The Boundary Lubrication of Chemically Grafted and Cross-Linked

Dec 7, 2011 - and synovial fluid at high concentrations; yet despite numerous studies, the role of ... The articular joint is an excellent water-based...
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The Boundary Lubrication of Chemically Grafted and Cross-Linked Hyaluronic Acid in Phosphate Buffered Saline and Lipid Solutions Measured by the Surface Forces Apparatus Jing Yu,† Xavier Banquy,† George W. Greene,‡ Daniel D. Lowrey,‡ and Jacob N. Israelachvili*,†,‡ †

Department of Chemical Engineering and ‡Materials Department, University of California, Santa Barbara, California 93106, United States ABSTRACT: High molecular weight hyaluronic acid (HA) is present in articular joints and synovial fluid at high concentrations; yet despite numerous studies, the role of HA in joint lubrication is still not clear. Free HA in solution does not appear to be a good lubricant, being negatively charged and therefore repelled from most biological, including cartilage, surfaces. Recent enzymatic experiments suggested that mechanically or physically (rather than chemically) trapped HA could function as an “adaptive” or “emergency” boundary lubricant to eliminate wear damage in shearing cartilage surfaces. In this work, HA was chemically grafted to a layer of self-assembled amino-propyl-triethoxy-silane (APTES) on mica and then cross-linked. The boundary lubrication behavior of APTES and of chemically grafted and cross-linked HA in both electrolyte and lipid 1,2-dioleoyl-sn-glycero-3-phosphocholine (DOPC) solutions was tested with a surface forces apparatus (SFA). Despite the high coefficient of friction (COF) of μ ≈ 0.50, the chemically grafted HA gel significantly improved the lubrication behavior of HA, particularly the wear resistance, in comparison to free HA. Adding more DOPC lipid to the solution did not improve the lubrication of the chemically grafted and cross-linked HA layer. Damage of the underlying mica surface became visible at higher loads (pressure >2 MPa) after prolonged sliding times. It has generally been assumed that damage caused by or during sliding, also known as “abrasive friction”, which is the main biomedical/clinical/morphological manifestation of arthritis, is due to a high friction force and, therefore, a large COF, and that to prevent surface damage or wear (abrasion) one should therefore aim to reduce the COF, which has been the traditional focus of basic research in biolubrication, particularly in cartilage and joint lubrication. Here we combine our results with previous ones on grafted and cross-linked HA on lipid bilayers, and lubricin-mediated lubrication, and conclude that for cartilage surfaces, a high COF can be associated with good wear protection, while a low COF can have poor wear resistance. Both of these properties depend on how the lubricating molecules are attached to and organized at the surfaces, as well as the structure and mechanical, viscoelastic, elastic, and physical properties of the surfaces, but the two phenomena are not directly or simply related. We also conclude that to provide both the low COF and good wear protection of joints under physiological conditions, some or all of the four major components of jointsHA, lubricin, lipids, and the cartilage fibrilsmust act synergistically in ways (physisorbed, chemisorbed, grafted and/or cross-linked) that are still to be determined.

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INTRODUCTION The articular joint is an excellent water-based lubrication system. In synovial fluid (SF), healthy cartilage surfaces slide on each other with extremely low friction coefficients of 0.0005− 0.04 (at least initially; see below), and showing little wear over the lifetime of a person.1 The top layer of the articular joint is the cartilage tissue. Water is the major component of the cartilage tissue, comprising about 70−80% of the cartilage by weight. In the remaining “solid fraction”, type II collagen is the most abundant component, constituting about 50−80% of the dry weight. The remaining components are hyaluronic acid (HA), glycosolated proteins (known as proteoglycans or lubricin), and various lipids, among which phosphatidylcholines (PCs) are the most common.2,3 The function of the cartilage is highly dependent on the porous (open network) structure of the collagen matrix, which leads to a lubrication process by the pressurization of the interstitial fluid of cartilage.4,5 The friction force varies nonlinearly with the load, generally increasing with the load, but the coefficient of friction (μ or COF, defined as μ = dF∥/ © 2011 American Chemical Society

dF⊥, where F∥ is the friction force and F⊥ is the normal load) can increase or decrease.6 When there is a finite adhesion force between two sliding surfaces, finite friction forces can be measured even at zero or negative loads, and the COF can be very large. As the load increases, the “load-dependent friction” dominates over the “adhesion-dependent friction”,7 resulting in a decreasing COF.8 Different surface textures and/or types of surface damage during sliding can also lead to an increase or a decrease of the COF with increasing load.9,10 Time also plays an important role in the friction of cartilage surfaces (as it does in other systems as well).8 Recent measurements of the COF of cartilage have shown that, although usually starting from very low values, typically μ = 0.0005−0.04, the COF steadily increased with the sliding/ Special Issue: Bioinspired Assemblies and Interfaces Received: September 30, 2011 Revised: November 28, 2011 Published: December 7, 2011 2244

