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Optoacoustic imaging of blood
for visualization and diagnostics of breast cancer
Alexander A. Oraevsky1,3, Elena V. Savateeva1, Serge V. Solomatin1, Alexander A. Karabutov1,3,
Valeriy G. Andreev1, Zoran Gatalica1, Tuenchit Khamapirad1, P. Mark Henrichs2
1University of Texas Medical Branch, Galveston TX 77555-0456
2LaserSonix Technologies, Inc., Galveston, TX 77550
ABSTRACT
Aggressive malignant tumors may be diagnosed based on relative concentration of oxyhemoglobin and deoxyhemoglobin in
the tumor microvasculature. Optoacoustic images of breast cancer and prostate cancer may be acquired at two laser
wavelengths matching maximum of oxyhemoglobin (1064-nm, Nd:YAG laser) and deoxyhemoglobin (760-nm, Alexandrite
laser). Two optoacoustic systems operating in forward and backward mode respectively for breast cancer and prostate cancer
detection, employing arrays of ultravide-band piezoelectric transducers and multichannel electronics was described. After
systems testing and calibration in phantoms, initial experiments were performed on patients with suspicious tumors.
Quantitative analysis of two-color optoacoustic images was correlated with biopsy and histology. Possibility for tumor
differentiation was demonstrated.
Keywords: Optoacoustic imaging, breast cancer, ultrawide-band acoustic transducer, angiogenesis, hemoglobin
1. INTRODUCTION
At nearly 600,000 new cases each year, breast cancer is an alarmingly common and serious health concern for women
worldwide. In North America, breast cancer accounts for over one quarter of all female cancers. About 12.8% of all
American women develop breast cancer during their life. It is estimated that 192,200 new cases was diagnosed and 40,200
deaths will occur as a result of breast cancer this year [1]. Because of the high incidence, even small improvements in breast
cancer diagnosis and treatment may save tens of thousand lives every year. Statistics indicates that detection and follow-up
in the early stages of cancer offer a much greater chance of survival than detection and follow-up at later stages. This is
particularly true for breast cancer.
Diagnostic imaging of tissues and organs has been a field of rapid advances, especially with the novel modalities of
ultrasound imaging (USI), nuclear medicine (NM), nuclear magnetic resonance (MRI), X-ray CT, bioelectric and
optical/spectroscopic cancer visualization strategies [2]. However, an ultimate imaging technology combining high
sensitivity and high specificity with safety and convenience for the patient is yet to be developed.
1.1. Current imaging techniques
The gold standard, X-ray radiography, and ultrasound are the two imaging techniques that are currently available for breast
cancer detection. However, tumor differentiation with these technologies is near to impossible, especially in the absence of
calcifications. It is recognized that the magnetic resonance imaging has a potential in breast cancer diagnostics, but at the
present time this promising technology is in its research and development stage. All these techniques have some merits for
diagnosis of a specific tissue abnormality or lesion. However, all of them are not free of difficulties in application [2].
X-ray radiography (mammography) has found a wide application in breast cancer screening. However, mammography has
several limitations for applications in diagnostics: (1) The minimal size of tumors detectable by mammography is about 5-10
mm; (2) There is a high percentage of false-positive (80%) and false-negative (30%) x-ray mammograms (total percentage
of false mammograms reported by various clinics varies [3]); (3) Images of radiographically dense breasts have low contrast;
(4) X-ray mammography utilizes carcinogenic ionizing radiation. While mammograms are good at detecting cancer in post-
menopausal women, spotting tumors in the breasts of women under 40 is more challenging because the breast tissue is so
dense [4].
Biomedical Optoacoustics III, Alexander A. Oraevsky, Editor,
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Ultrasound imaging is also being applied in clinics for breast cancer diagnosis. The main merit of ultrasonography is
utilization of pressure waves with 2.5-10 MHz ultrasonic frequencies, which can propagate substantial distances in biological
tissues with minimal attenuation and alterations. However, the limit of detection by current ultrasonic imaging systems at a
depth of up to 6 cm inside biological tissues is limited to ~10 mm due to low acoustic contrast between the tumors and
normal tissues [5]. In addition, heterogeneous structure of the breast yields complex ultrasound images with numerous
artifacts resulting from multiple reflections of the ultrasonic waves from tissue boundaries. Recently three major advances of
ultrasound imaging were reported: (1) high-frequency harmonic ultrasonic imaging [5], (2) color Doppler [6] and (3)
elasticity detection [7]. However, ultrasound technology did not find a parameter for measurement that directly relates to
breast cancer malignancy.
Magnetic Resonance Imaging is a potent imaging technology. The major merit of MRI is high resolution of less than 1-mm
[8]. This technique is capable of characterizing the tissue chemical content. However, the sensitivity of MRI in many
instances is not very impressive, and not all types of breast tissues can be differentiated by this technology, which is also very
expensive. In recent studies MRI demonstrated utility in assessing the density of microvasculature in breast cancer
angiogenesis [9]. MRI tuned to the tumor angiogenesis showed a two-fold improvement in sensitivity and a three-fold
improvement in specificity over mammography and physical examination [9]. Additionally, MRI has a 100% negative
predictive value, which makes it useful for detection of early cancer [10]. However, the accuracy of tumor differentiation
with MRI was not satisfactory. Dynamic nuclear magnetic resonance imaging was shown to have a better diagnostic utility
when applied with gadolinium contrast agent [11].
