Photo acoustic tomography

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Lihong Wang (2014), Scholarpedia, 9(2):10278. doi:10.4249/scholarpedia.10278 revision #138868 [link to/cite this article]
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Curator: Lihong Wang

Figure 1: First 3D photoacoustic microscope. Reprint with permission (Zhang, Maslov et al. 2006).

Photoacoustic tomography (PAT), sometimes referred to as optoacoustic tomography, is defined as cross-sectional or three-dimensional (3D) imaging of a material based on the photoacoustic effect (Wang 2009). Therefore, PAT possesses spatial resolution along the depth dimension and at least one of the other two dimensions. In PAT, light is absorbed by biological tissue and converted to transient heating, which is subsequently converted into an ultrasonic wave due to thermoelastic expansion. Detection of the ultrasonic wave yields a tomographic image. Combining rich optical contrast and scalable ultrasonic resolution, PAT is the only imaging modality capable of providing multiscale high-resolution structural, functional, and molecular imaging of organelles, cells, tissues, and organs in vivo. While functional imaging measures physiological parameters, such as oxygenation and blood flow, molecular imaging senses biomarkers to identify specific cancer cells or detects gene expression products to track gene activations.


Concept of PAT

PAT involves optical excitation, ultrasonic detection, and image formation. A short-pulsed laser is usually used to produce ultrasound in biological tissue efficiently. The amplitude of the photoacoustic pressure depends on the optical energy deposition as well as the thermal and mechanical properties of the tissue. Because either unscattered or scattered photons can produce photoacoustic signals, photoacoustic waves can be generated deeply in biological tissue. Because the ultrasonic scattering coefficient in tissue is 2–3 orders of magnitude less than the optical counterpart, high spatial resolution can be achieved by detecting the photoacoustic waves. Consequently, PAT breaks through the optical diffusion limit (~1 mm in the skin) for high-resolution optical-contrast imaging. The image formation is essentially triangulation of the photoacoustic sources according to the time-of-flight signals recorded at multiple locations.

For a sufficiently short laser pulse, the initial pressure induced by the laser excitation is given by \[ \tag{1} p_0 = \frac{\beta T}{\kappa} \]

Here, \(\beta\) denotes the thermal coefficient of volume expansion; \(\kappa\) denotes the isothermal compressibility; \(T\) denotes the temperature rise. In soft biological tissue, each mK temperature rise yields approximately an 800 Pa pressure, which is above the noise level of a typically used ultrasonic transducer. PAT reconstructs the distribution of the initial pressure.

3D photoacoustic microscopy

Figure 2: 3D photoacoustic microscopic image of a human palm. (a) Photograph of the palm. Red box: area to be imaged by PAM. (b) En face view of the PAM image. (c) Sagittal view of the PAM image along the dashed line in 2b. Reprint with permission (Favazza et al. 2011).

PAT can be implemented as 3D photoacoustic microscopy (PAM) (Figure 1) (Maslov et al. 2005; Zhang, Maslov et al. 2006). An optical fiber is used to deliver the excitation laser light, which is monitored by a photodiode for energy calibration. A conical lens is used to form a donut-shaped beam so as to reduce surface photoacoustic signals. The beam is then weakly focused into the tissue while the ultrasonic transducer focuses coaxially into the same region for photoacoustic detection. The photoacoustic time-of-flight signal is recorded at each lateral location of the ultrasonic transducer. Multiplying the time-of-flight by the speed of sound yields a 1D depth-resolved image (A-line), where the focusing of the ultrasonic transducer provides the lateral resolution. Linear or raster scanning over the tissue produces 2D or 3D tomographic images.

The broadband ultrasonic detector has a numerical aperture of 0.44 and a center frequency of 50 MHz. As a result, the lateral resolution, determined by the focal diameter of the ultrasonic transducer at the center frequency, measured ~45 µm. The axial resolution, determined by the ultrasonic bandwidth, measured ~15 µm. The maximum imaging depth, determined by the ultrasonic attenuation in tissue, measured >3 mm. This system is marked as AR-PAM in Figure 4. A representative PAM image acquired using a 584 nm laser wavelength is shown in Figure 2, where selected vessels are labeled with arrows for reference between the two views. Key anatomical features of the skin are also marked.

Photoacoustic computed tomography

Figure 3: First functional photoacoustic images of cerebral hemodynamic changes in response to one-sided whisker stimulations in a small animal, acquired with the scalp and skull intact. Reprint with permission (Wang et al. 2003).

