nihms159129.pdf

Multiscale photoacoustic microscopy and computed tomography

Lihong V. Wang

Optical Imaging Laboratory, Department of Biomedical Engineering, Washington University in St.

Louis, Campus Box 1097, One Brookings Drive, St. Louis, MO 63130-4899

Lihong V. Wang: LHWang@biomed.wustl.edu

Abstract

Photoacoustic tomography (PAT) is probably the fastest growing biomedical imaging technology

owing to its capability of high-resolution sensing of rich optical contrast

in vivo

at depths beyond the

optical transport mean free path (~1 mm in the skin). Existing high-resolution optical imaging

technologies, such as confocal microscopy and two-photon microscopy, have fundamentally

impacted biomedicine but cannot reach such depths. Taking advantage of low ultrasonic scattering,

PAT indirectly improves tissue transparency by 100 to 1000 fold and consequently enables deeply

penetrating functional and molecular imaging at high spatial resolution. Further, PAT holds the

promise of

in vivo

imaging at multiple length scales ranging from subcellular organelles to organs

with the same contrast origin, an important application in multiscale systems biology research.

1. Introduction

Definition of photoacoustic tomography: old physics and new technology

Photoacoustic tomography (PAT),

–

3 also referred to as optoacoustic tomography, is defined

as cross-sectional or three-dimensional (3D) imaging of a material based on the photoacoustic

effect. Planar imaging without depth resolution or one-dimensional (1D) depth-resolved

imaging does not belong to tomography. In the photoacoustic phenomenon, light is absorbed

by a material and converted to heat, and subsequently an acoustic wave is generated due to

thermoelastic expansion. PAT—a hybrid of optical imaging and ultrasound imaging—

combines rich optical contrast and high ultrasonic resolution in a single modality. PAT is

capable of providing high-resolution structural, functional,

–

7 and molecular

–

13 imaging

vivo

in optically scattering biological tissue at unprecedented depths. While structural imaging

cannot differentiate live and dead tissues, functional imaging can do so by measuring

physiological functions, such as the blood flow and oxygenation. In addition, molecular

imaging can sense biomarkers to identify specific cancer cells or detect gene expression

products to track gene activations.

The milestones of PAT and its precursor are briefly summarized here. One hundred years after

Bell reported the photoacoustic effect in 1880,14 Bowen envisioned its potential for imaging

with the excitation of ionizing radiation (high-energy electrons, x-ray photons, neutrons, and

other charged particles) or non-ionizing radiation (radiowaves, microwaves, and ultrasonic

waves).

He illustrated 1D depth-resolved imaging using radiowave excitation but never

mentioned the possibility of using optical excitation. Ten years later, laser-induced

photoacoustic depth-resolved imaging was achieved initially with no or poor (e.g., ~6 mm)

L. W. is the author to whom correspondence and requests for materials should be addressed.

9. Competing interests statement

The author has a financial interest in Microphotoacoustics, Inc. and Endra, Inc., which however did not support this work.

NIH Public Access

Author Manuscript

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Published in final edited form as:

Nat Photonics

. 2009 August 29; 3(9): 503–509. doi:10.1038/nphoton.2009.157.

NIH-PA Author Manuscript

lateral resolution in the direction perpendicular to the acoustic axis.16

–

20 To acquire cross-

sectional or 3D images that are both axially and laterally resolved, PAT based on an

approximate reconstruction algorithm was developed;21 alternatively, PAT based on tightly

focused ultrasonic detection was attained.22 In 2003,

in vivo

functional PAT was reported.

The subsequent rapid growth of research on PAT is highlighted in this article.

Rationale for optical imaging: rich optical contrast

Although many biomedical imaging modalities exist, their limitations call for new imaging

technologies. Optical imaging is a strong contender primarily because of its rich contrast. From

the safety perspective, light is nonionizing radiation with photon energy of the order of 2

electron volts, whereas x-ray has photon energy of thousands of electron volts and can

consequently cause carcinogenesis. From the physics perspective, optical signals are related

to the molecular constituents of tissue and hence can provide biochemical information. From

the optics perspective, optical absorption is related to various intrinsic contrast origins, such

as oxygenated and deoxygenated hemoglobin, melanin, lipids, and even water. As a result,

optical imaging can provide functional imaging—mapping of physiological parameters. By

using multiwavelength measurement, one can simultaneously quantify concentrations of

multiple chromophores of different colors, such as oxygenated and deoxygenated hemoglobin

molecules in red blood cells. Such quantification of hemoglobin can provide functional

imaging of the concentration and the oxygen saturation of hemoglobin—both parameters are

related to angiogenesis and hypermetabolism (hallmarks of cancer)—as well as brain activities

or injuries.

