Shadows occur when an ultrasound beam cannot pass through an area deeply due to the presence of a strongly reflecting or attenuating tissue. Figure 2-7 shows acoustic shadowing caused by the stones in the gallbladder. The shadows occur in regions of high acoustic impedance mismatch, such as soft tissue / gas or soft tissue / bone interfaces. These shadowing artifacts prevent visualization of the accurate anatomy on a scan by covering it with an anechoic shadow. This may cause misdiagnosis of the tissue anatomy.


Other causes of shadowing artifacts are improper scanning techniques, improper settings, or poor ultrasound systems.


Some possible ways to reduce these artifacts include taking images from several angles, changing the lateral resolution, or decreasing the frequency to avoid missing information.


Mirror Image Artifact (a.k.a. Ghost Artifact)


Multiple reflections (reverberation) often occur in regions with high impedance mismatches, such as air/fluid or flesh/bone interfaces. The multiple reflections duplicate a true reflector when the waves from a highly reflective surface are redirected toward a second structure. The redirected waves form a replica of the original structure, which appears on the image as a second structure.


Mirror image artifacts occur in both grayscale and color Doppler imaging. The true reflector and the artifact are equidistant from the mirror plane located between the two, as shown in Figure 2-8. This violates the assumptions that (1) ultrasound waves travel in a straight line and (2) waves travel directly to a reflecting tissue and are reflected directly back to the transducer.


A ghost artifact can develop on a color Doppler image when multiple reflections occur beyond borders. Ghost or mirror image artifacts can be reduced by decreasing the overall gain or changing the beam angle. During diagnosis, mirror images should be clearly separated from actual anatomy, especially in guided needle biopsy, where samples must be taken from a specific location.


Two possible ways by which mirror image artifacts can be formed are illustrated in Figure 2-9a.


Circular structures such as cysts can cause refraction, which produces shadows on the object’s edge due to a mismatch in acoustic impedance at the boundary or interface. The change in direction (bending) results from the change in the propagation velocity of the ultrasound waves. A schematic of this process is illustrated in Figure 2-9b.


Beams generated from the edges of a single-element transducer tend to spread from the primary beam, as shown in Figure 2-10. These lobe beams can be reflected into the primary beam, adding energy to the beam’s main axis. These artifacts violate the assumption that all reflections occur in the path of the beam’s main axis.


This duplication of the true anatomy with false reflection results from strong reflections that return to the transducer. Since the machine assumes all echoes to be coming from a true anatomic structure, incorrect images are displayed together with the correct ones.


A rapidly oscillating ultrasound beam produces multiple side lobe echoes that appear on the display as a curved line equidistant from the transducer. The two most important features that distinguish a side lobe artifact from an anatomical structure are that it is equidistant to the transducer along its length and it passes through anatomical structures.


The side lobe artifact is corrected by imaging the structure in multiple directions. The artifact will not appear in all viewing directions.


While the assumption is that all reflections are in the path of the beam’s main axis, as the beam is projected from the transducer into the tissue, some of the ultrasound waves spread outward, as shown in Figure 2-11. The presence of a strong reflector along the path of the diffracted waves produces echoes that are misinterpreted as being along the beam’s main axis. The grating lobe artifact generally appears weaker than the true reflector. This obscures the actual anatomy with a false reflection.


The artifact is corrected by taking multiple views of the structure under examination. An artifact will not appear in all views.


Multipath artifacts occur when the primary ultrasound beam reflects off anatomy at an angle such that a part of the echo returns to the transducer and at the same time, another echo also reaches the transducer after reflecting off a second boundary.


The echo from the secondary reflector takes a longer path and hence longer time to get back to the transducer. Since ultrasound machines measure the depth based on the time between the transmitted signal and the received echo, the machine will perceive a longer time and depth and position the image on the wrong spot, as shown in Figure 2-12.


In this case, the assumption that the ultrasound beam travels directly to the reflector and back to the transducer is violated. This phenomenon creates what is called a propagation path error. These artifacts give rise to the incorrect axial location of an object due to longer path lengths.


Multipath reflections may form images that appear deeper or misplaced. The problem can be reduced by taking multiple views at different angles.