The typical size of tumor detection by the conventional diagnostic modalities is usually about 10 mm. Such tumors are
already biologically advanced and capable of metastasizing. Furthermore, no one state-of-the-art imaging technology
provides certain diagnostic information. Therefore, surgical biopsy and needle biopsy are being applied for differentiation
between malignant and benign lesions and cysts [12]. High false-positive (80%) and false-negative (30%) rates represent a
significant problem for breast cancer screening. Therefore, an imaging modality highly specific for malignant tumors,
especially smaller tumors in radiographically dense breasts would be of great importance for successful and inexpensive
treatment and, eventually, for reduction of breast cancer mortality in the USA.
1.2. Angiogenesis as a marker for aggressively developing cancer
As discovered by J. Folkman, rapidly growing cancer cells need additional blood supply and gradually develop a dense
microvascular network inside or around tumors required for tumor growth and progression. Vascularization of the tumor
occurs through a series of sequential steps before or during the multistep progression to neoplasia [13,14]. Angiogenesis
appears to be a marker for future breast cancer development and may have clinical implications in diagnosis and treatment.
There is a constant requirement for vascular supply in solid tumors. Tumor-associated neovascularization allows the tumor
cells to express their critical growth advantage. Experimental and clinical evidence suggests that the process of metastasis is
also angiogenesis-dependent [15,16]. Tumor angiogenesis is employed as a target for breast cancer chemotherapy agents
[16,17].
In the early stage of a cancer development the tumor depends on blood vessels in the surrounding healthy tissue to support its
continued growth. A tumor without its own independent supply of new blood vessels may be restricted in its growth to 1 or 2
mm in size with just a few million cells and be non-life threatening. Many of these breast tumors will remain at this in situ
stage for years before switching into a rapid growth stage when the rate of the cell growth is much greater then the rate of
apoptosis and requires new capillary blood vessels. Most of the new vessel growth found around malignancies have no
smooth muscle. Thus, they are not enervated, rendering them unresponsive to the control of epinephrine (adrenaline). They
have fragmented basement membranes and are leaky, making these new vessels more penetrable by tumor cells than are
mature vessels. This new vascular phase of growth is usually followed by rapid tumor growth, bleeding and metastasis.
It was demonstrated that malignant breast tumors (carcinoma) not only have enhanced blood content, but also contain
noticeably hypoxic blood [18,19]. In contrast, benign tumors (fibroadenomas) have relatively normal level of blood
oxygenation. Spectroscopic imaging, i.e. optical imaging using several optical wavelengths allow the noninvasive, in vivo
imaging and quantification of oxygenated and deoxygenated hemoglobin that reflect physiologic functions of tumors [20,21].
1.3. Optical tomography: strong and weak sides of the emerging diagnostic technology
Advantages. Rapidly emerging optical imaging technologies based on time-resolved measurements of transmitted laser
pulses or phase-resolved detection of diffusely scattered photon density waves employ high optical contrast of deeply-
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penetrating near-infrared radiation [22,23]. This high optical tissue contrast is associated primarily with blood in tumor
angiogenesis. Hemoglobin is the major chromophore in biological tissues in the near infra-red spectral range. Blood
absorption coefficient equals about 5-10 cm-1 at l ≈ 750-1050 nm [24, 25], while background absorption in normal breast
tissue is only 0.03-0.05 cm-1 [26-28]. The increased concentration of strongly absorbing hemoglobin was found to yield 2-8
fold optical contrast between breast tumors and normal tissues in vitro and in vivo [20,21,26,27]. Significant optical contrast
between normal tissues and tumor angiogenesis was found also in other types of cancer, such as colon and prostate
carcinomas [29,30]. Naturally occurring optical contrast may be further increased with application of contrast-enhancing
agents, especially agents employing liposome encapsulated dyes and nanoparticles [31,32].
Having association with tumor angiogenesis, optical imaging possesses capacity to resolve functional and molecular
characteristics of breast cancer. Hence, malignancy differentiation can be based on a novel set of functional features that are
complementary to current radiological imaging methods. These features include concentration of hemoglobin in tumors and
the rate of oxygen consumption by tumors. Functional imaging may enhance diagnostic accuracy of breast cancer detection,
lower the current state-of-the-art detection limits, and play a vital role in therapeutic strategy and monitoring. Feasibility of
optical tomography as an adjuvant to ultrasonography for diagnosis of breast cancer was recently demonstrated [33].
Drawbacks. Near infrared diffuse optical spectroscopy and time-resolved optical imaging evolve as promising methods that
eventually may enhance diagnostic capabilities of existing technologies for breast cancer screening. Optical techniques are
based on quantitative measurements of optical and functional contrast between healthy and diseased tissue. Because of their
unique, quantitative information content, diffuse optical methods may play an important role in breast diagnostics and
improving our understanding of breast disease. However, ubiquitous light scattering in breast tissues presents a great obstacle
to pure optical imaging. The photon scattering substantially reduces resolution and accuracy of localization of optical
tomography (to about 1-cm) and limits its sensitivity and maximal depth of tumor detection (to about 3-cm). Because of these
limitations pure optical tomography has not yet found applications as a clinical modality for cancer detection and diagnosis.
