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IEEE TRANSACTIONS ON MEDICAL IMAGING, VOL. 28, NO. 8, AUGUST 2009

A 3-D High-Frequency Array Based 16 Channel

Photoacoustic Microscopy System for

In Vivo

Micro-Vascular Imaging

Rachel Bitton*

, Member, IEEE

, Roger Zemp

, Member, IEEE

, Jesse Yen

, Member, IEEE

Lihong V. Wang

, Fellow, IEEE

, and K. Kirk Shung

, Fellow, IEEE

Abstract—

This paper discusses the design of a novel photoa-

coustic microscopy imaging system with promise for studying the

structure of tissue microvasculature for applications in visualizing

angiogenesis. A new 16 channel analog and digital high-frequency

array based photoacoustic microscopy system (PAM) was devel-

oped using an Nd:YLF pumped tunable dye laser, a 30 MHz piezo

composite linear array transducer, and a custom multichannel

receiver electronics system. Using offline delay and sum beam-

forming and beamsteering, phantom images were obtained from

m carbon fiber in water at a depth of 8 mm. The measured

lateral and axial spatial resolution of the system was

m and

m, respectively. The dynamic focusing

capability of the system was demonstrated by imaging a composite

carbon fiber matrix through a 12.5 mm imaging depth. Next, 2-D

in vivo

images were formed of vessels around 100

m in diameter

in the human hand. Three-dimensional

in vivo

images were also

formed of micro-vessels 3 mm below the surface of the skin in two

Sprague Dawley rats.

Index Terms—

High-frequency ultrasound, multichannel re-

ceiver electronics, photoacoustic imaging, transducer array.

I. I

NTRODUCTION

TUDIES in oncology have shown that angiogenesis, the

formation of new blood vessels within a tumor, or the

growth of new blood vessels between a tumor and surrounding

tissues, plays a critical role in tumor growth and metastasis of

cancer [1]–[3]. Tumors need to be supplied by blood vessels,

delivering oxygen and nutrients while removing metabolic

waste in order to propagate. While the formation of new

micro-vessels (vessels smaller than one millimeter) can be part

of normal development and wound healing, it is also a key

Manuscript received April 01, 2008; revised September 25, 2008. First pub-

lished January 06, 2009; current version published July 29, 2009. This work was

supported in part by the National Institute of Health under Grant R01 EB000712

and Grant P41-EB2182.

Asterisk indicates corresponding author

*R. Bitton was with University of Southern California, Los Angeles, CA

90089 USA. She is now with the Department of Radiology, Stanford Univer-

sity, Stanford, CA 94305 USA(e-mail: rbitton@stanford.edu).

R. Zemp was with Washington University, St. Louis, MO 63130 USA. He is

now with the Department of Electrical and Computer Engineering, University

of Alberta, Edmonton, AB, T6G 2V4 Canada (e-mail: zemp@ece.ualberta.ca).

J. Yen and K. K. Shung are with the Department of Biomedical Engineering,

University of Southern California, Los Angeles, CA 90089 USA (e-mail:

jesseyen@usc.edu; kkshung@usc.edu).

L. V. Wang is with the Department of Biomedical Engineering, Washington

University, St. Louis, MO 63130 USA (e-mail: lhwang@biomed.wustl.edu).

Color versions of one or more of the figures in this paper are available online

at http://ieeexplore.ieee.org.

Digital Object Identifier 10.1109/TMI.2008.2011899

initial step in tumor progression, since tumor cells induce an-

giogenesis [4]. A high-resolution imaging technology capable

of visualizing micro-vessels would lend the ability to identify

part of the early angiogenic process.

The development of novel approaches to biomedical imaging

is stimulated by the manifest need for high-speed, high-resolu-

tion noninvasive techniques. Laser induced photoacoustic mi-

croscopy is an imaging modality based on the intrinsic optical

properties of biological tissue and ultrasonic detection at high

frequencies (

). Photoacoustic imaging uses short

laser pulses that are absorbed in the tissue to cause acoustic

pressure transients, which are detected with an ultrasonic trans-

ducer. Two factors provide exceptional motivation for the de-

velopment of photoacoustic methods as a diagnostic tool in vas-

cular imaging, the strong intrinsic optical absorption of blood,

and the resolution per image depth of ultrasound.

