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For discussions in this chapter, we concentrate only on the techniques and appli-
cations of thermal ablation.
15.3 Therapeutic Modalities for Thermal Ablation
15.3.1 RF Ablation
For tumors that are difficult to resect surgically, percutaneous tumor ablation, for
example, in situ radio-frequency ablation (RFA), as a minimally invasive method of
destroying tumors presents a preferred alternative. It has been suggested in recent
studies that tumor ablation for some tumors may be as effective as resectioning.
Commercial RFA systems have been available since 1996. RFA is effective in the
control of some heart arrhythmias as well as tumors of the kidney, lung, breast,
bone, and adrenal system with significantly increased survival rates. For larger and
infiltrative tumors (>3 cm in diameter), however, RFA may provide only palliative
and adjuvant treatment in debulking tumors, reducing pain, and enhancing the effi-
cacy of other modalities.
While passing through the biological tissue, the RF electrical currents (in the
range of 300- to 1,000-kHz frequencies) achieve a controlled heating of the zone.
Basically, two electrodes—an active electrode with a small surface area placed in the
target and another electrode with a larger dispersive electrode—are used to close the
electrical circuit. In a bipolar ablation system, two active electrodes are used. The
factors affecting the necrosed area include electrode length and diameter, tip tem-
perature, and treatment duration. The highest temperatures are achieved at the
proximal and distal ends of the electrodes.
For treatment of larger tumors, simultaneous application of RF energy to a
multiprobe array system with arrays of two to five electrodes, such as “hooked nee-
dle” or “umbrella” electrodes, are commercially available (RITA Medical Systems,
Mountain View, California; Radiotherapeutics, Sunnyvale, California). Internally
cooled RF electrodes (through which chilled perfusate is circulated), help in prevent-
ing tissue boiling and cavitation at the needle tip [7, 8].
The generation of high-frequency ac causes ionic agitation that is converted to
heat and induces cellular death due to coagulation necrosis. Internally cooled RF
ablation electrodes use dual-channel probes with an electrode at the tip, and cooling
fluids such as saline are continuously circulated through one channel and removed
through another. Various electrode geometries and a large variation in the electric
power in treatment protocols are reported in the literature. Furthermore, the values
of electrical and thermal characteristics for whole tissues/organs are reported in the
literature; however, the microscopic structures are often overlooked.
Basically, the electrothermal mechanism is explained by assuming the tissues to
be purely resistive, because the displacement currents are negligible. With J being the
current density (A/m
2
), and E the electric field intensity (V/m), the distributed heat
source q (Joule loss) in (15.1) is given by:
qJE=⋅
(15.4)
404 Modalities and Devices for Thermal Ablation
To illustrate the thermal-flow coupled problem in order to improve tempera-
ture prediction in the circulating blood, the mass and momentum equations are
employed. Fluid dynamics theory can be utilized to compute the velocity field for
the saline flowing out of the electrode in irrigated electrodes. The methods used in
creating large thermal lesions with cool-tip cluster electrodes include sequential
ablation, simultaneous activation of electrodes, and rapid switching of power
between electrodes. In the simultaneous method, electrical interference between
electrodes may lead to little heating at the center between the electrodes.
15.3.2 Microwave Ablation
Microwave energy has been used for several decades to ablate various types of tis-
sues, for instance, tumors of the uterus, breast, skeletal muscle, liver, prostate, and
cardiac tissue. Besides delivering similar results as that of RF energy for tissue abla-
tion, microwaves offer better and more consistent destruction of tumors. The elec-
tromagnetic fields can propagate through blood or desiccated or scar tissue and can
be deposited directly in the targets at depth. Several factors affect the penetration
depth, including the dielectric properties of the tissue, the frequency of the micro-
wave energy, and antenna design. The size of the ablation zone is found to be
logarithmically dependent on input power. Undesirable peripheral heating may
result by increasing the input power due to heat generated inside the catheters feed-
ing the antenna. Some commercial systems use water cooling of the needle shaft.
