The basic principle of all photoacoustic (and, more generally, thermoacoustic) techniques is that the absorption of electromagnetic energy causes a change of thermal state that leads to a change in temperature, density, and possibly pressure.5
In photoacoustic imaging, acoustic waves are produced by pulsed laser light.6
For biomedical applications, photoacoustic generation is due to the photothermal effect: Heat deposited by absorbed light produces acoustic waves through thermal expansion.9
Temperature elevations are very small—for example, if skin is irradiated with the maximum permissible exposure of laser light, tissue surfaces are warmed by a few hundredths of a degree, and the acoustic wave has an amplitude of hundreds of millibars.
Consider plane-wave illumination of an object by a pulsed laser source of high peak power but very low duty cycle—for example, a 10-ns pulse produced at a repetition rate of 20 Hz. If energy is absorbed faster than it can diffuse away, a local volume with elevated absorption is heated by the laser pulse, and a local acoustic transient is generated. The magnitude and shape of the acoustic transient depend on the duration of the laser pulse.
Under conditions typical for biomedical imaging, the photoacoustic wave equation is
is pressure, vs
is the longitudinal wave speed in the medium, Γ is the dimensionless Grüneisen parameter
β is the thermal coefficient of volume expansion, Cp
is the heat capacity at constant pressure, and H
is the heating function representing thermal energy deposited per unit volume and per unit time.9,10
Absorbed laser light and the corresponding temperature rise generate the source term in the acoustic wave equation. Localized regions of high optical absorption act as sources of propagating acoustic waves.
Under a so-called stress confinement condition—in which the duration of the laser pulse is less than a characteristic confinement time—the amplitude of the acoustic wave launched by an optical absorber depends only on the total amount of energy absorbed, not on its time profile, and so is given by ΓμaF
, where μa
is the optical absorption coefficient and F
is the optical fluence at the absorber. The acoustic wave’s specific shape is determined by the boundary conditions for all optical absorbers heated by the laser. In the limit of highly localized, instantaneous absorption, the acoustic pulse shape is approximated by the time derivative of the optical pulse.10
Propagating acoustic waves created by each optical pulse can be detected and recorded by an array of ultrasonic transducers, as illustrated in , and rapidly reconstructed to produce real-time images of the source distribution. If optical illumination and acoustic detection systems are properly designed, a complete 2D or 3D image can be produced from a single laser pulse. In addition, ultrasonic and photoacoustic imaging can share array and receiver electronics and can therefore be combined to obtain photoacoustic and ultrasonic images of the same object in rapid succession at real-time (that is, video) rates.
Figure 3 Photoacoustic generation and detection. Black dots in the left panel represent regions of high optical absorption. When heated by a laser pulse, they launch acoustic waves, which are picked up by an ultrasound detector. The ultrasound waveform shown in (more ...)
Ultrasonic imaging is the mostly commonly used real-time clinical modality capable of 2D and 3D images.11
Spatial resolution is directly related to the acoustic frequency, which is typically 1–20 MHz, and corresponds to acoustic wavelengths of 1.5 mm to 75 µm. The ultrasonic attenuation coefficient is also nearly linear with operating frequency, so there is a direct tradeoff between spatial resolution and penetration. Submillimeter spatial resolution is typical up to penetration depths of about 15 cm, and better than 100-µm resolution is possible at depths up to several centimeters. Contrast in ultrasonic images is related to mechanical characteristics of soft tissue. Since mechanical properties are not directly related to molecular function, ultrasonic systems provide mostly anatomical images and physiological measurements related to blood flow and tissue elasticity. Therefore, when ultrasonic and photoacoustic systems are combined, both morphology (that is, anatomy) and function (for example, blood oxygenation as illustrated in ) can be imaged simultaneously. Furthermore, multimodal imaging can be extended to cellular and molecular events.
shows a block diagram of a typical photoacoustic imaging system. To achieve the spatial resolution and field of view required for biomedical applications, most systems use nanosecond pulsed lasers such as Q-switched neodymium-doped yttrium aluminum garnet, alexandrite, and tunable pulsed lasers or high-peak-power pulsed diode lasers. The laser beam’s energy is typically delivered to the tissue via an optical fiber or a set of optical fibers tightly bundled on the laser-head side and distributed around the ultrasonic transducer on the other end. Each fiber may be attached to an optical diffuser positioned to produce nearly uniform optical illumination of the image plane probed by the transducer array. Light scattering in tissue also helps to spatially homogenize the laser irradiation. If a single fiber is used, the laser beam is distributed around the transducer and into the tissue using an assembly of lenses, prisms, and mirrors.
A typical photoacoustic imaging system consisting of a pulsed laser, ultrasonic pulser–receiver, processing unit, and fiber-optic light delivery system interfaced with the ultrasonic transducer.