PAT can also be implemented as photoacoustic computed tomography (PACT). In PACT, an array of unfocused ultrasonic transducers detects photoacoustic waves, which are used to reconstruct a tomographic image through an inverse algorithm. In circular-view photoacoustic computed tomography, ultrasonic detection positions follow a ring while a pulsed laser beam is expanded to illuminate the tissue in the region of interest. Representative functional PACT images are shown in Figure 3. The colorscale shows the differential optical absorption between the on and off states of stimulation, whereas the grayscale shows the blood vessels. The ultrasonic transducer has a broad bandwidth centered at 3.5 MHz, yielding an in-plane resolution of ~0.2 mm. The out-of-plane resolution, determined by the cylindrical focusing of the transducer, is ~1 mm.

Photoacoustic contrasts

If all absorbed optical energy is converted into heat and nonthermal relaxation such as fluorescence is negligible, the initial pressure is given by \[ \tag{2} p_0 = \Gamma \mu_a F \]

where \(F\) denotes the optical fluence (J/m2), \(\mu_a\) denotes the optical absorption coefficient (m-1), and \(\Gamma\) denotes the Grueneisen parameter that measures the conversion efficiency from optical energy deposition to pressure (\(\Gamma \propto \frac{\beta}{\kappa}\)).

According to Eq. (2), PAT is exquisitely sensitive to optical absorption, the most sensitive one among all optical imaging modalities. The photoacoustic excitation converts a small change in optical absorption coefficient to an equal fractional change in ultrasound signal, tantamount to a relative sensitivity of 100%. In principle, any molecules are optically absorbing; it is a matter of locating the absorption band. So far, PAT has been used to image the following endogenous biological molecules:

  1. Oxyhemoglobin
  2. Deoxyhemoglobin
  3. Melanin
  4. Water
  5. Lipids
  6. DNA
  7. RNA

PAT has also imaged various exogenous contrasts such as organic dyes, products from reporter genes, and nanoparticles.

PAT is also sensitive to any change in the Grueneisen coefficient. In particular, the Grueneisen coefficient increases with the equilibrium temperature (~5% per Kelvin), which enables PAT to image temperature with a sensitivity of the order of 0.1 K for monitoring of thermal therapy.

Like any other waves, photoacoustic waves are subject to the Doppler effect, which enables imaging of blood flow. In fact, PAT is the only modality that measures all the endogenous physically parameters—including the concentrations of oxyhemoglobin, the concentrations of deoxyhemoglobin, the diameters of blood vessels, the blood flow velocities, and the volume of the region of interest—required to quantify the metabolic rate of oxygen in vivo in absolute units.


The spectral dependence of optical absorption of absorbers enables photoacoustic spectroscopy voxel by voxel. Consequently, physiological parameters, such as the oxygen saturation of hemoglobin, can be quantified. For example, at the 584 nm isosbestic optical wavelength, where the molar absorption coefficients of oxyhemoglobin and deoxyhemoglobin are identical, the photoacoustic signal is sensitive to the total concentration of hemoglobin but insensitive to the oxygenation of hemoglobin. Tuning the laser to another wavelength where the two forms of hemoglobin have different molar absorption coefficients provides a second measurement. The two measures are combined to quantify the concentrations of both forms, from which the oxygen saturation of hemoglobin can be computed.

Multiscale PAT

Figure 4: Spatial resolution versus penetration achieved by various PAT systems denoted by horizontal bars. AR: acoustic resolution; LA: Linear array; Mac: Macroscopy; OR: Optical resolution; PA: Photoacoustic; SM: Sub-micron; SW: Sub-wavelength.

PAT provides deep imaging at high resolution. At depths beyond the optical diffusion limit, the center frequency and the bandwidth of the ultrasonic detection system predominantly determine the spatial resolution; the greater they are, the better the spatial resolution is, but the worse the ultrasonic penetration becomes. Such scalability empowers PAT for multiscale imaging (Figure 4).

Figure 5: Photograph of an integrated photoacoustic (PA) and ultrasound (US) imaging system modified from a Philips clinical US array system. Reprint with permission (Kim et al. 2010).

Handheld clinical ultrasound probes have been adapted for concurrent PAT (Figure 5), which corresponds to LA-PACT in Figure 4. The nominal bandwidth of the ultrasound probe is 4–8 MHz, providing an axial resolution of ~0.4 mm. Penetration of multiple centimeters in biological tissue has been demonstrated (Figure 6).

Figure 6: In vivo deeply penetrating photoacoustic (PA) and ultrasonic (US) imaging. PA images acquired (a) before and (b) 10 minutes after methylene blue injection. (c) Overlaid post-injection PA (pseudocolor) and US (grayscale) images. B: blood; SLN: sentinel lymph node. Reprint with permission (Kim et al. 2010).

