Temperature Controlled Radiofrequency Ablation
Olaf J. Eick, PhD.
Sr. Scientist, Bakken Research
Center Maastricht, the Netherlands
Address for correspondence: Olaf J.
Eick, PhD, Sr. Scientist, Bakken Research Center Maastricht, the
Netherlands. Email: olaf.eick@medtronic.com
Since its introduction in 1987, radiofrequency (RF)
ablation has developed to become the treatment of choice for symptoms
caused
by atrio ventricular (AV) reentrant tachycardia, isthmus related atrial
flutter, AV-nodal reentrant tachycardia and to some extent also for
certain
types of ventricular tachycardias. The introduction of new cardiac
activation mapping systems has further contributed to the successful
and safe application of RF ablation for various tachyarrhythmias.
Historical Background
In
1979, the first catheter ablation was performed coincidentally when
Fontaine et al. observed a complete AV block in a patient undergoing
defibrillation while one defibrillator electrode was in electrical
contact with a catheter electrode positioned at the bundle of His1. In the early 80's high voltage direct current
ablation was further developed by Gallagher and Scheinman and used as a
therapeutic approach to treat
supraventricular tachycardias2,3
. Although highly effective, direct current ablation was accompanied by
severe complications such as cardiac tamponade, hypotension following
shock delivery and the induction of ventricular arrhythmias, which
stimulated
the search for alternative energy sources. Huang et al. investigated in
1987 experimentally the use of radiofrequency energy for catheter
ablation4, followed by the first AV-node
ablation by Budde
et al.5 and the interruption of an
accessory
pathway by Borggreffe et al.6
The
advantages of radiofrequency energy as compared to direct current are
the generation of well circumscribed lesions without stimulation or
sensory effects. In order to more appropriate titrate radiofrequency
power, catheter tip temperature monitoring was investigated
experimentally in 1989 by Hindricks et al. and Haines et al.7. In 1992 Langberg et al. published their
observations of using temperature monitoring with a thermistor embedded
in the distal electrode during catheter ablation
of accessory pathways in 20 patients8.
As a result of a constant power delivery, the tip temperature increased
gradually to reach a steady state within a few seconds. Neither the
applied power nor the electrogram characteristics at the ablation site
(amplitude, signal stability) could predict tip temperature. Lower
temperatures were found when transient blocks had been induced (50
± 8° C) than when permanent blocks resulted in accessory
pathway ablation (62 ± 15° C). Complications were not
reported. Consequently, a closed loop temperature control system was
introduced to regulate the applied power such
that a preset target temperature could be reached and maintained during
radiofrequency delivery. Calkins et al reported in 1994 a multicenter
experience with 270 consecutive patients undergoing catheter ablation
of supraventricular arrhythmias9. Compared
to the power control mode, application of radiofrequency energy
delivered using the temperature control mode was associated with a
threefold reduction in the incidence of developing
a coagulum and a fivefold reduction in the incidence of automatic power
shutdown due to an impedance rise or an electrode temperature of
greater than 100° C. Whether the target temperature was reached or
not was related to the ablation site, with ablation of the
atrioventricular junction resulting in the highest temperatures and
ablation of AVNRT the lowest. In addition, the temperature of the
applications of radiofrequency energy that resulted in successful
ablation was not significantly different from the temperature of those
that failed.
Tissue Heating
In
RF ablation, the heating of cardiac tissue is mainly resistive. RF
current is applied to the tissue via a metal electrode at the tip of
the catheter, with a large skin electrode serving as indifferent
electrode. The current density patterns in the tissue are determined by
electrode size and geometry, electrode contact and local tissue
properties. Also, of course, the current density will be proportional
to the current (I) delivered by the RF generator, which, for constant
resistance (R) of the electrode-tissue volume conductor is proportional
to the square root of the RF power (P = I2 R).
Current flow through a resistive medium causes heating, which is
proportional to the square of the current or, locally, to the square of
the local
current density. The temperature increase will locally be proportional
to the energy applied per second (local power), and increases inversely
proportional to the heat capacity of the local medium. In addition,
when
temperature differences between adjacent areas develop because of
differences
in local current density or local heat capacity, heat will conduct from
"hotter" to "colder" areas, causing the temperature of the former to
decrease
and that of the latter to increase. Additionally, heat loss to the
blood
pool at the surface and to intramyocardial vessels determine the
temperature profile within the tissue.
Temperature Distribution
The
heating occurs especially in the proximity of the active electrode due
to its relatively small surface area causing locally high current
density as compared to the site of the indifferent electrode.
Typically, living tissue will be permanently destroyed at temperatures
of approximately 45° to 50° C sustained for several seconds10.
The
tissue surface is cooled by the blood flow and thus the highest
temperature during radiofrequency delivery occurs slightly below the
surface. In a numerical model, the temperature distribution during
radiofrequency energy was simulated 11.
