Ultrasounds are known since about 170 years. Firstly, echo sounding has been used in the First World War to fix the position of submarines, now we apply the echo sounding method in medical imaging to study organs and tissue structures, to learn about the body’s function, and to recognize the malfunction. That requires good knowledge about both physical and technological properties in ultrasonic and the tissue interaction with ultrasound waves (Fig. 4.1) [1].

../images/462633_1_En_4_Chapter/462633_1_En_4_Fig1_HTML.png — Fig. 4.1

Sound wave

Ultrasound (US) imaging has become an attractive technology for medical diagnosis in a variety of clinical settings, including obstetrics and gynecology, cardiology, and urology. This is due to its noninvasive nature, real-time capability, and cost-effectiveness. Unlike other diagnostic modalities, ultrasound systems do not emit ionizing radiation and scans may be conducted as often as necessary without the risks of repeated exposure to X-rays or radionuclides.

4.1 Basic Physical Principles of Medical Ultrasound

4.1.1 The Sound Wave

In ultrasound diagnosis, we apply the mechanical energy of the sound wave. The wave is physical appearance by which the mechanical energy is propagating through space, by particles oscillation.

Unlike X-rays, sound is not electromagnetic, matter must be present for sound to travel, which explains why sound cannot propagate through a vacuum. Sound propagation is the energy transfer from one place to another within a medium, and some energy is also imparted to the medium [1]. Sound is categorized according to its frequency, number of mechanical variations occurring per unit time, and the terms infrasound and ultrasound are used for frequencies below and above the audible range (Table 4.1):

Table 4.1

Different sound frequencies

Infrasonic	<20 Hz
Human audible range	20 Hz–20 KHz
Ultrasonic	>20 kHz

Ultrasound is sound whose frequency is above the range of human hearing (high frequency sound waves).

The use of high frequencies is limited by their greater attenuation in tissue and thus shorter depth of penetration.

For this reason, different ranges of frequency are used for examination of different parts of the body:

3–5 MHz for abdominal areas
5–10 MHz for small and superficial parts
10–30 MHz for the skin

Although ultrasound images are captured in real time, they can also show movement of the body internal organs as well as blood flowing through the blood vessels.

The characteristics of a sound wave can be described by the following parameters (Table 4.2, Fig. 4.2) [1–4].

Table 4.2

Characteristics of a sound wave

Length	λ	The distance between two consecutive, identical positions in the pressure wave (e.g., between two compressions or between two rarefactions) It is determined by the frequency of the wave and the speed of propagation in the medium through which it is travelling
Frequency	ƒ	The number of cycles per second performed by the particles of the medium in response to a wave passing through it. Expressed in Hertz, where I Hz = 1 cycle passing a given point each second, therefore 3 MHz = 3 million cycles per second
Period	T	The time taken for a particle in the medium through which the wave is travelling to make one complete oscillation about its test position
Amplitude	A	The maximal displacement of the oscillating particle from the equilibrium position
Velocity of propagation	v	The speed of sound with direction specified. When a sound wave travels through any medium, it is certain that parameters of that medium determine the speed of sound propagation

../images/462633_1_En_4_Chapter/462633_1_En_4_Fig2_HTML.png — Fig. 4.2

Representation of sound wave characteristics

The body is elastic medium in which the sound wave longitudinally propagates. The wave frequency is the source oscillation. The sound wave, in reason of the generated mechanical energy, causes the particles oscillation that determines local pressure and medium density regular changes [2].

The relationship between frequency, velocity, and wavelength is given by the following formula:

$\lambda =c/f$

The wave propagation velocity is determined by the medium properties: in the air the slowest propagation (344 m/s), increases in liquids and solids (water: 1520 m/s; liver: 1566 m/s; skull bone: 2717–4077 m/s).We can know that ultrasound propagates in soft tissue with an average velocity of around 1540 m/s.

The acoustic impedance, defined as the product of the acoustic velocity in the medium and the density of the medium (Z = p × v), is a parameter that describes how the sound wave differently propagates in tissues.

4.1.2 Generation of Ultrasound

The source of ultrasound waves is a piezoelectric crystal [2–4]. The principle of piezoelectricity is central to the ultrasound generation, it’s defined how the ability of some materials produce a voltage when deformed by an applied pressure and produce a pressure when deformed by an applied voltage.