Unlike RF ablation, microwave energy heats the tissue by using electromagnetic
fields and not by conductive heating, so it does not require a direct tissue contact for
creating a lesion. It is faster than RF ablation and creates more uniform heating pat-
terns without the risks of charring. Dielectric radiative heating is produced by
enhanced vibration and friction of the molecules as the energy penetrates the tissue
and excites the water molecules. After coagulation necrosis, the tissue is left as a
nonconductive scar. RF and microwave ablation may be performed in various
modes including open, laparoscopic, or percutaneous procedures [13, 14].
In millimeterwave/microwave ablation, localized heating (up to ~150°C) is
achieved by electromagnetic energy delivered via catheters to a precise location in
the tissue of interest. Desired temperature profiles can be obtained by controlling
various parameters such as the delivered power, pulse duration, and frequency. The
major components of the apparatus are the microwave source (up to 10W at a 2- to
300-GHz frequency), a catheter/transmission line (to deliver power to the antenna
situated at the treatment site), and an antenna at the distal end of the catheter that
focuses the radiated beam. Microwave radiation follows the rule of squares,
wherein energy density falls off with the square of the distance from the antenna.
Depending on the frequency of the incident energy, the power deposition in the tar-
get tissue can be more than 2 cm around the antenna.
Several such systems are available. AFx Corporation’s (Fremont, California)
Flex 2 Lynx device (and the Flex 4 and Flex 10 devices) emits microwave radiation
at 2.54 GHz. Systems available for prostate application include Targis and
Prostatron (Urologix, Minneapolis, Minnesota), Tmx-2000 (TherMatrx, Evanston,
Illinois), and CoreTherm (ACMI, Framingham, Massachusetts). Both single- and
multiple-antenna systems (driven simultaneously) are used for localized ablation.
15.3 Therapeutic Modalities for Thermal Ablation 405
The VivaWave microwave ablation system (Vivant Medical, Mountain View, Cali-
fornia) for the treatment of kidney tumors permits the ablation of large lesions and
faster therapeutic times by the use of simultaneous irradiation by multiple probes.
The ability of a single microwave antenna to ablate a given volume is determined
mainly by the input power, antenna efficiency, and application time. Multiple-
antenna ablation may be a better alternative that large, single antennas due to better
power distribution covering a larger area.
15.3.3 Laser Thermal Ablation
Various modes have been demonstrated for using laser energy in desired target (can-
cer/tumor) destruction: laser thermal ablation (LTA), interstitial laser
thermotherapy (ILT), interstitial laser photocoagulation (ILP), and laser-induced
interstitial thermotherapy (LITT) [15–17]. The underlying basic principle for all of
these methods remains the same: the use of localized laser energy deposition, result-
ing in lethal thermal injury to the tissue.
Laser fibers are directly placed in contact with the target site to be treated, and
the delivered energy produces conductive and radiant heat leading to cell death.
Depending on the input power, the mechanism of cell death could be different. At
temperatures above 60°C, local tissue heating results in irreversible damage. Pho-
tons from low-intensity laser energy interact with molecular chromophores that are
inherent in mammalian cells such as hemoglobin, myoglobin, bilirubin, melanin,
and xanthophyll. Tissue composition, vascularity, tissue optical properties, the
wavelength of light, and the amount of energy applied are some of the factors that
affect the way laser light is scattered, reflected, and absorbed.
Exogenous administration of photosensitizing agents such as photofrin may
also accentuate photochemical effects such as in photodynamic therapy (PDT).