1.4. Optoacoustic imaging combines merits of optical contrast and ultrasonic resolution
Optoacoustic tomography was recently proposed as a technique for quantitative tissue characterization and early cancer
detection [34-47] and specifically for imaging breast cancer [48-52,55-57]. This technology combines the most compelling
features of light and sound: pronounced optical contrast and high resolution of ultrawide-band ultrasound detection, yielding
a more sensitive imaging modality with substantially improved accuracy of the localization (<1-mm). The hybrid of optical
and acoustic technologies permit to overcome the problems of pure optical imaging associated with the loss of resolution and
sensitivity due to strong light scattering and problems of pure ultrasound imaging associated with poor tissue contrast and
low signal-to-noise ratio.
The proposed tomography utilizes the optoacoustic effect. Absorption of pulsed laser radiation leads to preferential energy
deposition in volumes with enhanced absorption (i.e. tumors). Absorbed laser energy yields temperature rise and
simultaneous pressure jump in the irradiated volume, thereby converting heated volume into a source of ultrasonic waves.
Short-pulse (nanosecond) energy deposition results in high thermoacoustic efficiency of pressure generation (up to 0.8) in
live tissues and enables superior spatial resolution of 10 µm to 500 µm, depending on the depth of monitoring from 1-mm to
5-cm [32-36,41-44]. The optically generated pressure profile resembles the distribution of absorbed energy in the irradiated
tissue. Stronger optical absorption in tumors leads to a higher pressure-amplitude in comparison with the amplitude of
background signals from normal tissue. The signal generated by the optoacoustic imaging system is initially a spatial
replicate of the tumor in the direction of detection. This signal mimics the spatial distribution of photon absorption creating
an ultrasonic wave that mimics the shape of the object. The propagation of this signal to the surface follows the rules of
ultrasound transmission, but the signal itself is distinct from typical ultrasound. Tumors with dimensions from 1 mm to 10
mm irradiated with laser pulses represent themselves as sources of pressure waves with ultrawide ultrasonic frequency range
of 150 kHz to 1.5 MHz. Such ultrasonic waves can propagate in tissues with insignificant attenuation and deliver spatially
resolved information to the surface of tissue [40]. However, undistorted detection of ultrasound within such a wide frequency
range represented a great challenge during initial development stage of the optoacoustic tomography [47].
In contrast to pure optical tomography, where diagnostic information about tissue structure is integrated over the entire
optical path, the optoacoustic imaging permits direct reconstruction of the absorbed energy distribution from the profile of
laser-induced pressure [34-36,40,43]. The sensitive time-resolved piezoelectric detection and analysis of the laser-induced
ultrasonic profiles offers a unique possibility to visualize objects at the depth of 8-cm in optically turbid and opaque tissues
and provide quantitative information about tissue structure with accuracy of better than 1-mm. In contrast to pure ultrasonic
imaging, where ultrasonic waves delivered to tissue from outside and only fraction of total ultrasonic energy is reflected back
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from tissue boundaries, the optoacoustic imaging converts tumors into bright acoustic sources standing out on a dark (low
absorbing) background of normal tissues, and detects 100% of ultrasonic energy emitted by the tumors. Thus, Laser
Optoacoustic Imaging System (LOIS) utilizing sensitive detection of laser-induced ultrasonic waves instead of the detection
of scattered photons is, therefore, free of the limitations associated with the pure optical and pure ultrasound technologies.
Experimental and theoretical studies of 1980s created a solid background for the laser optoacoustic imaging system to be
invented [51]. There are several research groups in the USA and in Europe that made material contributions to the field of
laser optoacoustic imaging. The group of Dr. Kruger focuses its attention specifically on development of the breast cancer
imager using microwaves as an excitation source [46,48]. Optoacoustic imaging technology utilizing green laser pulses is
being applied by a group of Dr. F. deMul for detection of hemorrhages, hematomas and leaky blood vessels in skin [39,40].
Dr. Beard’s group has made significant progress in the development of miniature endoscopic optoacoustic system for
monitoring of atherosclerotic lesions in blood vessels [37]. A modification of LOIS yielding measurements of ultrasonic
waves at the site of laser irradiation opens opportunity for optoacoustic imaging and quantitative characterization of various
human organs [42,44,47].
