Note: Descriptions are shown in the official language in which they were submitted.
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Description
Prosthesis Fixation To Bone
Technical Field
The present invention relates to the fixation,
or fastening, of artificial joints or other prosthesis,
including dental implants, to bone without the use of a
cement or grouting agent.
Background Art
Presently, orthopedic surgeons most commonly
use polymethyl methacrylate (PMMA) cement for fixing
artificial joint components to bone. This techni~ue
has the advantage that high fixation strength is attained
immediately postoperatively. The patient can undertake
physical activity invol~ing the newly implanted joint
within a few days postoperatively. This is beneficial
for the patient's physical well-being because it stimu-
lates circulation and respiration.
However, joint implantation using the PMMA
cement has not been entirely satisfactory in the long
term. Artificial joints, and their fixation, must
withstand large mechanical forces. Especially so in
the weight bearing joints: hip, knee and ankle.
Transfer of these large mechanical forces from the
prosthetic joint to the bone is through a complex
structural system when cement is employed. The tensile
and compression strengths and the moduli of elasticity
vary greatly among the elements in this system: bone,
cement and prosthesis, which is commonly metal. The
cement is the least strong and the most flexible of the
three, and cement is also subject to brittle failure.
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~2:~7SS3
Failure of prosthetic joint implants is often traceable
to failure in the cement fixation. Many hip femoral
prosthesis stems have fractured after the supporting
cement interface has failed in one way or another.
Because of the demonstrated long-term inade-
quacy of prosthetic fixation using cement, it has been
a continuing objective to achieve direct fixation
between the prosthetic structural component and bone.
Numerous attempts to achieve this goal have been made
over many years. Investigators have employed or pro-
posed:
- metal components press fit impacted into
prepared bone canals or cavities. These components
were tapered pins with both smooth or irregular
surfaces, acetabular cups with "petals" or teeth
for cutting into the bone, stems with sintered
porous metal surfaces. See, for instance, U.S.
Patent No. 3,996,625 to Noiles.
- metal components with porous plastic coat-
ings of several kinds.
- metal components coated with a biologically
active and aseptic glassy material, sometimes
called a "bioglass" coating.
- porous plate elements have been tried but
their mechanical strength is very low.
- ceramic components with and without porous
or threaded surfaces, and with biologically active
ionic surface treatments.
- metal components with threaded stems, and
threaded ceramic acetabular cups.
~se of most of the above structures and
methods does not permit the initial implantation to
achieve intimate mechanical load transmitting relation-
ships between the prosthesis and bone. That is, they
i~3~7~;S3
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are intended to permit bone ingrowth into the porosities
or irregularities of the surface of the prosthesis.
This bone ingrowth phenomenon is reported to take place
in about one to five months in order to achieve adequate
structural strength for patient physical activity
involving the affected joint. During this time the
prosthesis-to-bone interface must be maintained without
motion, because it is known that motion at this interface
will cause the body to develop soft non~bony tissue at
this interface which provides inadequate support for
the prosthesis. Therefore, most of the above proposed
techniques anticipate restricted patient activity for
extended periods. Such restricted activity is not
desired for reasons of the patient's overall physical
health.
The threaded prosthetic stem concept can
provide initial intimate load bearing prosthesis-to-bone
interface. ~owever, the threaded stem has surface
discontinuities which severly reduce the fatigue endur-
ance strength of the prosthetic component. There areadditional difficulties in screwing into the prepared
bony canal or cavity the entire prosthetic component to
achieve the correct depth of insertion and angle of
orientation. For instance, a part of the prosthesis
may interface with a part of the bone when attempting
to screw the prosthesis into position.
Results of recent experience with prosthetic
joint components with porous metal surfaces which
foster bone ingrowth have confirmed that bone reshapes
and redensifies itself, by a behavior called "remodel-
ing", to suit the path of load transmission from the
prosthesis to the bone. This same experience also
demonstrates that it is desirable to transfer a maximum
fraction of the total load as close as possible to the
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normal joint surface in order to encourage the retention
of a maximum amount of normal bone mass. For example,
a femoral stem prosthesis for a hip joint which provides
for bone ingrowth at the distal end of the stem may
promote load transfer at that part of the prosthesis
with the result that the bone adjacent to the proximal
part of ihe stem will not carry a physiological share
of the total load and therefore will become less dense
and less strong. While a prosthesis so fixed may
function satisfactorily, such a biological change is
undesired in the event that the femoral stem prosthesis
ever has to be replaced, for any of a number of reasons,
in which case the surgeon is forced to deal with an
abnormally reduced amount of bone stock in the proximal
femur.
There are three principles which are generally
accepted to apply to the successful fixation of joint
prostheses by direct bone contact and support of the
prosthesis. One, the prosthesis must be in contact
with sound bone. That is, the bone to which force is
transmitted by the prosthesis must have adequate strength
to support the applied stresses. This implies that the
stress applied to the bone will be within the physiolog-
ical stress carrying capability of the bone. Two, the
prosthesis must be a good fit in the prepared bony
cavity. And three, there must not be motion between
the prosthesis and the bone.
It is clear that the above three requirements
are closely interrelated and very much dependent on
favorable geometric relationship between the prosthesis
and the bone. It must be true that if a patient's
joint and bone structure functioned to any reasonable
extent prior to implantation of a prosthesis, then the
; patient's bone quality is somewhere adequate to support
~1237553
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the loads due to that degree of function of that
particular joint. The problem then becomes one of
providing a prosthesis of the correct shape and size to
contact the patient's bone at the optimum interface
surface for satisfactory transfer of force from the
prosthesis to the bone. Further, the prosthesis must
satisfy the above and also fill the space created in
the bone with the utmost of congruency in order to
inhibit motion between the prosthesis and the bone. It
has been reported that bone may grow to fill spaces
adjacent the implant of up to 2mm. Certainly, spaces
however small between the implants and the bone do not
favor the necessary absence of motion therebetween.
Because humans vary so remarkably in physical size and
shape, we begin to see that each prosthesis should be
custom sized and shaped to suit the bone into which it
is to be implanted. Aside from the economic cost of
providing a custom prosthesis for each joint of each
patient, there is an overriding practical impediment to
so doing. The exact dimension for an optimum size and
shape of prosthesis cannot be determined before the
time of surgery when the bone is opened and its true
nature is learned.
The truth of the above may be substantiated
by the relative success to date of implantation of
joint prostheses using polymethyl methacrylate cement.
The cement serves the function of providing a custom
prosthesis for the individual bone at the time of
implantation. The bone is opened, explored, reamed and
broached to create a cavity which is surrounded by bone
judged by the surgeon to be of ade~uate strength to
support the forces to be received by the bone. The
basic prosthesis, usually metal, is available in an
assortment of shapes and sizes, perhaps as many as two
~237~53
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dozen. The utilization of PMM cement to fill the
spaces between the prosthesis and the bone is, in fact,
the creation of a custom prosthesis for that particular
implantation. The mechanical properties of the cement
are inadequate to provide a satisfactorily high percent-
age of successful implants for long term use, however.
With regard to dental prosthesis fixation to
bone, for more than ten years, attempts to implant
devices in human jaw bones where natural teeth are
missing have not been successful to the point where
even one moderately well-accepted design exists.
Experience to date has demonstrated that remodeling of
the jaw bone to accommodate the non-physiological
stress patterns introduced by the artificial implant
generally causes an undesired reduction in the total
volume of bone. It is a principal of physiology that
bone develops shape and density according to the manner
in which load is imposed on it. A change of shape or
density on account of a change of loading is called
bone remodeling. Further, the loss of bone is generally
in that part of the jaw where the implant emerges from
the bone to support the artificial tooth, bridge or
other dental appliance. This loss of bone is at least
partly attributed to the reduced stress in that part of
the bone where the implant emerges from the bone. This
occurrence has been reported particularly in patients
fitted with blade type implants.
Functional loads imparted to a natural tooth
or an implant are principally compression and bending.
There is little likelihood of any significant torsion
load being present. Current practice in implanting
dental anchorage devices favors non-loading of the
implant for an initial period of 2 to 4 months during
which time the bone supporting the implant recovers
~237S:;3
from the trauma of the implantation procedure. This
has been conveniently accomplished by using a two or
more part device, where the bone anchorage part is
implanted wholely within the jaw bone and the gum
tissue is closed over the implant for the initial time
period. One surface of the implant is approximately
flush with the alveolar ridge of the mandible or maxilla,
and through this surface there has been provided a
female thread, into which a second part of the prosthesis
having a threaded male stem can be fastened when the
gum tissue is penetrated for so doing.