Higher and rapid energy exposure causes coagulative necrosis, which may reduce
optical penetration in the tissue. In LITT, to avoid carbonization and vaporization
of tissue at the distal end of the laser fiber, low-energy (3W to 20W) is applied over
elongated time (2 to 20 minutes). Temperatures at the commonly used bare-tip
quartz fibers (0.5 to 2.5 mm in diameter) may reach 300°C to 1,000°C, producing
rapid tissue coagulation and elliptical lesions. The carbonization severely impedes
light penetration and heat propagation into the tissue. The use of sapphire-tipped
laser fibers significantly reduces the charring effect at the fiber tip. Besides the input
energy in the beam, pulse width, pulse rate, and interpulse interval, the shape, size,
and material of the fiber tip has a direct influence on the shape and size of the
thermal lesions produced:
Intensity [W/cm
2
] = Peak power [W]/Focal spot area [cm
2
]
Peak power [W] = Pulse energy [J]/Plus duration [sec]
Larger coagulative tissue zones can be achieved by using multiple-fiber systems
with beam splitters and simultaneous energy deposition. A short pulse duration
maximizes the peak power and minimizes thermal conduction to the surrounding
tissue. Initial clinical usage of laser-induced tissue ablation included transuretheral
prostate application using a Nd:YAG laser fiber (with a wavelength of 1,064 nm)
under TRUS guidance. This, however, resulted in damage to the exposed urethral
406 Modalities and Devices for Thermal Ablation
tissue as well. In ILC, a needle-shaped laser fiber is placed in the prostate adenoma
through the urethra under cystoscopic guidance, and a diffusing diode laser beam
(830-nm wavelength) is applied interstitially (Indigo laser system, J&J Companies,
Cincinnati, Ohio).
Laser technology and computer control software have evolved significantly in
the past decade. For photorefractive keratotomy (PRK) of the eye lens, laser sources
(such as argon-fluoride excimer or solid-state lasers) are used. PRK is performed
with an excimer laser, which uses a cool ultraviolet light beam to precisely ablate
very tiny bits of tissue (at the submicron levels) from the surface of the cornea to
reshape it for better focusing of light into the eye. Endovenous laser ablation (ELA)
is a new technique for treating saphenous vein reflux by eliminating the reflux.
15.3.4 Focused Ultrasound-Based Ablation
Ultrasound waves are mechanical waves that can propagate through biological tis-
sue and can be brought to a tight focus. The physical mechanisms responsible for
the therapeutic effects of ultrasound are both thermal and mechanical stress. HIFU
is becoming an increasingly attractive modality for tissue ablation because of its
unique ability to noninvasively target deep-seated tissue volumes without affecting
intervening and surrounding cells [18–25]. However, various factors require atten-
tion in order to attain efficacious results from insonification of biological tissues.
First of all, the anatomy and functionality of the tissue under consideration should
be properly studied in order to better understand the location of the potential tar-
gets and to assess the severity of the risks that they may create. The specific location
of the tissue/organ in the human body also affects the design of applicator systems.
Based on the dimensions, location, and nature of the abnormality, the size and
shape of the applicators(and thusthe beam shape)can beevaluated by modeling the
ultrasonic field in front of the transducers. In certain cases, beam positioning in safe
regions can be crucial to the inherent functioning of the biological system, for
example, in brain tissue.
When ultrasound energy propagates through a biological medium, a part of it is
absorbed and converted into heat. This heat energy is responsible for various revers-
ible or irreversible biological changes depending on the irradiation conditions. For a
plane wave of intensity I
0
, traveling in a medium with amplitude absorption coeffi-
cient, α, the rate at which heat is generated, Q, is given by:
QI= 2
0
α
(15.5)
If there is no heat loss from the region, the rate of rise of temperature can be cal-
culated by the following relation:
dT
dt
I
c
m
=
2α
ρ
(15.6)
where ρ is the density of the medium and c
m
is its specific heat per unit mass. This
presents a very simplified relation that is applicable to situations where absolutely
no thermal conduction occurs and remains valid only some time after the applica
-
tion of a sonic pulse. In practical situations, thermal diffusion occurs at a rate pro
-
15.3 Therapeutic Modalities for Thermal Ablation 407
portional to the temperature gradient and heat transfer to adjoining areas becomes
significant very quickly.
Focal ultrasound surgery is a noninvasive means of surgery capable of selective
and precisely localized destruction in desired regions. A complete ablation of the
affected regions by inducing temperatures on the order of 60°C to 80°C is achieved;
the damaged region is called a lesion. This technique differs from hyperthermia
mainly on the basis of dosage levels (in particular, field intensity and exposure
times). The power levels involved here are on the order of 1,000 to 10,000 W/cm
2
.
