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BACKGROUND OF THE INVENTION
This invention pertains to a catheter designed to couple radiofrequency
(RF) energy to biological tissue surrounding the catheter tip. Typical
application is in thermal ablation of cardiac tissue.
Percutaneous ablation is a therapeutic procedure used with increasing
frequency for treatment of ventricular tachycardia. It works by destroying
cardiac tissue responsible for the disease. For example, this subject is
covered in Ablation in Cardiac Arrhythmias, G. Fontaine & M. M. Scheinman
(Eds.), Futura Publishing Company, New York, 1987. A recent review of this
field is given in a chapter by D. Newman, G. T. Evans, Jr., and M. M.
Scheinman entitled "Catheter Ablation of Cardiac Arrhythmias" in the 1989
issue of Current Problems in Cardiology, Year Book Medical Publishers.
Currently, catheter ablation is performed by delivering a high voltage
direct current pulse from a standard defibrillator through an electrode
catheter designed for temporary pacing. Radiofrequency (RF) ablation using
electrosurgical power units is in clinical investigation, as a safer
ablation alternative to high voltage direct current pulses. Continuous,
unmodulated RF output in the frequency range of 500 KHz to 1.2 MHz is
typically used. (RF without qualifiers refers to the electromagnetic
spectrum from 10 kHz to 100 GHz.) Laser energy is also being tested for
catheter ablation of arrhythmias.
Some experimentation has been reported with the use of microwave energy for
catheter ablation. U.S. Pat. No. 4,641,649 issued Feb. 10, 1987 to P.
Walinski, A. Rosen and A. Greenspon describes a catheter consisting of a
miniature coaxial line terminated in a protruding inner conductor antenna.
This system operates at 925 MHz. Another microwave ablation catheter
experiment has been reported by K. J. Beckman, & J. C. Lin et al,
"Production of Reversible Atrio-Ventricular Block by Microwave Energy"
abstracted in Circulation 76 (IV): IV-405, 1987. Technical details of a
folded dipole applicator catheter used by Beckman have been described by
J. C. Lin and Yu-jin Wang in "An Implantable Microwave Antenna for
Interstitial Hyperthermia" in Proceedings of the IEEE, Vol. 75 (8), p.
1132, August, 1987. An earlier microwave applicator which fits into a
blunt-ended mylar catheter has been described by B. E. Lyons, R. H. Britt,
and J. W. Strohbehn in "Localized Hyperthermia in the Treatment of
Malignant Brain Tumors Using an Interstitial Microwave Antenna Array",
IEEE Trans on Biomedical Engineering, Vol. BME-31 (1), pp. 53-62, January,
1984.
A general geometrical requirement of catheter-based applicators is that
they must be confined in slender cylindrical structure with a radius
commensurate with the catheter diameter. In the discussion of catheter
applicators which follows, it is convenient to adopt a cylindrical
coordinate system with the z-axis coincident with the catheter axis and
pointed toward the distal end. The radial component is at the direction
normal to the z-axis and the circumferential component has a direction
around the z-axis. Radius "r" is measured from the catheter axis. The
catheter diameter is "a".
A common feature of all of the above RF catheters (the laser catheter which
is an optical device will not be discussed further) is that the energy
delivery is predominantly via an electric field (E-field) originating at
the applicator's electrode/tissue interface. This class of catheter
applicator will therefore be referred to as E-field applicators. Although
the configurations of the E-field applicators described above vary, the
E-field coupling causes a rapid decrease in current density and therefore
tissue heating, as a function of distance from the electrode.
In order to represent the state of the art of RF heating catheters and to
compare it with the preferred embodiment of this invention, a simplified
E-field applicator is shown in FIG. 1A. Applicator electrode 10 is a wire
immersed in a lossy dielectric medium which has electrical properties
typical of muscle tissue. In spite of the simple geometry and low
frequency approximation used in the description, FIG. 1 retains the
salient feature of an E-field coupling.
In FIG. 1A, RF potential V14 is applied between applicator electrode 10 and
a remote boundary 15 which corresponds to a neutral electrode applied to
the skin. The exact location of boundary 15 is not important to the shape
of the E-field near applicator electrode 10. Radial electric field E16
coincides with current density vector J.sub.r =.sigma.E.sub.r in the
tissue, where .sigma. is the conductivity of the tissue.
