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Description  |
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TECHNICAL FIELD
This invention relates generally to a microwave antenna and more
particularly to an implantable helical coil microwave antenna for improved
localization of interstitial hyperthermia.
BACKGROUND ART
Hyperthermia constitutes an effective adjuvant treatment for malignant
tumors which are refractory to conventional therapy with surgery,
radiation or chemotherapy. Hyperthermia can be administered to a patient
either by an externally applied electromagnetic or ultrasound source, or
internally by an interstitial heating technique. Implantable microwave
antenna heating has proven the most popular of the three present
interstitial heating modalities.
However, major problems arise with the use of conventional half-wavelength
dipole antennas which severely limit the applicability and effectiveness
of interstitial microwave hyperthermia. Such problems include: variability
of heating profiles when the antenna is inserted to different depths in
the tissue to be treated, restricted range of possible heating lengths for
a given microwave frequency, and the presence of a so-called "cold region"
or "dead length" occurring at the tip of the antenna. Attempts have been
made to solve one or more of the above problems by providing implantable
radiating antennas having improved performance characteristics.
For example, two-node and three-node microwave antennas have been proposed
to expand the heating volume to as much as twice that provided by a
single-node dipole antenna. However, such antennas have exhibited an
inhomogenous heating pattern with three or four peaks along the antenna
axis and a failure to heat effectively out of the antenna tip. A variable
diameter dipole antenna has also been proposed to force the heating
current into larger diameter sections of the antenna which fit snugly
within the biocompatible plastic catheter. The larger diameter section at
the tip of the antenna appears to provide more effective tip heating, but
the antenna still exhibits considerable dependence of heating on insertion
depth and periodic excessive surface tissue heating.
Other types of dipole antennas, such as the sleeved coaxial slot and
balun-fed folded dipole, have been proposed for shifting the heating field
out to the antenna tip. The concept of multiple breaks in the coax outer
conductor of the antenna with each section being driven by a separate
microwave source has held some promise for closer control of the depth
heating profile, but at the expense of greatly increased equipment
complexity. Although often accomplishing an expansion of the effective
heating length for a given frequency and/or a reduction in dead length at
the tip, the above types of antennas are commonly plagued with the same
critical problem as that of the linearly polarized simple dipole antenna,
namely, a critical dependence of the heating pattern on the depth of
insertion.
Another proposed technique employs an "over-ride" reciprocated motion
system for linearly translating the dipole antenna during treatment.
Although this technique may potentially solve at least part of the axial
heating pattern problem, predictability and real time control of the
overall heating pattern would likely prove difficult due to power
deposition pattern changes at different positions within the range of
antenna movement.
A similar development of antenna designs has occurred for intracavitary
heating applications. Antennas of this type having somewhat larger
diameters (e.g., 1-1.5 cm v. 0.1-0.15 cm) have been used in the treatment
of tissues surrounding body cavities. An antenna of the latter type has
been constructed with a 1.0 cm diameter coax cable outer conductor cut in
a helical manner and pulled apart axially to form a helical extenson of
the antenna feedline outer conductor having ten turns extending 14 cm in
length. Thermal profiles of the antenna were found to be quite variable
for the different conditions studied and the antenna exhibited a strong
dependence on both source frequency and insertion depth. Most tests were
performed using insertion depths less than the 14 cm length coil.
A so-called "flexible leakage type" antenna has also been proposed for use
at 2450 MHz. This type of antenna consists of a helical structure composed
of 1.0 mm wide copper foil tape interconnected between the inner and outer
conductors of a 2.0 mm diameter flexible coaxial cable.
DISCLOSURE OF THE INVENTION
The improved microwave antenna of this invention is comprised of an inner
conductor, a dielectric insulator covering the inner conductor and an
outer conductor covering only a proximal portion of the insulator to form
a coaxial cable thereat and to leave a distal portion of the insulator
uncovered. A helical coil surrounds the distal portion of the insulator
and has its proximal end spaced axially from a distal end of the outer
conductor and its distal end connected to a distal end of the inner
conductor.