2. MATERIALS AND METHODS
2.1. Gel phantoms
Breast phantoms were made of 10%-gelatin and had cylindrical shape with dimensions of 120-mm x 120 mm. Water
absorption (µa = 0.15 cm-1 at 1064 nm [20]) provided the absorption coefficient of 0.13 cm-1 in the gel phantoms. The typical
effective attenuation coefficient, µeff, for normal breast tissue in vivo is about 1.0 to 1.2 cm-1 in the entire near-IR spectral
range. The gradual increase of water absorption in the wavelength range from 700-nm to 1100 nm is accompanied by
simultaneous decrease of scattering, making the effective optical attenuation almost constant. The effective optical
attenuation of the skin in the near-IR equals 1.3-1.4 cm-1. Polystyrene spheres were used to provide effective scattering
coefficient of 4 cm-1. Absorption by water and scattering by polystyrene microspheres yielded the effective attenuation µeff
of 1.3 cm-1 for our phantom at the wavelength of 1064-nm. Artificial blood vessels made of polyethelene tubes have been
embedded in the phantom gel and filled with the blood of sheep. The vessels had various shapes including straight pipes and
loops with absorption coefficient in the range of 0.45 to 1.0 cm-1 depending on the level of hemoglobin oxygenation.
Diameter of the vessels have been varied from 1.5-mm to 4.5-mm. Liquid gel at the temperatures very close to the
solidification temperature was poured inside a cylindrical container made of Plexiglas. After solidification in a refrigerator
the phantom was taken out from the Plexiglas container and examined with LOIS. The breast phantoms were made
acoustically homogeneous and dense, while the ratio of absorption coefficients of the blood vessels and surrounding gel was
ranging from 3 to 6, a value anticipated between breast tissues and cancerous tumors in vivo.
2.2. Breast cancer patients
Female patients screened with x-ray mammography and/or ultrasound, who had masses or lesions in the breast suspicious to
be cancerous, have been recruited to participate in this study. Approximate location of tumors was known prior to
optoacoustic imaging procedure. However, LOIS images have been obtained prior to biopsy that provided the final diagnosis
for comparison with that made noninvasively.
Patients were placed on the stereotactic examination table routinely used for biopsy procedures. The optoacoustic imaging
was performed by attaching the array transducer to one side of the breast approximately above the tumor and illuminating the
breast with a single optical fiber on the opposite side of the breast. The surface of the detector array was preliminary
lubricated with Sonogel™ in order to provide a good acoustic contact between breast tissues and the transducers. The
illuminating system with objective lens was moved along the breast surface to find an optimal position for the tumor
illumination. An optoacoustic image required at least 1 second for illumination. The entire imaging procedure including data
acquisition and image reconstruction took about 2 seconds, so that the optoacoustic imaging was performed in close-to real
time. Position of the detector array and the optical fiber was recorded during imaging. The LOIS images were compared with
x-ray mammograms, ultrasonograms obtained prior to LOIS procedure and with biopsy obtained after imaging.
2.3. Laser and light delivery system
Schematic diagram and a photograph of the laser optoacoustic imaging system (LOIS) is shown in Fig. 1a,b. Two compact
pulsed solid-state lasers were incorporated in the system enabling two-color irradiation for optoacoustic imaging. The two
lasers were coupled into one fiberoptic light delivery system. An Nd:YAG laser (Big Sky Lasers, MA) was operated at the
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wavelength of 1064-nm and an Alexandrite laser (Laser Energetics, NJ) was operated at the wavelength of 757-nm.
Repetition rate of laser pulses was 20 Hz for both lasers. A 1-mm diameter quartz optical fiber was employed for delivery of
laser pulses to the tissue surface. An objective lens attached to the housing of the optoacoustic transducer array was used to
expand laser beam from the fiber and deliver a circular beam with diameter of about 8-mm to the skin surface of the breast.
Array of Acoustic Transducers
Computer System
for Data Acquisition
and Processing
Array of Optical Fibers
Breast
with T um o r
LASER
a
Monitor
Transducer
Optical
Fiber
Laser
Data Boards
b
Figure 1. Schematic diagram of clinical LOIS (a) and a photograph of the 32-element arc transducer array (b).
The optoacoustic imaging procedure was performed using laser fluence of 10 to 20 mJ/cm2, which corresponds to one tenth
to one fifth of maximum permissible laser exposure for the skin in the wavelength from 700 nm to 1100 nm (see American
National Standards for Safe Use of Lasers ANSI Z136.1-2000). The laser irradiance of 10-20 mJ/cm2 can induce a
temperature rise of only 40-80 mK at the irradiated breast surface and less than 1-mK in tumor located at a 5-cm depth. The
laser irradiation was performed with 20-ns laser pulses. Such a short duration of laser pulses ensures that the laser-induced
pressure was confined in the target volume of tumor during laser heating. The irradiation conditions of temporal pressure
confinement within the irradiated voxels to be resolved on the image, will result in instant pressure generation and serve two
major goals. First, the profile of laser-induced pressure can mimic the shape and dimensions of a tumor only when laser
heating is much faster than thermal expansion of the tumor. Secondly, short laser pulses yield exceptional optoacoustic
efficiency (or otherwise called Grüneisen parameter) of 0.7-0.8 for conversion of absorbed optical energy into pressure
(ultrasonic waves). The method of laser irradiation under conditions of temporal pressure confinement yielding an
exceptional efficiency of the optoacoustic generation [38].
2.4. Array of ultrawide-band acoustic transducers
A specially designed array of PVDF acoustic transducers was employed in LOIS (LaserSonix Technologies, Inc., TX). The
array had 32 rectangular piezoelectric transducers with dimensions of 1.5 x 15 mm and 2 mm distance between transducers.