An ultrasound beam incident on a curved or oblique boundary is reflected in various directions. Some of the reflected waves are directed away from the transducer. This reflection is similar to the scattering process. In this case, reflectors do not appear on the image due to longer path lengths and increasing attenuation.


Figure 2-13 illustrates the possible reflections from an oblique surface in various directions. This observation contradicts the assumption that an ultrasound pulse travels directly to a reflecting boundary surface and back to the transducer.


The strength of the echo received by the transducer is less than the expected intensity, which gives false brightness, missing reflections, or an improper location of the anatomic structure. These artifacts are associated with weak, too-bright echoes or improperly located structures. This artifact can be reduced by changing the transducer angle or by using a large footprint.


While it may be convenient to assume that the beam width stays approximately equal to the transducer size, the ultrasound beam actually spreads out as it moves away from the transducer, as Figure 2-14 shows. Due to this divergence, the echoes generated from the edge of the beam appear to be coming from the center of the beam. The artifacts are most apparent when most of the beam travels through the fluid and part of it interacts with adjacent soft tissue, as shown in Figure 2-14.


While the ultrasound machine is based on the assumptions that (1) sound travels in a straight line, (2) all echoes are parallel to the transducer axis, and (3) sound waves travel at 1540 m/s in soft tissue, the ultrasound echoes may be reflected repeatedly between two highly reflective surfaces that are parallel to the primary ultrasound beam. This causes reflections that oscillate between the tissue and the transducer. The artifacts appear on the image as multiple stairways that are equally spaced apart from one another and tend to increase with increasing tissue depth.


Figure 2-15 shows an image with multiple reflections, indicated by arrows at the top. The first bright line at the top close to the transducer is the only real line image. The other bright images below the actual reflector are artifacts. Another way reverberation artifacts can occur is when the transducer behaves as another reflecting surface such that the returning echoes are rereflected back into the tissue-reflecting structure, resulting in the formation of an identical artifact located at twice the distance from the transducer.


Because of attenuation, each image formed due to subsequent echoes is weaker than the first, as shown in Figure 2-15. The artifact can be prevented by moving the transducer probe at various angles to see an area covered by the artifact.


There are various possible causes of comet tail artifacts. These are


The artifact typically appears as one or multiple solid bright “tails” parallel to the axis of the primary sound beam, as shown in Figure 2-16. The pattern may differ depending on the size, shape, and composition of the reflecting tissue structure and the scan orientation and distance from the transducer.


This artifact helps diagnose or rule out pneumothorax. If the pneumothorax is present, the air within the pleural space hinders the propagation of ultrasound waves, thereby preventing the formation of comet tail artifacts.


The major problem is that the tails cause significant attenuation such that the beam becomes significantly weak and cannot reach deeper regions of the tissue. These tails prevent the scan from imaging the underside of the reflecting structure. The artifact can be prevented by performing multiple scans at different angles to view the area obscured by the tail.


One of the other assumptions of ultrasound machine operation is that the speed of sound in soft tissue is exactly 1540 m/s. However, sometimes the waves may propagate through a medium at a speed other than that of a soft tissue. This produces the correct number of reflectors, which appear at incorrect depths. For speeds greater than 1540 m/s, the depths are underestimated due to the short transmission-reception time of the beam from the transducer to the reflector and back to the transducer. When the sound travels at a speed slower than 1540 m/s, the distances of the reflectors are overestimated.


These artifacts can be reduced by changing the beam’s angle, which may help minimize the difference in propagating speed. Speed error artifacts cannot be prevented entirely. It should be remembered that the artifact causes incorrect placement of the reflectors on the image display.


Care must be taken to identify incorrect placements to avoid misinterpreting the image. This is often achieved by taking images from different angles. If the placement of the reflectors cannot be duplicated at different angles, then the image is an artifact. These speed errors can make the image appear “split” or “cut.” The speed error degrades the quality of an image that relies on resolution, such as the differentiation of lesions and cysts and guided biopsy.


Two assumptions of the ultrasound machine are that echoes come from the main axis of the beam and that sound travels in a straight line. However, very small structures between the beamlines degrade the image detail and produce misleading echoes. The artifact can be prevented by having a higher spatial resolution or line density. Low line density per frame produces poor detail in images. Another cause is excess gain, which tends to create or obscure information and reduce lateral resolution.


