Photoacoustic imaging in biomedicine

  Photoacoustic imaging, as a hybrid biomedical imaging modality, is developed based on the photoacoustic effect. In photoacoustic imaging, non-ionizing laser pulses are delivered into biological tissues (when radio frequency pulses are used, the technology is referred to as thermoacoustic imaging). Some of the delivered energy will be absorbed and converted into heat, leading to transient thermoelastic expansion and thus wideband (e.g. MHz) ultrasonic emission. The generated ultrasonic waves are then detected by ultrasonic transducers to form images. It is known that optical absorption is closely associated with physiological properties, such as hemoglobin concentration and oxygen saturation ^[2]. 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 ^[3]. Fig. 1 is a schematic illustration showing the basic principles of photoacoustic 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 ^[4]. Since blood usually has orders of magnitude larger 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, and skin melanoma detection etc^[3].

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Advantages of photoacoustic imaging

In functional imaging, optical imaging is highly desirable because of the strong correlation between optical absorption and hemoglobin concentration and/or oxygenation. However, existing high-resolution optical imaging technologies, including confocal microscopy, two-photon microscopy, and optical coherence tomography (OCT), do not sense optical absorption directly (Table 1). Moreover, since these imaging modalities rely on ballistic photons for imaging, their imaging depths are limited (Table 1) due to the strong optical scattering in biological tissues (OCT imaging in relatively transparent tissues such as the eye is an exception). Photoacoustic imaging, however, does not rely on ballistic photons for excitation; and ultrasonic waves have 2-3 orders of magnitude weaker scattering than optical waves in biological tissues. Consequently, photoacoustic imaging provides high resolution at relatively large imaging depth (Table 1). Similar to ultrasound imaging, the resolution and imaging depth of photoacoustic imaging is scalable, depending on the frequency of the ultrasound transducer used (refer to the section of photoacoustic microscopy (PAM) for details). Therefore, photoacoustic imaging combines the advantages of optical absorption contrast with ultrasonic spatial resolution for deep imaging beyond the ballistic regime ^[5].

Photoacoustic imaging systems

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/thermoacoustic computed tomography (PAT/TAT)

General photoacoustic equation

Given the heating function , the generation and propagation of photoacoustic wave pressure in an acoustically homogeneous inviscid medium is governed by

where v_s is the speed of sound in medium, β is the thermal expansion coefficient, and 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^[5].

The forward solution of Eq. (1) is given by

In stress confinement, which occurs when the laser pulsewidth is much shorter than the stress relaxation time^[5], Eq. (2) can be further derived as

where p₀ is the initial photoacoustic pressure.

Universal reconstruction algorithm

In PAT system, the acoustic pressure is detected by scanning an ultrasound transducer over a surface that encloses the photoacoustic source. To reconstruct the internal source distribution, we need to solve the inverse problem of equation (3) (i.e. to obtain p₀). A representative method applied for PAT reconstruction is known as the universal back projection algorithm^[6]. This method is suitable for three imaging geometries: planar, spherical, and cylindrical surfaces.

The universal backprojection formula is

where Ω₀ is the solid angle subtended by the entire surface S₀ with respect to the reconstruction point inside S₀, and

A typical PAT/TAT system

A typical PAT system is shown as the left part in Fig. 3. The laser beam is expanded and diffused to cover the whole region of interest. Photoacoustic waves are generated proportional to the distribution of optical absorption in the target, and are detected by an ultrasonic transducer. TAT system is the same as PAT system except using a microwave excitation source instead of laser. A typical TAT system is shown as the right part in Fig. 3. Although single-element transducers have been employed in these two systems, the detection scheme can be extended to using ultrasound arrays as well.

