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Файл:Ординатура / Хирургия / Библиотека им академика М.И. Перельмана / Книга_5835_Библиотеки_им_академика_М_И_Перельмана.pdf
where I is the intensity and x is the depth in tissue. For short pulses, the power depo-
sition density is the main factor elevating the tissue temperature.
14.4 Attenuation Measurement as a Function of Thermal Dose
One of the most important issues for HIFU surgery is the understanding and control
of the thermal exposure in tissue. When tissue temperature is maintained between
50ºC and 100ºC for a few seconds (less than 5 seconds), tissue necrosis occurs, and
thermal lesions are created. We now have a very good understanding of how ther-
mal lesions vary with the amount of applied acoustical power and the pulse duration
for some tissues (brain and liver). However, we do not yet have enough information
about the effect of heating on absorption (tissue variable), which also affects the
size, shape, and placement of the thermal lesions. Absorption is an acoustical prop-
erty of tissue with wide variation from tissue to tissue, and it represents the rate at
which energy in tissue is converted to heat. For short pulses (less than 5 seconds) it is
one of the primary factors contributing to the temperature elevation in tissue
because the effect of blood flow is minimized [29]. Attenuation includes absorption,
scattering, and reflection, but when minimizing scattering and reflection, attenua-
tion will reflect losses due to absorption for the most part. In this section, the results
of attenuation are presented,which are very close to the results of absorption[30].
A number of studies have been conducted on the effect of temperature on tissue
attenuation or absorption. Dunn [31] studied the effect of absorption with tempera-
ture (up to 30ºC) in the spinal cord of mice, where an increase of absorption with
temperature was reported. Dunn and Brady [32] studied the absorption as a func-
tion of temperature in mammalian central nervous tissue at different frequencies,
and it was shown that the absorption decreased for frequencies lower than 1 MHz
and increased for frequencies higher than 1 MHz. Gammel et al. [33] studied the
effect of attenuation with temperature for excised porcine liver, kidney, backfat, and
spleen and for human liver. The frequency was varied from 1.5 to 10 MHz and three
different temperatures were studied (4º, 20º, 37ºC). It was found that basically,
above 5 MHz, the attenuation decreased with temperature for all tissues including
the human liver.
Pressure absorption was studied by Fry et al. [34]on rat liver, where for the tem-
peratures studied (30ºC, 37ºC, 41ºC), there were no statistically significant differ-
ences in absorption in this temperature range. However, for tissue ablation,
attenuation at higher temperatures (50ºC to 100ºC) is required. Robinson and Lele
[35] studied the changes in attenuation (from 30°C to90°C) in cat brain and showed
that the attenuation stays essentially constant up to 50ºC, and then increases rap-
idly. Bamber and Hill [36] studied the effect of temperature on attenuation in the
range of 5ºC to 65ºC on excised liver tissues. Many trends were observed, for exam-
ple the attenuation of bovine liver decreased and then increased until it reached
65°C, where it stayed constant. In human liver the attenuation dropped with tem-
perature, and for frequencies lower than 2 MHz, the change was insignificant.
Basically in the past years, some studies examined attenuation or absorptionas a
function of temperature (upto 37°C) and some at higher temperatures. However, no
study has examined both attenuation and absorption (at least for one tissue) as a
374 MR-Guided HIFU for Kidney and Liver Applications
function of temperature (with maximum temperature reached close to temperatures
seen in HIFU surgery) or as a function of thermal dose. The paper by Damianou et
al. [30] presents a complete study of both attenuation and absorption as a function
of thermal dose. This section reports on the effect of a thermal dose on the
attenuation of muscle, kidney, and fat.
14.4.1 Results of Attenuation Versus Thermal Dose
Figure 14.6 shows the attenuation of porcine kidney (cortex and medulla), muscle,
and fat as a function of thermal dose. Note that the attenuation increases rapidly
during necrosis and eventually stabilizes after total necrosis has been produced.
The trend of attenuation with dose shows a trend similar to that seen in the
results published by Damianou et al. [30]. The normal kidney tissue included both
medulla and renal cortex, for which we found no significant differences in terms of
ultrasonic attenuation. The research on muscle and fat is justified, because these
two tissues usually surround the kidney. The trend of attenuation with dose in all
tissues is similar. There is a rapid increase of attenuation with thermal dose and,
when necrosis occurred, attenuation stabilizes. At that point, the attenuation is
roughly double the attenuation at room temperature.
