Cement-free femoral prosthesis component and method of producing it

The present invention relates to a femoral prosthesis component for cement-free anchoring and to a method of producing it. The prosthesis component, which can be produced by means of CAD and image-analysis methods, provides for the largest possible surface for the transmission of forces, and its mass and rigidity can be adapted to the individual properties of the bone.

BACKGROUND OF THE INVENTION 
The present invention relates to a femoral prosthesis component which is to 
be anchored without using cement and a method of producing it. 
In the fields of surgery and orthopaedics with respect to the locomotor 
system, the artificial joint replacement has become a standard surgical 
intervention and is today one of the most frequently carried out 
operations of all. The long-term results of replaced joints have been 
quite variable and the life time of, for example, replaced hip joints 
ranges from few weeks up to 27 years (Draenert and Draenert 1992). 
It has been found out by scientific research that different factors are 
responsible for the loosening of an endoprosthesis component, such as 
infections, insufficient surgical skills, choice of the wrong implant and 
excessive strain. Nevertheless, so far many cases of loosening could not 
be explained in a satisfying manner. It has only been detected that 
certain combinations of factors frequently lead to a loosening, such as a 
massive cement-free implant used in combination with the bone of a 
rheumatic. Such types of prostheses which were to interlock within the 
bone and were implanted together with bone cement showed (Draenert 1988) 
that bone cement as a filling material between metal and bone cannot 
fulfil an anchoring function but is pulverised. 
The problem involved in the anchoring of prosthesis components could in the 
end be attributed to the phenomenon of bone deformability. This explained 
why an easily deformable bone of a rheumatic is deformed by a metal 
prosthesis anchored without cement such that rapid loosening ensued. On 
the other hand, it could be shown that a fragile or soft spongiosa as well 
as a normal spongiosa (cancellous bones) can be stiffened by means of 
polymethylmethacrylate (PMMA) bone cement and thus gets extremely rigid 
(Draenert and Draenert 1992). A thus stiffened bone structure could be 
found with all those implants which had successfully been used for 10 to 
20 years and could be histologically examined. On the other hand, quite 
compact femora could be provided with prosthesis components without using 
bone cement as an anchoring means, and these prosthesis components have 
successfully been implanted for about 10 years (Draenert and Draenert 
1992). However, in these cases, the results could not often be reproduced. 
It is an object of the present invention to provide a femoral prosthesis 
component which can be anchored without using cement and with which, after 
its implantation, good long-term results can be expected. 
SUMMARY OF THE INVENTION 
This object is achieved by the present invention. 
In connection with the invention, the problem has been investigated how the 
strength of a bone influences the life time of an implant. By means of 
histological studies it could clearly be proven that soft, deformable 
bones only show a stable anchoring if the implants used have a low mass. 
The present invention is based on the following findings regarding the 
anchoring of prostheses: Every bone exhibits an individual shape and an 
individual strength; therefore, both factors must be taken into account 
when selecting the prosthesis. A solid, compact bone is a good indication 
for metal-bone anchoring without using bone cement. Two factors are above 
all important in this context: 1. to obtain the best possible primary 
stability of the anchorage and 2. to provide the largest possible surface 
for the transmission of forces between the prosthesis and the bone. It 
has, however, to be considered that the various compartments of the bone, 
such as epiphysis, metaphysis and diaphysis, have completely different 
shapes and strengths. There were early trials to adapt the prosthesis 
shaft to the medullary cavity, cf. EP-A-0 038 908; however, it was rapidly 
found out that one single implant design could not be adjusted to the 
variety of different bone shapes (Noble et al., 1988); moreover, there was 
no possibility of determining the strengths of a bone and considering them 
when designing a prosthesis. 
In connection with the present invention, it has been found out that there 
is a good correlation between the density of a bone and its strength. 
According to the invention, the density of a bone can therefore be used as 
a measure for its strength. By combining various image-analysing and 
computer-aided calculations, a method could be found with which the 
morphology of the medullary cavity of the bone as well as the strength of 
the bone could be determined and taken into account for the design of a 
prosthesis component. These experiments resulted in a design of a 
prosthesis component which can be fit ideally into the medullary cavity 
and whose mass and/or stiffness can be selectively changed such that in 
each case the largest possible surface is provided for the transmission of 
force between prosthesis and bone. 
