Articular cartilage is composed primarily of water, collagen, and large proteoglycans (Fig. 3, left panel). By wet weight, cartilage is 68–85 % water, 10–20 % collagen, 5–10 % proteoglycan, and <5 % other matrix molecules [8]. The interstitial fluid contains dissolved electrolytes, predominantly Na⁺, Ca²⁺, Cl⁻, and K⁺. Chondrocytes account for less than 10 % of the total tissue volume [9] and are responsible for the metabolic activity of cartilage. The primary collagen in articular cartilage is fibril-forming type II collagen with variable orientation of collagen fibers depending on the depth of the tissue. In the superficial zone (top 10–20 %), fibers are oriented parallel to the articular surface; in the middle zone (middle 40–60 %), fibers are oriented randomly; and in the deep zone (bottom ~30 %), the fibers are oriented perpendicularly to the subchondral bone. Aggrecan accounts for 80–90 % of all proteoglycan in cartilage. Chondroitin sulfate, keratin sulfate, and hyaluronan are the primary glycosaminoglycan (GAG) side chains in cartilage. Chondroitin sulfate and keratin sulfate have negatively charged sulfate and carboxyl groups that make the chains anionic overall. Hyaluronan is not sulfated and interacts with aggrecan and link proteins to form large aggregates that are immobilized in the extracellular matrix and which consequently stabilize the extracellular matrix. Since these charged proteoglycans are immobilized within the extracellular matrix, the resulting charge is referred to as the “fixed charge density.” The proteoglycan distribution, and therefore the distribution of fixed charge density, varies through the cartilage depth. Aggrecan level is the lowest in the superficial zone and increases with depth [8].

The structure of articular cartilage produces complex material behavior (Fig. 3, right panel). Because cartilage structure and composition vary with depth, material properties also vary with depth [10–13]. The material behavior of cartilage varies between species within the same joint, between joints within species, and spatially within each joint within each species [14–16]. Cartilage exhibits nonlinear material behavior in both tension and compression. Under uniaxial tensile stress, cartilage material behavior is primarily determined by the collagen fibrils. The stress–strain curve in tension exhibits a toe region followed by an approximately linear region due to the uncrimping of collagen fibers followed by loading of straightened fibers [8]. Under uniaxial compressive stress, the material response of cartilage is governed by the proteoglycan matrix and fluid flow. The modulus of cartilage in tension is approximately one to two orders of magnitude lower than in compression, a discrepancy known as tension–compression nonlinearity [16]. This characteristic is important for most modes of cartilage deformation that are relevant to whole-joint mechanics.

Fluid–solid interactions, the intrinsic viscoelasticity of the solid phase, and fixed charge density causes flow-dependent viscoelasticity of the cartilage and swelling [17–19]. The cartilage swelling is caused by the fixed charge density of the tissue and charge–charge repulsion between closely packed GAGs attracting interstitial fluid counterions [8]. Collagen primarily bears tensile stress, and thus the fibrillar collagen resists expansion of the solid matrix during swelling [20].

Solutes, including nutrients and metabolic by-products, move through cartilage via diffusion. Solute diffusivity in cartilage is smaller than in aqueous solution [8], and diffusivity decreases as the tissue is compressed [21, 22]. The size of the solute also influences solute diffusivity. For large solutes, cyclic loading can enhance diffusion but it has no effect on small solutes [23].

Areas of contact on the articular surfaces vary somewhat with activity and motion during the activity in both normal and pathomorphologic hips. In the normal hips, the direction of loading and the location of chondral contact change from predominantly superior–posterior during ascending stairs, to more superior during walking, to superior–anterior during descending stairs [26]. Predicted peak stress in normal hips ranges from 7.52 ± 2.11 MPa for heel-strike during walking (2.3 times the body weight) to 8.66 ± 3.01 MPa for heel-strike during descending stairs (2.6 times the body weight). Across all activities of daily living, the contact area of the femur on the acetabular cartilage occupies about a third of the total surface area. Even in the normal hips, the distribution of contact stress is highly nonuniform due to local incongruities between the femoral and acetabular cartilage. In addition, contact stress varies more between different subjects for a single activity than between different activities for a single subject.