Because of the strong scattering of light in biological tissue,

optical imaging methods such as confocal microscopy and

optical coherence tomography (OCT) suffer from degraded

spatial resolution with increased depth. Confocal microscopy

and OCT are limited to resolutions between 1–2

mata

0.5 mm image depth, and around 16

m at a 2.5 mm image

depth, respectively. [5]–[8]. In ultrasound, the scattering of

energy is several orders of magnitude weaker than optical scat-

tering. Therefore, energy can penetrate deeper into the tissue,

providing submillimeter resolution at greater imaging depths

compared with the optical limit [9]. For example, ultrasound

backscatter microscopy (UBM) systems are capable of imaging

at depths of 20 mm with a 115

m spatial resolution, and

4 mm image depths with 50

m spatial resolution [10], [11].

Nevertheless, ultrasonic imaging suffers from reduced contrast

because the detection of the backscattered signals is based on

the differences of the acoustic properties in biological tissue.

This presents unique position for photoacoustics, juxtaposed

between two basis imaging modalities, ultrasonic and optical

imaging. Photoacoustic imaging brings optical based contrast

into the ultrasonic depth imaging range, attempting to capi-

talize on the strengths of both optical and acoustic techniques.

Inherent to the technology are capabilities for functional infor-

mation to be extracted [12]. By characterizing the spectrum of

different optical absorbers and irradiating the sample at multiple

wavelengths, photoacoustic experiments have distinguished

between oxygenated and deoxygenated hemoglobin [13].

The draw to image at higher frequencies is palpable; spa-

tial resolution improves with increased frequency. Much of

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: A 3-D HIGH-FREQUENCY ARRAY BASED 16 CHANNEL PHOTOACOUSTIC MICROSCOPY SYSTEM

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the work in high-frequency photoacoustic imaging has been

generally confined to mechanical scans with single element

transducers [14]–[17]. Methods of image reconstruction include

a backprojection approach or simple time of flight measure-

ments, as in traditional ultrasound. Backprojection techniques

are similar in concept to X-ray computed tomography (CT)

and positron emission tomography (PET) scans. Various algo-

rithms have been derived for the inverse solution to the wave

equation, based on diffraction optics, to form the reconstructed

image [18], [19]. This method has been used to image mouse

brain tissue in photoacoustic computed tomography (PAT)

[20]. In this approach, a tomographic circular scan is per-

formed using an unfocused wideband single element transducer

which is scanned mechanically. High-resolution PAT circular

scan images have been produced using offline backprojection

reconstruction [21]. Backprojection reconstruction has also

produced some trail artifacts inherent to the technique when

imaging small dark absorbers such as hair fibers in tissue

phantoms [17]. Unlike PAT, backward mode detection uses a

focused transducer. High frequency photoacoustic images of

rat micro-vessels

in vivo

have been produced by using a single

element transducer, backward mode detection, and the synthetic

aperture method. In one experiment, a 50 MHz, single element

transducer with a large numerical aperture was scanned linearly

across a single axis [18], [22]. There are two disadvantages to

this technique; a large numerical aperture will cost imaging

depth due to beam divergence past the focal zone, and the long

scan time needed for mechanical scanning of the transducer.

Experiments

in vivo

in situ

are affected by the passage of

time. The condition of the live animal will change over time,

and in some cases the animal may expire before the scans are

complete. This is certainly a problem for functional imaging as

oxygenation and blood perfusion vary after expiration.

The advantage of photoacoustic microscopy with transducer

arrays over single element photoacoustic tomography is the

potential to image in real time with higher frame rates. Trans-

ducer array technology facilitates fast acquisition times, as

well as electronic steering and focusing of the receive beam.