The resolution can be corrected by choosing the proper gain or focal zone setting to reduce erroneous echo contrast. An appropriate setting reduces the grainy pattern.


Near field clutter is an artifact that arises from multiple noise sources. Any acoustic noise near the transducer may cause high-amplitude oscillations of the piezoelectric crystals in the transducer. It involves the near field and may hinder the identification of structures that are close to the transducer. These oscillations cause the artifacts to appear and disappear. In general, artifacts change their appearance and appear or disappear depending on the view.


The artifact occurs due to acoustic noise near the transducer, resulting from high-amplitude oscillations of the piezoelectric elements. The appearance of additional echoes from high-amplitude reflections from the transducer in the near field can overshadow the weaker echoes from true anatomic structures. Nevertheless, when viewed from multiple angles, the real tissue structures will remain constant, while the artifacts will appear and disappear as the scan angle changes.


This artifact is caused by small gas bubbles, which produce reflections after the transducer receives the initial reflection. The transducer sees the echoes as though they are coming from structures in the deeper part of the tissue. The pocket of fluid and air continuously resonates, reflecting ultrasound and creating a hyperechoic region. This process commonly occurs in regions with air bubbles and water.


The artifact generally appears as very bright and continuous parallel bands extending to the image’s bottom. Its brightness makes it hard to view the region beneath it. The effect can be reduced by moving the scanning beam at different angles.


This artifact is the opposite of the shadowing effect due to low attenuation. It is common in fluid-filled structures such as the gallbladder, the urinary bladder, or cysts due to excessive brightness. The primary causes are improper ultrasound settings and scanning techniques.


Focal enhancement is formed due to attenuation effects, improper ultrasound settings, and improper brightness, which cause the sound waves to weaken as they propagate in the medium. This is because the amplitude decreases.


The change in the propagation speed of ultrasound due to differences in impedance causes the transmitted beam to change its direction after passing through the boundary. When the beam moves from one medium to another with a higher (or lower) impedance, the transmitted beam is refracted toward (or away from) the normal at the point of incidence.


This phenomenon occurs when a beam hits the interface obliquely, creating separate beams with different propagating speeds, contrary to the assumption that sound travels in a straight path. This can create a duplication of the anatomic features. For example, refraction may cause a single feature to appear as a double feature.


A propagating artifact appears when the beam dimension is far greater than the reflector size. Another possible cause is improper elevation resolution. The artifact “creates” debris that appears in cyst images, which can lead to false diagnoses. This is because the image plane is neither extremely thin nor uniform, as assumed. Using tissue harmonic imaging and thinner transducer arrays would reduce the problem.


Transducers are available in various shapes and sizes depending on users’ needs. As discussed earlier, a transducer consists of piezoelectric crystals—not one crystal but an array of multiple crystal elements. The various arrays and transducers are discussed below, and their schematics are shown in Figure 2-17.


Modern linear ultrasound transducers contain 256 to 512 elements. Their advantage is high sensitivity when directed perpendicular to the surface under examination. However, the beam cannot be steered, which limits the field of view.


The arrays are similar to those in the linear array but are curved, which gives them the advantage of scanning a wider field of view than linear arrays.


Phased arrays have much smaller elements than those in a linear array. Typically, they contain 128 elements to transmit and receive each data line. Linear phased arrays are typically used for viewing restricted acoustic windows.


Ultrasound images are generated when the transducer transforms the reflected wave or echo from the mechanical energy of vibration into an electrical signal that is converted into an image on the display. The image can be displayed in any of the following three modes: (1) amplitude (A) mode, (2) brightness (B) mode, or (3) motion (M) mode. These are discussed in detail in the sections below.


The A-mode is the first form of image display. The depth (reflector-time relationship) is represented on the horizontal axis, and amplitude is displayed on the vertical axis. The A-mode measures the reflectivity at different depths under the transducer. It is commonly used for examining the eyes, the liver, and the brain. Low frequencies of 2–5 MHz are used for abdominal, cardiac, and brain scanning. Organs such as the eyes and peripheral blood vessels use a 5–15 MHz frequency.


