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Файл:Ординатура / Хирургия / Библиотека им академика М.И. Перельмана / Книга_3835_Библиотеки_им_академика_М_И_Перельмана
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10 Therapeutic IVUS and Contrast Imaging 237
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Fig. 10.3 Representative histology images of porcine coronary arteries treated with rapamycin
microbubbles and IVUS. a Artery section treated with rapamycin microbubbles only. b Artery
section treated with rapamycin microbubbles and therapeutic intravascular ultrasound. “I” and “M”
denote vessel intima and media, respectively. Scale bars are 500 µm. Source Kilroy et al. (2015)
flow) direction. Clearly, for maximum ARF effect, the transducer should be as long
in the axial direction as possible consistent also with the desire to minimize the rigid
length near the device tip. A rigid tip creates challenges when negotiating tortuous
vessels. Using a similar ARF optimized IVUS transducer, enhanced delivery of a
model drug was demonstrated in ex vivo and in vivo arteries (Kilroy et al. 2014a).
The long duration, moderate pressure, and ARF pulses were used to affect model
drug delivery from the microbubbles accumulated on the vessel. It was observed
that drug delivery was restricted to within the −6 dB beamwidth (Fig. 10.4). This
demonstrates the potential for ARF to assist in providing a spatially controlled drug
delivery effect to a vessel wall. Using both axial and rotational translations, it follows that it is possible, in principle at least, to “paint” the vessel wall according to a
therapy plan individualized to any unique asymmetric lesion.
Microbubbles in IVUS Imaging
Shell-stabilized gas-filled microbubbles are widely used as the basis of a
contrast agent in ultrasound imaging. Comprehensive reviews of microbubble
design and clinical applications are available in the literature (Cosgrove and
Lassau 2010; Unnikrishnan and Klibanov 2012). As a blood flow tracer, in
their simplest usage, microbubbles provide a bright echo signal that delineates the
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Fig. 10.4 a Microscope image of DiI delivery along the vessel wall. Gray-scale bar measures the
pixel intensity in arbitrary units of fluorescence. The dotted yellow lines denote the −6-dB beam
width of the ultrasound transducer (1.3 mm). b Fluorescence intensity increase over background
in the artery plotted along the circumference in an ex vivo artery where delivery was performed in
phosphate-buffered saline solution. Plot axes were truncated to show that the fluorescence intensity
increase occurs along the entire vessel wall. A maximum fluorescence intensity increase of 10× was
measured in this artery with a mean of 5.3 ± 2.9 fold. This result was produced with a sonoporation
pulse with a peak negative pressure of 2 MPa and pulse repetition frequency of 1 kHz. The region
from 0 to 2 mm has a lower fluorescence intensity increase because the catheter was not centered
and may have blocked the flow of microbubbles to this region. Source Kilroy et al. (2014a)
extent of the vasculature and the blood filled chambers of the heart. In this context,
microbubbles have been used for assessing myocardial perfusion defects following
myocardial infarction (Wei et al. 1998). They have also been used to assess enhanced
perfusion and angiogenesis in cancer applications (Hohmann et al. 2003; Zhao et al.
2010). In additional to having a role in assessing anatomy (via enhanced of vessel
boundaries) and function (i.e., blood flow/perfusion), microbubbles are now finding
applications using ligand-based molecular targeting to vascular endothelial growth
factor receptor-2 (VEGFR-2) in cancer settings (Pochon et al. 2010). Microbubbles
have received regulatory approval in most advanced countries for a range of cardiovascular, abdominal imaging, and molecular imaging applications. In particular,
in the USA, microbubbles have been approved for better delineating the chamber
boundaries in the heart and, in April 2016, were FDA approved for the characterization of liver lesions (FDA 2016). Initial clinical trials have been completed on
VEGFR-2 targeted microbubbles in the USA and Europe and it seems probable that
10 Therapeutic IVUS and Contrast Imaging 239
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molecular targeted microbubbles will achieve regulatory approval for clinical usage
in coming years.
