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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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3 The Integration of IVUS and OCT 63
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the aperture of the GRIN lens is larger than GRIN fiber probe, the theoretical lateral
resolution of a GRIN lens probe is usually better; secondly, a GRIN lens usually
performs with a higher-quality gradient refractive index profile than that of a GRIN
fiber and less spherical aberration t han a ball lens, which makes its lateral resolution
even better.
After separately fabricating the OCT sub-probe and the ultrasound transducer,
these two components can be fixed inside a stainless steel (Yin et al. 2010) or poly-
imide (Yin et al. 2011) tube. On the tube, a window needs to be cut to allow the light
beam and the sound wave to exit. The different arrangements of the transducer and
the OCT sub-probe have been proposed, to minimize the size of the integrated probe
while preserving its imaging performance. As they are all designed for intracoronary
applications, we review them later in this chapter.
Outer Sheath Design
An outer sheath can help avoid cross-contamination between the imaging probe and
the biological tissue. Furthermore, the sheath is designed to remain stationary when
the probe rotates for scanning. Such a design can prevent the rotating probe from
contacting and damaging the tissue.
The design of an IVUS-OCT sheath is similar to that of a clinically used IVOCT
catheter, except that special attention needs to be paid to the selection of its material.
The sheath material has to allow both light beams and sound waves to pass through
without high attenuation or reflection (Li and Chen 2016).
Since the sheath is a cylindrical tube, its curved surface makes the light beam
diverge faster in one axis than in the other axis. Consequently, astigmatism is generated to the light beam leaving the sheath. Various methods of astigmatism correction
can be used, such as purging (Katwal and Lopez 2015) and using the ellipse ball-lens
design(Tanetal.2012). Purging means injecting a high refractive index medium
(such as a saline solution or a contrast agent) into the lumen between the imaging
probe and the sheath. Since the media in both s ides of the curved interface have similar high refractive indices, the amount of undesired focusing is reduced. In a later
design (Tan et al. 2012), a ball lens with different radii in two axes was employed
so as to focus light to different extents and compensate astigmatism caused by the
curved surface of the sheath.
Key Measures
To characterize the imaging performance of the system we built, two key measures,
resolution and SNR, are commonly used. Here, we present equations that relate
the imaging performance, the system speed and the probe size, and briefly discuss their implications. Technical approaches to maintain high performance without
significantly increasing the system complexity or reducing the imaging speed are
summarized later in this section.
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Resolution
Resolution is defined as the minimal resolvable distance between two-point objects.
Transverse resolution For both ultrasound and OCT technologies, the transverse
resolution is determined by the beam spot exiting the imaging sensor.
In ultrasound, the theoretical lateral resolution is typically expressed as (Biomedical 2006;Ng2011):
V
=
R
L
· F#(1)
f
0
where V is the speed of ultrasound, f
focal length l
and the outer diameter D of the transducer aperture.
f,
is central frequency, F# is the quotient of the
0
In OCT, the theoretical lateral resolution is related to the numerical aperture (NA)
of the imaging optics:
λ
NA
0
(2)
where λ
= 0.37
R
L
is the center wavelength of the OCT light source.
0
In reality, focusing aberration usually occurs. It deteriorates the beam quality
and, thus, makes the actual transverse resolutions worse than the theoretical ones
calculated by Eqs. (1) and (2).
Axial resolution Unlike transverse resolution, axial resolution is independent of the
focusing ability of an imaging sensor.
In ultrasound, the axial resolution can be estimated by the following equation
(Biomedical 2006;Ng2011):
R
A
=
2 · fo· BW
V
(3)
where BW is the fractional frequency bandwidth.
For an OCT sub-system with a Gaussian spectrum light source, the axial resolution
is determined by the center frequency and the bandwidth of the light source λ:
2
λ
2In2
R
=
A
nsπ
λ
0
(4)
where n
is the refractive index of the sample.
s
For an OCT sub-system with a non-Gaussian spectrum light source and/or with
chromatic dispersion between the sample and reference arms, estimation of its axial
resolution by the above equation can be inaccurate. A measurement of the fullwidth-at-half-maximum (FWHM) of the axial point spread function is usually used
to obtain the actual axial resolution.
SNR Classically, the SNR of an imaging system is the ratio of the peak signal to the
root-mean-square of the noise level.
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In an ultrasound sub-system, the SNR is related to the pulse duration, which is
inversely proportional to the imaging speed, and the lateral beam width (Werkmeister
et al. 2013):
A(f
SNR =
16
ρvwxwyt
n(f
)
0
)
0
(5)
where w
speed in the medium, t is the pulse duration, A(f
at the center frequency, and n(f
and wyare the beam width in the x-and y-axis, respectively, v is sound
x
is the amplitude of the signal
)
0
is the noise at the center frequency.
