U.S. patent number RE34,421 [Application Number 07/870,145] was granted by the patent office on 1993-10-26 for x-ray micro-tube and method of use in radiation oncology.
Invention is credited to Earl R. Parker, William J. Parker.
United States Patent |
RE34,421 |
Parker , et al. |
October 26, 1993 |
X-ray micro-tube and method of use in radiation oncology
Abstract
An apparatus and method for the treatment of a patient having a
tumor is disclosed. An X-ray generating source is positionable at a
location in close proximity to the tumor. The X-ray generating
source is operable at a voltage level in the range of approximately
10-60 KeV, thereby enhancing absorption of the generated X-rays by
the tumor and minimizing the side effects of radiation therapy on
the patient's normal tissue.
Inventors: |
Parker; William J. (West Hills,
CA), Parker; Earl R. (San Mateo, CA) |
Family
ID: |
24469277 |
Appl.
No.: |
07/870,145 |
Filed: |
April 17, 1992 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
Issue Date |
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Reissue of: |
616397 |
Nov 21, 1990 |
05090043 |
Feb 18, 1992 |
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Current U.S.
Class: |
378/121; 378/65;
378/122 |
Current CPC
Class: |
A61N
5/1001 (20130101); H01J 35/32 (20130101); H01J
35/00 (20130101); H01J 35/116 (20190501) |
Current International
Class: |
A61N
5/10 (20060101); H01J 35/00 (20060101); H01J
35/32 (20060101); H01J 035/32 () |
Field of
Search: |
;378/64,65,119,121-123,130,140 ;600/1-3,6,7 |
References Cited
[Referenced By]
U.S. Patent Documents
Other References
Dunlee OL-1-DL-7, Stationary Anode Isert, Brochures, Jun.
1972..
|
Primary Examiner: Porta; David P.
Claims
What is claimed and desired to be secured to Letters Patent of the
United States is:
1. An apparatus for the treatment of a patient having a tumor,
comprising:
an X-ray generating source insertable into the body of said patient
to a location in close proximity to said tumor, said X-ray
generating source comprising a substantially cylindrical glass
X-ray tube having a length on the order of one-quarter to two
inches and a diameter less than one inch and being operable at a
voltage level in the range of approximately 10-60 KeV, thereby
enhancing absorption of the generated X-rays by said tumor and
minimizing the side effects of radiation therapy on the patient's
normal tissue.
2. The apparatus of claim 1, wherein said X-ray generating source
includes a metal-jacketed micro-tube assembly, comprising:
a) an evacuated glass tube;
b) a cathode supported within said evacuated glass tube;
c) an anode supported within said evacuated glass tube;
d) support means for supporting said cathode and said anode;
e) a metal jacket enclosing said evacuated glass tube therein, said
metal jacket including a window for directing the generated
radiation in the desired manner; and,
f) electrical connection means for connecting said cathode and said
anode to a source of electrical power.
3. The apparatus of claim 1, wherein said glass X-ray tube contains
a stable vacuum of at most 10.sup.-6 Torr.
4. The apparatus of claim 3, wherein said glass X-ray tube is a
cold emission cathode tube.
5. The apparatus of claim 3, wherein said glass X-ray tube includes
a heated filament cathode.
6. The apparatus of claim 3, further including a power source for
operating said glass X-ray tube at a frequency range of d.c. to
1,000,000 cycles per second.
7. The apparatus of claim 3, further including a power source for
operating said glass X-ray tube at a frequency in the range between
70,000 and 1,000,000 cycles per second.
8. The apparatus of claim 1, wherein said glass X-ray tube has a
length of approximately one inch and a diameter of approximately
one-quarter inch or less.
9. The apparatus of claim 1, wherein said X-ray generating source
includes a liquid-cooled micro-tube assembly, comprising:
a) an evacuated glass tube;
b) a filament cathode supported within said evacuated glass
tube;
c) an anode supported within said evacuated glass tube;
d) support means for supporting said cathode and said anode;
e) a housing for supporting and enclosing said evacuated glass tube
therein, a liquid coolant chamber being formed between said housing
and said glass tube;
f) electrical connection means for connecting said cathode and said
anode to a source of electrical power; and,
g) coolant conduits for supplying coolant to said coolant
chamber.
10. An X-ray micro-tube for the treatment of a patient having a
tumor, comprising:
a substantially cylindrical evacuated glass X-ray tube having a
length on the order of one-quarter to two inches and a diameter
less than one inch and containing a stable vacuum of, at most,
10.sup.-6 Torr, said glass X-ray tube being locatable at a location
in close proximity to said tumor, said X-ray tube being operable at
a voltage level in the range of approximately 10-60 KeV, thereby
enhancing absorption of the generated X-rays by said tumor and
minimizing the side effects of radiation therapy on the patient's
normal tissue.