A somewhat similar challenge exists for contrast agents as for IVUS from a
clinical translation perspective: Both IVUS and contrast agents take a previously
non-invasive imaging modality and make it minimally invasive. This inevitably adds
cost, very moderate level of pain and a small risk (primarily of infection) associated
with the required needle-based veinous access. As in the case of IVUS versus transcutaneous ultrasound, the penalty involved in migrating into an invasive modality
is rewarded with expanded research and clinical information. Nevertheless, the slow
rate of adoption, at least in the USA, has been disappointing to those performing
research in the field. However, IVUS benefits via the fact that it is minimally invasive from the outset so the addition of microbubbles creates only marginal increases
to cost and risk.
The high echo signal obtained from microbubbles is a consequence of their very
low density and high compressibility relative to adjacent tissues and structures. However, because other targets in an ultrasound field can also give rise to a strong echo
signal (e.g., vessel/blood interface, etc.), it is highly desirable to exploit the unique
echo signal “signature” of a microbubble. The high compressibility of a microbubble gives rise to a nonlinear vibrational response to impinging compressional ultrasound wave energy. As a consequence, harmonics of the transmitted waveform are
present in the echo signal. Most often the first harmonic of the fundamental (transmitted) signal is used in signal isolation methods. It should be noted that in the
literature this harmonic signal is usually referred to as the “second harmonic” to
mean the first harmonic of the transmitted signal—i.e., centered at 2f
where f
o
is the fundamental, transmitted, ultrasound center frequency. However, higher harmonics and subharmonics are sometimes employed. Given the very high frequencies
typically used in IVUS, it can be observed that achieving transducer and system
bandwidth to encompass both transmitted and nonlinear harmonic signal presents an
obvious technical challenge. A number of signal isolation methods have been developed to isolate the nonlinear microbubble signal and thereby isolate it from adjacent
tissue signal. The most common approaches involve multiple pulse excitation along
a common acoustic beamline (Shen et al. 2005). These methods include “amplitude modulation” (AM) (Brock-Fisher et al. 1996), “pulse inversion” (PI) (Chapman
and Lazenby 1997), and “contrast pulse sequences” (CPS) (Phillips 2001; Phillips
and Gardner 2004). These methods all involve the use of the observation that serial
signals acquired of an identical, stationary, target behave as a linear system when
all components of the pulse-echo system are linear. More simply stated, in a linear
system, the echo signal is a scaled replica according to the amplitude/phase of t he
excitation (transmitted) signal. Thus, these acquired linear signals can be scaled and
added, or subtracted, in order that linear origin signals cancel. In the case of amplitude modulation, a scaled replica of an initial pulse is used (e.g., “1,” “½”). The two
received signals have a compensatory scaling applied to the, for example, smaller of
the two signals such that when subtract, perfect cancelation occurs (i.e., in this case
just described, the second pulse “½” is scaled by a factor of 2 resulting in a “1”). In
the case of pulse inversion, the two transmitted signals are inverted replicas of each
o
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other (i.e., “1,” “−1”). Upon summing the received signals, the linear origin signals
cancel. “Contrast Pulse Sequences” involves a combination of both AM and PI (e.g.,
½, −1, ½). In t his CPS example, the three received signals can be summed, without
scaling, to achieve cancelation of the linear origin signal. Notice that in all of these
three methods, any nonlinear origin signal will not perfectly cancel. It should also
be noticed that these nonlinear signal separation methods are sensitive, to varying
degrees, to both subharmonics and superharmonics—not just the second harmonic.
It is also possible to isolate nonlinear signals using conventional frequency filtering.
Additionally, cancelation techniques are frequently combined with frequency filtering (Shen et al. 2005). It may be observed that these multi-pulse methods, especially
the longer pulse methods (e.g., CPS), are susceptible to imperfect fundamental signal
cancelation whenever there is tissue motion between successive pulses (Shen et al.
2005). In the context of IVUS, this is a concern even in the presence of the very s hort
inter-pulse intervals implicit in very short pulse-echo transit durations.