Temperatures thus produced bring rapid, complete, irreversible damage to the
exposed tissue. The exposure times are very small (on the order of 1 to 10 seconds)
as compared to hyperthermia (10 to 40 minutes) [1–6]. The duration of exposure
also has an effect on the time of tissue repair after the treatment. Longer exposure
times result in higher damage, unpredictable lesion shapes, and delayed healing. The
problems posed by heat diffusion due to blood perfusion are, therefore, less
significant.
In the noninvasive approach, the acoustic energy is applied completely exter-
nally withoutany access incisions. The common types of applicators in this category
are spherically focused transducers (focused bowls). Electronic phased arrays have
been developed for various applications. Either type can be used in various modes:
extracorporeal, intracavitary, or directly placed in situ, using minimally invasive
techniques. Ultrasound frequencies on the order of 1 to 4 MHz and focal transduc-
ers with apertures up to 10 cm in diameter have been reported in the literature. The
design of focused ultrasound equipment and the planning of treatment require the
ability to model the system to predict in vivo outcomes. The acoustic field of a spher-
ical radiator at a single frequency has already been calculated by many researchers,
for example, by O’Niel [26]. A two-dimensional field model based on Huygens’
principle and the Fresnel-Kirchhoff diffraction theory has been developed for a large
aperture, axially symmetric, focused source of ultrasound. Field variations in the
radial as well as axial directions can be studied with this model [27, 28]. The acous-
tic intensity I at any arbitrary point P in the field at a distance r from the source and
time t is given by the following relation:
()
(
)
IA
r
e
itkr
m
=
−
∑
ρ
λ
π
ω
cm
2
2
2
(15.7)
where λ is the wavelength of acoustic waves, ω is the angular frequency, k is the
wave number, ρ is the density, c is the velocity of acoustic waves in the medium, m
represents the number of probes, and A is an arbitrary constant.
Fixed-focus, large-aperture, bowl-shaped transducers incur several limitations
during ablation procedures such as reduced flexibility of beam parameters, ineffec-
tive coupling for small targets, large treatment durations, and the presence of
off-focal hot spots in the intervening tissue while scanning the beam to cover the tar-
get region. For minimizing unwanted damage to the normal tissue due to repeated
overlapping of the beam energy while scanning a target region in 3-D and improved
flexibility of the beam parameters, the use of multiple transducers has been pro-
posed [27–30] by dividing the required power between two or more probes, such
that the multiple beams superimpose in the same focal region. The spatial configura
-
408 Modalities and Devices for Thermal Ablation
tions of these probes with respect to each other and with respect to the target tissue
were studied by developing 2-D and 3-D numerical models and verified in in vitro
and ex vivo tests.
The 3-D model, by analogy with optical field theory, uses Fresnel-Kirchoff dif-
fraction theory to evaluate the pressure amplitude at any location in front of the
transducer by taking a double integral over the entire surface of the source as
follows:
(
)
p
ick
UR
e
r
ddR
itkr
a
=
⎛
⎝
⎜
⎜
⎞
⎠
⎟
⎟
−
∫∫
ρ
π
φ
ω
π
2
0
0
2
0
(15.8)
where U
0
is the peak amplitude at the transducer face, a is the radius of the trans-
ducer, R is the radial location on the transducer face, φ is the angular location on the
transducer face, and other symbols have their usual meanings as described earlier.
The computer simulations help in treatment planning for dosage levels and to
determine the optimum range of interprobe distances and angles of orientation
(defining size and shape and intensity in the superimposed foci). From the simula-
tion studies, it is possible to decide on the range of suitable spatial configurations
(location and orientation)of theprobes based on the desired size, shape, and level of
achievable intensities in specific configurations, as suited to a specific application.
Probe positioning mechanisms and scanning of the focal region in the desired target
is done using dedicated manipulator systems, called focused ultrasound surgery
robots (FUSBOT), a representative example of which is described later in this chap-
ter. Robotic/mechatronic assistance and imaging guidance yield higher accuracy,
precision, reliability, and repeatability when manipulating surgical instruments in
desired locations.