Within the optical diffusion limit, the transverse spatial resolution of PAM can be achieved by optical focusing, which corresponds to OR-PAM in Figure 4. Focusing light through an objective lens with a numerical aperture of 0.1 yields a 2.6 µm lateral resolution, which is limited by the optical focal diameter. Such a resolution allows in vivo imaging of capillaries—the smallest blood vessels, showing single files of individual red blood cells (Figure 7). When the numerical aperture is increased, sub-micron and even sub-wavelength resolution down to 220 nm has been achieved as indicated by SM-PAM and SW-PAM in Figure 4.

Figure 7: Photoacoustic image of the vasculature in small animal skin. RBC: red blood cell. Reprint with permission (Hu, Maslov et al. 2011).

Imaging speed

PAT is fundamentally a high-speed imaging technology. Given a sufficient number of ultrasonic detection channels, the data acquisition is only limited by the acoustic propagation time across the region of interest, merely ~0.1 ms for a 15 cm range. In practice, frames rates of 10s of Hz have been achieved.


PAT has the following limitations. First, optical attenuation limits the penetration to ~5–7 cm in tissue when <1 mm resolution is desired. However, microwaves or radiowaves can be used for deeper excitation although the contrast origins differ. Second, ultrasound sustains strong reflection from gas–liquid or gas–solid interfaces due to the mismatch of acoustic impedances. Therefore, ultrasound signals cannot penetrate through gas cavities or lung tissues efficiently. For the same reason, ultrasonic detection requires direct contact between the ultrasonic transducers and the biological tissue. Usually, ultrasound coupling gel is applied to the tissue surface to avoid intervening air cavities. Third, ultrasound suffers significant attenuation and phase distortion in thick bones such as the human skull. Fortunately, unlike pulse-echo ultrasound imaging, PAT involves only one-way ultrasound attenuation through the skull. Sufficiently strong photoacoustic signals have been observed through Rhesus monkey skulls. The remaining challenge is to compensate for the phase distortion due to the skull.

Potential applications

PAT is expected to find broad applications in both biology and medicine. Preclinical applications include imaging of

  1. Non-fluorescent pigments (red blood cells & melanin)
  2. Angiogenesis and anti-angiogenic response
  3. Microcirculation physiology and pathology
  4. Drug response for screening
  5. Brain functions
  6. Biomarkers
  7. Gene activities through reporter genes

Clinical applications include

  1. Melanoma cancer screening
  2. Gastrointestinal tract endoscopy
  3. Intravascular catheter imaging
  4. Neonatal and adult brain imaging
  5. Breast cancer detection
  6. Prostate cancer detection
  7. Guided sentinel lymph node needle/core biopsy for breast cancer staging
  8. Early response to chemotherapy
  9. Dosimetry in thermal therapy
  10. In vivo label-free histology by photoacoustic imaging of cell nuclei
  11. Blood flow, oxygenation, and tissue metabolism imaging

The unique in vivo multiscale structural, functional, and molecular imaging capability of PAT should enable systems biology research at multiple length scales and accelerate translation of microscopic laboratory discoveries to macroscopic clinical practice.


Cox, B. T., J. G. Laufer, et al. (2009). "The challenges for quantitative photoacoustic imaging." Proc. of SPIE 7177: 717713.

Favazza, C. P., O. Jassim, et al. (2011). "In vivo photoacoustic microscopy of human cutaneous microvasculature and a nevus." Journal of Biomedical Optics 16(1): 016015 (1-6).

Hu, S., K. Maslov, et al. (2011). "Second-generation optical-resolution photoacoustic microscopy with improved sensitivity and speed." Optics Letters 36(7): 1134-1136.

Kim, C., T. N. Erpelding, et al. (2010). "Deeply penetrating in vivo photoacoustic imaging using a clinical ultrasound array system." Biomedical Optics Express 1(1): 278-284. [1].

Maslov, K., G. Stoica, et al. (2005). "In vivo dark-field reflection-mode photoacoustic microscopy." Optics Letters 30(6): 625-627.

Wang, L. V. (2009). "Multiscale photoacoustic microscopy and computed tomography." Nature Photonics 3: 503-509.

Wang, X. D., Y. J. Pang, et al. (2003). "Noninvasive laser-induced photoacoustic tomography for structural and functional in vivo imaging of the brain." Nature Biotechnology 21(7): 803-806.

Zhang, H. F., K. Maslov, et al. (2006). "Functional photoacoustic microscopy for high-resolution and noninvasive in vivo imaging." Nature Biotechnology 24(7): 848-851.

Additional readings

Wang, L. V., Ed. (2009). Photoacoustic Imaging and Spectroscopy. Taylor & Francis.

Wang, L. V. and H. Wu (2007). Biomedical Optics: Principles and Imaging. Wiley.

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