In the B-mode, the echoes are represented by bright dots. A shade of gray is assigned to each echo. The brightness of the dots indicates the strength of the echoes. The position of a dot on the screen represents the reflector distance and is determined by the transducer-reflector time relationship. Many diagnoses are made in the B-mode (in black and white) with a relatively simple probe and protocol. It finds use in the study of both stationary and moving structures. The B-mode is an electronic conversion of the A-mode and A-line information into brightness-modulated dots on the display screen, as illustrated in Figure 2-18.


The B-mode display can be used for the M-mode and 2D grayscale imaging. Modern B-mode ultrasound uses both the fundamental and the second harmonic frequencies. Harmonic imaging is most useful in patients with thick and complicated body wall structures. Figure 2-19 shows anatomical structures in the B-mode.


The B-mode is also used for early intima-media thickness analysis of the carotid arteries (located in the neck) to determine the potential for lethal cardiac events. Abnormal thickening of the arterial walls of the carotid arteries is an early indicator of vascular disease throughout the body. The thicker the arterial wall, the greater the risk of heart attack or stroke.


In the M-mode, the motion of an object points along the transducer axis and is revealed by a bright trace moving up and down across the image. This display is commonly used to evaluate the morphology, movement, and velocity of cardiac valves and walls.


Imaging the pattern of moving cardiac structures over time constitutes M-mode echocardiography, as Figure 2-20 shows.


Doppler imaging is based on the Doppler effect, which shows the relationship between velocity and frequency shift. This imaging system is mainly used to measure blood flow velocity to determine any narrowing of the arteries and assess the risk of stroke occurrence.


In the sonographic application of the Doppler effect, a typical moving source would be flowing blood, and a typical receiver would be a stationary transducer. When the source is moving away, the detected frequency is lower. Conversely, a source moving closer to the receiver would have a higher detected frequency.


For a stationary source (transducer) and a receiver (target) moving with velocity (v) at an angle (θ) relative to the direction of the incident wave of frequency (fs) from the transducer, the Doppler frequency is given by


where c is the speed of sound in the tissue, which is 1540 m/s.


The transducer transmits and receives sound waves in the form of sinusoidal signal waves. The magnitude of the Doppler shift is related to the velocity of the blood cells or moving tissue, and the polarity of the shift reflects the direction of blood flow. The blood flowing toward the transducer is positive, and the blood flowing away from the transducer is negative.


The Doppler shift (Δf) is directly proportional to the velocity (v) of the blood cells, the transducer frequency (fs), and the cosine of the angle of incidence (θ) and is inversely proportional to the velocity of sound in tissue (c = 1540 m/s). In cardiac applications, the angle of incidence in the Doppler equation is assumed to be 0 or 180 degrees.


Three modalities are currently used in Doppler echocardiography, and these are pulsed wave (PW) Doppler, continuous wave (CW) Doppler, and color flow (CF) Doppler imaging. Color flow imaging evaluates the Doppler flow information for its direction toward or away from the transducer based on the color display. A commonly used acronym for remembering the color and direction is BART—blue away, red toward. The PW Doppler does not continuously transmit and receive the ultrasound pulse. Multiple crystals in the transducer are excited in a quick burst, producing ultrasound waves. This transmission burst is then followed by a “listening” period during which the crystals detect the reflected signals. Signals from more superficial structures are received sooner than those from deeper structures with more extended “listening” periods. This feature allows signals only from specific depths to be processed, thereby controlling sample size and range resolution. Therefore, two vessels located above each other can be evaluated separately, and vessels can also be followed as their courses change. The PW Doppler wave is site-specific and can only measure low-flow velocities—it cannot correctly measure high velocities (above –m/s). The CW Doppler uses two piezoelectric crystals, one to emit ultrasound continuously and the other to receive the reflected waves continuously. This results in a fixed sample size and no range resolution or ability to place the sample volume at a specific depth. It also cannot create anatomic images. It is used for precise settings such as very high peak systolic velocities.1 The CW Doppler measures very high blood flow velocities, and the color flow CF Doppler is a PW Doppler with multiple gates that allow it to measure the flow velocity through the heart on the two-dimensional echocardiographic image. Physicians often use Doppler imaging to detect blockages to the blood flow (due to clots), constriction of vessels or tumors, and congenital vascular malformations. Figure 2-21 shows an ultrasound image of the right common carotid artery and the corresponding Doppler waveform.


