Biomedical Engineering Reference
In-Depth Information
result in the fracture of the metallic stem in the THR. Aggressive fatigue loading of cemented
artifi cial joints is responsible for the formation of microcracks in bone cement and stem/bone
micromotions [67].
Another issue is related to the fact that after insertion of prosthesis-a bone, prosthesis structure
is a composite one. It consists of bonded elements with different elastic and geometrical proper-
ties. These differences result in an altered load distribution in the artifi cial joint, as compared to
the natural one [67]. In natural joints, the loads are distributed over the entire cross-section of the
proximal part of the bone (i.e., femur). In the case of an artifi cial joint, the load is partially trans-
formed by shear forces across the bone-cement-prosthesis interfaces. This altered load transfer
leads to increased stresses at the cement-prosthesis interface and unloading of the bone away from
prosthesis. The interface shear stresses are further increased due to the stiffness ratio between the
prosthesis and the bone, typically at the order of 10:1 and higher. In addition, the bending displace-
ments in the bone surrounding the stem are reduced because of relatively high fl exural stiffness of
the prosthesis. The change in load distribution increases the stress in some regions and reduces it
in other regions. Areas with higher loads may experience an increase in bone mass, while areas
with reduced load may experience a decrease. Moreover, for an inadequate proximal fi t of the stem,
either initially as an effect of bone preparation, or gradually postoperatively as the effect of stem
subsidence, the proximal load transfer is bypassed in favor of distal one. This bypass mechanism as
well as stress shielding causes failure of the arthroplasty [67].
To improve joint replacement in terms of long-term component fi xation and wear properties,
future work should also be concentrated on the design of advanced prosthetic materials, which will
better mimic AC properties [33], such as water content, stiffness, shock absorption, promoting fl uid
fi lm lubrication, and low coeffi cient of friction.
One of the biomaterial design confl icts is between mechanical and biological compatibilities.
Many load-bearing implants require materials with a strength and durability stronger than bone,
because the implant lacks the ability of living bone to repair localized damage due to fatigue or
overloading. Strong materials should
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distribute high static and repetitive dynamic loads (up to 2500 N);
protect the cancellous bone from high stresses; and
allow to obtain a fi rm attachment to the underlying bones, leading to a long-term fi xation.
On the other hand, soft material for bearing is needed to
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improve wear properties of bearing surfaces in TJR;
provide a smooth, lubricated surface with fl uid fi lm and low coeffi cient of frictions;
reduce nominal contact pressure; and
increase joint congruence.
The soft material and its elastic deformation have been found to be the most signifi cant factor in
the prediction of the fi lm fl uid and low friction capability in artifi cial joints [71]. Cushion articulat-
ing surfaces consisting of low elastic modulus materials, which can articulate with full fl uid fi lm
lubrication, are needed to mimic the natural lubrication of the joint where synovial fl uid lubricates
the bone cartilage interface [74]. Moreover, when the fl uid fi lm breaks down, such as during periods
of heavy loading with little movement or at the start of movement, this biomaterial must give low
friction and a low wear rate in these conditions of mixed or boundary lubrication. Recent studies
have shown that polyurethane as a soft bearing material, which is articulated against a highly pol-
ished metallic surface, provides a lower coeffi cient of friction as compared with standard polyethyl-
ene versus metal bearings [75]. When polyurethane was replaced with water-swollen hydrogel, the
friction was considerably reduced [76].