Continuity of current in a cylindrical geometry in FIG. 1 results in
current density which decreases with the inverse square of the radius r.
Therefore, corresponding electrical power dissipation resulting in heating
of tissue decreases with the fourth power of a/r. Typically, an electrode
radius is limited by practical catheter size to a maximum of 1 mm. In
order to effectively ablate ventricular tachycardia (see Moran, J. M.,
Kehoe, R. F., Loeb, J. M., Lictenthal, P. R., Sanders, J. H. & Michaelis,
L. L. "Extended endocardial Resection for the Treatment of Ventricular
Tachycardia and Ventricular Fibrillation", Ann Thorac Surg 1982, 34:
538-43), it is desirable to heat tissue up to 10 mm from the catheter
axis. In the applicator represented by electrode 10 in FIG. 1A, heat
dissipation at the catheter surface is 10,000 times more intense than heat
dissipation at a 10 mm radius.
In order to examine the effect of this wide range of heat dissipation, it
is useful to divide the lossy medium in FIG. 1A into three cylindrical
shells: first shell R11 adjacent to the applicator electrode 10, followed
by shell R12, and R13 beginning at the 10 mm radius. Since the shells are
traversed by the same current and the potential drop across the shells is
additive, energy delivery can be represented by three resistances R11,
R12, and R13 in FIG. 1B, connected in series with the source of RF
potential V14.
A very steep heating gradient at the applicator electrode 10 tends to
desiccate blood or tissue close to the electrode, increasing the
resistivity of R11. This in turn further increases the relative power
dissipation in R11 in comparison with R12 and R13, until effective
impedance of the desiccated region R11 becomes, in effect, an open circuit
shutting off the flow of RF power to the tissue beyond R11. This indeed is
the problem of state-of-the-art RF ablation catheters which severely
limits effective heat delivery to more distant tissue.
Insulation of the applicator electrode 10 from the tissue does not reduce
the heat dissipation gradient: If the applicator electrode 10 is insulated
from the lossy medium by a thin dielectric tube, the effect of the
dielectric can be represented by capacitor (not shown) in series with the
source of RF potential V14. Now the applicator must be operated at a
frequency high enough so that the impedance of the sum of resistances R11
and R12 and R13 must be higher than the capacitive impedance of the
dielectric tube. R11 still dominates the heat distribution.
Therefore in biomedical applications, there is a need for a
catheter-compatible RF energy delivery system which dissipates heat more
uniformly to a specified depth, thereby leading to a more accurately
controlled and larger ablated region. It is also desirable to eliminate
the effect of desiccation of tissue adjacent to the electrode on heat
dissipation to surrounding tissue.
OBJECT OF THE INVENTION
Accordingly, the principal object of the invention is an RF energy
applicator which is housed in a biomedical catheter, typically of 2 mm
diameter. This applicator exhibits deeper and more uniform heat
dissipation and is not subject to power reduction from desiccation of
tissue in the proximity of the applicator, typical of state-of-the-art
devices.
A further object of the invention is a cardiac ablation system which
provides monitoring and control of RF power fed to the catheter and which
also provides signal processing, monitoring and display of the
intracardiac electrogram. In this application, the RF energy applicator is
configured to allow recording of intracardiac electrograms in proximity to
the catheter tip. This is important in order to accurately localize the
cardiac tissue to be ablated.
A still further object of the invention is a localized hyperthermia system
for cancer treatment, where the catheter with RF energy applicator offers
adjustable depth of heating compatible with a tumor size. This system also
provides monitoring and control of RF power fed to the catheter for proper
thermal dosimetry.
Further advantages of the invention will become apparent from the
consideration of the drawings and the ensuing description thereof.
BRIEF DESCRIPTION OF THE DRAWINGS
FIG. 1A shows the electric field (E-field) applicator represented by a
conductor immersed in a lossy dielectric medium (state of the art).
FIG. 1B is an equivalent circuit of heat delivery of an E-field applicator
(state of the art).
FIG. 2A shows a solenoidal applicator in the form of a helix immersed in a
lossy dielectric medium.
FIG. 2B is an equivalent circuit of heat delivery of a solenoidal
applicator.