Arrays provide parallel photoacoustic signal detection through

adjacent transducer elements which each posses their own

electrical connections. Although array technology is common

in frequencies below 20 MHz, high-frequency transducer ar-

rays pose challenges both in the fabrication process, perfor-

mance demands, and parallel multichannel electronic system

design [23]. The frequency of the photoacoustic echoes that

are generated are dependent on the size of the vessel or ab-

sorber, causing the bandwidth performance of the transducer

to be of particular importance. For high-frequency arrays,

the complexity of high speed electronic design increases as

channel number increases. Proper noise suppression, isolation,

sampling, and critical timing restraints set the boundaries for

high frequency photoacoustic systems. Until now, most of the

work in photoacoustic imaging has been executed using com-

mercially available components either in single element trans-

ducer/single channel systems using oscilloscopes, or using

lower frequency arrays with commercial ultrasound systems

[16], [17], [24]. Our aim was to create a novel multichannel

high-resolution photoacoustic microscopy system to visualize

micro-vascular structures in rats. In this paper, we present a

prototype 16 channel receiver system which was developed to

accommodate a custom made high frequency transducer array.

The photoacoustic microscopy receiver system uses multi-

channel parallel signal processing, in backward mode detection,

to acquire the raw data from each element in the transducer

array. This method is especially useful and differs from current

high-frequency ultrasound linear array systems in which active

channels are beamformed via hardware, followed by transfer of

the beamformed data to a PC [25]. Transmitting and capturing

the raw radio-frequency (RF) data from all active channels pro-

vides the distinctive ability to access RF data from each indi-

vidual element, allowing the most flexibility for beamforming,

and image reconstruction.

II. T

HEORY

Photoacoustic imaging has its physical basis in a phenom-

enon called the optoacoustic, or photoacoustic effect. This ef-

fect can be observed in a variety of media, including biological

tissue, wherever pulsed electromagnetic energy can be absorbed

[24]. Under conditions of thermal and stress confinement, short

laser pulses will induce acoustic waves most efficiently. As inci-

dent laser light in the visible spectrum interacts with tissue, it is

either absorbed by local chromophores, or scattered in various

extents depending on the inherent optical properties of the tissue

[27]. The photoacoustic effect occurs when the pulsed light en-

ergy is absorbed locally in biological tissue, and a small rapid

temperature rise in the medium causes thermoelastic expansion.

This expansion produces pressure transients, which propagate

as acoustic waves throughout the tissue omni-directionally [28].

The initial pressure generated,

, is related to the spatial por-

tion of heating function

at position

, and the Grüneisen

parameter (

). Initial pressure can then be written

(1)

In this form,

, where

is the local fluence

in (

), determined by the incident light as well as the scat-

tering and absorption parameters

and

, respectively [29].

Subsequent pressure generated at given time and position incor-

porates the initial pressure and obeys the common form of the

time retarded wave equation. Thus, the pressure profile of the

generated photoacoustic echo is based on the optical properties

of the target in the media, and does not necessarily mimic the

profile of the laser pulse itself.

Tissue absorbs light differently at specific wavelengths, con-

ditional to the optical properties of the medium. Photoacoustic

imaging is especially well suited for vascular imaging since in

the visible spectrum, light absorption in subdermal tissue is prin-

cipally due to the dominant chromophores oxy- and deoxy-he-

moglobin [30].

As a result of the high optical scattering of light in tissue,

photoacoustic imaging does not focus the transmit (light) beam

as in ultrasonic imaging techniques. Rather, photoacoustic

imaging utilizes the multiple scattering effects to its advantage.

