- Pub. Date:
Related collections and offers
|Publisher:||Cambridge University Press|
|Sold by:||Barnes & Noble|
|File size:||30 MB|
|Note:||This product may take a few minutes to download.|
About the Author
Bret P. Nelson MD, RDMS, FACEP is Director of Emergency Ultrasound and Associate Director, Emergency Medicine Residency Program, Department of Emergency Medicine, Mount Sinai School of Medicine, New York, NY, USA.
Read an Excerpt
Cambridge University Press
978-0-521-68869-7 - Manual of Emergency and Critical Care Ultrasound - Author's by Vicki E. Noble, Bret Nelson and A. Nicholas Sutingco
To become versed in the language of ultrasonography, it is necessary to review some of the basic principles of physics. The wave physics principles of ordinary (i.e., audible) sound apply to ultrasound (US) and its applications. Thus, to create a foundation for further discussions, a number of definitions and basic concepts are presented here.
Basic Definitions and Physics Principles
Amplitude is the peak pressure of the wave (Figure 1.1). When applied to ordinary sound, this term correlates with the loudness of the sound wave. When applied to ultrasound images, this term correlates with the intensity of the returning echo.Image not available in HTML version
Figure 1.1 Low‐ and high‐ amplitude sound waves.
Ultrasound machines can measure the intensity (amplitude) of the returning echo ; analysis of this information affects the brightness of the echo displayed on the screen. Strong returning echoes translate into a bright or white dot on the screen (known as hyperechoic). Weak returning echoes translate into a black dot on the screen (known ashypoechoic or anechoic). The “gray scale” of diagnostic ultrasonography is the range of echo strength as it correlates to colors on a black–white continuum (Figure 1.2).Image not available in HTML version
Figure 1.2 Most ultrasound machines have 256 shades of gray that correspond to the returning amplitude of a given ultrasound wave.
Velocity is defined as the speed of the wave. It is constant in a given medium and is calculated to be 1,540 m/s in soft tissue (i.e., the propagation speed of soft tissue is 1,540 m/s). Using this principle, an ultrasound machine can calculate the distance/depth of a structure by measuring the time it takes for an emitted ultrasound beam to be reflected back to the source (Figure 1.3). (This is likened to the use of sonar devices by submarines.)Image not available in HTML version
Figure 1.3 (a) The near field of the screen shows objects closest to the probe. (b) The far field of the screen shows images further from the probe. Courtesy of Dr. Manuel Colon, University of Puerto Rico Medical Center, Carolina, Puerto Rico.
Frequency is the number of times per second the wave is repeated. One Hertz is equal to one wave cycle per second. Audible sound has frequencies from 20 to 20,000 Hz. By definition, any frequencies above this range are referred to as ultrasound. The frequencies used in diagnostic ultrasound typically range from 2 to 10 MH
Figure 1.4 shows that high‐frequency sound waves generate high‐resolution pictures. High‐frequency sound waves use more energy because they generate more waves, which send back more echoes over short distances to the machine, creating detailed pictures of shallow depth. However, because they lose energy more rapidly, high‐frequency ultrasound does not penetrate long distances. Conversely, lower‐resolution waves conserve energy, and although not creating pictures of equally high resolution, they are able to penetrate deeper into tissue.Image not available in HTML version
Figure 1.4 Low‐ and high‐ frequency sound waves.
Wavelength is the distance the wave travels in a single cycle. Wavelength is inversely related to frequency because of the principle velocity = frequency × wavelength. Therefore, high frequency decreases wavelength (and thus penetration), and lower frequency increases wavelength (and thus penetration).
Attenuation is the progressive weakening of a sound wave as it travels through a medium. Following is the range of attenuation coefficients for different tissue densities in the body :Air 4,500 Poor propagation, sound waves often scattered Bone 870 Very echogenic (reflects most back, high attenuation) Muscle 350 Echogenic (bright echo) Liver/kidney 90 Echogenic (less bright) Fat 60 Hypoechoic (dark echo) Blood 9 Hypoechoic (very dark echo) Fluid 6 Hypoechoic (very dark echo, low attenuation)
Several factors contribute to attenuation : the type of medium, the number of interfaces encountered, and the wavelength of the sound. Diagnostic ultrasound does not transmit well through air and bone because of scatter and reflection. However, ultrasound travels well through fluid‐containing structures such as the bladder. Attenuation also occurs as sound encounters interfaces between different types of media. If a tissue is homogeneous and dense, then the number of interfaces is reduced and less attenuation occurs. If a tissue is heterogeneous and less dense, then more attenuation occurs.
