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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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104 J. Hui and J.-X. Cheng
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plaque enabled by a 2-kHz barium nitrite Raman laser. Sci Rep 4(6889). https://doi.org/10.1038/
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for co-registered intravascular imaging. J Biomed Opt 16(10):106001. https://doi.org/10.1117/1.
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jacc.2014.04.053
Chapter 5
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Contrast-Enhanced Dual-Frequency
Super-Harmonic Intravascular
Ultrasound (IVUS) Imaging
Jianguo Ma and Xiaoning Jiang
Background
Atherosclerotic cardiovascular disease is a leading cause of death worldwide, and
one which often manifests without warning (Naghavi 2003). According to the 2014
update of Heart Disease and Stroke Statistics by the American Heart Association
(Go et al. 2014), there are more than 2000 deaths every day in the USA on average,
which is 1 death every 40 s. For up to 75% of acute coronary syndromes, the underlying pathological mechanism is hypothesized to be atherosclerotic plaque rupture
(Naghavi 2003). Unfortunately, a high percentage of vulnerable plaques are also
angiographically occult, and these are responsible for a high proportion of ensuing
cardiac events resulting in either fatalities or requiring further interventional treatment (Glaser 2005; Goertz et al. 2007). For this reason, detection and characterization
of plaques which are rupture prone is one of the most active areas of research in cardiology and biomedical imaging (Constantinides 1990). The vasa vasorum is a network
of microvessels which supports larger vessels such as the aorta, and increased density
of the vasa vasorum has been associated with a plaque advancing from a stable state to
a rupture-prone state (Naghavi 2010; Moulton et al. 2003). Additionally, intraplaque
hemorrhage occurring from thin-walled, immature microvessels has been present in
plaques in many cases of sudden coronary death (Virmani 2005). Evidence suggests
that vasa vasorum proliferation and associated angiogenesis and inflammation is
J. Ma (
School of Instrumentation and Optoelectronic Engineering, Beihang University, Beijing 100191,
China
e-mail: majianguo@buaa.edu.cn
Beijing Advanced Innovation Center for Big Data-Based Precision Medicine, Beihang University,
Beijing 100191, China
X. Jiang
Department of Mechanical and Aerospace Engineering, North C arolina State University, Raleigh,
NC 27695, USA
© Springer Nature Singapore Pte Ltd. 2020
Q. Zhou and Z. Chen (eds.), Multimodality Imaging,
https://doi.org/10.1007/978-981- 10-6307-7_5
B
)
105
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associated with plaque instability and rupture (Virmani 2005; Kolodgie et al. 2003;
Milei et al. 1998; Moreno and Fuster 2004). As our ability to predict the instability
of atherosclerotic lesions remains a substantial challenge, there is an unmet need for
new imaging methods to identify, detect, and differentiate these pathologies (Jaffer
et al. 2006).
The new technology of ultrasound molecular imaging utilizes contrast agents
displaying targeting ligands to identify areas of inflammation and angiogenesis
associated with disease progression (targets that cannot be identified by B-mode
ultrasound) (Lindner 2004; Choudhury et al. 2004; Gessner and Dayton 2010). Prior
data suggest that ultrasound molecular imaging will provide a unique opportunity for
plaque biomarker evaluation (such as inflammatory or angiogenic markers) and for
identification of vulnerable plaques (ten Kate et al. 2010). Additionally, a new highfrequency contrast imaging technique, acoustic angiography (Gessner et al. 2013c),
takes advantage of exciting microbubbles near resonance and detecting their highfrequency, broadband harmonics with sufficient bandwidth separation to achieve
both high resolution and high contrast-to-noise ratio (CNR). Data have shown that
acoustic angiography enables detailed visualization and analysis of microvascular
structure (Gessner et al. 2012, 2013c), and will likely be applicable to vasa vasorum
imaging. Thus, we hypothesize that there is a role for contrast-enhanced ultrasound
imaging in the assessment of atherosclerosis at high-order harmonic frequencies.