While optical based methods rely on the signals of only singly

backscattered photons [8], which effect image speckle and limit

penetration depth, the cumulative effect of multiple scattering

aids to a better perfusion of energy, and maximizes irradiation

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IEEE TRANSACTIONS ON MEDICAL IMAGING, VOL. 28, NO. 8, AUGUST 2009

Fig. 1. PAM system setup. The laser is triggered and coupled by a lens to an

optical fiber to illuminate the sample. Photoacoustic waves are received through

48 elements in the transducer array. For each laser event, the system provides

the receive front end, filtering and amplification stages, and transfers raw data

from 16 elements of the array. The laser fires 3 times in order to collect the

photoacoustic data from all 48 elements in the array.

homogeneity of the tissue. Photoacoustic signals are generated

from the region of interest within the tissue. Consequently, they

are subject to only one-way, rather than round-trip, ultrasonic

image quality degrading aberrations. Additionally, the echo

wait time is approximately half that of ultrasound. For single

wavelength excitation, it is possible that a multichannel pho-

toacoustic imaging system with the proper parameters would

be capable of constructing

in vivo

images in real time with only

one nanosecond duration laser pulse.

The axial and lateral resolution of a photoacoustic image is

determined by the generated acoustic echo and transducer prop-

erties. The beam width of the active transducers in the array

scan determines the theoretical lateral spatial resolution of the

system. It can be related proportionally to the transducer wave-

length by

(2)

where the

number,

, is the ratio of focal distance to aperture

dimension.

III. M

ATERIALS AND

ETHODS

A. System Design

The photoacoustic microscopy system is comprised of three

main components: An Nd:YLF laser source used to irradiate the

tissue and induce photoacoustic waves, a 48 element piezo com-

posite transducer array which receives the photoacoustic waves,

and a custom 16 channel parallel receive electronic system. The

photoacoustic electronic system controls the laser operations,

and processes and digitizes the data from the transducer array

(Fig. 1). In the experimental setup both the transducer array and

optical fiber are mounted on a 3-D translation stage, and then

lowered into a water tank. The process initiates when a TTL

signal from the photoacoustic system motherboard triggers the

laser pulse to irradiate the sample through the optical fiber posi-

tioned obliquely to the sample and array. After each laser pulse,

the 16 element subaperture is processed through the analog and

Fig. 2. PAM receive system architecture. The transducer array elements all

receive preamplification before being multiplexed. The channels then pass

through filtering, variable, and fixed gain stages before A/D conversion. Eight

channel cards accommodating two channels each send digital data to the

computer via the PCI bus.

digital boards, and then stored to the PC. To fully sample the 48

element array, the laser fires three times for each image acquisi-

tion. Algorithms were written for the PAM system in Labview

and Matlab. They provide a user interface, and delay and sum

photoacoustic image reconstruction.

The laser setup includes a diode-pumped Nd:YLF

-switched laser (INNOSLAB Edgewave, Germany) and a

tunable dye laser (Cobra, Sirah Laser, Germany). The

-switch

is an externally triggered attenuator which modulates the

factor of the optical resonator cavity, producing 6.5 ns pulses

at 14 mJ when deactivated. The dye laser is used to tune the

light to 598 nm, and produced 2 mJ per pulse. The light is then

coupled by a beam shaper and microscope objective into a 600

m optical fiber which delivered a per pulse energy of 0.8 mJ.

Since the area of illumination on the skin surface was about

4 mm, the energy fluence was estimated as 10

(

The linear transducer array used is constructed of a 2–2 piezo

composite material measuring

mm in the azimuth

and elevation directions, respectively. It contains 48 rectangular

elements centered at 30 MHz, and is used in receive mode only.

The array elements bear a

(100

m) pitch, and use a lens to

form an elevational focal depth of 8 mm [29]. The simulated one

way

fractional bandwidth (full-width half-maximum) of

the array is 70%. The receive electronics were developed to in-

clude signal processing and data transfer in two separate stages;

analog and digital (Fig. 2). All 48 elements are connected to

receive circuits including preamplification. They are then mul-

tiplexed down to create 16 active channels which pass through

filtering, fixed and variable gain stages. The 16 analog channels

are then converted to 16 digital channels using eight channel

boards which digitize two channels each and store in temporary

memory before transfer to the computer.