FIG. 3 shows details of a catheter tip mounted solenoidal applicator with
intracardiac electrogram monitoring capability.
FIG. 3A and FIG. 3B show magnified details of the circled portions of FIG.
3.
FIG. 4 is a block diagram of RF heating and intracardiac electrogram
monitoring ablation catheter system.
DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT
FIGS. 1A and 1B, which illustrate the problems inherent to state-of-the-art
E-field applicators, have already been discussed in the Background section
above.
FIG. 2 shows a conductor in the form of a helix 20 traversed by RF current
I24. A helix radius in a catheter application is typically a=1 mm and
maximum desired radius of tissue heating for cardiac ablation is R=10 mm.
The resultant magnetic field typified by H21, H22, and H23 has primarily
an axial (z) component and induces a transverse electric field E.sub..phi.
typified by E21, E22, and E23 (and a proportional current density not
shown), primarily in the circumferential direction around the helix. The
circumference of E22 corresponds to R.
At an operating frequency of 915 MHz, and within a cylinder of a radius of
10 mm, the magnitude of induced electric field E.sub..phi. (and
corresponding current density J=.sigma.E.sub..phi.) decreases
approximately as a/r, where a is the radius of the helix. The z component
of electric field E.sub.z, which follows field lines similar to H.sub.21,
H.sub.22 and H.sub.23, decreases with radius r even more slowly than a/r.
Hence the heating power dissipation decreases no faster than (a/r).sup.2.
For a typical catheter radius a=1 mm and desirable depth of heat
penetration in ablation R=10 mm, the ratio of heat dissipation at the
catheter wall to heat dissipation at R=10 mm is approximately 100:1. It
should be noted that this is a major improvement over the ratio of heat
dissipation for the E-field applicator which for the same conditions is
10,000:1.
With a solenoidal applicator, the effective heat dissipation radius R can
be adjusted: R increases with decreasing frequency. For ablation of
cardiac arrhythmias, ISM (industrial, scientific, and medical) frequencies
of 915 MHz and 2450 MHz are of interest. For hyperthermia treatment of
cancer, a wider gamut of frequencies is needed depending on the size of a
tumor.
FIG. 2B shows an equivalent circuit of a helical heating applicator. Within
a volume inside a radius of 10 mm, a circular induced electric field
E.sub.100 multiplied by the length of its circumference gives a potential
around each cylindrical shell which is approximately equal. The shell of
lossy medium adjacent to the helix, energized by E21, the shell at the
intermediate distance energized by E22, and the shell corresponding to
R=10 mm energized by E22, appear in FIG. 2B as parallel resistances R21,
R22, and R23 respectively, exposed to the same potential. Current source
I24 feeds the three resistances.
Now, if desiccation occurs adjacent to the helix, resistance R21 increases.
This reduces power dissipation in R21 and increases power dissipation in
resistances R22 and R23. In general then, as power is increased to a point
of desiccation at a catheter surface, the heat delivered to a desiccated
volume decreases in a solenoidal applicator while it increases in an
E-field applicator. Thus, the solenoidal applicator is much less likely to
cause excessive desiccation but even if desiccation occurs, it will not
lead to a decrease in power dissipation in remote tissue at R=10 mm.
The helix in FIG. 2A is an example of a solenoidal applicator structure,
characterized in general by current loop or loops and an electrical short
as an end termination. Solenoidal applicator generates a magnetic field in
the surrounding tissue. This magnetic field by induction generates in
turn, an electric field and current which heats the tissue. In contrast,
the E-field applicator has an electrical open end termination and the
primary, rather than induced, electric field heats the tissue.
The preferred embodiment of the helical solenoidal applicator in an
ablation catheter is shown in FIG. 3. Coaxial line 43 consists of a center
conductor 44 (0.16 mm diameter), a dielectric 46 (1.35 mm outside
diameter), a metal braid 45 and insulating sleeve 57 (1.8 mm outside
diameter). Small diameter and flexible construction makes the coaxial line
43 suitable for a biomedical catheter application. Helical winding 50 is
wound on a ceramic or ferrite core 51. A heat-shrunk TEFLON(TFE:
Tetrafluoroethylene sleeve 53 covers the helical winding 50.