The above known implanted parts may be exter-
nally smooth or threaded posts or cylinders, or blades.
Any of which may be of metal, carbon, plastic or ceramic,
lS either solid or porous, and uncoated or coated with a
variety of biologically acceptable materials.
As best understood, all of the above are
designed for approximately equal or uniform bony attach-
ment to all imbedded surfaces, and certainly in no
instance is there provision for enhanced bony fixation
in the area near the alveolar ridge and for less enhanced
bony attachment to that part of the implant which
extends relatively more deeply into the jawbone,
either mandible or maxilla.
According to the present invention there is provided a
prosthesis for mounting in a bone comprising a sleeve intended
for mounting in the entry part of an elongated cavity defined in
the bone, said sleeve to provide an initial, intimate, load-
bearing, prosthesis-to-bone interface with the part of the cavity
in the region of the cavity's opening, said sleeve defining an
external geometric pattern of projections which cut into the bone
when implanted in the bone to engage the bone, and a member
received within the sleeve, said member having a joint motion
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surface at one end, said member and the internal surface of said
sleeve defining mutually coacting means to fix said member in
said sleeve, said member extending deeper into said cavity than
said sleeve to provide an initial, intimate, load-bearing,
prosthesis-to-bone interface in the depth of said cavity.
According also to the present invention there is a pf~c
joint prosthesis component means for fixation to bone, comprising
first means defining a joint motion surface;
a stem attached to said first means defining a joint
motion surface for extending into the central canal of the bone
into which the component means is to be fixed;
tha~ part of the prosthesis components means which is
intended to be located within the bone at the end of the bone
near the joint motion surface defining an external geometric
pattern of elongated projections spaced circumferentially around
said prosthesis component means which engage with the bone when
implanted in the bone, said elongated projections having a
thickness of from about 0.5 mm to about 2.0 mm, a height of at
least about 0.7 mm, a spacing of from about 1 mm to about 4 mm
and an effective length at least ten times their thickness; and
the effective length of said external geometric pattern
of elongated projections is less than one-half of that portion of
the prosthesis component means which is intended to be implanted
within the bone.
Also, according to the present invention there is
provided a joint prosthesis component for fixation to bone
comprising a sleeve intended to be located in the end of a bone
near a joint for holding a stem means carrying a joint motion
1237~;~3
g
surface in a fixed position, said sleeve defining an external
geometric pattern of effectively elongated projections spaced
circumferentially around said sleeve which engage the bone when
implanted in the bone, said elongated projections having a
thickness of from about 0.5 mm to about 2.0 mm, a height of at
least about 0.7 mm, a spacing of from about 1 mm to about 4 mm,
and an effective length at least ten times their thickness.
Yet further, in accordance with the present invention
there is provided a bone implant support means for a dental
prosthesis having a first portion intended to be implanted in
bone and lie nearest the alveolar ridge, said first portion
constituting no more than half the depth of the portion intended
to be implanted in bone, and a second portion intended to be
implanted in bone and lie furthest from the alveolar ridge, said
first portion having an external surface which substantially
increases the first portion's interface contact area with bone in
A comparison to the second portion's interface contact area with
bone.
There is further provided in accordance with the
present invention a component for use in a prosthetic joint
comprising a hollow tubular means which has a closed end and an
open end, the portion of said tubular means which is adjacent to
the open end having external features which increase its surface
area to at least twice the surface area of the portion adjacent
to the closed end and the inner surface of said portion ad;acent
the open end defining a female part of a self-locking taper; and
elongated pin means having a first section which is adjacent to
one end of said pin means defining a male cone of self-locking
taper which is received in said female part and coacts therewith
and a second section which ad;acent the other end of said pin
means which projects from the open end of said hollow tubular
means to serve as a support for a joint motion surface.
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D i sc losure of the Invention
The present invention for implantation of an
artificial joint prosthesis derives from a consideration of
the three principal distinct types of force which may be
transmitted from a structural component of a joint
prosthesis to the host's bone, the recognized desirability
of transferring a maximum part of the total load to that
part of the bone which is closest to the joint motion
surface, and the additional desirability of not using PMMA
cement in that part of the bone which is closest to the
joint motion surface, which at the same time providing
immediate fixation with sufficient initial strength to
prevent motion between the prosthesis and the bone during
early physical rehabilitation of the patient. It is further
desirable to provide a prosthesis to bone fixation geometry
which disrupts the normal physiological blood flow pattern
to the minimum possible extent. It is recognized that the
' strength of fixation will increase with time as the bone
remodels itself to accommodate the new stress pattern
createa by the implantation of the prosthesis if there is no
motion between the prosthesis and the bone.
The three principal types of forces transmitted
between the prosthesis and the bone are compression, torsion
and bending. While tensile forces do exist in the weight
bearing bones, they are generally the result of bending. It
is highly unlikely that a joint prosthesis would transmit a
net tensile force to the bone. Further, the present
invention contemplates the transmission of tensile force
from the prosthesis to the bone, as will be discussed later.
A preferred embodiment of the invention will be
described as embodied in a hip joint prosthesis of the
proximal femur although it is applicable to any joint
prosthesis including but not limited to those for a
shoulder, elbow, wrist, knee, ankle, finger and toe. The
hip prosthesis is provided with a stem which extends into
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the canal of the femur for a distance of approximately 5 to
8 inches, although this could be longer should conditions
dictate. The stem carries a collar or flange, transverse to
the stem, which abuts the excised proximal end of the femur
where the head and neck of the natural femur have been
excised for implantation of the prosthesis. Proximal of the
collar the prosthesis comprises a neck portion which
supports the ball or head of the prosthesis at a distance
from the extended centerline of the shaft of the femur.
12:~755~
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The stem also carries a number o~ relatively
short longitudinal fins or splines adjacent to the
flange on the side of the flange opposite from the hip
joint, which fins are, at the time of implantation,
embedded in the prepared cancellous bone which exists
at the end of the bone adjacent the joint surface.
Some of the outer edges of the fins may contact and cut
into the inside of the cortical wall of the bone which
surrounds the cancellous bone. The side of the flange
which is in contact with bone and all of the surfaces
of the fins may be coated with a porous sintered metal
layer or any other textured or treated surface designed
to enhance fixation to bone.
The force of compression is transferred from
the prosthesis to the bone principally by means of the
collar which abuts the excised proximal end of the
femur. The collar is preferably shaped to contact
essentially all of the excised surface, which is more
or less transverse to the shaft of the bone.
The force applied to the femoral prosthesis
is exerted downward on the head of the prosthesis by
the acetabulum or socket of the hip joint, and passes
through the center of the ball. When the line of
action of this force intersects the centerline, or
extended centerline, of the femoral canal, or is parallel
to this centerline, the forces transmitted from the
prosthesis to the bone are limited to those of compres-
sion and bending. When the line of action of this
force is other than just described, then there is a
component of this force which must be transmitted from
the prosthesis to the bone as torque. That is, any
force applied to the head of the prosthesis whose line
of action is not in a plane which contains the centerline
of the femoral canal will create a torque about this
centerline.
lZ37SS3
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The short longitudinal fins or splines on the
stem and emanating from the collar are driven into the
prepared cancellous bone of the femur adjacent to the
excised surface of the femur to establish a tight fit
therein at the same time the collar abuts the excised
surface. The numerous fins provide a relatively large
surface area through which torque is transmitted from
the prosthesis to the bone at that part of the bone
closest to the joint surface. The fins act as keys to
prevent the prosthesis from rotating about the axis of
the shaft of the femur. The applied torque is resisted
by compression and shear forces which are distributed
throughout a large volume of the cancellous bone.
Finally, the bending component of the force
system will be transmitted from the prosthesis to the
bone by two opposed forces, one of which acts perpen-
dicular to the centerline of the femoral shaft at the
proximal end of the femur, essentially in the area
occupied by the finned part of the stem; and the other
of which acts perpendicular to the centerline of the
femoral shaft at the distal end of the prosthesis. The
bending component is a large part of the complete force
system, and the forces which constitute the two forces
described above will be larger if the distance between
them is small.
Preferably the part of the prosthetic stem
contained within the bone will be 5 to 8 inches long.