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: A 3-D HIGH-FREQUENCY ARRAY BASED 16 CHANNEL PHOTOACOUSTIC MICROSCOPY SYSTEM

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Fig. 3. Photograph of the fabricated a) 16 channel analog receiver board im-

plemented on a four-layer PCB and b) digital PAM system with motherboard

and channelcards implemented on a six-layer PCB.

The first stage is comprised of the front end receive and signal

processing electronics (Blocks 1 and 2, Fig. 2). The array ele-

ments interface with the PAM system through RG-174 shielded

RF cables and SMA connectors [Fig. 3(a)]. On the front end

receive system; each of the 48 elements contains a fixed, ultra

low noise, 18 dB, preamplification stage (MAX4107, Maxim/

Dallas Semiconductor). Four-to-one multiplexers are used to

select between elements (AD8184, Analog Devices), forming

16 active channels. Following channel formation, a fourth-order

Butterworth band pass filter is used to remove spurious signals

with frequencies out of the desired transducer response. The

filter was designed to have a wide response and sharp cutoff

so that it can also double as an anti-aliasing filter before analog

to digital conversion stages. The signal is then amplified by low

noise, dual variable gain amplifiers (VGA) providing a 0–40 dB

gain range (AD8332, Analog Devices). The differential out-

puts of the VGA are converted back to single ended signals

through transformers (T1 6T, Minicircuits), which also provide

signal isolation. A final fixed gain stage boosts the signal another

20 dB. The measured system receiver gain provides a selectable

range from 33–73 dB.

The second stage of the receiver system digitizes, provides

the timing network for the system, and controls data transfer

through a motherboard-channelboard scheme (similar to the de-

sign within a PC) (Block 3, Fig. 2). The digital system supplies

the master triggering, providing clock synchronized triggers for

the laser, the receive electronics, and the PC data transfer card.

The laser is triggered first at a 5-Hz repetition rate. After a pro-

grammable delay, the acquisition is triggered to account for the

latency between the laser trigger and the optical delivery. The

16 channel group raw data is acquired in less than 11 ms after

each laser pulse. Then, handshaking control and clock signals

are provided to transfer data using a PCI based digital NI-6534

card (National Instruments). Each complete image frame re-

quired three laser shots.

The high-frequency array imposed challenges on electronic

component performance and on system design. High-speed

considerations for multichannel systems such as bus topology,

impedance matching, and clock synchronization influenced

both schematic design and manual board layout. To ensure

proper timing synchronization of multiple channels, all control

and clock signals are based on divisions of a single 100-MHz

oscillator. The digital system contains nine channelboards

that plug into a motherboard through high speed connectors

(

-Series, Samtec) [Fig. 3(b)]. Each channelboard receives

two analog channels, and contains two 8-bit, 100 MHz,

analog-to-digital converters (AD9054, Analog Devices), one

16-bit temporary memory storage FIFO with 1 K depth, pro-

viding an imaging depth of about 15 mm (SN74V225, Texas

Instruments). Each FIFO is shared between two channels, and

a line driver on each channelboard selects a particular board for

transfer. The first eight channelboards digitize each group of 16

elements from the array, while the ninth channel is dedicated

to laser energy measurements, provided by a photodiode. The

motherboard provides the 16-bit data bus; interconnect to

channelboards and to the PC, as well as the clock distribution

network. Using the 16-bit NI-6534 running at 12.5 MHz,

digital data is transferred to the PC at a rate of 25 MB/s for

each channelboard.

