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259
11
Photoacoustic Tomography of Microcirculation
Song Hu and Lihong V. Wang
11.1
Introduction
Microcirculation, the periphery of the cardiovascular system, is responsible for
supplying gases and nutrients to biological tissues. Disorders in microcirculation
break the homeostasis of metabolism and may ultimately lead to organ morbidity.
Thus, in vivo imaging of microcirculation is of physiological, pathophysiological,
and clinical significance.
While optical microscopy [1–5] has been widely used to dissect the delicate
features of the microvasculature, label-free microvascular imaging beyond the
soft penetration limit (ballistic or quasiballistic regime) of biological tissues [6]
has proved to be challenging. Photoacoustic tomography (PAT), by detecting
endogenous or exogenous optical absorption contrasts acoustically, extends the
penetration limit of high-resolution microvascular imaging to the optical diffusive
regime [7]. In label-free PAT, red blood cells (RBCs) are usually irradiated by a
short-pulsed laser beam. The resulting transient thermoelastic expansion induces
high-frequency ultrasonic waves (photoacoustic waves), which encode the spatial
distribution of RBCs and thus can be used to reconstruct a three-dimensional
(3D) image of the microvascular structure and hemoglobin concentration [8]. With
multiwavelength measurements, PAT can further differentiate oxyhemoglobin
(HbO
2
or O
2
Hb) and deoxyhemoglobin (HbR or Hb) to quantify hemoglobin
oxygen saturation (sO
2
) [9]. With improved temporal resolution, blood flow rate
can also be assessed by analyzing RBC dynamics [10].
This chapter provides a comprehensive review of microvascular aspects of PAT,
including system designs, imaging methods, and recent progress in biomedical
applications. Photoacoustic theories, technical details of other embodiments
of PAT, and more biomedical applications can be found in the literature
[6, 11–13]. For this chapter, all the microvascular images were produced with a
7 ns wavelength-tunable laser system unless otherwise specified.
Microcirculation Imaging, First Edition. Edited by Martin J. Leahy.
2012 Wiley-VCH Verlag GmbH & Co. KGaA. Published 2012 by Wiley-VCH Verlag GmbH & Co. KGaA.
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260 11 Photoacoustic Tomography of Microcirculation
11.2
PAT Systems
PAT has two major embodiments: photoacoustic microscopy (PAM) and pho-
toacoustic computed tomography (PACT). PAM is based on two-dimensional
(2D) raster scanning of the overlapping dual foci of optical illumination and
acoustic detection. According to the different mechanisms used to achieve
lateral resolution, PAM can be further classified into two forms: optical-resolution
photoacoustic microscopy (OR-PAM) [14] and acoustic-resolution photoacoustic
microscopy (AR-PAM) [15, 16]. Instead of using time-intensive raster scanning,
PACT uses a linear or curved ultrasound array to achieve dynamic acoustic focusing
over the entire region of interest (ROI), with the aid of a reconstruction algorithm.
Such array-based PACT systems enable real-time imaging of hemodynamics [17];
however, the image quality needs to be improved [17, 18].
11.2.1
High-Resolution PAM
In PAM, the dual foci of optical illumination and ultrasonic detection are usually
configured coaxially and confocally to maximize the detection sensitivity. A smaller
overlapped focal area provides higher lateral resolution. In OR-PAM [14, 19], a
diffraction-limited optical focus, which is much smaller than the ultrasonic focus,
provides micrometer lateral resolution within the optical quasiballistic regime
(∼1 mm) or submicrometer lateral resolution within the optical ballistic regime
(∼0.1 mm) [6, 20]. In AR-PAM, by contrast, the acoustic focus – smaller than
the optical focus [16] – affords 50 µm resolution in the optical diffusive regime
(>3 mm). The penetration depth and spatial resolution of PAM can be scaled
simultaneously, while the ratio of the penetration depth to the axial resolution
(referred to as relative spatial resolution) remains as high as 200, which is adequate
for high-resolution imaging [7].
An OR-PAM system is shown in Figure 11.1 [21]. To ensure laser safety, the
pulsed laser beam is first attenuated by a tunable neutral density filter. Then,
a fraction of the pulse energy (10 –100 nJ) is spatially filtered by a pinhole and
focused by a microscope objective into a diffraction-limited spot (5 µm in diameter)
for photoacoustic excitation. To monitor the fluctuation in laser energy, 5% of
the incident light is reflected by an optical beam sampler to a photodiode. The
remaining light passes to an acoustic-optical beam splitter, which is positioned
under the objective to separate the optical illumination in the vertical direction
and the acoustic detection in the horizontal direction. The acoustic-optical beam
splitter consists of two right-angle prisms and a thin layer of silicone oil. The prism
glass and silicone oil have similar optical refractive indices but very dissimilar
acoustic impedances, which makes the layer optically transmissive but acoustically
reflective [22]. Through time-resolved ultrasonic detection and 2D raster scanning
of the acoustic-optical dual foci or the experimental object along the transverse
plane, a complete volumetric view of the biological tissue is recorded [8].
