Chapter 3. Physics and Image Artifacts

Diagnostic ultrasound has experienced tremendous technological advances. Over the past 50 years, ultrasound has evolved from a single specialty tool with large bulky machines to a technology that is highly compact and portable. The development of smaller, less expensive ultrasound systems has increased the number of medical specialties utilizing ultrasound. Many are discovering the benefits of “point-of-care” diagnostic ultrasound. Medical students, nurses, mid-level providers, and physicians have embraced ultrasound as a tool to facilitate patient evaluation and improve outcomes of invasive procedures. The operator must have a basic understanding of the physical principles of ultrasound. It is these principles upon which ultrasound bases its ability to be an effective tool in medical imaging.

The simplest way to describe ultrasound is in the pulse-echo principle. Sonar can be used as an example of the forerunner of diagnostic ultrasound. A submarine that possesses sonar capability can precisely control when an acoustic pulse is generated. It assumes a relative propagation speed as it travels through a specific medium (water). The amount of elapsed time required for the “echo” to return subsequent to striking an object allows the relative distance to be calculated to the target of interest.

Diagnostic ultrasound uses the same concept of the pulse-echo principle. Electric current is passed through crystals in the transducer and generates a sound wave. This piezoelectric effect generates a constant pulse of high-frequency, longitudinal, mechanical sound waves that can be measured and used in calculations. This pulse travels at a relatively constant speed until it encounters a reflective surface, which causes a fraction of the sound to reflect back toward the transducer crystal. When the returning sound wave strikes the crystal, it generates an electrical impulse that is processed into a diagnostic image. Based on the assumption that sound travels at the same speed through all tissues (1540 m/s), a computer measures the round-trip time and intensity of each returning “echo.” The amount of time required for the returning echo determines its relative distance from the transducer while the returning intensity is proportional to the grayscale assignment of the pixel. Each returning echo is presented as a pixel (dot) of information on the display device.

Sound waves are actually a series of repeating mechanical pressure waves that propagate through a medium (Figure 3-1). These pressure waves are measured in hertz (cycles/second). Typically, audible sound ranges between 16,000 and 20,000 Hz. Ultrasound is technically defined as a “sound” having a frequency in excess of 20,000 Hz. In medicine, ultrasound used for diagnostic purposes incorporates frequencies that generally range between 2 and 15 million cycles/second, or 2 and 15 MHz, well above the range of human hearing.

Amplitude

Amplitude is the peak pressure of the wave (height). This may be simply interpreted as the loudness of the wave. Amplitude correlates with the intensity of the returning echo. A loud sound has large amplitude while a soft sound has small amplitude (Figure 3-1).

Period

Period is the time required to complete a single cycle.

Frequency

Frequency is the number of times per second the wave is repeated. The range of frequencies typically discussed here is between 2 and 15 MHz.

Spatial Pulse Length

A diagnostic ultrasound transducer generates an image by sending and receiving ultrasound waves. It receives the returning, or reflected, ultrasound waves; the generation of ultrasound waves usually occurs less than 1% of the time. The period where it generates ultrasound waves is termed a pulse. Spatial pulse length is the distance or length of each pulse. Spatial pulse length is determined by the frequency and pulse duration.

Transducer technology is based on the piezoelectric effect. Piezoelectric is defined as “pressure-electricity” and refers to materials that have a dual function of converting electric energy into mechanical energy (pressure) and conversely mechanical energy (returning echo) into electrical energy. While quartz is a naturally occurring crystal, the crystal elements in modern transducers are synthetic. The arrangement and the number of crystals within a transducer vary depending on the manufacturer, transducer design, and its intended application.

Transducer frequency has a direct effect on image quality and resolution. In general, high frequencies result in higher resolution and enhanced image quality. While resolution may increase, tissue penetration will decrease. Lower frequencies result in lower resolution, but have better tissue penetration.

Resolution is the result of the spatial pulse length as well. Higher frequency provides more reference points or pixels over a similar spatial distance and thus produces higher resolution by displaying smaller tissue segments. However, the trade-off is that the higher frequency will not travel as far or penetrate deep tissue.