11. The X-ray micro-tube of claim 10, wherein said glass X-ray tube
operates at a frequency of between d.c. and 1,000,000 cycles per
second.
12. The X-ray micro-tube of claim 11, wherein said glass X-ray tube
operates at a frequency in the range between 70,000 and 1,000,000
cycles per second.
13. The X-ray micro-tube of claim 12, wherein said glass X-ray tube
includes a coolant jacket positioned about its periphery.
14. The X-ray micro-tube of claim 12, further including protective
shielding circumscribing said glass X-ray tube for patient
protection in the event of glass breakage.
15. A method for the in-situ treatment of a patient having a tumor,
comprising:
(a) providing an X-ray generating source comprising a substantially
cylindrical glass X-ray tube having a length on the order of
one-quarter to two inches and a diameter less than one inch and
being operable at a voltage level in the range of approximately
10-60 KeV;
(b) positioning said X-ray generating source at a location in close
proximity to said tumor; and,
(c) applying power to said X-ray generating source in said range of
approximately 10-60 KeV, thereby enhancing absorption of the
generated X-rays by said tumor and minimizing the side effects of
radiation therapy on the patient's normal tissue.
16. The method of claim 15 wherein said step of positioning said
X-ray generating source includes inserting said source into the
body of said patient to a location in close proximity to said
tumor. .Iadd.
17. An apparatus for the treatment of a patient having a tumor,
comprising:
an X-ray generating source insertable into the body of said patient
to a location in close proximity to said tumor, said X-ray
generating source comprising a substantially cylindrical glass
X-ray tube having a diameter less than one inch and being operable
at a voltage level in the range of approximately 10-60 KeV, thereby
enhancing absorption of the generated X-rays by said tumor and
minimizing the side effects of radiation therapy on the patient's
normal tissue. .Iaddend. .Iadd.18. An X-ray micro-tube for the
treatment of a patient having a tumor, comprising:
a substantially cylindrical evacuated glass X-ray tube having a
diameter less than one inch and containing a stable vacuum of, at
most, 10.sup.-6 Torr, said glass X-ray tube being locatable at a
location in close proximity to said tumor, said X-ray tube being
operable at a voltage level in the range of approximately 10-60
KeV, thereby enhancing absorption of the generated X-rays by said
tumor and minimizing the side effects of radiation therapy on the
patient's normal tissue. .Iaddend. .Iadd.19. A method for the
in-situ treatment of a patient having a tumor, comprising:
(a) providing an X-ray generating source comprising a substantially
cylindrical glass X-ray tube having a diameter less than one inch
and being operable at a voltage level in the range of approximately
10-60 KeV;
(b) positioning said X-ray generating source at a location in close
proximity to said tumor; and,
(c) applying power to said X-ray generating source in said range of
approximately 10-60 KeV, thereby enhancing absorption of the
generated X-rays by said tumor and minimizing the side effects of
radiation therapy on the patient's normal tissue. .Iaddend.
Description
BACKGROUND OF THE INVENTION
1. Field of the Invention
This invention relates to the production and medical use of X-rays,
and more particularly to the production, by X-ray micro-tubes, of
low energy, highly absorable, polychromatic X-rays and to the use
of those X-rays in the treatment of tumors when such X-rays
micro-tubes are placed within, or adjacent to, mammalian bodies in
very close proximity to, or within, tumors.
2. Description of Related Art
Its has been stated that the goal of radiation therapy is to
achieve in a selected treatment volume, a dose distribution of
radiation that provides the patient with maximum tumor control and
the least possible effect on surrounding normal tissues.
(PRINCIPLES AND PRACTICE OF RADIATION ONCOLOGY, C. A. Perez, and L.
W. Brady, Editors; J. B. Lippincott Company, Philadelphia. 1987 Pg.
159).
To achieve that desired goal, many methods have been advanced over
the last 90 or so years which have focused on the use of very high
energy sources of radiation. Current radiation treatment of tumors
involves the use of large external high energy devices such as
X-ray machines, linear accelerators, betatrons, or microtrons, or
the use of very high energy emissions from radiosotopes.
the radioisotopes may be placed in large external machines such as
.sup.60 Cobalt teletherapy machine, or implanted near a tumor site.
Present Applicants are aware of a surgical procedure being utilized
in the treatment of brain tumors in which tiny holes are drilled in
the skull. The surgeon inserts thin tubes with closed bottom ends
into the tumor. Radioactive pellets the size of peas are inserted
into the tubes. The implants deliver strong radiation to the tumor.
They are removed few days later.
