1 Introduction

To look inside the human body, invasive techniques such as surgery canpotentially damage the body. Therefore, noninvasive biomedical imaging modal-ities, such as projection radiography, x-ray computed tomography (CT), positronemission tomography (PET), single-photon emission computed tomography(SPECT), magnetic resonance imaging (MRI), and ultrasound (US), are widelyused in clinics today for the diagnosis and treatment of various diseases. Someof these modalities (projection radiography, x-ray CT, PET, and SPECT) useionizing radiation, which exposes the human body to harmful radiation. On theother hand, much safer MRI is time consuming, expensive, and cannot be usedfor patients with metallic implants as it uses high-strength magnetic fields. USimaging offers a cheap, real-time, easier to operate, and safe alternative.However, poor soft tissue contrast and difficulty in imaging through skull/bonelimit its use in many body parts. Nonetheless, all these imaging modalities arecapable of imaging deeper (∼20 cm inside soft tissues in the case of US imaging),which is a major advantage for clinical applications.

Conventional optical microscopy provides images with high spatial resolutionand rich optical contrast, making it possible to visualize cellular and subcellularstructures. However, due to light scattering in biological tissue, imaging depth islimited to not more than ∼1 mm. Therefore, even though optical microscopy isa gold standard in pathology, it is not suitable for in vivo human imaging.Another high-resolution purely optical imaging modality, known as optical coher-ence tomography (OCT), can image tissue with high resolution (∼1 μm) and at animaging depth of up to ∼2 mm. However, OCT also has limited application to eyeretinal imaging to date. Therefore, researchers are trying to come up with varioustechniques to overcome these depth limitations of optical imaging. Diffuse opticaltomography provides imaging depth up to a few centimeters deep inside tissue,but the spatial resolution is relatively low—one-third of the imaging depth.Therefore, high-spatial-resolution images at deeper imaging depths are always achallenge for purely optical imaging techniques.

Over the past two decades, photoacoustic imaging (PAI) has gained signifi-cant attention from the research community. It is a hybrid, nonionizing modalitycombining the best of optical and US imaging. This hybrid modality provides richoptical contrast images with high US spatial resolution even several centimetersdeep inside the tissue. In this Spotlight, we discuss in detail various PAI tech-niques and highlight the importance of pulsed laser diodes (PLDs) for PAI. ThisSpotlight is divided into four sections. In Section 1, we introduce the PAI, itsbasic principle, photoacoustic (PA) signal detection, and then we discuss variousways of implementing the PAI, followed by resolution versus imaging depth lim-itations and imaging speed limits. Then we talk about several light sources beingused and how PA images are formed. Section 2 focuses on PA computed tomog-raphy/photoacoustic tomography (PACT/PAT) systems and their instrumentation

Kalva and Pramanik: Photoacoustic Tomography 1

For example: If d ¼ 10 μm, using typical soft tissue properties, we obtain the

thermal relaxation time ¼ τth ¼ d2

αth¼ ð10× 10−4 cmÞ2

1.3× 10−3 cm2∕s ¼ 7.69 × 10−4 s, and stress

relaxation time ¼ τs ¼ dvs¼ 10× 10−4 cm

0.15 cm∕μs ¼ 6.67 ns. Therefore, a pulsed excitation

width shorter than 6.67 ns will satisfy the above-mentioned conditions and willefficiently generate PA waves.