IVUS Specific Applications of Microbubbles for Imaging
Microbubbles have long been recognized in IVUS imaging as making a contribution toward better delineating the boundary between the lumen and the wall of a
blood vessel (Vavuranakis et al. 2005; Masuda et al. 2001). The increased contrast
obtained when using microbubbles affords improved detection of in-stent neointima and reduces observer variability when assessing stenosis and neointima area
(Masuda et al. 2001). It is now established that plaque microvascularity and plaque
vulnerability are closely correlated (Naghavi and Falk 2010;Howardetal.2015).
Consequently, the use of microbubbles with IVUS has been used to assess the vasa
vasorum that develops in coronary atherosclerotic lesions (Vavuranakis et al. 2005,
2008; Goertz et al. 2007a; Carlier et al. 2005) [As an aside, although IVUS is nec-
essary for assessing the vasa vasorum in the coronary arteries, a transcutaneous
ultrasound approach for assessing the vasa vasorum in carotid arteries is feasible
(Song and Zhang 2015)]. Microbubbles, even when using a non-optimized IVUS
device, yield enough vascular contrast enhancement to enable assessment of vascular density (Carlier et al. 2005). It is possible to measure intra-plaque perfusion
density by contrasting the gray-scale image intensity acquired before, during, and
after microbubble infusion (Vavuranakis et al. 2008). Using this approach, it has been
observed that increased microbubble image signal between the intima-media boundary and the adventitia, following infusion, was associated with increased plaque
vascularity (Fig. 10.5). Thus, it i s believed that a combination of IVUS plaque mor-
phological assessment (e.g., identifying cap thickness, calcification, lipid core) and
vasa vasorum density imaging using microbubble-enhanced IVUS provides a powerful instrument and method for identification of plaques susceptible to rupture—i.e.,
“vulnerable plaques” (Carlier et al. 2005).
Efforts have been made to improve the sensitivity and specificity of IVUS instrumentation for microbubble detection. In this context, the very high frequencies com-
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Fig. 10.5 Depiction of qualitative representation of enhancement. Unprocessed images are displayed a before, b during, and c after injection of microbubbles. Corresponding processed images
are displayed in d–f. Enhancement is graded from minimal (blue) to maximal (red). Values are a
percentage of the maximum grey level intensity difference (255). Arrows indicate points of intense,
stable enhancement at the media–adventitia border. Diffuse points of enhancement are present
nearby. Source Vavuranakis et al. (2008)
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monly used in IVUS present significant technical challenges. Super harmonics occur
at very high frequencies (e.g., 40–80 MHz) and require very high signal bandwidth. It
needs to be emphasized that this infers very high signal bandwidth at every “link” in
the processing chain—transmitter electronics, transducer, transducer interconnecting
cable, receiver electronics, and digitization. The mismatch between commonly used
IVUS frequencies and the natural resonant frequencies of commonly used microbubbles (<5 MHz typically), has motivated the consideration of subharmonic detection.
These challenges have motivated development of dual-frequency IVUS transducers
and methods to suppress echoes originating from tissue and preserve the nonlinear
signal originating from microbubbles (Goertz et al. 2007b). IVUS imaging using subharmonic and second-harmonic signals have been evaluated in vessel phantoms and
have yielded up to 15 dB of microbubble-specific contrast enhancement (Goertz et al.
2006a, 2007b). These approaches have also been in tested a fully developed rabbit
aortic plaque (in vivo) (Goertz et al. 2006a, b). Harmonic imaging provided a successful means of imaging the structure of the vasa vasorum with an ~8 dB increase in
sensitivity relativeto standard, fundamental frequency-based, IVUS imaging (Goertz
et al. 2006a, b) (Fig. 10.6). These early successes provided evidence that specialized
microbubble imaging is feasible at IVUS imaging frequencies and motivated further
development of newer techniques that improve microbubble-specific detection in t he
vasculature. These new methods include radial modulation (Yu et al. 2014), superharmonic (Ma et al. 2014), ultraharmonic imaging (Maresca et al. 2014), and chirp-based
imaging (Shekhar and Doyley 2013; Shekhar et al. 2016). For example, Maresca et al.
recently demonstrated the use of conventional band-limited IVUS transducers with
chirp transmit sequences for ultraharmonic contrast-enhanced imaging of microvasculature not readily visualized by conventional IVUS imaging (Maresca et al. 2013,
2014) (Fig. 10.7). The data provided by contrast-enhanced IVUS can also be used
to extend the quality and accuracy of existing methods for characterizing the vessel
wall. For example, Huntzicker et al. (Huntzicker et al. 2016) have recently demonstrated that in model-based IVUS elastography, the additional information gained in
relation to the geometry of detected microvessels can be used to improve the accuracy
of assessments of stress distribution in the vicinity of the fibrous cap on a vulnerable
plaque.