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greater than Rg, RF, or the brush layer thickness, i.e., significantly greater than molecular dimensions.8,18 HA on its own has been considered to be a potential boundary lubricant in cartilage;19 however, whether it functions as a sole cartilage lubricant (rather than working together or synergistically with other molecules) is still not clear.20 Various experiments including selective enzymatic digestion of SF have demonstrated that HA, acting alone, does not have any effect on the BL of cartilage.17,21 Experiments on cartilage sliding against cartilage also showed that just adding HA-containing SF as the lubricant does not make a significant difference to the COF.12 Similarly, surface forces apparatus (SFA) tests on free HA in solution between two mica surfaces showed that free HA has little or no effect on the (high) COF and poor wear resistance due to the electrostatic forcing out (“depletion”) of HA from between the two negatively charged surfaces when they are brought together, i.e., under compression.17 By contrast, similar SFA friction experiments showed that when HA is chemically grafted to a lipid bilayer and cross-linked, the bound HA layer provides excellent wear protection even at pressures up to 200 atm, although the COF was high (0.15− 0.3),17,22,23 which, interestingly, is close to the value of natural cartilage after prolonged shearing at high loads.6,7 The HAgrafted bilayer experiments indicate that HA has poor lubrication properties when judged by the COF, but may play a crucial role as an antiwear agent, effectively preventing damage to the cartilage surface. However, in the HA-grafted bilayer experiments, HA was grafted onto a physisorbed bilayer on mica,17 which could be very different from immobilizing HA on a cartilage surface.6 Recent SFA studies on digested cartilage demonstrate that nature has developed an “adaptive” mechanically controlled effective “boundary lubrication” mechanism.6 Starting from the zero load, when the two surfaces are well separated, the first lubrication mechanism in a joint (as it is for all fluid-lubricated surfaces) is the fluid-pressurization driven EHDL,11 which, due to the high viscosity of high MW HA solutions, prevents the two cartilage surfaces from coming into “contact”24 for up to 30 min6,11 even when the loading (lithostatic) pressure pushing them together is high.25 Under severe conditions such as under high loading pressures or loading durations, the cartilage undergoes severe deformation (strains of up to 70%),26 which squeezes the pore and interfacial fluid out from the spongy cartilage interior as well as radially away from the gap between the surfaces. This loss of interfacial fluid causes the failure of EHDL, and the system enters the BL regime. The severe deformation of cartilage mechanically traps (effectively binds) parts of the “free” HA molecules in the collagen network, while the parts exposed at the surface into the solution remain free, where they behave like a highly flexible polymer brush layer, or effective boundary lubricant layer,27 that lubricates the surfaces and protects them from becoming damaged. Lipids are abundant in the cartilage interfacial layer and SFs. The superficial zone is coated with a thick layer of lipids among which PCs are predominant.28 Recently, surface active phospholipids in the articular joint have been suggested as possibly playing a major role in the BL of joints.29 Some studies have shown that adding lipids reduces the COF of cartilage sliding on glass. Other studies have suggested that phospholipids can function as effective boundary lubricants, whereas still others have proposed a synergistic effect of HA with phospholipids.29−31 Yet, despite the various studies on the

shearing time under a constant load and sliding speed. For example, μ increased from 0.01 to 0.24 for cartilage sliding against glass over a sliding time of 75 min,11 and a similar trend was observed for cartilage sliding against cartilage, with μ increasing to 0.2−0.3 after increased the dwell time to 300 s.12 The significant increase in the COF suggests that under prolonged (>30 min) loading and/or sliding, the lubrication mechanism of cartilage undergoes a transition from hydrodynamic/elastohydrodynamic lubrication (EHDL)13 (also referred to as or related to pressurization lubrication, viscous lubrication, and weeping lubrication) to boundary lubrication (BL).6 HA is a long (1−6 MDa) linear, negatively charged biological polyelectrolyte (Figure 1a). It is abundant in cartilage and SF,

Figure 1. (a) The chemical structure of HA (MW = 1.6 MDa). (b) Schematic of chemically grafted and cross-linked HA on amino-propyltriethoxy-silane (APTES)-coated mica surfaces. (c) The normal forces F⊥ measured between two APTES-coated mica surfaces and between two HA physisorbed and grafted (chemisorbed) surfaces as functions of the surface separation distance, D.

