U.S. patent application number 11/491703 was filed with the patent office on 2006-11-16 for radiovisible hydrogel intervertebral disc nucleus.
This patent application is currently assigned to Howmedica Osteonics Corp.. Invention is credited to Christopher DeMaria, Paul Higham, Chau Ngo, Philip F. III Williams.
Application Number | 20060255503 11/491703 |
Document ID | / |
Family ID | 31991880 |
Filed Date | 2006-11-16 |
United States Patent
Application |
20060255503 |
Kind Code |
A1 |
Higham; Paul ; et
al. |
November 16, 2006 |
Radiovisible hydrogel intervertebral disc nucleus
Abstract
A spinal implant for replacing the natural nucleus of the disc
made from a polymer such as hydrogel having a radiopaque material
located within the polymer. The material may be in the form of a
powder dispersed throughout the polymer or may be in he form of a
powder dispersed in layers or in other specific locations within
the polymer. The radiopaque material is metal such as gold,
tungsten, titanium, tantalum or platinum. The metal may also be in
the form of a foil or wire located within the hydrogel such as
polyurethane, thereby making the implant visible on x-rays. Other
polymers besides hydrogel may be used with the radiopaque material
being dispersed therein.
Inventors: |
Higham; Paul; (Ringwood,
NJ) ; Ngo; Chau; (Secaucus, NJ) ; DeMaria;
Christopher; (Glen Rock, NJ) ; Williams; Philip F.
III; (Teaneck, NJ) |
Correspondence
Address: |
LERNER, DAVID, LITTENBERG,;KRUMHOLZ & MENTLIK
600 SOUTH AVENUE WEST
WESTFIELD
NJ
07090
US
|
Assignee: |
Howmedica Osteonics Corp.
Mahwah
NJ
|
Family ID: |
31991880 |
Appl. No.: |
11/491703 |
Filed: |
July 24, 2006 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
10244306 |
Sep 16, 2002 |
|
|
|
11491703 |
Jul 24, 2006 |
|
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|
Current U.S.
Class: |
264/255 |
Current CPC
Class: |
A61F 2002/3008 20130101;
A61F 2310/00149 20130101; A61L 27/52 20130101; A61L 2430/38
20130101; A61F 2250/0064 20130101; A61L 27/50 20130101; A61F 2/441
20130101; A61F 2310/00131 20130101; A61F 2002/30616 20130101; A61F
2310/00137 20130101; A61F 2310/00023 20130101; A61F 2/442 20130101;
A61F 2002/30971 20130101; A61F 2002/444 20130101; A61F 2/30965
20130101; A61F 2310/00155 20130101; A61F 2250/0098 20130101 |
Class at
Publication: |
264/255 |
International
Class: |
B29C 41/22 20060101
B29C041/22 |
Claims
1. A method for forming a radiopaque polymeric spinal nucleus
implant, comprising: obtaining a polymer in a liquid phase (e.g.
solution, melt, or pre-polymer); mixing radiopaque particles with
said polymer liquid phase; and placing a first layer of the polymer
solution in a mold, cooling the mold to solidify the layer, placing
a first layer of said mixed polymer solution and radiopaque
particles on top of the cooled first layer of polyvinyl and
thereafter cooling the layer containing the radiopaque metal to
form a layered implant.
2. The method as set forth in claim 1, wherein said radiopaque
material is a metal.
3. The method as set forth in claim 2, wherein the metal is
selected from the group consisting of gold, tungsten, tantalum,
platinum, titanium and a combination thereof.
4. The method as set forth in claim 3, wherein the polymer is a
hydrogel.
5. The method as set forth in claim 4, wherein the solution is a
polyvinyl alcohol solution.
6. The method as set forth in claim 5, wherein a second layer of
the polyvinyl alcohol solution is placed on top of said first layer
of mixed polymer solution, the second layer pf polyvinyl alcohol
solution cooled and a second layer of mixed polyvinyl alcohol and
radiopaque particles is placed on the cooled second layer of
polyvinyl alcohol solution and thereafter cooling the second mixed
layer.
7. The method as set forth in claim 1, wherein the polymer is a
polyurethane.
8. The method as set forth in claim 1 further comprising placing a
second layer of polymer solution in the mold on said first layer of
mixed polymer and radiopaque particles, cooling said second polymer
solution layer to solidify the same and placing a second layer of
said mixed polymer and radiopaque particles solution on said second
layer of polymer solution.
9. The method as set forth in claim 3 wherein the particle has a
diameter of between 10 and 100 .mu.m.
10. The method as set forth in claim 9, wherein the particle has a
minimum diameter of about 75 .mu.m.
11. The method as set forth in claim 1, wherein the polymer
solution is selected from the group consisting of a hydrogel, a
cross-linked protein and a polyurethane.
12. A method for forming a radiopaque polymeric spinal nucleus
implant, comprising: obtaining a polymer in a liquid phase (e.g.
solution, melt, or pre-polymer); mixing radiopaque particles with
said polymer liquid phase; and placing a first layer of the polymer
solution in a mold, freezing the mold to solidify the layer,
placing a first layer of said mixed polymer solution and radiopaque
particles on top of the solidified first layer of polyvinyl and
thereafter freezing the layer containing the radiopaque metal to
form a layered implant.
Description
[0001] This application is a divisional of U.S. application Ser.
No. 10/244,306, filed on Sep. 16, 2002, the disclosure of which is
incorporated herein by reference.
BACKGROUND OF THE INVENTION
[0002] This invention relates to a prosthetic intervertebral disc
nucleus. More particularly, it relates to an artificial disc
nucleus made of a hydrogel material having a radiovisible material
therein.
[0003] The intervertebral disc is a complex joint anatomically and
functionally. It is composed of three component structures; the
nucleus pulposus (the nucleus), the annulus fibrosus (the annulus)
and the vertebral end-plates. The biochemical composition and
anatomical arrangements within these component structures are
related to the biomechanical function of the disc.
[0004] The nucleus occupies about 25-40% of the total disc
cross-sectional area. It is primarily composed of mucoid material
containing mainly proteoglycans with a small amount of collagen.
The proteoglycans consist of a protein core with chains of
negatively charges keratin sulphate and chondroitin sulphate
covalently attached thereto. Due to these constituents, the nucleus
is a loose hydrogel which usually contains about 70-90% water by
weight. Although the nucleus plays an important role in the
biomechanical function of the disc, the mechanical properties of
the disc are not well known, largely because of the loose hydrogel
nature of the nucleus.