24 In addition, extrinsic optical absorption contrast agents can be used to

provide molecular imaging of gene expressions or biomarkers.8

–

Challenges in optical imaging: optical diffraction and diffusion

High-resolution optical microscopy faces two fundamental challenges: diffraction and

diffusion. Diffraction limits the spatial resolution, whereas diffusion limits the penetration. The

resolution of optical microscopy has been improved beyond the diffraction limit recently.

–

This article focuses on breaking the diffusion limit.

Despite light diffusion due to tissue scattering, a quick experiment by transmitting flash light

through the palm in the dark would prove the penetration capability of scattered light in

biological tissue. Although the flash light is broadband, only red light transmits substantially.

Unlike x-ray photons, optical photons do not cast a shadow of the bone. Such an observation

is well-understood in biomedical optics. Because hemoglobin—the primary chromophore in

biological tissue—absorbs light strongly in the green and blue spectral region, red light

propagates further before being absorbed than shorter wavelength light. Further, strong light

scattering by tissue components such as cell nuclei and mitochondria blurs any shadows of

internal structures. In fact, a red photon may experience hundreds to thousands of scattering

events before being absorbed. The reciprocal of the mean distance between equivalent isotropic

scattering events—called transport mean free path—is on the order of 1 mm in the visible and

near-infrared spectral range, around which photon propagation transitions from the ballistic

regime into the diffusive regime.

Because optical focusing is ineffective in the diffusive

regime, the transport mean free path is referred to as the soft depth limit for high-resolution

optical imaging. Beyond the soft depth limit, no pure optical imaging technologies have been

able to achieve high spatial resolution. Although scattering

per se

does not eliminate photons,

tortuous elongated photon paths boost the chance of tissue absorption. Even diffuse photons

are too few to be practically useful beyond a depth of ~50 mm, which is referred to as the hard

depth limit for optical imaging.

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Solutions to meet the challenges: photoacoustic tomography

PAT overcomes the soft depth limit and achieves deep imaging at high resolution. Ultrasonic

scattering coefficient in tissue is 2–3 orders of magnitude less than the optical counterpart. As

a result, PAT can provide high spatial resolution by detecting ultrasonic waves induced by

diffuse photons. In most cases, the center frequency and the bandwidth of the ultrasonic

detection system predominantly determine the spatial resolution; the greater they are, the better

the spatial resolution is, but the worse the ultrasonic penetration becomes. Such scalability

empowers multiscale imaging. It is worth mentioning that lower-frequency electromagnetic

waves including microwaves and radiowaves have been adopted for photoacoustic excitation

to overcome the optical hard depth limit.

Ultrasound imaging can provide high spatial resolution although ultrasound images appear

grainy owing to the speckle artifacts. The relative spatial resolution—defined as the ratio of

the penetration limit to the depth resolution (i.e., the pixel count in the depth direction without

artificial interpolation)—is more than 100, which is considered high enough to provide good-

quality images. From ultrasound imaging, PAT inherits the attribute of high spatial resolution

but rejects the trait of speckle formation. The speckle-free nature of PAT results from strong

correlations among the photoacoustic waves from the absorbers because all initial

photoacoustic pressure rises are positive. Consequently, prominent boundaries always build

up in photoacoustic images and suppress the interior fluctuations.

2. Fundamentals of PAT

Concept of PAT: optical excitation, ultrasonic detection, and image formation

Simply put, in PAT, light goes in and ultrasound comes out. For efficient ultrasound generation,

a short-pulsed laser is the usual choice. The strength of the photoacoustic pressure depends on

the optical energy deposition as well as the thermal and mechanical properties of the tissue.

Because either unscattered or scattered photons can produce photoacoustic signals as long as

photons are converted into heat, PAT can function in the optical diffusive regime. The image

formation is essentially triangulation of the photoacoustic sources according to the time-of-

flight signals recorded at multiple locations outside the tissue.

