Note: Descriptions are shown in the official language in which they were submitted.
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c
Method for producing a medical implant made of a beta-
titanium alloy, and a corresponding implant
The invention relates to a joint prosthesis having a
shaft made from a titanium alloy.
The major joints of the human body are subject to high
mechanical stresses. For example, the joints of the
locomotor apparatus have to bear a large part of the
body's weight, and moreover they are moved every time a
step is taken. Therefore, the bones which support the
joints have a powerful cortical structure. Their
integrity is important for sufficient functioning of
the joint. The same is true of the arm joints; although
the weight which they have to support is lower, they
are moved more frequently and are therefore exposed to
high levels of wear. Moreover, their dimensions are
smaller and they are more susceptible to injury.
Prostheses intended for permanent implantation
(endoprostheses) not only have to have sufficient
mechanical properties to ensure the desired
functionality, but also have to have a biocompatibility
that is as high as possible to ensure that they are
tolerated by the patient over a prolonged period of
time. In particular the latter aspect is very
important, since any incompatibilities which occur
generally require explantation of the prosthesis. This
equates to failure of the prosthesis.
It is known that inadequate transmission of load from
the prosthesis to the surrounding bone can lead to
degeneration of the bone tissue. This often leads to
the prosthesis coming loose. Therefore, to avoid this
degeneration, it is important to ensure loading that is
as physiological as possible by the prosthesis. Tests
have shown that hip prostheses with a lower modulus of
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elasticity produce a loading situation which is more
physiological than when using rigid prostheses. For
example, in the case of femoral prostheses, there has
been a move away from cobalt-chromium alloys, which
generally have a very high modulus of elasticity in the
region of approx. 200 000 N/mm2, toward titanium
alloys, which have a lower modulus of elasticity, such
as for example TiA16V4, the modulus of which is approx.
100 000 N/mm2. However, these levels are still well
above the modulus of elasticity of the cortical bone,
at approx. 25 000 N/mm2.
The invention is based on the object of improving a
joint prosthesis of the type described in the
introduction in such a way as to achieve more
physiological transmission of load.
The solution according to the invention lies in a joint
prosthesis having the features of the independent
claims. Advantageous refinements form the subject
matter of the subclaims.
According to the invention, in a joint prosthesis
having a shaft made from a titanium alloy, it is
provided that at least the shaft is investment-cast and
has a body-centered cubic crystal structure (known as
13-titanium alloy).
It has been found that the joint prosthesis according
to the invention can be used to achieve a significantly
lower modulus of elasticity. Depending on the titanium
alloy used and the heat treatment carried out, it is
possible to reach moduli of elasticity of approx.
60 000 N/mm2. This corresponds to virtually half the
modulus of elasticity which has previously been
achieved with titanium alloys. Furthermore, the
invention provides for at least the shaft to be
investment-cast. This allows more complex shaping of
the prosthesis. The forging processes which have
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hitherto primarily been used for titanium prostheses
only allow the production of relatively simple
structures. This restriction is overcome by the
invention. Consequently, the prostheses according to
the invention can be better matched to the loads which
are to be absorbed. For example, the shaping of the
prosthesis may vary more finely according to the local
stresses. The prosthesis only has to be of stronger and
therefore more rigid dimensions in specifically the
regions in which high stresses occur; in the other
regions, it can be of weaker and therefore more elastic
design. This allows the matching of the prosthesis to
the anatomical conditions to be further improved.
Moreover, it is easy for securing elements, such as
projections, to be formed integrally with the
prosthesis. It is possible to provide a greater number
of and more complex securing elements. Therefore, the
prosthesis is more suitable for cement-free
implantation. The benefit of the invention is that
complex shapes which cannot be practically realized by
forging processes can be achieved even for prostheses
made from 13-titanium alloys. In general, it will be the
case that the prosthesis together with the shaft is
investment-cast and heat-treated in one piece, although
the possibility of assembling the prosthesis from a
plurality of parts including the shaft should not be
ruled out.
The invention can advantageously be used for artificial
hip joints, in particular for femoral prostheses. These
are among the most highly stressed prostheses and have
a shaft of complicated shape for implantation in the
femur. It has been found that degeneration phenomena
readily occur in particular in the upper region of the
femur if a prosthesis that is too rigid has been
implanted. This often leads to failure of the
prosthesis. In the case of a femoral prosthesis
according to the invention, the modulus of elasticity
is considerably lower and therefore much closer to a
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physiological level of the bone material in the upper
region of the femur. The femoral prosthesis according
to the invention successfully counteracts the risk of
degeneration. The same applies to an embodiment in the
form of a knee prosthesis, which generally have very
long shafts.
It is preferable for the titanium alloy to be a
titanium-molybdenum alloy. The addition of molybdenum
stabilizes what is known as the Vphase of the titanium
alloy. This allows the formation of the desired
body-centered cubic crystal structure. Molybdenum as
alloying element has a lower toxicity than other
alloying elements which likewise stabilize the Vphase,
in particular niobium or vanadium. The reduction in the
toxicity is an important benefit of a prosthesis
intended for long-term implantation.