The antenna will function to confine heat to the immediate area surrounding
the helical coil and to produce substantially identical thermal profiles
at varying depths of insertion of the antenna into the treatment volume
tissue when the antenna is energized with microwave energy. In addition,
the antenna of this invention exhibits other advantages over existing
implantable microwave antenna designs. For example, the size of the heated
volume for a given frequency can be adjusted readily by simply changing
the length of the coil (within a preselected range). The dead length (cold
portion at the end of the implanted antenna) is eliminated which minimizes
the need for implanting antennas deeper than the lower extent of the
target region. Undesirable tissue heating along the antenna feedline is
eliminated with deep insertions, as well as overheating of the tissue
surface for shallow depth insertions of the antenna.
The antenna of this invention will thus provide well focused and controlled
interstitial hyperthermia to a given volume of tissue, regardless of
location within a larger structure without undesirable and uncontrollable
hot spots along the antenna feedline. Multiple antenna arrays may be used
to expand the effective heating volume laterally, using either coherently
or incoherently phased microwave sources. Use of the antenna is compatible
with existing radioactive seed brachytherapy for combination interstitial
hyperthermia and radiation therapy of malignant tumors. Although the
antenna is particularly adapted for improved heating uniformity in
interstitial hyperthermia therapy of cancer, the antenna is useful for a
variety of other applications wherein heat must be applied uniformly to
the interior of large lossy dielectric volumes. New applications in food
warming, material quick-thawing and tissue therapy are expected with
reduced pricing of antennas and microwave sources.
BRIEF DESCRIPTION OF THE DRAWINGS
Other objects and advantages of this invention will become apparent from
the following description and accompanying drawings wherein:
FIG. 1 is a longitudinal cross-sectional view of a microwave antenna
embodying this invention;
FIG. 2 graphically illustrates the effect of varying insertion depths on
the axial thermal profiles of a 2450 MHz, L=1 cm. helical coil antenna in
a tissue equivalent homogenous phantom;
FIG. 3 is a longitudinal cross-sectional view of a commonly used style of
implantable dipole antenna;
FIG. 4 is a graph, similar to FIG. 2, showing the effect of varying
insertion depth on the axial thermal profiles of a 2450 MHz, L=1 cm simple
dipole antenna (FIG. 3) for comparison with the antenna of this invention;
FIG. 5 is a comparison in dog thigh muscle in vivo of the radial
temperature fall-off from the antenna of this invention to that of the
standard dipole antenna, and to that of thermal conduction heating only;
and
FIG. 6 is a graph showing the essentially constant shape of the induced
temperature distribution for three different 915 MHz helical coil antenna
insertion depths in dog thigh muscle in vivo.
BEST MODE OF CARRYING OUT THE INVENTION
General Description
FIG. 1 illustrates an implantable helical coil microwave antenna 10
embodying the present invention. The antenna was constructed from a
miniature coaxial cable comprising a metallic inner conductor 11
surrounded by a tubular dielectric insulator 12 fully covering the inner
conductor. A distal section of a tubular metallic outer conductor 13 was
removed so that the outer conductor only covered a proximal portion of
insulator 12 to form a coaxial cable portion thereat and to leave a distal
portion of the insulator uncovered.
A metallic wire helical coil 14 was constructed to surround the distal
portion of insulator 12 and to have its proximal end spaced axially from a
distal end of outer conductor 13 to form a separation gap G therebetween.
A distal end of the helical coil was soldered and connected at 15 to a
distal end of inner conductor 11. As more fully explained hereinafter,
antenna 10 was found to confine heat to the immediate area surrounding
helical coil 14 and to induce substantially identical thermal profiles at
varying depths of insertion of the antenna into an organic subject when
the antenna was energized with microwave energy.
In use, antenna 10 is adapted to slip within a standard 16-gauge plastic
catheter 16, as is, commonly used for brachyradiotherapy treatments. In
use, the catheter entirely covers the implanted antenna and is closed at
is distal end. Antenna design parameters, such as the axial coil length L,
coil turn density and specific wire material composing helical coil 14,
are predetermined for proper impedance match to biological tissue at a
given microwave frequency.
The exact number of coil turns and separation distance of gap G are
predetermined for each antenna application, along with the coil turn
density and connection configuration of the coil to the coax feed line
which are essential determinates of the radiated field. In addition to the
ability of antenna 10 to effectively confine heat to the region
immediately surrounding helical coil 14 and to efficiently radiate such
heat independently of implanted depth, the size of the heated voluem can
be readily adjusted by adding more such antennas to the array, by changing
frequency and/or by changing length L of the helical coil for a given coil
turn density. Additional adjustments to the heating field may be
accomplished by using variable spacing and variable diameter of the coil
turns to shape the power deposition pattern along the antenna length.