significantly increases the viscosity of SF, and may play important roles in cartilage’s lubrication processes, including EHDL and BL. HA complexes (physically cross-links) with the protein lubricin (LUB) to form hydrogel aggregates in solution or on the cartilage surface to which LUB is often bound.14 Such gel-like layers are believed to be extremely important for both EHDL and BL processes.15 The role of brush polymer-like polyelectrolytes, such as HA and LUB, in the BL of joints was first proposed by McCutchen,4 and has recently drawn the attention of many studies.16−18 When two polymer- or polyelectrolyte-grafted or covered surfaces approach each other, they usually experience a repulsive “entropic” force once the outer segments begin to overlap. This interaction, often referred to as “steric” repulsion, is due to the unfavorable entropy associated with compressing the chains between surfaces, and can keep two surfaces well separated at distances 2245

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(friction or shear) force F∥ between two grafted HA layers was measured as a function of the load (normal force, F⊥), shearing velocity, and shearing distance in electrolyte and DOPC solutions. The radius R of the cylindrically curved mica surfaces was 1−2 cm (Figure 1b).

effects of phospholipids on cartilage, no clear picture has emerged. HA, LUB and other glycosolated proteins, and lipids are the three families of molecules that are considered to be the most important boundary lubricants in articular joints. However, most studies have focused on the BL property of only individual components.17,18,32 The cartilage surface is a complex multicomponent surface with anisotropic hierarchical structures. It is quite possible that some or all of its three classes of components (protein, polysaccharide, and lipid) function synergistically to provide the excellent tribological behavior found in joints. In this paper, we present the results of SFA friction-force measurements focusing on the BL behavior of chemically grafted and cross-linked HA on mica surfaces in electrolyte solutions, without and with added 1,2-dioleoyl-snglycero-3-phosphocholine (DOPC), one of the most abundant PCs in the SF and on the cartilage surface.33

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RESULTS Normal and Shear Forces of Free and Grafted HA Layers. The measured normal forces, F⊥(D), were purely repulsive (no adhesion; Figure 1c), and exhibited some hysteresis on separation, which is typical of most high MW polymer-mediated interactions. Upon the approach of two APTES grafted surfaces (no HA; Figure 1c, left curve, open circles), a weak repulsion was first detected at D ≈ 65 nm. The repulsive force increased sharply below about 30 nm, reaching a steric “hard wall” repulsion at D ≈ 20 nm (pressure P ≈ 0.1 MPa), corresponding to two compressed “hydrated” APTES layers, and indicating that the thickness of one (compressed) APTES layer on mica in PBS buffer was about 10 nm, with some of the polymer chains extending out as far as 30 nm into the solution. These thicknesses are much larger than that of a monolayer of APTES, a small saline molecule (MW = 221.37) with a molecular size of smaller than 2 nm, indicating that the APTES molecules were highly polymerized on the mica surfaces. The normal forces between two grafted (chemisorbed) HA layers on APTES in PBS buffer (Figure 1c, right curves) were repulsive both on approach and separation. The range of the repulsion varied in different experiments and different contact locations, and ranged from 100 to 300 nm, which was always of much longer range than that between two APTES layers, indicating the successful grafting of HA on the APTES layers. The right curves in Figure 1c show a typical force distance profile on approach and separation. On approach (“In” curves), the repulsive force started at D ≈ 120 nm and sharply increased to a hard wall at D ≈ 85 nm (compared to 65 and 20 nm, respectively, for two APTES layers). A purely repulsive force was also measured on separation of the surfaces (“Out” curves), very similar to the approach curves, indicative of a roughly reversible interaction, as expected for grafted polymer layers that cannot be squeezed out from between the two surfaces even under high loading conditions. By contrast, the forces between two physisorbed HA monolayers (Figure 1c, middle curves) were also repulsive and approximately reversible, but extended only from ∼40 to ∼60 nm, significantly less than for the grafted HA layers. The measured hard wall distances of 40− 100 nm are close to one or two radii of gyration (Rg ∼ 50 nm) of the HA molecules used in the experiments, suggesting that the interactions between two (physisorbed or chemisorbed) HA monolayers on each surface obey (at least semiquantitatively) the expectations from theories of polymermediated interactions.8 The friction forces F∥ were measured as a function of the normal force (load) F⊥ for both the physisorbed and grafted HA layers on APTES. Physisorbed HA layers showed very poor lubrication properties (Figure 2a): F∥ rapidly increased with increasing load, giving a COF greater than 1. This high F∥ could be due to the entanglement of the HA chains during shearing. Since the HA molecules only physically attach to the APTES layer through electrostatic interaction, the HA chains can be dragged back and forth in the process of shearing of the surfaces, always remaining in the entangled state and giving rise to the high F∥. Another likely contribution to the high F∥ is the