This length creates reaction forces to bending which
will not exceed the acceptable load capacity of the
bone which envelops the stem. Also, it is preferred
that the surface of the distal end of the stem not
transmit the force components of axial compression or
torque from the prosthesis to the bone; therefore, this
surface should not be textured, coated or treated to
~237S53
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enhance transfer of shear loads at the surface interface
with the bone. It is important that the distal stem
fit securely within the femoral canal to prevent any
transverse movement between the prosthesis and the
bone. It is also contemplated that polymethyl meth-
acrylate cement can be used advanta~eously to fix the
distal prosthesis stem in the canal of the femur. The
inventive structure limits the load transfer at this
point to a reaction force to bending. That is, the
principal stress in cement so used will be in compres-
sion between the prosthesis stem and the wall of the
canal of the femur. The cement is satisfactory for
this type loading. This technique adds to the ability
to provide custom fit with a limited number of component
sizes.
On the other hand, the surfaces of the collar
and the fins are preferably textured, coated or treated
to enhance transfer of shear loads at the prosthesis-to-
bone interface in this area. It can be seen that such
transfer of shear loads will contribute to maximizing
the transfer of all three load type components to the
proximal bone of the femur. The fins will thus transmit
some of the pure compression load as well as some of
the compression and tension loads resulting from bending.
The underside of the collar can transmit a part of the
torque by transmitting shear forces at the collar-to-
bone interface. The function of the fins will not be
diminished if some of the fins at this outer edge
contact the cortical wall of the femur. In fact, this
circumstance may be beneficial.
Additionally, it is known that the principal
avenues of blood supply within the femur are longitudinal.
The longitundinal fins provide the advantage of providing
a multitude of paths for stress transfer from prosthesis
53
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to bone in the proximal femur with a minimum disruption
of the blood supply within the femur, while also per-
mitting the regeneration of physiologically desirable
longitudinal blood paths.
Because human bones come in an endless variety
of diameter, length, wall thickness, taper, curvature,
etc., an alternative construction is proposed which has
the greater practical utility. In this alternative
embodiment, the collar and longitudinal fins are integral
with a thin wall truncated conical sleeve. The large
end of the sleeve carries the collar which extends
radially outward. The surfaces of the sleeve which
contact bone may be textured, coated or treated to
enhance fixation to the bone. Or, the entire sleeve
may advantageously be made from a suitable porous
metal.
With the alternative sleeve embodiment, the
femoral component has no collar or fins. The shaft of
the stem is smooth and tapered to lock within the
sleeve by the well-known principle of mechanical tapers.
This embodiment offers the advantage of permitting a
variety of size selections for the sleeve component and
for the stem component separately. Thus a smaller
number of total components is needed to achieve a given
number of total size combinations for the final assembly.
The two-component embodiment may be more economical to
manufacture, and will conveniently allow selection of
size and implantation of the sleeve before the stem is
implanted.
With either of the above embodiments, one
aspect of preparation of the femur consists in reaming
the intramedullary canal to a cross-sectional shape and
size and to a depth to accept the shaft of the stem of
the prosthesis so that the distal part of the prosthesis
1237~;~;3
1 5
will be securely held within the femur without the
possibility of transverse motion between the prosthesis
and the bone. There must be a selection of sizes of
reamers, and a selection of sizes for the distal stem
of the prosthesis so that the above condition of fit is
obtained. Alternatively or concurrently, use of PMMA
cement may be advantageously confined to fixing the
distal prosthetic stem in the femoral canal as described
above.
A second aspect of preparation of the femur
consists in creating an essentially transverse surface
of the proximal femur against which the collar will
abut to lie in a plane which is the same as the plane
which the underside of the collar will define when the
prosthesis is implanted. A bone cutter and guide can
be provided to permit this condition to be obtained. A
selection of prosthetic components with varied collar
areas must be available so that one can be chosen which
will closely match the shape of the bone against which
the collar fits. The prosthesis is intended to be
implanted with the collar fully seated against the
mating bone.
A third aspect of preparation of the femur
consists in broaching multiple slots into the cancellous
bone of the proximal femur, which slots are to accom-
modate by press or impacted fit the multiple fins of
the prosthesis. A selection of sizes of broaches must
be available so that the slots can extend radially as
much as the particular femur will allow. A selection
of prostheses with various sizes of multiple fin
envelopes must also be available to correspond to the
several sizes of broaches so that the prepared slots
can be filled with fins in tight proximity to cancellous
bone, and in some areas the edges of some of the fins
~23'7S53
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will be in tight proximity to the cortical wall o~ the
femur. If the slots in the bone are each 1.5mm wide,
each fin will be somewhat thicker, say 1.6mm to 1.7mm
thick, so that the fins must be driven into the slots.
In this manner the proximal femur is immediately in a
preloaded fit to the prosthesis, and motion between the
prosthesis and the bone is prevented during the early
physical rehabilitation of the patient. Angular location
of the broached slots is made consistent with the
desired angular orientation of the neck of the prosthesis
in the final implanted condition, if the embodiment
requires.
Alternatively, the longitudinal fins can be
made self-broaching and a selection of prostheses
provided with the volume envelope of the fins increasing
in a series of sizes. Successive prostheses are driven
into the femur bone and removed to be replaced by the
next larger prosthesis until the desired security of
fit achieved. This technique is preferred to be used
with the thin walled sleeve construction.
Thus the implantation of either embodiment
provides initial mechanically strong fixation to resist
motion between the prosthesis and the bone which could
result from the three principal forces. Motion due to
compression is resisted by contact between the collar
and the excised surface of the bone, including the
cortical margin of the bone, especially the region
known as the calcar. Motion due to torque is resisted
by the many securely implanted fins in the relatively
large volume of cancellous bone in the proximal femur,
as well as by some engagement between the edges of some
fins and the cortical wall. Motion due to bending is
resisted by a large fraction of the fins in the proximal
cancellous bone at the one force and reaction area, and
12375S3
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by the secure fit of the distal stem in the femoral
canal at the second force and reaction area.
The initial fixation is s-trong enough to
prevent motion between the prosthesis and the bone
during postoperative recuperation and rehabilitation.
As stated above, motion between the prosthesis and the
bone will cause the development of soft non-bony tissue
which is inadeauate to support the prosthesis.
~ne advant2c,es of the prostllesis of the-~re--er~-ed e~Gl~.ts
of this invention are fourfold. Cne, the prosthesis is designed,
sized and installed with immediate load bearing juxta-
position between the several elements of the prosthesis
and the associated bone. Bone does not have to grow
into the spaces between the fins as it has to grow into
the interstices of porous or other irregular surfaces.
Two, PMMA cement is not used in the highly loaded part
of the bone nearest the joint motion surface. This is
the area where the use of cement has proven to be the
least successful. Three, with a planned and controlled
program of increasing patient activity, the bone remodels
itself to accommodate the new force patterns, and the
fixation becomes stronger the more it is used. It is
to be emphasized that the prosthesis transmits a maximum
of load and stress to the bone at the most proximal
part of the femur, so that the density and strength of
this proximal bone may be preserved. Fourth, the
prosthesis creates a minimum disruption of the normal
blood supply paths within the proximal femur.
Yet another aspect of the present invention
is directed to the provision of a thin walled truncated
conical sleeve with a smooth inner surface. The outer
surface is preferably threaded in the manner of a
self-tapping bone screw. The thread is preferably of a
- high lead, multi-start configuration which permits
1237S~;3
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rapid advance during insertion in combination with a
large number of threads to provide a large load bearing
thread area.
Use of self-tapping threads is preferable to
prior tapping with a tap, because with self-tapping
threads bone chips stay in contact with the threads to
fill spaces which are bound to exist due to the irregular
and non-homogenous nature of bone, and because these
bone chips become nuclei for the growth of new bone in
the manner of a bone graft.
The conical taper of the sleeve and the high
lead multi-start thread provide a very practical benefit
in permitting the sleeve to be screwed home in relatively
few turns, preferably fewer than four turns. A sleeve
may be two inches long and may have threads in the
pitch range of 8 to 25 per inch. A straight single
start thread 2 inches long with 25 turns per inch will
require 50 turns for full insertion. A tapered single
start thread 2 inches long with 25 turns per inch will
require a number of turns to seat which depends on the
thread depth and taper angle in the following relation-
ship:
depth of thread tangent of angle of taper
per side
N turns x advance per turn
For example, a 3 taper per side (tangent of .05), a
thread depth of .03 inch and a 25 turns per inch single
thread will re~uire 15 turns for full engagement. A
thread with 5 starts of the same pitch will require
only 1/5 as many turns, or 3 turns to attain full
thread depth engagement.
~23~iS3
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The sleeve is screwed into the prepared canal
or cavity of the bone to achieve intimate and mechanical-
ly strong contact between its threaded outer surface
and the bone, and to be i.n desired axial aIignment and
depth of position with the bone to accept a load bearing
component of the prosthesis.