Bandwidth

Historically, ultrasound transducers emitted only one frequency. As ultrasound equipment became more sophisticated, each transducer could generate multiple different frequencies, but could only send one frequency at a time. Modern ultrasound transducers emit a “center” frequency during the transmit portion of the cycle. A range of frequencies exists on either side of the center frequency and is known as the bandwidth (Figure 3-2). The resulting frequency is actually an average of the frequencies in the bandwidth.

Many ultrasound systems make use of these bandwidth frequencies during the received portion of the cycle and thus incorporate broadband transducers. Technology may allow the operator to select one of multiple “center” frequencies available from a single transducer. This selection allows the operator to easily maximize the transmit frequency of the transducer that offers the best resolution or best penetration for the area of interest. Regardless of which type of transducer technology is utilized, the highest frequency should be used that will penetrate the area of interest and that offers the best resolution.

Image resolution is based on many transducer factors including the spatial pulse length of the wave. The spatial pulse length is dependent on specific transducer characteristics set by the manufacturer. This may explain why simply increasing the transmit frequency of a transducer may not consistently result in improved resolution or improved image quality.

Velocity

Velocity of sound is defined as the speed of the wave. The velocity of sound is dependent on the material through which the wave is traveling. Velocity is independent of frequency. Since the speed of ultrasound through a given medium is constant, the closer the molecules are in position to one another, the better the propagation. Therefore, sound travels faster in bone than in human soft tissue. When molecules become less dense (gases), the velocity of the sound slows even further or may not propagate at all as is the case with a vacuum.

Wavelength

Wavelength (propagation speed/frequency) is the distance the wave travels in a single cycle.

Attenuation

The attenuation of sound begins the instant the pulse is generated within the transducer and it continues throughout its round-trip path until the sound pulse returns to the transducer to be recorded as an “echo.” There are several factors that contribute to attenuation. These factors include the wavelength of the emitted sound, the medium through which the sound is traveling, and the number of interfaces it encounters.

The type and density of tissue combined with its degree of homogeneity or heterogeneity contribute to the rate of attenuation. Tissue of the same type and density facilitates the transmission of sound. Ultrasound travels best, with the least attenuation, through homogeneous fluid-filled structures. This is why transabdominal ultrasound of the uterus and ovaries is facilitated by a distended urinary bladder. The fluid inside the bladder provides an acoustic window for the sound wave and allows an efficient use of the transmitted sound to visualize the posterior anatomy.

Reflection is a form of attenuation. It is the redirection of part of the sound wave back to its source. Reflection is the foundation upon which ultrasound scanning is based. The ultrasound beam should evaluate the anatomy of interest at 90 degrees to maximize the reflection and visualize the anatomical structures. Manipulating the transducer so the area of interest is positioned directly under the transducer in the center of the display offers improved visualization and the ability to better appreciate the surrounding anatomical structures.

Refraction is the redirection of part of the sound wave as it crosses a boundary of mediums possessing different propagation speeds. This condition worsens with non-perpendicular incidence (Figure 3-3).

Scattering occurs when the ultrasound beam encounters an interface that is smaller than the sound beam or irregular in shape. The term Rayleigh scattering is specific to red blood cells (RBC) (<5 microns) that are not normally seen unless they are clustered together making a larger surface area. This may occur when blood movement slows during stasis, early clot formation, or asystole. Spontaneous contrast is another term used to describe this phenomenon, which is best noted during real-time B-mode imaging.

Absorption occurs when the energy of the sound wave is contained within the tissue. When the acoustic energy is converted into thermal energy, it dissipates as heat within the tissue. This forms the foundation for therapeutic ultrasound. For instance, physiatrists use therapeutic ultrasound to help patients recover from injuries by using an output power significantly greater than that of diagnostic ultrasound.

Interfaces

When sound crosses a boundary of tissues having different acoustic impedance, an interface is said to occur. In emergency ultrasound, a classic interface is Morison's pouch, where the liver abuts the hyperechoic renal fascia. This interface becomes more pronounced when anechoic fluid is seen in Morison's pouch.