Duplex ultrasonography combines physiologic information based on Doppler shift frequencies with anatomic information from real-time, high-resolution B-mode imaging.


Understanding blood flow dynamics is important in the study of vascular disease development, such as atherosclerosis, thrombosis, or aneurysms. The circulation system transports nutrients and waste around the body (delivering oxygen and nutrients to the cells and removing cellular wastes and carbon dioxide). Its other function is to maintain a constant temperature and potential or power of hydrogen (pH) in all organs of the body. The circulation system comprises the heart (the pump that drives the blood to all body tissues), blood vessels (delivery routes), and blood (the medium that transports the food and the waste materials). The blood flows continuously through two separate loops that originate and terminate at the heart: the pulmonary circulation and systemic circulation loops, as shown in Figure 2-22. Pulmonary circulation carries blood between the heart and lungs, and systemic circulation carries blood between the heart and the body’s organs and tissues. At any time, about 84% of the entire blood volume is in systemic circulation, 7% is in the heart, and 9% is in the pulmonary vessels.


Under normal conditions, the average resting heart rate of an adult between the ages of 18 and 80 is about 75 beats/min, with a stroke volume of 70 mL/beat (cardiac output of L/min). For vigorous-intensity physical activity, the heart rate can increase to as high as 200 beats/min, with a stroke volume of up to 150 mL/beat (cardiac output of about 25 L/min).[1] The arteries respond to varying pressure conditions by dilating or shrinking to accommodate the hemodynamic demands.


The presence of a pressure gradient between the aorta and the veins ensures the blood keeps moving to the peripherals. In mathematical form, the volume per unit time (Q) can be expressed using Dacy’s law: Q = ΔP/R, where ΔP is the pressure differential and R is the resistance. Figure 2-23 shows the systemic blood pressure throughout different paths of the body.


As the heart pumps the blood, the pressure varies between systolic pressure (pressure peak after ventricular systole) and diastolic pressure (pressure drop during ventricular diastole). In the aorta, the systolic average pressure is about 120 mm of Hg, while the diastolic average pressure is about 80 mm of Hg.


The velocity of the flow is mainly determined by three critical variables: radius (r), vessel length (λ), and viscosity (η), which are related to one another by Poiseuille’s equation:


where ΔP is the pressure differential and resistance is given by


This shows that since the radius changes with vasoconstriction and vasodilation, the effect will be a dramatic change in the resistance and flow of blood.


The blood flow through straight, long, and smooth vessels is almost linear, with each layer of blood remaining the same distance from the walls of the vessels. These different layers flow at different velocities. Speed is dependent on both the axial distance and pressure. At high pressure, the velocity is high, whereas at low pressure, the velocity is low. This leads to a decrease in pressure and velocity from the heart to peripheral circulation.


Due to the difference in systolic and diastolic pressures, pulse pressure is generated during systole. A pulsatile blood flow is created down the pressure gradient into systemic circulation. The pulse pressure is ~ 40 mm of Hg, the difference between systolic and diastolic pressures.


In a laminar flow, the motion of the fluid is very orderly, with all particles moving in straight lines parallel to the walls of the tube, as shown on the left of Figure 2-24. The velocity profile across the tube is parabolic, with the fluid’s highest velocity at the tube’s center, as shown in Figure 2-25. The parabolic profile arises because the fluid molecules touching the walls experience more resistance than those at the center.


When the velocity of the blood becomes too high as it passes through a constricted vessel or a rough surface, the flow may become irregular, resulting in random fluctuations in position and time, leading to turbulent flow.


The tendency for turbulent flow is measured using the Reynolds number, which depends on the velocity of the flow, the diameter of the vessel, and the density of the blood:


where ν is the average blood flow velocity (in cm/s), d is the vessel diameter (in cm), ρ is density, and η is the viscosity of the blood. In these units, turbulence occurs when Re > 200, resulting in the formation of eddies. Turbulence can occur in regions of stenosis with increased flow velocity. This type of flow is not common in healthy vessels.


where PSV and EDV represent the peak systolic velocity and end-diastolic velocity, respectively.