In one embodiment, the load bearing component
of the artificial joint prosthesis fits tightly into
the smooth inner conical surface of the sleeve and is
locked therein by the well-known principle of mechanical
machine tapers, and also extends through the sleeve
with a part of the load bearing component extending
further into the prepared bony canal or cavity for
additional fixation to the bone and stabilization of
the prosthesis. In another embodiment, the outer
surface is conical, truncated, and preferably threaded.
The cone is closed at the small end, and the inner
surface may be cylindrical, hemi-spherical or other
suitable geometry. The load bearing component of the
prosthesis fits within the inner surface of the sleeve
and is locked therein by screw threads or other suitable
means.
me preferred embcdiments of the invention provide immediate
intimate structural relationship between the prosthesis and the
bone of mechanical strength sufficient that the patient
can start limited weight bearing activity within a few
days postoperatively. The initial fixation is strong
enough to prevent motion between the prosthesis and the
bone during postoperative recuperation and rehabilita-
tion. As stated above, motion between the prosthesisand the bone will cause the development of soft non-bony
tissue which is inadequate to support the prosthesis.
Because it is a principle of physiology that
bone develops shape and density according to the manner
123~SS3
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in which load is imposed on it, and because the
prosthesis will transmit force to the bone in a manner
different from that imposed by the original natural
joint, it is true that the bone will have to reshape
and redensify itself before the new prosthetic joint to
bone fixation attains maximum strength. This is true
whenever the pattern of force application to a bone is
changed.
The material of the prosthetic sleeve must be
biologically acceptable to the development of bone in
intimate contact with the sleeve. Preferably the
sleeve is made of titanium alloy, specifically an alloy
known as Ti6 A1 4V. This alloy is highly resistant to
corrosion and is well tolerated by the body. It has
high mechanical fatigue endurance strength and a modulus
of elasticity, or stiffness characteristic which, while
approximately five times greater than that of bone, is
approximately half that of other metals commonly used
in artificial joint prostheses.
With regard to the dental implant there is
provided a two-part support for a dental prosthesis,
one part of which is implanted entirely within the
jawbone. Of the implanted part, that portion adjacent
the alveolar ridge has its external bone interface con-
tact surface area significantly increased by screw
threads, and that portion distant from the alveolar
ridge lacks the external thread which increases the
bone interface contact surface area of the adjacent
portion. The other part, which projects outward from
the crest of the jawbone through the gum tissue, fits
within the first part and is held therein by a self-
locki~g mechanical taper.
In the preferred two-part embodiment, the
part implanted in bone is an elongated hollow tubular
lZ37553
- 21 -
member closed at one end. The outer surface of the
closed end is smooth and the outer surface of the open
end is threaded in the manner of a self-tapping bone
screw. The self-tapping thread is preferred because
bone chips created during insertion stay in contact
with the threads to fill spaces, which exist due to the
porous nature of bone, to become nuclei for the growth
of new bone in the manner of a bone graft. The thread
may be tapered or straight. The opening in the threaded
end of the member provides the female cone of a self-
locking taper.
The part which extends outward from the bone,
through the gum tissue to support the prosthesis is a
pin, stud, or post having three zones. One end of this
part is a zone which is the male cone of a self-locking
taper which fits within the implanted hollow tubular
member. The center zone of this part is a smooth
cylinder which extends through the gum tissue. The
other end of this part is a zone on which is mounted
the prosthetic appliance, bridge or single tooth by any
suitable means, as for instance by a second self-locking
taper. A single tooth may be fused directly to this
other end.
A second embodiment of the dental implant
provides a threaded sleeve with a self-locking taper
therethrough and a prosthesis supporting post which
extends through the sleeve and into direct contact with
the bone. The length of the post in contact with the
bone has a smooth surface and is at least as long as
the threaded sleeve.
A third embodiment of the dental implant
provides increased interface contact area with bone by
means of multiple longitudinal fins or flutes rather
than screw threads. The fins or flutes are preferably
self-broaching.
1~375S3
- 22 -
For implantation of each of the above embodi-
ments, the jawbone is prepared by drilling and reaming
a hole in the jawbone accurately sized to receive -the
smooth end of the implant in a tight fit and is also
sized to accept the root diameter of the threaded or
fluted end of the implant. Thus, the threads or flutes
must cut their way into the bone. This action provides
immediate intimate structural relationship between the
prosthesis and the bone, thereby preventing motion
between the prosthesis and the bone during the postopera-
tive period. It has been shown that motion between the
prosthesis and the bone will cause the development of
soft non-bony tissue which is inadequate to support the
prosthesis.
As with the artificial joint prosthesis, the
implanted material for the dental prosthesis must be
biologically acceptable to the development of bone in
intimate contact with the prosthesis. Preferably the
inventive prosthesis parts are made of titanium or
titanium alloy, especially an alloy known as Ti6 Al 4V.
These metals are highly resistant to corrosion and are
well tolerated by the body. Their strength and stiffness
characteristics are appropriate to this use, as is well
known.
srief Description of the Drawinss
Embodiments of the invention will now ~e described in the follcwing drawings:
FIG. l is an oblique view of a femoral
prosthesis of an artificial hip joint showing a typical
force system acting on a prosthesis of the present
invention, including the reaction forces exerted on the
prosthesis by the femur;
FIG. 2 is the prosthesis of FIG. l showing
the force system components acting in the plane through
the centerline of the stem and the center of the sphere;
- 23 - 1237~53
FIG. 3 is the prosthesis of FIG. 1 showing the force
system components acting in a vertical plane through the
centerline of the stem;
FIG. 4 is a general view of an implanted artificial hip
joint embodying the teachings of the present invention;
FIG. 5 is a detail view of a prosthesis of the proximal
femur showing self-broaching longitudinal flutes according to
the present invention;
FIG. 6 shows an alternate form of the present
invention, used for implantation and fixation of a femoral
prosthesis;
FIG. 7 is a plan view of an alternate form of the thin
wall fluted sleeve shown in FIG. 6;
FIG. 8 is a section view of the sleeve of FIG. 7;
FIG. 9 is another view of the sleeve of FIG. 7;
FIG. 10 is a detail view of the femoral prosthesis of
FIG. 6;
FIG. 11 is found on the same page as FIG. 6 and is an
additional view of the sleeve of FIG. 6;
FIG. 12 is found on the same page as FIG. 6 and shows
an impact instrument used for implanting the self-broaching
sleeve of FIG. 8;
FIG. 13 is found on the same page as FIG. 7, 8, 9 and
10 and is a section through the sleeve of YI5. 8 along the line
XIII-XIII;
FIG. 14 is found on the same page as FIG. 7, 8, 9 and
10 and shows alternative self-broaching flutes on the sleeve of
FIG. 8;
FIG. 15 shows a hip joint fermoral prosthesis stem
implantation embodying the teachings of the present invention;
FIG. 16 is a detail view of the sleeve of the present
invention, partially in section;
- 24 -
~Z;~3
FIG. 17 is a view of a femoral prothesis embodying the
teachings of the present invention;
FIG. 18 is found on the same page as FIG. 15 and shows
a reamer used to prepare the femoral canal for the implantation
of a femoral prosthesis according to the present invention;
FIG. 19 is found on the same page as FIG. 15 and shows
an insertion and alignment instrument used to install the sleeve
of the present invention;
FIG. 20 shows a knee joint prosthesis implantation
using the present invention;
FIG. 21 is a rear view of the knee joint implantation
of FIG. 20;
FIG. 22 shows an alternate form of the present
invention used for the implantation of a hip joint acetabular
prosthesis;
FIG. 23 is found on the same page as FIG. 16 and 17 and
is a detail view of the sleeve of FIG. 22, partially in section;
FIG. 24 shows the implantation of a hip joint
acetabular prosthesis using an alternate form of the sleeve of
FIG. 22;
FIG. 25 shows additional detail of the sleeve of FIG.
24;
FIG. 26 is a perspective view of part of a lower
jawbone showing an implant embodying the teachings of the
present invention;
FIG. 27 is a sectional view through the implant and
jawbone in the plane indicated in FIG. 26;
FIG. 28 is a sectional view in the same plane as FIG.
27, showing a temporary screw plug in place in the tubular
member;
FIG. 29 is a sectional view through the jawbone and an
implant of an alternate embodiment;
FIG. 30 is a fragmentary perspective view of a lower
jawbone showing another alternate embodiment;
1;~37553
- 25 -
FIG. 31 is a section through the implant of
FIG. 30 in the plane indicated in FIG. 30; and
FIG. 32 is a sectional view similar to FIG.