So, when a voltage is applied to the faces of a crystal, it expands or contracts, then it resonates converting electricity to ultrasound. At the same time, when the crystal receives an echo, the sound deforms the crystal and a voltage is produced on its faces. This is analyzed by the system.

Piezoelectric crystal is the source and the detector of ultrasound and, consequently, plays a very important role in transducer technology.

Ultrasound probes are produced of very thin crystals because the best transformation in both directions by the piezoelectric effect is achieved when the crystal frequency is the same as the frequency of the sound wave. It is realized for thickness of λ/2.

The most used materials are ferroelectric ceramic, lead zirconate titanate, or plastic, polyvinylidene difluoride. The crystals are sandwiched by silver film electrodes on opposite sides.

The transducer is usually treated as a three-port network including two mechanical ports and one electrical port as shown in Fig. 4.3. The mechanical ports represent the front and back surfaces of the piezoelectric element and the electrical port represents the electrical connection of the piezoelectric element to the electrical source. The front layer is known as an acoustic matching layer, which can improve the transducer performance significantly [1–5].

../images/462633_1_En_4_Chapter/462633_1_En_4_Fig3_HTML.png — Fig. 4.3

Transducer as a three-port network (in pink: piezoelectric crystal; in glycine and violet: mechanical ports; in green: acoustic matching layer)

4.1.3 Interaction of Sound with Tissue

The ultrasound shape beam is important for the image quality. The beam profile is made up of three parts (Fig. 4.4):

1.

Near field (Fresnel zone)
2.

Far field (Fraunhofer zone)
3.

Transition point

../images/462633_1_En_4_Chapter/462633_1_En_4_Fig4_HTML.png — Fig. 4.4

The ultrasound shape beam

The near field or Fresnel zone is the part of the beam useful for imaging; however, this can be quite large in area, depending on the diameter of the crystal. The shape and dimensions of this first trait are determined by the shape and dimensions of the emitter and by the frequencies of the wave. Imaging requires a very narrow beam to produce high-resolution diagnosis. In the far field or Fraunhofer zone, the second trait, the intensity is uniformly decreasing with distance. Resolution is preserved in a better way if the near field is longer [3–5].

A beam profile of a simple transducer appears as in the diagram below and shows that the energy is not confined to a single lobe, but radiates off at various angles to the transducer face as off-axis energy. These off-axis areas are called side-lobes (Fig. 4.5).

../images/462633_1_En_4_Chapter/462633_1_En_4_Fig5_HTML.png — Fig. 4.5

The beam profile with energy divided into primary beam and side-lobes or off-axis areas

When the ultrasonic wave traverses body’s tissues, part of the energy will be transmitted but various factors can cause it to lose energy, decreasing in intensity and amplitude. This phenomenon is called attenuation, and it depends on distance, tissue characteristics, and beam frequency. Approximately, the attenuation rate in soft tissue is 1 Db/MHz/cm, and there are four types of processes which contribute to the attenuation (Fig. 4.6):

../images/462633_1_En_4_Chapter/462633_1_En_4_Fig6_HTML.png — Fig. 4.6

The main processes of attenuation of ultrasonic wave passing through tissues

Refraction: it’s the deviation in the pathway of the beam, generating from two different speeds of sound at the tissues interface, in which the angle of incidence isn’t 90°. Refraction can produce artifacts.
Reflection: it’s an acoustic impedance mismatch, occurring at the interface between soft tissues of different acoustic impedance, when the interface is large relative to the wavelength. At the interface, a percentage of the sound is reflected, the others are transmitted into the tissue (soft tissue/air interface—99% is reflected, soft tissue/bone—40% is reflected, liver/kidney—2% is reflected).
Absorption: it’s a transfer of energy of the beam to the medium in which sound is passing. This kind of attenuation of sounds increases with frequency, so we can understand why high frequency transducers cannot be used for examining deep structures in the body. Absorption increases with the viscosity of the medium. Bone adsorbs ultrasound much more than soft tissue, so that, in general, ultrasound is suitable for only bones surfaces. They can’t reach the areas behind bones; therefore, the black zone behind bones is called acoustic shadow.
Scattering: it occurs when interface is equal in size to wavelengths, it’s also named non-specular reflection and it’s characterized by scattering in many directions not with equal amount in all directions [8].

