|Publication number||US3852045 A|
|Publication date||3 Dec 1974|
|Filing date||14 Aug 1972|
|Priority date||14 Aug 1972|
|Publication number||US 3852045 A, US 3852045A, US-A-3852045, US3852045 A, US3852045A|
|Inventors||M Karagianes, K Sump, K Wheeler|
|Original Assignee||Battelle Memorial Institute|
|Export Citation||BiBTeX, EndNote, RefMan|
|Patent Citations (9), Non-Patent Citations (3), Referenced by (176), Classifications (19)|
|External Links: USPTO, USPTO Assignment, Espacenet|
United States Patent  Wheeler et al.
. [451 Dec. 3, 1974 1 VOID METAL COMPOSITE MATERIAL AND METHOD  Assignee: Battelle Memorial Institute,
 Filed: Aug. 14, 1972  Appl. No.2 280,266
 US. Cl .I 29/182, 3/1, 75/200, 75/214, 75/222, 128/92 C  Int. Cl C226 l/08, C226 l/O4, B22f 3/16, A61C 13/30, A61f 1/24  Field of Search 75/222, 200, 214; 29/182; 128/92 C; 3/1
 References Cited UNITED STATES PATENTS 2,943,933 7/1960 Lenhart 75/214 3,052,967 9/1962 Fischer 75/222 X 3,362,818 1/1968 Schwarzkopf et a1. 75/222 X 3,535,110 10/1970 Todd 75/222 X 3,605,123 9/1971 Hahn 128/92 C X FOREIGN PATENTS OR APPLICATIONS 527,192 7/1956 Canada 75/222 898,140 6/1962 Great Britain 75/222 784,124 10/1957 Great Britain.... 75/222 717,034 10/l954 Great Britain 75/222 OTHER PUBLICATIONS Schwarzkopf, P. Controlled Porosity of Microstruc- 11% VIBRATORY LOADlNG OF MICRQSPHERES 1N CUP MAcHM-io SHAPES OF VMC MATERIAL BY VAPORIZAT1ON tures by Powder Metallurgy, in Int. J. Powder Met, 2(4); pp. 3-1l, 1966.
Welsh, R. P. et al. Surgical Implants in J. of Bone & Joint Surgery, 53A(5); pp. 963-977, July 1971.
Galante, J. et a1. Sintered Fiber Mletal Composits as a Basis for Attachment of lmplants to Bone, in J. of Bone & Joint Surgery. January, 1971, 53A(5) pp. 101-114.
Primary Examiner-Leland A. Sebastian Assistant Examiner-R. E. Schafer Attorney, Agent, or Firm-Wells, St. John & Roberts [5 7] ABSTRACT A porous metallic material having controlled patterns of interconnected voids specifically adapted for tissue ingrowth application. The material can be produced having interconnected voids of spherical or other desired shape, as well as oriented voids arranged in a pre-selected spatial pattern. The method of producing the porous structure involves the arrangement of solid expendable void former within a receiving cavity of a mold or form in a pattern corresponding to the size, shape and spatial pattern of the voids desired in the final matrix. Metallic powder formed of a biocompatible material is packed about the expendable void former and the composite material is subjected to high energy rate forming pressures to densify its structure. The expendable void former is then removed and the remaining matrix of metal is sintered to further strengthen the web structure that remains after re moval of the void former.
11 Claims, 25 Drawing Figures VlBRATORY LOADlNG OF METALLlC POWDER AROUND MICROSPHERES HE RF IMPACTED BILLET PROMOTE BONDING CUP FULLY LOADED 1N lMPACTlON CONTAlNER HEATlN'G OF BlLLET 530 C-VACUUM PATENTEL DEC 31974 SHEET 2 OF 7 FIG4 FIG 3 FIG 2 PATENTEL BEE 31974 SHEET 3 OF 7 FIG 6 FIG 5 FIG 7 PAIENILL 553 31574 FIG 8 VI IC GOAT I.I I. PIN- CQMPRESSIVE SI-IEAR (LBS) LOAD DISTAL TEST PIN ---W up 275 IJ PUSH 460 PUSH 0ST @IJT FORCE FWCE FIG 9 CYCLIC LOADING CF VMC-TISSUE INTERFACE (8 WEEKS IMPLANT) E= ELASTIC FAILURE AT (5 P: PLASTIC PEAK LOAD g 50 E P= 04103 V I: PP=QOOIO75 E P=OOO1 E E 20.0255 9 /E/ P=0.0095
550 CYCIIIS AT E 10- EACH LOAD NANCE 0 0.0m oo'oz 0.0'03 0.0M uws DISPLACEMENT (I N.)
PATENIEL BEE 31974 SHEET 5 OF 7 FIG l3 mama; M 31914 3, 529m;
SHEH B 8? 7 FIG 17 HORIZONTAL FIG 18 TWO-WAY ANGLE FIG '19 ONE WAY DOUBLE ANGLE FIG 21 VERTICAL FIG 23 IIHIHH Illllllll HORIZONTAL LOOP HORIZONTAL LOOF PATENTEL BEE 31974 SHEET 7 0F 7 FIG 24 FIG 25 VOTE) METAL COMPOSITE MATERIAL AND METHOD BACKGROUND OF THE INVENTION This invention relates to a method of producing a porous structure of sintered metallic particles adapted specifically for tissue ingrowth applications. It arises primarily from research directed to ingrowth of bone tissue for prosthetic purposes. The material is designed to fill the existing need in anatomical restoration for a reliable long-term biocompatible material for attaching artificial elements to living tissues. It has specific application to areas of bone and joint restoration and artificial tooth implantation. Porous metal materials are one solution to this attachment problem because they permit tissue invasion of the porous metallic structure. Metals, such as certain stainless steels, and alloyed or unalloyed titanium are sufficiently passivated by their naturally occurring oxide coatings to have good biocompatibility and are well-suited to sintering technology. When necessary, passivation can be intentionally enhanced by increasing the thickness of the oxide layer by special treatment processes. Stainless steel and titanium base materials are further selected for such usage because of their high strength to density ratio and easy fabrication.