27 showing the reaction forces applied to the implant
by the jaw bone due to a bending force component applied
to a prosthetic tooth.
Best Mode for Carryinq Out the Invention
With reference now to the drawings, wherein
like references characters designate like or correspond~
ing parts throughout the several views, there is shown
in FIG. 1 a diagrammatic representation of a typical
force system acting on the femoral prosthesis 10 of an
artificial hip joint. The femur 12 is shown inclined
at approximately a 30 angle to the horizontal, a
position corresponding to that of a person arising from
a chair. This discussion will treat static forces
only, because they well illustrate the principles
involved. Also, the plane through the centerline 14 of
the femoral stem 16 and the center of sphere 18 is per-
pendicular to a vertical plane through centerline 14.
Also, for purposes of discussion the reaction forces
between the prosthesis 10 and femur 12 are shown acting
at a point where they are in fact each distributed over
some surface area.
The force FW being transmitted from the hip
joint socket, not shown, to the femur acts vertically
downward at the center of the sphere 18. In the vertical
plane containing FW and parallel to centerline 14, FW
can be replaced by two components FC and FT, where FC
is parallel to centerline 14 and FT at 90 to FC liesin a plane perpendicular to centerline 14.
Reaction forces exerted by the femur 12 on
femoral prosthesis 10 in resisting forces FT and FC are
12~7S5:~
- 26 -
assumed to act on centerline 14. These reactions can
be analyzed separately and se~uentially in the appro-
priate planes. FIG. 2 shows the reaction forces to FC
acting in the plane through centerline 14 and the line
of action of force Fc. Here, FC can be replaced by F
acting on centerline 14 and the moment Ml which equals
FC x d. This system is resisted by the reaction force
FCB and a couple whose forces FMl and FM2 act perpen-
dicular to centerline 14 at opposite ends of the
implanted prosthetic stem as shown. FCB, FMl and FM2
are forces exerted on the prosthesis 10 by the femur 12.
In FIG. 3, first consider the forces acting
in the plane containing the line of action of force FT
and perpendicular to centerline 14, here force FT f
FIG. 1 can be replaced by its equivalents, force FTl
acting perpendicular to centerline 14 at P, and moment
M2 which equals FT x d. Moment M2 works on stem 16 of
the prosthesis 10 and sets up the equal and opposite
reaction moment M3 exerted by the bone on the prosthesis
at the area of interface most resistant to rotation of
the stem within the femur 12. This area is specified
to be concentrated at point N.
In FIG. 3, next consider the forces in the
vertical plane through the line of FTl and the center-
line 14, here we show one reaction force FR to be nearthe proximal end of femur 12 at point N. A summation
of moments reveals that the remaining bone reaction
force FS will vary inversely as the length of stem 16.
A summation of forces perpendicular to centerline 14
reveals that reaction force FR is equal to FTl plus Fs.
Therefore, a short stem will cause the greater reaction
at FR and is undesirable.
Again, with reference to FIG. 1 in combining
forces FM2 and FS at the distal end of prosthesis 10
~237553
- 27 -
graphically, one sees the net reactive force at this
point to be FG- Combining forces FR, FM1 and FCB at
the proximal end of prosthesis 10 one sees the net
reactive force at point N to be FH. In addition, one
sees that the reaction moment M3 will exist at the
proximal end of the femur 12 provided the prosthesis lO
is designed to transmit torque to the bone at this
point, and only this point.
From the above it is clear that the forces
and torque which are transmitted from the prosthe~is to
the proximal femur may be considerably greater than
those transmitted at the distal end of the prosthesis.
This situation is advantageous if the prosthesis is
designed to transmit the larger forces to the proximal
bone in a manner which the bone accepts favorably. One
thrust of this invention is that the proximal bone will
best accept large forces from the prosthesis when the
prosthesis is configured so as to diffuse or dissipate
large forces into a large volume, or against a large
area, of bone, both cancellous and cortical. When
these large forces are so diffused to load the proximal
bone of the femur within its normal physiological
stress limits, the bone will respond by maintaining an
adequate volume and density or by remodelling to have a
volume and density which is greater than that attained
over time with prior art devices.
FIG. 4 shows a femoral prosthesis 10 of an
artificial hip joint implanted in femur 12 according to
the teachings of the present invention. Preparation of
the femur 12 includes excising the neck of the femur at
a surface which will abut the undersurface 24 of collar
22 when the prosthesis is implanted, reaming and broach-
ing the femoral canal to accept the stem 16 of the
prosthesis, and broaching slots in the proximal cancellous
1237553
- 28 -
bone of femur 12 to accept the longitudinal fins 26
which extend distally on stem 16 from collar ~2.
Fins 26 may alternatively be termed as ribs, splines,
flutes, keys, etc., as long as the result is an external
geometric pattern of elongated projections.
The fins 26 have a thickness as small as is
reasonable to manufacture and handle without damage,
approximately in the range from 0.5 to 2mm. The fins 26
have a height of at least 0.7mm and the spaces between
the fins are approximately 1 to 4mm. It is to be
emphasized that the fins provide a primary force trans-
mitting interface of this invention, and that this is
different from the bone-to-prosthesis interface of the
so-called bone ingrowth concepts using porous materials,
because according to the invention interdigitation is
created at the time of implant, and the bone projections
within the geometric envelope of the surface of the
finned part of the prosthesis have a minimum width of
lmm and a minimum height of approximately 0.7mm.
Further, these bone projections have a length dimension
longitudinally of the fins of 10 or more times their
width. That is, their length may be 10, 20, 30mm or
more. Further still, in the annulus space 52 in FIG. 13
where bone and fins are interdigitated when the prosthesis
is implanted, the ratio of volume space occupied by
bone to that occupied by fins is always greater than
1 to 1, and may be as high as 5, 6 or 7 to 1. ~ndeed,
the theoretically ideal ratio of respective volumes in
the interdigitated space is the inverse of the strengths
of the two materials, or for bone and implant grade
metals, approximately 20 or 25 to 1.
In contrast, bone ingrowth into porous material
takes at least several weeks and the bone projections
into the pores have a maximum dimension of 0.5 to lmm
~237~;3
- 29 -
in any direction. The porous material in U.S. Patent
3,855,638 specifies a maximum porosity of 40%. There-
fore, where the bone and porous material occupy the
same space, the ratio of volume space occupied by bone
to that occupied by metal is always less than 1 to 1.
The mechanics of the porous metal to ingrown bone is
not efficient, because the metal is stronger than the
bone by approximately 20 to 1 on a volume basis.
In the preparaton of the femur 12 preferably
each cavity in the bone is cut slightly smaller than
the part of the prosthesis which will fit in the
corresponding part of the cavity. Bone will accept the
prosthetic elements so driven into undersized cavities,
albeit at a great spread of allowable dimensional
interference. The soft cancellous bone will easily
yield to accept prosthetic intrusion, while the hard
cortical bone of the femoral shaft will yield only
slightly, and can be split if asked to accept too great
an interference fit.
The size relationships of the elements of the
prosthesis to the bone of FIG. 4 are very important.
Collar 22 preferably has a size and shape to cover the
entire area of excised bone in contact with the under-
surface of the collar at 24. The envelope of the
volume of the fins 26 preferably corresponds closely to
the size and shape of the cancellous bone at the proximal
end of femur 12. The cross section of the femur in
this area is more elliptical than round, having a
larger diameter medially to laterally than anteriorly
to posteriorly. Accordingly, the envelope of the
flutes should be elliptical on the same axes. It is
preferred that the flutes fill the cancellous bone
space sufficiently that the outer edges of the flutes
contact some cortical bone of the wall of the femur,
1237553
- 30 -
especially at the anterior, posterior and medial aspects.
The distal stem 20 must fit within the shaft o~ femur 12
so that there is no transverse movement or looseness in
any direction. Preferably the stem 16 fits tightly in
the canal of the femur as a result of femoral reaming
and prosthesis size selection. FIG. 4, however, shows
an alternate implantation technique where the distal
stem 20 is held securely within the femoral canal by
the use of PMMA cement confined to the area 32. As
explained above for the inventive construction, forces
transferring from the prosthesis 10 to the femur 12 are
much smaller at the distal end than at the proximal end
of the prosthesis. Under this circumstance, the PMMA
cement will provide satisfactory long-term security of
fixation of the distal stem 20 within the canal of
femur 12. Further, the invention tends to create
minimal axial shear and torque loads on cement so used.
FIG. 5 shows a detailed view of the femoral
prosthesis 10 where the fins at 28 are designed to be
self-broaching so as to cut their own path into the
cancellous bone of the proximal femur. FIGS. 4 and 5
are generalized drawings to illustrate the principles
of the invention.