Definitively, if interface is larger than wavelength, we can apply classical laws of reflection and refraction. For reflected waves reflected angle is equal to incident one, and for transmitted wave the angle depends on the velocity of propagation in the media. If the incident angle is zero, the transmitted wave does not change direction and the reflection is backward.

This is very important for diagnostic procedures. When the acoustic impedance is the same, the transmission is maximal, instead the reflection is maximal with very different acoustic impedance.

The greater the difference in acoustic impedance between two media, the higher the fraction of the ultrasound energy that is reflected at their interface and the higher the attenuation of the transmitted part.

Air and gas reflect almost the entire energy of an ultrasound pulse arriving through a tissue. Therefore, an acoustic shadow is seen behind gas bubbles.

For this reason, ultrasound is not suitable for examining tissues containing air, such as the healthy lungs. For the same reason, a coupling agent is necessary to eliminate air between the transducer and the skin.

Certain tissue densities, such as bone, diaphragm, pericardium, slow the ultrasound beam, reflecting the waves and producing a bright or hyperechoic image. Other tissues allow ultrasound beams to pass and reflect at moderate speeds, creating a gray image on the screen, such as muscle, liver, or kidney.

Some tissue allows ultrasound waves to pass easily and retain their strength, creating dark black or hypoechoic images on the ultrasound screen, such as blood, ascites, amniotic fluid, and urine (Fig. 4.7) [3, 4, 8].

../images/462633_1_En_4_Chapter/462633_1_En_4_Fig7_HTML.jpg — Fig. 4.7

The example of hypoechoic image (amniotic fluid) and hyperechoic image (the placenta)

4.1.4 Resolution

Resolution is defined, as the ability to distinguish echoes in terms of space, time, and strength and is critical for the high-quality image resolution.

Contrast resolution: it refers to the ability of an ultrasound system to demonstrate differentiation between different tissues (liver/kidney).
Temporal resolution: it is the ability to show changes over time. This is very important for particular exams, like echocardiography.
Spatial resolution: It is the ability to detect and display structures that are close, together. It is reasonable to consider two different types of spatial resolution: axial and lateral one. Axial resolution is dependent especially upon the length of the pulse used to form the beam (SPL: spatial pulse length), so the shorter the pulse length, the better the axial resolution. The SPL is determined by the number of cycles in one pulse and the length of each cycle.

We know that the wave length is inversely proportional to the frequency, so higher frequencies generate shorter wavelengths and therefore shorter pulse lengths. In conclusion, we can say that higher frequency transducers will show better axial resolution. On the other hand, the lateral resolution is the ability to distinguish between two separate targets perpendicular to the beam [3, 4, 8].

4.1.5 Artifacts

The term artifact is used to describe part of image that does not accurately represent the anatomic structures present in the subject evaluated. Consequently, features that result from incorrect adjustment of instrument settings, are not true artifacts.

Artifacts may variously affect the image quality, but, in the majority of cases, they are easy to recognize. Sometimes, they hinder the correct diagnosis or lead to false diagnosis of a pathological condition where none exists (Fig. 4.8a-b) [10].

../images/462633_1_En_4_Chapter/462633_1_En_4_Fig8_HTML.png — Fig. 4.8

(a–b) Artifact linked to the ultrasound contrast medium

We know that ultrasound systems operate on the basis of described assumptions relating to the interaction of the sound beam with soft tissue interfaces, such as:

A constant speed of sound in the body (assumed 1540 m/s).
All echoes detected by the transducer originate from the central axis of the beam.
The ultrasound beam travels in straight lines.
The time taken for an echo from a given interfaces to return to the transducer is directly related to its distance from the transducer.
The rate of attenuation of the beam is constant with depth and throughout the field of view.

If these assumptions are incorrect, the result is the image that does not accurately reflect scan plane anatomy. The list below describes some of the common artifacts, some of which offer useful diagnostic information.

Acoustic Shadowing is a total reflection on a strong reflector, such as gas or foreign body, or extensive absorption (bone), of the ultrasound wave. The shadow is seen underneath the bright object due to US waves reflection towards the probe. However, from a diagnostic point of view, it may limit the examination of body regions behind the gas or the bone, but on the other hand it is useful for diagnosing stones, calcifications, or foreign bodies [9, 10].

