U.S. patent number 3,852,045 [Application Number 05/280,266] was granted by the patent office on 1974-12-03 for void metal composite material and method.
This patent grant is currently assigned to Battelle Memorial Institute. Invention is credited to Manuel T. Karagianes, Kenneth R. Sump, Kenneth R. Wheeler.
United States Patent |
3,852,045 |
Wheeler , et al. |
December 3, 1974 |
VOID METAL COMPOSITE MATERIAL AND METHOD
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 removal of
the void former.
Inventors: |
Wheeler; Kenneth R. (Richland,
WA), Sump; Kenneth R. (Kennewick, WA), Karagianes; Manuel
T. (Richland, WA) |
Assignee: |
Battelle Memorial Institute
(Richland, WA)
|
Family
ID: |
23072352 |
Appl.
No.: |
05/280,266 |
Filed: |
August 14, 1972 |
Current U.S.
Class: |
428/566; 419/2;
428/567; 623/23.61; 428/553; 606/76 |
Current CPC
Class: |
B22F
3/1134 (20130101); A61F 2/28 (20130101); B22F
5/10 (20130101); B22F 7/002 (20130101); B22F
3/1112 (20130101); B22F 2998/00 (20130101); A61F
2002/30968 (20130101); B22F 2998/00 (20130101); Y10T
428/12153 (20150115); Y10T 428/1216 (20150115); Y10T
428/12063 (20150115); A61F 2002/30978 (20130101) |
Current International
Class: |
A61F
2/28 (20060101); B22F 5/10 (20060101); B22F
3/11 (20060101); A61F 2/30 (20060101); C22c
001/08 (); C22c 001/04 (); B22f 003/16 (); A61c
013/30 (); A61f 001/24 () |
Field of
Search: |
;75/222,200,214 ;29/182
;128/92C ;3/1 |
References Cited
[Referenced By]
U.S. Patent Documents
Foreign Patent Documents
|
|
|
|
|
|
|
527,192 |
|
Jul 1956 |
|
CA |
|
898,140 |
|
Jun 1962 |
|
GB |
|
784,124 |
|
Oct 1957 |
|
GB |
|
717,034 |
|
Oct 1954 |
|
GB |
|
Other References
Schwarzkopf, P. Controlled Porosity of Microstructures by Powder
Metallurgy, in Int. J. Powder Met., 2(4); pp. 3-11, 1966. .
Welsh, R. P. et al. Surgical Implants in J. of Bone & Joint
Surgery, 53A(5); pp. 963-977, July 1971. .
Galante, J. et al. Sintered Fiber Metal Composits as a Basis for
Attachment of Implants to Bone, in J. of Bone & Joint Surgery.
January, 1971, 53A(5) pp. 101-114..
|
Primary Examiner: Sebastian; Leland A.
Assistant Examiner: Schafer; R. E.
Attorney, Agent or Firm: Wells, St. John & Roberts
Claims
Having thus disclosed our invention, we claim:
1. A method of making a porous matrix structure having voids
therein for tissue ingrowth applications, comprising:
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 material 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 open-mesh 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.
11. 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; (d) 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.
Description
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 10
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
(HERF) 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 70-85 percent
dense. High energy rate impaction of the sphere-powder mixture at
elevated temperatures produces a compact that has high density
(>95 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 metal 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 one-fourth 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. 17, 18 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
finished 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 fine 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 80 percent of it theoretical density.
The composite material is subsequently subjected to a high energy
rate forming (HERF) 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.degree. 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 "Dynapak" 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,
electro-discharge 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 100 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. 1 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 11
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
sufficiently 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 container 18 shown
in the upper right-hand corner of FIG. 1. The excess volume above
plate 13 is filled with metallic powder 15 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.degree. c. under vacuum to prepare it for
High Energy Rate Forming. Impaction 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 (110X) 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 1
72.6 grams of -50 + 60 mesh (250-297 .mu.) magnesium spheres from
the Hart Metal Company and 92.4 grams of -50 + 60 mesh (250-297
.mu.) Ti-6Al-4V spheres from Whittaker Nuclear Metals Division were
loaded into a 2 1/2 inches O.D., 2 3/8 inches I.D., 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 3/8 inches deep.
A 70mesh screen disc (210 .mu. openings) was placed over the sphere
bed followed by a heavier perforated plate having 1/16 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-6A1-4V powder (<44 .mu. 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 powder 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
530.degree.C. for 40 minutes, at a pressure <1 .mu. Hg. The
billet was then transferred to the closed die tooling of a High
Energy Rate Forming (HERF) 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 1,000.degree. C. for 30
minutes. The magnesium was removed during the process.
The porous Ti-6Al-4V was sintered at 1350.degree. 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 .times. 5 .times. 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 100-125 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:
1. 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 1-2 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 of tissue
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 stand-in for man in the testing of 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 75-100 micron
sample will show complete penetration with connective tissue at 4
weeks, with new lamellar bone formation growing inward to a depth
of approximately 100 microns. With pore sizes greater than 100
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 needs.
