3.1 Basics of Ultrasonography

Ultrasonography (US) is performed using physical properties of ultrasound, i.e., acoustic waves with a frequency above 20,000 Hz (20 kHz) (Fig. 3.1). The normal range of human hearing of a healthy individual is between 16 Hz and 20 kHz. Therefore, ultrasound waves are all acoustic waves with frequency higher than the threshold of human perception of sound. In clinical practice acoustic waves with frequency from 2 to 30 MHz are propagated within a patient’s body (Ahuja and Evans, 2013, [1]).

Contrary to electromagnetic radiation such as X-rays, ultrasound waves require an elastic medium for propagation that can be deformed. When oscillation is transmitted, energy transfer occurs.

The source of ultrasound is a probe, also called a transducer, containing piezoelectric elements usually composed of barium titanate or lead zirconate. These crystals or ceramic elements are characterized by special properties, i.e., when electrical current is applied to the piezoelectric element, it contracts and at the same time emits an acoustic wave. The current is proportional to the force of squeezing. Change of the direction and the focus depth of ultrasound wave is adjusted using the phased array techniques. When the piezoelectric crystal is stretched, the voltage turns to the opposite.

The produced ultrasound wave enters the human body then returns from the examined tissues as mechanical oscillations of waves reflected off the given object, known as echoes (Fig. 3.2). When the wave comes back to the piezoelectric element, it serves as the receiver of acoustic signals producing an electric current. Obviously, the characteristics of the current incurred by returning acoustic wave is different from the initial one.

Acoustic waves are characterized by velocity, wavelength, frequency, and intensity. The velocity equals to the product of multiplication of frequency (in Hertz) and wavelength (in meters). The velocity of ultrasound wave in a human is estimated to be 1540 m/s which is an average value of velocity in tissues, similar to that of an acoustic wave propagated in water. In soft tissues and fluids transmission of ultrasound occurs in the form of longitudinal waves, i.e., the direction of wave propagation is the same or opposite from the direction of the displacement of the medium. In bone, longitudinal wave propagation is accompanied by transverse waves. The velocity of the acoustic wave is characteristic for the medium in which it is propagated therefore when crossing tissue interfaces only wavelength is changed and frequency determined by the source is unaffected [1].

Intensity of ultrasound beam influences the range of examination, i.e., the depth at which imaging still can be performed. Intensity determines the amount of energy transmitted by a wave in one second per unit of area at right angle to the direction of wave propagation.

Ultrasound field generated by a piezoelectric crystal is divided into two parts—near field and far field. In the near field, the width of the beam is constant and the shape of the beam is similar to a cylinder, while in the far field the beam becomes divergent. In the near field, the structure of ultrasound field is not homogenous due to overlapping and interfering of spherical partial waves emitted by different parts of the piezoelectric element. The size of the near field is increased in larger probes, it also increases together with probe frequency.

Since patient tissues are not homogenous, ultrasound encounters different tissue interfaces and various internal structures (fluid collections, calcifications, gas bubbles, discontinuities in tissues). The impedance of tissues is varied, and its values for soft tissues are similar to those of water. The interactions with areas differing in acoustic impedance lead to changes in the returning wave characteristics in comparison with the emitted original ultrasound wave. Inside the examined objects physical phenomena occur—analogical to those applied in optics. These phenomena include reflection, deflection, refraction, scattering, and absorption with the release of thermal energy (Fig. 3.3). Strong reflections occur when areas very distinct in impedance are imaged such as interfaces between soft tissue and bone as well as between soft tissue and air. Reflection is dependent on the angle of incidence of acoustic wave—when it hits the examined object at right angle, the reflection is strong. On the contrary, lower angulation leads to partial reflection and less echo coming back to the probe. Scattering and absorption result in attenuation of ultrasound wave, which decreases penetration depth of an acoustic wave. Scattering is responsible for creation of internal image of tissues, and more rough surfaces produce more scattering. Refraction is observed in case of differences in velocity of acoustic waves in two areas of the imaged object.

The part of acoustic wave picked up by piezoelectric elements in the transducer acting as the receiver gives rise to electric signal which is presented on screen in real-time as a black-and-white two-dimensional image (B-mode—brightness mode) (Fig. 3.4). Brightness of each pixel in the image on screen depends on the amplitude of the echo. As the sound velocity is assumed to be constant and the time required for the wave to travel forth and back known, it is possible to calculate the depth of tissues where wave reflection occurred. This way the final image in the screen can be compared to a sonar map of the sea bed, where structures located closer to the probe (thus, the surface of skin or mucosa) are visualized on top of the screen, and those lying deeper in scanned tissues further away from the probe towards the bottom of the screen [2].

In the B-mode, the so-called echogenicity is evaluated. An area which produces strong echoes is called hyperechoic (Fig. 3.4), on the other hand areas with no internal echoes are named anechoic. Hypoechoic lesions (Fig. 3.5) are characterized by echogenicity lower than structures in the vicinity, while areas of the same or comparable echogenicity are called isoechoic. Post-acoustic shadowing occurs when ultrasound beam is completely reflected off the outer surface of a structure such as, e.g., condyle or a very dense lesion, such as calcification (Fig. 3.4). When ultrasound beam travels through a very low-density lesion such as a cyst, the liquid content does not reflect ultrasound thus a bigger portion of the beam reaches tissues located below the fluid collection and more echoes are generated behind the lesion. This appearance is called post-acoustic enhancement (Fig. 3.5).

A-mode (amplitude mode) is more simple than B-mode as it does not depict distribution of echoes over a two-dimensional cross-section but presents them as plotted amplitude of spikes versus time as a function of depth. In this type of presentation, the probe is placed on skin surface and is not displaced during the examination. Therefore, only mobile objects will produce images in the form of amplitude of echoes. This mode is used in ophthalmology for estimation of distances between different parts of the eye [3].

In M-mode (motion mode, also called Time-Motion mode) the probe is also stationary and only a single chosen ultrasound line is emitted and received. All US reflecting objects are displayed on screen along the time axis. In this type of image, echoes are presented as pixels and their brightness corresponds to the magnitude of echo amplitude. Very high sampling rate in this mode is advantageous as it allows detection and quantification of very fast motions. M-mode is mostly used in cardiology.

Tissue Harmonic Imaging (THI) is based on the properties of nonlinear propagation of ultrasound in the examined tissues. The shape of the ultrasound wave is distorted due to uneven velocity of the wave propagation—i.e., faster high-pressure portion of the wave and slower low-pressure part of the beam. The difference in the form of the wave produces the so-called tissue harmonics which are multiples of the frequency—either fundamental or transmitted. Subsequent harmonics are characterized by decreasing amplitude therefore only the second harmonic is sufficient for generation of an image. THI technique increases signal-to-noise ratio, reduces artifacts coming from reverberations, as well as increases resolution, both axial and lateral.

US elastography offers a possibility of evaluation of stiffness of tissues on the basis of analysis of change of their shape when an external stimulus is applied such as exerted pressure (Fig. 3.6) or emission of an acoustic impulse propagated within tissues as the so-called shear wave (Fig. 3.7). A colored map represents areas with higher and lower stiffness in a qualitative manner (Fig. 3.8). In some US machines quantitative assessment is possible in the form of Young’s modulus values given in kilopascals (Yuan et al. 2016). Usefulness of the technique is already established for, e.g., breast lesions, thyroid nodules, and musculoskeletal application, and it is being investigated in the maxillofacial area including salivary glands, lymph nodes, muscles of mastication, palatal tumors, tongue carcinoma, and TMJ disk [4–8].

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