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BACKGROUND OF THE INVENTION
This invention relates to a hyperthermia apparatus and more particularly to
an apparatus having combined hyperthermia treatment and noninvasive
thermometry capabilities.
Known hyperthermia treatment systems include multiple applicators and
multiple temperature sensors for controlling the operation of the
hyperthermia system. The multiple applicators utilizing ultrasound or
microwave energy in the direct contact operating mode are placed directly
upon an elastic cooling belt containing a circulating cooling liquid to
carry the heat of hyperthermia treatment away from the surface of the
healthy tissue. The temperature sensors are implanted in the normal tissue
in the vicinity of the tumor, as well as within the tumor. Systems
involving the placement of temperature sensors within the body are
referred to as invasive thermometry systems. Those persons skilled in the
art desiring additional information for this system are referred to U.S.
Pat. No. 4,397,314 issued Aug. 9, 1983.
A known instrument for detecting microwave energy and for giving an
accurate measurement of the power density thereof is disclosed in U.S.
Pat. No. 3,919,638 issued Nov. 11, 1975. This instrument is substantially
unaffected by polarization or modulation of the electromagnetic waves and
includes a planar array of parallel connected diode detectors each having
a pair of antenna leads forming a dipole antenna. The diode array may
include groups of diodes having different antenna lead lengths to detect
different frequencies of microwave energy for a meter. The meter may be
selectively switched between the outputs of the different groups.
Further, the potential use of multiple frequency band radiometry as a means
of noninvasive sensing of one-dimensional temperature profiles is
presented in an article entitled "Noninvasive Thermometry Using
Multiple-Frequency-Band Radiometry: A Feasibility Study", Stavros D.
Prionas and G. M. Hahn, Bioelectromagnetic 6:391-404, 1985 Alan R. Liss,
Inc. The article discloses that microwave thermography has been
extensively used for the detection of cancerous nodules. Operating
frequencies in the range of 1.3 to 6.0 GHz (free space wavelengths in the
range of 5 to 23 Cm.) have been employed. At these long wavelengths
subcutaneous temperature measurement is possible and detection of
superficial tumors in the brain and thyroid is, in principle, feasible.
The article further discloses that a computer tomographic approach using 10
GHz microwaves has been proposed as an alternative to mammographic
examination. A self-balancing microwave radiometer for measuring the
energy emitted from a heated volume within a single frequency band has
been developed at the RCA laboratories. The power spectrum of thermal
noise generated by a given temperature depth distribution is governed by
Planck's law of blackbody radiation. The frequency spectrum of energy
received at the surface of the human body is affected by the frequency
dependent attenuation properties of the intervening tissues. Microwave
radiometry (the technique of measuring noncoherent electromagnetic energy,
in the microwave part of the spectrum, that is emitted or scattered by the
medium under observation) can be used to measure the thermal noise emitted
from a heated volume of biological tissue.
This article reports the analysis of the spectral content of this thermal
noise and the comparison of the magnitude of the signal to the inherent
threshold of noise detectability associated with an ideal microwave
radiometer. In the analysis a one-dimensional temperature distribution
model was assumed. In real situations, three-dimensional temperature
distributions will be encountered. It is clear that to resolve such a
three-dimensional temperature field with a reasonable amount of spatial
resolution additional information will be needed. This additional
information might be in the form of data acquired using different
orientations of a signal receiving aperture or a properly phased array of
receiving apertures. An intriguing alternative is to employ a phased array
of receiving apertures that coherently detect the signal emanating from a
point in space. In either case, well established signal processing
algorithms could be used to convert the measured data to reconstruct the
temperature distribution.
The use of "Radiometer Receivers for Microwave Thermography" was disclosed
in an article of the same title by D. V. Land, University of Glasgow,
Glasgow, Great Britain, published in the microwave journal, May 1983. A
comparison or Dicke radiometer configuration is used. The receiver
produces an output at the input switching or modulation frequency that is
proportional to the difference between the source temperature detected by
an antenna and the temperature of an internal reference load or noise
generator.
Finally, the application of board-band correlation techniques to medical
microwave thermography was studied and reported in an article entitled
"The Thermal and Spatial Resolution of a Broad-Band Correlation Radiometer
with Application to Medical Microwave Thermography", Joseph C. Hill et al.
, IEEE Transations on Microwave Theory and Techniques, Vol. MTT-33, No. 8,
August, 1985.
