Polymer-based composite hip prostheses
Introduction
Hip prosthesis (HP) designs for total hip replacement (THR) present a stem which fits into the medullary canal passing through the proximal part of the femur (epiphysis). Thus, a HP mainly replaces the cancellous bone, occupies part of the medullary canal and substitutes part of the femur head tissue which is discontinued in its proximal part according to the type of injury or fracture [1].
Finite element modelling (FEM) and follow-up on implanted metallic hip prostheses (MHPs) show that long-term clinical stability is related to bone remodelling stress-shielding and excessive stresses cause necrosis/resorption of bone [2], [3], [4], [5], [6].
Cemented stems caused less bone resorption and lower interface stresses than uncemented stems made from the same materials. Bone cement or other interlayers act as mechanical buffers, withstanding the mechanical stresses due to the mismatch between metal and bone properties. Critical stresses in the Charnley prosthesis are the tensile and shear stresses across the bone/acrylic interface. However, it appears that it is not possible to alter the metal and acrylic properties in a way that will significantly improve the stress situation in the metal/bone/acrylic system [7], [8], [9].
An appropriate stem may produce stress reductions in the range of 30–70% in the cement at the cement interface [10], [11]. Elastomer-coated prostheses (ECPs) have been developed in order to obtain the mechanical advantages of cemented THR [12], [13]. However, linear elastic models of transversally isotropic bone and isotropic MHPs suggest resorptive phenomena around the stems [2], [3], [14].
Calcar bone resorption is always a consequence suggested of MHP designs; the loss of proximal bone is estimated to be around 40–60% in 5 years. The higher the stiffness of the prosthesis, the greater is the bone remodelling, and hence the greater the bone loss due to the stress-shielding effect [8].
The biomechanics of bone growth, absorption, fracture healing, etc., are related to materials properties, structural properties and bonding characteristics of the implant. Metal stems are homogeneous isotropic materials; thus, MHP design focuses on the geometry [8], while composite HPs (CHPs) are material-structure designs, providing many new options and possibilities in implant design [15].
Fibre-reinforced composite materials can offer strength comparable to that of metals, and also more flexibility than metals. Mechanical properties of a CHP can be tailored to meet the bone mechanical behaviour [16].
Carbon fibre reinforcing thermoset resins such as epoxies were the first choice for composite prostheses [17]. Toxicity of eventual unreacted monomers suggests the use of thermoplastic polymers as a matrix for implant applications [18]. Polymers under investigation as a matrix for composite implants mainly include poly(sulfone) (PS), poly(ether-etherketone) (PEEK) and poly-etherimide (PEI) [15], [19], [20], [21], [22], [23], [24], [25], [26]. Material-structural designs may differ, although a common challenge is a stem more flexible than those of MHPs in order to improve proximal stress transfer [27], [28], [29]. These engineering polymers are characterized by high mechanical properties, thermal stability, very marginal water absorption and relatively easy processing. In addition, their high level of solvent and thermal resistance allow the production of a sterilizable medical device. Moreover the selected materials have demonstrated, at the same time, both positive and negative properties for the particular applications. PEEK has excellent mechanical stability but critical processing conditions due to its temperature-sensitive semi-crystalline structure. Polysulfone has shown a reduction of mechanical properties following saturation in Ringer’s solution.
Previous in vitro and in vivo [30], [31] studies have shown that PEI is an excellent substrate for cell spreading and growth, eliciting no cytotoxic response or haemolysis, coupled with both easy processability and resistance to sterilizability (γ rays and autoclave). These data suggested that PEI could be an attractive biomaterial, either alone or as a matrix for composite structure. In this work, PEI has been considered as a composite matrix reinforced with glass and carbon fibres for designing a composite hip joint prosthesis with adequate stability and mechanical properties.
By tailoring the stiffness of the prosthesis both along its length and through its thickness it is possible to change the pattern of load transfer between the prosthesis and the bone. Finite element modelling (FEM) combined with a mathematical description of adaptive bone remodelling suggests the high performance of a CHP in terms of mechanical stability and tissue conservation [8], [32], [33], [34]. In a previous work [34] the 3D FEM assumed linear-elastic behaviour of both CHP and bone. Materials properties of the composite stem were regarded as anisotropic and inhomogeneous, while the properties of the cortical bone were treated as orthotropic and homogeneous where the principal directions are those following lamellae orientation. The results led to drop-off plies of carbon/glass fibre-reinforced poly(ether-imide) (C-G/PEI) to constitute a symmetrical and balanced CHP, as characterized here.
Section snippets
Materials and methods
The composite structure was prepared by a pre-determined sequence of laminae constituted by carbon fibres and glass fabric pre-impregnated with PEI (Ten Cate Advanced Composites BV, Netherlands). The chemical structure of PEI is reported in Fig. 1. The fabrication of the carbon-glass reinforced PEI was performed in two separate stages. In the first stage carbon and glass reinforced plies were laid up in a curved female mould according to the macro-mechanical FEM design [34] and processed by
Results
Table 2 shows the average tensile, bending and torsional properties of the C-G/PEI standard specimens.
The distances of carbon-reinforced plies from the neutral plane do not affect the elastic modulus and strength measured in tension which are about 40 GPa and 600 MPa, respectively (zones I, II and III). Bending properties depend on the position of the carbon and glass plies from the neutral plane. The bending modulus (Eb) and the bending stiffness (EbI) increase in the zone IV–zone I direction
Discussion
The product between the elastic modulus (E) and the second moment of area is the bending stiffness (EI) or torsional stiffness (GJ) depending on the direction of the deformation. Loads acting on an implanted HP are characterized by in-plane (bending) and out-of-plane (torsion) behaviours which are related to bending and torsional stiffnesses, respectively.
The mechanical properties of standard specimens (Table 2) reflect a principle of material design which is a peculiarity of composite systems.
Conclusions
The axial stiffness of the compact femur is relatively constant over the length that was analysed, while the bending and torsional stiffnesses increase in the diaphysis–epiphysis direction. Fig. 4 shows stiffnesses and strength of bone compared to a titanium MHP (E=110 GPa) and to the proposed CHP. It can be noticed that using homogeneous materials such as metals the only shape design is not enough to realize a prosthesis characterized by stiffness similar to the surrounding bone and with
Acknowledgements
This work was supported by grants from Programma Nazionale di Ricerca per i Materiali Innovativi Avanzati and committed from the Ministero dell'Università e della Ricerca Scientifica e Tecnologica to the Biomateriali S.r.l.-Brindisi, Italy.
References (38)
- et al.
J. Biomech.
(1990) - et al.
J. Biomech.
(1993) - et al.
J. Biomech.
(1989) - et al.
Biomaterials
(1998) - et al.
J. Biomed. Eng.
(1990) - et al.
Biomaterials
(1998) - et al.
J. Biomech.
(1992) - et al.
J. Biomech.
(1998) - L.G. Gibson, M.F. Ashby, Cellular Solids, Pergamon, Oxford, 1992, pp....
Acta Orthop. Belg.
(1993)