BACKGROUND OF DISCLOSURE
The present invention relates to a miniaturized, low power, programmable
x-ray source for use in delivering substantially constant or intermittent levels
of x-rays to a specified region. More specifically, the invention relates to an
apparatus for delivering a uniform x-ray flux to an interior surface of a body cavity.
Most conventional medical x-ray sources are large, fixed position
machines. Generally, the head of the x-ray tube is placed in one room and the control
console in an adjoining area, with a protective wall, equipped with a viewing window,
separating the two. The x-ray tube typically is approximately 20 to 35 centimeters
(cm) long, and approximately 15 cm in diameter. A high voltage power supply is housed
within a container located in a comer of the room containing the x-ray tube. Patients
are brought to the machine for diagnostic, therapeutic, or palliative treatment.
Diagnostic x-ray machines are typically operated at voltages below
150 kilovolts (kV), and at currents from approximately 25 to 1200 milliamps (mA).
By contrast, the currents in therapeutic units typically do not exceed 20 mA at
voltages which may range above 150 kV. When an x-ray machine is operated at nominal
voltages of 10 to 140 kV, the emitted x-rays provide limited penetration of tissue,
and are thus useful in treating skin lesions. At higher voltages (approximately
250 kV), deep x-ray penetration is achieved, which is useful in the treatment of
major body tumors. Super voltage machines, operable in the 4 to 8 megavolt (MV)
region, are used to ablate or destroy all types of tumors, except superficial skin
A conventional x-ray tube includes an anode, grid, and cathode assembly.
The cathode assembly generates an electron beam which is directed to a target, by
an electric field established by the anode and grid. The target in turn emits x-ray
radiation in response to the incident electron beam. The radiation absorbed by a
patient generally is that which is transmitted from the target in the x-ray tube
through a window in the tube, taking into account transmission losses. This window
typically is a thin section of beryllium, or other suitable material. In a typical
x-ray machine, the cathode assembly consists of a thoriated tungsten coil approximately
2 mm in diameter and 1 to 2 cm in length which, when resistively heated with a current
of 4 amps (A) or higher, thermionically emits electrons. This coil is surrounded
by a metal focusing cup which concentrates the beam of electrons to a small spot
on an opposing anode which also functions as the target. In models having a grid,
it is the grid which both controls the path of the electron beam and focuses the
The transmission of an electron beam from cathode to anode is influenced
by electron space charge forces which tend to become significant in conventional
x-ray machines at currents exceeding 1 A. In such conventional machines, the beam
is focused on the anode to a spot diameter ranging anywhere from 0.3 to 2.5 millimeters
(mm). In many applications, most of the energy from the electron beam is converted
into heat at the anode. To accommodate such heating, high power medical x-ray sources
often utilize liquid cooling and a rapidly rotating anode, thereby establishing
an increased effective target area, permitting a small focal spot while minimizing
the effects of localized heating. To achieve good thermal conductivity and effective
heat dissipation, the anode typically is fabricated from copper. In addition, the
area of the anode onto which an electron beam is incident requires a material of
high atomic number for efficient x-ray generation. To meet the requirements of thermal
conductivity, effective heat dissipation, and efficient x-ray generation, a tungsten
alloy typically is embedded in the copper.
In use, the total exposure from an x-ray source is directly proportional
to the time integral of the electron beam. During relatively long exposures (e.g.
lasting 1 to 3 seconds), the anode temperature may rise sufficiently to cause it
to glow brightly, accompanied by localized surface melting and pitting which degrades
the radiation output. However, thermal vaporization of the tube's coiled cathode
filament is most frequently responsible for conventional tube failure.