Porous titanium base materials, each possessing certain properties and structure, and being fabricated by various current research groups employing unique methods of fabrication. One existing approach is to press and sinter titanium fibers in a high compliance (low stiffness) structure to produce random void concentrations between the fibers ranging between 30 and 70 percent. At 50 percent density, the pore diameters are typically a size range of 170 to 350 microns. Deformation is approximately percent at 2,000 psi. Compressed and sintered wire or fiber envelopes have also been bonded to solid cores to provide added strength in the resulting layer material.
An alternate method employed to produce porous metals has utilized flame spraying of titanium hydride powder on solid titanium base rods. The porosity of the resulting structure is irregular, with large pore sizes averaging 60 microns in diameter. The estimated density of the surface material is between 40 to 50 percent.
Various powder metallurgy techniques have also been proposed for constructing porous metals by combining and sintering metal powders of appropriate screen size to give structures of variable density and porosity. Powder size and densification parameters are varied to produce a range of pore sizes with random pore distribution. Pore size range has been reported generally between 50 to 200 microns in diameter.
SUMMARY OF THE INVENTION The disclosed invention relates to a method of making a porous structure for tissue ingrowth applications. The method involves the initial step of arranging an expendable void former within a mold cavity in the arrangement desired for the void portions in the final matrix. A metallic powder of a biocompatible material is packed about the void former to complete filling of the cavity. The composite material is next impacted to densify the composite material. The void former is then removed without disrupting the surrounding matrix of metallic material, which is subsequently sintered to further bond the metallic material particles in the web pattern that remains after removal of the void former.
According to this disclosure, a powder metallurgy process is used to create porous metals which have the unique features of reproducible and predetermined void structures and strong, biocompatible matrices. These features are of particular value in the use of such porous materials in tissue ingrowth applications such as prosthetic bone anchor devices, artificial tooth implants and other hard tissue substitutes. The resulting material herein is termed a Void Metal Composite." The particular value of the Void Metal Composite process lies in its ability to produce porous structures with controlled pore volume, pore size, and pore orientation.
The fabrication procedure used in production of Void Metal Composite materials first compacts appropriately sized powder of the chosen matrix metal by vibratory compaction about a volume: of expendable void former. The void former is arranged in the size, shape and spatial pattern of the voids desired in the final matrix. The composite material is subjected to a High Energy Rate Forming (HERE) process at moderate heat under vacuum. The densified composite material can then be further shaped by conventional machining processes. The void former is next removed. Improved bonding is achieved by a high temperature sinter. Final shaping and sizing processes that may be used include electro-discharge machining (EDM). The product can be completed by surface treatment, such as electrolytic anodization.
Void Metal Composite (VMC) materials are one of a class of materials called open cell structures. Other examples of this type of structure can be found in various foam-rubber and foam-polymers in wide use today. The open cell structure refers to the fact that the cells in the material are interconnected, or open, as opposed to other kinds of foam structures where the individual cells are enclosed by solid material.
VMC material fabrication begins by placing in a mold spheres of an easily vaporized material, such as magnesium or calcium. The diameter of these spheres will determine the pore size in the final VMC structure. A fine powder of the material or alloy that will become the final matrix is compacted by high frequency vibrations into the interstices between the spheres. After high frequency vibration, the bi-metal compact is approximately -85 percent dense. High energy rate impaction of the sphere-powder mixture at elevated temperatures produces a compact that has high density percent) and good forming characteristics. The high energy rate impaction process achieves higher densities in the compact because of the extreme pressures achieved (250 400 kpsi) in a very short time.
Following high energy rate forming, the bi-metal compact typically is machined into its final shape, and then heated in a vacuum at an elevated temperature to vaporize out the sacrificial sphere-molding material, leaving behind the spherical pores and a partially sintered meta] structure. A final high temperature vacuum treatment causes the metal structure to fully bond and achieve its best mechanical properties. Shaping of the porous metal structure may be accomplished by EDM after the sintering step.
In contrast, a typical powder metallurgy process involves cold pressing of non-spherical fine powders at room temperature under forces that are about onefourth that achieved in the high energy rate process. The density of a green compact from a cold press, powder metallurgy process is 60 70 percent compared to densities of 95 percent achieved in the high energy rate forming process.
The result of the VMC fabrication process is a porous metal structure that offers the potential of a wide range of structural modifications resulting from an extremely flexible manufacturing process. Under the close packing conditions, each large void-forming sphere theoretically touches twelve other spheres. Exclusion of the metal powder at the contact points produces openings or interconnectivity between all of the pores in the final porous metal. The size of these windows is related to the size of the pore-formers, the particle size of the fine metal powder, and the relative stiffness of the two materials at the impaction temperature. The pores can be varied in shape, e.g., from spheres to rods. In addition, the rod-shaped pores can be oriented with respect to the specimen surface. VMC material is also readily fabricated into complex shapes either by machining or by the use of a shaped mold.
It is a first object of this invention to provide a practical method of producing controllable porous metallic structures for tissue ingrowth applications.
Another object of this invention is to provide a practical method for insuring adequate interconnection between voids produced in a porous metallic matrix so as to permit ingrowth of living tissue and resulting bonding between the tissue and porous structure.
Another object of the invention is to provide a method of producing a biocompatible porous structure from metal with the ability to orient and locate the pores for proper design to match the physical requirements of implant applications.
These and further objects will be evident from the following disclosure, taken also with the accompanying drawings which illustrate the essential features of the disclosure.
DESCRIPTION OF THE DRAWINGS FIG. 1 is a schematic flow diagram illustrating the fabrication process of the present disclosure;
FIGS. 2, 3 and 4 are exterior photographs of intramedullary pins produced according to this disclosure;
FIGS. 5, 6, and 7 are enlarged photomicrographs of sections cut through materials produced according to this disclosure;
FIG. 8 is a plot of compressive shear test of a VMC- bone interface;
FIG. 9 is a plot of a cyclic loading test of a VMC- bone interface;
FIG. 10 is a schematic'perspective view of a test intramedullary pin having a solid core;
FIG. 11 is a schematic longitudinal section through the pin in FIG. 10;
FIG. 12 is a schematic illustration of the application of the pin shown in FIG. 10;
FIG. 13 is a schematic flow diagram illustrating production of a dental anchor according to this disclosure;
FIGS. 14, 15 and 16 are schematic transverse views through sample specimens having oriented pores;
FIGS. l7, l8 and 19 are schematic axial sectional views through sample specimens having oriented pores;
FIGS. 20 and 21 are schematic transverse and axial views, respectively, of a sample specimen having pores aligned parallel to the axis;
FIGS. 22 and 23 are schematic transverse and axial views respectively illustrating a sample specimen having looped transverse pores extending inwardly from its outer surface;
FIG. 24 is an elevation photograph of a test sample having pores oriented transversely; and
FIG. 25 is a photograph showing an end view of the sample in FIG. 24.