[0005] As the nucleus is surrounded by the annulus and vertebral
end-plates and the negatively charged sulphate groups are
immobilized due to the attachment of these groups to the polymer
matrix, the matrix has a higher concentration of counter ions than
its surroundings. This ion concentration results in a higher
osmotic pressure than the annulus e.g., ranging from about 0.1 to
about 0.3 MPa. As a result of the high fixed charge density of the
proteoglycan the matrix exerts an osmotic swelling pressure which
can support an applied load in much the same way as air pressure in
a tire supports the weight of a car.
[0006] It is the osmotic swelling pressure and hydrophilicity of
the nucleus matrix that offers the nucleus the capability of
imbibing fluid until it is balanced with the internal resistance
stresses, due to the tensile forces of the collagen network, and
the external stresses due to the loads that are applied by muscle
and ligament tension. The swelling pressure (Ps) of the nucleus is
directly dependent on the concentration and fixed charge densities
of proteoglycan, i.e., the higher the concentration and fixed
charge densities of proteoglycan the higher will be the swelling
pressure of the nucleus. The external pressure changes with
posture. When the human body is supine the compressive load on the
third lumbar disc is 300 newton (N) which rises to 700 N when an
upright stance is assumed. The compressive load increases, yet
again, to 1200 N when the body is bent forward by only 20.degree.
C. When the external pressure (Pa) increases the previous balance,
i.e. Ps=Pa, is upset. To reach a new balance, the swelling pressure
has to increase. This increase is achieved by increasing the
proteoglycan concentration in the nucleus which is achieved by
reducing the fluid in the nucleus. That is why discs lose about 10%
of their height, as a result of creep, during the daytime. When the
external load is released i.e., Ps is greater than Pa, the nucleus
will imbibe fluid from its surroundings in order to reach the new
equilibrium value. It is this property of the nucleus that is
mainly responsible for the compressive properties of the disc.
[0007] The annulus forms the outer limiting boundary of the disc.
It is composed of highly structured collagen fibers embedded in an
amorphous base substance which is also composed of water and
proteoglycans. The amount of proteoglycans is lower in the annulus
than in the nucleus. The collagen fibers of the annulus are
arranged in concentric laminated bands or lamella, (about 8-12
layers thick) with a thicker anterior wall and thinner posterior
wall. In each lamella, the fibers are parallel and attached to the
superior and inferior vertebral bodies at an angle of about
30.degree. form the horizontal plane of the disc in both
directions. This design particularly resists twisting because the
half of the fibers cocked in one direction will tighten as the
vertebrae rotate relative to each other in the other direction. The
composition of the annulus along the radial axis is not uniform.
There is a steady increase in the proportion of collagen from the
inner to the outer sections of the annulus. This difference in
composition may reflect the need of the inner and outer regions of
the annulus to blend into very different tissues while maintaining
the strength of the structure. Only the inner lamellae are anchored
to the end-plates forming an enclosed vessel for the nucleus. The
collagen network of the annulus restrains the tendency of the
nucleus gel to absorb water from surrounding tissues and swell.
Thus, the collagen fibers in the annulus are always in tension, and
the nucleus gel is always in compression.
[0008] The two vertebral end-plates are composed of hyaline
cartilage, which is a clear, "glassy" tissue, that separates the
disc from the adjacent vertebral bodies. This layer acts as a
transitional zone between the hard, bony vertebral bodies and the
soft disc. Because the intervertebral disc is avascular, most
nutrients that the disc needs for metabolism are transported to the
disc by diffusion through the end-plate area.
[0009] The intervertebral joint exhibits both elastic and viscous
behavior. Hence, during the application of a load to the disc there
will be an immediate "distortion" or "deformation" of the disc,
often referred to as "instantaneous deformation." It has been
reported that the major pathway by which water is lost, from the
disc during compression, is through the cartilage end-plates. Since
the water permeability of the end-plates is in the range of about
0.20 to about 0.85.times.10.sup.-17 m.sup.4N.sup.-1 sec.sup.-1 it
is reasonable to assume that under loading, the initial volume of
the disc is constant while the load is applied. Because the natural
nucleus of the disc is in the form of a loose hydrogel, i.e., a
hydrophilic polymeric material which is insoluble in water, it can
be deformed easily, the extent of deformation of the disc being
largely dependent on the extensibility of the annulus. It is
generally believed that hydrostatic behavior of the nucleus plays
an important role in the normal static and dynamic load-sharing
capability of the disc and the restoring force of the stretched
fibers of the annulus balances the effects of the nucleus swelling
pressure. Without the constraint by the annulus, annular bulging of
the nucleus would increase considerably. If the load is maintained
at a constant level, a gradual change in joint height, commonly
referred to as "creep" will occur as a function of time.
Eventually, the creep will stabilized and the joint is said to be
in "equilibrium." When the load is removed the joint will gradually
"recover" to its original height before loading. The creep and
relation rates depend on the amount of load applied, the
permeability of the end-plates and the water binding capability of
the nucleus hydrogel. Creep and relaxation are essential processes
in pumping fluid in and out of the disc.
[0010] Degeneration of the intervertebral disc is believed to be a
common cause of final pathological changes and back pain. As the
intervertebral disc ages it undergoes degeneration. The changes
that occur are such that, in many respects, the composition of the
nucleus seems to approach that of the inner annulus. Intervertebral
disc degeneration is, at least in part, the consequence of
compositional changes in the nucleus. It has been found that both
the molecular weight and the amount of proteoglycans in the nucleus
decrease with age, especially in degenerated discs, and the ratio
of keratin sulphate to chondroitin sulphate in the nucleus
increases. This increase in the ratio of keratin sulphate to
chondroitin sulphate and decrease in proteoglycan content decreases
the fixed charge density of the nucleus from about 0.28 meq/ml to
about 0.18-0.20 meq/ml. These changes cause the nucleus to lose
part of its water binding capability which decreases the maximum
swelling pressure it can exert. As a result, the maximum water
content drops from over about 85%, in preadolescence, to about
70-75% in middle age. The glycosaminoglycan content of prolapsed
discs has been found to be lower, and the collagen content higher,
than that of normal discs of a comparable age. Discs L-4-L-5 and
L-5-S-1 are usually the most degenerated discs.