The safety of these above-identified high energy radiation sources
has been a constant concern for health professionals. Not only can
the tumor and surrounding normal tissue within the patient be
affected by these high energy radiation sources, but the health
professionals working near the patient can be adversely affected if
adequate safeguards are not taken.
Although the high energy devices have been designed to produce a
maximum of antitumor activity with a minimal effect on a patient's
normal tissues, the side effects of the radiation therapy on the
patient's normal tissue can still be the limiting factor in a
course of therapy.
SUMMARY OF THE INVENTION
The present invention is a method and apparatus for treating tumors
by low energy, highly absorbable, polychromatic X-rays (also called
Bremsstrahlung or White radiation) produced by small X-ray
micro-tubes placed within, or adjacent to, a patient's body in
close proximity to, or within, a tumor. The design of the X-ray
micro-tube can be relatively simple: a miniature X-ray production
source, generally a glass tube a fraction of an inch in diameter
and with a length of approximately one-half of an inch, to several
inches, containing at least an anode and a cathode. The cathode may
be a pointed cold cathode, or a heated filament, and the tube must
be evacuated to, at most, 10.sup.-6 Torr. The target portion of the
anode may be formed of tungsten, as in conventional X-ray tubes.
The glass tube may be surrounded by a plastic envelope so as to
prevent injury to the patient or health professional should the
glass break. A metal jacket containing a window may be placed
around the tube so as to allow the X-rays to travel only in the
direction of the tumor. The X-ray micro-tube may be disposable or
re-sterilized.
The depth of X-ray penetration into tissues can be easily and
accurately controlled by adjusting the voltage applied to the X-ray
micro-tube. Tissue penetration depths within a few centimeters from
the surface of the tube are characteristic of the White radiation
produced by the micro-tubes. This reduces damage to normal tissue
except in the immediate vicinity of the tumor. The X-rays are
produced by applied voltages between 10 kilovolts and 60
kilovolts.
The voltage applied to the X-ray micro-tube may have an operable
frequency between direct current and 1,000,000 cycles per second,
the higher frequencies providing greater patient safety. The
current through the X-ray micro-tube is generally much lower than
that conventionally used and is in the micro-ampere range. Patient
safety is assured by ground fault interrupters and current limiting
circuitry.
The micro-X-ray tubes may be place in-situ by a number of methods,
including, but not limited to: implantation during surgery;
insertion through a normal body orifice; insertion in conjunction
with a fiber-optic scope through a normal body orifice; in
conjunction with a fiber-optic scope through a surgical incision;
insertion through a trocar catheter; or insertion through a
catheter contained within a surgical incision.
Other objects advantages and novel features of the present
invention will become apparent from the following detailed
description of the invention when considered in conjunction with
the accompanying drawings.
BRIEF DESCRIPTION OF THE DRAWINGS
FIG. 1 is a graph of Mass Absorption Coefficient, .mu./.rho.,
versus Energy, MeV, showing the ranges for Photoelectric and
Compton Scattering absorption.
FIG. 2 is a graph illustrating the spectral distribution of the
X-ray energies and relative intensities emitted from a tungsten
anode for various constant d.c. potentials.
FIG. 3 is a graph illustrating examples of Penetration-depth curves
of constant d.c. voltage, KeV vs penetration depth in body tissue,
in centimeters, for monochromatic X-rays, showing distance for
percentage of intensity remaining.
FIG. 4 includes schematic curves showing the changes in the
intensity distribution of the White radiation spectrum from a
tungsten anode at 50 KeV constant potential d.c. after penetrating
through various depths of tissue.
FIG. 5 includes schematic curves showing the changes in the
intensity distribution of the White radiation spectrum from a
tungsten anode at 40 KeV constant potential d.c. after penetrating
through various depths of tissue.
FIG. 6 includes schematic curves showing the changes in the
intensity distribution of the White radiation spectrum from a
tungsten anode at 30 KeV constant potential d.c. after penetrating
through various depths of tissue.
FIG. 7 includes schematic curves showing the changes in the
intensity distribution of the White radiation spectrum from a
tungsten anode at 25 KeV constant potential d.c. after penetrating
through various depths of tissue.
FIGS. 8A to 8J illustrate short X-ray micro-tube designs.
FIG. 8A illustrates a short micro-tube with a heated filament
cathode.
FIG. 8B illustrates a short micro-tube with a cold cathode electron
emitter.
FIG. 8C illustrates a very short micro-tube having the cathode and
anode positioned on the same end, with an internal glass tube for
insulating the anode connecting wire.
FIGS. 8D(a) and 8D(b) illustrate the use of a thin internal anode
film on the inside of the glass tube through which X-rays can
penetrate to produce a cylindrical transmitted X-ray beam.
FIGS. 8E(a) and 8E(b) illustrate the use of a thick internal anode
film on the inside of the glass tube, the resulting back-scattered
X-ray beam producing a hemicylindrical X-ray pattern.