14.4 Attenuation Measurement as a Function of Thermal Dose 375
14
12
10
8
6
4
2
0
−3 −2 −1012345 6
log (dose referenced at 43°C)
Attenuation (dB/cm)
14
12
10
8
6
4
2
0
−3 −2 −1012345 6
log (dose referenced at 43°C)
Attenuation (dB/cm)
14
12
10
8
6
4
2
0
−3 −2 −1012345 6
log (dose referenced at 43°C)
Attenuation (dB/cm)
14
12
10
8
6
4
2
0
−3 −2 −1012345 6
log (dose referenced at 43°C)
Attenuation (dB/cm)
16
Kidney cortex
Kidney medulla
Muscle
Fat
(a) (b)
(c) (d)
Figure 14.6 Attenuation versus thermal dose for (a) kidney cortex, (b) kidney medulla, (c) muscle, and (d)
fat.
14.4.2 Simulation Results Using Variable Attenuation
Table 14.1 shows the simulated length and width using the pure thermal model and
using the model that accounts for the variable attenuation. The transducer parame-
ters used are frequency = 4 MHz, radius of curvature = 10 cm, and transducer diam-
eter = 4 cm. The focal in situ spatial average intensity of 2,000 W/cm
2
was applied
for 5 seconds at a focal depth of 15 mm. The zero perfusion represents the case of in
vitro kidney, whereas the in vivo case is modeled by using the perfusion of 70
kg/m
3
-s. Having established a rough estimate of the lesion size using simulation,
experiments were carried out in kidney in vitro and then in vivo.
14.5 Kidney Ablation
For applications such as the treatment of benign prostate hyperplasia (BPH), it is
sufficient to destroy as much tissue as possible for the purpose of tissue debulking.
For applications in oncology, destruction of all viable tumor cells is required and,
therefore, protocols in this area must be very accurate and reliable. In this work,
methods suggested by Malcolm and ter Haar [24] were used to produce complete,
reliable, and consistent ablation of renal tissues (creation of a contiguous array of
touching lesions of thermal origin, avoiding boiling and cavitation, and use of cool-
ing to avoid merging of lesions in front of the focal point).
14.5.1 Results
Figure 14.7 shows the length of a lesion created in excised porcine kidney using the
in situ spatial average intensity of 2,000 W/cm
2
for 5 seconds. The lesion produced is
376 MR-Guided HIFU for Kidney and Liver Applications
Figure 14.7 Thermal lesion placed in the cortex demonstrating the excellent repeatability.
Table 14.1 Simulated Length and Width for the Pure Thermal Model
and the Model of Variable Attenuation
Quantity
Pure Thermal Model Variable Attenuation Model
Perfusion (kg/m
3
-s) Perfusion (kg/m
3
-s)
0 (in vitro) 70 (in vivo) 0 (in vitro) 70 (in vivo)
Length (mm) 17.2 16.7 20.6 19.7
Width (mm) 2.6 2.4 3.3 3.1
Note: The transducer parameters used are frequency = 4 MHz, radius of curvature = 10 cm,
and transducer diameter = 4 cm. The focal intensity of 2,000 W/cm
2
was applied for 5 sec-
onds at a focal depth of 15 mm.
placed entirely in the cortex of the kidney. The length is measured along the trans-
ducer central axis, whereas the width is measured perpendicularly to the transducer
central axis. The lesion length is about 20 mm and the width is about 3 mm. In a cer-
tain row of lesions, only one dimension of the lesion (length or width) was mea-
sured, because the sample was cut either along the length or the width.
Lesions were created all the way to the medulla, provided that there were no air
spaces in the medulla. Air spaces are created because of the absence of blood that
normally flows in this region. However, ablation of the cortex tissue is of primary
concern, since the renal carcinoma (most important renal cancer) grows inside the
cortex and then extends to the peritoneal fat.
Figure 14.8 shows the lesion width with 2,000 W/cm
2
applied for 5 seconds.
The lesion measured was between 2.8 and 3.4 mm. The lesions appeared to be
repeatable for a given acoustic exposure. Some differences that are observed are
attributed to the variation of focal depth arising from the kidney curvature.
Creation of thermal lesions in the cortex is very consistent and repeatable. For
the given transducer (f = 4 MHz, R = 10 cm, and d = 4 cm) and the given exposure
(in situ spatial average intensity = 2,000 W/cm
2
, pulse duration 5 seconds, and
15-mm focal depth),the lesion length is around 20 mm and the lesion width is about
3 mm.
In most of the in vitro experiments, the propagation inside the medulla was dif-
ficult due to the air spaces in the medulla. Normally in an in vivo kidney, there is
blood inside the medulla, making the medulla a possible target for ultrasound. Fig-
ure 14.9 shows lesions that propagated inside the medulla region. Thus, in the case
for which there are no bubbles inside the medulla, ablation is feasible.