The mass and/or the stiffness of the femoral prosthesis component according 
to the present invention can be adjusted to the individual properties of 
the bone. In the medial, in particular the medioproximal portion of the 
prosthesis, the bending strength of the prosthesis is the decisive factor. 
In the lateral portion of the prosthesis, tensile stresses are predominant 
in the distal as well as in the proximal portion so that there the tensile 
strength of the prosthesis is also of importance. In the distal portion, 
the tensile strength is of particular importance. Due to the muscular 
attachments not covering the neck of the femur and the head of the neck of 
the femur, there are also torsional forces. In the present invention, the 
aforementioned properties of the prosthesis material are mostly summarised 
as "rigidity". According to the present invention, the rigidity of the 
prosthesis and/or its mass is adapted to the individual properties of the 
bone. 
There are several ways of adaptation; for example, the material of the 
prosthesis can be selected according to the individual properties of the 
bone. In case of a dense bone, a material having a higher specific mass 
and a higher rigidity can be selected whereas in case of a bone with a low 
density, the material to be selected has a low specific mass and rigidity. 
CoCrMo alloys, Ti, Ti alloys, steel, plastics or composite materials can 
for example be used as materials of the prosthesis. 
It is also possible to select an inhomogeneous material for the prosthesis 
component, in the sense that in portions of higher bone density a material 
of higher specific mass and/or rigidity is used than in portions of lower 
bone density. In this connection, it has to be considered that the bone 
density can greatly vary, and that the density of the spongy portion of 
the bone can be merely 15 to 20% of that of the compact substance of the 
bone. When using a porous prosthesis material, the desired inhomogeneity 
of the material can for example be obtained by varying the pore size and 
reducing it in portions of higher bone density. Composite materials can 
also be used as material for the prosthesis component wherein, for 
example, the fibre content of the composite material can vary along the 
axial length of the prosthesis component. Thus, in particular the rigidity 
of the prosthesis component can be varied and adapted to the bone density. 
Furthermore, the mass and/or the rigidity of the prosthesis component can 
be adapted to the individual properties of the bone by a suitable 
selection of the shape of the prosthesis component, particularly of the 
cross-section of the prosthesis component in various bone portions. If, 
for example, at least an essential part of the length of the femoral 
prosthesis component is U-shaped or horseshoe-shaped in its cross-section, 
as proposed in WO 90/02533, the cross-sectional area and thus the 
prosthesis mass can be adapted in various sections by a suitable selection 
of the size and depth of the groove or the slot between the two arms of 
the U-shaped cross-section. A transition from a solid shaft to a U-shaped 
cross-section with very thin arms is conceivable according to the 
invention. The largest possible surface for the transmission of forces 
between the bone and the prosthesis is guaranteed by the fact that the 
prosthesis component forms an uninterrupted surface or closed contour in 
its medial, dorsomedial and anteromedial portions. 
The mass and/or the rigidity can for example also be changed, in particular 
in order to reduce the mass and/or rigidity, by providing bore holes which 
partly pass through the prosthesis shaft, such as blind holes, or bore 
holes which completely pass through the prosthesis shaft. On the other 
hand, ridges and/or reinforcing elements provided at the outer and/or 
inner contours of the prosthesis, for example of a U-shaped prosthesis 
shaft, can increase the mass and/or rigidity of the prosthesis component. 
Such elements can be provided either on portions or over essentially the 
whole length of the prosthesis component. 
The mass and/or the rigidity of the prosthesis component is adapted to the 
bone density preferably by a linear correlation between the bone density 
and the mass and/or rigidity of the prosthesis component; that means, for 
example, that the mass or rigidity of the prosthesis in the respective 
portion of the prosthesis component is increased proportionally if the 
bone density is doubled. 
In detail, it can be proceeded as follows in order to design and produce 
such an individual prosthesis component: 
A patient having a deformed hip joint changed due to arthrosis is examined 
in a CT scanner, and stacked images of both hip joints are digitized and 
stored as cross-sectional images. So-called binary images are produced of 
the cross-sectional images by means of image analytical methods, i.e. 
black and white contrast images whose inner and outer contours can be 
analysed and which depict the femur. The inner contour is put together in 
a 3D model. The centre of rotation of the hip joint is determined and 
depicted as the centre of a sphere together with the contour model by 
means of the image analysis (cf. FIGURE). 