Although there has been considerable success using existing commercial IVUS
transducers with conventional levels of signal bandwidth (<80% −6 dB pulse-echo),
there exists a high level of interest in creating transducers true multiple frequency
capability. Although 50% −6 dB bandwidth (i.e., approximating a two-cycle pulse)
is sufficient to achieving λ axial resolution—and lateral resolution is only able to
approach λ resolution in near ideal focusing conditions (i.e., f/number ~1, at focus),
higher bandwidth has immediate value if attempting to assess sub or super harmonic
signals. Additionally, the higher bandwidth will improve any signal analysis dependent on spectral qualities of the received signal. For example, “virtual histology,”
using radiofrequency analysis (Nair et al. 2002, 2007) has encountered significant
limitations in terms of reported accuracy (Thim et al. 2000) and these are probably
attributable at least in part to limited signal bandwidth. Higher transducer bandwidth can be achieved using high coupling coefficient single crystal materials, more
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Fig. 10.6 In vivo contrast imaging results in an atherosclerotic rabbit aorta. a Fundamental mode
before agent injection, where “C” is the catheter and “VC” is the vena cava. b Fundamental mode
10 seconds after injection where changes in adventitial enhancement are not evident, except for
in a region at 4 o’clock and within the vena cava. c Harmonic mode before injection shows the
tissue signals to be largely suppressed. d At 10 seconds after injection, the harmonic mode shows
significant adventitial enhancement, consistent with the detection of adventitial microvessels. Scale
of images is 12 mm across. The dynamic range of the fundamental and harmonic images are 40 and
25 dB, respectively. Source Goertz et al. (2006b)
complex matching layer schemes or cMUT technology. However, in the IVUS
field, the small dimensions (physical fabrication difficulties, electrical matching
challenges, cable loss effects, etc.) present more extreme practical challenges to
the transducer designer. It also worth reiterating that due to the importance of tissue/contrast signal arising immediately adjacent to the transducer, using an offset
transmitter/receiver approach is not a viable option. In one approach, multiple passive layers are used or provide 20 and 40 MHz operation from a single element
transducer (Vos et al. 2005). Stacked, high- and low-frequency element-based, transducer designs have been proposed. Generally, the two components of the stack will
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Fig. 10.7 Images acquired in the tissue-mimicking phantom using the chirp reversal sequence
and the pulse inversion sequence. a IVUS chirp image without contrast. b IVUS chirp image
with contrast. c IVUS chirp reversal contrast image. d 22 MHz conventional IVUS image without
contrast. e 22 MHz conventional IVUS image with contrast. f Ultraharmonic IVUS contrast image.
Intensity scales in contrast modes are displayed relative to their respective B-mode intensity. In both
the chirp reversal (c) and ultraharmonic (f) contrast modes, small model vessels (0.5 mm vessel
to left) can be identified above background. In the ultraharmonic example, (f), a 0.2 mm vessel to
bottom is readily visible. Source Maresca et al. (2013)
mechanically cross couple making for a challenge in obtaining a “clean” spectral performance at each of the two desired center frequencies. In one design, low-frequency
(35 MHz) and high-frequency (90–150 MHz) transducer elements were stacked to
provide for conventional resolution and imaging depth performance at 35 MHz and
very fine, but shallow, imaging performance at the higher frequency (Ma et al. 2015).
In a second stacked design, a low-frequency transducer element (6.5 MHz) was
placed under a high-frequency transducer element (30 MHz) (Fig. 10.8) (Ma et al.