In the hips with acetabular dysplasia, the acetabulum is shallow. Dysplasia is typically diagnosed radiographically by an anterior and/or lateral center edge angle (CEA) less than 20–25° and an acetabular index greater than 10°. These measures are indicative of a shallow acetabulum and an upwardly sloping sourcil, respectively. Patient-specific finite element (FE) predictions of contact area show that only the superior region of the acetabulum exhibits significantly different labral contact areas when comparing normal and dysplastic subjects [25] (Fig. 4). This suggests that during activities of daily living, the superior labrum in the dysplastic hip is loaded more than other portions of the acetabular labrum. Predictions of labral load support corroborate this finding – the labrum in dysplastic hips supports loads that are 2.8–4.0 times larger than that of normal hips. These results are consistent with clinical observations of labral hypertrophy, as well as the superior or anterosuperior location of labral tears in the dysplastic hip [27–29].

In the hips with acetabular retroversion, the acetabulum opens more posterolaterally than normal. This is recognized on anteroposterior radiographs by the presence of a crossover sign, which indicates a prominent anterior acetabular wall, a deficient posterior acetabular wall, or both. Cartilage contact in retroverted hips is focused in the superomedial acetabulum, whereas in normal hips, contact patterns are distributed over more of the entire acetabulum (Fig. 4) [30]. Additionally, retroverted subjects tend to have patches of more medial contact.

Focal cartilage defects often result from cartilage delamination, typically due to shear loading that causes failure at the osteochondral interface. In vitro, both articular surface contact stress and maximum shear stress have been shown to predict cartilage fissuring under impact loads [31, 32]. The maximum shear stress during activity for the normal hip tends to occur at the interface of the acetabular cartilage and subchondral bone, near the junction of the cartilage with the labrum [33].

For a specific joint, chondral defects greater than a certain size tend to increase in size, generate an inflammatory response in the joint, and ultimately lead to OA, while other smaller defects will remain stable and not progress to disease [34]. Furthermore, a number of geometric factors contribute to the variability in clinical success treating osteochondral defects , as these factors often produce unfavorable conditions for the formation of new cartilage or the survival of implanted plugs, scaffolds, or cells [35]. The major geometric factors are defect size, joint curvature, joint congruence, cartilage thickness, and status of the rim of the defect. Increased joint curvature generally causes the rim of the defect to experience higher stresses and strains during joint loading and articulation. This is further exacerbated if the joint surfaces are incongruent. Thinner regions of articular cartilage have less ability to increase the local congruency during deformation since deformation is limited by the reduced thickness. And, finally, if the rim of the defect is well defined and intact, the defect will be more likely to remain stable and not progress in the joint, whether it is repaired or not.

As illustrated in the upper panels of Fig. 5 for a defect in the femoral condyle of the knee, increasing defect size exposes the rim of the defect to higher stresses and strain. Furthermore, in larger defects, healing neocartilage or implanted plugs or cells are more exposed to deformation during healing, resulting in a progressively worse outcome independent of other factors. Because the congruency of diarthrodial joints can vary as a function of joint kinematics, a defect may be subjected to different degrees of loading as a function of joint articulation. This is illustrated in the lower panels of Fig. 5 for the femoral condyle of the knee as a function of knee flexion.

For many patients with chondral defects, a trial of nonoperative measurement is appropriate. This often consists of activity modification or weight management to decrease the mechanical load across the damaged area of cartilage. Physical therapy focusing on hip and core strengthening may also help improve the muscular dynamics around the hip, particularly if there is a component of instability [40, 41]. Patients frequently ask about nutritional supplements. The most well known of these is a combination of glucosamine and chondroitin. Individual trials have shown mixed results, and in meta-analysis, there was no clinically relevant effect on joint pain or joint space narrowing [42].

Intra-articular cortisone injections may be used to reduce symptoms from inflammation. However, each injection has some risk of infection, soft tissue weakening, and possibly a microscopic effect on the cartilage [43]. Multiple cortisone injections should not be the definitive management strategy for young patients who are potential candidates for cartilage restoration procedures. Platelet-rich plasma (PRP) has also been touted as a possible treatment for cartilage defects or osteoarthritis. Advocates for intra-articular PRP injections promote it as a topical application of biological factors that promote healing of the defect. However, preparations of PRP vary and although the effective agent has been proposed, it is not definitively known [44]. There are two studies of PRP injection for hip arthritis; other studies of intra-articular PRP are for the knee or for the talus. Most patients have some initial improvement with gradual worsening over time [45]. Hyaluronic acid or viscosupplementation improves symptoms by a combination of antiinflammatory effects, restoration of the viscosity of synovial fluid, and normalization of synovial cell hyaluronate synthesis [46]. It appears to be safe, without potential adverse effects [46]. The results of hyaluronic acid injections are mixed. In the hip, 45–60 % of people have pain relief 6 months after injection [46].