When the laser pulse duration is shorter than both the thermal relaxation timeand stress relaxation time, the laser excitation is said to be in thermal and stressconfinements. In that case, we can neglect the fractional volume change.Therefore, initial pressure rise (p0) can be written in terms of the local temperaturerise ΔT (in the order of few millikelvin) as

p0 ¼ β × ΔT∕k, (1)

where β is the thermal expansion coefficient, and k is the isothermal compressibil-ity. The local temperature rise can be expressed as

ΔT ¼ ηthAe

ρCv, (2)

where ηth denotes the percentage of light being converted into heat, Ae denotes thespecific optical energy absorption, Cv denotes the specific heat capacity at constantvolume, and ρ is the mass density. After substituting Eq. (2) into Eq. (1) and defin-ing a dimensionless Gruneisen parameter Γ ¼ β

kρCv, we can express Eq. (1) as

p0 ¼ ΓηthAe, (3)

where Ae is proportional to the local optical fluence F. Then, Eq. (3) becomes

p0 ¼ ΓηthμaF, (4)

where μa is the optical absorption coefficient. Although Γ and ηth are dependent ontype of the tissue, they are usually considered to be constants.10 Hence, the initialpressure rise p0 is proportional to μa and F.

For example: If the laser fluence is 20 mJ/cm2, assuming μa ¼ 0.1 cm−1 and

ηth ¼ 1, the temperature rise ¼ ΔT ¼ ηthAeρCv

¼ 1× 0.1cm−1 × 20mJ∕cm2

1 g∕cm3 × 4 Jg−1K−1 ¼ 0.5mK, and the

initial pressure rise will be ¼ p0 ¼ ΓηthAe ¼ 0.2 × 1 × 0.1 cm−1 × 20mJ∕cm2 ¼4mbar (1 J/cm3= 10 bar). Typically, each millidegree rise (1 mK) in temperature willproduce 8 mbar or 800 Pa of pressure rise.

1.3 Photoacoustic signal detection

Upon laser excitation, there is the generation of the initial pressure p0 [Eq. (4)],then an acoustic wave starts to propagate within the tissue at the speed of

Kalva and Pramanik: Photoacoustic Tomography4

object (circular scanning) using one SUT. For each laser pulse excitation, a time-resolved PA signal (also known as an A-line) can be acquired (similar to anA-scan in US imaging). There are two methods of collecting the A-lines in a fullcircle using one SUT: (i) continuous scanning method, and (ii) stop-and-go scan-ning method. In the continuous scanning method, the SUT is moved around theimaging object at a constant speed and multiple A-lines (one A-line per laserpulse) are collected throughout the motion of the transducer. The SUT stops onceit completes a full 360-deg rotation, whereas in the stop-and-go scanning method,the SUT rotates in steps of predefined length. After each step, the SUT stops andcollects the required number of A-lines and then moves to the next step. Whilethe transducer is in motion, there will be no A-line acquisition. This process con-tinues until the SUT completes one full 360-deg rotation. Due to such differencesin scanning methods, the scan time for stop-and-go scan is much higher comparedto continuous scan to collect same number of A-lines.14 This is because in con-tinuous scan, the imaging speed is dependent on the number of A-lines collected,whereas in stop-and-go scan, the imaging speed is limited by the motor to move

Figure 1 Optical illumination and acoustic detection configurations for (a) SUT-PAT system,(b) CAT-PAT system, (c) LAT-PAT system, and (d) VAT-PAT system. (e) Schematic pro-jected view of VAT showing the individual piezoelectric elements and cylindrical cavity.(d) Redrawn from Ref. 12 and (e) reprinted with permission from Ref. 13.

Kalva and Pramanik: Photoacoustic Tomography 7

1.7 Photoacoustic image formation

Photoacoustic imaging provides mainly absorption contrast, so we can recover theabsorption coefficient μa distribution inside the tissue. However, we cannotdirectly get μa distribution. First, we recover the initial local pressure rise p0 insidethe tissue. From Eq. (4), the absorption coefficient μa can then be calculated pro-vided we know the local optical fluence F. In practice, the local fluence inside thetissue is not very easy to compute (or known). However, we can assume that theadjacent fluence F in the tissue is usually similar, but μa varies significantly. Forexample, blood absorbs light strongly compared to the surrounding tissues in thevisible wavelength. Thus, if we assume F to be locally homogeneous, then p0

can be used to directly map the relative absorption coefficient μa. Once the pres-sures at the tissue boundaries pðr0!, tÞ are measured, we can recover the originalp0 distribution inside the tissue in two ways: reconstruction-based image forma-tion and focused-scanning image formation. Reconstruction-based image forma-tion is used in PACT/PAT, whereas focused-scanning image formation is usedin PAM.