Acoustic Impedance

Acoustic impedance refers to the resistance of the tissue to molecular movement. Acoustic impedance is directly related to tissue density. Blood, urine, fat, and muscle all have enough difference in acoustic impedance to generate a reflection when the ultrasound wave passes from one tissue to the next. Greater density difference between tissues produces greater or stronger reflection. The intensity of the reflection (how loud the echo is) is determined by how much of a difference in impedance exists between the tissues in contact. A small density difference (acoustic impedance) results in a small echo being generated. A large difference in density results in a large echo generated with much of the energy lost to reflection. Therefore, little energy remains available to continue for visualization of deeper structures. This explains why diagnostic ultrasound cannot “see” through bowel gas or bone, as there is too large a difference in acoustic impedance between these types of interfaces and soft tissue. This phenomenon is referred to as “acoustic mismatch” and will be covered further in the section on artifacts.

Image Resolution

Resolution refers to the quality of the image produced by the machine. It is the ability to differentiate the details of anatomical structures. While there are several factors that contribute to the overall image quality, we will limit our discussion to axial, lateral, temporal, and contrast resolution.

Axial Resolution

Axial resolution is the ability to differentiate two closely spaced structures that lie in a plane parallel to the direction of sound wave propagation; that is, objects above one another on the viewing screen.

There are several factors that contribute to the quality of the axial resolution; however, the rate-limiting step under control by the operator is the ultrasound beam and transducer frequency. In general, higher transducer frequency produces better axial resolution and resultant image quality (Figure 3-4).

Lateral Resolution

Lateral resolution is the ability to differentiate two closely spaced structures that are positioned perpendicular to the direction of propagation of the ultrasound beam; that is, objects that are side by side on the viewing screen. Generally, lateral resolution will be inferior to axial resolution. The primary equipment control that improves lateral resolution is adjusting the “focal zone” to the area of interest. However, some popular point-of-care ultrasound machines do not have the ability to separately adjust the focal zone. On these machines, adjusting the depth or resolution mode will indirectly adjust the focal zone. Additionally, ultrasound beam width contributes to this side-to-side resolution. The beam may be as wide as 5–10 mm at the surface of the transducer, but tapers to as little as approximately 1–1.5 mm at the focal zone depth. If this “focal zone” is positioned on the screen adjacent to the area of interest, then the ultrasound beam will in theory be the narrowest at this point, allowing for improved lateral resolution. Multiple focal zones may be selected on a single ultrasound image in an attempt to maximize the resolution at specific depths. Since this action requires additional processing time, a slower frame rate will occur and the image will begin to appear “choppy.” Single focal zone capabilities are generally sufficient for most abdominal and cardiac examinations. Multiple focal zones are often of greater value when examining superficial structures when transducer movement is at a minimum and additional signal processing time is of less concern.

Adjusting the focal zone also narrows the “in and out of plane dimension” of the beam or the “slice thickness.” The ultrasound beam is not quite two dimensional (2D), but has a thickness similar to a credit card. Adjusting the focal zone can prevent “slice thickness artifacts” that result from the beam being thicker than the structure being imaged and averaging larger volumes of tissue than intended. For instance, when trying to visualize a blood vessel (like the inferior vena cava), if the beam thickness is wider than the blood vessel, it may get reflected signal from the tissue on either side of the blood vessel and the averaged signal display may make the lumen appear to have some tissue component producing a slice thickness artifact (Figure 3-5). The operator can also affect frame rate by adjusting the width of the sector from wide to narrow or vice versa. As the sector width narrows, the frame rate increases (Figure 3-6). This adjustment can be especially important when scanning to detect motion such as in early pregnancy using a wide-angled endovaginal transducer.

Temporal Resolution

Temporal resolution refers to the acquisition rate of a composite frame expressed as frame rate (frames per second) or sometimes expressed as hertz (cycles per second). A frame rate of 15 Hz is required to see structures move in “real time” such as a fetal heartbeat or other adult cardiac structures. Adjusting sector width or decreasing line density will have the greatest impact on increasing temporal resolution or frame rate because in both situations the system is scanning fewer lines that require less time, making the image appear less choppy.