Biomedical applications of PAT/TAT

Intrinsic optical/microwave absorption contrast and diffraction-limited high spatial resolution of ultrasound make PAT and TAT promising imaging modalities for wide biomedical applications:

Brain lesion detection^[9]

Soft tissues with different optical absorption properties in the brain can be clearly identified by PAT. For example, the absorption contrast between the lesion area and the background parenchyma is significant as shown in Fig. 4(a). Fig. 4(b) is the corresponding open-skull photograph after experiment.

Hemodynamics monitoring ^[7][9]

Since HbO₂ and Hb are the dominant absorbing compounds in biological tissues in the visible spectral range, multiple wavelength photoacoustic measurements can be used to reveal the relative concentration of these two chromophores. Thus, the relative total concentration of hemoglobin (HbT) and the hemoglobin oxygen saturation (sO₂) can be derived. Therefore, cerebral hemodynamic changes associated with brain function can be successfully detected with PAT.

Fig. 5 is the functional imaging of cerebral hemodynamic changes in response to whisker stimulation. In Fig.5, (a) is the noninvasive PAT image of the vascular pattern in the superficial layer of the rat cortex acquired with the skin and skull intact. (b) and (c) are functional PAT images corresponding to left-side and right-side whisker stimulation, respectively. (d) is the open-skull photograph of the rat cortical surface (B, bregma; L, lambda; M, midline; A, activated regions corresponding to whisker stimulation).

Fig. 6 shows functional changes of sO₂ and HbT in the rat cerebral cortex as a result of the physiological modulations. Under the hyperoxia status, the averaged sO₂ level, 2>, in the areas of the imaged cortical venous vessels in the rat brain, is higher than that under the normoxia status, and the difference <ΔsO₂> is ~10%; while under the hypoxia status, the 2> is lower than that under the normoxia status, and the difference <ΔsO₂> is ~13%. While the averaged HbT value, , presents a relative increase of ~12%, denoted by <ΔHbT>/, as a result of the change from normoxia to hypoxia, which was comparatively larger than the ~4% relative increase in as a result of the change from hyperoxia to normoxia.

Breast cancer diagnosis^[10]

By utilizing low scattered microwave for excitation, TAT is capable of penetrating thick (several cm) biological tissues with less than mm spatial resolution. Since the cancer tissue and normal tissue have very different response to radio frequency, TAT has great potential in early breast cancer diagnosis. Fig. 7 shows the TAT image of a mastectomy specimen, where malignant breast tissue generates a much stronger thermoacoustic signal than the surrounding benign tissue due to its high microwave absorption.

Photoacoustic microscopy (PAM)

Fig. 8 shows a representative PAM set-up ^[1]. A tunable dye laser is pumped by a Q-switched pulsed Nd:YAG (neodymium: yttrium aluminum garnet) laser. A short laser pulse at a certain wavelength between 532-770 nm is generated to irradiate the target tissue to induce acoustic pressure waves. Laser pulses of ~6 mJ/cm² at the focus will be delivered at 10 Hz repetition rate. An Optical fiber of 0.6 mm core diameter is coaxially positioned on a three-dimensional mechanical stage with changeable ultrasound transducers between 25-75 MHz.

The imaging depth of PAM is mainly limited by the ultrasonic attenuation. The spatial and lateral resolutions depend on the ultrasound transducer used. An ultrasound transducer with high central frequency and broader bandwidth are chosen to obtain high axial resolution. The lateral resolution is determined by the focal diameter of the transducer. For instance, a 50-MHz ultrasound transducer provides 15 micron axial and 45 micron lateral resolution with ~3 mm imaging depth.

PAM has multiple important applications in functional imaging. Two examples are shown in Figs. 9 and 10. PAM can detect changes in oxygenated/deoxygenated hemoglobin in small vessels. As shown in Fig. 9, arterioles (red) and venules (blue) are clearly delineated with high spatial resolution. ^[11]. Fig. 10 shows the capability of PAM to image skin melanoma by using dual wavelength to obtain the morphological relationship between the melanoma and its surrounding blood vessels structures in vivo ^[1].

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