Nuclear binding energy

Fig. 1. Schematic illustration of photoacoustic imaging.

File:Melanoma3DMovie.gif

3D photoacoustic imaging of melanoma in vivo.

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. 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. Fig. 1 is a schematic illustration showing the basic principles of photoacoustic imaging.

Fig. 2. Absorption spectra of oxy- and deoxy-hemoglobin.

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 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.

Advantages

	Δf	Primary contrast	Δz	δz	δx	Speed
	Hz		mm	µm	µm	Mvx/s
Photoacoustic microscopy	50 M	Optical absorption	3	15	45	0.5
Photoacoustic microscopy	5 M	Optical absorption	50	700	700	0.5
Confocal microscopy		Fluorescence, scattering	0.2	3-20	0.3-3	10-100
Two-photon microscopy		Fluorescence	0.5-1.0	1-10	0.3-3	10-100
Optical coherence tomography	50 T	Optical scattering	1-2	0.5-10	1-10	20-4.000
Scanning Laser Acoustic Microscopy	300 M	Ultrasonic scattering	1-2	20	20	10
Acoustic microscopy	50 M	Ultrasonic scattering	20	20-100	80-160	0.1
Ultrasonography	5 M	Ultrasonic scattering	60	300	300	1
Table 1. Comparison of contrast mechanisms, penetration depth (Δz), axial resolution (δz), lateral resolution (δx=δy) and imaging speed of confocal microscopy, two-photon microscopy, optical coherence tomography (50 THz), ultrasound microscopy (50 MHz), ultrasound imaging (5 MHz), photoacoustic microscopy (50 MHz), and photoacoustic tomography (3.5 MHz). Speeds in Megavoxel per second of non-parallel techniques.

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/optoacoustic computed tomography (PAT/TAT/OAT)

General equation

Given the heating function $H({\vec {r}},t)$, the generation and propagation of photoacoustic wave pressure $p({\vec {r}},t)$ in an acoustically homogeneous inviscid medium is governed by

$\nabla ^{2}p({\vec {r}},t)-{\frac {1}{v_{s}^{2}}}{\frac {\partial ^{2}}{\partial {t^{2}}}}p({\vec {r}},t)=-{\frac {\beta }{C_{p}}}{\frac {\partial }{\partial t}}H({\vec {r}},t)\qquad \qquad \quad \quad (1),$

where $v_{s}$ is the speed of sound in medium, $\beta$ 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.

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

$\left.p({\vec {r}},t)={\frac {\beta }{4\pi C_{p}}}\int {\frac {d{\vec {r'}}}{|{\vec {r}}-{\vec {r'}}|}}{\frac {\partial H({\vec {r'}},t')}{\partial t'}}\right|_{t'=t-|{\vec {r}}-{\vec {r'}}|/v_{s}}\qquad \quad \,\,\,\,(2).$

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

$p({\vec {r}},t)={\frac {1}{4\pi v_{s}^{2}}}{\frac {\partial }{\partial t}}\left[{\frac {1}{v_{s}t}}\int d{\vec {r'}}p_{0}({\vec {r'}})\delta \left(t-{\frac {|{\vec {r}}-{\vec {r'}}|}{v_{s}}}\right)\right]\qquad \,(3),$

where $p_{0}$ is the initial photoacoustic pressure.

Universal reconstruction algorithm

In a 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 $p_{0}$). A representative method applied for PAT reconstruction is known as the universal backprojection algorithm. This method is suitable for three imaging geometries: planar, spherical, and cylindrical surfaces.