The detector elements were placed on an arc with diameter of 120-mm. Piezoelectric polymer PVDF of 110-µm thickness
was used for transducer fabrication. Low acoustic impedance and ability to operate in the ultrawide ultrasonic frequency
band without strong resonance and reverberations are the advantages of PVDF for detection of optoacoustic profiles. Special
polymer backing material was used for acoustomechanical matching of the transducer, widening the ultrasonic band of
detection and damping reverberations after detection. The transducer array was properly electrically shielded. The front
surface of transducer array was covered with aluminum foil to reduce external electric noise. The thickness of Aluminum
layer was optimized to achieve optimum between ultrasound attenuation and electric noise reduction. The housing of the
transducer array was also made of aluminum. Acoustically absorbing resin was placed between piezoelectric elements and
housing to prevent ringing of the housing that follows the vibration of the piezoelectric elements themselves.
3. EXPERIMENTAL RESULTS
3.1. Optoacoustic detection of single erythrocytes.
A droplet of whole blood was put in a cuvette with aqueous solution of polystyrene microspheres in PBS buffer with
effective optical scattering coefficient µeff =5cm−1 at the wavelength of 532-nm. An ultrawide-band acoustic transducer
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with sensitivity of 1V/bar made of 0.1-mm thick PVDF piezo-copolymer film was placed at the bottom of the aqueous
solution. A free surface of the solution open to air was irradiated with single nanosecond pulses from Nd:YAG laser second
harmonic with incident energy of 2.5-mJ (Figure 2a). Butterfly-shaped optoacoustic signals typical of spherical ultrasonic
waves were detected on the screen of the digital oscilloscope (Fig. 2b). The solution was diluted to the point when only one
or a few separate optoacoustic signals were detected.
Laser Optoacoustic Detection of Single Erythrocytes
Nd-YAG
laser
=532 nm
red
blood
cell
piezo-
transducer
transient
acoustic
wave
transient
acoustic
wave signal
to
scope a
-20
-15
-10
-5
0
5
10
0 0.5 1 1.5 2 2.5 3 3.5
012345
Optoacoustic Signal, mV
Time, µs
Depth, mm
laser-induced
acoustic wave
in single red
blood cells
b
Figure 2. Schematic diagram of the experiment (a) and optoacoustic signals detected in strongly diluted solution of blood (b).
Single read blood cells were detected with single laser pulse and optoacoustic detection.
Interpretation of the experimental result is as follows. Solution of blood contained EDTA that eliminated aggregation of
erythrocytes. The red blood cells were freely moving in solution driven by Brownian motion. At the moment nanosecond
pulse propagation through the solution, only a single erythrocyte or just a few erythrocytes were covered by the area of laser
beam. Therefore, only a single optoacoustic signal (black trace on Fig. 2b) or a just a few separate optoacoustic signal
(orange trace in Fig. 2b) were detected with acoustic transducer.
This experiment demonstrated ultimate sensitivity of optoacoustic detection with a single optoacoustic transducer. Such an
exceptional sensitivity of the pulsed optoacoustic detection method can be explained by two factors. First, our acoustic
transducer had large piezoelectric element resulting in high electrical capacity of 1 nF and corresponding low thermoelectric
noise level of just 4 µV. Secondly, we discovered that erythrocytes having exceptionally high optoacoustic efficiency
(probably due to lipids in their membranes) serve as a perfect optoacoustic target. The amplitude of optoacoustic signals in
erythrocytes is about 1.7 times higher than that in aqueous solution of a dye with similar absorbance, which can be explained
by the corresponding difference of thermoacoustic efficiency. Thus, red blood cells may be used a natural contrast agent for
the optoacoustic imaging. Neither optical tomography nor ultrasound imaging can utilize thermoelastic properties of blood
constituents, which represents yet one more advantage of the optoacoustic tomography as compared to other imaging
modalities.
0
0
1
10
100
1000
10000
400 600 800 1000 1200
Wavelength in nm
Absorption Coefficient, 1/cm
Oxyhemoglobin (HbO
2
)
Hemoglobin (Hb)
Water (H
2
O)
Figure 3. Absorption spectra of water, oxygenated and deoxygenated blood from visible to near-IR spectral range.
Optical properties of blood are well established (see for example [53]). Deoxygenated hemoglobin possesses a local
maximum of about 10-cm-1 at 760-nm followed by the range of relatively constant value and then sharply decreases to 1-cm-1
at 1064-nm. In contrast, oxyhemoglobin has minimum at about 650-nm and then its absorbance gradually increases to
approximately 8-cm-1 at 1064-nm. Absorption spectrum of blood presented in Fig. 3 demonstrates that optoacoustic
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measurements might be employed for accurate quantitative measurements of blood oxygenation. Optoacoustic images based
on high-resolution measurement of profiles of absorbed optical energy distribution in tissue can be employed for obtaining
quantitative information about the level of blood oxygenation in breast tumor microcirculation.