An essential difference between the present invention and the prior art is
the use of elements common to a hyperthermia treatment system and a heat
detecting system to produce a combined hyperthermia treatment and
noninvasive thermometry apparatus.
An advantage of the combined hyperthermia treatment and noninvasive
thermometry is the cost savings. Another advantage is the commonality of
parts which enables utilization of the apparatus in the heating mode and
the temperature measurement mode using the same apparatus parameter
settings to produce quickly the desired results through selectively
switching a power source and a radiometer into a circuit containing the
common parts. Without combining of the devices it may be either impossible
or impractical to non-invasively monitor deep temperatures during
hyperthermia treatments.
SUMMARY OF THE INVENTION
Accordingly it is an object of the present invention to provide an
apparatus capable of creating hyperthermia in body tissue and measuring
temperature distributions within the treated volumes during the course of
hyperthermia treatments. The tissue regions to be treated and monitored
include even deep tissues in the central area of the human torso, limbs,
and brain.
Another object of the invention is to provide a cost effective hyperthermia
and thermometry apparatus.
Still another object of the invention is to provide a combined apparatus
for heating selected tissue areas whose set up parameters can be used also
for measuring the temperatures of the selected tissue areas.
Yet another object of the invention is to provide a hyperthermia apparatus
which can be effectively used in either a treatment mode or a diagnostic
mode or both.
Briefly stated the hyperthermia apparatus comprises a combined
transmitter/receiver system for inducing hyperthermia in body tissue and
for measuring noninvasively the temperature of the corresponding body
tissue.
BRIEF DESCRIPTION OF THE DRAWINGS
The novel features which characterize the present invention are defined by
the appended claims. The foregoing and other objects and advantages of the
invention will hereinafter appear, and for purposes of illustration, but
not of limitation, a preferred embodiment is shown in the accompanying
drawings.
FIG. 1 is a schematic in block form of a first embodiment of the
hyperthermia system constituting the subject matter of the invention;
FIG. 2 is a diagram showing the method of operation of a system according
to FIG. 1;
FIG. 3 is a diagram showing relative electric field amplitudes within a
homogeneous target specimen;
FIG. 4 is a diagram showing relative power density in a homogeneous target
specimen;
FIG. 5 is a partially cut away perspective view of one preferred applicator
for use with the present invention;
FIG. 6 is a top diagrammatic view showing operation of the applicator of
FIG. 5;
FIG. 7 is a diagram of a dipole antenna for use with the present invention;
FIG. 8 is a perspective view of a folded dipole array for use with the
system of the present invention;
FIG. 9 is a perspective view of a third preferred applicator for use with
the present invention;
FIG. 10 is a partial view of the hyperthermia system showing the
applicators with equal path to target with equal phase setting on phase
shifters, and the non-equal path to non-target tissue; and
FIG. 11 is a schematic in block of a second embodiment of the invention.
DESCRIPTION OF THE PREFERRED EMBODIMENTS
FIG. 1 shows a block diagram of a hyperthermia treatment apparatus 10
having a subsystem 12 for creating hyperthermia in a target specimen 14 by
means of electromagnetic radiation (EMR), a subsystem 16 for measuring
noninvasively the temperature of the target specimen, and elements common
thereto 18. The common elements include a central processor unit 20 which
controls the system 10, and is in an interactive feedback relationship
with each of its elements. The central processor unit (CPU) 20 accepts a
plurality of inputs describing the present condition of the target 14.
A control panel or console 26 is coupled to the CPU 20 and is used by an
operator to control a treatment and monitor its progress. The control
panel 26 can be used to display any information obtained from the target
14 as well as all indicators of system operation. Various memory devices,
represented by a single memory block 28, are coupled to the CPU 20. The
memory 28 stores the result of pretreatment calculations which are used by
the CPU 20 to control the progress of the treatment. Also, all pertinent
operating data are stored in another part of the memory 28 as generated in
order to have a complete record of the treatment process and results for
future use.