DESCRIPTION OF THE PREFERRED EMBODIMENT The present method and materials relate to fabrication of a controlled metallic matrix having predetermined interconnected voids of controlled volume. When applied to prosthetic anchoring devices or hard tissue substitutes for tissue ingrowth applications, such as root implants for tooth structures or anchors for artificical bone joints, the controlled interconnected voids provide open space in which healthy tissue can grow and remain viable in order to secure the device permanently.
Examples of biocompatible metals suitable for such application are titanium and titanium alloys, such as Ti- 6Al-4V alloy, and stainless steels, AISI as AISi 304 stainless steel. A suitable expendable void former is spherical magnesium or calcium. The sphere diameter of the void former determines the pore size of the tinished metal matrix. Other easily vaporized sacrificial materials can be used in differing structural shapes such as rods or screens, depending upon the nature of the void pattern desired in the final product.
In the Void Metal Composite (VMC) basic fabrication process, the versatility that results from coupling powder metallurgy techniques and void forming processes by use of sacrificial material, offers the opportunity to create custom designed materials uniquely suited for prosthetic use in bone.
To begin the process, the void former material is first packed within the cavity of a form or die. When using spheres, the'spheres of magnesium 'or other material are hand tamped in the form. Original density of the magnesium material in the form after packing is about 65 percent of its theoretical density. A screen is applied across the form to hold the spheres in place within its cavity and restrain the column of spherical particles against movement during later operations.
A time mesh matrix powder of the metallic material desired in the final matrix is then passed through the restraining screen and forced into the sphere interstices by variable frequency vibration. The average particle size for the matrix powder should be approximately one-sixth the diameter of the spheres or less. The vibration packing results in the density of the composite material in the form being about percent of it theoretical density.
The composite material is subsequently subjected to a high energy rate forming (I-IERF) process to densify its structure to approach the theoretical density of the composite material and from bonds between the metallic particles. This is preferably carried out following preheating of the composite material under vacuum pressure, the vacuum pressure serving to remove air and prevent oxidation or contamination of the materials when heated. When magnesium is used as the void former, the composite is heated to a temperature of approximately 530 C., which is sufficiently below the Mg melting point and the eutectic temperatures at which reactions occur between the materials of the composite. This temperature should approach the melting temperature of the void former, but cannot be so high as subsequently permit the material to rise above its melting temperature or eutectic temperature in response to its subjection to impaction. The high energy rate forming process can be accomplished in a pneumatic impaction machine commonly known by the trade name Dynapald or by alternate methods such as explosive forming. These processes result in pressure welding bonds being formed between adjacent metallic particles. The heat and high pressure of the HERF process deforms the void former material and assures surface contact between abutting particles of the void former, resulting in larger windows or openings between adjacent voids after removal of the void former as a result of the greater surface contact between spheres. The pressure also serves to impress the matrix material particles into the larger spheres, resulting in roughened interior void surfaces following removal of the void former.
After final compaction, the composite material is removed from the form by machining processes or other removal methods At this point the density of the composite material should be at least 95 percent of its theoretical density. The expendable void former is then removed by selective heating, leaching or other removal processes under conditions which do not adversely affect the matrix material. The use of heat is of particular advantage in that it initiates sinter bonding between the matrix particles. Heating can be continued upward after removal of the void former to attain a sintering temperature at which solid state diffusion bonds are created. The final sintering temperature is below the melting point of the metallic matrix material, thereby maintaining the integrity of the pores or voids created by prior removal of the expendable void former material.
Either prior to removal of the void former or after sintering of the matrix, the material can be shaped to the desired structural configuration for application purposes. If accomplished before removal of the void former, surfacing can be done by machine tools, using a lathe, milling machine, drill, cutter, grinder, electrodischarge machining apparatus or abrasive saw. After removal of the expendable void former, physical machining would probably damage the surface porosity of the matrix, Surface treatment would normally then be accomplished by electro-discharge machining processes, grinding or abrasive sawing.
Electro-discharge machined surfaces appear to be best for promotion of tissue ingrowth and fixation due to increased numers of surface pore openings and greater surface roughness, which promote earlier fixation and development of vascularization between implant surface and bone.
The final sintering process strongly bonds the porous metal particles. The spherical voids or pores are interconnected in the metallic matrix by smaller pores or windows produced at points of contact between adjacent spheres of the expendable void former. The matrix has a regular, predictable configuration of interconnected voids and surrounding metal. With pore sizes in the range of 275 to 460 microns, the size of the average interconnecting pore window or opening is about 50 to microns.
The strength and ductility of the final Void Metal Composite (VMC) structures are very sensitive to density, pore size, material type and purity, and fabrication parameters. Pore size has also been found to influence tissue ingrowth in the structure. it has previously been recognized that pore size in porous materials for such use should not be less than 100 microns for development of haversian systems or 50 microns for soft tissue ingrowth. Optimum upper limits of about 500 microns have also been reported. When using spherical pores, interconnected porosity appears essential to successful ingrowth. Elongated pores formed by void formers such as wires, rods, or screens need not have interconnectivity to sustain deep, penetrating healthy tissue.
In general, a wide range of pore sizes and shapes can be fabricated by the VMC process. Shapes include the use of spheres, cylinders, wires, perforated sheets and irregular particles as pore formers. Size ranges for pore diameter extend from 50 microns to upper size determined by need. Orientation can be carefully controlled by use of such materials as wire or perforated sheets as the void former. The void former materials can also be arranged to produce structures with metal volumes which grade from solid to porous in a controlled way.