2014). The lower-frequency transducer insonated microbubbles near their resonance
while emitted superharmonics were detected using the high-frequency element. This
“transmit low, receive high” (TLRH) technique permits much high-resolution imaging of microbubbles than is possible when limited to imaging near the microbubble
resonance frequency (Hu et al. 2010). Using this approach, an in vitro microbubble SNR of 12 dB has been demonstrated by Goertz et al. (2006a). More recently,
the stacked element design has been extended to use lateral-mode transmission to
yield yet higher versatility and performance (Wang et al. 2016). The transmit low,
receive high imaging approach has been extended recently by Lindsey et al. (2016)
to include adaptive (minimum variance) beamforming to achieve further refinements
to imaging resolution. In vitro, combining minimum variance and phase coherence
factor processing improved the spatial resolution by 41.7% while improvements of
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Fig. 10.8 a Design of the dual-frequency IVUS transducer. b Prototype transducer housed on the
tip of a 20-gauge hypodermic needle. Source Ma et al. (2014)
2.2 dB in contrast to tissue ratio and 37.2% in spatial resolution were achieved in vivo
(Lindsey et al. 2016).
Molecular Imaging in the IVUS Field
Ultrasound molecular imaging is based on the detection of microbubbles that have
adhered to a vessel wall resulting from a binding event between molecule specific
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ligand on the microbubble surface and a matched molecular receptor expressed on
the vessel surface. Single microbubble origin signal to noise does not present a major
challenge—certainly for the in vitro case (Klibanov et al. 2004). However, reliable
separation of signals arising from tissue, freely circulating (non-adhered) microbubbles and adhered microbubbles (“true” molecular signal) is challenging. The underlying approaches used in conventional (non-IVUS) applications have been largely
mapped over to use in IVUS. The same targeting and signal separation methods can
be used—except that the signal bandwidths create more challenges. The potential
of ultrasound molecular imaging to elucidate the molecular status of cardiovascular
(and other) disease and thus enable the earlier detection, or staging, of disease has
been recognized for many years now (Lindner 2004).
Early in vitro reports have proven that the signals from molecular targeted, and
adhered, microbubbles possess sufficient signal to noise when using both fundamental mode 40 MHz (Phillips et al. 2012) and when using nonlinear signal-based
sensing approaches (Goertz et al. 2007a). An early (1999) in vivo IVUS molecular
imaging experiment involved echogenic liposomes (ELIP) targeted to intercellular
adhesion molecule-1 (ICAM-1), a Yucatan mini-swine model with atherosclerotic
carotid arteries and 20 MHz IVUS (Demos et al. 1999). In 2000, Lanza et al. studied
perfluorocarbon nanoemulsions for detection of tissue factor expression in swine
carotid arteries. A more extensive study by Hamilton et al. (2004), also using ELIPs,
reported specific detection of ICAM-1, vascular cell adhesion molecule-1 (VCAM-
1), fibrin, fibrinogen, and tissue factor in swine carotid arteries subsequent to angioplasty balloon injury. IVUS imaging (20 MHz) indicated that all artery segments
producing increased IVUS signal intensity were matched with histology consistent with increased expression of the target molecule. More recently, Phillips et al.
(2012) presented in vitro 40 MHz IVUS imaging of VCAM-1 targeted microbubbles (Fig. 10.9). Together, these in vivo studies suggest IVUS-based identification of
the molecular-specific information may enable early characterization of endothelial
injury following angioplasty balloon injury and identify atheroma components that
may guide the eventual choice of therapeutic intervention.
Integrated IVUS and Microbubble Generation Devices
Usually, when IVUS and microbubbles are combined, the microbubbles are introduced via a separate intravenous injection or from a lumen and port on the catheter
device itself. Of course, using the device to dispense microbubbles directly increases
the local concentration of microbubbles allowing for greater potential for efficacy
with minimal systemic-related risk. Recently, it has been proposed that microbubbles
actually be formed in the catheter device itself (Dhanaliwala et al. 2013). Microbubbles generated using a flow focusing microfluidics device havebeen tested in contrastenhanced imaging (Hettiarachchi et al. 2007; Kaya et al. 2010; Dhanaliwala et al.
2013), molecular imaging (Talu et al. 2007; Seo et al. 2010) and therapeutic delivery
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