Physical Evaluation of VMC Material
In evaluating the mechanical behavior of such, the porous structure
produced by the VMC process, four properties are especially
important: the elastic molulus, the "usable" strength of the
material, the stress at which fragmentation or spallation occurs,
and the ductility of the material. While we have performed a few
tensile tests of the VMC material, the majority of work has been
done using compression tests on small cylinders. This approach was
chosen to minimize the cost of preparing the greater amounts of
material and the more complicated specimens needed for tensile
testing. In addition, the compression test is similar to the shear
test typically used to evaluate tissue-implant bond strength.
Mechanical properties of various VMC titanium materials are shown
in Table 1 along with comparative figures for dental materials. The
various titanium VMC materials exhibited elastic moduli of 0.4 -
0.5 .times. 10.sup.6 psi, compared to values of 1.7 and 6.7 .times.
10.sup.6 psi for dentin and enamel,, respectively. Thus, the
present titanium VMC is considerably more compliant than either the
tooth materials or wet bone which has a modulus of about 2.2
.times. 10.sup.6 psi. The value for wrought titanium is 14 .times.
10.sup.6 psi. It should be noted that very few structural
materials, metals and non-metals alike, have elastic modulus values
as low as bone.
TABLE I
__________________________________________________________________________
ULTIMATE FRAGMENTATION RATE EX 10.sup.6 (psi) STRENGTH (psi)
EVIDENCE %.DELTA.L/L CONDITION IN./MIN. TEN.* COMP.** TEN. COMP.
COMP. (psi) TEN. COMP.
__________________________________________________________________________
TOOTH ENAMEL+ -- 6.9 -- 1400 140,000 -- -- -- TOOTH DENTIN+ -- 1.6
-- 6500 50,000 -- -- -- 275.mu. 0.002 -- 0.4 -- 2,190 1430 -- 7.6
460.mu. 0.002 0.3 0.5 2080 3,150 2420 2.1 2.5 650.mu. 0.002 -- 0.5
-- 2,550 2490 -- 2.0 Ti-6A1-4V 0.002 -- 0.4 -- 3,140 -- -- 1.97
(460.mu.)
__________________________________________________________________________
*TENSILE ** COMPRESSION + REFERENCE: L. W. MORREY DDS, R. J. NELSEN
DDS, DENTAL SCIENCE HANDBOOK, U.S. GOVT. PRINTING OFFICE, 1970, p.
38
The elastic modulus is important because it is one of the primary
factors controlling stress concentrations where mechanical forces
must be transferred from one material to another. Nature controls
these stress concentrations in teeth by the geometry of the tooth
root and by insertion of the highly elastic periodontal membrane
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 moduli 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 obtain a favorable
compliance match with bone.
The strength of the VMC is also substantially lower than enamel or
dentin, Table I. For comparison, wet bone exhibits strength values
of about 11,000 psi in tension and 22,000 psi in compression. The
275 and 650 micron materials had somewhat lower strengths than the
460 micron material and we believe this can be traced to the
iron-containing second phase described above in the structural
analyses. The strength of the VMC materials noted in Table I are
low with values approximately one-tenth that of wet bone. This is
not surprising because of the low gross density. Strength is a
critical property for prosthetic materials and generally needs to
be at least as high as bone to be acceptable.
The final mechanical property of interest is ductility. Titanium
VMC exhibits apparent good ductility with measured elongations
higher than tooth materials and higher than wet bone (.about.1 1/2
percent).
In subsequent product refinement an increase in the mechanical
properties of VMC materials was accomplished by using a strong
alloy, Titanium-6Aluminum-4 Vanadium, increasing material density,
and by closer attention to powder chemistry to minimize embrittling
impurity content. The effects of alloying density, and pore size on
VMC material compressive properties are shown in Table II. The
marked improvement is evident by a more than ten-fold increase in
strength with a three-fold increase in deformability (% .DELTA.L/L)
in comparison 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
__________________________________________________________________________
FRAGMEN- ULTIMATE TATION %.DELTA.L/L (AT STRUCTURE DENSITY
DEFORMATION STRENGTH EVIDENCE ULTIMATE MATL. TYPE % THEOR. RATE
IN/MIN EX 10.sup.6 (psi) (psi) STRENGTH)
__________________________________________________________________________
WET CORTICAL -- 0.001 2.2 21,800 -- 1.65 BONE (FEMUR) Ti 275.mu. 18
0.002 0.4 2,190 1430 7.6 Ti-6-4 275.mu. 50 0.002 0.5 24,600 NONE
21.2 Ti-6-4 360.mu. 50 0.002 0.6 23,600 NONE 17.2 Ti-6-4 460.mu. 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.mu. 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 15
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. 10-12, 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 11, 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 of a 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 controlled
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 700.degree. C 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. 17, 18 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 two-way 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 100 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 1 was again followed. The screen was cut into 1 5/8 inch
diameter discs and stacked 2 inches high (approx. 130 discs) in a
compaction can, each layer rotated 45.degree.. Fine Titanium -- 6
Aluminum - 4 Vanadium powder (325 mesh) was vibratory compacted
around the column of screens and heated to 700.degree. C 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 HNO.sub.3. Final bonding was done by
vacuum sinter at 1350.degree.C.
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.degree. 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 .times. 10.sup.5 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.
* * * * *