FIG. R of the drawings diagrammatically illustrates the basic steps used in production of the VMC material structures. Beginning at the upper left-hand corner, a cylindrical form or mold in the shape of an upwardly open cup 10 is loaded wiwth microspheres 11. The microspheres 11 are composed of the void former material, with or without additional microspheres of the material chosen for the metallic matrix. The ratio of microspheres of void former material and metallic matrix material must be calculated in determining the resulting density of the final product. For a product structure of least density, all of the microspheres 11 will constitute void former material. By substituting random microspheres of the metallic material in place of those of the void former, one can increase the final density, while obviously decreasing somewhat the number of pores and number of interconnections between adjacent pores.
As can be seen in the upper left-hand corner, the microspheres it are initially retained in place by a covering circular screen disc 12. the disc 12 is placed above and in contact with the microspheres 11 to prevent then from moving upwardly during subsequent vibration.
It is preferably to securely retain the microspheres 11 in their selected positions within cup 10. This is accomplished by a perforated hold down plate 13 which is placed immediately above and abutting the disc 12. The plate 13 has apertures formed through it suffi ciently large to permit the passage of metallic powder through the plate 13 and the screen disc 12. Metallic powder 14 desired in the matrix is then loaded between and around the microspheres 11, using vibratory compaction techniques. During vibratory loading, the cup 10 and plate 13 must be securely clamped by conventional devices used with such equipment (not shown).
The fully loaded cup is placed in an impaction contallic powder and a lid 16 is welded across the top of container 18. The fully loaded container 18 is provided with an exhaust stem 17 connected to a suitable source of vacuum (not shown) to insure evacuation of the cup during heating of the material within it.
The fully loaded container or billet 18 is then heated to a temperature of about 530 c. under vacuum to prepare it for High Energy Rate Forming. lmpaction occurs along the axis of cup 10 and results in a significant reduction in its total height, as seen at the center of FIG. 1. During impaction, the cylindrical diameter of the billet is restrained from expansion, and the axial force to which it is subjected results in rapid densification of the material within it.
The impacted billet 18 is next machined to remove the cup 10 and the disc 12 and plate 13 and to produce the desired machined shape in the VMC material. Typical machined shapes are illustrated in FIG. 1 at 19 and 20.
The machined shapes 19, 20 are subsequently heated or otherwise treated to remove the void former material. Finally, the metallic matrix is sintered at a temperature adequate to bond the matrix powder without melting, thereby retaining the void areas that resulted from removal of the void former material. As a final step, when desired, the shapes 19, 20 can be coated by oxidation or other suitable processes to insure their biocompatible properties.
FIGS. 2-7 are photographs illustrating the structure produced according to this process. FIGS. 2, 3 and 4 show sample VMC intramedullary pins produced according to the above process from unalloyed titanium. The pin in FIG. 2 has an average void diameter of 275 microns. The pin in FIG. 3 has a void diameter of 460 microns and the pin in FIG. 4 has a void diameter of 650 microns. The photographs illustrate the uniformity of the spherical voids which are exposed at the outer surfaces of the pins, and the windows between adjacent voids. The uniformity of the voids and windows between adjacent voids is clearly evident in FIGS. 5, 6 and 7 which are enlarged photomicrographs (llOX) showing VMC structures having pore sizes of 275 microns and 650 microns, respectively. These photomicrographs illustrate the substantially solid nature of the interconnecting web of the metallic matrix constituting the final structure. Due to the high degree of compaction used in this process, very little void space remains within the web structure other than the designed voids left by removal of the void former.
The following Example describes actual production of a representative VMC material. Proportions and sizes of spherical and powder materials may be varied to achieve density and strength properties in the final matrix.
EXAMPLE I 72.6 grams of 50 60 mesh (250-297 11.) magnesium spheres from the Hart Metal Company and 92.4 grams of 50 60 mesh (250-297 p.) Ti-6Al4V spheres from Whittaker Nuclear Metals Division were loaded into a 2 1% inches O.D., 2 inches ID, 4 inches long stainless steel container. The spheres were throughly hand blended and the container was tapped a few times to settle and level the sphere bed which was 1 inches deep.
A 70mesh screen disc (210 u openings) was place over the sphere bed followed by a heavier perforated plate having l/l6 inch diameter holes. the container was placed in a chuck mounted vertically on a 500 pounds-force electrodynamic shaker.
61.2 grams of 325 mesh Ti-6Al-4V powder 44 pt particle size) was placed on the perforated plate. During vibratory compaction the powder went through the perforated plate and the screen to fill the void areas in the sphere bed. The vibrator frequency was continuously varied between 200 and 5,000 cps. 35 minutes were required for the loading.
The billet was removed from the chuck and the void area above the perforated plate filled with metal pow der as filter material. A stainless steel lid with an outgas stem was welded to the top of the container.
The billet was attached to a vacuum system and preheated at 530C. for 40 minutes, at a pressure l p. Hg. The billet was then transferred to the closed die tooling of a High Energy Rate Forming (HERE) type Dynapak Machine. It was impacted at 319,000 psi. The billet was ejected from the die and allowed to air cool.
The impact billet was opened by turning in a lathe. Dimensions of the densified disc were 2.382 inches dia. by 1.010 inches long. It weighed 214 g. and had a bulk density of 2.90 g/cc.
Specific shapes were then machined from the billet.
The cookout was done in a vacuum furnace at l,000 C. for 30 minutes. The magnesium was removed during the process.
The porous Ti-6Al-4V was sintered at 1350 C. for 1 hour in an Abar vacuum furnace. This heat treatment increased the grain size and bonding between particles. The resultant density was 44.8 percent leaving over 55 percent interconnected spherical voids.
Biological Evaluation of VMC Material Three VMC titanium structures, with nominal gross pore sizes of 275,460, and 650 microns and 18 percent density were selected for biological evaluation. Fifteen 2.5 X 5 X 3 mm VMC coupons, five of each pore size, were fabricated for in vitro tissue culture study, and 24 intramedullary pins of the selected pore sizes were made for acute and chronic tissue ingrowth studies in cat femora. The objective was not so much the development of devices but rather to evaluate tissue ingrowth into and performance of VMC under the actual stress and load conditions of a biological environment.