MATERIALS AND METHODS

APTES grafting. Mica surfaces were first modified with APTES (Sigma-Aldrich) using chemical vapor deposition (CVD).34,35 Prior to CVD, thin mica sheets of uniform thickness 2−5 μm were glued onto two cylindrical silica disks. The surfaces were then treated with water− argon plasma for 5 min at 40 W. The plasma-activated mica surfaces were transferred into the CVD chamber and stored under vacuum. APTES vapor was pumped into the CVD chamber and allowed to react with and be deposited on the activated mica surfaces for 2 h, ideally resulting in a monolayer. During the reaction, a low pressure was maintained to minimize the condensation of microscopic droplets of APTES on the surfaces. Following the deposition, covalent APTES grafting was done by annealing the surfaces in a vacuum oven for 30 min at 80 °C. HA Grafting. The APTES-grafted mica surfaces were then immersed in a 3 mg/mL HA (average MW = 1.6 MDa, SigmaAldrich) solution for 3 h. The right amounts of N-(3-dimethylaminopropyl)-N′-ethylcarbodiimide hydrochloride (EDC) and N-hydroxysulfosuccinimide (NHS) (Sigma-Aldrich) were added into the HA solution to bring the EDC and NHS concentrations to 1 M for each component (50/50 EDC/NHS mixture). EDC/NHS can chemically graft HA onto APTES as well as cross-link the grafted HA layer, forming a gel-like HA layer (Figure 1b).17 The surfaces were then rinsed thoroughly using phosphate buffered saline (PBS) buffer and kept in the buffer solution before they were used in the SFA experiments. All glassware was cleaned in ethanol and then rinsed thoroughly with Milli-Q water and PBS buffer. Milli-Q water was used in all the glassware cleaning and buffer preparations. Preparation of DOPC Solution. DOPC (Avanti) solutions were prepared by dissolving DOPC in chloroform (Sigma-Aldrich) at a concentration of 1 mg/mL. The chloroform was evaporated by blowing filtered nitrogen gas over the solution. DOPC was then hydrolyzed at a concentration of 1 mg/mL in PBS buffer at pH 7.4 and 22 °C. The solution was then sonicated for 4 h to dissolve (break up) the DOPC lamellar phase particles formed during the nitrogen bubbling. Prior to experiments, the solution was filtered (pore size ∼0.2 μm) to remove any large particles (vesicles or liposomes) that were formed during this process. The SFA. The SFA has proven to be an excellent technique for studying the tribological properties of varies materials including polymers, nanoparticles, lipid monolayers, bilayers, and biological macromolecules.17,36,37 An SFA 2000 was used to measure the normal and lateral (or shear) forces between the two curved cylindrical surfaces as a function of the separation distance D (gap width or water film thickness) between the HA-grafted mica surfaces (Figure 1b).38,39 The SFA is capable of measuring the distance between the HA-grafted mica surfaces D at the angstrom level and the normal force F⊥ with an accuracy of 10 nN. The bottom surface, mounted to a bimorph sliding attachment, was sheared back and forth laterally with a travel distance of about 80−100 μm at various shearing velocities, and the lateral 2246

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the tensile pressures associated with shearing (stretching) the entangled HA chains. Our shearing distance (∼80 μm) was much longer than the contour length of the HA molecules (∼1 μm) that were chemically bonded to the APTES layers. The shearing could also drag the HA chains out of grafting spots. Therefore, mechanical failure could occur either between the HA and the APTES layers or between the APTES layers and the mica substrates, resulting in surface damage in either case. Normal and Shear Forces of DOPC Lipid Multilayers. The lubrication properties of two mica surfaces in a pure DOPC solution were also measured. Fifty microliters of the DOPC solution was injected between the mica surfaces after measuring the bare mica−mica contact. After allowing adsorption for 1 h, the two surfaces were brought together, and the normal forces F⊥ between the surfaces were recorded as a function of distance (Figure 3a). The interaction between the

Figure 2. Measured friction forces F∥ between two HA grafted mica surfaces in PBS buffer as a function of (a) the normal force (load) F⊥ at a constant sliding speed of V = 10 μm, and (b) sliding speed V at a constant load of F⊥ = 1 mN. Inset of panel a shows that LUB has a low COF but poor wear resistance, as reported in ref 18.