Feinstein has illustrated the potential of contrast-enhanced transcutaneous ultrasound imaging on the carotid artery (Feinstein 2006), but the potential of transcutaneous ultrasound has limitations with resolution and motion artifacts (Staub et al.
2010), especially if the target is the deeper coronary arteries. This may present an
opportunity for intravascular ultrasound (IVUS) (Slager et al. 2000), which has been
widely utilized for the characterization of coronary vessel walls (Tobis et al. 1991),
morphology of plaques (Jang et al. 2002), and so on. However, conventional IVUS
transducers are not optimized for contrast imaging (Nissen et al. 1991), and, therefore, are ineffective for vasa vasorum imaging. This absence of technology may be
due to the fact that nonlinear detection strategies for contrast imaging are most effective near the resonant frequency of microbubble contrast agents, which is typically
between 1 and 10 MHz (Kasprzak et al. 1999). Thus, conventional contrast imaging
strategies are not very effective with high-frequency ultrasound (35–50 MHz) that
is typically used with IVUS. To overcome this challenge, Goertz and collaborators
have been evaluating both subharmonic and harmonic contrast IVUS imaging, with
the goal of vasa vasorum imaging (Goertz et al. 2006, 2007). Their research showed a
contrast-to-tissue ratio (CTR) of 28 dB in subharmonic imaging with a fundamental
frequency of 30 MHz (Goertz et al. 2007) and 25 dB in second-harmonic imaging
with a fundamental frequency of 20 MHz (Goertz et al. 2006).
We hypothesize that resolution and contrast-to-tissue ratios can be further
improved over subharmonic or harmonic contrast IVUS imaging by excitation of
microbubbles near resonance and detecting their backscatter at a bandwidth substantially higher than that of the transmission, previously called “super-harmonic,”
“ultra-broadband,” or “transient” imaging. In prior work, de Jong et al. (2002),
Bouakaz et al. (2003), and Kruse and Ferrara (2005) demonstrated that substantial
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improvements in CTR could be achieved by detecting the high-frequency energy
produced by microbubbles excited at lower frequencies. Bouakaz et al. utilized a
dual-frequency transducer to illustrate that the CTR of the fourth and fifth harmonic
could be 15 and 7 dB higher than that of the second harmonic, respectively (Bouakaz
et al. 2003). Kruse et al. also utilized a dual-frequency transducer arrangement,
and demonstrated that substantial scattered energy from microbubbles excited near
2 MHz could be detected at a frequency as high as 45 MHz. More recently, Gessner
et al. (2010, 2012, 2013c) utilized mechanically scanned dual-frequency transducers
(transmit at 2 or 4 MHz, and receive at 30 MHz) on the VisualSonics Vevo770 to perform high-frequency 3-D contrast imaging of in vivo microvasculature and achieved
resolution on the order of 100 microns with a CTR high enough so that microvessels
could be readily segmented from the images and their morphology analyzed.
Despite the promising CTR and vessel imaging capability of this imaging
approach, there is a substantial challenge for “ultra-broadband” contrast-enhanced
intravascular ultrasound (CE-IVUS), which is likely why it is yet relatively undeveloped. The primary limitation is the large frequency span, which is outside of
the current bandwidth of commercially available single-frequency transducers. Such
difficulty could be surmounted by using multiple confocal transducers as described
by Gessner et al. (2010), however, such transcutaneous exposure method is almost
impossible to be used for coronary vasa vasorum imaging due to penetration depth
limitation and existence of ribs at the chest. Furthermore, decreasing the frequency to
<20 MHz is not a solution for a larger penetration depth because the high resolution
is needed to image the vasa vasorum with an average diameter of 161 µmforthefirst
order and 68 µm for the second order (Kwon et al. 1998). Intravascular ultrasound
(IVUS) imaging may be a solution to access the vasa vasorum deep inside the body,
and has been widely utilized for the characterization of coronary vessel walls (Tobis
et al. 1991), the morphology of plaques (Jang et al. 2002), and so on. For the CEIVUS super-harmonic vasa vasorum imaging, a specifically designed dual-frequency
IVUS transducer is essential.