Scanning electron microscopy view of the VMC structure per se, have permitted physical study of the major pore construction and interconnecting channels in the evaluated samples. These passages theoretically number 12, but typically reveal fewer openings. Their diameters are of the order of -125 microns, an interconnectivity providing ample passageways for tissue ingrowth.
Fifteen VMC coupons, in the form of small cylindrical discs were incubated with tissue culture cells for 28 days, with culture observations being made at two day intervals. The results indicated that titanium VMC had minimal effect on cells in an in vitro environment, with no differences observed between the various pore sizes.
A total of seven cats were implanted with VMC intramedullary pins of pore sizes ranging from 275 to 650 microns, one pin in each femur. The animal were serially sacrificed at intervals of 3, 6, and 9 weeks postoperation for complete necropsy examination. Results were as follows:
l. There were no visible aftereffects, including postoperative infection in any of the surgical implantations.
2. Radiograms were taken of several of these animals just after surgery and compared to those taken prior to necropsy. No bone changes or inflammatory response attributable to the VMC intramedullary pins could be detected.
3. No gross pathologic lesions were seen in any of the tissue.
4. All I. M. pin implants consistently showed good tissue ingrowth. Grossly, after 3 weeks postimplantation, there were no distinguishable differences in ingrowth between the various pore sizes, or between the 3, 6 and 9 weeks samples. All pins appeared to show cortical bone invasion at the bone/metal interface. 5. Only one set of 650 micron pore size pins were implanted. This type of pin was discontinued due to its brittleness and lack of strength. These adverse mechanical properties resulted primarily from use of metal powder containing high impurity levels. Also, tissue ingrowth in the smaller pore sizes was equally as good as that found in the 650 micron size.
The tissue culture test coupons previously mentioned were examined at 12 and 21 days post-exposure by scanning electron microscopy. After 12 days, only a minimal tissue deposit was detectable on the interior pore surfaces. At 21 days exposure, however, all pore surfaces were covered by a thin, nearly structureless film. By comparison with the as-fabricated pretest micro-structure, it was estimated that this tissue thickness was about l-2 microns.
In contrast to the large open pores in the in vitro tissue culture tests, profuse tissue ingrowth was evident in the 21 day cat femur-pin samples. Much of the intrapore deposit was fibrous in appearance, generally spanning opposite walls of the large pores. The granular structure of the metal pin and interpore openings were largely obscured by a continuous tissue layer. At the juncture between the pin and cortical bone, the actual bone-pin interface was indistinguishable due to tissue ingrowth.
As a corollary to these observations, and to aid in the interpretation of the results, reflected light microscopy of thin ground sections of certain of these samples has been undertaken. These studies have shown conclusively that tissue invasion, bone lacunae formation, and calcification is taking place in the VMC material.
This behavior of the VMC femur pins is encouraging from the standpoint of microstructural changes and evidence of tissue ingrowth. The tissue penetrated the entire pin within 21 days, and gave indication of progress toward completely filling the pin voids. Bone-pin interfaces appeared to be well bonded, and there was no evidence of tissue incompatability.
In order to retain the desirable characteristics of pure titanium, such as inertness, biocompatibility and response oftissue ingrowth, and still increase the strength of the VMC material, a titanium alloy was subsequently evaluated. This alloy (titanium plus aluminum 6 percent and vanadium 4 percent was fabricated in small discs of 275 and 460 micron pore size and placed in a tissue culture system. These screening tests for cellular compatibility and potential toxicity demonstrated that this material was very similar to pure titanium in its ability to allow excellent cellular ingrowth with no detectable toxicity.
The new metal was further evaluated in vivo. Intramedullary pins were implanted in cat femora and the data obtained was compared with similar experiments using pure titanium VMC. For all intents and purposes, the new material was accepted by the host identically to that found in the previous work. No adverse biologic reactions were noted; yet great increases in strength were obtained.
Two additional prototypic intramedullary pins were fabricated from unalloyed titanium VMC material and implanted for 8 weeks in goat femora in a later testing program. Postmortem static shear and cyclic loading tests were conducted on these test pins. FIG. 8 shows the shear loads and interface deformation values of these short-term, two different pore size implants. Forces were seen to peak and slowly decay to push-out frictional values.
FIG. 9 shows the cyclic loading response of these implants. The top curve corresponds to cyclic loading of a specimen implanted in the femoral midshaft near endosteal cortical bone. The lower curve is associated with an implant in the trabecular bone region of the proximal end of the femur. Elastic displacement increased with increasing load, and the plastic or permanent displacement damage produced by load cycling markedly increased with increased load level.
Titanium VMC endosseous dental anchors were inserted for experimental purposes in prepared alveoli of freshly extracted mandibular second premolar teeth of Hanford Miniature Swine. Previous work in dental research established that these animals are a good standin for man in the testingof new dental materials and operative procedures. Immobilization of the implant was accomplished by splinting the anchor to the first and third premolar by means of stainless steel bonds, wire, and acrylic cement. One VMC dental anchor, was removed and studied at 6 weeks post-implanation to study tissue ingrowth into the device. The anchor, splint, and associated premolars and mandibular bone was removed in block, fixed in 10 percent neutral buffered formalin, imbedded in Maraset (Marblette Corp., Long Island, N.Y.). and. 6 to 10 micron thin ground sections and diamond saw thick sections were examined by reflected light and light microscopy.
There is some difference of opinion as to whether ankylosis or a cushioning effect from a false peridontal membrane would be the best form of fixation for dental anchor implants. Tissue type and extent of ingrowth can be controlled by the primary pore size of the material, as we and others have found out. For instance, with pore sizes of 15 to 20 microns in diameter, tissue penetration in bone is restricted to the first one or two layers of pores on the external surface of the material. There is hardly enough room in these pores for a single fibroblast or osteocyte, let along the connective tissue fibers that do penetrate the surface. At about 45 micron pore size, fibrous connective tissue will infiltrate the material to a depth of about. 20-30 microns, while at -100 micron sample will show complete penetration with connective tissue at 4 weeks, with new lamellar bone formation growing inward to a depth of approximately l00 microns. With pore sizes greater than microns, lamellar bone ingrowth will penetrate to a depth of about 300 microns in 10-12 weeks. It appears that the critical pore diameter necessary for capillary ingrowth is approximately 30-40 microns, and for lamellar bone formation about 50-75 microns. Thus, by proper design of the material, tissue type and extent of ingrowth can be controlled to conform to specific brane between the tooth and the alveolus. We believe that control of the tissue ingrowth through control of pore and window size offers another solution to the problem of stress concentration. Since the elastic modneeds uli of the VMC materials are approximately one-fourth to one-fifth the value of bone, we have substantial design choices available in stiffening this material to ob- Physlcal Evaluauon of VMC Mammal tain a favorable compliance match with bone.