The model described a 7
F electrode, 4 mm in length positioned perpendicularly to the tissue,
and
considered heat convection within the tissue and heat loss due to
convective cooling (Figure 1). Figure 2a, b, and c illustrate the temperature behavior at the
electrode tip and 1, 3 and 5mm deep within the tissue for different
flow conditions with a target temperature of 70° C, i.e. a
temperature increase of 33° C, and a power limit of 50 W.
The catheter tip temperature rises rapidly at the beginning. The
temperatures within the tissue rise more slowly than the tip
temperature because it
takes time for heat to transfer from the hotter rim around the
electrode
into the tissue. After 10 s the temperature at 1 mm becomes higher than
the electrode temperature for high flow conditions (Figure 2a) and after 50 s with intermediate
flow (Figure 2b) whereas the temperature is
always highest at the tissue surface if there is no convective cooling (Figure 2c). In addition, the temperatures at a
certain depth are higher under high flow condition resulting in deeper
lesions as compared to those under moderate flow and no flow,
respectively.
Figure 1:
Model for radiofrequency ablation simulating a 7 F electrode, 4 mm in
length and positioned perpendicularly to the tissue. Heat convection
within the tissue and heat loss due to convective cooling were
considered and temperatures over time and distance were calculated by
solving the appropriate heat equations.
Figure 2: Temperature
behavior over time at the electrode tip and 1, 3 and 5mm deep within
the tissue for different flow conditions (a, high flow; b, intermediate
flow and
c, no flow) with a target temperature of 70° C, i.e. a temperature
increase of 33° C and power limit of 50 W.
Figure 2a
Figure 2b
Figure 2c
This
temperature behavior was confirmed experimentally by Cao et al. who
performed temperature measurements in vitro with multiple thermocouple
needles at different depths and recorded the temperature behavior
during ablation 12. The catheter tip
temperature rose rapidly at the beginning. It approached the target
temperature (60° C) in about 2 to 3 s. Then the RF generator
adjusted the delivered power to maintain the target temperature. When
the ablation stopped, tip temperature dropped rapidly due to convective
cooling by the blood flow. The temperatures measured within the tissue
(0.8 mm, 2.0 mm, and 2.9 mm) rose more slowly and after 30 s the
temperature at 0.8 mm depth became higher than the
electrode temperature.
The
heating of cardiac tissue by radiofrequency current thus depends on the
current density distribution in the tissue, the tissue's heat capacity
and heat conduction properties and heat convection.
Catheter Tip Temperature
The
catheter tip temperature in turn depends on the tissue temperature but
also on convective cooling by the surrounding blood, and the tissue
contact of the ablation electrode 13, 14, 15. In addition, the
measured catheter tip temperature depends on the electrode material
with its heat capacity and the type and location of the temperature
sensor. There are mainly 2 different types of temperature sensors, i.e.
a thermocouple and a thermistor. The thermistor requires a driving
current and the electrical resistance changes as the temperature of the
electric conductor changes. More frequently used are thermocouples
based on the so called 'Seebeck effect'. When 2 different metals are
connected (sensing junction) a voltage can be measured at the reference
junction that is proportional to the temperature difference between the
2 junctions. For temperature measurements during radiofrequency
ablation typically type T thermocouples are being used that consist out
of copper and constantan wires and are incorporated in the center of
the ablation electrode.
Relation between Tip temperature, Tissue Temperature and Lesion
Size
Radiofrequency current heats cardiac tissue and in turn the catheter
electrode is being heated. Consequently, the catheter tip temperature
is always
lower - or ideally equal - than the superficial tissue temperature.
The
catheter tip temperature, which is measured and used to control the
radiofrequency power output, can be significantly lower than the tissue
temperature 16,17.
Kongsgaard et al. evaluated the differences between tip temperature and
tissue temperature in vitro for different catheter electrode dimensions
and flow rates during temperature controlled radiofrequency delivery16. With a thermosensor placed approximately 1 mm
beneath a 4 mm catheter, the
temperature within the tissue was on average 42 ± 6° C
higher
than at the electrode tip after 30 seconds radiofrequency delivery with
a preset target temperature of 70° C.
Only
under standardized and stable experimental conditions of flow and
electrode-tissue contact the tip temperature correlates to lesion size
because increasing target temperature then increases power consumption
and in turn increases lesion size 7,14.
During temperature controlled radiofrequency ablation tip temperature,
tissue temperature and lesion size are affected by electrode-tissue
contact and cooling effects by the blood flow. Petersen et al.
investigated the effect of convective cooling on lesion dimensions 18. In vivo, 2 different application sites in the
left ventricle of pigs were ablated with a target temperature of
80° C using a 4 mm tip catheter. With higher convective cooling at
septal sites, the delivered power to reach target temperature was
higher resulting in deeper and greater lesions than at apical sites.
The applied power was positively related to lesion volume (r= 0.66)
whereas the measured tip temperature was not (r= -0.49).