[0011] It is known that although the nucleus only occupies about
one third of the total disc area, it takes about 70% of the total
loading in a normal disc. Thus, it has been found that the
compressive load on the nuclei of moderately degenerated discs is
about 30% lower than in comparable normal discs but the compressive
load on the annulus increases by 100% in the degenerated discs.
This load change is primarily caused by the structural changes in
the disc as discussed above. The excess load on the annulus, of the
degenerated disc, causes reduction of the disc height and excessive
movement of the spinal segments. The flexibility of the disc
produces excessive movement of the collagenous fibers which in
turn, injures the fiber attachments and causes delamination of the
well organized fibers of the annulus ring. The delamination annulus
can be further weakened by stress on the annulus and in severe
cases this stress will cause tearing of the annulus. This whole
process is very similar to driving on a flat tire, where the
reinforcement layer will eventually delaminate. Because the
thickness of the annulus is not uniform, with the posterior
portions being thinner than the anterior portions, delamination and
lesions usually occur in the posterior area first.
[0012] The spinal disc may also be displaced or damaged due to
trauma or diseases. In these cases, and in the case of disc
degeneration, the nucleus may herniate and/or protrude into the
vertebral canal or intervertebral foramen, in which case it is
known as a herniated or "slipped" disc. This disc may in turn press
upon the spinal nerve that exits the vertebral canal through the
partially obstructed foramen, causing pain or paralysis in the area
of its distribution. The most frequent site of occurrence of a
herniated disc is in the lower lumbar region. A disc herniation in
this area often involves the inferior extremities by compressing
the sciatic nerve.
[0013] There are basically three types of treatment currently being
used for treating low back pain caused by injured or degenerated
discs: conservative care, discectomy and fusion. Each of these
treatments has its advantages and limitations. The vast majority of
patients with low back pain, especially those with first time
episodes of low back pain, will get better with conservative
treatment. However, it is not necessarily true that conservative
care is the most efficient and economical way to solve the low back
pain problem.
[0014] Discectomy usually provides excellent short term results in
relieving the clinical symptoms, by removing the herniated disc
material, usually the nucleus, which causes the low back pain
either by compressing the spinal nerve or by chemical irritation.
Clearly, a discectomy is not desirable from a biomechanical point
of view. In a healthy disc, the nucleus takes the most
compressional load and in a degenerated disc this load is primarily
distributed onto the annulus ring which, as described above, causes
tearing and delamination of the annulus. Removal of the nucleus in
a discectomy actually causes distribution the compressive load onto
the annulus ring thereby narrowing the disc spaces. It has been
reported that a long-term disc height decrease might be expected to
cause irreversible osteoarthritis-like changed in the facet joint.
That is why discectomy yields poor long term benefits and results
in a high incidence of reherniation.
[0015] Fusion generally does a good job in eliminating symptoms and
stabilizing the joint. However, because the motion of the fused
segment is restricted, the range of motion of the adjoining
vertebral discs is increased possibly enhancing their degenerative
processes.
[0016] Because of these disadvantages, it is desirable to use a
prosthetic joint device which not only is able to replace the
injured or degenerated intervertebral disc, but also can mimic the
physiological and the biomechanical function of the replaced disc
and prevent further degeneration of the surrounding tissue.
[0017] Artificial discs are well known in the prior art. U.S. Pat.
No. 3,867,728, to Stubstad et al., relates to a device which
replaces the entire disc. This device is made by laminating
vertical, horizontal or axial sheets of elastic polymer. U.S. Pat.
No. 4,309,777, to Patil, relates to a prosthetic utilizing metal
springs and cups. A spinal implant comprising a rigid solid body
having a porous coating on part of its surface is shown in Kenna's
U.S. Pat. No. 4,714,469. An intervertebral disc prosthetic
consisting of a pair of rigid plugs to replace the degenerated disc
is referred by Kuntz, U.S. Pat. No. 4,349,921. U.S. Pat. Nos.
4,772,287 and 4,904,260 to Ray et al., teach the use of a pair of
cylindrical prosthetic intervertebral disc capsules with or without
therapeutical agents. U.S. Pat. No. 4,911,718 to Lee et al.,
relates to an elastomeric disc spacer comprising three different
parts; nucleus, annulus and end-plates, of different materials. At
the present time, none of these inventions has become a product in
the spinal care market. Bao et al., in U.S. Pat. Nos. 5,047,055 and
5,192,326 (assigned to the assignee of this invention and
incorporated herein by reference) describe artificial nuclei
comprising hydrogels in the form of large pieces shaped when fully
hydrated, to generally conform to the disc cavity or hydrogel beads
within a porous envelope, respectively. The hydrogels have an
equilibrium water content (EWC) of at least about 30% and a
compressive strength of at least about 1 meganewtons per square
meter (1 MNm.sup.-2) when subjected to the constraints of the
annulus and end plates of the disc. Preferably, the compressive
strength of the nucleus is about 4 MNm.sup.-2 or even higher.
[0018] The primary disadvantage of the invention of Substad et al.,
Patil, Kenna and Lee et al., is that use of their prosthesis
requires complete replacement of the natural disc which involves
numerous surgical difficulties. Secondly, the intervertebral disc
is a complex joint, anatomically and functionally, comprising the
aforementioned three component structures, each of which has its
own unique structural characteristics. Designing and fabricating
such a complicated prosthesis from acceptable materials, which will
mimic the function of the natural disc, is very difficult. A
further problem is the difficulty of preventing the prosthesis from
dislodging. Fourthly, even for prostheses which are only intended
for replacing the nucleus, a major obstacle has been to find a
material which is similar to the natural and is also able to
restore the normal function of the nucleus. Hydrophobic elastomers
and thermoplastic polymers are not desirable for use in the
prosthetic nuclei due to their significant inherent differences
from the natural nucleus e.g., lack of hydrophilicty, in the
elastomers, and lack of flexibility in their thermoplasts.
[0019] These problems are not solved by Kuntz, who uses elastic
rubber plugs, or by Froning and Ray et al., who use bladders, or
capsules, respectively, which are filled with a fluid or
thixotropic gel. According to the Ray and Froning patents, liquid
was used to fill the capsules and bladders, respectively, thereby
requiring that their membranes be completely sealed to prevent
fluid leakage. As a consequence, those devices cannot completely
restore the function of the nucleus which allows body fluid to
diffuse in and out during cyclic loading thereby providing the
nutrients the disc needs.