)
0
In shot-noise limit FDOCT (i.e., the system used in the integrated IVUS-OCT
system) the SNR can be expressed by the following equation (Bouma et al. 2008):
ηP
hυ f
s
A
(6)
where f
SNR =
is the A-line speed, Psis the signal power of the sample arm, hυ is the
A
single-photon energy, and η is the detector sensitivity. As illustrated by this equation,
the power of the sample arm is usually positively related to the SNR. Accordingly,
an OCT sub-system usually uses a 90:10 or 80:20 coupler that transmits more lights
to the sample arm than to the reference arm, so as to improve the SNR.
Similarly, as can be seen from Eqs. (5) and (6), SNRs of both ultrasound and
OCT sub-systems will drop as the imaging speed goes up. Thus, many challenges
must be overcome so as to achieve an ultrafast high-SNR IVUS-OCT system, which
is introduced in the section ‘Technical Advances: high-speed imaging system’. In
addition, the aperture size of an imaging sensor influences its lateral resolution and
the SNR according to Eqs. (1), (2), and (5). A summary on approaches to overcome
the trade-off between the performance and the size of an integrated probe is provided
in the section ‘Technical Advances: miniaturized IVUS-OCT probe’.
Case Study: IVUS-OCT to Detect Vulnerable Plaques
Background
It was briefly mentioned in the introduction that the development of integrated IVUSOCT technology was driven by the need for identifying vulnerable plaques. A thincap fibroatheroma (TCFA), the most common type of vulnerable plaques, is prone
to rupture and trigger thrombus formation. Thrombus can block blood flow within
coronary arteries and subsequently results in acute coronary syndrome (ACS). However, the majority of plaques in arteries are stable and do not cause ACS (Farooq et al.
2009; Sanz and Fayad 2008; Nakano et al. 2012; Ferrante et al. 2010). Treating a
plaque without assessing its vulnerability may lead to overtreatment of stable plaques,
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which is not recommended by clinical guidelines, and could result in an increase of
complications (Waxman et al. 2006;Falketal.2013; Montalescot et al. 2013). On
the contrary, identifying vulnerable plaques and optimizing therapies accordingly
(Fassa et al. 2005; Kereiakes et al. 2003; Wykrzykowska et al. 2012) are expected to
improve clinical outcomes. A tool that has both a high positive predictive value and
a high negative predictive value in detecting vulnerable plaques, i.e., tell cardiologists to locally treat a plaque or not (Mintz 2016), is needed clinically. It holds great
promise in preventing plaque rupture and life-threatening sequelae (Fleg et al. 2012;
Waxman et al. 2006). Many imaging techniques have been used to image plaques
(Lusis 2000; Narula and Strauss 2007), including conventional fluoroscopic angiography, magnetic resonance imaging (MRI), X-ray or computed tomography (CT),
OCT, IVUS with virtual histology (VH), angioscopy and near-infrared fluorescence
(NIRF) imaging (Puri et al. 2011; Sanz and Fayad 2008; Narula and Strauss 2007).
However, none of these currently available technologies is sufficient to accurately
evaluate the vulnerability of a plaque (Sanz and Fayad 2008;Purietal.2011).
Particularly, a stand-alone intravascular imaging modality can visualize a certain
feature of a TCFA,but not all key features. IVUS, with its deep penetration, can image
the large necrotic pool and the total plaque volume of a TCFA. The limitation of IVUS
is its resolution (~100 µm). On the contrary, OCT, with its high resolution, permits
a definite measurement of the cap thickness of a plaque. This critical parameter for
characterizing TCFA is otherwise not accessible by other in vivo imaging techniques
(Farooq et al. 2009). However, the primary limitation of OCT for this application is
its penetration depth (~1.5 mm in tissue) which is not enough to reveal the necrotic
core size. Clinical evidence has also proven these limitations. A large cohort IVUS
trial confirmed that IVUS alone is not sufficient in identifying vulnerable plaques
(Stone et al. 2011, 2012). Similarly, clinical data (Mintz 2016) also revealed that the
diagnostic accuracy of OCT to identify vulnerable plaque is limited mainly due to its
shallow penetration depth and low molecular specificity (Virmani 2011; Prati et al.
2010; Bezerra et al. 2009).