1.7.1 Image reconstruction in photoacoustic computed tomographyThe objective of image reconstruction in PAT is to estimate the initial pressurerise p0ð~rÞ inside the tissue from a set of measured acoustic/PA signals pðr0!, tÞ attissue boundaries. There are many methods of reconstructing the initial pressurefrom the boundary measurements, e.g., simple backprojection, frequency-domainreconstruction, system matrix-based reconstruction, etc.60–63

Here, we will take a close look into the simple backprojection method. Due toits simplicity, the use of an approximate modified backprojection algorithm in thetime domain is quite common.64,65 Using backprojection, the initial pressure risecan be obtained as follows:64

p0ð~rÞ ¼ZΩ0

bðr0!, t ¼ j~r − r0!j∕vsÞ

dΩ0

Ω0, (10)

where Ω0 is the solid angle subtended by the entire measurement surface S0 withrespect to the reconstruction point ~r inside S0. For planner geometry, Ω0 ¼ 2π andΩ0 ¼ 4π for spherical and cylindrical geometries. bðr0!, tÞ is the backprojectionterm. dΩ0 is the solid angle subtended by detection element dS0 with respect toreconstruction point at ~r. The term dΩ0∕Ω0 is a factor weighting the contributionto the reconstruction from the detection element dS0 (refer to Fig. 5). The back-projection term bðr0!, tÞ is given as

bðr0!, tÞ ¼ 2pðr0!, tÞ − 2ct∂pðr0!, tÞ

∂t: (11)

The reconstruction simply projects bðr0!, tÞ backward via a spherical surface(in our case, a circular arc, as our geometry is in two dimensions) centered at

Kalva and Pramanik: Photoacoustic Tomography16

PA simulated data generation code:% Generate simulated PAT scan data using k-wave

% Scan with single element ultrasound transducer (SUT)

% Full circular scanning, clockwise, Point target numerical phantom

% k-Wave Version 1.2.1 and Matlab R2017a

clear all; close all; clc;

%% parameters

medium.sound_speed = 1500; % sound speed in water [m/s]

outputfilename= ‘PAT_scan_data.txt’;

radius= 15e-3; % SUT radius in [m]

num_A_lines= 100; % number of A-lines

sensor.frequency_response = [2.25e6 70]; % 2.25 MHz center frequency SUT

% 70% nominal bandwidth

fs= 20; % sampling frequency in [MHz] or [MS/s]

Nsample= 400; % number of sample in each A-line data acquisition

object_sim.Nx= 341; % number of grid points in the x (row) direction

object_sim.Ny= 341; % number of grid points in the y (column) direction

object_sim.x= 34.1e-3; % total grid size [m]

object_sim.y= 34.1e-3; % total grid size [m]

% UST_pos = [0, 0]; % First SUT position (coordinates) [m]

%%

time.dt= 1/(fs*1e6); % sampling time in sec

time.length= Nsample; % number of points in time

time.t_array= 0:1:time.length-1; % time array of Nsample time steps

time.t_array= time.t_array*time.dt;

Nx= object_sim.Nx;

Ny= object_sim.Ny;

dx= object_sim.x/object_sim.Nx; % grid point spacing in the x direction

Figure 6 (a) Circular scanning geometry typically used for PAT. SUT, single-element ultra-sound transducer (detector). (b) Conventional delay-and-sum algorithm.