[0020] The Bao et al., prosthetic lumbar disc nuclei are made from
hydrogels. Hydrogels have been used in biomedical applications,
such as contact lenses. Among the advantages of hydrogels is that
they are more biocompatible than hydrophobic elastomers and metals.
This biocompatibility is largely due to the unique characteristics
of hydrogels in that they are soft and contain water like the
surround tissues and have relatively low frictional coefficients
with respect to the surrounding tissues. The biocompatibility of
hydrogels results in prosthetic nuclei which are more easily
tolerated in the body. Furthermore, hydrophobic elastomeric and
metallic gels will not permit diffusion of aqueous compositions,
and their solutes, therethrough.
[0021] An additional advantage of some hydrogels is their good
mechanical strength which permits them to withstand the load on the
disc and restore the normal space between the vertebral bodies. The
aforementioned nuclei of Bao et al. have high mechanical strength
and are able to withstand the body loads and assist in the healing
of the defective annuli.
[0022] Other advantages of the hydrogels, used in Bao et al.
nuclei, are their excellent viscoelastic properties and shape
memory. Hydrogels contain a large amount of water which acts as a
plasticizer. Part of the water is available as free water which has
more freedom to leave the hydrogel when the hydrogel is partially
dehydrated under mechanical pressure. This characteristic of the
hydrogels enables them to creep, in the same way as the natural
nucleus, under compression and to withstand cyclic loading for long
periods without any significant degradation or loss of their
elasticity. This is because water in the hydrogel behaves like a
cushion whereby the polymeric network of a hydrogel with a high
equilibrium water content (EWC) is less susceptible to damage under
mechanical load.
[0023] Another advantage of hydrogels is their permeability to
water and water-soluble substances, such as nutrients, metabolites
and the like. It is know that body fluid diffusion, under cycle
loading, is the major source of nutrients to the natural disc. If
the route of this nutrient diffusion is blocked, e.g., by a
water-impermeable nucleus, further deterioration of the disc will
ensure.
[0024] Hydrogels can be dehydrated and the resultant xerogels
hydrated again without changing the properties of the hydrogels.
When a hydrogel is dehydrated, its volume decreases, thereby
facilitating implantation of the prosthetic nucleus into the
nuclear cavity in the disc. The implanted prosthetic nucleus will
then swell, in the body, by absorption of body fluid up to its EWC.
The EWC of the hydrogel depends on the compressive load applied
thereto. Thus, the EWC of a specific hydrogel in an open container
will differ from the EWC of the same hydrogel in a closed vessel
such as an intervertebral disc. The EWC values, referred to below,
are for hydrogels subjected to compressive loads under the
conditions found in an intervertebral disc. The expansion factor of
a dehydrated hydrogel, in turn, is dependent on its EWC. Thus, it
may vary from 1.19 for a hydrogel of 38% EWC to 1.73 for a hydrogel
of 80% EWC. For an 80% EWC hydrogel, the volume of the dehydrated
prosthetic nucleus is usually about 20% of that of the hydrated
one. The ability to be dehydrated and then return to its original
shape upon hydration, up to its EWC, makes it possible to implant
the device posterior-laterally during surgery, thereby reducing the
complexity and risk of intraspinal surgery as traditionally used.
The danger of perforation of the nerve, dural sac, arteries and
other organs is also reduced. In addition, the incision area on the
annulus can be reduced, thereby helping to heal the annulus and
prevent the reherniation of the disc. Hydrogels are also useful for
drug delivery into the disc due to their capability for controlled
release of drugs. Various therapeutic agents, such as growth
factors, long term analgesics and anti-inflammatory agents can
attach to the prosthetic nucleus and be released in a controllable
rate after implantation of the nucleus in the disc.
[0025] Furthermore, dimensional integrity can be maintained with
hydrogels having a water content of up to about 90%. This
dimensional integrity, if the nucleus is properly designed will aid
in distributing the vertebral load to a larger area on the annulus
ring and prevent the prosthetic nucleus from bulging and
herniating.
[0026] However, it is normally difficult to implant a fully
hydrated hydrogel prosthesis in the cavity, of a disc, through the
small window provided in the disc, for removing the herniated
nucleus, especially in a percutaneous surgery by virtue of their
bulkiness in a fully in a fully hydrated state. Therefore, such
prosthesis must be implanted, in the disc in relatively dehydrated
states which requires long periods to achieve their EWCs due to
their low surface areas. Other hydrogels, having high surface
areas, do not completely conform to the shape of the nuclear
cavity. Other polymers such as those disclosed in WO 97/268407
(PCT/US97 00457), the teachings of which are incorporated herein by
reference, can also be used to fill the disc nucleus.
[0027] It is desirable to provide a hydrogel implant which is
inherently radiopaque, i.e., radiovisible so that surgeons could
view the placement of the implant in the cavity produced by the
removal of a spinal nucleus. It is advantageous if the radiovisible
material could be incorporated into the polymeric or hydrogel
material making up the prosthetic nucleus implant. It is desirable
to have a method of making the hydrogel or polymer radiopaque which
would allow dimensional changes in the hydrogel implant during
processing and after implantation without compromising the
mechanical integrity of the implant.
[0028] Various methods are used to implant a hydrogel or other
polymeric nucleus implant. Such a method is shown in U.S. Pat. No.
5,800,549, the teachings of which are incorporated herein by
reference.
SUMMARY OF THE INVENTION
[0029] It is an object of the invention to provide a novel hydrogel
or other polymeric implant for replacing the resected natural
nucleus of a spinal disc.
[0030] It is a further object of the invention to provide a novel
hydrogel or other polymeric prosthetic nucleus which has
radiovisible material contained therein.
[0031] It is yet a further object of the invention to provide a
method for incorporating radiovisible material within the hydrogel
or polymer either dispersed throughout the implant or in discreet
locations therein.
[0032] Such objects are achieved by the spinal implant for
replacing the natural nucleus of the disc made of a hydrogel having
radiovisible material located within the hydrogel. The material may
be a metal such as gold, tungsten, tantalum, platinum, titanium or
combinations thereof.
[0033] The material may be in powder form and may be distributed
throughout the hydrogel in a uniform manner or may be in powder
form and placed in the hydrogel or polymeric implant in discreet
layers or locations. The radiopaque powder preferably has a maximum
diameter of between 10 and 100 .mu.m and more preferably the powder
has a diameter of about 75 .mu.m
[0034] Alternately, the metal may be in the form of foil, either in
strip form in the form of flakes scattered throughout the implant.