FIGS. 8F(a) and 8F(b) illustrate the use of a thin internal anode
film on the inside of the glass tube resulting in a hemicylindrical
transmitted X-ray pattern.
FIG. 8G illustrates a short micro-tube with a spot-type anode
(thick or thin).
FIGS. 8H(a) and 8H(b) illustrate the use of a thin internal
hemispherical anode film used to produce a hemispherical X-ray
pattern.
FIG. 9 illustrates a long micro-tube design with an enlarged end
for evacuation and sealing.
FIG. 10 is a schematic illustration of a short liquid-cooled
microtube assembly.
FIG. 11 is a schematic illustration of a short metal-jacketed
microtube assembly.
The same elements or parts throughout the figures of the drawings
are designated by the same reference characters.
DETAILED DESCRIPTION OF THE INVENTION
An understanding of X-ray interaction with tumors requires some
background understanding of the physical phenomena involved.
Electromagnetic radiation extends over a very wide range of wave
lengths, i.e. from radio waves (3.times.10.sup.4 m to 5 m),
microwaves (5.times.10.sup.-2 m to 1.times.10.sup.-4 m), infrared
(1.times.10.sup.-4 m 7.times.10.sup.-7 m), visible
(7.times.10.sup.-7 m to 4.times.10.sup.-7 m), ultra-violet
(4.times.10.sup.-7 m to 1.times.10.sup.-8 m), X-rays and gamma rays
(1.times.10.sup.-8 m to 1.times.10.sup.-14 m). The wave lengths of
X-rays are frequently expressed in angstroms (.ANG.) with 1 .ANG.
being equal to 10.sup.-8 cm. These selected wave length bands of
radiation have been classified as ranges which interact with matter
in familiar ways. The value of the wavelength determines the size
of an object with which the electromagnetic radiation will react.
Radio waves will react with large electrical conductors, visible
and ultra-violet light react with the outer shell electrons in
atoms, and X-rays interact with the innermost orbital electrons.
The shorter the wavelength the higher the energy of the radiation.
The reciprocal of the wavelength is the frequency, with the
wavelength commonly being represented by the symbol .lambda., and
the frequency by the Greek letter .nu.. The energy, E, is equal to
.eta..nu., where .eta. is Planck's constant.
The nature and properties of the radiation within the X-ray band
vary with the energy (or wavelength) of the X-ray, just as the
characteristics of the light in the visible range vary with the
wave length of the radiation, the shorter wave lengths appearing as
blue colors and the longer ones as orange-red colors. A brief
discussion of how the different wave lengths of X-rays are produced
and how they differ in their interactions with body tissues
provides an understanding of the uniqueness and value of the
present invention.
Medically useful X-rays are normally produced in evacuated tubes
(usually made of glass) containing two elements, a cathode and an
anode. The cathode is typically a tungsten filament that is heated
to a temperature sufficiently high to cause electrons to reach
velocities permitting them to escape from the filament. The
escaping electrons are attracted to the anode at the opposite end
of the tube (also typically formed of tungsten), which exists at a
high positive potential, commonly in the range of 50,000 to 2
million volts.
The electrons are accelerated during their passage toward the
anode, reaching a high velocity before colliding with the anode and
causing inner shell electrons to be ejected from the tungsten
atoms. When the high energy ejected to be ejected electrons return
to normal positions in electron shells, X-rays are created. The
energy gained by the participating electrons is measured in
electron volts (eV) where e is the electrical charge of the
electron and V is the voltage difference between the cathode and
the anode. This measurement of energy is commonly used in the field
of X-ray diagnosis and therapy. (For the purpose of comparison, one
electron volt of energy per atom is equivalent to 23 kilocalories
per mole of atoms).
In the X-ray range of wave lengths, several different phenomena
occur when matter is exposed to X-rays. These phenomena are known
as Elastic (or Coherent) Scattering, Photoelectric Absorption,
Compton Scatter, and Electron Pair Production (one electron and one
position). When the energy of the photon is less than the binding
energy of the outer shell electrons of the exposed material, the
photons cause the orbiting electrons to oscillate in phase with the
X-ray and thus to emit electromagnetic radiation of the same
frequency as the incident ray. The re-radiated X-rays are scattered
with no absorption in the irradiated matter. This phenomenon,
called Elastic or Coherent Scattering, is of no consequence in
X-ray diagnosis or therapy. Electron Pair Production only occurs at
extremely high energies i.e., above one million electron volts (1
MeV), which is much higher than the energies used in this
invention. Thus, these two phenomena will not be further discussed
herein.