Contrast Resolution

Contrast resolution refers to the ultrasound system's ability to assign a grayscale value to returning echoes of varying amplitudes. Most modern diagnostic ultrasound systems use 256 shades of gray, which allows recognition of subtle changes in tissue density. A higher contrast (less shades of gray) may be more pleasing to the human eye but may, in fact, contain less diagnostic information. This dynamic range of information (measured in decibels or db) is often a programmable feature and may be examination specific. The optimum setting allows a clear differentiation between the area of interest and the surrounding anatomy and typically the more shades of gray the better the contrast resolution. Abdominal imaging often uses a wide dynamic range with many shades of gray. Cardiac imaging uses a narrow dynamic range to create clarity between the moving endocardium and the anechoic blood (Figure 3-7).

A-mode, or “amplitude-mode,” provided one of the original evaluations of the human body using sound. A-mode ultrasound included an oscilloscope display for returning amplitude information and the traditional image did not exist. The peak amplitude information on the horizontal axis provided information regarding the strength or “loudness” of the wave, while the vertical axis provided reflector distance information from the transducer (Figure 3-8A). A-mode imaging is used in ophthalmology when measuring precise distances to the retina for calculating intraocular lens implantation. A-mode is generally not used in emergency ultrasound.

B-mode, or “brightness-mode,” converts the amplitude of the returning echo into a grayscale image allowing better correlation with anatomical structures. This “2D” tomographic slice is what most emergency ultrasound exams use and is referred to as B-mode imaging (Figure 3-8A).

M-mode, or “motion-mode,” permits a simultaneous display of the 2D B-mode image and a characteristic waveform (Figure 3-8B). This waveform depicts the motion or deflection of the tissue relative to the transducer on the vertical axis and represents time or changes in the cardiac cycle on the horizontal axis. M-mode technology can be of value in the emergency and acute care setting during pregnancy examinations and permits measurement and documentation of fetal cardiac activity. It can also be useful to demonstrate timing of events during changes in the cardiac cycle such as identifying right ventricular diastolic collapse secondary to cardiac tamponade. Similarly, documentation of lung sliding can also be achieved using M-mode.

Doppler

Doppler is presented in a few different forms. Doppler technology relies on the interpretation of the “frequency shift” that exists between the transmitted and received Doppler signal, while the anatomy (blood within the vessel) is moving as it is imaged. For example, as a train whistle is engaged, the pedestrian at the crossing will experience an increase in the pitch (Doppler shift) of the whistle as the train approaches and a decrease in the pitch as the train continues to move away.

Doppler ultrasound technology makes use of this “frequency shift.” Sources (e.g., RBC) moving toward the receiver (transducer) produce a higher reflected frequency, while sources moving away produce a lower reflected frequency, allowing the system to display flow direction and velocity. The angle of interrogation of blood flow is a prominent factor in the quality and accuracy of the Doppler signal.

Spectral Doppler provides a characteristic waveform and allows a quantitative assessment for blood flow analysis consisting of continuous wave or pulsed wave technologies. Pulsed wave Doppler produces short bursts of sound. It uses the same crystal to generate and receive the signal. By the transducer listening at specific intervals, it can control precisely where the reflected sound is coming from and thus has “range gate resolution.” The limitation of pulsed wave techniques is that it can only display certain maximum or peak velocities (known as the Nyquist limit) before the signal will alias and become nonquantifiable. The maximum velocity it can display is limited by the transducer's Doppler frequency and the depth of the moving target. Aliasing or wraparound occurs if the velocity is too high to display. The peaks are cut off and displayed below baseline. Continuous wave Doppler uses different crystals to send and receive signals. One crystal constantly sends signals while another receives the reflected signal. Therefore, a live B-mode image is not seen on the viewing screen. Continuous wave Doppler has no depth or range gate resolution but it does not alias and can quantify much higher velocities.

Color Doppler utilizes the pulse-echo principle to generate color images. The color image is superimposed on the 2D image. The red and blue displays provide information regarding direction and mean velocity of flow. It cannot display instantaneous peak velocities and is not truly quantifiable. The color at the top of the display represents flow toward the transducer, and color at the bottom of the display represents flow away from the transducer. It is sensitive to transducer position (Figure 3-9).