Phased-array transducers use the same basic acoustic principles as acoustic
lenses that focus a sound beam. Variable delays are applied across the transducer
aperture. The delays are electronically controlled and can be changed instanta-
neously to focus the beam in different regions. Phased arrays [25] can be used for
electronically steering the beam over the area of interest without mechanically mov-
ing the transducer, and thus provide more flexibility in the shape and size of the
resulting lesions. With phased arrays, various focal patterns can be planned. A
treatment routine can be considered such that the entire tumor region is ablated by
creating distant and nonadjacent ablations to minimize cooling time between
sonications. However, the inherent disadvantages include increased complexity of
scanning electronics (particularly for large arrays), higher cost of transducers and
scanners, formation of constructive interference zones that might produce
pseudofoci beyond the actual focal regions, and a large number of elements, all of
which lead to greater control complexity. With improvement in array technology,
both linear and 2-D arrays are now available that can provide high-quality HIFU
applicators with increased bandwidth and reduced cross-coupling between the
array elements.
There has been good growth of commercial devices and systems utilizing HIFU
technology in the past decade. A commercial HIFU device called the Sonablate
200TM (Focal Surgery Inc. of Indianapolis, Milpitas, California) was developed for
the treatment of benign prostatic hyperplasia (BPH) and prostate cancers. In the
15.3 Therapeutic Modalities for Thermal Ablation 409
Sonablate system, a 4-MHz transducer is used for both lesion induction and imaging
purposes and can target treatment depths of 2.5 to 4.0 cm.
Another commercial device called the Ablatherm (Technomed International,
Lyon, France) utilizes separate transducers: 7.5 MHz for imaging and 2.25 MHz for
therapy purposes. This is also targeted for BPH treatment. Another version of this
system also called Ablatherm (EDAP TMS S.A., France) is commercially available
for treatment of prostate cancers. Insightec Ltd., Israel, commercialized an
MRI-guided noninvasive HIFU surgery system called the ExAblate 2000; it has been
FDA approved for the treatment of uterine fibroids. It utilizes MRI to visualize treat-
ment planning and monitor the treatment outcomes in real time. A Model JC Haifu
focused ultrasound based tumor therapeutic system has been developed by the
Chonqing Haifu Technology Company, China, for clinical ablation of solid tumors.
China Medical Technologies has developed a HIFU system initiated by the PRC
Ministry of Health for treatment of solid tumor/cancers.
15.4 Focal Ultrasound Surgery Robots—FUSBOTs
According to the Robotic Institute of America, a robot is “a reprogrammable,
multifunctional manipulator designed to move materials, parts, tools, or other spe-
cialized devices through various programmed motions for the performance of a vari-
ety of tasks.” The essential components of a robotic system, akin to a human body,
include the main structure, called a mechanical manipulator; an end-effector to
carry, position, and orient working tool(s); internal and external sensors to acquire
information about link positions and the environment; a central processing system
for planning and controlling tool trajectories in a desired fashion within the robot’s
workspace; and power source(s) for supplying power to various parts.
The use of robotic principles in surgical assistance, particularly in the area of
MIS, provides several advantages. These include higher accuracy, precision, and
repeatability in positioning surgical tools and maneuvering controlled trajectories
[27, 29, 30–32]. Minimally invasive and noninvasive surgical techniques have
been—and remain—prime areas of research in the biomedical arena in the past two
decades due to their numerous advantages over conventional surgery methods.
Advanced manipulation and computational tools can be used in preplanning, regis-
tration, and navigation of surgical devices based on the image data so as to spare the
surrounding and intervening healthy tissue. Most of the surgical robotic systems
work in interactive semiautonomous or assistive modes under supervisory control of
the surgeon.
At the Biomechatronics Research Group’s Robotic Research Center, we have
devised a range of medical robotic prototypes for focal ultrasound surgery (FUS),
named FUSBOTs, such as for breast surgery (FUSBOT
BS
), urological organ and
stomach cancer surgeries using transabdominal and suprapubic routes (FUSBOT
US
),
and neurosurgery (FUSBOT
NS
) for image-guided FUS applications.