Photoelectric Absorption is the dominant form of X-ray absorption
only in the lowest range of X-rays (e.g. 10 to 60 kilovolts). It
occurs when the energy of the X-ray photons is equal to the energy
binding the innermost shell electrons of the atoms in the exposed
matter. In this case, the X-ray photons interact with the electrons
orbiting closest to the nucleus, causing them to be ejected from
the atoms, and causing the photons to loose all of their energy and
disappear. This reaction cannot occur at X-ray beam energies below
the electron-atom binding energy. It is a maximum when the two
energies are equal, and decreases rapidly with increasing X-ray
energy above the maximum value.
Absorption of higher energy X-ray radiation (e.g. above 100 KeV)
occurs almost entirely by Compton Scattering. This process involves
the interactions of X-ray photons with any of the electrons in the
cloud surrounding the nucleus of the interacting atom. In this
case, an incident photon loses only part of its energy when it
reacts with an electron, which acquires the energy lost by the
photon and is ejected from the electron cloud of the atom. The
resulting photon with diminished energy is scattered and moves
forward at some angle to the line of the incident beam. The energy
of the scattered photon is not absorbed locally within the exposed
material during this event, but the energy acquired by the
participating electron is completely absorbed within the material.
Thus, the energy change of the incident X-ray beam is divided into
two parts, only one of which is directly absorbed.
Between 10 KeV and 100 KeV the absorbed fraction of Compton energy
increases from a low value at 10 KeV, finally reaching a high value
at 100 KeV, thereafter changing by a relatively small amount in the
range between 100 KeV and 1 MeV as shown on the mass absorption
coefficient vs energy curves in FIG. 1. (FIG. 1 is prior art
adapted from "Principles of Radiological Physics" by Robin J.
Wilkes. Second Edition, Pg. 483. Churchill Livingstone, N.Y. 1987.)
In this figure there are curves showing how the absorption of
X-rays by tissue and bone vary with photon energy. Below about 50
KeV energy (in the Photoelectric Absorption range), bones absorb
about 13 times as much of the incident energy as does tissue. Above
about 100 KeV (in the Compton range) the difference is very small,
with bone absorption being only about 1.7 times as much as tissue.
The higher energy range is generally used for cancer therapy.
The photoelectric absorption of 10 KeV X-rays by tissue is very
high, but the absorption coefficient decreases drastically as the
energy is increased. At 50 KeV the absorption coefficient has
decreased to about one percent of the 10 KeV value. In the 10 KeV
to 50 KeV range in which Photoelectric Absorption is dominant.
These lower energy X-rays penetrate only short distances into
tissue and consequently this is not normally considered to be
useful range for either medical diagnosis or therapy. It is common
practice in radiation therapy to remove as much as possible of this
skin damaging radiation from the X-ray beam. This is accomplished
by inserting thin sheets of aluminum or copper in the path of the
beam to absorb as much as possible of the low energy component of
the beam. The present invention makes it possible to utilize some
of the X-rays in this low energy range to treat some easily
accessible cancerous tissue more effectively than the higher energy
methods now in use, but without the sometimes severe damage to
normal body tissue that occurs with currently used practices.
Before entering into the detailed description of the new method and
equipment which are the subject of this invention, a brief
description of some of the well known characteristics and methods
of utilizing X-rays is considered. For example, traditional plots
of absorption coefficients versus X-ray energies, such as
illustrated in FIG. 1, refer to monochromatic X-rays, whereas the
radiation from an X-ray tube actually comprises a broad spectrum of
energies, as illustrated by FIG. 2. (FIG. 2 is prior art adapted
from "X-Ray Metallography" by A. Taylor, pg. 18, John Wiley and
Sons, New York, 1961)
In this figure relative (X-ray) intensity is plotted against
(X-ray) energy for various constant d.c. tube voltages, and
spectral curves of Bremsstrahlung or White radiation are thus
produced. The shapes of these curves are affected by tube
potential. The shapes of the curves are also affected by the nature
of the power source. Depending upon the design of the power supply,
the output may be constant d.c., a half-wave rectified a.c., or a
full wave rectified a.c.. The maximum intensities and the X-ray
energy distribution will be different for each type of power
source. Also, it must be remembered that the mass absorption
coefficient for tissue varies sharply with the energy (or wave
length) of the X-ray, as shown in FIG. 1. In reality, then, the
effective absorption coefficient for a beam consisting of a broad
spectrum of X-rays is the fraction of the beam that is absorbed
during the passage of the beam through one centimeter of material
and is composed of two parts: one being the component contributed
by Photoelectric Absorption, and the other being due to the Compton
effect.