The RI index is used as an indicator for detecting and managing renal artery stenosis, evaluating risk in chronic kidney disease, compiling differential diagnoses in acute and chronic obstructive renal disease, and predicting renal and global outcomes in critically ill patients.


Recent studies have shown that an increased RI reflects changes in intrarenal perfusion and systemic hemodynamics and the presence of subclinical atherosclerosis and, therefore, may provide valuable prognostic information for patients with primary hypertension.[2]


Ultrasound has a wide range of medical applications. These include its use in obstetrics (monitoring the progress of pregnancy), oncology (monitoring the growth of tumors), cardiology (visualizing heart function and physiology), and biopsy (guiding needles in various procedures) and as a rehabilitation modality. For example, today, an estimated 60–70% of pregnant women in the United States undergo ultrasound examinations during pregnancy. An estimated 250 million fetal ultrasound examinations are performed annually in the United States.[1] The examination aims to assess fetal abnormalities, confirm the site of pregnancy within the uterus, and determine gestational age.


Due to the increasing use of ultrasound in health care, most medical schools are transforming curricula to include medical ultrasound applications. In addition, most medical professions require successful completion of a Sonography Principles and Instrumentation (SPI) examination and some specialty examinations administered by the American Registry for Diagnostic Medical Sonography (ARDMS). The SPI examination requires a sound knowledge of the physical principles of ultrasound and imaging, which includes an understanding of the physics of ultrasound and imaging techniques. This requirement is appalling to most clinicians with no physics background. On the other hand, depending on the individual’s interest, a specialty examination will be given in one or more of the following areas: abdomen, breast, echocardiography, obstetrics, gynecology, pediatric, vascular technology, or musculoskeletal ultrasound. Most hospitals recognize ARDMS credentials as requirements for clinical sonographers and physicians.


Ultrasound images in medical imaging are generated from sound waves reflected from different tissues and organs and converted into electrical signals, which a computer processes to create an image displayed on a screen. This technology helps health care professionals visualize internal structures and diagnose medical conditions. To make it easier to understand the operation of a medical ultrasound machine, we will first discuss some basic physics principles.


A mechanical vibration is a back-and-forth motion. When vibrations affect the media around them, waves are generated. These waves transport energy from one point to another. If a single vibratory disturbance moves from one point to the other, it is called a pulse. A back-and-forth motion that occurs repeatedly is called a periodic motion.


A mechanical wave requires a material medium (such as a solid, liquid, or gas) to propagate through; its speed depends on the properties of that medium. Mechanical waves fall into two classes: longitudinal and transverse waves. For a transverse wave, the displacement of the medium is perpendicular to the direction of the motion of the wave. In a longitudinal wave, the displacement of the medium is in the same direction as wave motion. One example of a longitudinal wave is sound. These waves are similar to the motion of a pulse on a slinky, as illustrated in Figure 1-1.


A sound wave comprises alternating regions of low and high pressures. The waveform is a sinusoidal wave function in which the crests and troughs represent high- and low-pressure regions, respectively, as shown in Figure 1-2.


The maximum displacement or height from the horizontal axis, the equilibrium position, is the amplitude (A) of the wave. The distance between two successive points in the same phase is the wavelength (λ). For example, the wavelength is the distance between neighboring peaks, neighboring troughs, or any two points where the wave returns to the same shape, as shown in Figure 1-3.


The time for one complete cycle is the period (T); it is related to the number of oscillations (complete cycles) per unit time, called the frequency (f), as follows: T = 1/f. The unit of frequency is the cycle/second = 1/s or s-1. One cycle per second = 1 hertz (Hz).


The rate at which the waveform changes position with respect to time is called the velocity (v) of the wave. In mathematical form, v = f λ. The wave velocity depends on the characteristics of the medium in which it travels. The velocity has a unit of meters per second if the frequency (f) is in hertz and the wavelength (λ) is in meters. The speed of sound is fastest in solids, slower in liquids, and slowest in air. Sound does not travel through free space or vacuums due to the lack of a medium.


Sound waves are classified into three categories based on their frequency: infrasonic, audible, and ultrasonic waves.


Infrasonic waves: These are waves with frequencies of less than 20 Hz; they are not audible to the human ear. Various natural and man-made sources produce them, including earthquakes, volcanoes, thunderstorms, and industrial machinery. Animals such as elephants, whales, and alligators also produce these waves to communicate over long distances.


