**4.2 Compliant bearings**

There has long been interest in developing bearing materials that exhibit friction and wear behavior similar to that of articular cartilage. Cartilage is an example of a compliant bearing that has a low modulus but is capable of large deformation without failure. The friction coefficient between cartilage surfaces in a synovial joint is less than 0.01. This low friction is

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achieved via three lubrication mechanisms: elastohydrodynamic (EHD) lubrication, µEHD, and squeeze-film lubrication. Elastic deformation of the articular surfaces under load assists in allowing joint fluid to separate the surfaces, avoiding solid-solid contact with consequent wear. µEHD has a similar effect confined to the surface asperities (roughness) of the cartilage. EHD and µEHD predominate during the stance phase of walking, when pressure is generated in the synovial fluid by an entraining motion between the joint surfaces. Squeeze-film action predominates during heel strike, as the two cartilage surfaces move toward each other, squeezing the joint fluid from between the surfaces. Deformation of the articular cartilage assists in retention of the synovial fluid film. Healthy cartilage has a thickness of 2 to 4 mm at the femoral head and acetabulum. The modulus of elasticity is 10 to 50 MPa with a Poisson ratio of 0.42 to 0.47. The surface roughness is about 2 µm. Polyurethanes are synthetic polymers having properties comparable to those of articular cartilage. There have been extensive studies to determine the suitability of polyurethanes as bearing surfaces. The intent is to have a bearing in which the surfaces are separated by the pressure developed in the joint fluid as well as by the deformation of the articular surfaces, resulting in low friction and low wear. This is a different approach from that of using UHMWPE, CoCr, or alumina bearings that operate under mixed lubrication conditions with higher friction and a degree of solid to solid contact.(Scholes 2000)

### **5. References**


achieved via three lubrication mechanisms: elastohydrodynamic (EHD) lubrication, µEHD, and squeeze-film lubrication. Elastic deformation of the articular surfaces under load assists in allowing joint fluid to separate the surfaces, avoiding solid-solid contact with consequent wear. µEHD has a similar effect confined to the surface asperities (roughness) of the cartilage. EHD and µEHD predominate during the stance phase of walking, when pressure is generated in the synovial fluid by an entraining motion between the joint surfaces. Squeeze-film action predominates during heel strike, as the two cartilage surfaces move toward each other, squeezing the joint fluid from between the surfaces. Deformation of the articular cartilage assists in retention of the synovial fluid film. Healthy cartilage has a thickness of 2 to 4 mm at the femoral head and acetabulum. The modulus of elasticity is 10 to 50 MPa with a Poisson ratio of 0.42 to 0.47. The surface roughness is about 2 µm. Polyurethanes are synthetic polymers having properties comparable to those of articular cartilage. There have been extensive studies to determine the suitability of polyurethanes as bearing surfaces. The intent is to have a bearing in which the surfaces are separated by the pressure developed in the joint fluid as well as by the deformation of the articular surfaces, resulting in low friction and low wear. This is a different approach from that of using UHMWPE, CoCr, or alumina bearings that operate under mixed lubrication conditions with

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**11** 

*USA* 

*1Pioneer Surgical* 

*2SUNY Upstate, New York* 

**The Use of PEEK in Spine Arthroplasty** 

Cervical and lumbar disc arthroplasty are one component in the continuum of treatment for symptomatic degenerative disc disease (DDD) that is unresponsive to conservative care. In the lumbar spine, this may be accomplished via nucleus replacement or total disc replacement, and in the cervical spine by total disc replacement. The goal of both lumbar disc arthroplasty treatments are the same; relieving discogenic back pain through removing the pain source and restoring or maintaining motion segment function. For cervical arthroplasty, the goal is to relieve radicular pain as a result of nerve root compression, and/or myelopathy as a result of spinal cord compression, in addition to preserving motion. From a design standpoint, nucleus replacement technology consists of elastomers and nonelastomers, both preformed and *in-situ* cured, and can incorporate articulation similar to total disc replacements, with the intent of replicating to various extents the natural nucleus and preserving most of the annulus, thereby relying on a biomechanically intact annulus to share the compressive load. Also, most artificial nucleus devices are not fixed to the vertebral endplates, and therefore allow small, relative motion between their external surfaces and the vertebral endplates. In contrast, the majority of total disc replacement technology consists of articulating designs and material combinations that have been developed based upon the wealth of scientific and clinical information produced by the success of total joint arthroplasty. An artificial disc is designed to replace the entire disc tissue by excising almost all the disc materials, and therefore removing all the natural constraints in the anterior column. In addition, all artificial discs have a superior plate and an inferior plate, which are fixed to the two adjacent vertebrae. These represent key design

A key challenge for a disc arthroplasty device is selecting the proper material(s) for the various components that consitute its design. Unlike total joint replacement, a candidate for disc arthroplasty is on average 40 years of age with a target indication of 18 to 60 years (Zigler et al., 2007; Murrey et al., 2009). As a consequence, these devices are expected to last much longer than those of total joint recipients, whose average age is 70 years (Bergen, 2011, Garellick et al., 2010). Therefore, there are stringent requirements for long term implantable materials, and this will significantly limit the selection of materials available. Biocompatibility and biodurability, otherwise known as the abilities of a material to maintain its physical and chemcial integrity under *in vivo* applications without eliciting an aggressive host immune response for a given application, are essential for permanent

**1. Introduction** 

differences between the two technologies.

T. Brown1, Qi-Bin Bao1 and Hansen A. Yuan2