If the foil is in strip form, it should be relatively thin, i.e.,
in the range of 1-100 .mu.m thick so that when used with a
hydrogel, it expands and contracts as the hydrogel is hydrated and
dehydrated. In the preferred embodiment, a thickness at the lower
end of this range is desirable, for example, 2 .mu.m thick.
[0035] In an additional embodiment, the implant may be in the form
of a metal wire or coil placed in the hydrogel implant during its
formation. Again, the wire is of such a diameter as to be able to
fold upon itself during hydration and dehydration of the
hydrogel.
[0036] If the polymer used is formed in situ then the metal that
imparts radiovisibility to the implant is dispersed throughout the
polymer prior to injection and curing.
[0037] If the polymer is processed in the melt then the metal is
blended into the polymer when it is above its melt temperature.
[0038] Methods of making the hydrogel, including the radiopaque
material, include dissolving the polymer powder to form a
homogeneous solution and then mixing the metal powder with metal
flakes therein while the solution is still a liquid and then
freezing the solution to form a solid. Usually the solution is
poured into a mold to form the hydrogel and then the mold is placed
in the freezer.
[0039] Alternately, the hydrogel implant can be formed sequentially
by placing a homogeneous solution with radiopaque material in the
mold but filling only a portion of the mold, freezing the solution,
placing an additional layer of hydrogel, including the radiopaque
metal upon the first layer of solidified hydrogel and thereafter
freezing the second layer to form a solid. This process may be
repeated to form alternate layers in the prosthetic nucleus which
are either radialucent or radiopaque.
[0040] In yet another embodiment, the polymeric implant can be
formed by melting a polymer such as poly(acrylonitrile) and
blending the radiopacifying agent prior to molding the implant.
Sequential molding operations performed on one implant can result
in the radiopacifying agent being localized into discreet areas of
the implant. The polymeric implant can also be formed by injecting
both a crosslinkable material (e.g. monomer and/or prepolymer) and
a crosslinking agent, and then allowing the crosslinkable material
to cure in situ such as polyurethane. Radiovisibility can be
imparted to this implant by adding the radiopacifying agent to the
crosslinkable material, the crosslinking agent, or both.
[0041] If a coil or foil is used to produce the radiopaque portions
of the hydrogel, the foil or coil may be placed in the liquid
hydrogel prior to solidifying it by cooling.
BRIEF DESCRIPTION OF THE DRAWINGS
[0042] FIG. 1 is an elevational view of the vertebral disc absent
its nucleus with its associated vertebra;
[0043] FIG. 2 is an elevated view of the intervertebral disc and
associated vertebra of FIG. 1 from which the nucleus has been
removed;
[0044] FIG. 3 is an elevational view of the disc of FIG. 2 with the
polymeric nucleus of the present invention implanted therein
showing a radiopaque powder dispersed throughout the implant;
[0045] FIG. 4 is an elevational view of the disc of FIG. 2 with a
polymeric implant having a foil strip therein;
[0046] FIG. 5 is an elevational view of the disc space of FIG. 2
with a polymeric implant having a metal coil therein; and
[0047] FIG. 6 is an elevational view of the disc of FIG. 2 having a
polymeric implant, including three layers with each layer including
radiopaque metal powder dispersed therethrough.
DETAILED DESCRIPTION
[0048] Referring to FIGS. 1 through 6, in the preferred embodiment
the hydrogel prosthetic nucleus of the present invention, generally
denoted as 10, conforms when hydrated to its EWC, generally to the
shape of the natural nucleus. Alternately, the hydrogel can be
constrained in a polymer jacket. Such is taught in U.S. Pat. Nos.
5,674,295 and 6,132,465. The prosthetic nucleus is implanted within
cavity 11 the disc 12 of the vertebrae 14 and is surrounded by the
natural annulus 16. Vertebral end plates 20 and 22, as shown in
FIG. 1, cover the superior and inferior faces of nucleus 10
respectively. The implant is inserted through an opening 62 in
annulus 12.
[0049] Referring to FIG. 3, there is shown the prosthetic nucleus
of the present invention, including metal particles 30 dispersed
uniformly throughout. Uniformly is used in a relative sense not to
indicate that the particles are exactly spaced within the
hydrogel.
[0050] Referring to FIG. 4, there is shown the prosthetic disc
nucleus of the present invention with a radiopaque foil strip 40
located therein.
[0051] Referring to FIG. 5, there is shown a polymeric implant of
the present invention having a radiopaque or radiovisible coil 50
located therein.
[0052] Referring to FIG. 6, there is shown a prosthetic disc
nucleus 10 of the present invention having radiopaque layers of
metal particles or foil particles 60 located therein. While three
layers are shown in FIG. 6, one layer or two layers or more than
three layers may be utilized.
[0053] Hydrogels useful in the practice of the invention include
lightly cross-linked biocompatible homopolymers and copolymers of
hydrophilic monomers such as 2-hydroxylalkyl acrylates and
methacrylates, e.g., 2-hydroxyethyl methacrylate (HEMA); N-vinyl
monomers, for example, N-vinyl-2-pyyrolidone (N-VP); ethylenically
unsaturated acids, for example, methacrylic acid (MA) and
ethylenically unsaturated bases such as 2-(diethylamino) ethyl
methacrylate (DEAEMA). The copolymers may further include residues
from non-hydrophilic monomers such as alkyl methacrylates, for
example, methyl methacrylate (MMA), and the like. The cross-linked
polymers are formed, by known methods, in the presence of
cross-linking agents, such as ethyleneglycol dimethacrylate and
methylenebis (acrylamide), and initiators such as 2,2-azobis
(isobutyronitrile), benzoyl peroxide, and the like, and radiation
such as UV and .gamma.-ray.
[0054] Methods for the preparation of these polymers and copolymers
is well known to the art. The EWC of these hydrogels can vary,
e.g., from about 38% for Polymacon.TM. (poly HEMA) to about 79% for
Lidofilcon.TM. B (a copolymer of N-VP and MMA) under ambient
conditions.