B. Image Reconstruction

Images are created using offline delay and sum beamforming

and beamsteering. To calculate a single image A-scan line the

RF data is used from all 48 elements, rather than the 16 element

subaperture. This establishes a narrow beam focus and thus, a

more desirable lateral spatial resolution. To form the beam and

steer the beam, delays are applied to each element using a simple

geometric model. Assuming a distance

to the target point at

a lateral distance

and angle

from the element in question

(relative to the center element reference), the time delay

for element number

can be written as

(3)

where

is the sound velocity through the medium. This is iden-

tical to conventional ultrasound time delay calculations except

that the one way trip distance for received echoes is accounted

for applying

During beamformation, coherence factor weighting was ap-

plied to the phantom image data. This technique aids to reduce

focusing errors resulting from sound velocity inhomogeneities,

as well as steering errors. The coherence factor equation is given

(4)

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IEEE TRANSACTIONS ON MEDICAL IMAGING, VOL. 28, NO. 8, AUGUST 2009

Fig. 4. PAM receiver frequency response and 30 MHz transducer array fre-

quency response.

where

is the number of elements and

is the channel data for a given scan line angle

, after the time

delays for steering and focusing have been applied [32].

C. Experimental Setup

A spectrum analyzer was used to illustrate the frequency re-

sponse of the entire receive system. To demonstrate function-

ality and to characterize system performance, phantom images

were obtained. The phantom consisted of a 6

m carbon fiber

in water, imaged at the transducer array focal depth of 8 mm.

In vivo

rat images were obtained using the same experimental

setup as the phantoms for two different rats. A Sprague Dawley

rat was prepared by depilating a section on the back to reduce

excessive signal loss from the fur and then fixing the animal po-

sition below the water tank. Small subcutaneous vessels were

imaged at the transducer focus (6–10 mm), and a few millime-

ters below the surface of the skin. After imaging, the animal was

sacrificed, and the area imaged was excised for comparison and

verification of vessel structure.

IV. R

ESULTS

A. System Characterization

To assess the bandwidth of the assembled receiver system, the

frequency response was measured between the front end (trans-

ducer element input stage) and the final RF output stage using

a spectrum analyzer (E4401B, Agilent). The passband of the

measured receiver response encompasses the

one way

transducer response and additionally displays a wider band per-

formance (Fig. 4). The transducer array bandwidth dictates the

bandwidth for following stages. However, the PAM receiver al-

lows flexibility if the system is later paired with a high frequency

array of greater fractional bandwidth, which is currently under

development. A 1.2 dB ripple in the measured frequency re-

sponse is observed due filter trade off relationships that exist

between bandwidth, roll-off, and complexity.

To characterize the sensitivity of the front end, the minimum

detectable signal was measured. A function generator providing

a 30 MHz sine wave was connected to a variable gain attenuator

Fig. 5. Line spread function of PAM 6

m carbon fiber phantom constructed

from a

-scan projected on the

-axis (distance [m]).

and then fed into the front end of the receive electronics. The

signal was measured at two subsequent test points, and the min-

imum detectable signal was calculated given the known input

signal and attenuation. Test point 1 was measured using an os-

cilloscope at the output of the analog signal processing board.

Test point 2 was measured after the digital boards, using the RF

signal at the PC display stage, to quantify the effect of digital

switching noise to the system. The minimum detectable signal

was 316 and 500

, approximating the noise floor at test points

1 and 2, respectively. Using the KLM model, the minimum de-

tectable signal of 500

at 30 MHz corresponds to an esti-

mated minimum detectable transducer pressure of

Spatial resolution information was extracted by imaging a

single 6

m carbon fiber in water and projecting the image data

onto one axis to construct the line spread function, showing the

signal strength versus spatial distance (Fig. 5). Both axial and

lateral resolutions were based on the

width of the carbon

fiber in each image direction. The lateral and axial spatial reso-

lution of the system at the transducer focal point was measured

m and

m, respectively. We have previously

reported an axial resolution of 25

m with a similar system [33].