In this chapter, small aperture, dual frequency, and IVUS transducers are introduced (Ma et al. 2013a, b), which provide sufficient bandwidth separation for high
CTR, high-resolution CE-IVUS imaging. Design, fabrication, characterization, and
super-harmonic imaging capabilities are demonstrated. The imaging performance
was compared among transducers with different materials and excitation frequencies. High-resolution (70 µm), high CTR (23 dB) images of microbubbles could be
generated by the dual-frequency transducers, indicating the potential capability of
imaging the second-order vasa vasorum.
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Dual-Frequency IVUS Transducer
Dual-Frequency Transducer Structure Design
There are several alternative s tructures of the dual-frequency transducer design
(Fig. 5.1a–d). One structure is to put the high-frequency element and the lowfrequency element side by side, which is named as interleaved structure (Fig. 5.1a)
(Van Neer et al. 2010; Martin et al. 2014). Advantage of this structure is the low
cross talk between the high-frequency and the low-frequency elements. However,
this structure has misaligned beam profile between the two frequencies and the
transmitting pressure was too low in the sensitive region of the high-frequency beam
(x =−0.2, y = 0inFig.5.2a). If the elements are stacked in layers (Fig. 5.1b),
then the transmitting and receiving beam would have the best overlap because of the
spatial alignment. However, if the two active layers were directly bonded together,
then the high-frequency receiving signal would continue propagation after being
detected by the receiving element, the reflection of which at the back side of the
transducer would cause aliasing echoes that show up after the real signal. Such aliasing echoes could either cause fake imaging signal (if each echo is considered as a
signal) or reduce the resolution (if the series of echoes are considered as one signal). An acoustic filter (Fig. 5.1c) with a quarter wavelength for high frequency was
placed between the two piezoelectric layers to enhance the low-frequency transmitting wave penetration, while reflecting the high-frequency receiving wave to suppress
the aliasing echo (Azuma et al. 2010). Detailed analysis of the acoustic filter will be
described in the next subsection. Furthermore, it is impossible to make the electrical
impedance match for both elements if the two elements are same in the aperture.
The low-frequency element is thicker than the high-frequency element, so that the
capacitance is lower. Consequently, the impedance ( Z =
element is significantly higher than the high-frequency element due to low ω and
low C . Furthermore, the natural focus points (near-field and far-field transition) of
the two beams are also significantly different if the apertures of the two elements are
the same. Based on these concerns, the structure of the transducer was designed as
stacked layers with different apertures between the two active elements (Fig. 5.1d).
With the above consideration, the dual-frequency transducer was designed as a
dual layer transduction structure, composed of a low-frequency transmitting layer
and a high-frequency receiving layer (Fig. 5.1e) (Ma et al. 2013a). The transmission
layer was placed behind the receiving layer (with respect to a forward traveling sound
wave) since low-frequency transmitted acoustic waves could propagate through the
smaller, high-frequency element. The placement of the receiving layer counted in
the matching network of the low-frequency transmission element, which showed
positive effect that amplifies the low-frequency propagation wave. The selection of
6.5 MHz as the transmitting layer’s center frequency was because it was both close to
the contrast agents’ resonant frequency and since the piezoelectric material for this
element was readily available. The high-frequency receiving layer was positioned
in the front of the transducer to achieve an optimal receiving performance of the
1
) of the low-frequency
jωC
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Fig. 5.1 Structural design of dual frequency transducer design alternatives. a Interleaved structure.