Refer to FIGS. 6 through 11 which show alter-
native preferred embodiments where collar 122 and
flutes 126 and 128 are attached to the thin wall conical
sleeve 34. In this case, the femoral prosthesis 40 has
no collar or fins, but rather has the conical taper 42
which fits within the tapered inner bore 36 of sleeve 34
by the well-known principle of self-locking tapers. In
the implanted condition of FIG. 6, the sleeve 34 and
the stem of prosthesis 40 are locked together as a
single unit and will respond to the applied force
system as described above.
123~$5~
- 31 -
This embodiment provides numerous practical
advantages. To cover a range of size and shapes of
femurs 12, fewer sizes of femoral prostheses 40 are
required when compared with the construction of FIG. 5.
A large assortment of fluted sleeves 34 is reguired to
permit selection for optimum fit in the proximal femur.
The sleeves 34 are much cheaper and smaller than the
prostheses 10 or 40, however, and therefore present
much less of an inventory problem for the manufacturer
and the user. The sleeve 34 may be fabricated from
porous metal, or have its outer surfaces coated or
treated by any of a number of techniques designed to
enhance fixation to bone. The prosthesis 40 becomes an
uncompromised structural member, and can be designed
and fabricated to that purpose. The procedure for
implantation is facilitated for the surgeon, because it
can be done in a seguence of simple steps.
To implant the prosthesis embodiment of
FIG. 6, the surgeon excises the head and neck of the
femur to provide access to the femoral canal. The
canal is reamed to a depth and diameter to accept the
stem 16 of the prosthesis 40. Reamers are provided
which are matched to the lengths and diameters of the
several prosthesis stem sizes available. The reamer
also removes bone to accommodate the wall thickness of
sleeve 34. Generally, the surgeon will select the
largest reasonable stem size which will fit withln a
given femur. An appropriately designed instrument,
located in the reamed canal, is used to cut the proximal
surface of the femur around the canal at 90 to the
centerline of the reamed canal. This surface will then
abut accurately with the undersurface of the collar 24
when the fluted collar is driven into position. A
multi-fluted broaching tool can be used to prepare a
1237~i53
- 32 -
bed for the fluted sleeve 34, where each groove cut in
the femoral bone will be slightly smaller than the
corresponding flute which will fit in the groove. of
course, the broached grooves must be in correct angular
orientation for the construction of FIG. 6 where anti-
rotational lugs 46 engage the shoulder 48 of prosthesis
40. Anti-rotation lugs 46 may be furnished as an
assurance to the surgeon, but they are not necessary to
prevent rotation of stem 16 within sleeve 34 when the
stem 16 is solidly seated within the taper 36 of sleeve
34.
FIG. 8 shows an embodiment of sleeve 34 with
self-broaching flutes 128 and a short internal thread 38
at the small end of the tapered bore 36. Each flute or
rib 128 is shown to extend the entire length of sleeve
34 and stepped teeth 35 are shown as a means of making
the flute self-broaching. Alternatively, the broaching
teeth on each longitudinal flute can be formed by cuts
or notches 29 shown in FIG. 14 which interrupt the
lengthwise continuity of the flute 128. The individual
flute segments 31 lie in lengthwise alignment as shown
at 33 in FIG. 9. The effective length of an indi~idual
flute or rib 128 is specified to be the total length of
a series of segments 31 which are in longitudinal
alignment, and which segments follow one another into
the same space in the bone as the sleeve is being
implanted in the bone. This sleeve may conveniently be
used with the impact instrument 60 of FIG. 12 fo~
implantation into the proximal femur when the femur has
not previously been broached with a fluted broach.
Sleeve 34 fits on taper 62 o instrument 60, with
collar 22 engaging shoulder 64 and thread 38 engaging
thread 66. Stem 68 aligns instrument 60 in the reamed
femoral canal. Slide hammer 70 is then reciprocated to
123~ ;3
- 33 -
drive the fluted sleeve 34 into the proximal femurO
Sleeve 34 is available in a series of increasing sizes
of envelope of the flutes 28 for each given tapered
bore size 36. The surgeon first implants the sleeve
with the smallest envelope. The energy re~uired to
seat the sleeve gives an indication of the security of
seating. I f the tightness of fit is judged inadequate,
the sleeve is removed by operating the slide hammer 70
in the outward direction. Threads 68 engaged in threads
38 permit the sleeve to be so extracted. The next
larger fluted sleeve is implanted, and so on until the
surgeon is satisfied that a sleeve 34 is securely fixed
in the proximal femur 12. It is recommended that at
least four sizes of sleeve be furnished for each stem
size, and that the increase in sleeve fin envelope
diameter be approximately 1.5mm per size.
When a sleeve 34 has been securely seated in
the proximal femur, the surgeon selects a femoral
prosthesis 40 of correct taper diameter 42 and of
diameter of distal stem 50 to fit securely with the
reamed femoral canal. The distal end of stem 16 of
FIGS. 6 and 10 is shown with longitudinal flats 50.
These flats are designed to increase the latitude of
diametral fit of the stem in the bone for which there
will be no lateral motion between the stem and the
bone, and to reduce the hazard of splitting the femoral
shaft by an overly tight fit. The distal end 50 of the
stem 16 is made with a smooth surface and is neither
intended to transmit axial shear load from the stem to
the bone, nor intended to transmit tor~ue from the stem
to the bone.
Should there be any reason for the distal
stem 50 to not engage securely with the femoral canal,
the surgeon may elect to place PMMA cement in the canal
~237S53
- 34 -
in the area indicated at 32 in FIG. 4 prior to the
final insertion of the femoral prosthesis 40, and the
prosthesis is driven solidly into engagement with the
internal taper 36 of sleeve 34. Again, it must be
èmphasized that cement is not used in the cancellous
bone of the proximal femur. Note that in the sequence
just described, the finned sleeve is fully implanted
before cement would be delivered to the femoral canal
for anchoring the distal stem. This sequence prevents
cement from entering the interface between the fins and
the cancellous bone.
The femoral prosthesis 40 implanted according
to the above description is fixed to the femur with
adequate strength to permit early physical therapy and
i5 rehabilitation of the patient. The pattern of load
transfer to the femur creates stresses which favor the
retention of and development of sound bone in proximal
femur, and increased activity by the patient will tend
to improve the bone structure in accordance with the
above.
With regard to the aspect of the present
invention relating to the thin walled truncated conical
sleeve with a smooth inner surface, there is shown in
FIG. 15 the proximal end of a human femur 110. Phantom
line 112 shows the normal head of the femur which has
been excised for implantation of the prosthesis 120.
The hard outer shell 112 of the femur is known as dense
or cortical bone, and the less dense inner bone 114 is
known as spongy or cancellous bone.
Fermoral stem prosthesis 120 is shown implanted
in femur 110 with threaded sleeve 140 being used to
provide initial high force resisting fixation; force C
is axial compression and force B1 is the bending force
due to the patient's weight force W being offset from
the femoral shaft.
i237S53
- 35 -
The femoral prosthesis 120 implanted in com-
bination with sleeve 140 may be said to be one half of
a total hip joint prosthesis. The sleeve 182 and cup
184 shown in FIG. 22 are the other half of a total hip
joint prosthesis. Note that sleeves 140 and 182 do not
perform any of the motion functions of the joint, while
stem 120 and cup 184 each perform part of the motion of
the joint at the contact area between sphere 129 and
cup 184.
Stem prosthesis 120 and sleeve 140 are
mechanically locked by mating tapers 122. Mating
metallic cone tapers are self locking when the included
angle of taper is less than approximately 15, depending
on the particular metals, finish and coefficient of
friction. The preferred taper is approximately 6
included angle, or 3 taper per side. It is preferable
to make both sleeve 140 and stem 120 of titanium alloy,
which metal against itself has a high coefficient of
friction and galling tendency, further enhancing the
locking ability of taper 122. The stem 120 is seated
firmly in sleeve 140 by means of mallet blows to surface
128, or preferably to an intermediate protective device
used between the mallet and the prosthesis.
FIG. 18 shows reamer 130 which is used to
prepare the canal of femur 110 for implantation of stem
prosthesis 120 and sleeve 140. Reamer 130 has various
diameters and tapers as follows. Portion 132 of reamer
130 is tapered and sized to provide optimum function of
the external self-tapping threads 144 of sleeve 140.