Audible waves: These are sound waves with frequencies that the human ear can hear. The audible frequencies fall in the range of about 20 Hz to 20 kHz. The audible range of sound can vary between individuals due to various factors. Age is one factor that can affect a person’s ability to hear some frequencies, as the ear’s sensitivity decreases with age. For example, younger people are likely to hear sounds of up to 20 kHz, while older people mostly hear frequencies of far less than 20 kHz. Certain diseases and medical conditions, such as otitis media, otosclerosis, or Meniere’s disease, can also affect a person’s hearing ability. Additionally, individuals with hearing impairments or disabilities may have a reduced hearing range due to damage to the ear or nerve pathways involved in hearing.


Ultrasonic waves (also called ultrasound): These are waves with a frequency of more than 20 kHz. Humans cannot hear ultrasound, but bats use these waves for navigation. However, medical professionals use these waves to examine or image the different parts of the human body, a practice also known as sonography. Medical ultrasound imaging typically uses the to 20 MHz frequency range.


The amount of sound energy flux per unit of time is called the sound intensity (I). For a point source generating sound with acoustic power (P), the intensity (I) at distance (r) from the source obeys the inverse square law:


The acoustic power of an ultrasound wave is the quantity of energy generated per unit of time. The standard unit of acoustic power is the watt (W), and 1 watt = 1 joule per second. Therefore, the unit of sound intensity is W/m2. The intensity equation shows that sound intensity decreases as the square of the distance from the point source. We all know that sound loudness (ear perception of sound intensity level) decreases as we move away from the source, as illustrated in Figure 1-4.


The sound intensity level (also called the sound acoustic level) is commonly measured relative to the standard threshold of hearing intensity (Io) in decibels. A decibel is a dimensionless quantity (no units) represented as dB, which is based on the logarithmic scale. In mathematical form, the sound intensity level (ß) is expressed as


where Io = 10-12 W/m2, which is the faintest audible sound intensity.


Clinicians examine body tissue structures using ultrasound waves with 2 to 20 MHz frequencies. Inside the tissue structures, the waves propagate through the medium by vibrating molecules of the medium. In soft tissues, the propagation velocity is relatively constant at 1540 m/s. This velocity value is used by ultrasound machines for all human tissue. Ultrasound waves propagating through tissue undergo reflection, refraction, attenuation, scattering, and diffraction.


Like any wave, ultrasound waves are reflected at tissue boundaries and interfaces. The transducer detects these reflected waves, and piezoelectric signals are generated and processed into an image form via a computerized processing unit. These signals form the basis of all ultrasound imaging. The number of reflected waves detected by the transducer depends on the angle of incidence at the band boundary and the difference in acoustic impedance between the two tissues traversed by the beam. More details about the acoustic impedance will be discussed later. However, it represents the resistance of a tissue to the passage of ultrasound. Typically, a propagating ultrasound wave is split into two components, as Figure 1-5 illustrates.


If the wave traverses from medium 1 (with acoustic impedance Z1) to medium 2 (with acoustic impedance Z2), the reflection coefficient is


where θi is the angle of incidence, θr is the reflected angle, and θt is the angle of transmission. Pr and Pi represent the reflection and incident probability amplitudes, respectively.


Reflections can also be classified into two categories: specular and diffuse, as illustrated in Figure 1-6.


The ultrasound beam that succeeds in penetrating the boundary layers or interface is called the transmitted wave. The transmission coefficient is mathematically expressed in the following form (refer to Figures 1-5 and 1-7):[2]


The ratio of the speed of the transmitted wave (v2) to that of the incident wave (v1) is related to the ratio of the sines of the angles of transmission and incidence, a relationship called Snell’s law:


When the waves are reflected from a perfectly flat surface or boundary, the reflected waves tend to be uniformly parallel to each other. In contrast, they tend to be diffuse for rough surfaces. This phenomenon is commonly observed as a “mirage” when driving on a hot summer day, and the road appears to have a wet surface that disappears as one gets closer. This leads to two different kinds of reflections, specular and diffuse, which are illustrated in Figure 1-6.