b Stack layer structure. c Stack layer structure with acoustic filter sandwiched between the high–low-
frequency piezoelectric layers. d Stack layer structure with acoustic filter and different apertures
of the high–low-frequency elements. e Cross-sectional view of the final dual-frequency ultrasonic
transducer design
Fig. 5.2 Beam profile of a the low-frequency transmitter in the interleaved structure transducer,
b the low-frequency transmitter in the stack layer transducer, and c the high-frequency receiver
high-frequency element. Transducer dimensions were optimized using a KLM model
(Krimholtz et al. 1970) to validate the thickness of layers, length, and width of each
component for ideal thickness mode excitation. The aperture of the high-frequency
(0.5 mm × 0.6 mm) receiving layer was designed to be similar to that of commercial
IVUS transducers (Garcia-Garcia et al. 2010), and thus significantly smaller than that
of the transmission layer (3 × 0.6 mm
frequency component was designed to obtain reasonably low electrical impedance
2
). The relatively large aperture of the low-
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at low frequencies for higher acoustic pressure transmission. The modeling of the
30 MHz element considered the existence of the 6.5 MHz element at Side B (back
side) as backing according to the structure of the design in Fig. 5.1e (Ma et al. 2014b).
Acoustic Filter for Stack Layer Transducers
Function of Acoustic Filter
If the dual-frequency transducer is designed with stacked layers as illustrated in
Fig. 5.1b, then one major concern would be the acoustic interference between the
two elements. Such interference might cause advantages or disadvantages. Unfortunately, the direct bonding of the two active layers leads to interferences between the
two layers which are disadvantages for the dual-frequency intravascular ultrasound
transducer.
Isolation of High-Frequency Ultrasound Wave
For high-frequency ultrasound, the direct bonding of piezoelectric layers causes
an additional receiving signal, which is named as aliasing signal (Fig. 5.3). After
the receiving signal is detected by the high-frequency element, the acoustic wave
further propagates from the high-frequency piezoelectric layer to the low-frequency
piezoelectric layer (red solid arrow in Fig. 5.3a). After reflected from the back side of
the low-frequency piezoelectric layer, the echo signal (green dash arrow in Fig. 5.3a)
excites the high-frequency receiving element again and generates extra pulses, i.e.,
aliasing signal (Fig. 5.3b). The aliasing signal is indistinguishable from the real
target, leading to imaging artifacts. Hence, the aliasing signal must be removed or
suppressed to a level which is significantly lower than the real target signal.
Fig. 5.3 Aliasing echo
generation. a Schematic
diagram of dual-frequency
ultrasonic transducer with
high- and low-frequency
piezoelectric layers directly
bonded; b signal detected by
the receiver
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The traditional method of removing the aliasing signal was to use high absorptive
material to attenuate the signal transmitted backward, which does not work well in
intravascular dual-frequency transducer application except that the two frequencies
are significantly different ( f
> 14 fL). First, there are barely any high absorptive
H
materials that can eliminate the signal within a couple of wavelengths according to
the space allowed in the intravascular ultrasound transducer. Second, high absorptive
material with large thickness attenuates the low-frequency transmission ultrasound
as well. High attenuation on the low-frequency transmission is not acceptable for
the non-focused small-aperture transducer, which needs to generate high enough
pressure for microbubble nonlinear excitations. Taking these concerns into account,
the aliasing echo should not be removed using the absorption method if the two
frequencies are not significantly (>10 times) different.
Reflection can be used to suppress the backward wave by changing the acoustic impedance matching conditions at the boundary. With two active l ayers bonded
together (Fig. 5.3a), the impedance of them are well matched (PZT, PMN-PT …) or
perfectly matched (for the same material of the two layers). Most ultrasound energy is
transmitted backward and hence, causing the aliasing echo (Fig. 5.3b). If an interme-
diate layer is inserted between the two active layers, then the boundary condition can
be changed, which could possibly reflect most of the energy directly without propagating to the low-frequency element. Impedance of the intermediate layer should be
different from the active layers as much as possible. Low impedance material, with
impedance much lower than piezoelectric materials, was chosen in this application
because such materials are much more easily available. With a thickness of a quarter
wavelength, the resultant impedance at the interface can be much lower than that of
piezoelectric materials, so that most energy is reflected and little energy propagates.