Two major modes of PAT: focused scanning tomography and computed tomography

PAT manifests in two major modes. In the mode of focused scanning PAT, such as dark-field

confocal photoacoustic microscopy (PAM),

33 a focused ultrasonic transducer is

mechanically scanned.29 At each scanned position, the acoustic focusing yields lateral

resolution while the acoustic time-of-flight provides axial resolution. In the mode of

photoacoustic computed tomography, an array of unfocused ultrasonic transducers detects

photoacoustic waves in parallel. Therefore, high frame rates—such as 50 Hz—can be attained.

–

However, an inverse reconstruction algorithm is required to reconstruct a

tomographic image. The early reconstruction was based on the assumption that each ultrasonic

transducer receives signals simultaneously from photoacoustic sources on a plane instead of a

spherical surface,

which can be difficult to satisfy in practice. Later, various groups

developed accurate reconstruction algorithms.

More recently, an exact closed-form

reconstruction formula was found for three detection geometries: planar, cylindrical, and

spherical configurations.

Interestingly, one may think of a focused ultrasonic transducer used

in the first mode as an analog computer that performs approximately the inverse algorithm in

the second mode. While hardware focusing does not present discretization artifacts, the focal

position cannot be adjusted as readily as in photoacoustic computed tomography. As a result,

the optimal lateral resolution is achievable only within the focal zone unless a virtual-detector-

based synthetic aperture algorithm is used.

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Contrast mechanisms of PAT: optical absorption and Grueneisen coefficient

PAT is exquisitely sensitive to optical absorption. The photoacoustic excitation converts a

small change in optical absorption coefficient to an equal fractional change in ultrasound signal.

For the same reason, PAT is also sensitive to any change in the Grueneisen coefficient—the

conversion efficiency from optical energy deposition to pressure, which can also provide image

contrast. In particular, the Grueneisen coefficient increases with the equilibrium temperature

(e.g., ~5% per Kelvin in water), which enables PAT to image temperature with a sensitivity of

0.16 °C for monitoring of thermal therapy.

–

Each mK of laser-induced temperature rise in soft tissue yields ~800 Pa initial pressure, which

is on the order of the sensitivities of typical ultrasonic transducers. The conversion from

temperature rise to pressure is related to two parameters: the volume expansion coefficient and

the compressibility. It is fortunate that the two parameters conspire to enable PAT to be

applicable in tissue with a safe temperature rise.

3. Multiscale photoacoustic microscopy

Dark-field confocal photoacoustic microscopy: 15

m axial resolution, 3 mm depth

PAM can image optical-absorption contrast at depths beyond the optical transport mean free

path with a ratio of penetration limit to depth resolution greater than 100 [Fig. 1(a)].

For

ease of system operation, an optical fiber is used to deliver the excitation laser light, which is

monitored by a photodiode for energy calibration. A conical lens is used to form a donut-shaped

beam (i.e., dark field) so as to minimize the effect of surface photoacoustic signals on the deeper

signals. The beam is then weakly focused into the tissue while the ultrasonic transducer focuses

coaxially into the same region for confocal excitation and detection. The photoacoustic time-

of-flight signal is recorded at each location of the ultrasonic transducer. By assuming the speed

of sound to be 1.54 mm/

s, the time-of-flight is converted into depth, and consequently a 1D

depth-resolved image (A-line) is formed. Linear or raster scanning over the tissue produces

2D or 3D tomographic images. Here, the broadband ultrasonic detector has a numerical

aperture of 0.44 and a center frequency of 50 MHz. As a result, the lateral resolution,

determined by the focal diameter of the ultrasonic transducer at the center frequency, measured

~45

m. The axial resolution, determined by the ultrasonic bandwidth, measured ~15

m. The

maximum imaging depth, determined by the ultrasonic attenuation in tissue, measured >3 mm.

33 A representative PAM image is shown in Fig. 1(b).

Photoacoustic macroscopy: 144

m axial resolution, 30 mm depth

The image resolution and the penetration limit are scalable with the ultrasonic frequency. For

example, employing a 5-MHz ultrasonic transducer and an 804-nm near-infrared light source

scaled the penetration to 30 mm, whereas the resolution is scaled up to 144 and 560

m in the

axial and lateral directions, respectively.

