfaces as shown in Figure 9a. Consequently, the articulating surfaces are separated by a fluid film, leading to a localized depression where the lubricant becomes trapped supporting that load [154]. Similarly, Maroudas and Walker et al., proposed boosted lubrication which suggested when two articulating surfaces are compressed, the solvent component of SF enters the pores of the AC surface as shown in Figure 9b. While the solute or concentrated part of SF remains behind providing lubrication and this phenomenon was called boosted lubrication [94,155-158]. Both squeeze film and boosted lubrication fails in the case of OA due to the loss of proteoglycans which leads to a decrease in the dynamic and equilibrium compressive modulus of the AC [159,160]. This reduction allows for increased hydraulic permeability, resulting in decreased interstitial fluid load support during loading, ultimately leading to more cartilage damage [159].

Previous lubrication mechanisms were unable to provide a satisfactory explanation for the time-dependent frictional behavior exhibited by cartilage [74]. This behavior can be understood by examining the stress-relaxation or creep nature of AC which are directly associated with the release of interstitial fluid from the surface during tissue creep [67,123,124,161-163]. These properties of AC were elucidated by the biphasic theory proposed by Mow et al. When AC is subjected to compression or tension, the flow-dependent and flow-independent behaviors, as well as the mechanism of viscoelasticity, can be understood based on the change in the modulus values and the stress relaxation experiments (Figure 9c) [126,164]. This theory provides an explanation for the viscoelastic behavior of AC in various configurations, including tension, compression, and indentation [165]. According to the biphasic theory, AC is composed of a binary mixture consisting of an incompressible elastic solid (collagen and extracellular matrix) and an inviscid-incompressible fluid (interstitial fluid) [166,167]. Various researchers reported experimental measurements of interstitial fluid pressures calculating the aggregate elastic modulus of AC from creep experiments to be around 0.7 ± 0.09 MPa and from stress relaxation experiments to be around 0.76 ± 0.03 MPa [166-168]. Oloyede and Broom demonstrated this phenomenon through confined compression creep loading ex-situ experiments [169], while Soltz and Ateshian explained it using AC behaviours such as creep, stress relaxation, and dynamic loading in both confined compression (aggregate compression modulus = 0.64 ± 0.22 MPa and axial permeability = 3.62 ± 0.97 × 10^-6 m⁴/Ns) and unconfined compression ex-situ experimental setups (tensile modulus = 12.75 ± 1.56 MPa and radial permeability = 6.06 ± 2.10 × 10^-16 m⁴/Ns ) [125,167]. These findings have established that the viscoelastic behavior of AC is solely influenced by the intrinsic mechanical properties [170] such as porosity and permeability of solid phase in AC, as well as the frictional resistance arising from the drag exerted by the interstitial fluid. Furthermore, a state of stress relaxation equilibrium can be achieved when the flow of interstitial fluid comes to a halt. In this equilibrium state, the entire load is borne by the solid matrix, providing an explanation for the load deformation response observed in AC tissue [171,172].

As an extension of biphasic theory, triphasic theory was put forth which introduces the concept of an additional phase known as the ionic phase, distinct from the interstitial fluid. This model examines the mechanical properties of AC by incorporating three distinct phases: the solid phase, the interstitial fluid phase, and the ionic phase. In this model, the solid phase of the tissue is described as a homogeneous, isotropic, linearly elastic material that experiences infinitesimal strain within the interstitial fluid phase [126,127,167]. Considering these models, effective lubricant-based solutions can be identified depending on the properties of the AC. However, in the case of degraded cartilage, the alterations, including elevated water content within the AC structure, hinders the generation of proper fluid pressure during applied loading. Furthermore, the degenerative process impacts the hydraulic permeability of the superficial zones, thereby influencing the flow-dependent and flow-independent viscoelastic behavior of AC.

3. Relevance of AC lubrication theories to OA

4. Conclusion

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