The universal backprojection formula is

$\left.p_{0}({\vec {r}})=\int _{\Omega _{0}}{\frac {d\Omega _{0}}{\Omega _{0}}}\left[2p({\vec {r_{0}}},v_{s}t)-2v_{s}t{\frac {\partial p({\vec {r_{0}}},v_{s}t)}{\partial (v_{s}t)}}\right]\right|_{t=|{\vec {r}}-{\vec {r_{0}}}|/v_{s}},\qquad \quad (4),$

where $\Omega _{0}$ is the solid angle subtended by the entire surface $S_{0}$ with respect to the reconstruction point ${\vec {r}}$ inside $S_{0}$, and

$d\Omega _{0}={\frac {dS_{0}}{|{\vec {r}}-{\vec {r_{0}}}|^{2}}}{\frac {{\hat {n}}_{0}^{s}.({\vec {r}}-{\vec {r_{0}}})}{|{\vec {r}}-{\vec {r_{0}}}|}}.$

A simple PAT/TAT/OAT system

A simple PAT/TAT/OAT system is shown in the left part of 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 a single scanned ultrasonic transducer. A TAT/OAT system is the same as PAT except that it uses a microwave excitation source instead of a laser. A typical TAT system is shown in the right part of Fig. 3. Although single-element transducers have been employed in these two systems, the detection scheme can be extended to use ultrasound arrays as well.

File:Setup of PAT and TAT.png

Fig. 3. Schematic of single-element transducer PAT/TAT system.

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

File:Brain lesion PAT.png

Fig. 4. PAT imaging of rat brain lesion in situ.

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

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.

File:Brain stimulation PAT.png

Fig. 5. Cerebral hemodynamic changes in response to whisker stimulation.

File:O2 saturation PAT.png

Fig. 6. Image of the sO2 and HbT change in the cortical venous vessels between different physiological status.

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, <sO₂>, 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 <sO₂> is lower than that under the normoxia status, and the difference <ΔsO₂> is ~13%. While the averaged HbT value, <HbT>, presents a relative increase of ~12%, denoted by <ΔHbT>/<HbT>, as a result of the change from normoxia to hypoxia, which was comparatively larger than the ~4% relative increase in <HbT> as a result of the change from hyperoxia to normoxia.

Breast cancer diagnosis

File:Breast cancer TAT.png

Fig. 7. Thermoacoustic image of a mastectomy specimen.

By utilizing low scattered microwave for excitation, TAT is capable of penetrating thick (several cm) biological tissues with less than mm spatial resolution. Since cancerous tissue and normal tissue have very different responses to radio frequency radiation, 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. 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.

File:Setup of PAM.png

Fig. 8. Experimental set-up of dark field reflection mode PAM system.

File:O2 saturation PAM.png

Fig. 9. Multi-wavelength PAM imaging for blood oxygenation mapping of rat skin.

File:Melanoma PAM.png

Fig. 10. Multi-wavelength PAM imaging for in vivo melanoma detection.

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 micrometre axial and 45 micrometre 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. 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.

See also

• Photoacoustic tomography
• Multispectral Optoacoustic Tomography

References

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External links

• [1] Optical Imaging Laboratory, Department of Biomedical Engineering, Washington University in St. Louis
• [2] Photoacoustic Imaging Group, Department of Medical Physics and Bioengineering, University College London
• [3] Optoacoustic Imaging Lab, Department of Biomedical Photonics, Institute of Applied Physics, University of Bern
• [4] Biomedical Photonic Imaging Group, BioPhysical Engineering Group, Faculty of Science and Technology, University of Twente, Enschede, The Netherlands
• [5] Functional Optical Imaging Laboratory (FOIL), Department of Biomedical Engineering, Northwestern University
• [6] Biophotonics and Imaging Laboratory (BAIL), Department of Biomedical Engineering, Oregon Health & Science University
• [7] OptoSonics, Inc.
• [8] Tomowave Laboratories, Inc.
• [9] Biomedical Optical Imaging Laboratory, Department of Electrical & Computer Engineering, University of Connecticut
• [10] An interdisciplinary consortium of experts in photoacoustic imaging in Austria where research in physics and mathematics is heavily driven by input from biology and medicine
• [11] Recent advances in application of acoustic, acousto-optic and photoacoustic methods in biology and medicine
• [12] Ultrasound Imaging and Therapeutics Research Laboratory, Department of Biomedical Engineering, University of Texas at Austin