While the efficiency of x-ray generation is independent of the electron
beam current, it is highly dependent on the acceleration voltage. Below 60 kV, only
a few tenths of one percent of the kinetic energy from an electron is converted
to x-rays, whereas at 20 MV that conversion factor rises to 70 percent. An emitted
x-ray spectrum is composed in part of discrete energies characteristic of transitions
between bound electron energy levels of the target element. The spectrum also includes
an x-ray energy continuum, known as bremsstrahlung, which is caused by acceleration
of the beam electrons as they pass near target nuclei. The maximum energy of an
x-ray cannot exceed the peak energy of an electron in the beam. Further, the peak
of the bremsstrahlung emission curve occurs at approximately one-third the electron
Increasing the electron current results in a directly proportional
increase in x-ray emission at all energies. However, a change in beam voltage results
in a total x-ray output variation approximately equal to the square of the voltage,
with a corresponding shift in peak x-ray photon energy. The efficiency of bremsstrahlung
radiation production increases with the atomic number of the target element. The
peak output in the bremsstrahlung curve and the characteristic spectral lines shift
to higher energies as the atomic number of the target increases. Although tungsten
(Z=74) is the most common target material used in modern tubes, gold (Z=79) and
molybdenum (Z=42) are used in some specialty tubes.
One disadvantage of most x-ray devices used for therapy is the high
voltage, and consequent high energy radiation required when directed to soft tissue
within or beneath bone. One example is in directing x-rays to areas of the human
brain, which is surrounded by bone. High energy x-rays are required to penetrate
the bone, but often damage the skin and brain tissue between the radiation entry
site and the tumor. Another example in radiation therapy is in directing the x-rays
to soft tissue located within the body cavity, couched among other soft tissue,
or within an internal calciferous structure. Present high-voltage x-ray machines
are limited in their ability to selectively provide desired x-ray radiation to such
Another disadvantage of the conventional high voltage x-ray sources
is the damage caused to skin external to the affected organ or tissue. Therefore,
prior art high voltage x-ray sources often cause significant damage not only to
the target region or tissue, but also to all surrounding tissue between the entry
site, the target region, and the exit site, particularly when used for human tumor
therapy. However, since present devices apply x-ray radiation to target regions
internal to a patient from a source external to the target region, such incidental
tissue damage is practically unavoidable.
Conventional radiation treatment of the soft tissue that lines body
cavities, such as the bladder, vagina and cervix, urethra, uterus, colon and rectum,
involves application of x-radiation from an extracorporeal source. Consequently,
such techniques of radiation therapy have the disadvantage that they necessarily
radiate areas of the patient between the radiation entry site, the target tissue,
and the exit site, causing damages to such tissue.
Conventional methods of radiation treatment for body cavities also
have the further disadvantage of failing to provide the ability to establish a uniform
dose of radiation to the target tissue. In some cases, it is desirable that radiation
treatment of the tissue lining a body cavity should provide the same dose of radiation
to every segment of the tissue, i.e., a uniform, or other desired, dose. In other
cases, specifically contoured non-uniform doses may be desired. The prior art x-ray
sources cannot accomplish this for interior body cavities. As used herein, the term
"uniform dose" refers to an isodose contour, i.e., a surface over which the flux
density is substantially constant.
Some of these disadvantages can be overcome through the use of miniaturized
low power x-ray sources, such as the one described in the above-referenced U.S.
Patent No. 5,153,900 issued to Nomikos et al. These sources can be inserted into,
and activated from within, a patient's body. Thus, these sources can generate x-rays
local to the target tissue. When such x-ray sources are used to treat the tissue
lining a body cavity, the x-rays need not pass through the patient's skin, bone
and other tissue prior to reaching the target tissue. However, even utilizing these
sources there is no previously known method of providing a uniform, or other desired,
dose of radiation to the target tissue, particularly where the geometry of the target
region is not fixed, for example, as in the bladder which has a flexible inner wall
without a well-defined shape.
By way of example, some miniature sources of the type disclosed in
U.S. Patent No. 5,153,900 generally act as point sources of x-ray radiation. Therefore,
the strength of the radiation field decreases uniformly in air with approximately
the square of the distance from the source (i.e., 1/R2). Since body cavities
are not generally spherically symmetrical, a point source within a body cavity will
not deliver a uniform dose of radiation to the tissue lining the cavity.
WO-A-92/22350 discloses a completely implantable apparatus which is
provided for treatment of tissue surrounding a cavity left by surgical removal of
a brain tumor. The apparatus includes an inflatable balloon constructed for placement
in the cavity. A subcutaneously implantable treatment fluid receptacle is provided
for receiving a transdermal injection of a treatment fluid. A catheter connects
the receptacle to the inflatable balloon Various embodiments provide for simultaneous
application of heat therapy and/or radiation therapy and/or chemotherapy to the
remaining tissue surrounding the cavity from which the tumor was removed.
