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.
~f ~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 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 ~z~ss~
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 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 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 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.
~237S~;3 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 - lOa -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.
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 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 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 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 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 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 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 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 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 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 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 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;
~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.
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 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:~
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 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 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 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, 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.
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 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 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 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 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~
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 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 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 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 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 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 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 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;~
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 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.