The level of the molybdenum or molybdenum equivalent in
the alloy is expediently in the range from 7.5 to 25%.
The result of this, in particular in the case of a
molybdenum content of at least 10%, is sufficient
stabilization of the Vphase all the way down to the
room temperature range. The content is preferably
between 12 and 16%. This allows a meta-stable 3-phase
to be achieved by rapid cooling after casting. The mean
grain size of the crystal structure is at least 0.3 mm,
preferably 0.5 mm. There is generally no need to add
further alloy-forming elements. In particular, there is
no need to add vanadium or aluminum. The elimination of
these elements has the advantage, which has already
been mentioned above, that it is possible to avoid the
toxicity emanating from these alloy-forming elements.
The same applies to bismuth, the biocompatibility of
which likewise does not match that of titanium.
Furthermore, the titanium-molybdenum alloy has the
advantage of having improved mold filling properties
compared to known alloys such as TiA16V4. This makes it
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,
possible to form sharper-edged structures by the
investment casting process.
It has proven particularly suitable for at least the
shaft of the prosthesis according to the invention to
be hot isostatically pressed and solution annealed. It
has been found that considerable improvements with
regard to brittleness are achieved with a material
which has been heat-treated in this way. The hot
isostatic pressing counteracts undesirable effects of
concentrating the molybdenum in dendrites while
deflecting the remaining melt by dissolving
inter-dendritic precipitations. A temperature below the
P-transus temperature, specifically at most 100 C below
the P-transus temperature, is expedient. Temperatures
in the range from 710 C to 760 C, preferably of approx.
740 C, have proven suitable for a titanium-molybdenum
alloy with a molybdenum content of 15%. The solution
annealing improves the ductility of the alloy.
Temperatures of at least 700 C up to 880 C, preferably
in the range from 800 C to 860 C, have proven suitable
for this purpose. There is no need for a preliminary
age-hardening before or after the hot-isostatic
pressing. For cooling after the solution annealing, the
shaft is expediently quenched with water.
The invention is explained in more detail below with
reference to the drawing, which illustrates an
advantageous exemplary embodiment and in which:
Fig. 1 shows a diagrammatic view of a first exemplary
embodiment of a joint prosthesis according to
the invention;
Fig. 2 shows a diagrammatic view of a further
exemplary embodiment of a joint prosthesis
according to the invention;
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The exemplary embodiment illustrated in Fig. 1 shows a
femoral prosthesis for an artificial hip joint. The
femoral prosthesis 1 consists of a P-titanium alloy,
namely TiMo15. This alloy has a body-centered cubic
crystal structure at room temperature.
The femoral prosthesis 1 is intended for implantation
at the upper end of the femur. It can interact with an
acetabulum component 2 which has been implanted in the
pelvic bone. The femoral prosthesis 1 has an elongate
shaft 10 as bone anchoring element and a neck 11 which
adjoins it at an obtuse angle. At its end remote from
the shaft there is arranged a joint head 12 which,
together with a bearing insert 22 of the acetabulum
component 2, forms a ball joint. Implantation involves
complete or partial resection of the head of the
thighbone neck, opening up access to the medullary
cavity of the femur. This access is used to introduce
the shaft 10 of the femoral prosthesis 1 into the
medullary cavity, where it is anchored. Depending on
the particular embodiment, cement is provided as
anchoring means or the fixing is effected without the
use of cement.
The femoral prosthesis 1 introduces mechanical loads
acting on the hip joint, whether static loads when
standing or dynamic loads when walking, into the femur.
Physiologically compatible transmission of loads is
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important for permanent reliable anchoring of the
femoral prosthesis 1 in the bone material of the femur.
If the femoral prosthesis 1 is of very rigid design, it
absorbs a considerable portion of the load, thereby
relieving the load on the bone material in particular
in the upper region of the femur. In the longer term,
this leads to degeneration of the femur in this region.
This leads to the risk of the femoral prosthesis 1
coming loose and ultimately of the prosthesis failing.
To prevent this failure mode, it is known per se for
the femoral prosthesis 1 to be of less rigid, i.e. more
elastic with a physiologically favorable low modulus of
elasticity design. In particular the shaft 10 of the
femoral prosthesis 1 is critical in this respect. In
the cortical region, the bone material of the femur has
a modulus of elasticity of approx. 20 000 to
000 N/mm2. According to the invention, the femoral
prosthesis 1 has a modulus of elasticity of approx. 60
000 N/mm2. This is a favorable modulus which is much
20 lower than that of materials which are conventionally
used, such as TiA16V4. These materials have a modulus
of elasticity of approx. 100 000 N/mm2 or even 200 000
N/mm2 in the case of cobalt-chromium alloys.