The near-infrared window is conducive for deep

imaging because of the low hemoglobin absorption. Here, photoacoustic microscopy turns into

macroscopy, which can potentially provide penetration up to the optical hard depth limit (~50

mm) at <1 mm resolution. Such a system can potentially image deep human tissue, such as the

sentinel lymph node (SLN), which has an average depth of 12±5 mm in ultrasound images.

in vivo

image of a rat overlaid with additional tissue was acquired with this system [Fig.

1(c)], where the SLN accumulated injected dye through the lymphatics.

This imaging

capability followed by needle biopsy can potentially transform SLN mapping—used for breast

cancer staging—from a surgical procedure into a minimally invasive procedure.

Optical-resolution photoacoustic microscopy: 5

m lateral resolution, 0.7 mm depth

The spatial resolution of the original PAM system can also be scaled down in dimension.

Focusing light through an objective lens with a numerical aperture of 0.1 yields a 5

m lateral

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resolution, which is limited by the optical focal diameter. Such a resolution allows

in vivo

imaging of capillaries—the smallest blood vessels—as shown in Fig. 1(d), where single files

of red blood cells are imaged.

While light delivery and ultrasonic detection are combined

through a beam combiner in this work,

they were also combined earlier through a ring

ultrasonic transducer.

If the numerical aperture is increased, sub-

m resolution can

potentially be achieved. Of course, optical focusing is effective only within the soft depth limit.

Thus far, a 0.7 mm penetration has been achieved.

Multiscale photoacoustic microscopy/macroscopy: enabling systems biology research

Multiscale systems biology, recognized as an important future direction of biomedical sciences,

needs the support of multiscale imaging

in vivo

with common signal origins. To understand

the workings of a whole biological system, biological components spanning multiple spatial

scales—subcellular organelles (sub-

m scale), through cells (

m), to organs (cm)—must be

integrated. As demonstrated above, photoacoustic microscopy and macroscopy are uniquely

positioned for such multiscale imaging with the same contrast mechanism—optical absorption.

Single-cell

in vivo

photoacoustic flow cytometry: melanoma detection

Photoacoustic sensing has recently been shown to detect single melanoma cells in lymph

vessels

in vivo

in small animals.

This technology, coined

in vivo

photoacoustic flow

cytometry, has potential for

in vivo

detection of circulating tumor cells in humans. Cells flowing

through a lymph valve are naturally focused into a single file, which facilitates single cell

detection in a lymph vessel. The lymph transparency presents low background and

consequently facilitates high-sensitivity detection. Melanin in the melanoma cells provides

high optical absorption for photoacoustic detection without the use of exogenous labeling,

which can be toxic. Nonetheless, exogenous contrast agents can be used to provide molecular

specificity for targeted biomarkers.

Several optical wavelengths can be used to detect

multiple types of cells of different characteristic optical absorption spectra. Toward human

applications, it was suggested that superficial lymphatics in the leg be first targeted.

The

concept can also be applied to blood vessels although background due to erythrocytes can

compete with signals from other absorbers.

–

Currently, the technology is limited to non-

imaging sensing; however, PAM can be explored for imaging.

4. Photoacoustic computed tomography

Circular-view computed tomography: 1D circular scan or ring array

In circular-view photoacoustic computed tomography, ultrasonic detection follows a ring.

A pulsed laser beam is expanded to illuminate the object. An ultrasonic transducer scans

around the object in a circle to acquire the photoacoustic data. An inverse algorithm then

converts the photoacoustic data to the initial photoacoustic pressure distribution, which is an

image of the cross section containing the scanned circle. Representative functional images are

shown in Fig. 2(a).

The ultrasonic transducer has a broad bandwidth centered at 3.5 MHz,

yielding an in-plane resolution of ~200

m. The out-of-plane resolution, determined by the

cylindrical focusing of the transducer, is ~1 mm. To accelerate the imaging speed, an array of

ultrasonic transducers in a ring has been adopted.

Planar-view computed tomography: Fabry-Perot detection array

In planar-view photoacoustic computed tomography, ultrasonic detection follows a plane.

Optical interferometry, instead of the piezoelectric effect, can be used to sense the ultrasonic

displacement

–

or pressure.