The transducer picks up the reflected waves and converts the echoes into images. The strength of the echoes depends on the acoustic impedance between the two tissues through which the waves pass. Typically, boundary reflections occur on blood vessel walls and organ boundaries.


When an ultrasound beam strikes a tissue boundary obliquely, the transmitted component of the beam undergoes a change in direction. This change is due to the differences in the velocities of the incident and transmitted beams. This bending process, called refraction, is illustrated in Figure 1-8 and is often related to the formation of artifacts during ultrasound image acquisition.


In ultrasound imaging, refraction can result in the formation of artifacts such as double image artifacts, as shown in Figure 1-9. This artifact is caused by the differential refraction of the ultrasound beam while passing through the relatively different echogenic tissues, such as muscle and fat tissues, and the difference in velocities in those tissues.


Inside the human body, scattering is mostly a result of small changes in the density, compressibility, and absorption properties of the tissues. Ultrasound scanners detect these scattered waves to show the backscattered signal in the form of images. Scattered waves are primarily caused by red blood cells, or erythrocytes, due to their relatively small diameter of 6–8 µm, compared to the commonly used ultrasound wavelength of about mm. For ultrasound frequencies below 20 MHz, the backscattering signal from blood is about 10 to 27 dB lower than from the surrounding tissue.[3] This difference makes it possible to image the blood flow inside the tissue.


As the ultrasound beam traverses the tissue structures, it also loses energy through absorption. The lost energy varies depending on the tissue’s characteristics and the ultrasound wave’s frequency. For example, bones absorb more ultrasound energy than soft tissue does. Absorption of ultrasound in tissue is frequency dependent—it increases with increasing frequency.


As the ultrasound moves through tissues, some of the ultrasound energy is lost due to absorption through heat, reflection, refraction, and scattering. The beam weakens with increased depth into the tissue, increasing acoustic impedance mismatch. Another factor is the presence of air bubbles inside the tissue, which tend to form virtually impenetrable barriers to ultrasound. Attenuation becomes higher not only with increasing distance from the transducer but also because of the heterogeneity caused by acoustic impedance mismatch as well as the higher frequency of the transducer, as illustrated in Figure 1-10. This is because air has a higher resistance to ultrasound propagation than fluids.


The intensity (Ix) of an ultrasound beam at tissue depth x can be estimated using Beer’s law:


where Io is the incident intensity at the tissue surface and μ is the intensity attenuation coefficient. Of this attenuation, absorption contributes about 60–80%.[4]


Attenuation increases with increasing gas and fat. The higher the tissue density (or impedance), the lower the reflection. For example, blood has an attenuation coefficient value closer to dB/MHz.cm, while the typical value for bone is around dB/MHz.cm.2


Attenuation generally increases linearly with increasing frequency among different body tissues. Fluid-filled structures have much lower attenuation than solid structures. Hence the transmitted pulse from a fluid-filled structure is usually more substantial than that from passing through an equivalent amount of solid tissue.


The ultrasound beam spreads out with distance from the transducer as it passes through the tissue, causing diffraction, as shown in Figure 1-11. This results in the reduction of beam intensity.


This diffraction pattern is highly dependent on the shape and size of the transducer relative to the wavelength of ultrasound. This phenomenon causes a decrease in the intensity of the ultrasound beam. To achieve a parallel beam, the diameter of the crystal face is designed to be approximately 10 to 20 times the wavelength of ultrasound.


Acoustic impedance (Z) is the resistance of a tissue to the passage of ultrasound. It depends on the density of the tissue (ρ) and the velocity of propagation (v) of the ultrasound through the tissue: Z = ρv. If the density is in kg/m3 and the velocity is in m/s, then the specific acoustic impedance is expressed in the unit of rayl (Ry), which is equivalent to 1 kg/(m2s). A typical value of acoustic impedance is rayls for air, rayls for the liver and blood, and around 5 rayls for the bone.[5]


The greater the difference between two tissues (media), the more the ultrasound is reflected and the lesser the transmission through the tissue. For example, more ultrasound beams will be reflected at soft tissue / bone and soft tissue / air interfaces than at soft tissue / blood interfaces. The difference in acoustic impedance is called acoustic impedance mismatch. A greater acoustic mismatch results in a greater reflection and a lower transmission.[6]