FIG. 3 shows examples of approximate depth of penetration curves
(i.e. the distance through material which causes a specific
decrease in the X-ray intensity to, for example, 1/2, 1/4, 1/10
etc. of the initial beam intensity. I is the intensity, while
I.sub.o refers to the initial intensity). The values shown in this
figure are for monochromatic radiation. From accurate plots of the
mass absorption coefficients, .mu./.rho., versus energy curves
shown in FIG. 1, the linear absorption coefficient, .mu., needed
for calculating the decrease in X-ray intensity versus penetration
distance, as shown in FIG. 3, can be obtained by assuming that the
intensity, .rho., of tissue used for defining mass absorption
coefficient, .mu./.rho., is equal to 1. FIG. 3 highlights the fact
that in the 10-50 KeV range X-rays simply do not penetrate very
far.
As an example of the method used to obtain the curves in FIG. 3,
assuming the voltage to be 50 KeV, the value of I/I.sub.o =0.5, and
.mu., the absorption coefficient=0.24, the interrelationship of
these parameters is given by I/I.sub.o =e.sup.-.mu.x where x is the
penetration distance below the surface of the absorbing material (x
is a negative number). Substituting the numbers given above in the
equation yields a value of x of 2.87 cm (corresponding to the value
for the x=2.87 cm at 50 KeV and 50% absorption in FIG. 3).
The actual radiation being emitted from the anode on an X-ray tube
with, for example, 50 kilovolts applied potential, is actually a
whole spectrum of radiant energies, producing what is commonly
called "white" radiation. This is illustrated in FIG. 2, which
shows the distribution of energies (in KeV) for a tube with a
tungsten anode, when exposed to various constant tube voltages. In
spectra of this sort, each individual value would have an
absorption coefficient differing from those of the other energies
in the spectrum. The variation in absorption coefficients within
the spectrum must be taken into account when effective absorption
is calculated for the beam. In the lower end of the voltage range,
i.e., below about 50 KeV tube voltage, the absorption coefficients
may vary by more than 100 to 1 (see FIG. 1). However, in the higher
voltage range currently in general use (i.e., about 100 KeV to 1
MeV) the absorption coefficient is relatively constant. With white
radiation in the energy range of this invention, then, both the
absorption coefficient and the X-ray intensity are strongly
dependent upon the energy of the photons involved. For example, the
mass absorption coefficient for tissue varies from about 0.33 for
37 KeV to 3.3 at 12 KeV.
FIGS. 4, 5, 6, and 7 show intensity decreases in the initial
intensity values for white radiation from a tungsten anode X-ray
tube at different tube potentials, as the beam penetrates into
tissue. The shape of the Relative Intensity versus Energy (KeV)
curve changes as the distance from the outer surface of the tissue
being penetrated is increased. At each X-ray energy level the
absorption coefficient differs, as shown in FIG. 1, with a much
higher fraction of the lower X-ray energy components of the white
radiation beam being absorbed than is the case for the higher
energy components. The curves in FIG. 4 illustrate this effect at
various depths in tissue.
Examples of how and why the shapes of the white radiation energy
distribution curves change at different depths in tissue will be
helpful in understanding the curves plotted in FIGS. 4, 5, 6, and
7. FIG. 4 illustrates examples of the nature of the changes that
occur at a high energy value within the 50 KeV generated white
radiation spectrum, to those of a low energy value of the same
spectrum, by comparing initial intensities on the I.sub.o curve to
intensities on the same depth in tissue curve.
In this example, let the high energy case by chosen as 42 KeV and
the low energy be selected as 20 KeV. For the high energy case (42
KeV), let: I.sub.o =5.5 (from FIG. 4), .mu.=0.32 (from FIG. 1), and
x=1 cm.
By rearranging the formula, I/I.sub.o =e.sup.-.mu.x, we can solve
for the value of I at 1 cm using our parameters:
which is shown on FIG. 4.
Now, for the low energy case (20 KeV), let: I.sub.o =9.5 (from FIG.
4), .mu.=0.87 (from FIG. 1), and x=1 cm.
In this case, we have:
which is shown on FIG. 4.
This demonstrates that for equal intensities at a penetration depth
of one centimeter, the initial intensity, I.sub.o, had to be only
5.5 for the higher voltage (42 KeV) while the I.sub.o value for 20
KeV had to be much higher at 9.5. Thus, FIG. 4 further highlights
the advantage of the use of these low energy X-rays. They do not
penetrate very far. The low energy portions of the spectra
illustrated in FIG. 4 are almost entirely absorbed at 3 cm.
FIGS. 5-7 are generated in the same manner as FIG. 4 but for
constant potential d.c. voltages of 40 KeV, 30 KeV, and 25 KeV,
respectively. Information of this kind is necessary for determining
suitable voltages for treating tumors of different varieties and
sizes while minimizing damage to nearby tissues.