The shape of the shaft of the prosthesis component can subsequently be 
adapted to the shape of the medullary cavity. By means of several, for 
example six to ten, preferably nine sections which are evenly distributed 
along the length of the proximal femur, the density per unit area of the 
bone is determined via the binary image and compared with the 
corresponding section of a normal femur which has been previously 
analysed. This comparison results in a correlation factor as a measure of 
the strength of the individual bone, on the basis of which the ratio 
between the cross-section of the prosthesis and that of the medullary 
cavity can be determined. If the specific bone density corresponds to that 
of the normal femur, the contour model of the medullary cavity is 
eccentrically and/or concentrically reduced by 1 to 20%, preferably by 5 
to 10%, in order to determine the cross-section of the prosthesis 
component in the respective section. If the specific bone density is lower 
than that of the normal femur, the contour model is correspondingly more 
reduced to determine the cross-section of the prosthesis. The values 
between the individual sections can be interpolated. The set of data of 
the contour model is transferred together with the position of the centre 
of rotation to a CAD unit. In the CAD unit, the axis of the contour model 
is determined and undercuts in the design are corrected such that the 
prosthesis component can be inserted in a press fit manner into the 
medullary cavity with a rectilinear movement and/or with a slight screwing 
movement. The design such obtained is retransferred to the image-analysis 
unit in which a double contour model of the outer and the inner contours 
of the femur is produced, into which the prosthesis component can be 
fitted. Finally, while considering and correcting the enlargement ratio, 
the prosthesis component is projected into the ap X-ray image (path of the 
rays anterior-posterior) and the axial X-ray image, and inserted along its 
implantation axis. The mass and/or rigidity of the prosthesis is 
determined to be proportional to the bone density. Then the CAD data set 
is completed with the standard constructional data of the cone of the 
prosthesis neck for receiving the spherical head and of the implantation 
instrumentarium (drive-in/knock-out position of the prosthesis), and 
transferred to a milling unit. In the milling unit, the prosthesis 
component is milled from a blank, which is for example made of V.sub.4 A 
steel. Upon a surface treatment, the prosthesis component is washed and 
sterilised and can then be inserted.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS 
The FIGURE depicts a front view of the (implanted) prosthesis. 
The prosthesis according to the FIGURE, which is schematically depicted in 
the femur, comprises an attachable spherical head 1 which sits on a cone 2 
of the neck portion 3 of the prosthesis. Reference sign 4 designates the 
centre of rotation. The neck portion 3 is fixedly connected to a shaft 5 
of the prosthesis. In the sectional planes which are approximately evenly 
distributed over the length of the proximal femur and in which the bone 
density per unit area is determined, the optimum shaft cross-sections 5' 
obtained as described above are drawn. Eight hatched shaft cross-sections 
5' are drawn into the FIGURE and, for further clarification, three shaft 
cross-sections are additionally drawn at the side of the femur. 
Preferably, the bone density and the optimum shaft cross-section ensuing 
therefrom are determined in six to ten, for example nine sectional planes. 
Reference sign 6 designates the outer contour model and reference sign 7 
the inner contour model of the femur in each of the sectional planes which 
are obtained by the image analysis. The mass and/or rigidity of the 
prosthesis in the individual sectional planes can be adjusted by designing 
the shaft cross-sections suitably. If, for example, the mass is to be low, 
the slot or recess in the U-shaped shaft cross-section is enlarged, 
wherein at the same time the largest possible surface for the transmission 
of forces between prosthesis and bone is provided in the medial portion of 
the prosthesis. If the specific mass in a sectional plane is changed, the 
rigidity of the prosthesis component in this portion also changes, as a 
rule. Reference sign 8 designates the constructional axis of the 
prosthesis which is at the same time the axis of the medullary canal and 
the implantation axis. 
Literature: 
Draenert K. (1988), Forschung und Fortbildung in der Chirurgie des 
Bewegungsapparates 2, zur Praxis der Zementverankerung, Munich, Art and 
Science. 
Draenert K. and Draenert Y. (1992), Forschung und Fortbildung in der 
Chirurgie des Bewegungsapparates 3, die Adaptation des Knochens an die 
Deformation durch Implantate, Munich, Art and Science. 
Noble PC, Alexander JW, Lindahl LJ, Yew DT, Granberry WM, Tullos HS, 
Clinical Orthopaedics and Related Research, No. 235, October 1988, pp. 
148-163.