When a solid planar Fabry-Perot interferometer was

constructed,

a dielectric or polymer spacer was sandwiched between a pair of mirrors. A

focused laser beam was optically scanned across the surface of the interferometer. The detected

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power fluctuation of the reflected laser beam was converted to the optical thickness of the

interferometer, which was further converted to pressure. Then, an inverse algorithm was used

to reconstruct a photoacoustic image [Fig. 2(b)].

Optical detection offers primarily three advantages.

First, the optical detector can be made

optically transparent, facilitating light delivery to the object for full-field imaging. Second, the

spatial sampling frequency can be increased without sacrificing the sensitivity, allowing high-

resolution imaging. Sensitivities of tens of Pascals have been achieved at 20-MHz bandwidth.

Such sensitivities are comparable with those of 1-mm-diameter piezoelectric receivers and

better than those of smaller-diameter ones. Third, the frequency response can be stretched to

low frequencies, potentially enabling the detection of the interior of a large heterogeneity.

Presently, the technology is limited by the system complexity and the read-out speed.

Spherical-view computed tomography: rotation of object and detection by arc array

In spherical-view photoacoustic computed tomography, ultrasonic detection follows a

truncated spherical surface.

In a recently constructed small-animal whole-body

photoacoustic tomography system,

the animal, immersed in coupling liquid, is rotated inside

a virtual spherical surface, on which a concave arc-shaped array of 64 piezo-composite

ultrasonic transducers is situated. Two expanded counter-propagating laser beams illuminate

the small animal orthogonally to the plane of the array. At 755 nm wavelength, the internal

organs of a nude mouse were imaged [Fig. 2(c)].

Linear-view computed tomography: 1D linear ultrasonic array

In linear-view photoacoustic computed tomography, ultrasonic detection follows a line.

–

Linear ultrasound arrays in medical ultrasound imaging systems can be leveraged. The

ultrasound probes are typically handheld and can therefore adapt to various anatomical sites

conveniently. In ultrasound imaging, a subset of the array elements is switched on at a time

for ultrasound transmission and detection. Shifting the selection of the subset across the array

forms a two-dimensional image. In PAT, the entire array can receive ultrasonic signals

simultaneously because photoacoustic excitation usually covers a large volume. Naturally,

ultrasound and photoacoustic images can be displayed side-by-side to present complementary

contrasts.

Computed tomography based on linear or planar ultrasonic detectors

Photoacoustic computed tomography can be based on linear or planar ultrasonic transducers

instead of point detectors.

–

In this case, the detected signals represent an integration of

pressure over cylindrical or planar surfaces with varying distances to each transducer. Planar

ultrasonic transduction is of particular interest for primarily two reasons. First, the forward

problem represents a standard Radon transformation used in x-ray computed tomography.

Therefore, the associated inverse algorithms can be used directly for PAT. Second, the spatial

resolution is limited only by the ultrasonic bandwidth. In conventional photoacoustic computed

tomography, point ultrasonic detectors are approximated by finite-size detectors, resulting in

a tradeoff between the detection sensitivity and the lateral resolution. If the detectors are too

small, the sensitivity is compromised; if too large, the resolution becomes poor. Such a

compromise is devoid in PAT based on planar ultrasonic transducers.

Photoacoustic breast imaging

PAT lends itself to breast imaging. Breast tissue has relatively homogeneous (~±5%) speed of

sound, and breast tumors tend to present higher total blood concentrations and lower oxygen

saturations of hemoglobin than surrounding tissues. Figure 3(a) shows an

in vivo

photoacoustic

image of a human breast,

where the ultrasonic detection follows an arc.

The imaged lesion

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measured ~18 mm and centered at 23 mm deep from the laser-illumination surface; it was

confirmed by biopsy to be an invasive carcinoma. Figure 3(b) shows another

in vivo

photoacoustic image of a human breast,82 where the ultrasonic detection follows a plane.

The lesion was identified as a ring-shaped structure having an average optical absorption

contrast of 1.6 relative to the background. The lesion was estimated to be 30 mm in diameter,

which matched the pathologically estimated size (26 mm).

5. Photoacoustic molecular imaging

Small-molecule contrast: biomarkers imaged

in vivo

PAT is well suited for molecular imaging. While functional imaging is based on endogenous

contrast, molecular imaging is based on exogenous contrast—which essentially stains invisible

biomarkers

in vivo

. Functional imaging maps physiological activities usually at the tissue or

organ level, whereas molecular imaging measures biological and pathophysiological processes

at the molecular level. Simultaneous functional and molecular PAT of glioblastoma brain

tumors was demonstrated in live mice.