WO-A-94/08350 discloses an apparatus according to the preamble of
claim 1. The present invention is characterized by the features of the characterizing
portion of claim 1 with optional features recited in the dependent claims.
The present invention to provides an apparatus for delivering a uniform
dose of radiation to the tissue that lines a body cavity. The apparatus may include
a miniature low power x-ray source, for delivering a uniform or other desired dose
of radiation to the tissue that lines a body cavity.
Other advantages provided by embodiment of the present invention will
become apparent upon consideration of the appended drawings and description thereof.
SUMMARY OF THE INVENTION
In one form, the invention is a kit for applying a predetermined x-ray
flux to an interior surface of a body cavity including an x-ray source, intended
for insertion into the cavity, and an inflatable balloon assembly.
The x-ray source includes a tubular element, a beam source, and a
controller. The tubular element has a target end that contains an electron activated
x-ray source. The beam source is disposed near a beam source end of the tubular
element and is operative to generate an electron beam. The controller selectively
activates the beam source such that the electron beam is incident on the target
end. The target end of the x-ray source is positioned within the body cavity-to-be-irradiated.
The balloon assembly includes an inflatable balloon positioned at
the target end of the tubular element of the x-ray source. When inflated, the balloon
defines an interior region adjacent to the target end.
With this configuration, the balloon may be inflated so that it is
in contact with the lining of the body cavity, and displaces that cavity to define
a desired shape for that lining. The target end is positioned within the interior
region defined by the inflated balloon. By way of example, the balloon may be positioned
with a bladder of a patient and, when inflated, it may define a spherical interior
region and the target may be positionable at the center of the sphere. Also according
to this aspect, the electron activated x-ray source may generate an x-ray field
having an isodose contour coincident with the surface of the inflated balloon, thereby
providing a uniform dose to the lining of the bladder.
Thus, with the invention, a surface of a body cavity is conformed
to a predetermined contour and then the x-ray source is adjusted to establish a
uniform dose at that surface (i.e., an isodose contour, over which the flux density
is substantially constant). The flux density decreases with distance from the source
beyond the cavity lining, permitting treatment of the lining surface and diminishing
effects in tissue beyond that lining.
BRIEF DESCRIPTION OF DRAWINGS
The foregoing and other objects of this invention, the various features
thereof, as well as the invention itself, may be more fully understood from the
following description, when read together with the accompanying drawings in which:
- FIGURE 1 is a perspective view of a kit embodying the present invention;
- FIGURE 2 is a schematic block diagram of the x-ray source of FIGURE 1;
- FIGURE 3 is a cross-sectional view of the end of a probe having an alternate
target assembly which includes an x-ray shield and x-ray target for producing a
stable and reproducible source of x-rays;
- FIGURES 4A-4F show examples of various isodose contours that can be achieved
with the invention;
- FIGURE 5 shows the probe and balloon assembly of the kit of FIGURE 1 with its
- FIGURE 6 shows another embodiment of the probe and balloon assembly of the kit
of the Invention with its balloon inflated;
- FIGURE 7 shows another embodiment of the invention in which the x-ray probe
is inserted to be proximal to a wall of the balloon; and
- FIGURES 8A and 8B are cross-sectional views of a flexible probe which incorporates
a photoemitter located within the target assembly.
Like numbered elements in each FIGURE represent the same or similar
DESCRIPTION OF THE PREFERRED EMBODIMENTS
The present invention is a relatively small, electron-beam activated,
low power x-ray kit. The kit may be used for medical purposes such as therapeutic
radiation of the soft tissue lining body cavities, for example, the bladder or other
Generally, the apparatus of the present invention includes an electron-beam
(e-beam) activated x-ray source which operates at relatively low voltages, i.e.
in the range of approximately 10 kV to 90 kV, and relatively small electron beam
currents, i.e. in the range of approximately 1 nA to 1mA. At those operating voltages
and currents, the x-ray source may be made quite small and be adapted for implantation
in medical therapeutic applications. Adequate tissue penetration and dosage may
be attained by locating the x-ray source adjacent to or within the region to be
irradiated. Thus, the x-rays are emitted from a well-defined, small source located
within or adjacent to the region to be irradiated.