For the hyperthermia subsystem 12, a high frequency energy source 30 is
coupled to and controlled by the CPU 20. The source 30 is coupled through
a switch 31 to a receiver such as a radiometer 33 for the receiver
subsystem 16, and to power handling means including a power splitter 32,
which constitutes a branch point in the power handling means in that it
divides energy into a plurality of lines each having the same phase and
power, a phase adjuster, and applicator connecting switches 37. The phase
of the energy in each line shifter 34 manually or automatically. The
output of each phase shifter 34 is coupled to one individual applicator 36
or a group of applicators. The actual delivery of power to the applicators
36 is controlled by switches 37 located between the phase shifters 34 and
the applicators 36. The switches 37 may be simple on-off switches such as
relays or solid state switches, or they may be digital or continuously
variable attenuators or the switches 37 can also represent variable gain
amplifiers. Like the source 30 and the phase shifters 34, the switches 37
are preferably controlled by the CPU during system operation, although the
switches 37 may be manually operated. FIG. 1 shows only four phase
shifters 34, applicators 36 and switches 37, but an actual system 10 may
employ more of each to provide steering in several directions.
Referring to FIG. 2, eight individual applicators 36 are shown coupled
together in an octagonal arrangement and surrounding a circular target 14.
Each applicator 36 is diagrammatically represented by a rectangle. In
reality, each applicator 36 would have a shape suitable for the emission
of microwave EMR, several embodiments of which are shown in later
drawings. FIG. 2 is a two-dimensional representation of a
three-dimensional phenomena, with both the radiators 36 and target 14
extending for some distance perpendicular to the plane of the drawing. The
radiation emitted from each applicator 36 is aligned so that the electric
field component is perpendicular to the plane of the drawing and the
magnetic field component forms circular equal potential lines which lie in
the plane of the drawing. The stylized wave fronts shown in FIG. 2 by
arrows approximate the direction of EMR emitted by the various applicators
which is perpendicular to the electric and magnetic field component.
As the radiation emitted by the various applicators 36 converges on the
target 14, it is seen that the electric fields of the radiation are lined
up so that the target 14 sees, approximately, a converging circular wave
front. The energy of the various wave fronts converges in the center of
the target 14, where the electric field adds constructively and heats the
center regions 38 of the target 14 to a greater degree than that caused by
any one of the applicators 36 alone. This improved deep internal heating
is caused without dangerously increasing the radiant energy density at the
surface of the target 14, as the incoming energy is normally spread
equally over the entire target surface. Thus, the energy imparted to the
target 14 is concentrated near the center, when this is desired, and
minimized to the extent possible at the target surface. Changes in
amplitude and phase can displace the central energy focus to better heat a
non-central target.
As described in connection with FIG. 1, the energy radiated by each
applicator 36 has a constant phase relationship with that emitted by the
other applicators 36. This creates a synergistic result in the center area
38 of the target 14, whereby the center 38 is heated to a degree greater
than that of a simple sum of the energy of the various applicators 36. The
synergistic result will be described in more detail with relation to FIGS.
3 and 4. With all of the applicators 36 operating precisely in phase, the
central heating area 38 will be symmetrical around the center point of a
homogenous target 14. If the shape or location of the central heating
region 38 is desired to be other than symmetrical about the center,
changing the relative phase of the EMR emitted by the various applicators
slightly will cause the central heating region 38 to move generally toward
the applicators 36 which are phase-lagging the remainder (FIG. 10). By
controlling the phase of energy emitted by the applicators 36 as described
in connection with FIG. 1, it is therefore possible to manipulate the
location of the central heating region 38 to best achieve the desired
result. Manipulation of the central housing region 38 can also be
accomplished through control of the power levels by controlling the
switches, attenuators, or amplifiers 37. The power to each applicator 36
can be controlled as desired, through either on-off switches or
continuously variable switches or attenuators as described above. Lowering
or cutting off power to individual applicators 36 changes the shape of the
central heating region 38, and the power absorbed at various points in the
target 14 displace from the center 38 toward the higher power applicators.
FIGS. 3 and 4 show the mechanism by which the greatly increased power
deposition in the central heating region 38 occurs. Considering any pair
of diametrically opposed applicators 36 of FIG. 2, and a non-lossy
homogenous target 14, the drawing of FIG. 3 shows the standing wave
amplitudes of the E-field component of the EMR generated by such opposing
pair on a line through the center of the target 14. The horizontal axis
represents the distance between the opposing applicator emission faces,
shown as points F.sub.1 and F.sub.2, and the vertical axis represents the
amplitude of the alternating E-field standing wave at each distance. The
points S.sub.1 and S.sub.2 represent the opposite surfaces of the target
14 with no consideration presently being made of the E-field external to
the target 14.