[0055] Another type of hydrogel, useful in the practice of the
invention, is illustrated by HYPAN.TM. and poly(vinyl alcohol)
(PVA) hydrogels. These hydrogels, unlike the aforementioned
hydrogels, are not cross-linked. Their insolubility in aqueous
media is due to their partially crystalline structures. HYPAN.TM.
is a partially hydrolyzed polyacrylonitrile. It has a multiblock
copolymer (MBC) structure comprising hard crystalline nitrile
blocks, which provide the hydrogel with good mechanical properties,
and soft amorphous hydrophilic blocks to provide the hydrogel with
good water binding capability. The methods of preparing HYPAN.TM.
hydrogels of different water contents and mechanical properties
have been disclosed in the U.S. Pat. Nos. 4,337,327, 4,370,451,
4,331,783, 4,369,294, 4,420,589, 4,379,874 and 4,631,188. The
pre-nuclear forms of this material, for use in this invention, can
be prepared by melt processing using solvents such as DMF and DMSO,
as melting aids or by solution processing.
[0056] Other types of polymers useful in the practice of the
invention include medical grade polyurethanes and materials formed
by crosslinking protein precursors. These materials may or may not
form hydrogels in their final form but are still useful as
materials to form prosthetic nucleus replacements. Such materials
are shown in U.S. Pat. Nos. 5,888,220, 6,189,048, 6,183,518 and in
Publication US 20020049498 A1, the teachings of which are
incorporated herein by reference.
[0057] A preferred hydrogel for use in the practice of this
invention is highly hydrolyzed crystalline poly (vinyl alcohol)
(PVA). The amount of hydrolyzation may be between 95 and 100
percent depending on the desired EWC which will be from about 60%
to about 90%. Generally, the final hydrogel water content increases
with decreasing hydrolyzation of the initial PVA which results in
decreased crystallinity.
[0058] Partially crystalline PVA hydrogels may be prepared, from
commercially available PVA powders, by any of the methods disclosed
in the U.S. Pat. No. 4,663,358, the teachings of which are
incorporated herein by reference. Typically, 10-15% PVA powder is
mixed with a solvent, such as water, dimethyl sulfoxide (DMSO),
ethylene glycol and mixtures thereof. A preferred solvent is 15%
water in DMSO. The mixture is then heated at a temperature of about
100 to about 120.degree. C., until a viscous solution is formed.
The solution is then poured or injected into a tubular metal, glass
or plastic mold and allowed to cool to below -10.degree. C.,
preferably about -20.degree. C.
[0059] The solution is maintained at the temperature for several
hours, preferably about 20 hours, during which time crystallization
and, therefore, gelation of the PVA occurs. The shaped gel is
soaked with several portions of water which are periodically
replaced, over a period of at least two days, until all the organic
solvent in the gel has been replaced by water. The hydrated gel can
then be partially or completely dehydrated for implantation. The
hydrogels thus prepared have EWC's between 60-90% and compressive
strengths of at least 1 MNm.sup.-2, preferably about 4 MNm.sup.-2,
when subject to the same constraints as the natural nucleus in an
intervertebral disc. In general, any polymer that can be used for
biomedical purposes can be used as long as the polymer exhibits the
desired stiffness characteristics.
[0060] Completion of the solvent exchange is determined by known
methods. For instance, when the solvent is DMSO its removal, from
the gel, is determined as follows:
[0061] 50 .mu.L of a 0.01 N KMnO.sub.4 are added to 50 mL aliquots
of the water which has been separated from the gels. The presence
of DMSO, in the water, will be indicated by disappearance of the
characteristic pink color of the KMnO.sub.4. When the DMSO has been
completely removed, the pink color will not disappear. This method
has a detection limit of 0.3 ppm, for DMSO, when compared to a
blank and 0.3 ppm aqueous DMSO standard.
[0062] In general, any hydrogel that can be used for biomedical
purposes can be used as long as the hydrogel exhibits an EWC from
about 30 to 90% and a compressive strength of at least about 1
MNm.sup.-2, preferably 4 MNm.sup.-2, when subjected to the
constraints of the annulus and end plates of the disc. A rod or
tube made from these materials, in a dehydrated form, i.e., as
xerogels, can be prepared either by cast molding or lathe cutting.
In cast molding, the liquid monomer mixture, with initiator, is
poured into a mold of predetermined shape and size, and cured. If
desired, the casting mixture may include water, or another aqueous
medium. Under those circumstances the resultant rod or tube will be
partially hydrated, i.e., a hydrogel. In the case of lathe cutting,
the xerogel can be prepared, in a similar manner to the above, in
the form of a block or rod which is larger than needed to form the
prosthetic nucleus. The xerogel is then cut to the shape and size
required for implantation into the disc cavity. In both cases, the
hydrogel expansion factor, due to polymer swelling upon hydration,
has to be taken into account in designing the mold or in cutting
the block, rod or tube.
[0063] The exact size of the prosthetic nucleus, at its EWC, can be
varied for different individuals. A typical size of an adult
nucleus is about 2 cm in the semi-minor axis, about 4 cm in the
semi-major axis and about 1.2 cm in thickness.
[0064] Polymers curable within the body can also be used to replace
the natural nucleus and strengthen the annulus which is made of
cartilage. Natural cartilage is a non-vascular structure found in
various parts of the body. Articular cartilage tends to exist as a
finely granular matrix forming a thick incrustation on the surfaces
of joints. The natural elasticity of articular cartilage enables it
to break the force of concussions, while its smoothness affords
ease and freedom of movement. Preferred biomaterials, therefore,
are intended to mimic many of the physical-chemical characteristics
of natural cartilage. Biomaterials can be provided as one component
systems, or as two or more component systems that can be mixed
prior to or during delivery, or at the site of repair. Generally
such biomaterials are flowable in their uncured form, meaning they
are of sufficient viscosity to allow their delivery through a
cannula of on the order of about 2 mm to about 6 mm inner diameter,
and preferably of about 3 mm to about 5 mm inner diameter. Such
biomaterials are also curable, meaning that they can be cured or
otherwise modified, in situ, at the tissue site, in order to
undergo a phase or chemical change sufficient to retain a desired
position and configuration.
[0065] When cured, preferred materials can be homogenous (i.e.,
providing the same chemical-physical parameters throughout), or
they can be heterogenous. An example of a heterogenous biomaterial
for use as a disc replacement is a biomaterial that mimics the
natural disc by providing a more rigid outer envelope (akin to the
annulus) and a more liquid interior core (akin to the nucleus). In
an alternative embodiment, biomaterials can be used that provide
implants having varying regions of varying or different
physical-chemical properties. With disc replacement, for instance,
biomaterials can be used to provide a more rigid, annulus-like
outer region, and a more fluid, nucleus-like core. Such di-or
higher phasic cured materials can be prepared by the use of a
single biomaterial, e.g., one that undergoes varying states of
cure, or by using a plurality of biomaterials. Examples of suitable
biomaterials includes, but are not limited to, polyurethane
polymers.