It
is important to understand that radiofrequency power delivered to the
tissue determines lesion size and that the catheter tip temperature is
poorly correlated to lesion size in vivo. With good contact between
catheter tip and tissue and low cooling of the catheter tip the target
temperature can be reached with little power resulting in fairly small
lesions although a high tip temperature is being measured. In contrast,
a low tip temperature can be due to a high level of convective cooling
resulting in high power consumption to reach target temperature
yielding relatively large lesions. This is best illustrated with active
cooling of the catheter tip using irrigation during radiofrequency
energy delivery. The tip temperature is usually below 40° C but
that allows the application of high power output for longer durations.
Nakagawa et al. compared the lesion dimensions produced with
saline-irrigated electrodes with those produced with non-irrigated
electrodes in a canine thigh muscle preparation19.
With saline irrigation at 20 ml/min high radiofrequency voltage (66V)
could be applied over 60 s resulting in lesion volumes of 700 ±
217 mm3, whereas constant radiofrequency delivery was terminated
prematurely due to an impedance rise at tip temperatures of nearly
100° C resulting in lesions of
only 135 ± 33 mm3 . In the temperature controlled mode with a
target
temperature of 80° C lesions with a volume of 275 ± 55 mm3
were
produced. These results were further confirmed in an animal study
performed by Skrumeda et al. who reported significantly larger lesions
using irrigation during constant power delivery as compared to
non-irrigated ablations either
in temperature or power controlled mode 20.
With
high irrigation flow rates catheter tip temperature is not
representative for tissue temperature and therefore feedback cannot be
used to control power output. However, Petersen et al. demonstrated the
feasibility of
tip temperature controlled irrigated radiofrequency delivery with low
irrigation flow rates yielding larger lesions compared to non-irrigated
ablations 21.
A
similar effect can be observed with large tip electrodes. With large
electrode tips a larger area of the electrode tip is exposed to the
blood flow than with standard tip lengths resulting in higher power
consumption and lower tip temperatures yielding larger lesions 22.
In
addition to cooling effects, the electrode tissue-contact affects the
catheter tip temperature and the difference between tip temperature and
tissue temperature. Avitall et al. evaluated the influence of electrode
tissue contact on temperature rise and lesion size under stable flow
conditions in an in vivo animal model during constant power delivery 23. They demonstrated a higher temperature rise
within the first 10 seconds and an increase in lesion volume with
improving electrode-tissue contact.
With
poor contact between electrode tip and tissue less electrode material
is in contact with the tissue and heating of the tip by the tissue
occurs at a smaller rate resulting in relatively low tip temperatures.
As a consequence, the RF generator increases power output in order to
reach target temperature.
If
the electrode-tissue interface temperature increases above the boiling
point due to inadequately high power delivery, any liquids under the
electrode will boil and vapor will form causing a popping sound. This
is often
associated with sudden impedance rise and catheter dislodgment and may
cause significant tissue damage 17,24.
As evaporation may also occur intramurally,
a gas bubble may develop within the tissue under the electrode.
Continuous
application of radiofrequency energy will cause the bubble to expand
and
its pressure to increase, which may lead to eruption of the gas bubble
through
the weakest path, leaving behind a gaping hole. The release of the gas
bubble
is associated with a popping sound and likely with tearing of cardiac
tissue 25. Furthermore, if blood is
between the
electrode and the tissue during delivery of radiofrequency energy, a
rapid increase in temperature may result in coagulum formation on the
metallic
surface of the electrode. Coagulum formation, in turn, causes a rapid
increase
in voltage across the rising impedance, which may result in arcing and
tissue charring.
Outlook
Since catheter tip temperature is affected by cooling effects and
electrode-tissue contact it would be ideal to monitor in real time the
tissue temperature during radiofrequency energy delivery. Recently, it
was shown in an in vitro study that the lesion size becomes less
dependent from cooling effects and electrode-tissue contact if the
radiofrequency power is controlled
by the tissue temperature instead of the catheter tip temperature 26. An extractable needle that measures the
tissue temperature at a certain depth may allow tissue temperature
feedback
to the RF generator. Erdogan et al. suggested to measure an
electrochemical potential at the ablation electrode that detects the
release of free radicals during lesion generation 27.
In vitro they could demonstrate that this potential correlates to
lesion size. In addition, new technologies may be implemented in the
future including infrared sensors and ultrasound transducers to monitor
tissue temperatures during radiofrequency delivery.
Conclusion
Monitoring catheter tip temperature and closed loop control of power
output are useful to avoid excessive heating at the tissue surface,
that may result
in coagulum formation, and to accomplish effective heating at the
target
area. However, the catheter tip temperature is affected by cooling
effects
and electrode-tissue contact and thus poorly correlated to lesion size.
New means should be developed to assess tissue temperatures during
radiofrequency
delivery to achieve predictable and reproducible lesions.
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