The micro-tubes are similar in principle to standard X-ray tubes
except that they are much smaller and require only a small fraction
of the tube current required in conventional commercial machines
(i.e. microamperes vs milliamperes). The physical size of a tube
can be a fraction of an inch in diameter and with a length as small
as one-half of an inch, to as long as several inches. A variety of
useful tube designs is possible.
Referring now to FIG. 8a, a schematic illustration of an embodiment
having a filament cathode is shown, designated generally as 20. An
evacuated glass tube 22 contains a stable vacuum of at least
10.sup.-6 Torr. A heated filament cathode 24 (preferably a small
tungsten filament) and an anode 26 are provided which are connected
to an appropriate power source (as will be discussed below). The
anode 26 can be made of any one of a number of different metals
typically used, but tungsten is preferred (as it is conventional
commercial tubes).
Referring to FIG. 8b, a second type of X-ray tube that is suitable
for micro-tube use is schematically illustrated, designated
generally as 28. Tube 28 is a cold emission (or field emission)
cathode tube (which has no filament). The electrons are emitted
from a sharply pointed electrode 30, preferably formed of tungsten.
A very high potential gradient develops between the sharply pointed
tip of the electrode and the anode 32 when a high voltage is
applied across the X-ray tube 28. For micro-tube use in radiation
oncology, use of a cold electron emitter tube has some advantages.
This type of tube is simpler to make in smaller sizes than the
heated filament type.
The depth of penetration of the present X-ray microtubes can be
easily controlled by varying the tube voltage. The total desired
radiation exposure can be controlled by selecting an appropriate
time of exposure. A great advantage of the micro-tube is that it
can be placed on or very near the surface of the tumor, or within
the tumor, so that radiation damage to normal tissue is
minimized.
The source to tumor distance for the micro-tubes, therefore, is
extremely small compared with the source to skin distance (SSD) of
20 to 50 centimeters with the X-ray units now in common use. Since
the intensity of the X-ray beam varies inversely with the square of
the distance from the beam source, for a given tube voltage, the
same effective intensity of the X-ray beam at the site of a tumor
can be produced by the micro-tube with about one one-thousandth of
the current required for the proper operation of a large external
tube. Therefore, the present invention operates in the low
microampere range, rather than the low milliampere range required
for currently used large tubes.
Micro-tubes are relatively inexpensive and may be manufactured to
be re-sterilized or disposable. The exterior surface of the tube is
preferably covered with a thin tough biocompatible plastic
material, as will be described below, to guard against damage to
handlers or patients should accidental breakage of the glass tube
occur. The plastic tube cover can also have a built-in water
coolant jacket to dissipate the small amount of heat generated by
the tube (operating at a small fraction of a watt).
The power supply required is relatively simple and inexpensive
because of the low current required (microamperes vs conventionally
used milliamperes) and because of the relatively low tube voltages
required (generally less than 60 kilovolts compared 60 kilovolts to
one million volts for conventional equipment). Aside from the
micro-tube itself, only state of the art equipment is necessary.
However, one difference in detail is necessary. Conventional 60
cycle a.c. destroys the many normal nerve functions. The inventive
concepts of the present invention provide for use of a frequency
that would be sufficiently high so that the normal nerve functions
of the body would not be affected if an inadvertent contact of the
high voltage lead with body tissue did occur. The use of high
frequency currents in electrosurgical cutting and coagulation
machines is common and has long been known to be safe (see for
example, U.S. Pat. No. 3,699,967, entitled "Electrosurgical
Generator", and Chapter 3, Electrosurgery, Handbook of Biomedical
Engineering, 1988, Academic Press). Furthermore, use of such high
frequencies provides the ability for the patient to act as a
conduit for return of the anode current back to the power supply if
that anode is not directly connected to the power supply.
FIG. 8C illustrates the placement of the anode 34 and the cathode
36 on the same side of a microtube 38 to accomplish a reduced
length. An internal glass tube 40 is used to support the anode 34
and to insulate the anode connecting wire 42.
The micro-tubes of the present invention may be manufactured in a
variety of different ways to optimize their use. For example, in
FIG. 8D(a) a glass micro-tube 42 is illustrated with a thin
internal anode film 44 formed on its inner surface (preferably
vacuum deposited tungsten). An axially extending filament cathode
46 is provided. Thus, when operated, a cylindrical X-ray pattern,
illustrated by the arrows 48 in FIG. 8D(b), results. This design is
particularly useful if the microtube is desired to be inserted near
the center of the tumor.
FIGS. 8E(a) and 8E(b) illustrate a relatively thick film anode 50
deposited or otherwise formed on portions of the inside of the
glass tube 52. This results in a backscattered X-ray beam which
produces a hemicylindrical X-ray pattern, as illustrated by arrows
54. The X-ray pattern is established at portions of the micro-tube
which do not have the thick film formed thereon.