[0066] The in situ cured polymer may comprise a thermosetting
polyurethane polymer based on a suitable combination of
isocyanates, long chain polyols and short chain (low molecular
weight) extenders and/or crosslinkers. Suitable components are
commercially available and are each preferably used in the highest
possible grade, e.g., reagent or preferably analytical grade or
higher. Examples of suitable isocyanates include 4,4'-diphenyl
methane diisocyanate ("MDI"), and 4,2'-diphenylmethane diisocyanate
("TDI"). Examples of suitable long chain polyols include
tetrahydrofuran polymers such as poly(tetramethylene oxide)
("PTMO"). Particularly preferred are combinations of PTMO's having
molecular weights of 250 and 1000, in ratios of between about 1 to
1 and about 1 to 3 parts, respectively. Examples of suitable
extenders/crosslinkers include 1,4-butanediol and trimethylol
propane, and blends thereof, preferably used at a ratio of between
about 1 to 1 and about 1 to 7 parts, respectively. Such performance
can be evaluated using procedures commonly accepted for the
evaluation of natural tissue and joints, as well as the elevation
of biomaterials.
[0067] In particular, the in situ cured polymer forms, exhibit
mechanical properties that approximate those of the natural tissue
that they are intended to replace. For instance, for load bearing
applications, preferred cured composites exhibit a load bearing
strength of between about 50 and about 200 psi (pounds per square
inch), and preferably between about 100 and about 150 psi. Such
composites also exhibit a shear stress of between about 10 and 100
psi, and preferably between about 30 and 50 psi, as such units are
typically determined in the evaluation of natural tissue and
joints.
[0068] Biomaterials provided as a plurality of components, e.g., a
two-part polyurethane system, can be mixed with the radiopaque or
radiovisible metal powder at the time of use using suitable mixing
techniques, such as those commonly used for the delivery of
two-parts adhesive formulations. More preferably, the metal powder
can be added during melt processing of the polymer. A suitable
mixing device involves, for instance, a static mixer having a
hollow tube having a segmented, helical vein running through its
lumen. A two-part polyurethane system can be mixed by forcing the
respective components through a lumen, under pressure.
[0069] The hydrogels and polymers of the present invention have a
much higher structural integrity than the natural nucleus, i.e.,
they are deformed with greater difficulty under a mechanical
compressive load (shaped gel vs. loose gel). That is because,
unlike the loose gel of the natural nucleus, the shaped gel has
shape memory due to the cross-linking or strong hydrogen bonding in
the polymer matrix. However, it would still have extensive lateral
bulging under high compressive load if there were no boundaries to
constrain the deformation. Because use of the prevent invention
does not involve removal of the disc annulus and/or end-plates, the
lateral bulging of the hydrogel nucleus will be restricted by the
restoring forces of the stretch fibers. Also, due to its superior
structural integrity, the hydrogel nucleus will not herniate or
bulge through the previously herniated areas or the incision which
was made to remove the degenerated nucleus.
[0070] If a hydrogel is used, since the natural nucleus is also
primarily, a hydrogel, the implanted prosthetic nucleus can easily
restore all the biomechanical functions of the nucleus which had
been removed. Unlike the prior art prosthetic discs, the hydrogel
nucleus of the present invention will restore the viscoelastic
behavior of the disc due to the water binding capability of the
prosthetic hydrogel.
[0071] The implantation of a prosthetic nucleus 10 can be performed
in conjunction with a discectomy or chemonuclealysis. Because the
properties of the prosthetic nucleus of the present invention are
similar to those of the nucleus material, the herniated nucleus can
be partially or totally replaced by the hydrogel prosthetic
nucleus. Due to the small size of the prosthetic it can be
implanted into the disc by means of a posterior lateral approach,
thereby significantly reducing the difficulty and the risk of the
operation.
[0072] The volume of a hydrogel nucleus of about 80% EWC will be
reduced by about 80% (to about 20% of its original volume) when
dehydrated. Consequently, the surgeon does not need to jack apart
the vertebrae adjacent to a damaged disc as required by, for
example, the device disclosed in U.S. Pat. No. 4,772,287. The
height of the dehydrated prosthetic nucleus, when inserted, is
smaller than the disc space. Furthermore, the rigidity of the
dehydrated prosthetic nucleus will help the surgeons to manipulate
the prosthetic nucleus during the operation. After implantation,
the hydrogel nucleus of the present invention swells in the body to
a predetermined height which is enough to maintain the space
between the vertebral body. The swelling process normally takes
several hours to two days depending on the size of the prosthetic
nucleus and type of hydrogel.
[0073] The preferred method for making the radiopaque hydrogel or
polymer involves incorporating a metallic element into the
structure of the implant. The metallic element is in a form that
will allow it move with the polymeric structure as the implant
changes dimensions and/or geometry. This property is important
because it minimizes any internal stress amplification that could
be caused by incorporating the metallic component into the dynamic
(i.e. non-fusion) spine implant.
[0074] Three embodiments of the invention are described. The first
embodiment involves incorporating radiopaque materials such as a
metal powder, for example, gold or tungsten, into the polymer while
it is a liquid either due to the use of solvents, heat if the
polymer can be processed in the melt, or is in a pre-polymer form
prior to a curing step. The metal powder has a nominal diameter of
10-100 .mu.m, with a preferred maximum size of 75 .mu.m. The powder
is incorporated into the liquid polymer solution/melt preferably in
a concentration between 0.02 and 0.5 g per cc of polymer, with a
preferred concentration of 0.1 g/cc. The powder may be evenly
dispersed throughout the entire implant, or incorporated into the
implant in specific areas only. For example, by combining the use
of liquid-phase polymer that contains no metal powder in a mold
with liquid polymer that does contain metal powder it is possible
to form radiovisible areas of an implant in a variety of geometries
(e.g. lines, discs, spheres). An embodiment of the layered
polymeric implant has one or more planes in the implant comprised
of the radiovisible polymer. It is possible to incorporate 1 to 5
or more bands of powder-filled polymer in any plane across an
implant that also has polymer regions where no metal powder exits
using techniques that prevent the powder from migrating until the
liquid-phase polymer has formed a solid, as discussed below or
creating an even dispersion of metal powder throughout the entire
implant.