FIGURE 1 shows an x-ray kit embodying the present invention. That
kit includes an x-ray source 10 and a balloon assembly 400. A suitable x-ray source
10 is described in detail in the above referenced U.S. Patent No. 5,153,900, entitled
"Miniaturized Low Power X-Ray Source. The balloon assembly 400 is described in detail
below in conjunction with FIGURES 5, 6 and 7.
X-ray source 10 includes a housing 12 and an elongated cylindrical
probe 14 extending from housing 12 along a reference axis 16 and having a target
assembly 26 at its distal end. The housing 12 encloses a high voltage power supply
12A. The probe 14 is a hollow tube having an electron beam generator (cathode) 22
adjacent to the high voltage power supply 12A. Cathode 22 is located in close proximity
to an annular focusing electrode 23 typically at nearly the same potential as the
cathode 22. An annular anode 24 is positioned approximately 0.5 cm or more from
the annular focusing electrode 23. A hollow, tubular probe 14 extends along the
same axis as the cathode, grid, and the hole in the anode. Probe 14 is integral
with the housing 12 and extends toward a target assembly 26. In various embodiments,
parts of the probe 14 may be selectively shielded to control the spatial distribution
of x-rays. In addition, the probe 14 may be magnetically shielded to prevent external
magnetic fields from deflecting the beam away from the target.
The electron beam generator 22 may include a thermionic emitter (driven
by a floating low voltage power supply) or a photocathode (irradiated by an LED
or laser source). The high voltage power supply establishes an acceleration potential
difference between the cathode of generator 22 and the grounded anode 24 so that
an electron beam is established along the reference axis 16, through the center
hole of the anode and to the target assembly 26, with the region between anode 24
and the target assembly 26 being substantially field free. The beam generation and
acceleration components are adapted to establish a thin (e.g. 1 mm or less in diameter)
electron beam within the probe 14 along a nominally straight axis 16.
In a preferred embodiment, the probe 14 is a hollow, evacuated cylinder
made of a beryllium (Be) cap and a molybdenum-rhenium, (Mo-Re), molybdenum (Mo)
or mu-metal body and a stainless-steel base extension. The cylinder has a length
which is determined in view of the cavity to be treated. For example, for use with
a bladder, the probe may be 40 cm in length, with an interior diameter of 4 mm,
and an exterior diameter of 5 mm. For use with other cavities, different geometries
may be used.
The target assembly 26 includes an emission element consisting of
a small beryllium (Be) window 26A coated on the side exposed to the incident electron
beam with a thin film or layer 26B of a high-Z element, such as tungsten (W), uranium
(U) or gold (Au). By way of example, with electrons accelerated to 30 keV-, a 2.2
micron thick tungsten film absorbs substantially all the incident electrons, while
transmitting approximately 95% of any 30 keV-, 88% of any 20 keV-, and 83% of any
10 keV- x-rays generated in that layer. In the preferred embodiment, the beryllium
widow 26A is 0.5 mm thick with the result that 95% of these x-rays generated in
the layer 26B in directions normal to and toward the window, and having passed through
the tungsten target layer 26B, are then transmitted through the beryllium window
26A and outward at the distal end of probe 14.
The apparatus of FIGURE 1 is normally used in a manner where only
the probe 14 is inserted into a patient while the housing remains outside the patient.
In this form, some or all of the various elements shown within housing 12 may alternatively
be remotely located.
In one embodiment of the apparatus as shown in FIGURE 1, the main
body of the probe 14 can be made of a magnetically shielding material such as a
mu-metal. Alternatively, the probe 14 can be made of a non-magnetic metal, preferably
having relatively high values for Young's modulus and elastic limit. Examples of
such material include molybdenum, rhenium or alloys of these materials. The inner
or outer surface of probe 14 can then be coated with a high permeability magnetic
alloy such as permalloy (approximately 80% nickel and 20% iron), to provide magnetic
shielding. Alternatively, a thin sleeve of mu-metal can be fitted over, or inside
of, the probe 14. The x-ray apparatus 10 can then be used in environments in which
there are low-level dc and ac magnetic fields due to electrical power, the field
of the earth, or other magnetized bodies nominally capable of deflecting the electron
beam from the probe axis.
