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Sonoluminescence tomography of

turbid media

Lihong V. Wang, Qimin Shen

Lihong V. Wang, Qimin Shen, "Sonoluminescence tomography of turbid

media," Proc. SPIE 3597, Optical Tomography and Spectroscopy of Tissue III,

(15 July 1999); doi: 10.1117/12.356828

Event: BiOS '99 International Biomedical Optics Symposium, 1999, San Jose,

CA, United States

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Sonoluminescence Tomography of Turbid Media

Lihong V. Wang and Qimin Shen

Optical Imaging Laboratory

Biomedical Engineering Program

Texas A&M University

College Station, Texas 77843-3120, USA

ABSTRACT

A novel optical imaging technique was developed for noninvasive cross-sectional imaging of tissue-like turbid media. By use

of a sonoluminescence signal generated internally in the media by continuous-wave ultrasound, two-dimensional images

were produced for objects embedded in turbid media by raster scanning the media. Multiple objects of different shapes were

resolved using this imaging technique. The images showed a high contrast and good spatial resolution. The spatial resolution

was limited by the focal size of the ultrasonic focus.

Keywords: Sonoluminescence, tomography, scattering media, ultrasound, tissue optics

1. INTRODUCTION

Optical imaging, also known as optical tomography, in strongly scattering media, has become an active research field

because of its advantages of noninvasion, nonionization, and functional contrast for biomedical diagnosis.1'2 Several optical

imaging techniques being investigated include time-resolved optical imaging, frequency-domain optical imaging, optical

coherence tomography, optoacoustic tomography and ultrasound-modulated optical (acousto-optical) tomography. In these

approaches, time-resolved and frequency-domain techniques have achieved a comparable resolution of millimeters. Optical

coherence tomography has achieved 1O-im resolution but is limited to a penetration of <2 millimeters into biological tissues.

Optoacoustic tomography and acousto-optical tomography have achieved millimeter resolution and have potential to image

thick biological tissues. The image contrast is based on the difference in optical properties between abnormal and the

surrounding normal biological tissues. All these optical approaches use an external light source, mostly a laser.

Because biological tissues are optically turbid media, light is quickly diffused in tissues due to strong scattering.

Light transmitted through tissues consists of three types: ballistic light, quasi-ballistic light, and diffuse light. Ballistic light

travels straight through tissue with no experience of scattering by the tissue and hence carries direct imaging information as

x-ray does. Quasi-ballistic light experiences minimal scattering in the forward

direction and carries some imaging

information. Diffuse light follows tortuous paths, carries little direct imaging information and overshadows ballistic or quasi-

ballistic light. For a 5-cm-thick biological tissue with the assumed absorption coefficient jia

0.1

cm and reduced scattering

coefficient t'

= 10

cm', the ballistic light and quasi-ballistic light does not exist for practical purposes.3 Therefore, diffuse

light is the only carrier of imaging information for thick biological tissues. All optical tomography for thick biological tissues

must overcome the light scattering problem to obtain optical images.

Ultrasonic generation of light in a medium, known as sonoluminescence (SL), was first discovered in 1934. The

initial observations were multiple-bubble sonoluminescence (MBSL). SL has attracted an extraordinary amount of attention

in this decade since single-bubble sonoluminescence (SBSL) was reported in l99O.' Although the full explanation of SL is

still in development, it is well known that light is emitted when tiny bubbles driven by ultrasound collapse. The bubbles start

out with a radius of several microns and expand to —50 microns due to a decrease in acoustic pressure in the negative half

a sinusoidal period. After the sound wave reaches the positive half of the period, the situation rapidly changes.

The resulting

pressure difference leads to a rapid collapse of the bubbles accompanied by a

broadband emission of light —

sonoluminescence.

This process repeats with each cycle of sound.'3 The flash time of SL was measured to be in the tens of

picoseconds.5 SBSL is so bright that it can be seen by the naked eye even in a lighted room, whereas MBSL is visible only in

a darkened room.'0 The spectrum of SL contains molecular emission bands associated with the liquid, mostly water,

in which

the sonoluminescence occurs.'° A typical spectrum of SL is a broadband emission with peaks near 300-500

nm.'3 There are

also spectral peaks reported at 590 nm, 670 nm and 770 nm while alkali-metal salt of Na, Li and K were dissolved in

water, respectively.6 From the spectrum of SL in water, the local temperature within the cavities was

estimated on the order

Part

of the SPIE Conference on Optical Tomography and Spectroscopy of Tissue Ill

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San Jose, California • January 1999 SPIE Vol. 3597 • 0277-786X/991$1O.OO

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of 5000 K.'° Researchers have envisaged possible applications of SL in sonofusion, sonochemistry, and building ultrafast

lasers using the ultrafast flash of light in SL, which is the only means of generating picosecond flashes of light besides a

laser.