Since the dead length (cold portion at the end of the implanted antenna)
is eliminated, the need for implanting the antenna deeper than the lower
extent of the target region is minimized. Also, undesirable tissue heating
along the antenna feedline is eliminated along with overheating of the
tissue surface for shallow depth insertions.
Comparative testing of antenna 10 against the simple dipole antenna 17
(FIG. 3) showed that the power deposition characteristic of the standard
antenna varied significantly for varying insertion depth (FIG. 4) in
contrast to the power deposition pattern of antenna 10 which remained
essentially constant regardless of insertion depth (FIG. 2). In use,
implanted antenna 17 is also entirely covered with a plastic catheter 16
(FIG. 1), closed at its distal end. The ordinates "% maximum temperature
rise" in FIGS. 2 and 4 also depict "relative specific absorption rate."
DETAILED DESCRIPTION
The specific antenna embodiment 10 used for comparative testing purposes
described below constituted a 0.095 cm OD semi-rigid coaxial cable sized
to slip within plastic catheter 16 (FIG. 1). The outer diameters of inner
conductor 11, insulator 12 and outer conductor 13 were 0.02 cm, 0.061 cm
and 0.095 cm, respectively. Gap G, between the distal end of outer
conductor 13 and the proximal end of helical coil 14, was 0.1 cm whereas
length L of the helical coil was 1.0 cm. The distal end of the helical
coil was soldered to the distal end of the inner conductor at 15. The
inner and outer conductors were composed of a standard copper based alloy
having high electromagnetic wave energy transmission properties whereas
insulator 12 was composed of a standard Teflon (polytetraflouroethylene)
material.
Helical coil 14 was also composed of a metallic conductor having high
microwave energy transmission properties. For example, antennas
constructed with 0.032 cm nichrome, 0.032 cm varnish insulated copper, and
0.0203 cm silver-plated copper wire all have been used successfully. The
outside diameter of the helical coil closely approximated 0.12 cm to
facilitate insertion of antenna 10 into standard 16-gauge plastic catheter
16.
Helical coil 14 was formed by wrapping a wire tightly around a stainless
steel wire form, dimensioned to provide the desired diameter, length and
turn density of helical coil 14. After extracting the wire form from the
formed helical coil, the helical coil was installed carefully over the
bare dielectric insulator portion of insulator 12, as shown in FIG. 1, and
soldered at 15. Gap G was set at 0.1 cm.
In the preferred embodiments of this invention, axial length L and the turn
density of helical coil 14 provide an impedance match of a microwave
generator to an organic subject (e.g., tissue to be treated) at a
microwave frequency ranging from approximately 0.1 to 3.0 GHz. Helical
coil 14 preferably has a length L selected from the approximate range of
from 1.0-10.0 cm and has equally spaced turns in the approximate range of
7 to 16 turns per cm. More sophisticated applications of this invention
are anticipated with variable turn density along the coil to customize the
heating field shape.
SURGICAL PROCEDURE
The following discussion briefly summarizes a typical surgical procedure
using the above-described helical coil microwave antenna 10 for improving
the localization of interstitial hyperthermia. An appropriate length
antenna is first chosen to provide the desired heating pattern in
accordance with the above discussions. For 915 MHz operation, a nichrome
or copper wire coil with axial length L in the range of 1.0-10.0 cm may be
selected for proper heat localization to the coil tip. Depending on target
size, multiple-antenna array operation is possible to expand the effective
heating volume.
A parallel array of 16-gauge plastic catheters 16 are inserted into the
target region of a patient with the aid of a .ltoreq.1.0 cm grid template
or stereotactic surgical frame. Metal needles or stainless steel stylets
are normally used to guide the catheters in place. A standard CT scan or
simple X-ray will verify proper location of the catheters relative to the
tumor or other tissue to be treated.