Because the two oncoming wave fronts are of identical frequency and phase
and have their E-fields aligned parallel to the center axis of the target
14, the electric field at each point between the applicators is the sum of
the E-field vectors of each wave. When the frequency of the emitted
radiation is chosen so that the wavelength in the target 14 is
approximately three-fourths the diameter of the target 14, the amplitude
of the standing wave caused by two applicators 36 in the target 14 is
shown in FIG. 3. The maximum amplitude is located in the center region 38,
with minimums being located one-fourth wavelength to either side of the
center. The amplitude at the center is the sum of the amplitudes from each
applicator 36, which for the two opposed apertures is twice the E-field
created by a single applicator 36. When more than two applicators 36 are
used, as shown in FIG. 2 for example, the resultant E-field sum is of
course larger.
Testing has shown that best results are normally obtained when the
wavelength of emitted EMR is between approximately 3/4 and twice the
target 14 diameter. This gives a relatively well-defined central heating
region 38 and a good impedance match between the applicators 36 as
described below and the target 14. Thus, for a target 14 diameter d, the
preferred range of wavelengths can be found from the expression:
0.5.lambda..sub.m .ltoreq.d.ltoreq.1.3.lambda..sub.m (1)
where .lambda..sub.m is the wavelength in the tissue medium being heated.
For high water content tissues, such as muscle and blood, the wavelengths
at 100, 300 and 915 MHz are approximately 27, 11.9 and 4.5 cm.
respectively. For low water content tissues the respective wavelengths at
the frequencies are approximately 106.41 and 13.7 cm. If both types of
tissue are present in a target, it is preferable to select a wavelength
which satisfies equation (1) for the most prevalent tissue type (normally
muscle tissue). A wavelength larger than suggested by equation 1 can be
used if adequate impedance matching is obtained or provided by external
matching techniques and nearly uniform surface vs. central heating can be
expected.
FIG. 4 shows the relative power density at each point in the target 14
corresponding to FIG. 3. The power density is proportional to the square
of the electric field strength, so that the power density curve shows a
relatively sharp peak in the central heating region 38 for a
non-attenuating medium. Heating at any point is due to the power absorbed
at that point, which is in turn directly proportional to the power density
at that point. Therefore, a heating cross section of the target has the
same distribution as the power density curve of FIG. 4 when heat transfer
effects are neglected. However, a medium capable of absorbing the radiant
power is attenuating and will substantially reduce the central power
density peak as represented in FIGS. 3 and 4 and increase the power
density somewhat at the surface, resulting in near uniform heating being
dependent on frequency, tissue diameter, and tissue conductivity. The
centrally higher power density is still possible by proper frequency
selection and antenna size.
Since the power density is proportional to the square of the E-field, a
simple additive increase in the electric field at a given point results in
an increase in the power density at that point by the square of the
electric field. For example, in FIGS. 3 and 4 the electric field in the
central heating region 38 resulting from 2 apertures is twice that due to
a single applicator 36. Therefore, the power density of the central region
is 2.sup.2 =4 times the power density that would be caused by a single
applicator 36. When, as in FIG. 2, more applicators 36 are used, the
increase in power density becomes much greater than that caused by a
single applicator 36. When eight applicators 36 are used, the E-field at
the center is 8 times that caused by a single applicator 36 alone, and the
power density at the center is therefore 8.sup.2 =64 times the power
density caused by a single applicator 36. This enormous increase in power
density, and thus power absorbed, in the center of the target 14 is
obtained without significantly increasing the power density at any one
point on the surfce of the target 14. This phenomena, a synergistic result
due to all applicators 36 operating at the identical frequency and with a
predetermined phase relationship, allows deep heating of the target 14
without undesired excessive heating of the surface portions.
The above discussion of FIGS. 3 and 4 applies to non-lossy targets 14. In
such targets 14 there is little energy absorption by the medium, so that
the amplitude of the EMR from any given applicator 36 is undiminished as
the radiation passes through. HOwever, actual targets 14 are lossy, so
that the amplitude of EMR decreases as it passes through the target 14. In
a typical case, the amplitude of the E-field at the center region 38 may
be approximately 1/7 the E-field amplitude at the surface of the target
14. For the case of eight applicators 36, the power density in the center
region 38 is approximately 8.sup.2 /7.sup.2 or approximately 1.3 times
that at the surface of the target 12. It is important to note, however,
that the power density in the center region 12 is still sixty-four times
that which would be caused by a single applicator 36 alone. The general
shape of the E-field and power density wave forms of FIGS. 3 and 4 still
apply, with the actual peak value in the center region 38 being diminished
with a lossy medium to nearly the same level as the surface in many cases.