14.5 Kidney Ablation 377
Figure 14.8 Lesions created in pig kidney in vitro.
Figure 14.9 Lesions propagating in the medulla of pig kidney in vitro.
Figure 14.10 shows a lesion created by moving the transducer in a patterned
movement (square grid of 8 × 8 with a 3-mm step) in both directions. The intended
target was the round shape overlapped to the image. The power used was 2,000
W/cm
2
for 5 seconds. A delay of 10 seconds was used between the pulses in order to
eliminate the near-field heating [37]. Based on the width obtained in Figure 14.8, a
3-mm step will produce overlapping lesions. Necrosis with a 3- to 5-mm margin
around the target is desired. For the in vitro case, a larger step can be used
because ultrasound penetration is excellent and there is no blood flow. Based on the
simulation study forthe invivo case,a 3-mm step must be used. Figure 14.11 shows
the corresponding depth of the large lesion. The lesion was extended up to the
medulla.
The threshold of cavitation was found by detecting the in situ spatial average
intensity, which produces a signal in the receiver indicative of cavitation. Thus, sin-
gle lesions were created for this purpose, by varying the in situ spatial average inten-
sity from 2,000 to 3,500 W/cm
2
. For the given transducer and for a pulse duration of
5 seconds, it was observed that cavitation took place for in situ spatial average inten-
sity levels greater than 3,200 W/cm
2
(error was ±100 W/cm
2
). Figure 14.12 shows
the lesion length obtained using a 5-second pulse and 4,000 W/cm
2
in situ spatial
average intensity in kidney tissue in vitro. Note that, at this acoustic level, a
cavitational effect occurred, which was verified by the cavitational receiver. Tissue
378 MR-Guided HIFU for Kidney and Liver Applications
Figure 14.10 Large lesion in pig kidney in vitro. The transducer was moved in a grid pattern (8 ×
8) (top view).
Figure 14.11 Large lesion in pig kidney in vitro. The transducer was moved in a grid pattern (8 ×
8) (side view).
examination revealed gross tissue disruption, verifying the occurrence of an
acoustic cavitation event.
With a 4-MHz transducer, 5-second pulses, and in situ spatial average intensity
levels below about 3,200 W/cm
2
, the lesions created are based solely on thermal
effects. Data from Hynynen [38] suggest that the threshold spatial peak intensity at
4 MHz is about 6,000 W/cm
2
(extrapolated because data up to only 2 MHz is pro-
vided) or 3,333 W/cm
2
(in situ spatial average intensity). Thus, with the in situ spa-
tial average intensity of 4,000 W/cm
2
used in some experiments, cavitation effects
were not avoided.
Figure 14.13 shows the length of an 8 × 8 matrix of lesions created using 2,000
W/cm
2
for 5 seconds in pig kidney in vivo. The transducer was moved with a step of
3 mm. The lesions were usually firm and white and usually no adjacent hemorrhagic
zone was created. Despite the thick fat layer, necrosis of the renal cortex was reli-
able. The lesion length in the in vivo case appeared to be smaller (about 18 mm com-
pared with 20 mm in the in vitro case). Single lesions produced in the in vivo kidney
revealed a lesion width of about 3 mm for 2,000 W/cm
2
at 5 seconds. Thus, the step
of 3 mm produced overlapped lesions covering a large target volume.
Lesions were smaller in the in vivo experiments compared with lesions created
in dead tissue for the same exposure. The reduced size of the in vivo exposures is
attributed to the blood flow and possible reduction of the energy due to attenuation
at various interfaces including the skin.
14.5 Kidney Ablation 379
Cavitation
Figure 14.12 Large lesion in pig kidney in vivo showing cavitational activity. The transducer was
moved in a grid pattern.
Figure 14.13 Large lesion in pig kidney in vivo (through fat layer) by moving the transducer in a
grid pattern.
14.6 Kidney Ablation Under MRI Guidance
The use of MRI to detect changes due to heat was motivated by several studies,
which showed that T
1
and T
2
relaxation constants were sensitive to temperature.
One example is the study by Le Bihan et al. [39], which showed that the relaxation
constants were sensitive to temperature. Other representative examples are the stud-
ies by Dickinson et al. [40] and by Matsumoto et al. [41] who demonstrated the tem-
perature sensitivity of T
1
and diffusion.
The use of MRI to guide thermal therapy was fully investigated by groups
involved in laser therapy. The most representative examples are the studies by Bleier
et al. [42], Matsumoto et al. [41], Mumtaz et al. [43], Kettenbach et al. [44], and
Law et al. [45].