Some interfaces causing such effects are:

Soft tissue/gas (bowel, lung)
Soft tissue/bone, calcium (ribs, calculi)
Normal tissue/fibrous tissue (scars, ligaments)

Mirror image artifact is a multiple reflection of the ultrasound beam. Consequently, on the screen there is false representation showing two images of a single object. However, it is possible to distinguish the “mirrored” object from the real image because it appears deeper and will disappear if the probe position is changed.

Examples of these are:

Diaphragm/lung interface (show mirroring of the liver texture, above the diaphragm)
Bladder/rectum interface, when the rectum is gas filled (bladder wall is mirrored together with any bladder contents, such as ureterocele, catheter balloon, and mass)

Posterior acoustic enhancement occurs if ultrasound waves reach quickly a low density medium, such as urine in the bladder, and then reflect back quickly from the next structure encountered, higher in density, such as the posterior bladder wall. In this case, the deepest structure appears falsely hyperechoic, and the gain adjustment may be needed.

The basis for this artifact is that fluid-filled structures (cysts, gall bladder, etc.) attenuate sound to a much lesser degree than solid organs (liver, spleen, etc.). Therefore, there is more sound transmitted to structures deep to the fluid-filled structure and the resulting echoes from this deeper tissue are brighter than those from a similar depth in adjacent solid tissue [5, 10].

4.1.6 Ultrasound Techniques

The echo principle forms represent the basis of all common ultrasound techniques. The distance between the transducer and the reflector in the tissue is measured by the time between the emission of a pulse and reception of its echo. Additionally, the intensity of the echo can be measured. With Doppler techniques, comparison of the Doppler shift of the echo with the emitted frequency gives information about any movement of the reflector. The various ultrasound techniques used are described below [1, 3, 8].

4.1.6.1 A-Mode

A-mode (amplitude modulation, A-scan) is a one-dimensional examination technique in which a single crystal transducer is used. This mode presents amplitudes of reflected ultrasound in dependence of time. The echoes are displayed on the screen along a time axis (distance) as peaks proportional to the intensity of each signal (amplitude) (Fig. 4.9). The use of this technique today is very limited [8].

../images/462633_1_En_4_Chapter/462633_1_En_4_Fig9_HTML.jpg — Fig. 4.9

Schematic representation of A-mode and B-mode

4.1.6.2 B-Mode

B-mode (brightness modulation) is similar to the precedent technique, but in this case echoes are displayed as points of different gray-scale brightness corresponding to the intensity of each signal (Fig. 4.10). The brightness of the spot is a semi-quantitative measure of impedance relationship for the media building the reflecting interface [8].

../images/462633_1_En_4_Chapter/462633_1_En_4_Fig10_HTML.png — Fig. 4.10

B-mode or brightness mode technique

This presentation has a good resolution, but it can’t give a lot of data about interfaces nature, presenting only the interfaces of great difference. By adding the gray-scale the image provides more information about the structures. Indeed, in gray scale we have different echo intensities presenting in shades of gray.

While the A-scan is a single line investigation, B-scan is a 2D image and is most commonly used in practice. A probe with more than one crystal is used, and the images are produced by groups of piezoelectric elements. Another aspect is that B-mode images are presented in real time and can even follow the movement inside the tissues (blood flow, heart).

The images we use in diagnostic ultrasound are produced with B-Mode or brightness mode techniques. As a result of the signals caused by returning echoes, we obtain a pattern of dots on the screen. The strength of the echo received determines the brightness of each dot proportionally. For each pulse emitted, there is a line of sight as a result of the echoes derived from the tissue interaction. The image, known as a “frame” on the monitor, is composed of multiple lines of sights and, consequently, of multiple thousands of dots. Finally, the real-time image is a succession of several frames, generally 10–30 per second (Figs. 4.11 and 4.12).