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11.2 PAT Systems 261
Neutral density
filter
Photodiode
Pinhole
Mirror
Wavelength-
tunable laser
Trigger
Computer
Amplifier
Condenser
lens
Optical beam
splitter
Microscope
objective
Acoustic-optical
beam splitter
Ultrasonic
transducer
Figure 11.1 Schematic of an optical-resolution photoacoustic microscopy system.
OR-PAM, similar to other ballistic imaging modalities (confocal microscopy,
two-photon microscopy, and optical coherence tomography), is strongly subject
to optical scattering, which degrades the optical focusing and thereby limits the
maximum penetration to ∼1 mm in biological tissues. However, this limitation
can be overcome by taking advantage of the ‘‘weak’’ acoustic scattering in tissues
(1000 times smaller than the optical scattering). Maslov et al. [15] have developed
AR-PAM, which maintains a relatively high lateral resolution (∼50 µm) beyond the
quasiballistic regime.
An AR-PAM system is shown in Figure 11.2 [16]. The light from a tunable
dye laser, pumped by an Nd:YAG laser, is delivered through a multimode optical
fiber, expanded by a conical lens, and then weakly focused into the biological
x
y
z
Motor driver
Translation
stages
Conical lens
Sample holder
Base
Tunable laser Nd: YAG pump laser
Photodiode
Amplifier
Optical illumination
Ultrasonic
transducer
Mirror
AD
Computer
Heater &
temperature
controller
Dual foci
Sample
Annular illumination
with a dark center
Figure 11.2 The dark-field acoustic-resolution photoacoustic microscopy system.
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262 11 Photoacoustic Tomography of Microcirculation
sample by an optical condenser. A ring-shaped optical illumination was formed on
the sample surface with a dark center. The use of an optically dark zone greatly
reduces surface signals and provides an ideal location for transducer placement. A
high-frequency ultrasonic transducer with a custom-made spherical focusing lens
is placed coaxially with the optical lenses. The ultrasonic and optical foci overlap
each other to produce a confocal configuration. Note that although the lateral
resolution of AR-PAM is achieved by acoustic focusing, optical illumination is still
weakly focused here to enhance regional photon deposition.
11.2.2
Real-Time PACT
Although PAM offers label-free microvascular imaging with high resolution and
sensitivity, its current implementation is subject to a time-intensive image acquisi-
tion process caused by the limited rates of raster scanning and laser repetition. To
enable real-time monitoring of acute hemodynamics, array-based PACT has been
actively studied [17, 23–25]. Raster scanning can be eliminated by numerically
focusing the photoacoustic signals detected by the ultrasound array to essentially
everywhere within the optical illumination. The limitation in laser repetition rate
can be partially overcome by using a full-field optical illumination of the entire
ROI, where a complete tomographic image can be acquired with a single laser shot.
A 512 element ring-shaped array-based PACT system is shown in Figure 11.3
[17, 24]. A pulsed laser beam is expanded by a concave lens and then homogenized
by a diffuser to provide uniform incident fluence over the entire ROI. The
ring-shaped array encircles the ROI to detect photoacoustic signals within the
plane perpendicular to the optical axis. The universal back-projection algorithm
[26] is applied to reconstruct tomographic views of the biological tissues in the
detection plane. This array-based PACT system can achieve a maximum imaging
Acquisition
computer
Trigger
Nd:YAG Ti:Sapphire
Lateral x
Lateral y
Elevation
Diffuser
or lens
64-ch, 40 MSPS DAQ
Sync
(×4)
Backplane & FPGA
Channel select Gain control
Amp 1
Amp 2
Amp 64
Amp ...
8 : 1 MUX
(×64)
PreAmp 2
PreAmp 1
PreAmp 512
PreAmp ...
Receiver electronics
H
2
O in bag
Transducer
array
3-Axis
Linear
stepper
motor
Figure 11.3 Schematic of the 512 element ring-shaped
array-based photoacoustic computed tomography system.
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11.3 Microcirculation Parameters Quantified by PAT 263
rate of 1 fps. Improvements in the data acquisition system can further increase the
temporal resolution [24].