In the above-described embodiments, the x-ray emission element of
the target assembly 26 is adapted to be adjacent to or within the region of a patient
to be irradiated. The proximity of the emission element to the targeted region,
e.g. the body cavity, eliminates the need for the high voltages of presently used
machines, to achieve satisfactory x-ray penetration through the body wall to the
body cavity. The low voltage also concentrates the radiation in the targeted tissue,
and limits the damage to surrounding tissue and surface skin at the point of entry.
Generally, when treating a body cavity with radiation therapy, it
is desirable to uniformly radiate the entire surface of the soft tissue lining the
cavity such that an isodose contour is coincident with the surface of the body cavity.
An isodose contour is a surface in which the absorbed radiation energy is equal
at every point on the surface.
A preferred method of uniformly radiating a body cavity, such as the
bladder of a patient, is to use a device to first stretch the cavity into a spherical
shape, and then position an omnidirectional x-ray generating probe tip (i.e., a
point source) at the center of the cavity. With that configuration, an isodose contour
can be established which is coincident with the surface of the body cavity. One
device useful for stretching a body cavity to a spherical shape is an inelastic
FIGURE 2 is a block diagram representation of the x-ray source apparatus
10 shown in FIGURE 1. In that preferred configuration, the housing 12 is divided
into a first portion 12' and a second portion 12". Enclosed within the first housing
portion 12' is a rechargeable battery 12B, a recharge network 12D for the battery
12B, which is adapted for use with an external charger 50, and a telemetry network
12E, adapted to be responsive to an external telemetry device 52 to function in
the manner described below. That portion 12' is coupled by cables to the second
housing portion 12". The second housing portion 12" includes the high voltage power
supply 12A, controller 12C and the probe 14, as well as the electron beam generating
element 22. In the illustrated embodiment, the electron beam generator includes
a photoemitter 22 irradiated by a light source 56, such as a diode laser or LED,
powered by a driver 55. The light is focused on the photoemitter 22 by a focusing
In the illustrated embodiment, device 52 and network 12E cooperate
to permit external control (dynamic or predetermined) control over the power supply
12A and temporal parameters. Controller 12C may directly be used to control operation
in which case there is no need for network 12E.
In an alternate form, the beam generator may include a thermionic
emitter 22 driven by the power supply 12A. In operation of that form, power supply
12A heats the thermionic emitter 22, which in turn generates electrons which are
then accelerated toward the anode 24. The anode 24 attracts the electrons, but passes
them through its central aperture toward the target assembly 26. The controller
12C controls the power supply 12A to dynamically adjust the cathode voltage, the
electron beam current, and temporal parameters, or to provide pre-selected voltage,
beam current, and temporal parameters. Other suitable power supply configurations
are disclosed in U.S. Patent No. 5,153,900 and in Patent No. WO 94/08350.
Incident electrons generally cause target 26 to act as a point source
of x-rays. However, further specificity in treating the surfaces body cavities,
and of tumors on or adjacent to such surfaces, may be achieved by tailoring the
target and shield geometry and material at the emission site. This tailoring facilitates
the control of energy and the spatial profile of the x-ray emission to ensure more
homogenous distribution of the radiation throughout the targeted body cavity. This
tailoring is described in detail in the above referenced U.S. Patent No. 5 422 926,
entitled "X-ray Source with Shaped Radiation Pattern".