3.2. Phantoms with blood vessels
In order to demonstrate the utility of optoacoustic imaging for quantitative monitoring of hemoglobin oxygenation in
tumor angiogenesis, we performed pilot experiments of gel phantoms with embedded artificial “arteries and veins”. As
shown in Figure 4, a artificial blood vessel (polyethylene pipe shaped as a loop) was placed at substantial depth of 60 mm in
the phantom with optical absorption and scattering properties mimicking the breast tissues. Position of the transducer array
was at the bottom of the phantom, so that the field of view of the array overlapped with the blood vessel loop. A single fiber
was used for irradiation at only one location on the phantom surface as depicted on the top of Fig.4. Phantom surface
position is visible as white dash with the conus of optical energy spreading into the depth. This optoacoustic image has
relatively dark background with some lighter structures due to the phantom heterogeneity. Optoacoustic contrast (~500%) on
the image in Fig. 4 corresponds to the difference in background optical absorption and that in blood at λ=1064 nm.
Figure 4. Optoacoustic image of artificial “blood vessel” loop in a milky phantom. Optical properties of this phantom were
similar to the optical properties of the breast with real arteries at the wavelength of 1064 nm.
An exemplary two-dimensional optoacoustic image of a blood vessel loop filled with sheep blood having different level of
oxygenation is shown in Fig. 5a. Whole milk was diluted by water to obtain the effective optical attenuation coefficient 1.2
cm-1 in order to prepare the gel phantom shown in Fig. 5b with optical scattering mimicking the upper limit of that in the
breast tissue in the near-infrared spectral range [54]. Optical absorption coefficients of blood in blood vessels at the
wavelength of laser illumination of 757 nm varied from 0.8 to 4.0 per cm depending on oxygen content. In contrast, gel
simulating normal breast tissue had absorption of 0.09.13 cm-1, generating background signal with exponential trend ~exp-
µeff•z in optoacoustic images. The image reconstruction procedure includes optoacoustic signal integrals filtered through a
low frequency filter in order to remove the exponential trend associated with effective optical attenuation. Due to the limited
number of acoustic detectors in our system (32) and due to the fact that illumination usually performed at one site on the
surface, some artifacts associated with uneven optical field within the irradiated medium may be observed. Nevertheless, the
final image contains quantitative information on optical energy absorbed in tumors and blood vessels, which was verified
with analysis of image brightness relative to theoretically expected values.
Both loops of blood vessels were visualized and resolved and their shape and dimensions were depicted accurately when
compared with their relative position, shape and dimensions shown on the drawing Fig. 5b, and their depths within the
phantom were in good agreement with their real locations. The images of blood vessel cross-sections resemble circular
shapes. Note that our image reconstruction algorithm allows accurate reproduction of spherical objects and cylindrical
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objects. Thus, we may predict that objects with arbitrary shapes can be correctly reproduced on LOIS images. There is a
noise enhancement near the acoustic detectors due to 1/r-normalizing of spherical acoustic wave amplitude in detected
signals, visible at the bottom of Fig.5a. We are developing a new low frequency filtration procedure and plan its
implementation in the course of the proposed project, in order to completely eliminate all possible image artifacts, such as
thermoelectric noise artificially enhanced near the acoustic detectors and heterogeneity of optical fluence within the irradiated
breast.
-33 -17 0 17 33
0
10
20
30
40
50
60
70
80
90
Width, mm
Depth, mm
0.0 0.3 0.5 0.8 1.0
Contrast Gradient
Single Fiber, 757 nm
a
37 %
O2Hb
92 %
O2Hb
Laser Irradiation
Array of 16 transducers
757 nm
28 %
O2Hb
Field of
Illumination
Gel Phantom
b
Figure 5. Schematic diagram (a) and two-dimensional optoacoustic image of two blood vessel loops in a gel phantom with
optical properties similar to those of breast tissue in vivo in the near infrared spectral range.
The brightness of the blood vessel images was proprtional to the amount of absorbed optical energy, which in turn was
proportional to the level of blood oxygenation in each vessel. At the wavelength of 757-nm absorption coefficient in fully
oxygenated hemoglobin is about 4 times lower than the absorption of deoxyhemoglobin. Therefore, in Fig. 5a the highest
optoacoustic brightness can be seen in the upper part of the right loop filled with 28% oxygeneted blood. At the given
incident illumination fluence, the lower part of the vessel on the right filled with 92% oxygenated blood are poorly visible.
All vessels displayed in Fig. 5a were placed within the imaging plane. Therefore, their brightness correctly represents what
expected from blood oxygenation level. This imaging experiment demonstrated that the sensitivity of LOIS is sufficient for
detection of variation of blood oxygenation level with high resolution in optically strongly scattering and absorbing phantoms
with diameter similar to the real breast.
3.3. Optoacoustic imaging of breast cancer in vivo.
Eligible patients were chosen from the group of patients either with breast cancer diagnosed with x-ray mammography or a
combination of x-ray mammography and other imaging modalities (ultrasound, MRI) and biopsy or with breast tumors
suspected to be cancerous. Two-dimensional optoacoustic images were acquired in several locations on the breast. Two
images at two laser wavelengths of 1064-nm and 757-nm were acquired one after the other. Each image required irradiation
with 16 nanosecond pulses performed during 0.8 sec and collection of the imaging information on the 32-element array of
acoustic transducers. Image reconstruction procedure took about 1 sec on a 700 MHz PC. Laser pulses were delivered at one
site of the breast surface, and the optoacoustic transducer array was placed on the opposite side of the breast approximately
beneath the point of irradiation. Irradiation point was placed approximately in the center of the imaging field suspicious of
having a tumor. The exact location of tumors was not known prior to optoacoustic imaging procedure, however, approximate
location could be determined from the mammogram. Most tumors were not palpable and located several cm deep within the
breast.