bridging adhesion between the HA chains and the opposing surfaces. In PBS buffer, the HA molecules are negatively charged, and the APTES layer is positively charged, therefore leading to an adhesive electrostatic interaction between the APTES layers. Additionally, the physisorbed HA layers showed no wear resistance: damage of the mica surfaces was observed already at a fairly low load of ∼1 mN (pressure, P < 0.4 MPa). The friction forces F∥ between chemically grafted HA layers were significantly lower than those between physisorbed layers (Figure 2a), indicating that grafting did improve the lubrication properties of HA. F∥ increased linearly with the applied load with a (fairly constant) COF of μ = 0.52, characteristic of pure “load-controlled” friction. At a load of 5.6 mN (pressure ∼2 MPa), damage of the mica surface was observed after sliding some tens of cycles, that is, not immediately, but after sliding over a distance of 2000 to 4000 μm. Thus, although chemically grafted HA still had a considerably higher COF than that of cartilage, it showed a much improved lubrication performance in comparison to physisorbed HA, probably due to the shearinduced alignment (shear thinning) of the chains and lack of any adhesion contribution to the friction force. However, it is worth noting that the excellent lubrication of cartilage occurs only initially and is due to EHDL and the porous structure of the collagen network.11,40 Under prolonged loading and/or sliding, the friction coefficient of the cartilage increases up to as high as 0.2 to 0.5,1,6,11,12,41 which is much closer to the value of the friction coefficient of the grafted HA measured in our studies. More importantly, chemical grafting significantly increased the wear resistance of the HA layer: grafted HA can be sheared under a 5 times higher pressure (up to 2 MPa) without noticeable surface damage. The effect of sliding speed V on the friction force between two grafted HA layers was also investigated under a constant load of 1 mN and a sliding distance of ∼80 μm over a speed range of 2 decades, from V = 0.18 μm/s to V = 36.4 μm/s (Figure 2b). The friction force showed only a weak dependence on the sliding speed. At low V, the friction force increased logarithmically with V. Damage to the mica surface occurred at a sliding speed of V = 10 μm/s, and the friction force significantly increased after the surfaces became damaged. For both high loads and high sliding speeds, surface damage was observed at the end of the test. Surface damage was first noticed as discontinuities in the FECO fringes during the experiment. Through an optical microscope, wear tracks were seen to occur uniformly over the contact area along the sliding direction. The damage of the mica surfaces is probably due to

Figure 3. (a) Measured normal force F⊥ as a function of separation distance D between two mica surfaces in 1 mg/mL DOPC solution. a, e, and f are the corresponding F⊥/R values for the same points in panel b. (b) Friction force F∥ of DOPC on mica as a function of normal load F⊥. (c) F∥ of DOPC on mica vs sliding speed V at a constant load of F⊥ = 1.6 mN (regime a in panel b).

mica surfaces in the DOPC solution was purely repulsive with a range of more than 80 nm. The repulsive force increased rapidly at D ≈ 50 nm and increased roughly exponentially with distance. The strong repulsion starting at a fairly large distance and the thick hard wall at D ∼ 40 nm indicate that the DOPC lipids have formed multilayers on the mica surfaces under these experimental conditions. 2247

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Friction measurements of DOPC on mica were conducted with the same geometry. A fairly high COF, μ = 0.71, was measured at a constant sliding velocity of V = 5 μm/s (Figure 3b). Damage of the mica surfaces was observed at high F⊥ (∼7 mN), and the friction force F∥ immediately increased after the damage occurred. The velocity dependence of the friction of DOPC on mica was also studied at a constant load of F⊥= 1.6 mN (Figure 3c), corresponding to point a in Figure 3b. The shearing (sliding) velocity was varied by nearly 2 orders of magnitude, from V = 0.02 to V = 10 μm/s. F∥ increased logarithmically with V until damage occurred at a sliding speed of V = 3.5 μm/s. Our results suggest that DOPC does not have good BL properties on mica surfaces. This is very surprising and contradicts many studies showing that lipids and liposomes might function as major boundary lubricants.31,42 In our system the strong and steeply repulsive interaction measured between the mica surfaces, along with a “stiff” hard wall even under very high compression, indicate an ordered “solid-like” state. This “solid-like” state can lead to the high friction force measured, and could explain the inconsistency between other studies and our results. In our study, only DOPC was tested as the model lipid, whereas there are more than 20 different lipids in the cartilage system.33 These lipids, with varied chain lengths and head groups, could impart much better lubrication properties; and it is also likely that the lipids interact synergistically, for example, with LUB or HA, when they participate in the BL of cartilage surfaces under physiological conditions.17,18 The SFA results suggest that DOPC, acting alone, does not function as a good boundary lubricant, at least in the highly confined geometry between two molecularly smooth mica surfaces. The Lubrication Behavior of Grafted HA and DOPC Mixture. To test the synergistic effect of the grafted HA and DOPC, the normal interaction between two chemically grafted HA layers was first measured in PBS buffer. Fifty microliters of DOPC solution (1 mg/mL) was then injected into the gap solution between two grafted HA layers when the two surfaces were well separated by a few micrometers. After allowing adsorption for 1 h, the surfaces were then brought into contact without changing the contact position from the previous measurement. The injection of the DOPC solution had a strong effect on the normal interaction force between the HA layers. The range of the repulsive force changed dramatically after the injection, increasing from 150 to 280 nm (Figure 4a). The hard wall of the two HA gel layers was about 130 nm