In general little is gained by using over 8 antennas if a non-lossy
coupling medium is used to separate the antennas from the target.
The above-described synergistic increase in power density in the central
heating region 38 occurs only when all applicators 36 radiate at the same
frequency with the E-fields of the emitted EMR aligned. This E-field in
phase alignment preferably occurs along the central axis of the target 14,
which is perpendicular to the page as shown in FIG. 2. The E-field
alignment of the various applicators 36 is preferably made as accurate as
possible. However, some misalignment can be tolerated without unreasonably
reducing the performance of the system 10. The vector sum of the E-field
at any one point is equal to the sum of the individual E-field vectors. If
a particular applicator 36 is misaligned, the EMR power emitted by that
applicator contributes less to the synergistic power increase than an
aligned field by a factor equal to the square of the cosine of half the
angle between the misaligned E-field and the remaining E-fields. For small
angles the cosine is close to one, so that small mismatches in the E-field
alignment do not substantially degrade the synergistic power density
enhancement which occurs in the central heating region 38.
When the various applicators 36 radiate energy at slightly different
frequencies, it will become apparent to those skilled in the art that the
various E-fields will not always add constructively, and the power density
enhancement described above will not occur. In fact, in such a case, the
power density in the central region 38 will be, at best, the simple sum of
the individual applicator power densities. It is therefor important that
the frequency of radiation emitted by all aplicators 36 be absolutely
identical. For this reason, the preferred embodiment utilizes a single
power source 30 and a power splitter 32 whereby the power supplied to each
applicator 36 has the same frequency. While it is possible to use multiple
sources providing that they can be precisely phase locked to emit
identical frequencies, the practical considerations to accomplish this
would complicate the system and add expense with little benefit, thus the
preferred configuration requires only a single source 30 and splitter 32.
It is not necessary that the frequency supplied to the applicators 36 be
invariant with respect to time. In fact, it is desirable that the source
30 have a controllable frequency so that it may be adjusted to optimize
performance with various target 14 materials as described above.
The shape and location of the central heating region 38 is determined by
the distribution of applicators 36 in operation, the relative phase
between them, the diameter of the tissue, the positioning of the tissue
and the EMR frequency. It has been determined that the use of four or more
uniformly spaced radiating applicators 36 will provide an approximately
circular (ellipsoidal in three dimensions) central heating region 38.
Eight radiating applicators 36 are used in FIG. 2 instead of four, because
it has been determined that the power density, and thus heating, at the
surface of the target 14 is more uniform than when only four radiating
applicators 36 are used. An increase in the number of applicators 36 above
eight does not appear to make a material difference in the operation of
the system 10.
When all of the applicators 36 are precisely in phase, a symmetrical
heating region 38 is formed in the precise center of a homogeneous target
12. Varying the relative phase of energy emitted by the various
applicators 36 will cause the central heating region 38 to be shifted
somewhat away from center toward the applicators 36 with lagging phase.
The ability to alter the relative phase and amplitude between the
applicators 36 is extremely useful, for example when a non-homogenous
target 14 (such as an animal torso) is used. The wavelangth of the EMR
will vary slightly in the different tissues of the target 14, and an
alteration in emitted phase can compensate for the phase shifts thereby
induced. Thus, when the target 12 has a known cross section of different
tissues having known properties, the phase between the various applicators
36 can be adjusted to position the central heating region 38 at the
desired location. Off-set positioning of the tissue can also compensate
for non-homogenous effects on the heat pattern. The observations have
shown, however, that the non-homogeneous tissues of the body have not
significantly altered the central phase focus zone.
One applicator array 40 suitable for use is shown in FIGS. 5 and 6. This
annular array 40 comprises a battery of sixteen horn-type parallel plate
waveguide antennae forming a folded dipole array being coupled together
into two layers of eight radiators 36 each. For simplicity in
construction, each input to the applicator array 40 feeds a two by two
array of individual applicators 36. Thus, only four power inputs are
needed for this sixteen applicator array 40.