The idea of using MRI to guide HIFU was first cited by Fry and Johnson [46].
However, the idea was never explored in the research setting at that time. The com-
bination of ultrasound and MRI was first cited by Jolesz and Jakab [47] who dem-
onstrated that an ultrasonic transducer can be used inside an MRI scanner. The
concept of using MRI to monitor the necrosis produced by HIFU was demonstrated
in the early 1990s by Hynynen et al. [48] in canine muscle. In the following years
similar studies were conducted by Hynynen et al. [19, 20], Cline et al. [49], and
Darkazanli et al. [50]. In these studies it was shown that the contrast between
necrotic tissue and normal tissue was excellent. This was a great enhancement for
HIFU systems because it allowed the therapeutic protocols to be accurately moni-
tored. Therefore, the interest in using MRI as a diagnostic modality to guide HIFU
has increased.
In the preceding studies the T
1
- and T
2
-weighted fast spin echo (FSE) pulse
sequences were evaluated extensively. For fast imaging, various gradient pulse
sequences were used. It was found that the dimensions of the lesions measured using
MRI agreed well with the dimensions measured during the gross examination.
Therefore, research efforts were directed toward investigating whether HIFU and
MRI can be an effective pair for therapy/diagnosis of tissues accessible by
ultrasound.
During the evolution of the HIFUtherapy withMRI guidance technology, it was
realized that in order to apply this system in the clinic, temperature maps needed to
be created. Several studies have been published in this direction [51–56].
Hynynen et al. [15] was the first to guide HIFU surgery in the kidney using MRI.
In their study pulse sequences were proposed that are sensitive enough to show tem-
perature elevation during sonication, thereby indicating the location of the beam
focus. Fast pulse sequences such as gradient recalled acquisition in the steady state
(GRASS) were used with imaging times of a few seconds. The temperature elevation
could be clearly located in the gradient echo images.
The effectiveness of MRI in the brain was studied by Hynynen et al. [57],
Morocz et al. [58], Chen et al. [59], and Vykhodtseva et al. [60]. In liver the concept
has been explored by Rowland et al. [61]. The case of fat was monitored by
Hynynen et al. [62]. In the prostate it was explored by Nakamura et al. [63], Smith
et al. [64], and Sokka and Hynynen [65]. Therefore, the ability of MRI to guide
HIFU has been investigated for almost all tissues accessible by ultrasound.
380 MR-Guided HIFU for Kidney and Liver Applications
This technology was also investigated for clinical applications. For example, in
the breast it was studied by Pomeroy et al. [5] and Hynynen et al. [66]. HIFU/MRI
technology has also proven very useful in measuring some basic quantities of ultra-
sound. Acoustical streaming was monitored by Starritt et al. [67], ultrasonic fields
by Walker et al. [68], and ultrasonic pressure by Sharf et al. [69].
Therefore, the conclusion is that, in just a decade of research, MRI has become
an integral part of every aspect of therapeutic ultrasound. Generally, the advantages
of using MR for guiding and monitoring HIFU have been demonstrated. These
advantages include the following:
1. MR provides good tissue necrosis contrast.
2. The temperature elevation in the focal zone during sonication can be
detected with temperature-sensitive techniques.
The MRI images provide feedback data that can be used to monitor dynami-
cally the shape of the HIFU lesion, and therefore HIFU technology can be applied in
the clinical setting with confidence.
In this section, the goal was to investigate the effectiveness of MRI to monitor
the therapeutic protocols of HIFU use in the kidney. In Section 14.5 the target was
to create consistent and reliable necrosis in the cortex of the kidney. In that section,
the protocol was applied without any diagnostic guidance. In this section, MR
imaging can be used to monitor the therapeutic protocols.
Several MRI pulse sequences were investigated. For high-quality imaging,
which can be used at the end of a therapeutic protocol or at some instances of the
protocol, FSE techniques (T
1
- and T
2
-weighted and proton density) were investi-
gated. For fast imaging, the T
1
-weighted fast spoiled gradient (FSPGR) pulse
sequence was used
14.6.1 HIFU Parameters
In the study in this section moving the transducer in a grid pattern created large
lesions. The spacing between successive transducer movements was chosen accord-
ingly so as to create full necrosis coverage of a specific target. The in situ spatial
average intensity used was 2,000 W/cm
2
for 5 seconds, which ensures lesioning
using thermal mechanisms. This intensity is below the threshold ofcavitation in kid-
ney for this particular frequency and pulse duration [70]. The delay between succes-
sive ultrasound firings was 10 seconds. Although the study by MacDannold et al.