before the DOPC injection, and moved out to 250 nm after the injection. It has been shown that DOPC molecules can interact with HA molecules, forming a HA−lipid complex.43,44 The shift of hard wall in our SFA measurements again showed that there is very strong adsorption of DOPC on the grafted HA layer. This adsorption can cause swelling of the HA layer and increase the range of repulsion. Injecting the DOPC solution, however, did not significantly change the BL properties of the grafted HA layers. F∥ increased linearly with F⊥ with a COF of μ = 0.66, a value that is fairly close to that of grafted HA layers measured in PBS buffer (Figure 4b). Damage of the mica surface was observed after sliding a few cycles at F⊥ = 4 mN, and the friction force F∥ again increased significantly at the same F⊥ after the damage occurred. The similar COFs measured between grafted HA layers in PBS buffer and in DOPC solution suggest that, although there are probably some interactions between HA and DOPC, the two components together do not function as effective boundary lubricants on the mica surface.

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DISCUSSION Our experiments show that covalently grafted HA does not have low friction forces in electrolytes and DOPC solutions. This is consistent with various other studies.17,23 In cartilage, HA is not the primary boundary lubricant; however, it functions as an EHDL and an “emergency” boundary lubricant, and perhaps more importantly, acts as an effective wear protector of cartilage surfaces.6 The fact that covalently grafted HA showed much improved wear protection in comparison to adsorbed “free” HA also points to a wear protection role for HA. The Tribological Effects of Chemically Grafted HA on Mica. Free HA has been shown to not have good BL properties for mica surfaces, partially due to the repulsive electrostatic interactions between HA and mica.23 Our measurements on chemically grafted HA on mica demonstrate that grafted HA has a much better performance, both by lowering the COF and providing much improved wear resistance. The high F∥ and final damage of the surface are probably due to the entanglements of the shearing HA chains under high compression. Under low compression, the entropic steric repulsion between the polymer brushes gives rise to a strong entropic repulsion that keeps two surfaces well separated; however, higher compression leads to entanglement and/or interdigitation of the brushes, which significantly increases the viscosity of the film and possibly causes some bridging adhesion, leading to higher friction forces.16,42 The damage to the mica surfaces is likely due to these entanglements and bridging adhesion. High local pulling forces caused by the polymer entanglements and adhesion have previously been found to induce delamination and cavitation damage of mica surfaces.45,46 Comparison and Relevance of Results to Physiological Conditions. One important difference between our SFA experiments and natural joint lubrication processes is that HA is not chemically grafted to the cartilage surface. Recent SFA enzymatic digestion experiments have suggested that the role of high MW HA in cartilage lubrication is “adaptive”: HA molecules on the cartilage surface partially entangle with the collagen fibrils network and partially extend to the interfacial fluid. Under normal conditions, HA is free to move, although HA also weakly interacts with LUB, forming a weakly entangled or physically cross-linked gel.14 Under low loads or pressures, or shear velocities V, this gel layer acts as an effective boundary

Figure 4. (a) The normal forces F⊥ of pure DOPC on mica (red), grafted HA on mica (black), and grafted HA with DOPC (blue). (b) Friction force F∥ of grafted with injected DOPC solution in the gap. High F∥ were measured with a COF of μ = 0.66 before the mica surfaces got damaged. After surface damage, an increase in F∥ at the same F⊥ was observed. 2248

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lubricant and keeps two cartilage surfaces apart. As V increases, the system transits into the EHDL regime in which hydrodynamic and EHD effects play the most important role.40 A second transition takes place at more severe conditions, such as under high pressures or prolonged compression times. The interfacial fluid gets depleted due to the large deformation of the cartilage, leading to the breakdown of the EHDL. The system translates from EHDL to the second BL regime in which the deformation of the collagen decreases the lateral pore size of the cartilage collagen matrix while preserving the pore size in the normal direction.26 In this anisotropic deformation regime, the cartilage fibrils are able to mechanically trap the HA in the collagen matrix, especially the HA in the superficial zone. Although the weak physical interaction between LUB and HA fails to maintain the gel network due to the high pressure and shear force,14 the “trapped” HA can function as the “emergency” boundary lubricant. In this regime, HA serves mainly as the last mechanism of defense against the surface damage and wear. Our grafting HA on mica surfaces mimics this mechanical “trapping” process in the cartilage. Although the chemically grafted HA in our study is not equivalent to the trapped HA, it nevertheless captures some key properties of the mechanically immobilized HA. Our measurements demonstrate that HA fulfills this “emergency” lubrication role fairly well; and despite not having a low friction coefficient but still almost as low as natural joints after prolonged shearing, the grafted HA showed much better protection of the mica surface from damage under significantly higher loads compared to the physisorbed “free” HA. The cartilage in the joint operates over a great range of pressure, sliding speed, and shear distance. Our SFA experiments were done on mica surfaces, not on cartilage, and covered a smaller range of operating conditions of the joint, especially the sliding velocity and sliding distance. Our SFA tests focus on the BL behavior of grafted HA under low sliding speeds and shear distances, which is the regime in which BL plays the main role. Moreover, there are many physiological situations in which normal and shear stress are applied slowly or quasi-statically, for example, when a person is sitting, standing, or sleeping, to which our results have significant physiological relevance. In many cases, slow and quasi-static motions often cause great pain to arthritis patients, such as the pain and stiffness people feel when they first wake up in the morning.47 We have successfully developed a new SFA setup, which allows us to test the BL behaviors of grafted HA and other molecules, such as LUB and lipids, with a wider range of sliding speed and distance.48 The Relationship between COF and Wear. It is worth noting that high COFs have not been directly or quantitatively related to pain in arthritic joints, whereas damage that has occurred within the joints has (along with prolonged static loading, e.g., after first getting out of bed in the morning, and/ or after prolonged use of a joint). When friction is mentioned as the cause of arthritic or joint pain, it is almost always associated with “abrasive friction”, i.e., the friction that “degenerates”, “damages”, “wears away” or “destroys” the cartilage. Accordingly, to prevent surface damage (wear, degeneration, abrasion, destruction, etc.), which is generally accepted to be the main morphological manifestation of arthritis, the traditional wisdom has been that one should aim to reduce the COF.47