Exemplary optoacoustic images of breast ductal-lobular carcinoma and benign fibroadenoma obtained prior to biopsy are
presented in Fig. 6 and Fig. 7 respectively. Optoacoustic images depict dimensions, shape, location and structure of tumor
angiogenesis network of microscopic blood vessels with enhanced optical absorption. However, optical scattering and
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thermoelastic properties of the tumors could also contribute to the optoacoustic contrast. Location of the tumors inside the
breast and the tumor dimensions can be accurately determined from the optoacoustic images. However, direct comparison
with x-ray mammography or ultrasound images based on tissue density is not appropriate. Resolution of image visualization
with LOIS (~1-mm in depth and 1.5-mm in lateral dimension) was comparable with that of x-ray mammography and
ultrasound.
Figure 6. Two-dimensional optoacoustic in vivo images of a ductal-lobular carcinoma in the left breast (LB). Laser pulses
were delivered to a single point on the skin surface above the tumor. Location of the piezoelectric detector array was at the
bottom of the imaging field presented. Contrast between a tumor and surrounding normal tissue strongly depends on the laser
wavelength. At the wavelength of 1064-nm the image brightness is significantly lower that the brightness at the wavelength
of 757-nm, indicating tumor with reduced level of blood oxygenation.
Figure 7. Two-dimensional optoacoustic in vivo images of a benign fibroadenoma in the right breast (RB). Laser pulses
were delivered to a single point on the skin surface above the tumor. Location of the piezoelectric detector array
was at the bottom of the imaging field presented. Contrast between a tumor and surrounding normal tissue
strongly depends on the laser wavelength. At the wavelength of 1064-nm the image brightness is significantly
higher that the brightness at the wavelength of 757-nm, indicating tumor with normall level of blood oxygenation.
Two main conclusions may be drawn from clinical experiments on breast cancer patients. The first is that optoacoustic
tomography provides substantially enhanced contrast between normal tissues and tumors. The second is that the optoacoustic
contrast correlates with the level oxygen saturation of hemoglobin in the tumor microcirculation. The tissue contrast was
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naturally variable, however, no tumors examined had contrast less than 2 fold relative to background in normal glandular or
fatty tissue. Optoacoustic contrast in the breast tumors being angiogenesis dependent can serve as means for noninvasive
diagnosis of malignancy. A correlation between microvessel density in tumor angiogenesis (i.e. total concentration of
hemoglobin) and brightness of tumors on optoacoustic images is currently being studied using immunohistology.
4. DISCUSSION
4.1. Diagnostic utility of LOIS
We report the first case of the laser optoacoustic imaging system (LOIS-2) application in imaging a patient with
tumors detected with mammography but not diagnosed with biopsy. The optoacoustic part of the in vivo clinical study has
been implemented as described below. Two-dimensional in vivo optoacoustic images have been obtained and analyzed in
order to determine the level of oxygen saturation in the tumor microvasculature. The total optical absorption at a given
wavelength, λ, of laser irradiation in tumor angiogenesis can be described by the following expression:
µHb [Hb]+µ
HbO 2[HbO2] +µ
H2O[H2O]=µ
a(1)
where the total absorption, µa, is composed of three absorption coefficients of (deoxy)hemoglobin, oxyhemoglobin and
water, (the brackets indicate concentration of the corresponding molecule).
Assuming concentration of water is approximately equal in all tumors, one may generate a diagnostic curve based the value
of absorbed optical energy at two wavelengths measured from optoacoustic images. Each detected tumor can be assigned a
point on the plot of absorbed optical energy at the wavelength of 757-nm vs absorbed optical energy at the wavelength of
1064-nm as shown in Fig. 8. All measured points should be located below the line indicating maximum possible absorbance
determined as absorbance of the whole blood. Based on quantitative measurement of blood oxygenation level in detected
tumors utilizing two colors of laser illumination, malignant tumors should be located in the right lower corner of the plot
where detected masses possess high concentration of deoxyhemoglobin, benign tumors should be located in the upper left
corner where detected masses possess high concentration of oxyhemoglobin, and normal tissues should have low absorbance
at both wavelengths. A more sophisticated analysis of the two-color optoacoustic images may be designed at the time when
statistically significant data will be collected.
0
1
2
3
4
5
6
7
024681012
Line of maximum absorption
Malignant
Benign
Normal
Absorption at 1064-nm, 1/cm
Absorption at 757-nm, 1/cm
Figure 8. Tumor absorption at 757-nm vs tumor absorption at 1064-nm
The clinical experiment demonstrated potential capability of the laser optoacoustic imaging to detect and diagnose breast
tumors in vivo. Differential diagnosis of the tumors based on the contrast in optical properties of the tumors measured at two
laser wavelengths in the near IR spectral range was confirmed with biopsy performed after the optoacoustic imaging
procedure. Optimal characteristics and specifications of LOIS technical components: laser parameters, light delivery fibers,
detector array, electronics and software for image reconstruction have been improved as a result of this study.