Portion 134 is sized to provide a press fit of portion
124 of stem 120 of 0.5 mm. or more. This portion is
preferably tapered approximately 2 included angle for
both the reamer 130 and the stem 120. Portion 136 is
sized to provide bone which will be broached by the
~2:~7SS~
- 36 -
axial self-broaching flutes on portion 126 of stem 120.
This portion is preferred to be cylindrical or non-
tapered on both stem 120 and reamer 130. The secure
fixation of portion 126 of stem 120 in femur 110 is
required to resist the second bending force B2. Firm
fixation of stem portion 126 is also importantly required
to resist torsion forces on stem 120, thereby relieving
sleeve 140 of the need to resist torsion forces during
the early course of bone healing by redensifying and
reshaping to accept the new loading pattern.
Thin walled threaded sleeve 140 is shown in
detail in FIG. 16 in partial section. Bore 142 mates
with taper 122 of stem 120. Screw thread 144 is pre-
ferably a multi-start (2-6) thread of 1 to 3 mm. pitch
and 2 to 10 mm. lead. The taper, plus the high lead,
allows the sleeve to be screwed home in relatively few
turns, with the sleeve becoming ever tighter with addi-
tional turning. The outside diameter of sleeve 140 can
have one or more tapers, and can taper differently from
the bore 142. Indeed, for a given bore 142, a selection
of different thread O.D.'s and tapers should be made
available to the surgeon so threads 144 can be chosen
to suit the density and thickness of a particular
patient's cancellous bone 114, much as a wood screw
which does not hold well can be replaced by a screw of
larger diameter. Thread depth can also vary.
In self-tapping threaded sleeve 140, multiple
cutting flutes, as at 146, may be provided; and multiple
slots 148 are provided to permit driving the threaded
sleeve 140 into position. In order that insertion of
sleeve 140 is made in proper geometrical alignment with
the cavity prepared by reamer 130, insertion tool 150
is provided as shown in FIG. 19. Sleeve 140 is placed
on tool 150 where bore 142 is a free fit on taper 152,
3 237~;S3
and slots 148 of the sleeve are engaged by pins 158.
By using tool 150 to drive sleeve 140 during insertion,
axial alignment is maintained between tapered bore 142
of the sleeve and prepared canal diameter 116 in the
femur. This alignment prevents inducing unwanted
bending stresses in the system when stem 120 is implant-
ed, as would happen if sleeve 140 were not aligned with
the prepared canal.
After sleeve 140 has been implanted, stem 120
is inserted into the prepared canal of femur 110 through
bore 142. The head 129 of the femoral prosthesis 120
must be positioned in correct rotary angular location
relative to femur 110 before the self-broaching longi-
tudinal flutes 126 penetrate the prepared canal 116.
Surface 128 is provided for contact by an appropriate
instrument for driving prosthesis 120 into final seated
position.
Femoral prosthesis 120 is furnished in a
selection of sizes, and a size must be chosen so that
fluted portion 126 engages at least some of the cortical
bone 123 of femur 110. Of course, each different size
of prosthesis 120 requires a different set of mating
sleeves 140, and a different reamer 130.
The advantages of the invention would not be
attained by putting the external thread 144 of sleeve
140 directly on tapered portion 122 of prosthesis stem
120. The first reason is that a threaded prosthetic
stem 120 could not be screwed home in the femur as
shown in FIG. 15. The neck 127 of the prosthesis would
interfere with the greater trochanter 118 of the femur
110 preventing installation to the desired geometry as
shown. The second reason is that the torque required
for screwing home during installation must be within
the capability of the surgeon to apply, and is estimated
1237S53
- 38 -
to be approximately 125 lbs feet of tor~ue. The
magnitude of tor~ue applied to the prosthetic stem by
even limited patient activity is estimated to be
approximately the same. It is also apparent that were
a stem 120 with integral threads at 122 to be used,
there could not be at the same time axial broaching
flutes as at 126. Therefore, were the only torque
resistance to rotation of prosthesis 120 to be afforded
by threads as proposed immediately above, the prosthesis
would be inadequately fastened to be sure it would not
move relative to the bone were the patient to initiate
even limited use of the joint within a few days post
operatively. Thirdly, a threaded outside diameter on
the prosthesis stem 120 would significantly reduce the
fatique bending strength of the stem.
FIGS. 20 and 21 show the implantation of an
artifical knee joint using the inventive externally
threaded thin wall sleeve. The femoral prosthesis 162
has a stem 164 which locks within the mating taper of
externally threaded sleeve 166 shown partially in
section. The mechanics of installation are identical
with those described for installing the femoral hip
prosthesis and related sleeve shown in FIGS. 15 through
19. The necessity for being readily able to achieve
correct angular position of femoral prosthesis 162 is
apparent in this application because the condylar
portions of the prosthesis are positioned within a
cutout in the distal condylar portion of the natural
femur. It would not be possible to screw home such a
femoral prosthesis were the screw threads on the stem
of the prosthesis.
On the tibial side of the knee joint of FIG.
20, an externally threaded thin wall sleeve 170 is
anchored by being screwed into the prepared tibial
1237553
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canal. In this instance, however, the part within the
threaded sleeve is a thin walled bearing 172 which has
the integral flange 174. This bearing 172, with its
integral flange 174, is preferably made from ultra
high molecular weight polyethylene plastic and bearing
172 and integral flange 174 serve as an axial and
thrust bearing for and receives metal tibial stem
prosthesis 176. Tibial stem prosthesis 176 is free to
rotate within bearing 172 and also free to distract
from thrust bearing flange 174 as described in U.S.
Patent No. 4,219,893 to Noiles.
FIGS. 22 and 23 show a tapered externally
t~readed sleeve 182 and a bearing cup 184 received
within sleeve 182 forming the acetabular prosthesis 180
of an artificial hip joint. The multi-start thread 181
is interrupted by multiple flutes 183 which enhance the
self-tapping ability of the threads. Multiple slots
185 are provided to engage a driving tool during inser-
tion. A properly proportioned reamer, not shown, is
provided to prepare the bony acetabular cavity in the
pelvis 186 for implantation of the sleeve 182.
After sleeve 182 has been implanted, bearing
184 which has screw threads at 188 is assembled into
sleeve 182 by engagement with screw threads 187 in
sleeve 182. Bearing 184 may be of metal or plastic. It
may encompass more than 180 of the sphere 129 of
femoral prosthesis 120 of FIG. 17. Such a construction
is described in U.S. Patent No. 3,848,272, now U.S.
Reissue Patent No. Re. 28,895 issued to Noiles. Or the
bearing 184 may encompass 180 or less of the sphere
129, in which case it may have been assembled with
sleeve 182 prior to surgery.
An alternate advantageous embodiment is shown
in FIG. 24, comprising a tapered externally threaded
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sleeve 192 into which is received acetabular bearing
194 by means of a threaded attachment like that shown
in FIG. 22. A part of the acetabular bearing 194 and
sleeve 192 assembly has been cut obliquely at 191, at
an angle of from 10-30, to allow additional range
of motion of femoral prosthesis neck 127 in the segment
of the cut 191. The cutout portion of cut 191 extends
for less than 180 of the threaded periphery of bearing
194. Orienting the oblique cut at 191, anteriorly
permits additional motion to the anterior of the body,
and indeed more closely mimics the natural anatomy.
Because the cut l91 is desired to have a specific
angular orientation when the threads 193 are firmly
seated in the bone, the thread taper angle, lead, and
depth must permit knowing in advance very closely how
much rotation is required to achieve full depth thread
engagement from the starting condition where the crowns
of the threads first contact the prepared cavity. For
instance, the rotation to achieve tightening can be
made to be very close to 360 if the depth of the
thread divided by the lead equals the tangent of the
angle of taper of one side.
FIG. 25 shows radial flutes 195 formed in the
surface 197 of sleeeve 192 which have two functions.
One, the reamer used to prepare the cavity for insertion
of sleeve 192 has a convex surface to cut the bottom of
the acetabular cavity with a concavity at 197 and,
therefore, to leave bony material in place in the area
198 which interferes with the full depth seating of
sleeve 192. Flutes 195 can be forced into this inter-
ferring bone to provide gradually increasing resistance
to the final seating of sleeve 192 when threads 193 are
` in approximately full engagement with bone. This
feature allows a small angular range of rotation at
1237~i~j3
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final seating to assist locating cutout 191 in the
desired position. Second, the bone chips cut into the
flutes 195 will develop with new bone generation to
improve the fixation of sleeve 192.
With reference generally to FIGS. 26-32, and
FIG. 26 in particular, there is shown the two-part sup-
port for a dental prosthesis 210 implanted in jawbone
212 between two natural teeth 214. The drawings,
except where noted, omit showing the soft gum or gingival
tissues.