[0075] A second embodiment involves the use of a metal foil 0.001
to 0.1 mm thick for example gold, tantalum or tungsten foil. The
foil can be placed into the liquid polymer in the form of sheets or
strips, or it can be chopped into small pieces and incorporated
into the implant in the manner described for metal powders. Small
pieces of foil may provide an advantage for some polymer systems
with a lower viscosity in the liquid phase because each piece may
have less mass than a metal particle, which would result in less
tendency for migration through the liquid-phase polymer. In
addition, the geometry of a piece of foil with multiple irregular
folds may also have less tendency to migrate through liquid-phase
polymer than, for example, a smoother spherical particle. Examples
include suspending 1 to 5 or more strips of metal foil in any plane
in the implant, or incorporating 1 to 5 or more bands of polymer
that contain small pieces of chopped foil that may or may not be
"crinkled" as described above for powder, and creating an even
dispersion of pieces of chopped metal foil that may or may not be
"crinkled" throughout the implant.
[0076] A third embodiment involves the use of a metal wire or coil,
preferably 0.01 to 1.0 mm in length, for example, gold wire,
tungsten wire or platinum wire. The metal wire is suspended in the
liquid-phase polymer in such a manner that the wire will be
completely encapsulated by solid polymer at the end of the
manufacturing process. Alternatively, the wire can be formed into a
coil or other shape that may provide better radiographic
information about the implant and have less of an ability to
migrate through the solid polymer. Examples include suspending 1 to
5 or more pieces of wire in the liquid-phase polymer in such a way
as they will not be exposed to the surface of the implant, or
incorporating 1 to 5 or more bands of polymer that contain small
pieces of chopped wire or creating an even dispersion of pieces of
chopped metal wire throughout the implant using techniques as
described below.
EXAMPLE I
[0077] A PVA solution was formed by mixing 15 g of PVA powder
(Kuraray 117 or equivalent), having a molecular weight about 78000
and about 99.7% hydrolysed (Cat. No. 15129, Polysciences Inc.,
Warrington, Pa.), with 85 ml of a solvent comprising 15% water in
DMSO. The mixture was heated at about 110.degree. C. until a
homogenous viscous solution formed.
EXAMPLE II
[0078] 0.1 gram of gold powder (maximum diameter 75 .mu.m) per cc
of liquid-phase of PVA solution of Example I were mixed. The two
ingredients were combined in the following manner to create a
metal-filled polymer solution. A plunger from a 5 cc first syringe
was removed and the first syringe was slowly filled half way with
PVA solution, 0.5 g of the gold powder was poured into the syringe.
The first syringe was then completely filled with PVA solution, and
the plunger replaced. A two-way luer connector was screwed onto the
tip of the first syringe and the connector was primed with PVA
solution from the syringe. The first syringe and a second syringe
of equal size were connected using the connector. The gold powder
solution from the first syringe was squeezed into the empty
syringe. This was repeated until the solution was uniformly mixed.
A third syringe was filled with PVA solution without any gold. 5 cc
of the PVA solution of the third syringe was injected into a
nucleus mold having a total volume of about 20 cc and a diameter of
about 1.5 cm. The mold was cooled down in freezer (4.degree. C.)
for about 20 minutes. 1 cc of gold powder-filled PVA solution was
injected on top of the cooled solution in the mold. 9 cc of the PVA
solution was slowly injected into the mold, on top of the gold
powder solution. The mold was again cooled down in the freezer for
about 15 minutes. About 5 cc of gold powder-filled solution was
again injected on top of the cooled solution in the mold to form a
second radiopaque layer. The mold was then completely filled with
the PVA solution from the third syringe.
EXAMPLE III
[0079] About 20 cc of the gold solution was prepared as described
in Example II using two 20 cc syringes and 2 grams of gold powder.
The metal filled polymer solution was slowly injected into a 20 cc
implant mold (for a #5 size implant), filling the mold completely.
The mold reservoir was capped so that the metal-filled PVA solution
will not leak out of the mold if the mold is inverted. The mold was
pressurized and placed in a Turbula Mixer which was placed into a
programmable freezer. The mixer used must be able to both rotate
and tip the mold during gelation of the metal filled PVA solution
to keep the powder in the solution uniformly distributed. The mixer
can rotate the mold about a central axis, tilt the mold back and
forth through a 90.degree. arc about an axis perpendicular to the
central axis. This kept the metal particles uniformly suspended in
the solution until it gelled.
EXAMPLE IV
[0080] The PVA solution prepared for the third syringe of Example
II without the gold powder was used in this Example. A strip of
gold metal foil was suspended in an empty mold using a thin
monofilament of nylon so that the strip is positioned in
approximately the center of the mold. The nylon filament was smooth
and non-porous. The mold was slowly filled with the PVA solution.
The mold was then placed in the freezer and after completing the
process, the monofilament was pulled out of the implant leaving the
strip of metal foil intact.
EXAMPLE V
[0081] The gold foil of Example IV was cut into small pieces using
a food processor or other convenient method to create metal flakes.
0.03 g of metal flakes per cc of the liquid phase polymer of
Example I. 20 cc of solution were placed in a mold and the solution
was then frozen as above to form the implant.
EXAMPLE VI
[0082] A gold coil was suspended in an empty 20 cc mold using a
thin monofilament so that the coil is positioned in approximately
the center of the mold. The filament was smooth and non-porous. A
barb fitting through a small loop at the top of the metal coil is a
convenient way to attach the filament and the coil. Slowly fill the
molds with the PVA solution of Example I. After completing the
gelation process by freezing, the monofilament was pulled out of
the implant leaving the gold coil intact.
[0083] After gelation, the implants were sterilized as described in
my Ion Treated Hydrogel copending application, U.S. Ser. No.
10/020,389 and then packaged.
[0084] The implants are formed into 10 different sizes by volume
with a size range of 1.1 to 5.2 cc for future implantation.
[0085] Although the invention herein has been described with
reference to particular embodiments, it is to be understood that
these embodiments are merely illustrative of the principles and
applications of the present invention. It is therefore to be
understood that numerous modifications may be made to the
illustrative embodiments and that other arrangements may be devised
without departing from the spirit and scope of the present
invention as defined by the appended claims.
* * * * *