The x-ray spatial distribution can also be shaped by incorporating
an x-ray transmissive shield, having a variable thickness profile, into the target
assembly 26. FIGURE 3 shows a probe 14 having an alternate target assembly 126,
for use with the x-ray apparatus 10 shown in FIGURE 1, which incorporates such a
shield. In the illustrative embodiment, the probe 14 is substantially similar to
the probe 14 shown in FIGURE 1, except for the target assembly 126. Target assembly
126 includes a probe tip 126A made of a material (e.g. Be) which is nearly transparent
to x-rays, and an x-ray target 126B for generating a source of x-rays upon irradiation
with an electron beam, attached to the probe 14 along a probe axis 16 at the end
distal to the cathode 22 and anode 24 (shown in FIGURE 1). In the preferred form,
the outer surface of the probe tip 126A is convex, and preferably hemispherical,
as in the illustrated embodiment, although other convex shapes can be used. The
target assembly 126 is fabricated such that the outer diameter of the probe tip
126A is less than the outer diameter of the probe 14. A variable thickness x-ray
shield (or shadow mask, as it is sometimes referred to in the art) 128 and an underlying
shield carrier 128A are positioned over the probe tip 126A of the target assembly
126. At the junction of the target assembly 126 and probe 14, the outer diameter
of the target assembly 126 substantially matches that of probe 14.
The x-ray shield 128 is made from a material which is not completely
x-ray transparent (i.e. at least partially x-ray absorptive, such as heavy metals),
and is supported by the shield carrier 128A. The x-ray flux from any point of the
target assembly 126 is dependent in part upon the thickness of the x-ray shield
128 along an axis extending from the target 126B and passing through that point.
Thus, in accordance with an embodiment of the invention, a selective restriction
in thickness of the x-ray shield 128 is used to generate spatially-variable x-ray
FIGURES 4A-4F depict examples of various isodose contours that can
be achieved with the present invention. Specifically, FIGURE 4A shows the probe
14 adapted to deliver isodose contours which form a sphere of radiation 300 centered
about the probe tip 126. FIGURE 4B shows the probe 14 adapted to deliver a sphere
of radiation 302, wherein the probe tip 126 is offset from the center of the sphere
302. FIGURE 4C shows the probe 14 having a tip 126 adapted to deliver a radiation
field in the shape of an oblate ellipsoid (i.e., a "pancake" shape), as shown in
perspective at 304A and looking along axis 305 at 304B. FIGURE 4D depicts the probe
14 having a tip 126 adapted for delivering a radiation field in the shape of a prolate
ellipsoid (i.e., a "cigar" shape), as shown in perspective at 306A and along axis
307 at 306B. As shown in FIGURE 4D, the probe 14 enters the ellipsoid 306A along
its minor axis. FIGURE 4E shows the tip 126 also adapted for delivering a radiation
field in the shape of a prolate ellipsoid. The ellipsoid is shown in perspective
at 308A and along axis 309 at 308B. As can be seen, the probe 14 enters the ellipsoid
308A along its major axis. FIGURE 4F depicts the probe tip 126 adapted for delivering
an asymmetric radiation field shown in perspective at 310A and along axis 311 at
Broad-area radiation can be easily obtained by placing the target
assembly 26 of the probe 14 at a distance from the surface to be irradiated. The
solid angle of forward radiation from the target assembly 26 can be controlled with
an x-ray shield as described in the above referenced patent application. The thickness
of the shield at each point is determined so that a substantially uniform radiation
pattern is obtained.
Another application for such a broad-area x-ray source is intracavity
radiation within the body, such as the inside of the bladder. In such a case the
interface between the tissue and the broad-area x-ray source can be an inflatable
balloon, extending down the probe 14 so that the target assembly 26 is at the center
of the balloon.
Often, when treating a body cavity with radiation therapy, it is desirable
to uniformly radiate the entire surface of the soft tissue lining the cavity. In
other words, it is desirable to insure that an isodose contour is coincident with
the surface of the body cavity. One method of uniformly radiating the cavity is
to determine the three dimensional shape of the cavity through conventional methods
(e.g., by observation through a catheter, or through diagnostic procedures such
as a CT scan or Magnetic Resonance Imaging) and then to fabricate a probe tip 126
that will deliver isodose contours that match the shape of the cavity. Uniform radiation
is delivered to the cavity by inserting the probe, and activating it from within
the body cavity. This method can be difficult to practice since body cavities rarely
have a uniform shape, and further because body cavities vary in shape among individuals.
A preferred method of uniformly radiating a body cavity is to use
a device to first stretch the cavity into a known, uniform shape. In the preferred
embodiment for treatment of a bladder, the cavity is stretched to a spherical shape.