../images/462633_1_En_4_Chapter/462633_1_En_4_Fig11_HTML.jpg — Fig. 4.11

Hepatorenal recess or Morison’s pouch, the space that separates the liver from the right kidney

../images/462633_1_En_4_Chapter/462633_1_En_4_Fig12_HTML.png — Fig. 4.12

Two examples of optimization of scan image, modifying PRF in the study of umbilical cord. Note the aliasing in (a–c), and the absence of that phenomenon in (b–d)

4.1.6.3 M-Mode (TM-Mode)

This technique analyzes moving structures, such as heart valves. It consists of a continuous recording over time of the echoes generated by the transducer. On the ultrasound screen, there will be shown motion towards and away from the ultrasound probe at any depth along that line [8].

4.1.7 Doppler

In medical ultrasounds, the Doppler effect is used to measure the velocity of organs or fluids in the body, for example, time motion of heart muscle or the speed of blood in arteries. It was initially discovered by Austrian Mathematician Christian Doppler, who hypothesized that the pitch of the sound would change if the source was moving [7].

The practical description of the Doppler effect is that of the change in pitch of a train or siren as they come towards or away from you when you are standing still. The pitch increases as the train is coming towards you and decreases as it is moving away. In medical application, the Doppler effect is applied mostly for detecting the blood flow. In particular, as the blood is moving towards the transducer, the received frequency will be higher than the initial frequency; on the other hand, as the blood is moving away from the transducer, the received frequency will be lower. The Doppler shift frequency is the difference between transmitted and received frequencies. It depends on the insonating frequency ( ), velocity of moving blood (V), speed of sound in tissue (c), and the angle between the ultrasound beam and the flow direction (φ) as shown in the formula below:

$\cos\ \varphi$

Flow speed calculations based on the Doppler shift measurements can only be accomplished correctly with knowledge of the Doppler angle which is in the form of a cosine. For a given flow, the greater the Doppler angle is, the lesser the Doppler shift is. Consequently, a zero degree angle gives the greater Doppler shift and at 90° the Doppler shift is reduced to nil [6, 7].

There are three different types of Doppler equipment used for the detection of flow: Continuous Wave and Pulsed Wave spectral instruments and Color Mapping instruments.

Continuous Wave: the transducer consists of two crystals, one permanently emitting ultrasound and the other receiving all the echoes. It is possible to measure high flow velocities. Because there is no depth range gate, the signals can be very confusing and include a broad range of frequency shifts (Table 4.3).
Pulsed Wave: pulsed wave as well as for the continuous wave, the transducer works both as a transmitter and a receiver. Moreover, ultrasound is emitted in short pulses, thus the echoes arrive to the transducer between pulses, in an interval named gate. The receiver gate, setting by the operator at a given depth, reproduces only echoes arising from moving reflectors at that depth. A common problem with all pulsed Doppler techniques is the analysis of high velocities: the range for the measurement of Doppler frequencies is limited by the pulse repetition frequency (PRF). Other terms describing PRF are flow rate or scale. This determines the sampling time that is required to process the Doppler information. Since low flow must be sampled for a longer period of time to accurately analyze the information, low PRF must be selected. Higher flows can be sampled more quickly so high PRF can be selected. The preset will be adequate during the exam.

Table 4.3

Advantages and disadvantages of continuous wave

Advantages	Disadvantages
Excellent detection of high velocity flows	No range gate location
No aliasing limits	Spectral broadening

Thus, when the Doppler frequency is higher than the pulse repetition frequency, high velocities are displayed as low velocities in the opposite direction (spectral Doppler) or in the wrong color (Color Doppler). This phenomenon is known as “aliasing” (Fig. 4.10), and is directly comparable to the effect seen in movies where car wheels rotating above a certain speed appear to be turning backward.

This is probably the most common artifact occurring in Doppler studies. It simply means that the speed of the flow being examined is faster than half of the PRF set. Color aliasing can be useful in certain applications to map changes in frequency shift in a vessel with blood flowing at a constant angle with respect to the transducer. The aliasing will be displayed as the wrong color. Color changing from red, through yellow and green, to blue, as a result of changing in direction.

A correct display is possible only for Doppler frequencies within the range ± one half of the pulse repetition frequency, known as the Nyquist limit, as a consequence, increasing the PRF will unwrap the spectrum and display it in the correct direction, above or below the baseline. Doppler examination of higher velocities requires lower ultrasound frequencies and a high pulse repetition frequency, whereas low velocities can be analyzed with higher frequencies, which allow better resolution (Table 4.4).

Table 4.4

Advantages and disadvantages of pulsed wave

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