25 The invention allows simple production of even complex
shapes by investment casting. For example, the femoral
prosthesis 1 has a multiplicity of recesses and
sawtooth-like projections on its shaft 10. These are
used to improve anchoring of the femoral prosthesis 1
in the femur, allowing cement-free implantation. A .
plurality of grooves 14 are provided running in the
longitudinal direction of the shaft 10. They are
arranged on both the anterior and posterior side of the
shaft 10 but may also be provided on the lateral sides.
A plurality of rows of sawtooth projections 15 are
provided in the upper region of the shaft 10.
Furthermore, an encircling ring 13 is provided at the
transition to the neck 11. It can be designed as a
separate element, but the invention means that it may
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also be integral with the shaft 10 and neck 11. In
general, a single-piece design of the prosthesis is
preferred, with the exception of exchangeable or
optional attachment parts or wearing parts.
Furthermore, a fixing projection 16 is provided on the
shaft 10 adjacent to the ring 13 to prevent rotation.
Such complex shapes of joint prostheses can
conventionally only be produced from T1A16V4. However,
this material has a different, less favorable crystal
structure and therefore an undesirably high modulus of
elasticity.
The invention can advantageously also be used for other
types of joint prostheses. Fig. 2 illustrates a knee
prosthesis 3 as a further exemplary embodiment. It
comprises a femur component 31 and a tibia component
30. The femur component 31 has a long shaft 33 as bone
anchoring element. It is designed for implantation in
the medullary cavity of the femur, which has been
opened up by section of the natural knee joint. As in
the case of the femoral prosthesis, in this case too
the problem of degeneration of the surrounding cortical
structure occurs if the knee prosthesis 3, in
particular its shaft 33, is made too rigid. The same
applies to a shaft 32 of the tibia component 30.
The joint prosthesis according to the invention can
also be used for other joints, for example at the elbow
or the shoulder.
The text which follows describes a way of carrying out
the invention.
The starting material is a 13-titanium alloy with a
molybdenum content of 15% (TiMo15). This alloy is
commercially available in the form of small billets
(ingots).
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A first step involves investment casting of the parts
of the hip prosthesis. A casting installation is
provided for the purpose of melting and casting the
TiMol5. The casting installation is preferably a cold-
wall crucible vacuum induction melting and casting
installation. An installation of this type can reach
the high temperatures which are required for reliable
melting of TiMol5 for investment casting. The melting
point of TiMol5 is 1770 C plus a supplement of approx.
60 C for reliable investment casting. Overall,
therefore, a temperature of 1830 C needs to be reached.
The investment casting of the melt is then carried out
by means of processes which are known per se, for
example using wax cores and ceramic molds as lost mold.
Investment casting techniques of this type are known
for the investment casting of TiA16V4. The result is a
body-centered cubic crystal structure.
The castings, from which the casting molds have been
removed after the investment casting, are subjected to
a heat treatment. This involves hot isostatic pressing
(HIP) at a temperature just below the P-transus
temperature. This temperature may be in the range from
710 C to 760 C and is preferably approximately 740 C at
an argon pressure of 1100 to 1200 bar. It is expedient
for this purpose for a surface zone which may have
formed during casting in the form of a hard, brittle
layer (known as the a-case) to be removed by pickling.
This layer is usually approx. 0.03 mm thick.
Following the hot-isostatic pressing, the castings have
only a low ductility. It is assumed that this
embrittlement is attributable to
secondary
precipitations during the hot isostatic pressing and
the subsequent, generally slow cooling from the hot
isostatic pressing temperature.
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To dissolve these disruptive precipitations, the castings
are annealed in a chamber furnace under argon shielding gas
atmosphere. A temperature range from approx. 700 C to 860 C
for a duration of several hours, generally two hours, is
selected for this purpose. In this context, there is a
reciprocal relationship between the temperature and the
duration; a shorter time is sufficient at higher
temperatures, and vice versa. After the solution annealing,
the castings are quenched using cold water.
The mechanical properties achieved after solution annealing
are reproduced in the following table:
Solution Tensile 0.2%
Proof Elongation at
annealing strength Rm stress Rp break A5 [%]
temperature [N/mm2] [1\l/mm2]
[ C]
700 920 916 2.1
740 841 665 7.5
760 790 545 18.5
780 735.3 520 27.4
800 725 505 37.6
Solution Reduction of Modulus of Hardness HB30
annealing area after elasticity E
temperature fracture Z [kN/mm2]
[ C] [%]
700 10 68 285
740 19.3 66 278
760 23.4 65.4 268
780 40 63.7 260
800 52 59.4 255
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It can be seen that the modulus of elasticity drops as the
temperature rises during the solution annealing,
specifically down to levels of as low as 60 000 N/mm2. The
ductility values improve with decreasing strength and
hardness. For example, after solution annealing for two
hours at 800 C, the result is a modulus of elasticity of
60 000 N/mm2 with an elongation at break of approx. 40% and
a fracture strength Rm of approx. 730 N/mm2.