[71] recommends a delay of around 50 to 60 seconds in order to avoid buildup in
front of the target, for this application use of a transducer geometry that used a
10-second delay eliminated most of the built-up heating.
14.6.2 MRI Parameters
The various MRI parameters used for the various pulse sequences are listed in Table
14.2.
14.6 Kidney Ablation Under MRI Guidance 381
14.6.3 Results
T
1
-weighted FSE was explored by using different TR (220, 300, 400, 500, and 700
ms). Figures 14.14(a–e) show MR images using T
1
-weighted FSE at different TR.
Moving the transducer using a 4 × 4 grid pattern, with a step of 3 mm, created this
large lesion volume. This image shows the lesion in a plane perpendicular to the
transducer beam. Figure 14.14(f) shows a photograph of the kidney after slicing. All
of the parameters used are shown in Table 14.2, row 1.
The contrast between the lesion and normal kidney tissue is excellent with
T
1
-weighted FSE. The signal intensity of the lesion is homogeneous using this
382 MR-Guided HIFU for Kidney and Liver Applications
(e) (f)
(c) (d)
(a) (b)
Figure 14.14 Coronal MR images of HIFU ablation using T
1
-weighted FSE pulse sequence. (a) TR
= 220 ms. (b) TR = 300 ms. (c) TR = 400 ms. (d) TR = 500 ms. (e) TR = 700 ms. (f) Photograph
after slicing.
Table 14.2 Parameters Used for the Various MRI Pulse Sequences
Series Name TR (ms) TE (ms) Slice
Thickness
(mm)
Matrix FOV
(cm)
NEX BW
(kHz)
ETL Flip
Angle
1 T
1
-weighted
FSE
220, 300,
400, 500,
700
9.2 3, 5 (gap
0.3 mm)
256 × 256
16 1 31.25 8 —
2 T
2
-weighted
FSE
2500 16,32,
48,64
5 (gap 0.3
mm)
256 × 256
16 1, 2,
3,4
31.25 8 —
3 T
1
-weighted
FSPGR
50,100,
150,300
3.7 2 (gap 0.3
mm)
256 × 256
16 1 62.50 — 75º
Note: TR = time of repetition, TE = time of echo, FOV = field of view, NEX = number of excitations, BW = bandwidth, ETL = echo
train length.
method. The best contrast is observed for TR between 200 and 400 ms. This was
proved by plotting the signal intensity versus TR. The difference in the signal of the
lesion and kidney cortex is significant in this range. Any air spaces created due to
cavitational activity are not visible, because air spaces and lesions have almost the
same signal. MRI could detect the lesion immediately after the ablation. In [72], the
lesions in the brain appeared in MRI in 15 to 60 minutes using T
2
-weighted FSE.
Next, T
2
-weighted FSE was explored by using four different TEs. Figures
14.15(a–d) show MR images using T
2
-weighted FSE with TE of 16, 32, 48, and 64
ms. These images shows the lesion in a plane parallel to the beam. This large lesion
volume was created by moving the transducer using a 4 × 4 grid pattern with step of
3 mm. Note that the lesion was created entirely in the cortex, leaving the medulla
unaffected. Figure 14.15(e) shows the photograph of the kidney after slicing. All the
parameters used are shown in Table 14.2, row 2.
A similar evaluation was done for the case of a T
2
-weighted FSE pulse sequence.
The best contrast can be achieved for TE between 16 and 32 ms. The T
2
-weighted
FSE was proven to be the best pulse sequence that can detect cavitational activity.
This advantage is attributed to the significant difference in signal intensity between
the water-filled air spaces and necrotic tissue. Water-filled air spaces appear
brighter than thermal lesions. Therefore, for therapeutic protocols created using the
cavitational mode, T
2
-weighted FSE may be the optimum pulse sequence to be used.
For a therapeutic protocolusing only thermal mechanisms, T
1
-weighted FSE may be
suitable, although T
2
-weighted FSE can be used as well.
Figure 14.16(a) shows the signal intensity plotted against TR for a single kid-
ney. Figure 14.16(b) shows the graph of signal intensity plotted against TE with the
data derived from the images of Figure 14.15. The region of interest (ROI) was cir-
14.6 Kidney Ablation Under MRI Guidance 383
(e)
(c) (d)
(a) (b)
Figure 14.15 Axial MR images of HIFU ablation using T
2
-weighted FSE pulse sequence. (a) TE =
16 ms. (b) TR = 32 ms. (c) TE = 48 ms. (d) TE = 64 ms. (e) Photograph after slicing.