In that study, instead of using the

width, the axial reso-

lution was measured by superimposing the signal from a single

carbon fiber with the signal from the same fiber translated verti-

cally, and examining the envelope. The separation distance be-

tween two distinguishable peaks was used as the axial resolution

figure of merit, 25

To express the dynamic focusing capability of PAM, a com-

posite image was constructed by imaging the 6

m carbon fiber

in evenly spaced positions. The images from each carbon fiber

represented a matrix of target locations and were combined to-

gether to create a composite image. The carbon fiber matrix is

made up of 5 rows with a 1 mm separation in depth by nine

columns with 1.1016 mm lateral separations (Fig. 6). To ob-

tain this data set, the optical fiber is fixed relative to a single

carbon fiber, while the transducer position is moved. This en-

sures a more uniform illumination of the target regardless of the

matrix position. The data was taken with 598 nm wavelength

light in water. The two dead elements of the array may explain

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: A 3-D HIGH-FREQUENCY ARRAY BASED 16 CHANNEL PHOTOACOUSTIC MICROSCOPY SYSTEM

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Fig. 6. PAM phantom image composite of a 5

9 matrix of 6

m carbon fibers

in water. Image displayed with no averaging and 33 dB dynamic range.

Fig. 7. PAM system

in vivo

image of a cross section of blood vessels in the

lower portion of a human hand.

a slight variation in sensitivity (upper right corner); however, it

is shown that the targets are well focused throughout the region

of interest.

B. Images

In vivo

images

-scans were obtained from the lower portion

of a human hand using a wavelength of 568 nm and a laser flu-

ence of 7

. A number of bright signals from microves-

sels less than 100

m in diameter can be seen in the center of

the image (Fig. 7).

This progression led to 3-D photoacoustic rat microvessel im-

ages

in vivo

. Spaced at 0.005 in (0.127 mm) intervals in the ele-

vation direction (perpendicular to the

-scan plane), 100

-scans

were acquired along the image plane using a three axis transla-

tion stage for the transducer array. With 10

incident flu-

ence at 598 nm and an averaging index of 16, subcutaneous ves-

sels were imaged at depths of 3 mm below the skin’s surface (in

the 8 mm focal zone) in two Sprague Dawley rats (Fig. 8). Offline

image reconstruction was executed in

. Because of the

dense vasculature within the tissue, the images shown are recon-

structed from a small truncated portion of the data in the depth

direction, in order visualize the overlayed vessel structures at a

given depth. Microvessels of different diameters and vessel bi-

furcation can be identified at varying depths within the truncated

range. These vessels were not visible from the skin’s surface. The

animals were then sacrificed and portions of skin excised to

verify vessel structures.

V. D

ISCUSSION

The noise visible in the 3-D image could be related to the

3-D reconstruction static thresholding techniques were used to

form the boundaries of the vessel structures as 3-D surfaces.

A more robust approach might include a dynamic thresholding

technique which could refine the surface boundary threshold for

areas outside of the focal zone of the transducer. Additionally,

system noise contributions can be related to the lack of shielding

of the electronic system. The RF cables used for each transducer

element carry small signals that travel from the array to the con-

nector box, and then to the front end of the board. These ca-

bles are bundled closely and can contribute to signal loss and

potential crosstalk. Future array designs could integrate lines

into a single shielded RF connector which plugs directly into

the printed circuit board (PCB).

The greatest improvement in signal-to-noise ratio (SNR)

may be made in optimization of the light delivery technique.

These experiments were conducted with no more than half the

laser fluence limit (20

). A more efficient light delivery

scheme could improve the amount of laser energy delivered to

the tissue, yielding greater photoacoustic signal strength. Other

approaches to photoacoustic sensitivity enhancement are being

investigated by the introduction of exogenous contrast agents,

such as gold nanoshells [34].

In Fig. 8(c), the upper vessel could not be compared to the

photo since that portion of the tissue was not recovered during

the skin excising process. Some vessels are also disconnected

by a missing slice which may have been caused by missing

data sets, motion artifacts, or could possibly be improved by

decreasing the slice intervals at which the scans are taken. It is

also important to mention that vessels of larger diameters will

produce photoacoustic echoes of lower frequency compared

with those of smaller diameters. The large vessel low frequency

echoes may be difficult to resolve if they are outside of the

transducer bandwidth.