FIG. 27 shows, in section, the hollow tubular
member 216 which has external thread 218 at one end, the
smooth exterior at the closed end 220, and the self-
locking internal taper 222. Thread 218 and the smooth
exterior of closed end 220 each occupy about half the
length of hollow tubular member 216. Member 216 also
has a hexagonal or splined socket at 224 with which a
suitable inserting tool engages for screwing member 216
into the prepared cavity in jawbone 212.
The prepared cavity in jawbone 212 accepts
member 216 so that the smooth closed end 220 is a tight
fit. That is, the diameter of the reamed hole is
somewhat smaller than the diameter of smooth portion
220. The nature of the bone within jawbone 212 yields
to accept portion 220 in the manner in which a piece of
wood accepts a nail driven into an undersized pilot
hole. The prepared bone cavity is reamed to approxi-
mately the root diameter of screw thread 218 of member
216 to the depth that the screw thread 218 penetrates
the jawbone. The hole in the jawbone may be tapped
prior to implanting member 216, but preferably thread
218 is of self-tapping design. When such a self-tapping
thread is inserted once and left in position, the bone
chips created during the insertion lie in the interface
lZ37~i53
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between the threads and the bone, filling small spaces
which exist in the bone, and serve as nuclei for the
growth of new bone cells which more firmly anchor the
implant.
S FIG. 27 shows the second member, pin 230,
positioned within tubular member 216. Pin 230 has
three zones, each with a different function. The zone
at 232 is a male self-locking taper which mates with
the corresponding female self-locking taper 222 within
member 216. The central zone 234 is generally cylin-
drical and smooth and is that portion which passes
through the gum tissue which covers jawbone 212. The
third zone 236 serves as the fastening area for the
prosthetic bridge, tooth, or appliance. The zone 236
is shown as a male self locking taper, but could have
any suitable configuration such as a male or female
screw thread, or a grooved configuration as shown in
FIG. 32 where a single artificial tooth is fused or
cemented directly to pin 230.
The structure of FIG. 26 can be created at
the time of implantation and the pin 230 used for
limited function while the bone and gum tissue heal
around the implant because the implanted system has
sufficient mechanical strength to prevent unwanted
motion between the implant and the bone. However, some
practitioners currently favor a more conservative
postoperative course using the means shown in FIG. 28.
In FIG. 28, tubular member 216 is provided
with a female screw thread 226 formed in a recess at
end of socket 224 in the closed end. A temporary screw
plug 227 is inserted to seal the opening in tubular
member 216, after which gum tissue 228 is closed over
the implant by suture 229. This condition is maintained
for the desired time, perhaps 2 to 4 months, during
~1~37~i3
- 43 -
which time the bone and gum tissue recover from the
trauma of surgery and the bony anchorage of member 216
becomes even more secure. After this time of healing,
the gum tissue is opened, plug 227 removed, and pin 230
is securelv inserted.
When the implant is functional, loads are
transmitted to the bone by the implant in a manner
different from that done by the natural tooth. There-
fore, the jaw bone will remodel its shape and density
to accommodate the new load distribution pattern. The
desirable load transmission pattern of the invention
will be discussed with reference to FIG. 32. The total
load on the artificial tooth 280 and the implant 210 is
shown as FT. The largest component of FT is known to
lS be downward compression force Fc. Compression force FC
is transmitted from the implant 210 to the bone 212 by
shear and bending of the bone adjacent screw threads
218, by shear at the interface between the bone and
smooth surface 220, and by compression of the bone
beneath closed end 221. The greatest area of contact
between implant 210 and bone 212 is at screw thread
218. Therefore, the greatest amount of load will be
transmitted to the bone at the screw thread. This load
transfer will cause an increase in the amount and
density of bone adjacent the screw threads. This
occurrence is most desirable because recession of bone
at this alveolar ridge area has been an ever present
problem inhibiting long-term success in most dental
implants to date. Design proportions and postoperative
activity must be controlled to keep the stress on the
bone within physiological limits, because excessive
stress is also reported to cause destruction of bone.
However, with generous surface area of the threads and
gradually increasing functional loads, the jawbone
~23'7553
- 44 -
will remodel itself to support the prosthesis in a
favorable manner.
With reference to FIG. 32, the benefit of
increased implant to bone interface area adjacent the
alveolar ridge is vividly illustrated with respect to
the bending load component FB applied to the artificial
tooth 280. Bending component force FB will cause two
reactive bending forces to occur between the bone and
the implant, shown at RT operating on the threaded area
218 of implant 210 and RS operating on the smooth
closed end portion 220 of implant 210. Depending on
geometry, RS may be approximately equal to FB. Force
RT must equal FB plus Rs, because the summation of
horizontal forces must be equal to zero. From this we
see that RT must be approximately equal to twice Rs.
Accordingly, the interface between implant and bone
should be larger at RT that it is at Rs. The inventive
construction satisfies this requirement.
An alternate embodiment is shown in FIG. 29,
comprising a tapered external threaded sleeve 238 into
which is received pin 250. The four functional zones
of pin 250 are the smooth extended end portion 252
which fits securely in a prepared cavity in bone 212,
the self-locking male taper 254 which fits within
sleeve 238, the generally smooth cylinder 256 which
penetrates the gum tissue, and the bridge, tooth or
appliance mounting portion 258, which is again shown as
a male self locking taper. This embodiment has the
advantage that several external size variations of
sleeves 238 can be combined with several length varia-
tions of pins 250 to provide a greater number of overall
size combinations with fewer parts than can the con-
struction of FIG. 27. This makes for more economy in
manufacturing and in sales and hospital inventory
12375S;~
- 45 -
storage. To provide for closed early healing, a
temporary stub pin having only zones 252 and 254 can be
implanted during the time the gum tissue remains closed
over the implant. Alternatively, only sleeve 238 can
be implanted initially with a temporary plug of zone
254 shape in place while the implant is covered. In
this case, the cavity for zone 252 of pin 250 would be
prepared after the time of initial bone healing around
sleeve 238. Sleeve 238 can be made with fluted or hex
splines 239 in the small end of its bore to provide
means for driving into place, and for removal if
necessary.
A second alternative embodiment of the tubular
member is shown as member 260 in FIGS. 30 and 31. The
principle of increased bone to implant interface area
adjacent to the alveolar ridge is provided by the multiple
longitudinal fins 262. This embodiment permits a somewhat
larger pin to be used because the walls of tubular member
260 can be thinner at the buccal and lingual regions 266
due to the absence of external threads. In this case,
the compression load component FC is transmitted to the
bone adjoining fins 262 more by shear than by compression
or bending, and providing a porous or textured surface
on the fins is advantageous. The prepared bony cavity
for this implant is sized to accept the tapered portion
270 and the extended portion 264 with a secure tight fit,
as described above, and the fins are preferably shaped
to broach or cut their own path into the bone 212. Again
bone chips created by the broaching act as nuclei for
new bone growth. Thread 268 is provided for attachment
of an inserting tool for use during the implant proce
dure. It permits removal of one size of tubular member
260 when the clinician believes that use of a larger
size would be desirable.
12375~3
- 46 -
Smooth cylinder 220 (FIG. 27), smooth cylinder
252 ~FIG. 29) or smooth cylinder 264 (FIG. 30) may be
substituted by a cruciform shape (four flutes) or any
irregular or other regular cross section (including a
varying cross section). The important feature is that
the surface area of the upper half of member 216 (FIG.
27), and its corresponding part in the other figures,
is at least twice the surface area as the lower half.
Numerous modification and variations of the
present invention are possible in the light of the
above teachings. For instance, use of a single thread
is within the contemplation of the invention which has
characteristics which fall within the parameters given
for multi-start threads. Also, the threads on the
outer surface of the thin wall sleeve may be of a
porous metal or ceramic or may be treated with a
biologically active coating.
In addition, with regard to the dental implant,
certain porous coatings could be applied to the implanted
surface adjacent the alveolar ridge, while the deeper
implanted surface could be uncoated and smooth. Or the
closed end of the pin may have a cruciform cross section
to increase its flexibility and thereby perhaps improve
the force transfer pattern between prosthesis and bone.
Also the outer surface of the implant as illustrated
may be a porous metal, ceramic, plastic or carbon or
may be treated with a biologically active coating.
The invention is applicable for other implants
in the human or animal skeleton, as for instance for
artificial joints for most joints of the body, i.e., the
ankle, shoulder, elbow, wrist and finger. It is there-
fore understood that, within the scope of the appended
claims, the invention may be practiced otherwise than
as specifically described.