As illustrated above, the characteristics of TCFA highlight the necessity of an
imaging technique with a high resolution to identify the thin cap and with a deep
enough penetration depth to visualize the necrotic core s imultaneously. Moreover,
several clinical studies were reported and validated the usefulness of a combined
use of IVUS and OCT in characterizing plaque compositions (Kawasaki et al. 2006;
Rieber et al. 2006) and detecting vulnerable plaques (Sawada et al. 2008; Zhang et al.
2014; Fujii et al. 2015; Mintz 2016). All these findings motivate the development of
a fully integrated intracoronary ultrasound-OCT system.
Technical Advances
In this section, we reviewthe technical development of the IVUS-OCT technology for
imaging vulnerable plaques in coronary arteries. Apart from common considerations
of an integrated IVUS-OCT system previously described, there are several particular
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technical challenges that need to be addressed to translate this technology for clinical
intravascular imaging. These challenges include miniaturizing the IVUS-OCT probe,
increasing the imaging speed and optimizing the flushing process.
Miniaturized IVUS-OCT Probe
The inner diameter of a human coronary artery is 2–5 mm. Patients with atherosclerosis may have coronary arteries that are even narrower. Accordingly, an intracoronary
imaging catheter (including the imaging probe and the outer sheath) is usually made
with an OD smaller than 1 mm and with a soft tip to avoid scraping and damaging
the coronary artery wall during imaging procedures. However, over-reducing the size
(either the thickness or the aperture area) of an OCT or ultrasound probe will sacrifice the performance of the integrated imaging system. The aperture size of a sensor
(either a transducer or an OCT lens) limits the lateral resolution and SNR that a system
can achieve, based on Eqs. (1), (2), and (5). In addition, as described previously, the
thicknesses of the piezoelectric element of a transducer affect its central frequency,
while thicknesses of the matching and backing layers influence the sensitivity and
the impedance of a transducer. Thus, the challenge of making an IVUS-OCT catheter
is to maintain good imaging quality of IVUS and OCT without making the crossing
profile of the IVUS-OCT catheter larger than 1 mm. To address this technical challenge, an innovative probe arrangement that can make good use of the space has to
be applied.
In 2010, our groups published the first demonstration of an integrated IVUSOCT probe (Yin et al. 2010). This probe applied a side-by-side arrangement and
achieved an OD of 2.4 mm, see Fig. 3.3a. The size of the GRIN lens used in the
OCT probe had an OD of 0.5 mm, and the size of the PZT transducer was around
2
0.4 × 0.4 mm
. To better use space and provide coaxial IVUS-OCT imaging ability, an inside-outside arrangement was proposed and achieved (Li et al. 2010)by
placing a focused ring transducer (made by LiNbO
) around a 0.7 mm-OD GRIN
3
lens, see Fig. 3.3b. Although the OD of this prototype was the same as the former
design (2.4 mm), apertures for the OCT and IVUS sensors were both larger, enabling
higher lateral resolutions. Particularly with a large-aperture focused transducer, the
measured lateral resolution of the IVUS sub-system was as high as 22 µm.
In 2011, we reported the first IVUS-OCT probe suitable for in vivo imaging in
a coronary-sized artery (Yin et al. 2011). A sequential arrangement was applied to
reduce the probe profile. The OD of the integrated probe was 0.68 mm. It utilized
a relatively large GRIN lens (0.35 mm OD) to maintain high lateral resolution of
the OCT sub-system. PMN-PT material, which has high k
, εs/ε0,and d33,was
t
used to fabricate a miniaturized transducer while maintaining high sensitivity of the
ultrasound sub-system. However, there is a 2-mm longitudinal offset between the
transducer and the OCT sensor due to the sequential arrangement, see Fig. 3.3c.
Although the obtained ultrasound and OCT signal could be co-registered by postprocessing, a real-time automatically co-registered arrangement would still be ideal
for providing real-time guidance for interventions in the cardiac catheter laboratory
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Fig. 3.3 Schematics of different IVUS-OCT probe designs. a Side-by-side arrangement (Yin et al.