In the micro-tube illustrated in FIG. 8F(a) and 8F(b) a thin anode
film 56 is formed on only a portion of the micro-tube 58. This
results in a hemicylindrical X-ray pattern 60.
An alternate anode design is illustrated in FIG. 8G, the micro-tube
being designated as 62. In this instance, the anode 64 is a spot
type of thin or thick film.
In FIGS. 8H(a) and 8H(b) a hemispherical X-ray pattern 66 results
from formation of an anode film 68 near the end of the micro-tube
70.
The short micro-tubes illustrated in FIGS. 8A-8H are typically from
one-fourth inch to two inches in length, preferably approx. 1/2".
Diameters may range from 1/8" to 1", preferably 1/4". As noted,
these microtubes may be placed in-situ by a number of methods,
including, implantation during surgery; insertion through a normal
body orifice; insertion in conjunction with a fiber-optic scope
through a normal body orifice; insertion in conjunction with a
fiber-optic scope through a surgical incision; insertion through a
trocar catheter; or insertion through a catheter contained within a
surgical incision. The micro-tubes may also be placed adjacent to
the body next to the skin.
Longer micro-tubes may alternately be used which may be up to
several inches (i.e. 2"-8") in length. FIG. 9 illustrates a
schematic of a design of such a long micro-tube 72. The lead wire
74 for the cathode 76 and the lead wire 78 for the anode 80 connect
to a power supply (not shown). Long micro-tube 72 is particularly
useful in the brain and is made thin, for example, in the range of
1/8" to 1/4" in diameter. End 82 is enlarged and extends outside of
the body, serving as a "compass" for accurately rotating the
micro-tube and directing the X-rays in the desired manner. It is
understood that the various features shown in the previous Figures
may be implemented in the longer tubes of FIG. 9. Further, it is
understood that FIGS. 8 and 9 are meant only to be schematic
representations of possible micro-tube designs. Obviously,
biocompatible safety shields would be utilized to enclose the tubes
in actual applications.
The principles of the present invention are preferably implemented
with the micro-tubes being used as part of a mechanically shielded
and electrically insulated assembly. Referring now to FIG. 10, such
an implementation in the form of a short liquid-cooled micro-tube
assembly, designated generally as 84, is illustrated. Liquid-cooled
micro-tube assembly 84 includes a filament cathode 86 supported by
a filament support structure 88 within an evacuated glass tube 90.
Similarly, an anode 92 is supported by another filament support
structure 94 within the evacuated glass tube 90. Glass tube 90 may
be formed as described in the above-discussion regarding FIGS. 8
and 9. Glass tube 90 is positioned within a liquid coolant chamber
96 which is supplied by coolant hoses 98,100. (Water would be a
suitable coolant. The walls of the coolant chamber may be formed
of, for example, glass or plastic). Coolant chamber 96 is, in turn,
enclosed within the main housing 102 of the micro-tube assembly 84.
Housing 102 is preferably formed of plastic. Filament lead wires
104,106, anode lead wire 108, and coolant hoses 98.100, extend
through the main housing 102. These five elements are preferably
radially spaced and separated by walls to confine any leaks to a
specific portion of the assembly. Dashed lines 110 schematically
illustrate these walls. Additionally, approximately water seals 112
are provided.
A metal-jacketed micro-tube assembly, designated generally as 114,
is illustrated in FIG. 11. Assembly 114 includes a cathode 116
supported by a cathode support structure 118 within an evacuated
glass tube 120. Similarly, an anode 122 is supported by another
support structure 124 within the evacuated glass tube 120. Glass
tube 120 is contained within a metal jacket 126. Tube 120 may be
formed as described in the above-discussion regarding FIGS. 8 and
9. A window 128 is provided in the metal jacket 126 for directing
the radiation in the desired manner. Anode and cathode cables
130,132 including lead wires are provided for attachment to an
external power source (not shown).
The power supply needed for X-ray micro-tube operation is unique in
that it is a low energy device that can be made easily portable and
is less costly to make than those now supplied with deep therapy
equipment, which require much higher levels of energy. Modern
electronic designs and equipment capable of providing the currents
and voltages needed for the operation of the micro-tubes are state
of the art, and a variety of designs are available. Another
essential feature of the invention is that the power supply circuit
must contain a rapidly acting safety circuit interrupter that will
immediately operate should anything happen to cause the tube
current to suddenly increase to, for example, one milliampere,
which is still very safe for a patient but undesirable for
micro-tube operation.
Obviously, many modifications and variations of the present
invention are possible in light of the above teachings. It is
therefore to be understood that, within the scope of the appended
claims, the invention may be practiced otherwise than as
specifically described.
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