Distal end of the helical winding 50 is connected at distal peripheral
terminal 58 to distal electrode 56 and to bypass capacitor 55. Bypass
capacitor 55 is connected to braid 45 through metallized coating 52 on the
inside of core 51. The function of the bypass capacitor 55 is to ground
the RF energy. Thus during RF current flow through helical winding 50,
distal electrode 56 has no RF voltage thereby preventing E-field heating.
Distal electrode 56 in conjunction with a proximal ring electrode 47 picks
up a cardiac electrogram voltage between them. The distance from the
beginning of proximal ring electrode 47 to the end of the distal electrode
56 is 20 mm.
A number of turns on the helical winding 50 is chosen so that at an
operating frequency of 915 MHz, the helix is somewhat short of being a
quarter wavelength resonator. The proximal end of the helical winding 50
is connected to a variable tuning capacitor 48 at proximal peripheral
terminal 49. Variable tuning capacitor 48 is moved with respect to neutral
electrode 47 during manufacture for tuning to a precise quarter wavelength
resonance. Details of the tuning capacitor 48 are shown magnified in FIG.
3A.
RF power is coupled into the helical resonator by connecting the center
conductor 44 to the helical winding 50 at feed terminal 54. The connection
at feed terminal 54 is shown magnified in FIG. 3B. The position of feed
terminal 54 on the helix is selected for good match between the
characteristic impedance of the line and the impedance of the resonator.
The choice of an axial quarter wavelength resonator is by no means unique.
One could just as well select any multiplicity of quarter wavelengths
e.g., half wavelength or full wavelength resonators.
In some applications, it may be desirable to distort the axisymetrical form
of the induced E-field. This can be accomplished by partially covering a
dielectric sleeve 53 with metal foil (not shown). Currents induced in such
foil modify the shape of a heating pattern and so serve as an aperture
antenna. An asymmetrical field pattern can also be accomplished by a loop
antenna.
In cardiac ablation, it is highly desirable to be able to monitor
intracardiac electrogram just before and after the application of heat.
FIG. 4 shows a block diagram of a system which combines dosimetry control
of the solenoidal heat delivery with monitoring of intracardiac
electrograms.
The RF power is generated in an RF power source 41. The RF power is
controlled and monitored in controller 42 which couples the RF power to
the coaxial line 43 through capacitor 62, which for RF represents
substantially a short-circuit.
The center conductor 44 is attached at feed terminal 54 to the helical
winding 50 wound on a core 51. Quarter wavelength resonance tuning is
accomplished by adjustment of variable tuning capacitor 48 connected to
the helical winding 50 at proximal peripheral terminal 49. The RF ground
is maintained by the bypass capacitor 55 connected to the distal electrode
56 and then to helical winding 50 at distal peripheral terminal 58.
Distal electrode 56 in conjunction with the proximal ring electrode 47
picks up the local intracardiac electrograms and feeds this electrogram
signal through the coaxial line 43 to capacitor 62. Capacitor 62
represents a short circuit for the RF power and an open circuit for the
much lower frequency band (typically 0.1 Hz to 100 Hz) associated with
intracardiac electrogram activity. The electrogram signal appears
therefore on lines 63 and 64 at the input to the low-pass filter 61.
Filter 61 has a high input impedance to the RF and hence has no effect on
transmission of RF power between controller 42 and coaxial line 43. Filter
61 blocks the transmission of the RF power to switch 60 while allowing
passage of the electrogram signal. Switch 60 is closed simultaneously with
application of RF power, thus providing additional protection for monitor
59. Electrogram signal processing, display, and recording is accomplished
by monitor 59. Standard existing equipment is suitable for application as
monitor 59.
Solenoidal catheter applicator for hyperthermia treatment of tumors follows
largely the same design as the one represented in FIG. 3 except that in
this case, there is no need for the distal electrode 56 and the proximal
ring electrode 47. Since the depth of heat penetration depends inversely
on the square root of frequency, the frequency of operation and the
helical winding design can be tailored to the required depth of
penetration depending on tumor size.
While certain specific embodiments of improved RF heating applicator and
systems have been disclosed in the foregoing description, it will be
understood that various modifications within the scope of the invention
may occur to those skilled in the art. Therefore it is intended that
adaptations and modifications should and are intended to be comprehended
within the meaning and range of equivalents of the disclosed embodiments.
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