A small-molecule organic dye conjugated with a

cyclopeptide serves as the molecular imaging probe to target biomarker

integrin, which

tends to over-express in angiogenesis. The functional image shows tumor hypoxia, whereas

the molecular image shows tumor over-expression of integrin.

Nanoparticle contrast: biomarkers imaged

PAT can track various nanoparticles, such as metallic nanoshells

, nanorods

–

nanocages

, carbon nanotubes

, and dye-embedded nanoparticles

. The large optical

absorption cross sections of nanoparticles

provide excellent contrast for PAT. Unlike small-

molecule organic dyes, nanoparticles resist chemical or thermal denaturation and

photobleaching. However, further toxicity studies toward human applications are needed.

The surfaces of gold nanoparticles exhibit good biocompatibility and can accommodate

functionalities by conjugating with targeting peptides or antibodies.

So far, biomarkers

targeted in PAT include EGFR

, HER-2

and CXCR-4

, which have a propensity to

overexpress in cancers such as the breast cancer, as well as ICAM-1 in inflammatory responses.

In metallic nanoparticles, the optical absorption is based on surface plasmon resonance, and

the peak absorption wavelength can be tuned throughout the visible and near-infrared spectral

region by varying the physical dimensions.

Dye-embedded nanoparticles can also function as an effective contrast agent.

Indocyanine

green (ICG) is an organic dye that has been approved for human use. It has strong optical

absorption around the 800-nm near-infrared wavelength. Although free-form ICG has been

demonstrated to enhance photoacoustic contrast

in vivo

the half-life of ICG in the blood

stream is only several minutes, limiting the time window for sustained monitoring.

Encapsulation of ICG in nanoparticles increases the circulation time significantly.

addition, encapsulation provides surfaces for adding functionalities, increases the dye

concentration for greater optical absorption, and shields the dye molecules from destabilizing

effects due to the biological environment. ICG-embedded nanoparticles conjugated with

HER-2 antibodies have been shown to target breast cancer and prostate cancer cells

in vitro

Reporter gene contrast: gene expression imaged

in vivo

Molecular PAT of the lacZ reporter gene was demonstrated in living small animals.

Reporter

gene imaging, a form of molecular imaging, can facilitate understanding the nature of gene

expression during

in vivo

development and pathogenesis. While conventional reporter gene

imaging is based on fluorescent proteins such as the well-known green fluorescent protein,

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photoacoustic reporter gene imaging is based on optical absorption that is converted into heat.

The lacZ gene—one of the most widely used reporter genes—encodes

-galactosidase. A

sensitive chromogenic assay with X-gal—an optically transparent substrate—yields a stable

dark blue product, which serves as an optical absorption contrast agent for PAT. Because the

absorption of the blue product peaks around 650 nm, multi-millimeter penetration can be

achieved. However, if multi-centimeter penetration is desired, new reporter genes that produce

gene-expression products absorbing in the near-infrared spectral region need to be developed.

By providing high spatial resolution at depths beyond the optical transport mean free path

vivo

, PAT is expected to become a complementary tool to conventional optical imaging of

reporter genes.

Sensitivity to contrast agent

The sensitivity of PAT is competitive among the various molecular imaging modalities

although such a comparison depends on many parameters as well as the biomedical

applications.

When intrinsic biological contrasts due to such as red blood cells or melanoma

cells are sensed, single-cell sensitivity has been achieved.

When extrinsic small-molecule

contrast agents are imaged, PAT can provide a sensitivity of the order of ~10 nmol/L or ~fmol/

voxel at ~0.3 mm resolution.

Pure optical fluorescence imaging can provide a sensitivity

of the order of ~nmol/L or ~fmol/voxel at ~1 mm resolution,

where the resolution is worse

than that of PAT and the sensitivity is slightly better. Nuclear imaging such as PET can provide

a sensitivity of the order of ~pmol/L or ~0.1 fmol/voxel at ~3 mm resolution, where the

resolution is much worse than that of PAT and the sensitivity is better.