2010); b inside–outside arrangement (Li et al. 2010); c sequential arrangement (Yin et al. 2011);
d back-to-back arrangement (Li et al. 2013b); e side-by-side 90° apart design (Li et al. 2013a)
in a hospital. Moreover, the offset made this probe have a long rigid part [5-mm long,
similar to previously reported designs (Li et al. 2010; Yin et al. 2010)]. It would be
difficult to navigate this probe through a coronary artery that has a narrow lumen and
sharp bends. To overcome these drawbacks, a modified version was reported in 2013
(Li et al. 2013b). This probe applied a “back-to-back” arrangement, see Fig. 3.3d,
to ensure a coaxial positioning of the transducer and the OCT sensor, while keep a
small OD. In addition, a ball lens (OD: 0.28 mm) instead of a miniaturized GRIN
lens (OD: 0.35 mm) was used to further reduce the size of the OCT probe. The OD of
this probe was 0.9 mm, and the rigid part of this probe was only 2.5 mm long, which
was the shortest ever reported. Apart from these designs proposed by our group, the
ultrasound group led by F. Stuart Foster and the OCT group led by Victor X. D. Yang,
in Canada have also reported an IVUS-OCT probe design, see Fig. 3.3e. The probe
used a side-by-side 90° apart design and a ultrasmall GRIN lens (OD 0.14 mm) to
reduce the size of the IVUS-OCT probe (Li et al. 2013a). The OD of this probe was
1 mm. F. Foster and V. Yang also briefly reported a new probe design in a review
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paper (Bourantas et al. 2016). Although limited information was provided, it was
stated that this new IVUS-OCT probe allowed optics to reside within the ultrasound
transducer to facilitate miniaturization and the IVUS and OCT beams travel in the
same direction. These four designs (Yin et al. 2011;Lietal.2013a, b; Bourantas
et al. 2016) described above ensure that the OD of the probe is smaller than 1 mm
and are likely to provide safe access to coronary arteries.
High-Speed Imaging System
During intracoronary imaging, the imaging catheter stays inside a coronary artery and
may partially block the normal blood flow. Coronary arteries supply blood to cardiac
muscle that pumps oxygen-rich blood throughout the body. Due to this critical role
of the coronary artery, the time needed for imaging the artery, i.e., the time to keep a
catheter inside it, should be as short as possible. Furthermore, a short imaging time
will also reduce the risk of flushing-induced complications. Because blood causes
high optical attenuation, flushing an artery is often necessary to generate clear IVOCT
images (Li et al. 2015a). However, the longer it takes to image, the more flushing
agents will be injected, which can be harmful for the patient (Ozaki et al. 2012). To
ensure the safety of using an IVUS-OCT catheter in patients, a vast amount of effort
was put in reducing the time needed to image a coronary artery and, accordingly,
increasing the imaging speed.
By innovatively applying state-of-the-art devices that were originally developed
for other fields (including optical communication and computer engineering) in the
development of an IVUS-OCT system, the imaging speed of IVUS-OCT has been
greatly increased. The first reported IVUS-OCT system (Yin et al. 2010) was able to
image at 1 fps by using separated DAQ cards for US and OCT sub-systems. Utilizing
a more powerful dual-channel DAQ card and a GPU for parallel computing, 4 fps
was accomplished (Yin et al. 2011). Subsequently, a 20-fps system was achieved (Li
et al. 2014a, c) by using a modified probe design that ensures a more uniform rotation
and smoother transmission of the rotation from the proximal to the distal end. With
a recently developed ultrasound pulser and a more advanced design of a customized
slip ring, the achievement of 72 fps was reported (Li et al. 2015b). By using such a
system, IVUS-OCT imaging of a coronary artery can be conducted within 4 s. As a
clinically used IVOCT system takes 3–6 s to image a coronary artery, it is reasonable
to believe that the imaging speed of this IVUS-OCT system is sufficient for clinical
applications.
Flushing
Intracoronary imaging is conducted in an environment naturally filled with blood.
As blood is a strong attenuation source for OCT and high-frequency ultrasound
signals, finding an optimal flushing agent that works well for both IVUS and OCT
can improve in vivo imaging quality.
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Flushing media with high viscosity can result in efficient blood displacement and
clear OCT images (Suter et al. 2015). However,flushing media with too high viscosity
will reduce the SNR of an ultrasound sub-system (Li et al. 2015a). This is because
friction and thermal consumption of energy are induced when sound propagates
through such media (Stokes 1849). To evaluate the effectiveness of different flushing
agents and speeds and optimize flushing for both ultrasound and OCT sub-systems,
we performed a well-controlled quantitative test in an in vitro model that mimics the
human arterial system and in vivo in a rabbit (Li et al. 2015a).
Apart from considering the effectiveness of a flushing medium, toxicity is also a
crucial concern. As the flushing agent is injected into blood and can be circulated
through the whole body of the patient, we need to avoid using a medium with high
toxicity (McCullough 2008; Ozaki et al. 2012). In consideration of both flushing
effectiveness and toxicity, it was found that Dextran is the best among three flushing
media we used (mannitol, Iohexol, and dextran). Other OCT flushing media, such as
glucose solution and propylene glycol (Tuchin et al. 2002) and iodixanol (Suter et al.
2015), may work too but need to be validated with an IVUS-OCT system in vivo.