We report here a comprehensive study of a novel application of sonoluminescence in optical imaging:

sonoluminescent tomography (SLT) of dense turbid

SLT, minimally scattering ultrasound that generates an

internal light source in the media is used to image optically scattering media. SLT contains information not available in the

traditional ultrasonography. The major advantages of SLT include: 1) high signal-noise ratio (SNR) due to the internally

generated SL signal; 2) high contrast of imaging; 3) good spatial resolution, which is limited by the ultrasonic focal size; and

4) low cost of equipment. The cost of equipment to generate SL is as low as hundreds of U.S. dollars.'5 This paper gives a

full account of our investigation while our previous Letter demonstrated the concept of

METHODS AND MATERIALS

An Intralipid phantom was prepared by mixing 8 ml of dominantly scattering Intralipid (Pharmacia Inc., 20%) and 3.25x108

mol of dominantly absorbing Trypan Blue dye (Sigma, T5526) in 360 ml of distilled water. The reduced scattering

coefficient t' and the absorption coefficient ta were respecüvely 6.15 cm and 0.014 cm at the wavelength of 584 nm,

which is the absorption peak of Trypan Blue dye. The reduced scattering coefficient t' is equal to x (1 —

g),

where

the scattering coefficient defined as the probability of scattering per unit infinitesimal path length, and g is the scattering

anisotropy defined as the average cosine of the single-scattering deflection angle. Since the spectrum of SL is broadband,'6

the optical scattering properties of the scattering media were also given at the wavelength of 400 nm, corresponding to the

spectral peaks of the SL'7 and the maximum-sensitivity spectral range of the optical detector in our experiment. The p.s' and

the Intralipid phantom at 400 nm were 8.5 cnf' and 0.002 cm', respectively.'8

The turbid solution was contained in a 400-mI fused-quartz beaker (Quartz Scientific Inc., QBKLO400) and was

held on an x—y translation stage (Figure 1). An objects made of rubber was buried in the phantom. An ultrasonic transducer

(Panametrics, V314-SU) with a focal length of 3.68 cm, a focal diameter of 0.3 cm, and a focal zone of 3.44 cm transmitted

vertically an ultrasonic wave into the scattering medium. The ultrasonic transducer was driven by an amplified 1-MHz

sinusoidal signal from a function generator (Stanford Research System, DS345). The amplification was achieved by use of a

power amplifier (Mini-Circuits, TIA-1000-1R8) and a transformer. The height of the ultrasonic transducer was adjusted such

that the focal zone of the ultrasonic wave enclosed the buried object in the vertical direction. The height between the middle

plane of the object and the bottom of the beaker was 4.5 cm. The motorized translation stage, controlled by a personal

computer (PC), was able to scan along both the x and the y axes, which formed an x—y plane perpendicular to the ultrasonic

axis. The SL signal was detected by a photomultiplier tube (PMT) (Hamamatsu, R928) beneath the beaker, and then was

differentially amplified by a low-noise preamplifier (Stanford Research System, SR560). The amplified signal was a dc

voltage representing a time-averaged SL intensity with a time constant of —10 ms for the detection system. The amplified

signal was recorded by a digital oscilloscope (Tektronix, TDS 640A) and was subsequently acquired by the PC through a

GPIB interface (National Instruments, PCI-GPIB).

Figure 1. Schematic diagram of the experimental setup for sonoluminescence imaging.

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While raster scanning the beaker in the x-v plane with a step size of 1 mm. the PC recorded the dc signals of SL

versus the values of x and v.

The

optical and ultrasonic systems were fixed while the beaker was scanned. Two-dimensional

images of the objects buried in the scattering media were plotted with the acquired data.

3. RESULTS AND DISCUSSION

We previously reported the SL column in a clear solution that was imaged with a CCD camera [Figure 2(a)].'4 We further

modeled the distribution of acoustic pressure underneath the ultrasonic transducer using the following equation.

p(r,t)=C5

P(s)exp[I(wtT2]dS

where

p(r,

t) is

the pressure as a function of the observation point r

and

time t,

C is

a constant, p(s)

the pressure on the

surface of the ultrasonic transducer, oi

the angular frequency of the ultrasonic wave, 4

the phase delay from a point on

the transducer surface and the point of observation, and d

the distance between a point on the transducer surface and the

point of observation. The integration is over the surface of the ultrasonic transducer. The focal length in the calculated sound

field matched that specified by the manufacturer [Figure 2(h)1. The sound column strongly correlated with the SL column.

0.0 0.1 0.2 0.3 0.4 0.5 0.6

Acoustic Pressure (a.u.)

(b)

Figure 2. (a) Sonoluminescence column measured with a CCD camera. (b) Modeled sound column of the ultrasonic

transducer.

A rubber cube in the Intralipid phantom was imaged by SLT (Figure 3). The spatial resolution of the edges was

estimated to be 2-3 mm. and an excellent imaging contrast was observed.