Antennas 10 are inserted into the desired catheters 16 and extended to the
appropriate insertion depth. Standard temperature sensing probes are
inserted into other plastic catheters to monitor and control the
temperature distribution. If it is found that one or more of the antennas
are not well matched electrically to the tissue load, minor adjustments
can be made to the helical coil turn spacing or a double stub tuner can be
used to obtain the best match of antenna to generator. Each antenna is
connected to a microwave generator via a flexible coaxial cable in a
conventional manner.
Microwave power can be controlled either manually or automatically by a
computer feedback system to maintain the desired minimum tissue
temperature (typically 43.degree. C. for one hour, or equivalent). The
temperature probes are manually translated inside the respective plastic
catheters to monitor temperatures at approximately 1 cm increments within
the heated tissue volume several times during the treatment to provide
information on the internal temperature distribution of the tumor volume.
At the end of treatment, microwave power is terminated and all antennas
and sensors are removed from the catheters. Interstitial hyperthermia
therapy induced via the implanted antennas can be repeated before and
after radioactive seed brachyradiotherapy without additional surgery,
using the same implanted plastic catheters 16.
COMPARATIVE EVALUATION WITH EXISTING TECHNOLOGY
Standard dipole antenna 17 of FIG. 3 was constructed from a 0.095 cm
semi-rigid coaxial cable having an inner conductor 18 covered by an
insulator 19. An outer conductor 20 was cut circumferentially at 21 to
form an axial separation gap of 0.1 cm between the cut portions of the
outer insulator. A metallic connector 22 was soldered between the distal
ends of the inner and outer conductors. Previous tests have shown this
simple "Dipole" type structure operates identically to other dipole styles
having a soldered connection (not shown in FIG. 3) of inner conductor 18
to outer conductor 20, adjacent to gap 21.
The specific polarization pattern of HCS style antennas is highly dependent
on the source frequency, length of helix, spacing and diameter of coil
turns 14.
Antenna 10 with a coil turn density of 10 turns/cm, a gap G of 0.1 cm, a
coil diameter of 0.12 cm and a coil length of .lambda./4 of 1.0 cm (3.5
cm) was found to generate apparently circularly polarized electromagnetic
waves which effectively localized the heating to the region surrounding
the coil when driven at 2450 MHz (915 MHz). Other similar helical coil
antennas with lengths L ranging from 1-5 cm, turn densities of 7-16
turns/cm and gap G from 0.1-5 cm also have been tested successfully at the
two frequencies.
For comparative dosimetry study purposes and to match heating efficiencies
of the antennas, each antenna was tuned to the coaxial feedline with a
double stub tuner. The tuners were capable of precisely matching the
antennas to the source frequency and feedline characteristics. This
procedure enabled a direct comparison of antenna performance under optimum
matched conditions, regardless of tissue properties or antenna insertion
depth.
In order to study the antenna heating characteristics in a reproducible,
homogenous tissue medium, soft tissue phantom was used initially to obtain
the relative heating profiles of the different antenna configurations as a
function of insertion depth. The phantom was composed of a mixture of
distilled water (75.2%, base), TX-150 (15.4%, gelling agent), sodium
chloride (1.0%, to adjust electrical conductivity), and polyethylene
powder (8.4%, to lower dielectric constant). The mixture is known to have
approximately the same electrical properties as those of human soft tissue
at 915 MHz. A similarly appropriate phantom mixture was used for studies
at 2450 MHz.
The material was contained in an 8.times.8.times.11 cm plexiglass box
transversed by a 0.5 cm array of 16-gauge Teflon catheters for holding the
antennas and multi-sensor temperature probes. Five phantom models were
constructed during the course of the experiments and the reproducibility
of thermal profiles in each phantom was verified, using both antennas 10
and 17. To evaluate the difference in axial thermal profiles (FIG. 2 vs.
FIG. 4) produced by each antenna type, single antennas were inserted into
a catheter which was immersed in the phantom so that gap G (FIG. 1) or gap
21 (FIG. 3) was located 1.0 cm below the phantom surface. The total
insertion depth of 2.1 cm (approximately .lambda./2 in tissue) was
considered near optimum for the standard dipole antenna 17.