The robotic manipulator design and thus kinematics and dynamics of its
mechanical configuration are based on specific applications. Common features
include real-time image guidance and interactive and supervisory control by the sur
-
geon. Once the surgical protocol is decided in the preplanning phase, the robots
410 Modalities and Devices for Thermal Ablation
accurately position theHIFU transducer(s) at specified locations coincident with the
planned lesion position on a given 2-D image. During the procedure, a real-time
lesion positioning and tracking algorithm updates the formation of the lesions in the
desired target with a predictive temperature map processed from online imaging
data. The dimensions and range of motions of the robotic system correspond to
human anthropometry data. The potential of mechatronic/robotic assistance in the
operating room and system overview of one representative of these systems, the
FUSBOT
BS
, for noninvasive ablative procedures is described next.
Remote ablation of deep-seated abnormalities by various modalities such as the
use of HIFU can provide completely noninvasive procedures if the energy in the
beam is carefully targeted. For breast applications, a custom designed five
degrees-of-freedom (three for positioning, one for orientation of the end-effector,
and one for imaging) robot, FUSBOT
BS
, for guiding an end-effector through a pre-
determined and image-guided trajectory has been developed and tested at our center
(Figure 15.1).
The end-effector has a purpose-built jig for mounting the HIFU transducer(s)
and it operates in a degassed water tank. HIFU probes are positioned such that the
focal zone (of a single probe) or the joint focus (of multiple probes) overlaps within
the affected target area. Fragmentation of energy into multiple low-energy beams
helps to minimize hot spots in overlying structures [27, 28, 30, 33–35]. For a target
tumor area larger in size than the focal zone of the beam, the HIFU probe(s) need to
sweep the entire volume of the lesion in 3-D. The probe manipulation modules and
robotic work envelope encompass the human torso region and thus are capable of
reaching and treating cancers/tumors other than in the breast, such as through
acoustic windows using transabdominal and suprapubic routes incorporating a
sliding window opening at the top. The end-point compensated accuracy of this sys-
tem is tested to be within ±0.2 mm. Various laboratory trials in tissuein vitro and ex
vivo using the system validate its excellent precision and repeatability.
15.5 Concluding Remarks
The principles of operation for various modalities used in the thermal ablation of
remote tissue have been discussed in this chapter. A comparison of methodologies
and techniques was made where necessary. All of the ablative modalities presented
15.5 Concluding Remarks 411
Figure 15.1 The FUSBOT
BS
system for breast and abdominal route FUS applications.
here have been used in several clinical and preclinical applications; however, it was
not the intent here to present a historic review of the clinical results.
Furthermore, online imaging is having a profound impact on ensuring better
and safer surgical outcomes. It is an invaluable tool for planning and navigating a
complex trajectory of surgical tools into a patient’s anatomy in 3-D, through safer
routes, and for monitoring the progress of thesurgery. Temperature is a prime target
parameter in ablative procedures. Monitoring the ablative procedure may help in
assessing the tissue response to the thermal dose targeted in the region of interest.
MR images show the changes in tissue characteristics induced by temperature eleva-
tion during ablation procedures. The use of B- or M-mode diagnostic ultrasound
images is a common practice for localizing target lesions in various procedures. But
during HIFU ablation sequences, barely any changes can be monitored in such
images apart from echogenic regions formed by gaseous cavities/bubbles near cavi-
tation thresholds. Such a situation, however, should be avoided in carefully
controlled treatments.
A brief overview of novel image-guided surgical robotic systems, FUSBOTs,
dedicated to FUS applications of various parts/organs of the human body, in partic-
ular, for breast surgery, was also provided in this chapter. To our knowledge, this is
the first ever attempt to devise dedicated biomechatronic systems for FUS world-
wide (including the previous study of the author for brain tissue ablation; Ph.D. the-
sis, Imperial College, London, 1999). The range of benefits reaped in other medical
procedures by the use of robotic technology should be extended to noninvasive abla-
tive procedures that share extended problems of image guidance, targeting, and
control.
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