Indeed, there does not appear to be any systematic study relating COF to either wear or osteoarthritic or joint pain. Some studies indicated that adding liposome additives reduced the wear rate of cartilage, but the correlation between the friction force or COF and wear was not investigated.49 By chemically cross-linking various components found in joints, including HA, either to each other or to the cartilage surfaces, we have found that a high COF can be associated with good wear protection, while a low COF can have poor wear resistance, as also demonstrated in the LUB test.18 Both of these properties depend on how the molecules are attached to and organized at the surfaces (i.e., in the boundary lubricant layer), but they are not directly or simply related. The various molecules on the cartilage surface can be physisorbed or chemisorbed, cross-linked, or un-cross-linked, leading to these four different combinations of the ways in which the molecules can be organized on the surface. Depending on different situations or requirements, each of the combinations can be useful in terms of reducing the friction force and/or providing wear resistance. This work, together with our recent cartilage digestion study,6 suggests that the primary function of a healthy cartilage and joint lubrication system may be more to prevent wear than lower the friction coefficient. Interestingly, experiments on cartilage sliding against various materials also showed that after some initial running in, the COF rises to steady state values of 0.2−0.5,1,6,11,41 rather than remaining below 0.01. This “high” value is close to what we measure in our crosslinking and trapped HA experiments. The Interaction between HA and Lipids. Lipids have long been considered as candidates for the boundary lubricants of cartilage. In the cartilage superficial zone, various lipids form a gel-like layer together with collagen fibrils, with HA and LUB. Our study indicates that DOPC could interact with HA and cause a swelling of the grafted HA layer, which increases the thickness of the HA gel. This swelling is also likely to happen for the partially entangled and weakly LUB-cross-linked-HA on the cartilage surface. The increased thickness of the HA layer could help keep the collagen fibrils network of two cartilage surfaces well separated (at the molecular level), and therefore reduce the wear of the cartilage during shearing. The lubrication properties of lipids depend highly on the chemical structure of the lipid molecules. Although the DOPC used here as a model lipid did not show very good BL behaviors on mica surfaces, it may work synergistically with HA and LUB, functioning as a very good boundary lubricant.

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CONCLUSIONS Articular cartilage has a synergistic lubrication mechanism that provides excellent lubrication properties and wear resistance.6 Although EHDL is the main lubrication mechanism for cartilage under low loads and short loading and shearing times, BL is extremely important for the proper function of cartilage under severe conditions, i.e., at high loads and/or prolonged loading and shearing times. Chemically grafted HA on mica surfaces mimics the mechanically trapped HA in the cartilage collagen matrix, and behaves very differently from free HA in the solution or SF. Our normal and friction force measurements indicate that chemically grafted HA could provide strong repulsive interactions between two grafted HA layers and a significant improvement of wear resistance in comparison to free HA, although with a still fairly high COF. This is consistent with the role of HA as an “emergency” boundary lubricant in the “adaptive” joint lubrication model.6 2249