In evaluating the mechanical behavior of such, the The strength of the VMC is also substantially lower porous structure produced by the VMC process, four than enamel or dentin, Table I. For comparison, wet properties are especially important: the elastic molulus, bone exhibits strength values of about 11,000 psi in the usable strength of the material, the stress at tension and 22,000 psi in compression. The 275 and which fragmentation or spallation occurs, and the duc- 650 micron materials had somewhat lower strengths tility of the material. While we have performed a few than the 460 micron material and we believe this can tensile tests of the VMC material, the majority of work 15 be traced to the iron-containing second phase dehas been done using compression tests on small cylinscribed above in the structural analyses. The strength ders. This approach was chosen to minimize the cost of Of the VMC materials noted in Table I are low with valpreparing the greater amounts of material and the more ues approximately one-tenth that of wet bone. This is complicated specimens needed for tensile testing. In not surprising because of the low gross density. addition, the compression test is similar to the shear Strength is a critical property for prosthetic materials test typically used to evaluate tissue-implant bond and generally needs to be at least as high as bone to be strength. acceptable.
Mechanical properties of various VMC titanium ma- The final mechanical property of interest is ductility. terials are shown in Table I along with comparative fig- Titanium VMC exhibits apparent good ductility with ures for dental materials. The various titanium VMC measured elongations higher than tooth materials and materials exhibited elastic moduli of 0.4 0.5 X 10 psi, higher than wet bone (-l A percent). compared to values of 1.7 and 6.7 X 10 psi for dentin In subsequent product refinement an increase in the and enamel,, respectively. Thus, the present titanium mechanical properties of VMC materials was accom- VMC is considerably more compliant than either the plished by using a strong alloy, Titanium-6Aluminum-4 tooth materials or wet bone which has a modulus of Vanadium, increasing material density, and by closer about 2.2 X 10 psi. The value for wrought titanium is attention to powder chemistry to minimize embrittling 14 X 10 psi. It should be noted that very few structural impurity content. The effects of alloying density, and materials, metals and non-metals alike, have elastic pore size on VMC material compressive properties are modulus values as low as bone. shown in Table II. The marked improvement is evident TABLE I ULTIMATE FRAGMENTA- TIoN RATE Ex I0" (psi) STRENGTH (psi) EVlDENCE %AL/L CONDITION IN./MIN. TEN.* COMP." TEN. COMP. COMP. (psi) TEN. COMP.
TOOTH 6.9 I400 l40,0()() ENAMEL+ TOOTH I.() 6500 50,000 DENTIN+ 27511. 0.002 0.4 2.I90 i430 7.6 460,2 0.002 0.3 0.5 2080 3,| 2420 2.1 2.5 650p. 0.002 0.5 2,550 2490 2.0 Ti-6Al-4V 0.002 0.4 3,140 1.97
'TENSILE' COMPRESSION REFERENCE: L. W. MORREY DDS. R. .I. NELSEN DDS. DENTAL SCIENCE HANDBOOK, U.S. GOVT. PRINTING OFFICE, I970, p.
by a more than ten-fold increase in strength with a three-fold increase in deformability AL/L) in com-' parison to similar early VMC samples of unalloyed titanium. VMC mechanical properties are now adequate to help support the bone during load transfer at the interface.
TABLE II VMC MATERIALS COMPRESSION PROPERTIES FRAG- MEN- ULTI- TATION 7cAL/L (AT MATE STRUCTURE DENSITY DEFORMA- STRENGTH EVIDENCE ULTIMATE TION MATL. TYPE 70 THEOR. RATE lN/MIN EX (psi) (psi) STRENGTH) WET CORTICAL 0.00l 2.2 21,800 1.65 BONE (FEMUR) Ti 275 18 0.002 0.4 2,l90 I430 7.6 Ti-(1-4 275 1. 50 0.002 0.5 24,600 NONE 2 l .2 Ti e-4 36011. 50 0.002 0.6 23,600 NONE l7.2 Ti-6-4 460p. 22 0.002 0.36 5,200 2.5
Increasing the density of VMC material causes loss of porosity, with more random distribution of the remaining pores, when compared to lower density structures. Microscopic studies of the 18 percent and 50 percent dense structures of 275,12 pore size analyzed in Table II showed more massive webbing of the higher density material and slight loss of structure periodicity. The high density material, however, maintained pore interconnectivity, although the number of interconnecting passages appeared to be reduced.
Centerless grinding was originally employed as a VMC shaping method, with resultant smearing of surface metal and closure of surface porosity. VMC is now shaped by Electro-Discharge Machining, which produces a much improved surface. The EDM method also produces a roughened surface that is beneficial for initial bone ingrowth, and which provides space between the bone and metal for tissue vascularization.
Although titanium and its alloys are highly passivated because of a thing naturally occurring oxide surface coat, conservatism suggests that a thicker oxide coat should be applied as additional insurance against breaking and possible metal-tissue reaction. For this reason, an electrochemical surface treatment (anodizing) is now used to produce a much thicker, tightly adherent oxide coat on our titanium alloy material. With this treatment, anodizing occurs uniformly throughout the VMC structure.
SOLID CORE STRUCTURES Another property of the VMC material under investigation is the bonding of the VMC to a solid core. The bond between a VMC coating and a solid core shaft has been tested using two specimen designs. One of the specimens had smooth interfaces while the other had threaded interfaces to increase the shear area. By applying tensile loads to the end pieces shear loading was developed in the VMC-solid metal interface. Both of these specimens failed through the VMC structure itself and not at the interface members indicating a good bond between VMC and the wrought titanium. Both specimens showed an extension of approximately 0.008 inch based on the length of the effective gage section and this amounts to percent elongation.