Whether using pulsed wave Doppler or color Doppler imaging, if the sound pulse strikes the blood vessel or blood flow at an angle over 60 degrees, it can result in inaccurate velocity readings and may even suggest that there is no flow when in fact it is due to poor Doppler angle.

Power Doppler images are based on amplitude or strength of the motion. Color maps in power Doppler are represented by one continuous color. It provides better sensitivity for slow flow or low blood volume states like ovarian or testicular torsion. It accomplishes this because it compares or averages several frames storing the accumulated flow over a number of cardiac cycles. This does take additional processing time and slows the acquisition frame rate. It is less angle dependent, but more sensitive to motion artifacts.

Each form of Doppler consists of benefits and limitations; however, its operation appears deceptively simple. A comprehensive discussion of Doppler physics and velocity measurements is beyond the scope of this chapter.

Two-dimensional images (also referred to as B-mode or grayscale images) are created when the computer places pixels on a two-dimensional monitor based on the direction from which the signal was received and the time interval between emission and return of the signal. Echogenicity refers to the amplitude (brightness) of the signals reflected from a given structure compared to the amplitude of the signals from surrounding structures. If a structure presents as hyperechoic, it is said to be more echogenic (of increased amplitude) than the surrounding anatomy. Conversely, hypoechoic structures appear less echogenic (of decreased amplitude). Isoechoic information has the same echogenicity as the surrounding structures. Finally, anechoic refers to the absence of echoes. Typically, fluid-filled structures appear anechoic (devoid of echos or black) (Figures 3-10 and 3-11).

Unrecognized artifacts are frequently the source of misleading information and misdiagnosis. However, some artifacts are the key to identifying certain pathologies. Accurate interpretation of ultrasound images requires recognition of anatomical reflections that we expect to visualize as well as non-anatomic signals that appear as the result of artifacts. Artifacts will be defined as any echo information that does not correspond to accurate anatomical information. The origin of these artifactual echoes may occur from within the patient, as a result of attenuation or refraction, from an external source, or as the result of operator error.

Shadowing

Acoustic shadowing is one of the most common and frequently encountered imaging artifacts in diagnostic ultrasound. Acoustic shadowing frequently occurs when the sound encounters a highly reflective (high-attenuation) surface. The reflected energy is returned to the transducer with little acoustic energy available to continue traveling to deeper structures. “Clean shadows” are caused by ribs, gallstones, and other calcified structures. “Dirty shadows” result from acoustic mismatch at tissue–air interfaces and are most commonly caused by bowel gas (Figures 3-12 and 3-13).

Refraction—“Edge Artifact”

Shadows may also occur because of changes in the ultrasound beam. A change in the sound beam direction results when there is an oblique incidence of the sound beam as it crosses a boundary of tissue with different propagation speeds or strikes a curved structure. As the sound crosses this boundary, a change in the beam direction occurs and an acoustic shadow results (Figure 3-10 and Figure 3-14).

Acoustic Enhancement

Sound waves traveling through areas of lower attenuation cause posterior acoustic enhancement. The effects of acoustic enhancement are quite simply the opposite of high-attenuating objects. As sound passes through ultrasound-friendly structures (e.g., simple cysts, distended normal gallbladder filled with bile, urinary bladder, and some types of solid tissue), less attenuation of the signal occurs, which results in a greater amount of acoustic energy available to continue its journey along the same path. This increase in acoustic energy results in a similar increase in echogenicity (hyperechoic) immediately posterior to the area where less attenuation of the signal occurred. This type of imaging artifact is commonly used to confirm the presence of areas suspected to be fluid-filled, such as hemorrhage, joint effusions, and tissue necrosis or abscess and can be used to see the spread of anesthetic as it is injected. It is most commonly referred to as increased through-transmission or posterior acoustic enhancement (Figure 3-15).

Gas

The presence of gas is often the enemy of ultrasound. The large difference in density that exists between gas and soft tissue disperses the sound waves. Diagnostic ultrasound is not equipped to handle large differences in density and much of the acoustic energy is scattered with little or no appreciable diagnostic information visualized. Often, bowel gas can prevent visualization of some portion of the abdominal aorta. Transducer pressure or a change in patient and/or transducer positioning may minimize this obstacle (Figure 3-16).