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4.2. Technical characteristics of LOIS
Real-time imaging capability of LOIS is very important for enabling capability of functional imaging with high spatial
resolution. The laser pulse repetition rate of 20 Hz and parallel multichannel data acquisition provides close to real-time
imaging capability. Therefore acquisition and averaging of 16 signals from each transducer will take less than a second. Data
transmission time was minimal using with Enhanced Parallel Portport. Currently, less than 2 seconds are needed for data
processing, back-projected image formation and image filtration using a personal computer with a 700 MHz processor. This
will result in close-to-real-time image acquisition and processing in our clinical imaging system.
Buttons on the front panel of LOIS computer interface allow acquisition of two separate images, processing and conditioning
of each image and further operations with the two images such as expansion of a portion of each image and additional post-
processing. Differential image and division of images also may be produced. Current design of LOIS allows operator to
position the acoustic detector array manually through the breast and to obtain images of chosen parts of the breast.
Five major characteristics of the laser optoacoustic tomography that make it attractive for diagnostic imaging of breast cancer
are: (1) high detection sensitivity which results from efficient generation of acoustic sources in tumors by short laser pulses
(efficiency of ~70-80%) and sensitive detection of acoustic waves with piezoelectric transducers, (2) high in-depth resolution
of ~500 µm, which results from short-pulse laser excitation and fast time-resolved detection (ultrawide-band ultrasound
detection), (3) high (sub-mm) lateral resolution that results from geometry of the detection array, (4) high optical contrast (up
to 6 fold) between tumors and normal tissues which most probably results from enhanced concentration of blood in malignant
tissues due to angiogenesis and correctly chosen laser wavelength to penetrate tissue and target tumors, and (5) substantial
depth of imaging without noticeable loss of resolution (up to 7-cm), which results from an optimal laser excitation
wavelength and efficient conversion of laser energy into ultrasonic waves propagating in soft tissues with insignificant
attenuation and minimal distortion.
The two-color optoacoustic tomography method will provide noninvasive diagnosis being inexpensive compared with
existing biopsy methods. This combination makes the new method commercially competitive. Our preliminary study
compared total (fixed and variable) facility costs in hospitals, ambulatory surgery centers and imaging centers for minimally
invasive vacuum-assisted biopsy (VAB) and wire-localized open surgical biopsy (OSB) vs anticipated costs of the proposed
optoacoustic technology. The study found that in each setting the overall cost savings of OAT vs VAB would be about $500
and that for OAT vs OSB would be about $1000 depending on exact procedure employed and facility type. Recent
breakthroughs in all-solid-state laser technology and miniaturization and integration of multifunction electronic components
allow a portable and compact design of the commercial laser optoacoustic imaging system.
5. CONCLUSIONS
Four years ago our group began basic research and development of the breast cancer detection system supported by the
National Cancer Institute and two years ago received additional support from the US Army Breast Cancer Research Program
[55-57]. Presently, LOIS is being evaluated at the UTMB hospital in patients with breast cancer. Future advancements of
LOIS are envisioned in the direction of developing diagnostic utility and improved imaging parameters (sensitivity and
resolution), accommodating the needs of continuous monitoring for ambulatory subjects, including those with implants, and
in producing displays that intelligently combine structural, chemical and functional information. Emphasis continues to be on
maximizing accuracy of quantitative information, on minimizing image acquisition and processing time, costs, eliminating
patient discomfort, and improving easiness of image interpretation. The estimated manufacturing cost of the laser
optoacoustic imaging system will not exceed the cost of an ultrasound imaging machine. An affordable price of LOIS and its
portability will permit wide application of the optoacoustic technology in diagnostic clinics. The application of low-level
non-ionizing near-infra red radiation will allow continuous, safe, and frequent screening. The noninvasive diagnostics is
painless and much less expensive compared with biopsy. In addition, LOIS can be used for image-guided biopsy and image-
guided therapy, especially for monitoring the effect of chemotherapy with anti-angiogenesis drugs. This hybrid technology
employs the most useful features of optical spectroscopy and ultrasonic imaging: nonionizing near-infrared radiation (to
achieve maximum tissue contrast) and low-noise piezoelectric transducers for ultrawide-band ultrasonic detection (to
visualize both, masses of abnormal tissue and fine details of tumor microcirculation). The diagnostic capability of LOIS to
differentiate noninvasively breast carcinoma, fibroadenoma and cyst will soon be tested in statistically significant number of
patients. LOIS promises to become the first real-time, high-resolution imaging system that ustilizes tissue contrast based on
tumor angiogenesis.
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6. ACKNOWLEDGMENTS
This work was supported by the National Cancer Institute (grant #R29-CA80221), DOD Breast Cancer Research Program,
US Army (grant #DAMD17-99-1-9404) and Fogarty International Center of NIH (FIRCA grant R03-TW05776).
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