In many applications, a combination of solid metal and porous structure is desirable. In the design of protruding members for insertion into tissue, such as end joints for-bones or dental anchors, it might be desirable to have a protruding bearing surface or connector of solid material and an implanted area of porous material. For instance, in a prosthetic hip joint, an implanted cylinder of porous material might be affixed to a protruding spherical bearing member forming the exterior elements of the joint.
It also might be desirable to use a solid core within an implanted member for structural reinforcement. Such an arrangement is schematically illustrated in FIGS. 104.2, which shows an intramedullary pin having a solid core and cylindrical porous areas at each axial end of the pin. The pin itself is illustrated in FIGS. 10 and Ill, the core being designated by the numeral 21 and the cylindrical porous areas being designated at 22, 23. The general manner by which such a pin would be used when implanted in a femur is illustrated in FIG. 12, the posterior view of the femur being outlined at 24. The bond between the machined solid core 21, which might be constructed of any biologically compatible material. As an example, core 21 might be constructed of alloyed or unalloyed titanium to match the metallic structure of the areas 22, 23. The bond between the core 21 and areas 22, 23 is either formed during impaction of the VMC structure by positioning of the core material in the billet before compaction, or can alternately be machined and bonded to the VMC material after formation of it. The areas 22, 23, can be internally threaded to complement mating threads along the reduced extensions of core 21. The core 21 might have an exterior grooved surface to promote mechanical attachment of the porous structure if bonding is accomplished during compaction of the VMC structure. This bond is further enhanced by the cookout and sintering process steps. The solid core 21 can be formed during impaction by use of metallic powder to produce a solid volume surrounded suitably by the porous composite material described above. In addition, by suitable placement of void former material and metal powder, one can attain a gradation in density from a core area to an exterior porous area in the impacted material.
FIG. 13 schematically illustrates production ofa dental anchor according to the general process described above. A mixture of metal matrix powder and void former is subjected to high energy rate impaction to form a composite billet 25. A section machined from the billet 25 is drilled and tapped to receive a solid titanium alloy piece 26 which includes a protruding pin 27 designed for attachment of an artificial crown (not shown). The assembled anchor 28 is heated to remove the void former, and further heated to sinter the VMC material and bond the porous material to the solid metal portion of the anchor. After a final shaping operation, VMC portion 29 of the anchor is implanted beneath the .gumline for ingrowth of surrounding bone and soft tissue. The illustrated dental anchor is considered to be only one example of a practical manner of combining solid and porous materials according to this disclosure in a usable implant structure.
VARIATION OF PORE PARAMETERS Some of the major advantages of the VMC concept are the ability to directly control pore size and orientation, and to vary the modulus (stiffness) of the material. Production of a wide range of structures that are tailored to optimize stress distribution and transfer of load at the metal/tissue interface is extremely important to long range implant stability. Widening use of porous implants will require materials whose structures are customized to respond to requirements established by stress analysis of the interface zone. Direct control of the pore size, shape, orientation, and spatial distribution can be achieved by use of wire as the void former in VMC fabrication instead of magnesium spheres.
Nature has constructed bone with a composite structure whose properties are not only highly directional but sensitive to loading rate. This anisotropy and viscoelasticity is very apparent in the variation of bone properties with loading direction and rate of deformation. Evidence of natural bone modeling in response to tensile and compressive stress is seen in the configuration of the trabeculae in spongy bone whose woven network is oriented to maximize its mechanical strength. It is reasonable to conclude that orientation of the ingrowing bone trabeculae in porous materials with respect to applied load might have a marked effect on the trabeculae strength. The orientation of the porous material structure which determines trabeculae orientation might, therefore, be of considerable importance to overall attachment strength.
The most direct method for forming oriented porosity is the use of wire as an expendable void former. Removal of the wire will leave uniform, cylindrically shaped pores without sharp corners and notches which create stress risers. Pore length can be of any value, and, most importantly, the wire direction can be con trolled to give a pre-selected pore orientation. Some of the many possibilities are seen in FIGS. 14-23. When it is realized that many combinations of grouping and orientation with the structure axis can be done, the concept of controlled pore orientation becomes a valuable tool in providing custom designed materials to match bone structure.
A typical cylindrical pore size for oriented voids is in the 275 460 microns range where we already have preliminary spherical pore data and experience.
The horizontal design (FIG. 17) is fabricated-by assembling a stack of screens or wires in an impaction container followed by vibratory compaction of Ti-6Al- 4V powder in the void areas of the assembly. All of the screens or wires are to be oriented the same so that relatively large vertical webs of Ti-6Al-4V are obtained. The billet should be preheated to approximately 700C and impacted at approximately 300,000 psi.
The double angle design (FIG. 19) requires a shaping process for the screen or wire material. The square screen pattern is somewhat distorted; however, orientation is maintained to permit relatively large, strong webs of Ti-6Al-4V to be formed.
FIGS. 14-23 graphically illustrate four arrangements that might be used in a typical sample specimen of the VMC material. The transverse patterns illustrated in FIGS. 14, 15 and 16 are usable in any of the axial patterns illustrated in FIGS. l7, l8 and 19. For instance, the radial pattern of wire or rod formers in FIG. 14 can be arranged either in a horizontal or transverse pattern (FIG. 17), and angular or oblique pattern (FIG. 18) or in a double angle or conical pattern (FIG. 19). A twoway pattern as shown in FIG. 15 can be produced by use of screens or by alternate placement of parallel wires or rods in adjacent axial layers.
FIG. 20 illustrates a VMC structure having vertical or axial pores parallel to one another. FIGS. 22 and 23 illustrate exterior looped pores open to the exterior surface of the sample and terminating short of its interior surface.
- A more detailed description of the VMC fabrication process employing wire as the void forming material is given below. Use of wire offers one method for producing uniform size pores with close directional control.
Screen material is selected for its ability to hold shape during compaction, to not alloy readily with the powder matrix, and to be removed easily from the matrix by vaporization or chemical reaction. Screen materials of choice are magnesium, brass, or copper, but other materials may be used if suitable. Lower practical limits of screen size are in the mesh range, determined primarily by the relative ease of vibratory compacting 325 mesh powder through and around layers of screens.