Then, an omnidirectional x-ray generating tip of probe kit (i.e., a point source)
is positioned at the center of the now spherical cavity. With that configuration,
an isodose contour is coincident with the surface of the body cavity. One device
useful for stretching a body cavity to a known, shape is a surgical balloon positioned
at the target end of probe 14.
FIGURE 5 shows the probe 14 of the x-ray source of FIGURE 1, with
a balloon assembly 400 including a balloon 410 disposed about the target assembly
26. As shown in FIGURE 5, the balloon 410 is deflated and compactly folded about
the target end of probe 14.
FIGURE 6 shows an embodiment with a probe 14 and a balloon assembly
400 including an elongated source guidance tube 402 extending along a central axis
and having a proximal end 404 and a distant end 406. The source guidance tube 402
has an interior channel 408 extending along the central axis. An inflatable balloon
410 is affixed to the outside of the distal end 406 of tube 402. The probe is slidably
positionable within the tube 402 so that the target end of probe 14 is positionable
within the interior region of balloon 410 when the balloon is inflated. With the
balloon 410 inflated, defining a spherical region 404, as shown in FIGURE 6 the
target assembly 26 is substantially at the center of the balloon 410.
Inflation and deflation of balloon 410 can be controlled from proximal
end 404 of probe 14, as will be discussed in detail below. Combinations of balloons
and catheters are well known and are described, for example, in U.S. Patent No.
For the embodiment of FIGURE 5, in practice, balloon 410 is initially
deflated and packed around the distal end of probe 14, as shown in FIGURE 5. The
distal end 406 of probe 14, with the folded balloon 410, is then inserted into the
body of the patient such that the distal end is positioned within the body cavity
to be treated. Proximal end 404 remains external to the patient during the entire
procedure. After distal end 406 has been inserted into the body cavity, balloon
410 is inflated to stretch the body cavity to a spherical shape.
As noted above, FIGURE 6 shows balloon 410 positioned within body
cavity 420 (shown in dotted lines). Body cavity 420 could be, for example, the bladder.
Initially cavity 420 defines a non-uniform shape, but inflating balloon 410 stretches
the lining of cavity 420 into a substantially spherical shape where the bladder
provides relatively little resistance to the inflation. Preferably, after inflation,
substantially all of the exterior of balloon 410 contacts the interior surface of
FIGURE 6 also shows a channel 408 extending along probe 14, establishing
a gas flow path by which the balloon 410 can be inflated from outside the patient.
In the preferred embodiment, probe 14 is inserted such that target assembly 26 is
positioned at the center of balloon 410. Since balloon 410 has stretched cavity
410 into a spherical shape, the center of balloon 410 is coincident with the center
of the cavity 420. Accordingly, positioning target assembly 26 at the center of
the inflated balloon 400 also centers target assembly 26 within the body cavity.
Once target assembly 26 is centered, the electron beam generator may
be activated to direct an electron beam to be incident on target assembly 26, resulting
in generation of x-radiation with an isodose contour coincident with the inflated
balloon, and the lining of the deformed body cavity.
FIGURE 7 shows another embodiment of the invention useful for treating
one section of a body cavity 420, such as a region bearing a tumor. In FIGURE 7,
target assembly 26 is shielded so only x-rays traveling in the direction of the
forward solid angle, shown by arrows 422, are emitted from target assembly 26. In
this embodiment, only region 424 of cavity 420 is radiated.
The above discussion has described the invention in connection with
spherical balloons, however, as those skilled in the art will appreciate, the invention
can be practiced with balloons of many shapes, including elliptical and cylindrical
shapes, which can be used in treating, for example, the colon, or other internal
cavities such as the urethra, vagina and cervix, uterus, colon, or rectum.
As discussed above with respect to FIGURE 1, the apparatus 10 includes
beam generation and acceleration components to generate and accelerate electrons,
prior to those electrons entering the probe 14. The generated electron beam then
flows through probe 14, impacts the target 26, and thereby produces x-rays. In the
absence of magnetic fields, the electrons flowing through the probe 14 follow a
straight-line trajectory. Consequently, the probe 14 is typically rigid without
However, in certain medical applications it is beneficial to use a
flexible probe. One such application involves threading the x-ray source down an
existing pathway, such as the trachea. Another such application involves maneuvering
the x-ray source around critical structures, such as a nerves or blood vessels.