Reverberation

When sound “bounces” between two highly reflective objects, reverberation artifacts appear as recurrent bright arcs displayed at equidistant intervals from the transducer. This artifact will frequently appear when performing procedural guidance, with distal replication of the needle. This reverberation artifact may be severe enough to obliterate the B-mode image distally. Adjusting the depth, time gain compensation (TGC), transducer positioning, patient positioning, or transducer transmit frequency may reduce the appearance of reverberation artifacts. However, their presence is not likely to be confused with a pathologic condition (Figure 3-17).

Reverberation is also seen when imaging the pleural line of the lung. In the setting of normal lung, this reverberation produces horizontal hyperechoic lines distal to the pleural line, termed A-lines. In the setting of the interstitial syndrome, the dominant reverberation artifact becomes B-lines, or comet-tail artifacts, that originate from the pleural line and extend to the far field of the image (see Trauma, Figure 5-15).

Mirroring

Mirror artifacts are displayed as objects that appear on both sides of a strong reflector. These artifacts can be confusing in appearance and occur because of changes in the reflected beam. Ultrasound assumes that sound is traveling in a straight line and the distance (or depth) of the reflector is proportional to the travel time necessary to make the return trip. When the ultrasound beam undergoes multiple reflections as it returns, an incorrect interpretation of the signal timing ensues and results in a duplication of structures. The more posterior echo information is the “false” echo. Mirror artifacts are common around the diaphragm and may depict hepatic structures appearing on both sides of the diaphragm. During the FAST exam, mirror artifact of the liver or spleen above the diaphragm argues against a pleural effusion or hemothorax. Changes in transducer or patient positioning should alleviate the potential for misdiagnosis (Figure 3-18).

Side Lobes

It is easy to envision one ultrasound beam being emitted from the transducer as the thickness of a credit card. In reality, ultrasound beams of lower intensity, called “side lobes,” may originate at angles to the primary beam and are generally of little consequence. Highly reflective interfaces return echoes via this pathway of side lobes and present as false information. This false information may be introduced as an oblique line of acoustic reflection. Changes in the scanning angle often confirm these returning echoes as side lobes (Figure 3-19).

Ultrasound system and transducer performance may change over time. Diagnostic ultrasound phantoms exist in part to allow a reproducible standard to be documented, to ensure that the ultrasound system and its components are operating at the performance level defined by the product manufacturer. Ultrasound phantoms allow evaluation of transducer parameters, measurement calibration, focal zone, axial and lateral resolution, sensitivity, functional resolution, grayscale displays, transducer beam profile, and Doppler measurements.

Risks associated with diagnostic ultrasound are merely theoretical and clinically significant adverse effects have not been reported. Nonetheless, the American Institute of Ultrasound in Medicine (AIUM) has adopted an acronym termed ALARA: As Low As Reasonably Achievable. The term reminds operators that the amount of time for an examination, along with equipment settings, contributes to prudent use of diagnostic ultrasound. The receiver controls (i.e., TGC and Gain) need to be optimized before increasing acoustic power capabilities in an attempt to secure or improve the desired image. Some point-of-care ultrasound machines with simplified controls do not allow the user to independently adjust the acoustic power. The power may be automatically reduced when certain presets are selected such as “Early Obstetrical” setting.

Further reading materials include the AIUM publications listed below on ultrasound safety and are available by contacting AIUM:

American Institute of Ultrasound in Medicine

www.aium.org

800-638-5352

14750 Sweitzer Lane Suite 100

Laurel, MD 20707

Ultrasound Safety Publications

Bioeffects and Safety of Diagnostic Ultrasound

Evaluation of Research Reports: Bioeffects Literature Reviews (1962–1982)

Evaluation of Research Reports: Bioeffects Literature Reviews (1985–1991)

Evaluation of Research Reports: Bioeffects Literature Reviews (1992–1998)

How to Interpret the Ultrasound Output Display Standard for Higher Acoustic Output Diagnostic Ultrasound Devices (technical bulletin)

Mechanical Bioeffects from Diagnostic Ultrasound: Consensus Statements

Medical Ultrasound Safety

What You Should Know about the Safety of Your Ultrasound Examination (patient pamphlet).