EXAMPLE 2 A VMC structure was fabricated using 60 mesh brass screen as the expendable void former. The basic process described above in Example I was again followed. The screen was cut into 1 inch diameter discs and stacked 2 inches high (approx. discs) in a compaction can, each layer rotated 45. Fine Titanium 6 Aluminum 4 Vanadium powder (325 mesh) was vibratory compacted around the column of screens and heated to 700C under dynamic vacuum. The material under vacuum was compacted at this temperature by high energy rate forming to near-theoretical density. Impaction pressures were approx. 300,000 psi. Screen was removed by reaction in warm I-INO Final bonding was done by vacuum sinter at l350C.
The resulting structure was approximately 50 percent dense with pore size of 180 microns. VMC webbing was approximately 250 microns in diameter. The cylindrical pores were formed to be oriented at 90 to the long axis with good interconnectivity in the plane of each disc but less in the connectivity between layers. Mechanical properties in compression of this prototype structure were as follows:
Ultimate Compressive Strength 44,465 psi Deformation to Failure 8.4% Modulus of Elasticity 6 X 10 psi FIGS. 24 and 25 show elevation and end photographs of a sample of VMC material produced by use of an axial stack of parallel screen void formers as described above in Example 2. The axial layering of the resulting voids is clearly evident in FIG. 24 and the transverse interconnections produced by removal of the screen former material in the various layers is clearly shown in FIG. 25. The ability to control the size, shape, spatial positioning and orientation of the pores in the disclosed VMC material provides almost unlimited design ability in matching the pore structure to the strength and directional properties of the tissue which is to be either replaced or reinforced by the material.
Further changes are obviously available to those utilizing the concepts set out above. The specific limitations set out herein are not intended as limitations, but are examples of the current practical status of this development.
Having thus disclosed our invention, we claim:
1. A method of making a porous matrix structure having voids therein for tissue ingrowth applications,
arranging a plurality of elements of an expendable void former of a solid substance within a form cavity, the elements being arranged within the form cavity in a pattern corresponding to the size, shape and spatial pattern of the voids desired in the final matrix;
packing a fine powder of biocompatible metallic particles into the form cavity about the elements of the expendable void former to complete filling of the form cavity with a composite material comprising the expendable void former and metallic particles;
subjecting the composite material to a high energy rate forming process to cause the density of the resulting densified composite material to approach its theoretical density and form a densified matrix of metallic particles having initial bonds between the metallic particles;
removing the expendable void former without disruption of the densified matrix of metallic particles;
and subsequently sintering the matrix of metallic particles so as to cause the metallic particles to further bond and thereby form a solid metal matrix about the voids left by removal of the expendable void former.
2. A method for constructing a high strength biocompatible bone implant element having interconnected voids of a desired size and pattern to enable live bone tissue to progressively grow into and through the voids to integrally interconnect the implant element with a bone structure comprising:
arranging a plurality of solid expendable void former elements of a selected size within a form cavity and in a selected spatial pattern corresponding to the desired size and pattern of the interconnected voids;
packing a fine powder of biocompatible metallic particles into the form cavity about the void former elements to fill the form cavity with a composite ma:
terial of the expendable void former and metallic particles;
subjecting the composite material to a high energy rate forming process to cause the density of the resulting densified composite material to approach the theoretical density and form a densified metallic structure of metallic particles having initial bonds between the metallic particles;
removing the expendable void former elements without disruption of the densified metallic structure of metallic particles to form the interconnected voids of the desired size and pattern within the densified metallic structure suitable for enabling the live bone tissue to grow into and through the voids in the densified metallic structure; and
subsequently sintering the densified metallic structure so as to cause the metallic particles to further bond to form the bone implant element.
3. The method set out in claim 2 wherein removal of the expendable void former is accomplished by heating the densified composite material to a temperature above the vaporizing temperature of the expendable void former and below the melting point of the matrix of metallic particles under vacuum for a time sufficient to remove the expendable void former by vaporization.
4. A method as set out in claim 2 further comprising the step of machining the densified composite material prior to removal of the void former.
5. A method as set out in claim 2 further comprising the step of machining the metallic material after the sintering step.
6. A method as set out in claim 2 wherein the expendable void former comprises solid spheres of preselected diameter.
7. A method as set out in claim .2 wherein the expendable void former comprises an o-pemmesh screen.
8. A method as set out in claim 2 wherein the biocompatible metallic material is titanium, titanium alloy, or stainless steel.
9. A method as set out in claim 2 wherein the biocompatible metallic material is titanium, Ti-6Al-4V alloy or 300 series stainless steel.
10. A method as set out in claim 2 wherein the expendable void former is magnesium, brass or copper.
lll. A biocompatible bone implant element having a high strength densified sintered metallic structure with a network of interconnected voids of a desired size and pattern formed therein to enable live bone tissue to progressively grow into and through the interconnected voids to integrally interconnect the implant element and a bone structure;
in which the implant element is formed by (a) arranging a plurality of solid expendable void former elements of a selected size in a selected spatial pattern within a form cavity in which the selected size and spatial pattern of the former elements corresponds to the desired size and pattern of the interconnected voids; (b) packing a fine powder of biocompatible metallic particles into the form cavity about the void former elements to fill the form cavity with a composite material of the expendable void former and metallic particles; (c) subjecting the composite material to a high energy rate forming process to cause the density of the resulting densified composite material to approach the theoretical density and form a densified metallic structure of metallic particles having initial bonds between the metallic particles; ((1) removing the expendable void former elements without disruption of the densified metallic structure of metallic particles to form the interconnected voids of the desired size and pattern within the densified metallic structure; and (e) subsequently sintering the densified structure to further bond the metallic particles together to form the high strength sintered structure.
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|U.S. Classification||428/566, 428/553, 606/76, 419/2, 428/567, 623/23.61|
|International Classification||B22F5/10, B22F3/11, A61F2/28, A61F2/30|
|Cooperative Classification||A61F2002/30978, B22F2998/00, A61F2/28, B22F3/1134, B22F5/10, A61F2002/30968|
|European Classification||A61F2/28, B22F3/11D4, B22F5/10|