Kalva and Pramanik: Photoacoustic Tomography18

cross-sectional PAT images at various anatomical locations (in small animals) ofbrain, liver, kidneys, and bladder, respectively. Using this system, an imagingspeed of 1 frame in 1.6 s (0.625 Hz) was obtained. A state-of-the-art circulararray-based PAT system utilizes a parallel 512-channel DAQ system and providesa frame rate of 50 Hz (50 frames/s), i.e., we can acquire a single cross-sectionalPA data in 0.02 s.76 This high imaging speed is very helpful for visualizingshort-duration physiological processes. Because of the faster speed, one can alsoeasily acquire different plane images by moving the sample in the vertical direc-tion (z plane) and collect PAT data for different horizontal planes leading to full3-D imaging ability. Using this state-of-the art CAT-PAT system, label-freePAT imaging of a small-animal whole body from brain to trunk was achieved ata high-frame rate of 50 Hz.76 These CAT-PAT systems were also demonstratedfor breast cancer screening in females77 and imaging blood vasculature in humanextremities, such as arms, wrist, biceps, legs, and foot.78

Circular scanning geometry-based PAT systems (SUT-PAT and CAT-PAT)have limited usage due to their nonclinically relevant designs. They can be usedin small-animal whole-body imaging, human breast imaging, imaging of human

Figure 10 (a) Schematic of the CAT-PAT system. Reprinted with permission from Ref. 72.In vivo noninvasive cross-sectional PAT images of mouse acquired at anatomical locationsof (b) brain, (c) liver, (d) kidneys, and (e) bladder. EY, eyes; CV, cortical vessels; LV, liver;VC, vena cava; SC, spinal cord; GI, gastrointestinal tract; PV, portal vein; SP, spleen; KN,kidney; BM, backbone muscle; BL, bladder. Reprinted with permission from Ref. 75.

Kalva and Pramanik: Photoacoustic Tomography 25

be very difficult to operate by clinicians. Thus, for the PLD-PAT system to beacceptable to clinicians and effectively used for point-of-care diagnostics, there isa need for a handheld PA probe that can be used in a similar fashion as the hand-held US probe of clinical US machines. Hence, the clinical US probe needs to bemodified/redesigned to integrate the laser within the probe. But integrating theexisting PLD laser with the clinical US probe increases the overall size and weightof the probe, thereby making it difficult to operate normally. Hence, there is a needto miniaturize the PLD source to be able to fit within the handheld US probe with-out changing any of its dimensions considerably.

Steenbergen’s research group has designed a handheld PA and US (PA/US)probe in collaboration with industrial partners Quantel and Silios by integratinga PLD and a US transducer.114 This probe is used with a modified commercialportable US system for dual-modality imaging. A picture of this system is shownin Fig. 18(a). The PLD source is specifically developed by Quantel (Paris,France). It delivers laser pulses of ∼130-ns pulse duration at 805-nm wavelengthwith a PRR of 10 kHz and per-pulse energy of ∼0.56 mJ. But the PLD outputbeam suffers from a wide divergence angle compared to conventional lasers.Using cylindrical microlenses and diffractive optical elements provided by SiliosTechnologies (Peynier-Rousset, France), the beam was collimated and reshapedto reduce the divergence with 80% efficiency. The US probe (Esaote SL3323)consists of 128 elements, each about 5 mm in length and 0.245 mm in pitch.These array elements have a center frequency of 7.5 MHz and 100% bandwidthat −6 dB. Proper care was taken to shield the probe from electromagnetic noisegenerated by the laser driver. The schematic of the handheld probe can be seen

Figure 18 (a) Photograph of portable PA/US imaging system. (b) Schematic of handheldprobe. US, ultrasound array transducer; P, deflecting prism; CR, aluminum cooling rim;DOE, diffractive optical elements; MCL, microcylindrical lenses; DS, diode stack. PA/USreconstructed images of human PIP joint in (c) sagittal plane and (d) transverse plane.Reprinted with permission from Ref. 114.

Kalva and Pramanik: Photoacoustic Tomography 39