It has been determined that the stacking of two or more applicators 36 in
the E-field direction (perpendicular to the page in FIG. 2) still provides
a substantially uniform electric field vertically but reduces the
applicator size. It hs also been determined that stacking of the
individual applicators 36 in an annular array 40 along the H-field (in the
plane of the page in FIG. 2) provides a substantially uniform electric
field around a target 14. Use of an array of individual applicators 36
allows each one to be sized and constructed so as to provide a good
impedance match between the applicators 36 and the target 14. A detailed
discussion of methods for fabricating horn-type applicators 36 suitable
for use with the present invention is contained in pending U.S. Patent
Application Ser. No. 136,506, filed on Apr. 2, 1980, now U.S. Pat. No.
4,462,42 and titled Annular Electromagnetic Radiation Applicator For
Biological Tissue And Method, which disclosure is herein incorporated as
if set forth verbatim.
The applicator array 40 is surrounded by a casing 42, which serves to
support the individual applicators 36 in place and to decrease stray
radiation, which can be hazardous. As shown in FIG. 5, the casing 42 is
partially cut away, exposing portions of four separate applicators 36. A
coaxial power input line 44 is coupled to a parallel plate waveguide 46.
The waveguide 46 is coupled to four feed guides 48. The feed guides 48 are
in turn coupled to four individual applicators 36, and have the same
dimensions so that power is split evenly to the applicators 36. Each set
of four applicators 36 therefore radiates energy having the same phase,
power and E-field alignment.
FIG. 6 shows a top view of the applicator array 40. A target 14 is
suspended interiorly of the array 40, and is surrounded by a bolus 50. The
bolus 50 preferably contains deionized water, and is made from a flexible
material so as to seal tightly around the target 14. Air gaps 52 may be
left between portions of the bolus 50 and applicators 36, or the bolus may
be manipulated to fill these gaps as desired.
The use of a bolus 50 has several important advantages. The fluid therein
can be circulated through an external heat exchanger (not shown) to cool
surface regions of the target 14. When deionized water is used in the
bolus 50, there is very little power loss in the bolus 50, so that the
full power raidated by the applicators 36 is delivered to the target 14.
Use of the bolus 50 improves the impedance match between the applicators 36
and the target 14. At the frequencies of interest, the impedance of a
typical biological target 14 is approximately 44 ohms. The impedance of
the applicators 36 and other electrical portions of the system is
preferably 50 ohms in order to be compatible with standard components. The
impedance of deionized water at the frquencies of interest is also
approximately 44 ohms, so that all parts of the system 10 are inherently
closely matched. If the water-filled bolus 50 were not present, a larger
mismatch would occur at the radiating face of the applicators 36 and at
the surface of the target 14. This mismatch occurs because the impedance
of air is approximately that of free space, or 377 ohms. Any impedance
mismatches cause reflections at the boundaries, lowering the percentage of
radiated energy delivered to the target 14 and increasing stray radiation
hazards.
As can be seen in FIG. 6, the applicator array 40 emits energy in the
pattern discussed hereinbefore in connection with FIGS. 2 and 3, resulting
in the power density pattern shown in FIG. 4. Thus, the applicator array
40 provides heating in the central region 38 without excessive heating of
the surface regions.
An alternate embodiment of an applicator suitable for use with the present
system 10 is shown in FIG. 7. This applicator 54 is essentially a dipole
antenna pair sized for use with EMR of the frequencies contemplated. Each
arm 56 of the upper and lower radiating portions 58, 60 acts as a single
radiator in a manner similar to that of the annular array 40 of FIGS. 5
and 6. A coaxial feed line 61 is coupled to the center of the upper and
lower radiating portions 58, 60, a balun can also be used to transform the
coax-line to a balanced feed. When this applicator 54 is driven in a
conventional manner, the E-field of the emitted radiation is aligned with
the length of the arms 56.
The shape and size of the antenna arms 56 determine the optimum frequencies
of operation and impedance characteristics of the dipole 54. It has been
determined experimentally that a dipole 54 having tapered arms, wherein
the ratio of arm width (W) to length (L) is maintained constant at
approximately 0.087, gives a good impedance match with remainder of the
system 10. When the dipole applicators 54 are combined into a cylindrical
array 66 as shown in FIG. 8, and a water bolus (not shown) as similar to
that discussed with FIG. 6 is used, a reasonably good 50 ohm impedance
match is achieved. This approach tends to be more narrow band than the
annular array of FIG. 5, but is much simpler to build.