The suppression further happens when the tiny back-reflected wave interface with
the boundary again. As a result, there is very little aliasing echo from the back side of
the transducer. Such an intermediate layer changes the equivalent impedance at the
boundary from well matched to mismatched. As a result, this particular intermediate
layer is denominated as anti-matching layer.
Promotion of Low-Frequency Ultrasound Wave Propagation
For low-frequency ultrasonic wave, the direct bonding of the piezoelectric layers
usually causes low wave t ransmission efficiency. In super-harmonic imaging, the frequency of receiving ultrasonic waves is at least three times of the transmitting wave.
Therefore, the frequency-specific matching mechanism (like quarter-wavelength
matching layer) does not match both f requencies. Matching layer in front of the
high-frequency piezoelectric layer is usually designed to match the high-frequency
wave because of its short wavelength and it is more sensitive to the thickness. In this
case, the transmission efficiency is low without proper matching layer. Actually, both
layers in front of the low-frequency piezoelectric layer, the high-frequency piezoelectric layer and the high-frequency matching layer, are thin and almost negligible
for the low-frequency wave propagation. Consequently, the low-frequency ultrasonic
wave suffers from low transmission efficiency.
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A properly designed low impedance layer sandwiched between the high
impedance piezoelectric layers promotes the low-frequency ultrasonic wave transmission. This low impedance layer can be the same physical layer as the antimatching layer for high-frequency ultrasonic wave. It functions oppositely for
high–low ultrasonic waves, and is denominated as acoustic filter.
Microwave Analysis of Piezoelectric Transducers
Both microwave and mechanical wave (ultrasound) share the same wave propagation
properties although they are in different field. General wave equation u(x , t) =
j (ωt −kx)
Ae
and refraction properties are also the same for both types of waves. The two waves
could be studied together with shared theories and methods (Mason 1930).
Mass-spring-damper model in mechanical vibration is equivalent to the lumped
element resistor–inductor–capacitor (RLC) circuit in electrical vibration (Fig. 5.4).
Equivalences of the parameters are shown in Table 5.1. Lumped element analysis is
usually used for conceptual design of transducers, but is not widely used for detailed
parameters of the transducers.
+Be
j (ωt +kx)
applies for the propagating wave in both domains. Reflection
Fig. 5.4 a Lumped element equivalence of the dynamic behavior of the piezoelectric material,
b the electrical equivalent circuit, and c impedance response of the vibrating system
Table 5.1 Equivalence
matching table
Mechanical properties Electromagnetic properties
Force Vo l t a g e
Velocity Current
Mass Inductance
Spring constant Inversion of capacitance (1/C)
Damping factor Resistance
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Fig. 5.5 Schematic diagram
of transmission and
reflection in a a three-layer
structure and b its equivalent
microwave transmission line
Continuous variations of strain and stress are the principal features of piezoelectric
transducers, which is typically equivalent to a distributed circuit like transmission
line for electromechanical wave. Electromechanical wave equations are defined as
2
∇
−
2
∇
−
2
∂
1
c
1
c
E = 0, (5.1)
2
2
∂t
2
∂
B = 0, (5.2)
2
2
∂t
where E and B are the electric field and magnetic field, respectively, and c is the wave
speed in the medium defined as c = 1/
and dielectric constant in the medium.). Such wave equations are mathematically
identical to a mechanical wave
where u is the particle displacement from still state and c is the sound speed defined
as c =
√
E/ρ. Electromagnetic wave reflection coefficient is also identical as that
for mechanical wave. The cascade transmission line theories could be applied to the
ultrasound wave propagation in multiple layers.
A typical multi-layer wave propagation problem is a thin intermediate layer placed
between two infinitely large media (Fig. 5.5a), and the equivalent circuit is a short
section of transmission line connecting two infinitely long lines (Fig. 5.5b). If Z
and Z03are long, then Z1= Z01and Z3= Z03. Neglecting the loss, the input
impedance Z
is calculated as
in
∂
∂x
√
με (μ and ε are the magnetic permeability
2
u
2
∂2u
1
−
c
= 0, (5.3)
2
2
∂t
01
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