However, the

sensitivity of PAT can be improved potentially by increasing parameters such as the local

optical fluence, the optical absorption cross section of the contrast agent, the number of signals

for averaging, and the voxel size. For example, when extrinsic nanoshell contrast agents are

imaged, PAT can provide a sensitivity of the order of ~10 pmol/L or ~0.0001 fmol/voxel at

~0.06 mm resolution.

Because different modalities have dissimilar imaging capabilities, the

sensitivity should not be the only parameter governing the selection of a tool for a particular

application. In additional, nuclear imaging uses ionizing radiation, whereas PAT and optical

fluorescence imaging use nonionizing radiation.

6. Limitations of PAT and future outlook

PAT has limitations as does any imaging technology. First, optical attenuation limits the

penetration to ~5 cm when <1 mm resolution is desired in actual tissue.

100

101

However,

microwaves or radiowaves can be used for deeper excitation although the contrast origins

differ. Second, ultrasound sustains strong reflection from gas–liquid or gas–solid interfaces

due to the strong mismatch of acoustic impedances. Therefore, ultrasound signals cannot

penetrate through gas cavities or lung tissues efficiently. For the same reason, ultrasonic

detection requires direct contact between the ultrasonic transducers and the biological tissue.

Usually, ultrasound coupling gel is applied to the tissue surface to avoid intervening air cavities.

The development of noncontact optical detection of acoustic displacement can potentially

overcome this problem.

–

Third, ultrasound suffers significant attenuation and phase

distortion in thick bones such as the human skull. Fortunately, unlike pulse-echo ultrasound

imaging, PAT involves only one-way ultrasound attenuation through the skull. Sufficiently

strong photoacoustic signals have been observed through Rhesus monkey skulls.

102

103

The

remaining challenge is to compensate for the phase distortion due to the skull.

The research field on PAT has gained significant momentum recently. The number of research

articles on PAT has more than tripled since year 2003.

104

PAT is expected to find broad

applications in both biology and medicine. In the author’s opinion, the following

in vivo

applications hold particular promises and may come to fruition in the near future: melanoma

detection,

intravascular catheter imaging,

105

–

107

endoscopic imaging, simultaneous

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functional and molecular imaging,

reporter gene imaging,

sentinel lymph node mapping,

breast imaging,

108

–

110

and multiscale imaging.

Acknowledgments

The author acknowledges the support by the National Institutes of Health Grants R01 EB000712, R01 NS046214,

R01 EB008085, and U54 CA136398. Thanks to Song Hu for providing Fig. 1(d).

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Figure 1. Multiscale scanning photoacoustic imaging of small animals

in vivo

(a) Annotated photograph of the dark-field confocal photoacoustic microscope.32 (b)

Photoacoustic image of a melanoma and blood vessels acquired with a 50-MHz photoacoustic

microscope.32 M: melanoma. Axial resolution: 15

m. Penetration limit: 3 mm. (c)

Photoacoustic image of a sentinel lymph node (SLN) 18 mm below the laser-illumination

surface acquired with a 5-MHz photoacoustic macroscope.50 Axial resolution: 144

Penetration limit: 30 mm. (d) Photoacoustic image of the vasculature, including capillaries,

acquired with an optical-resolution photoacoustic microscope.51 Lateral resolution: 5

Penetration limit: 0.7 mm.

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Figure 2. Non-invasive photoacoustic computed tomography of small animals

(a) Functional images of a rat brain acquired

in vivo

with left- and right-side whisker

stimulations, respectively.

Color: differential absorption

due to brain activation. Gray:

structure of the cortex. Field of view: 20 × 20 mm. (b) Left: Schematic of superficial cerebral

vascular anatomy: A, superior sagittal sinus; B, lateral sinus; C, inferior cerebral vein. Right:

Image of a mouse brain acquired

ex vivo

with a system based on Fabry-Perot interferometry.

in vivo

, where a 10-mm thick section of the

abdomen is shown.

Wang

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Figure 3.

In vivo

photoacoustic computed tomography of the human breast acquired at 1064-nm

laser wavelength

(a) Image of a human breast containing a poorly differentiated infiltrating ductal carcinoma.

The slice is 21 mm deep from the laser-illumination surface.

Field of view: 120 × 120 mm.

(b) Image of a breast containing an invasive ductal carcinoma. The slice is 13.5 mm deep from

the laser-illumination surface.82

Wang

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