Photoacoustic imaging or optoacoustic imaging is a biomedical imaging modality based on the photoacoustic effect. Non-ionizing laser pulses are delivered into biological tissues and part of the energy will be absorbed and converted into heat, leading to transient thermoelastic expansion and thus wideband (i.e., megahertz-order bandwidth) ultrasonic emission. The generated ultrasonic waves are detected by ultrasonic transducers and then analyzed to produce images. It is known that optical absorption is closely associated with physiological properties, such as hemoglobin concentration and oxygen saturation. As a result, the magnitude of the ultrasonic emission (i.e. photoacoustic signal), which is proportional to the local energy deposition, reveals physiologically specific optical absorption contrast. 2D or 3D images of the targeted areas can then be formed.
Biomedical imaging
The optical absorption in biological tissues can be due to endogenous molecules such as hemoglobin or melanin, or exogenously delivered contrast agents. As an example, Fig. 2 shows the optical absorption spectra of oxygenated hemoglobin (HbO2) and deoxygenated hemoglobin (Hb) in the visible and near infrared region. Since blood usually has orders of magnitude higher absorption than surrounding tissues, there is sufficient endogenous contrast for photoacoustic imaging to visualize blood vessels. Recent studies have shown that photoacoustic imaging can be used in vivo for tumor angiogenesis monitoring, blood oxygenation mapping, functional brain imaging, skin melanoma detection, methemoglobin measuring, etc.
Two types of photoacoustic imaging systems, photoacoustic/thermoacoustic computed tomography (also known as photoacoustic/thermoacoustic tomography, i.e., PAT/TAT) and photoacoustic microscopy (PAM), have been developed. A typical PAT system uses an unfocused ultrasound detector to acquire the photoacoustic signals, and the image is reconstructed by inversely solving the photoacoustic equations. A PAM system, on the other hand, uses a spherically focused ultrasound detector with 2D point-by-point scanning, and requires no reconstruction algorithm.
Photoacoustic computed tomography
General equation Given the heating function H ( r , t ) {\displaystyle H({\boldsymbol {r}},t)} , the generation and propagation of photoacoustic wave pressure p ( r , t ) {\displaystyle p({\boldsymbol {r}},t)} in an acoustically homogeneous inviscid medium is governed by
∇ 2 p ( r , t ) − 1 v s 2 ∂ 2 ∂ t 2 p ( r , t ) = − β C p ∂ ∂ t H ( r , t ) ( 1 ) , {\displaystyle \nabla ^{2}p({\boldsymbol {r}},t)-{\frac {1}{v_{s}^{2}}}{\frac {\partial ^{2}}{\partial {t^{2}}}}p({\boldsymbol {r}},t)=-{\frac {\beta }{C_{p}}}{\frac {\partial }{\partial t}}H({\boldsymbol {r}},t)\qquad \qquad \quad \quad (1),}
where v s {\displaystyle v_{s}} is the speed of sound in medium, β {\displaystyle \beta } is the thermal expansion coefficient, and C p {\displaystyle C_{p}} is the specific heat capacity at constant pressure. Eq. (1) holds under thermal confinement to ensure that heat conduction is negligible during the laser pulse excitation. The thermal confinement occurs when the laser pulsewidth is much shorter than the thermal relaxation time. The forward solution of Eq. (1) is given by
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