Horizontal Axis (cm)

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x (mm)

SL Intensity (V)

Figure

3. Two-dimensional sonoluminescence

image

of the object in the Intralipid phantom.

The contrast of the SL images was based on the difference between the optical and ultrasonic properties of the

objects and those of the surrounding medium. These objects were optically opaque and ultrasonically absorbing. When the

object was moved toward the ultrasonic focus, the SL intensity dropped quickly for several reasons. First, the ultrasonic field

was reduced below the focus because of the slight acoustic attenuation of the rubber object. Second, the object yielded no SL

signal. Third, the SL signal above the object was partially blocked by the object and hence was more difficult to reach the

PMT.

To observe the spatial resolution for distinguishing multiple objects, we buried two cubic objects in the Intralipid

phantom [Figure 4(a)I. The distance between the two objects was x.

which

was varied from 1 mm to 6 mm. One-

dimensional SL images across the centers of the two objects were obtained for various distances x

[Figure

4(b)]. From these

figures, we observed the spatial resolution of 2-3 mm. which was similar to that observed in the single-object SL images. The

spatial resolution was limited by the focal size of the ultrasonic transducer, which was 3 mm.

To observe the spatial resolution for revealing the shapes of buried objects. we buried two objects of different shapes

in the Intralipid phantom: a square and a triangular object [Figure 5(a)]. A two-dimensional SL image of the two objects was

acquired [Figure 5(b)]. The 2-mm separation between the tip of the triangular object and the square object was barely

resolved in the SL image. The hypotenuse of the triangular object looked zigzagged because of the I-mm step size in the

raster scanning dunng the data acquisition. However, both objects were clearly imaged with the correct shapes and sizes.

0 2

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Object 1

Object 2

_f fx

8mm

6mm

(a)

Ax = 6 mn

Ax=5mrr

Ax =4 mn

Ax=3mn

Ax = 2 mn

= 1 mn

ià"i 'Ô '"6 4b

x(mm)

(b)

Figure

4. (a) Schematic diagram of two rubber objects buried in the Intralipid phantom with a separation Ax; (b) One-

dimensional sonoluminescence images of the two objects along the x axis when the separation Ax was varied from 1 mm to 6

mm.

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25:

10:

0- -

20-

10-

Figure 5. (a) Schematic diagram of two rubber objects with different shapes buried in the Intralipid phantom; (b) Two-

dimensional sonoluminescence image of the two differently shaped objects buried in the Intralipid phantom.

Although the objects used in this experiment had both optical and ultrasonic contrast relative to the background

medium, images of objects may be obtained using SLT based on several contrast mechanisms in general. First, when an

object has ultrasonic contrast relative to the background, the SL signal originating from the object will differ from that

onginating from the background medium. The SL generation is affected by the local ultrasonic intensity. Second, when an

object has contrast in optical properties, the SL signal from the object will be attenuated differently because the SL light must

propagate through the object. Third, when an object has contrast in ability to generate SL, the SL signal from the object will

be different even if the local ultrasonic pressure is the same.

Using the present ultrasonic system, we obtained a SL column of —3.5 cm in length and —0.3 cm in diameter [Figure

2(a)1.'4 The length of the SL column limits the imaging resolution along the ultrasonic axis. Similarly, the diameter of the SL

column determines the imaging resolution on the x-v plane. A more tightly focused ultrasonic transducer may be used to

reduce the size of the SL column significantly. When the SL column is reduced to a desired size, one may acquire three-

dimensional images of scattering media by scanning in all three directions.

SL light propagates outward in the scattering medium in all directions. We may improve the signal-noise ratio of the

detection system dramatically by integrating the SL light over a large detection area or by a light collection system similar to

an integrating sphere. Because all the SL light is useful for imaging, integrating the SL signal would allow an increased

imaging depth as well.

Although there are potentially harmful effects caused by cavitation, the threshold of ultrasound intensity leading to

volume lesions is very high. The damage threshold in spatial-peak-temporal-peak (SPTP) power was reported to be 400

WIcm2 and 900 W/cm2 at 1 MHz for dog brain tissue and dog thigh muscle, respectively.'9 The peak pressure in our

expenment was —2.0 bars at the ultrasonic focus, corresponding to an SPTP power of 1.3 WJcm2, which was two orders of

magnitude less than the damage threshold. The peak pressure was also far less than the 23-bar safety limit set by the U.S.

Food and Drug Administration, which is usually conservative.20 SLT uses ultrasonic waves to drive pre-existing

microbubbles between 5 pm to —50 pm in size: hence, formation of new bubbles with ultrasound is not necessary. The

threshold acoustic power to generate SL through pre-existing bubbles is much less than that required to form new bubbles. It

349

8 mm

(a)

—,

5 10

x (mm)

0.0

0.5

1.0

1.5

SL Intensity (V)

(b)

2.0

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