Axial power deposition profiles of antennas 10 and 17 were compared for
total insertion depths of 1.35 cm, 2.1 cm and 3.1 cm in phantom, as
illustrated by each of the three curves in FIGS. 2 and 4. These tests were
intended to model the three clinically relevant conditions of antenna use:
too shallow, optimum implant depth, and too deep. Radial power deposition
profiles in several planes perpendicular to the antennas were obtained for
a single insertion depth of 2.1 cm. These profiles were then compared to
the radial temperature fall-off obtained using a heated water circuit of
similar dimensions as a control for strictly thermal conduction heating.
The antennas were both driven at either 915 MHz or 2450 MHz using a
continuous wave microwave power source (Model CA 2450, manufactured by
Cheung Laboratory, Inc., Lanham-Seabrook, Md. Power fed to each antenna
was tuned with a double stub tuner (Model 1729, Maury Microwave,
Cucamonga, Calif.) for optimum impedance match to the generator, since no
attempt was made to trim each antenna to exactly 50 ohms. Since the
phantom material had no cooling effect from circulating blood, all
experiments consisted of short 30 sec. heat trials during which the rate
of change of temperature was determined at all internally monitored points
to represent the power deposition characteristics of the antennas.
The thermal profile information was obtained using a multiple-sensor
optical fiber probe with four sensors spaced 0.25 cm apart with each
inserted in a catheter parallel to and 0.5 cm away from the antenna axis.
Using a separate stationary single sensor probe located mid-depth in a
second parallel catheter as control between trials, longer axial heating
profiles were measured by moving the multi-sensor probe 1 cm and repeating
the heat trial after cool-down of the phantom to initial conditions.
All temperatures were recorded every 10 seconds by a computerized fiber
optic thermometry system and displayed in tabular and graphic forms on a
color monitor. The increase in temperature above baseline [.DELTA.T was
calculated for each point and the measure of power deposition (Specific
Absorption Rate or SAR) was determined from the time rate of change of
temperature following power on as SAR=c.multidot.d.DELTA.T/dt, wherein
c=specific heat of phantom tissue].
The axial thermal power deposition profile of each antenna was determined
independently at four different sites within the phantom box for each
experiment. The antennas were tested in more than one phantom to minimize
erroneous conclusions that might arise from slight catheter placement
variations at depth in the phantom or other systematic test errors. Axial
profiles from corresponding trials were averaged together by first
selecting the maximum SAR of each linear distribution as 100% SAR and
normalizing the profile to a percentage of the peak (Relative SAR).
FIGS. 2 and 4 compare the effects of varying insertion depth on the axial
power deposition profiles of antennas 10 and 17. As noted in FIG. 4, the
thermal profile of standard antenna 17 varied significantly, depending on
insertion depth. With a 3.1 cm total insertion depth, the thermal profile
was almost symmetrical with the peak located 1.5 cm below the surface and
a 50% HL and Dead Length of 2.04 and 0.68 cm, respectively. With a shorter
insertion of 1.35 cm (gap 21 depth of 0.25 cm), the 50% HL and Dead
Lengths were both drastically reduced to 1.21 cm and 0.17 cm,
respectively, but the antenna entrance point was overheated with 78% of
the peak SAR obtained near the surface.
In contrast, the power deposition profiles induced by antenna 10 were
essentially identical regardless of insertion depth. The heat peak moved
correspondingly deeper with increasing insertion, remaining 0.5 cm
proximal to the antenna tip midway along the axis of helical coil 14 (FIG.
1). The 50% HL for 1.35, 2.1, and 3.1 cm insertion depths was a constant
1.2 cm and the Dead Length remained essentially 0.0. Studies on the
reproducibility of profiles for antennas 10 and 17 tested identically in
five different phantoms disclosed no significant variation in the location
of Peak Depth, 50% HL, or Dead Length.
Relative radial power deposition profiles perpendicular to the axes of
antennas 10 and 17 were also obtained and compared with the radial
temperature fall-off from a heated water circuit of similar dimensions.
The corresponding comparative radial temperature distributions for L=3.5
cm antennas 10 and 17 driven at 915 MHz in dog thigh muscle tissue in vivo
are shown in FIG. 5. FIG. 6 gives the absolute temperature distributions
in dog thigh muscle along the axial length of antenna 10 (at R=0.5 cm
distance) for three different clinically relevant antenna insertion
depths. Note the very similar 50% HL's and slopes of the individual
temperature distributions for the three different implant conditions.
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