dx.doi.org/10.1021/la203851w | Langmuir 2012, 28, 2244−2250

Langmuir

Article

buoyancy or gravitational displacement force), and is called the lithostatic pressure, which we here refer to as the loading pressure. This is the pressure that gives rise to elastohydrodynamic deformations of the cartilage and the flow patterns of the fluid out of the cartilage interior, through the thinning water gap, and into the synovium (the SF “reservoir”). These dynamic pressures and processes are quite different from the hydrostatic pressure that is uniformly isotropic throughout the fluid phase under static conditions, i.e., when there is no motion of the surfaces or flow of fluid. (26) Greene, G. W.; Zappone, B.; Soderman, O.; Topgaard, D.; Rata, G.; Zeng, H. B.; Israelachvili, J. N. Biomaterials 2010, 31 (12), 3117− 3128. (27) It should be noted that this “layer” is highly “hydrated”, typically composed of more than 90% water. (28) Hills, B. A. Proc. Inst. Mech. Eng., Part H 2000, 214 (H1), 83− 94. (29) Nitzan, D. W.; Nitzan, U.; Dan, P.; Yedgar, S. Rheumatology 2001, 40 (3), 336−340. (30) Forsey, R. W.; Fisher, J.; Thompson, J.; Stone, M. H.; Bell, C.; Ingham, E. Biomaterials 2006, 27 (26), 4581−4590. (31) Sivan, S.; Schroeder, A.; Verberne, G.; Merkher, Y.; Diminsky, D.; Priev, A.; Maroudas, A.; Halperin, G.; Nitzan, D.; Etsion, I.; Barenholz, Y. Langmuir 2010, 26 (2), 1107−1116. (32) Hills, B. A.; Crawford, R. W. J. Arthroplasty 2003, 18 (4), 499− 505. (33) Sarma, A. V.; Powell, G. L.; LaBerge, M. J. Orthop. Res. 2001, 19 (4), 671−676. (34) Banquy, X.; Zhu, X. X.; Giasson, S. J. Phys. Chem. B 2008, 112 (39), 12208−12216. (35) Liberelle, B.; Giasson, S. Langmuir 2007, 23 (18), 9263−9270. (36) Maeda, N.; Chen, N. H.; Tirrell, M.; Israelachvili, J. N. Science 2002, 297 (5580), 379−382. (37) Israelachvili, J. N.; Mcguiggan, P. M.; Homola, A. M. Science 1988, 240 (4849), 189−191. (38) Israelachvili, J. N J. Colloid Interface Sci. 1973, 44 (2), 259−272. (39) Israelachvili, J.; Min, Y.; Akbulut, M.; Alig, A.; Carver, G.; Greene, W.; Kristiansen, K.; Meyer, E.; Pesika, N.; Rosenberg, K.; Zeng, H. Rep. Prog. Phys. 2010, 73 (3), 036601. (40) Caligaris, M.; Ateshian, G. A. Osteoarthritis Cartilage 2008, 16 (10), 1220−1227. (41) Forster, H.; Fisher, J. Proc. Inst. Mech. Eng., Part H 1996, 210 (2), 109−119. (42) Goldberg, R.; Schroeder, A.; Barenholz, Y.; Klein, J. Biophys. J. 2011, 100 (10), 2403−2411. (43) Crescenzi, V.; Taglienti, A.; Pasquali-Ronchetti, I. Colloids Surf., A 2004, 245 (1−3), 133−135. (44) Crockett, R.; Grubelnik, A.; Roos, S.; Dora, C.; Born, W.; Troxler, H. J. Biomed. Mater. Res. A 2007, 82A (4), 958−964. (45) Kuhl, T.; Rutgs, M.; Chen, Y. L.; Israelachvili, J. N. J. Heart Valve Dis. 1994, 3 (Suppl. I), S117−S127. (46) Chen, Y. L.; Israelachvili, J. Science 1991, 252 (5009), 1157− 1160. (47) MedlinePlus [Internet]. Bethesda (MD): National Library of Medicine (US); Available from: http://www.nlm.nih.gov/ medlineplus/arthritis.html. (48) Lowrey, D. D.; Tasaka, K.; Kindt, J. H.; Banquy, X.; Belman, N.; Min, Y.; Pesika, N. S.; Mordukhovich, G.; Israelachvili, J. N. Tribol. Lett. 2011, 42, 117. (49) Verberne, G.; Schroeder, A.; Halperin, G.; Barenholz, Y.; Etsion, I. Wear 2010, 268, 1037−1042.

The significant swelling of the grafted HA layer in DOPC solution further indicates that there may be some interactions between HA and lipids; however, these interactions did not improve the BL properties of HA in our SFA tests. Together with LUB, HA and lipids can form a gel-like structure on the cartilage surface, which can be crucial in the EHDL and BL processes in joints. Our results suggest that none of the four major components of jointsHA, LUB, lipids, and collagen fibrilsact alone as the “magic bullet” of lubrication, either in EHDL or BL, and that all or some of these molecules must work synergistically under physiological conditions. This synergy requires further investigation.

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ACKNOWLEDGMENTS This work was funded by the McCutchen Foundation. J.Y. acknowledges the CSP Technologies fellowship from the Materials Research Laboratory, University of California, Santa Barbara.

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dx.doi.org/10.1021/la203851w | Langmuir 2012, 28, 2244−2250