A future development for the system is to image in real time

by increasing the pulse repetition frequency of the laser (up to

1 KHz), increasing the acquisition transfer speed through hard-

ware, and increasing the speed of the processor which executes

beamformation and running the image reconstruction algorithm

in C programming, rather than Matlab. The limitation of time

between laser shots (analogous to pulse repetition frequency,

PRF) is the execution of the Labview program. The Labview

program which controls the FIFO to PC transfer creates a data

transfer bottleneck as it is not based on a real-time platform, the

timing is nondeterministic, and may produce PRF fluctuations in

the ms error margin. Although the diode-pumped Nd:YLF laser

permits triggering at arbitrary rates unlike flashlamp-pumped

lasers, a faster deterministic PRF could provide more uniform

energy deposition between laser pulses. The transfer speed con-

trol timing issue can be overcome through hardware by pro-

viding control signals from a dedicated processor or CPLD. This

would alleviate the problem, since both the sampling of RF data

(100 MHz) and each 16 channel PC transfer (12.5 MHz) on this

PAM system are fast enough for 1 KHz pulse repetition rates

and real time imaging. This would eliminate the Labview bar-

rier, delivering completely deterministic control between laser

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IEEE TRANSACTIONS ON MEDICAL IMAGING, VOL. 28, NO. 8, AUGUST 2009

Fig. 8. Three-dimensional PAM images of micro-vessels below the surface of the skin in two Sprague Dawley rats graphed on axis

, representing lateral,

scan direction, and depth dimensions, respectively. a) Rat1 PAM image showing mircro-vessels with corresponding photo of excised skin b) alternate

viewing

angle for Rat1 image, and c) Rat2 PAM image with corresponding photograph of excised skin. Markers

, and

denote micro-vessels within the tissue.

events and provide data fast enough for 16 channel real time

PAM imaging.

VI. C

ONCLUSION

A new 30 MHz array based photoacoustic imaging system

was developed in which raw RF photoacoustic data is acces-

sible for all 48 elements and images were formed after three

laser pulses. The 16 channel photoacoustic receive system was

fabricated and characterized, then used to create phantom im-

ages of 6

m carbon fibers in water. The

axial and lateral

spatial resolution of the system was measured as

m and

m, respectively. The dynamic focusing capability was

demonstrated through a 12.5 mm depth using a composite image

of a carbon fiber matrix. Two-dimensional

in vivo

images were

formed of micro-vessel structures in the human hand. Three-di-

mensional

in vivo

images were also formed of micro-vessels

below the surface of the skin in two Sprague Dawley rats.

Photoacoustic microscopy is a nonionizing modality with

much room for growth, holding great promise for medical

imaging. Some of the most attractive qualities include a scalable

resolution and imaging depth dependent upon the ultrasonic

transducer frequency, no speckle artifacts, high contrast, and

the extension towards real time imaging. To compensate for the

limitations of light scattering in deeper tissues and to provide im-

ages based on the acoustic and optical properties of the medium,

both conventional ultrasound images and photoacoustic images

can be used for comparison,

and to provide two different types

of contrast. A bimodal imaging system could produce a simul-

taneous display of photoacoustic and pure ultrasonic images

acquired from the same cross sections of tissue, supplying an

anatomical reference for the photoacoustic signals. Furthermore,

these images could be constructed by utilizing the same trans-

ducer array for both modalities, offering added diagnostic value

without the need for two separate systems.

BITTON

et al.

: A 3-D HIGH-FREQUENCY ARRAY BASED 16 CHANNEL PHOTOACOUSTIC MICROSCOPY SYSTEM

1197

The high-frequency PAM array system presented in this paper

offers a new perspective on the capabilities of photoacoustic

imaging. High-resolution photoacoustic images can have appli-

cations in oncology by tracking the important vascular struc-

tures associated with angiogenesis.

CKNOWLEDGMENT

The authors would like to thank J. M. Cannata for his work

in the design and fabrication of the high-frequency transducer

array.

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