Referring to FIG. 8, four dipole pair radiators 54 are assembled on a
rigid, non-conducting frame 64 to form a cylindrical array 66. Each dipole
radiator 54 is separately connected to the source 30 and power splitter 32
through a separte coaxial feed line 61. Each applicator 54 launches
microwave EMR towards the center of the array 66 where the target specimen
(not shown) is located. Preferably, a deionized water bolus (not shown)
surrounds the target so as to better couple energy from the applicators 54
to the target, and to minimize reflections. The phase of the energy to
each dipole applicator 54 can be controlled to vary the location of the
central heating region 38 or to compensate for wavelength variations in
non-homogenous targets as described in connection with FIG. 2.
Referring to the alternate embodiment of FIG. 9, a cylindrical dipole
applicator 68 comprises two coaxial conducting cylinders 70 placed close
together. These concentric cylinders 70 act as a single dipole applicator
wherein the radiating arms comprise a flat radiating sheet which has been
folded around to make contact with itself. The cylindrical dipole 68 will
radiate toward its central axis, and constructive interference of the
E-field will cause the synergistically increased power absorption in the
central region 38 as described with the previous applicators. A single
coaxial feedline 72 is sufficient to drive the cylindrical dipole 68.
However, there will be some phase lag in the EMR emitted from the portions
of the dipole 68 diametrically opposite the feedline 72. This will cause
the central heating region 38 to shift somewhat away from the feedline 72
contacts. While this may be desirable in some cases, the preferred
embodiment includes four coaxial feedlines 72 equally spaced around the
dipole 68. When all four feedlines 72 are driven at the same phase, the
central heating region 38 will be centered around the axis of the
cylindrical dipole 68. Some manipulation of the central heating region 38
location can be made by varying the phase to the coaxial feedlines 72, but
in general the degree of control will be less than that experienced with
either the dipole array 66 or the horn radiator array 40.
Since the effective radiating aperture width of the cylindrical dipole 68
is equal to its circumference, and the height is limited by the size of
the target to typically two feet or less, inherent impedance matching as
was obtained with the dipole array 66 is difficult to achieve. Since the
cylindrcal dipole radiator 68 will not be inherently matched with the
impedance of the remainder of the system at most desirable frequencies, a
conventional impedance matching device (not shown) should be used to
minimize losses and reduce reflected power.
Both the folded dipole array 62 and the cylindrical dipole 68 emit
radiation from both their inner and outer surfaces. An internal water
bolus will increase the proportion of radiation emitted centrally, due to
the better impedance match into the lower impedance fluid media. In order
to reduce further the hazard of stray radiation, an outer conducting
cylinder (not shown) can be placed around the cylindrical dipole 68 or
dipole array 66. This outer shield can be grounded or left floating in
order to reflect such radiation and reduce the outwardly emitted
radiation. The reflecting shield must be spaced a sufficient distance from
the cylindrical dipole 68 or dipole array 66 so that the ground plate will
not capacitively load the aperture so much as to interfere with the
primary emitted radiation distribution and reduce heating in the central
region 38 or cause undesired non-central heating. In the preferred
embodiment, the outer conducting cylinder is grounded and displaced so
that minimal power pattern changes occur. Normally the space between the
outer cylinder and the dipoles is filled with air or another low
dielectric material thereby reducing the amount of energy coupled to the
shorting cylinder. The grounding of the outer conducting cylinder is
preferably done with a second coaxial outer shield dielectrically spaced
from and outside of the coaxial outer conductor connecting to the dipole
radiators.
In order to provide an effective hyperthermia treatment, the operator must
be able to accurately determine the internal status of the target 14. For
a living target 14, monitoring the vital signs gives a general indication
of the health of the target 14 and indicates adverse events affecting its
health. However, these signs, such as pulse, respiration, blood pressure
and oral temperature, do not indicate whether enough heat is being applied
to the region of interest to be effective.
Two additional measurements provide a fairly complete picture of the
internal local effects of the hyperthermia treatment. The first of these
is the measurement of actual temperature at selected points within the
target 14. A real time thermal profile allows the operator to determine
whether the desired regions of the target 14 are being heated to
temperatures wh | | |