FIGURE 8A shows a diagram of apparatus 200 including a flexible probe
214. The apparatus 200 includes a high voltage network 218, a laser source 220,
a probe assembly 214, and a target assembly 226. According to one embodiment of
the invention, the apparatus 200 provides the required flexibility, without using
strong magnetic fields, by locating electron generating and accelerating components
in the target assembly 226. The probe assembly 214 couples both the laser source
220 and the high voltage network 218 to the target assembly 226. The probe assembly
includes flexible fiber optical cable 202 enclosed in a small-diameter flexible
metallic tube 204.
The target assembly 226, which can be for example 1- to 2- cm in length,
extends from the end of the probe assembly 214 and includes a shell which encloses
the target 228. According to one embodiment, the target assembly 226 is rigid in
nature and generally cylindrical in shape. In this embodiment the cylindrical shell
enclosing the target assembly can be considered to provide a housing for the electron
beam source as well as a tubular probe extending from the housing along the electron
beam path. The inner surface 226A of the assembly 226 is lined with an electrical
insulator, while the external surface 226B of the assembly 226 is electrically conductive.
According to a preferred embodiment, the target assembly is hermetically sealed
to the end of the probe assembly 214, and evacuated. According to another embodiment,
the entire probe assembly 214 is evacuated.
The terminal end 202A of the fiber optical cable 202 is preferably
coated, over at least part of its area, with a semitransparent photoemissive substance
such as, Ag-O-Cs, thus forming a photocathode 216. A high voltage conductor 208,
embedded in the fiber optical cable 202, conducts electrons to the cathode 216 from
the high voltage network 218. Similarly, the flexible tube 204 couples a ground
return from the target 228 to the high voltage network 218, thereby establishing
a high voltage field between the cathode 216 and the target 228. The fiber optical
cable 202 acts as an insulating dielectric between the high voltage conductor 208
and the grounded flexible tube 204.
In one embodiment, to eliminate scattering of the light in the fiber
optic cable 202 by the high voltage wire 208, the fiber optic cable 202 can have
an annular configuration, as shown in cross-section in FIGURE 8B. The light from
the laser 220 travels down the annular core 250 of the fiber optic cable 202. Cladding
260 on each side of the core 250 has an index of refraction so as to refract the
light beam incident on the interface back into the core 250. A grounded flexible
metal tube 204 surrounds the outer cladding 260.
As in previously described embodiments, the target 228 can be for
example, beryllium, (Be), coated on one side with a thin film or layer 228A of a
higher impedance element, such as tungsten (W) or gold (Au).
In operation, the small semiconductor laser 220 shining down the fiber
optical cable 202 activates the transmissive photocathode 216 which generates free
electrons 222. The high voltage field between the cathode 216 and target 228 accelerates
these electrons, thereby forcing them to strike the surface 228A of target 228 and
produce x-rays. In order to generate, for example, 20 µA of current from
an Ag-O-Cs photocathode 216 with a laser 220 emitting light at a wavelength of 0.8
µm, the 0.4% quantum efficiency of this photocathode 216 for this wavelength
requires that the laser 220 emits 7.5 mW optical power. Such diode lasers are readily
commercially available. According to one embodiment of the invention, the photoemissive
surface which forms cathode 216 can, in fact, be quite small. For example, for a
current density at the cathode 216 of 1 A/cm2, the photoemitter's diameter
need only be approximately 50 µm.
One difficult fabrication aspect of this apparatus is the fabrication
of the photocathode 216, which for practical substances, with reasonable quantum
efficiencies above 10-3, should be performed in a vacuum. This procedure
can be carried out with the fiber optical cable 202 positioned in a bell jar, where
for example, an Ag-O-Cs photosurface is fabricated in the conventional manner. Subsequently,
without exposure to air, the optical cable 202 can be inserted into the tube 204.
The end 202B can be vacuum sealed to the flexible tube 204.
The invention may be embodied in other specific forms. The present
embodiments are therefore to be considered in all respects as illustrative and not
restrictive, the scope of the invention being indicated by the appended claims rather
than by the foregoing description, and all changes which come within the meaning
and range of equivalency of the claims are therefore intended to be embraced therein.