11.3
Microcirculation Parameters Quantified by PAT
Taking advantage of the endogenous optical absorption contrast of hemoglobin –
the major constituent of RBCs and the primary oxygen carrier in the blood
stream – PAT can quantify important multiple functional parameters, such as the
concentration of total hemoglobin ([HbT]), sO
2
, and blood flow, within individual
microvessels. These quantification methods form the technical core of the
microvascular aspects of PAT and have been used in neurovascular and metabolic
studies [9, 27, 28]. Note that PAT can also visualize the metastasis of tumor cells
in the microcirculation by capitalizing on other endogenous chromophores (such
as melanin) or exogenous contrast agents [29, 30].
11.3.1
Total Hemoglobin Concentration and Microvascular Structure
In the visible spectral range, hemoglobin is the predominant optical absorber in
the blood stream. The photoacoustic signal generated within a blood vessel is pro-
portional to hemoglobin absorption. However, the two major forms of hemoglobin
(HbO
2
and HbR) have distinct absorption spectra, and thus the measured photoa-
coustic signal usually depends on hemoglobin concentration as well as oxygenation.
Fortunately, in the visible and near-infrared regions of the hemoglobin absorption
spectra, there are many isosbestic points (390, 422, 452, 500, 530, 545, 570, 584,
and 797 nm), where the molar extinction coefficients of HbO
2
and HbR are equal
[11]. Thus, photoacoustic signals acquired at such wavelengths reflect [HbT] only,
regardless of sO
2
. These isosbesticwavelengths are alsopreferred for imaging of mi-
crovascular structure because both oxygenated arteries and deoxygenated veins can
be imaged simultaneously without any oxygenation-induced bias in signal strength.
The [HbT] mapping of an entire mouse ear was acquired by OR-PAM at the
isosbestic wavelength of 570 nm (Figure 11.4). The signal strength, represented by
the pixel brightness in grayscale, indicates [HbT]. The lower brightness in small
capillaries indicates a lower [HbT] compared with that in the large arterial and
venous trunks. From the microvascular structure delineated by OR-PAM, we can
clearly distinguish the artery (open arrow) and the vein (diamond arrow) that form
an arterial–venous pair. We can further trace the venous branching down to the
capillary network through as many as 11 bifurcations (closed arrows).
11.3.2
Hemoglobin Oxygen Saturation
While an isosbestic wavelength is used to sense the [HbT], the dependence
of light absorption on hemoglobin oxygen saturation is used to measure sO
2
.
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264 11 Photoacoustic Tomography of Microcirculation
High
Relative
HbT
Low
500 µm
Figure 11.4 Optical-resolution photoacoustic microscopy
of the total hemoglobin concentration and microvascular
branching in a living mouse ear.
Multiwavelength PAT can image the relative concentrations of both HbO
2
and
HbR and consequently measure sO
2
within single microvessels.
In the visible spectral range, HbO
2
and HbR are the two dominant absorbers in
blood. Thus, the blood absorption coefficient can be expressed as [31]
µ
a
λ
i
, x, y, z
= ε
HbR
(
λ
i
)
·
[
HbR
]
+ ε
HbO
2
(
λ
i
)
·
[
HbO
2
]
(11.1)
where µ
a
λ
i
, x, y, z
is the blood absorption coefficient at wavelength λ
i
and position
(x, y, z); ε
HbR
(
λ
i
)
and ε
HbO
2
(
λ
i
)
are the molar extinction coefficients of HbR
and HbO
2
, respectively; and [HbR] and [HbO
2
] are their relative concentrations.
Assuming that the optical fluence is wavelength independent, we can use the
detected photoacoustic signal amplitude
λ
i
, x, y, z
to represent µ
a
λ
i
, x, y, z
in
Eq. (11.1) and calculate [HbR] and [HbO
2
] as relative values. Afterward, the sO
2
can
be computed as
sO
2
=
[
HbO
2
]
[
HbR
]
+
[
HbO
2
]
(11.2)
Note that [HbT] = [HbR] + [HbO
2
].
The sO
2
of a 1 × 1mm
2
region in a living nude mouse ear was mapped by
OR-PAM with dual wavelengths (570 and 578 nm) as shown in Figure 11.5. We
can clearly distinguish paired arteries and veins by their sO
2
values (∼0.95 vs
∼0.75). Note that the sO
2
values varied in the same blood vessel by ∼10% in
some specific regions (arrows in Figure 11.5), and the shapes of these regions look
like sebaceous glands. The inaccuracy in the sO
2
calculation is likely due to the
wavelength-dependent optical scattering of the glands.
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Соседние файлы в папке Библиотека им академика М.И. Перельмана