These above developments have made the intracoronary application of IVUSOCT technically possible. Meanwhile, experiments in animal and human specimens
have been conducted to evaluate the clinical adaptability and the safety of these
developments which are reviewed below.
In Vivo and Ex Vivo Validations
In Vivo Validations
Rabbit and porcine models are commonly used (Ohtsuki et al. 2001) prior to human
experiments to validate the design of an intravascular imaging system. The aortas
of rabbits and the coronary arteries of pigs are comparable to the caliber of human
coronary arteries. In addition, rabbits and pigs can grow lesions similar to human
atherosclerotic plaques (Kolodgie et al. 2003; Schoenhagen et al. 2001).
The first in vivo demonstration of an integrated IVUS-OCT probe was achieved
in a rabbit (Yin et al. 2011). IVUS-OCT image pairs of a rabbit aorta were acquired
with a 4-fps system and a 0.68 mm-OD probe. The first in vivo intracoronary IVUSOCT imaging was conducted in a pig (Li et al. 2014a). Imaging was performed in
a swine left anterior descending artery by using a 20-fps system and a 3.7 F IVUSOCT catheter (Fig. 3.4). Successively, an in vivo rabbit experiment validated the
effectiveness of using dextran as the flushing agent for simultaneous IVUS-OCT
imaging (Li et al. 2015a). In 2015, the safe use of an ultrafast IVUS-OCT system
(72 fps) was demonstrated in live rabbits (Li et al. 2015b).
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Fig. 3.4 An OCT-IVUS image pair obtained in a normal swine coronary artery in vivo. a OCT
image, b IVUS image, c corresponding H&E histology. G: guide wire. T: tissue. V: vessel. I: intima.
E: EEL. A: adventitia. Scale bar: 1 mm. Reprint from (Li et al. 2014a) with permission
Ex Vivo Validation
Although animal experiments can evaluate the clinical adaptability and the safety
of these developed systems, few TCFA-like plaques can be found in animal models
(Brezinski 2006; Rekhter 2002; Kolodgie et al. 2003; Rekhter et al. 1998). Thus,
human cadaver samples were used to evaluate whether a device can acurately detect
TCFA. In 2015, we reported our work on demonstrating the usefulness of IVUSOCT in distinguishing a TCFA from false TCFAs (Li et al. 2015b), i.e., plaques that
are easily misdiagnosed by stand-alone IVUS or OCT. Fibrous caps were shown as
signal high regions in the OCT images (Fig. 3.5Ia, IIa, and IIIa). In IVUS and OCT
images, the signal-low regions behind the caps revealed the presence of necrotic or
lipid pools. As anticipated, the OCT imaging capability of the IVUS-OCT catheter
differentiated thin cap (Fig. 3.5Ia) and thick cap (Fig. 3.5IIa) whereas the ultra-
sound imaging capability in the IVUS-OCT catheter differentiated large (Fig. 3.5Ib)
and small (Fig. 3.5IIIb) necrotic cores. IVUS-OCT with simultaneous dual-imaging
capabilities clearly discerned differences between TCFA and false TCFA (plaques
with a thick cap or a small necrotic core).
Other Studies Related to IVUS-OCT
IVUS-OCT for Angioplasty Planning and Follow-Up
Apart from being used for the detection of TCFA, IVUS-OCT may also play an
important role in angioplasty planning and follow-up.
Angioplasty is a commonly used non-surgical procedure to treat arteries, such
as coronary or peripheral arteries (Arthurs et al. 2010), that were narrowed by
atherosclerotic plaques. During this procedure, stents that open up arteries are usually
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Fig. 3.5 Characterizing human atherosclerotic plaques by the ultrafast IVUS-OCT system. First
row: Example of a TCFA. (Ia) OCT image in which arrows point at the fibrous cap; (Ib)corresponding IVUS image indicates the location of necrotic core; (Ic) photograph of the corresponding
histology slide with CD 68 stain, highlighting macrophages and necrotic core. Middle row: A falsepositive case of TCFA diagnosis based on IVUS only (IIb) was produced due to the insufficient
resolution and sensitivity. Size of the thick cap can be determined by the corresponding OCT (IIa)
and CD 68 histology (IIc). Bottom row: A false-positive case of TCFA diagnosis based on OCT only
(IIIa) was produced due to OCT’s limited penetration depth. A small lipid pool can be determined
by IVUS (IIIb) and CD 68 histology (IIIc). Arrows denote the fibrous cap. NC: necrotic core; SLP:
small lipid pool. Scale bar: 0.5 mm. Reprint from the reference Li et al. (2015b) with permission
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