Reflection is the redirection of part of the sound wave back to its source. Refraction is the redirection of part of the sound wave as it crosses a boundary of different media (or crosses tissues of different propagation speeds such as from muscle to bone). Scattering occurs when the sound beam encounters an interface that is relatively smaller or irregular in shape (e.g., what happens when sound waves travel through air or gas). Absorption occurs when the acoustic energy of the sound wave is contained within the medium.
Resolution refers to an ultrasound machine's ability to discriminate between two closely spaced objects. The following images represent two points that are resolved as distinct by a machine with higher resolution (the paired dots) and the same structures visualized by a machine with lower resolution (the two dots are seen as a single indistinct blob). Axial resolution refers to the ultrasound machine's ability to differentiate two closely spaced echoes that lie in a plane parallel to the direction of the traveling sound wave. Increasing the frequency of the sound wave will increase the axial resolution of the ultrasound image. Lateral resolution refers to the ultrasound machine's ability to differentiate two closely spaced echoes that lie in a plane perpendicular to the direction of the traveling sound wave (Figure 1.5). In most portable ultrasound machines, the machine self‐adjusts the focal zone (or narrowest part of the ultrasound beam) automatically over the midrange of the screen. However, some machines have a button that allows you to shift that narrow part of the beam up and down.Image not available in HTML version
Figure 1.5 Axial resolution improves with higher frequency. Lateral resolution improves with narrow bandwidth (focal zone).
Finally, acoustic power refers to the amount of energy leaving the transducer. It is set to a default in most machines to prevent adverse biologic effects, such as tissue heating or cell destruction. This is to adhere to the ALARA or “as low as reasonably acceptable” principle – meaning the lowest amount of energy is used to obtain the information clinically needed to care for the patient. Therapeutic ultrasound operates differently from the diagnostic ultrasound discussed so far in that it purposely uses the heating properties of ultrasound to affect tissue. Often, therapeutic ultrasound is used in physical therapy or rehabilitation after orthopedic injuries to help mobilize tissue that has been scarred.
Ultrasound devices all use the same basic principle for generating ultrasound waves and receiving the reflected echoes. This principle is made possible by a property that quartz (and some other compounds, natural and synthetic) possesses called the piezoelectric effect. The piezoelectric effect refers to the production of a pressure wave when an applied voltage deforms a crystal element. Moreover, the crystal can also be deformed by returning pressure waves reflected from within tissue. This generates an electric current that the machine translates into a pixel. As mentioned, this pixel's gray shade depends on the strength or amplitude of the returning echo and thus the strength of the electric current it generates.
Many different arrangements of this basic piezoelectric transducer/probe have been developed (Figure 1.6). For example, a convex probe has crystals embedded in a curved, convex array. The farther the beams have to travel, the more the ultrasound beams fan out. This reduces lateral resolution in deeper tissue. It also produces a sector‐ or pie‐shaped image.Image not available in HTML version
Figure 1.6 Curvilinear probe on left, and microconvex probe on right.
A linear array probe (Figure 1.7) has crystals embedded in a flat head. As a result, the ultrasound beams travel in a straight line. Because the ultrasound beams are directed straight ahead, a rectangular image is produced.Image not available in HTML version
Figure 1.7 Linear probe.
Probes also come in different sizes or “footprints” because sometimes you will need smaller probes to sneak through ribs or other structures that are not ultrasound‐friendly. Finally, each probe has a range of frequencies it is capable of generating. Usually, linear probes have higher frequency ranges, and curved probes have lower frequency ranges. One exception to this is the intercavitary probe used in obstetric and gynecologic ultrasound (Figure 1.8). Although it has a curved footprint, it also uses higher‐frequency ultrasound to obtain high‐resolution pictures of smaller structures close to the probe.Image not available in HTML version
Figure 1.8 Intercavitary probe.
Using the Transducer/Probe
When scanning with the transducer, use adequate amounts of ultrasound gel to facilitate maneuvering the transducer and to optimize the quality of images obtained. Any air between the probe and the surface of the skin will mean that sound waves traveling through that space will scatter and the strength of the returning echoes will decrease. In addition, several scanning planes should be used whenever imaging any anatomic structure. This means that it is always important to image structures in two planes (i.e., transverse and longitudinal) because we are looking at three‐dimensional structures with two‐dimensional images.
One of the first principles to remember is that every probe has a raised marker or indentation on it that correlates to the side of the screen with a dot, the ultrasound manufacturer's logo, or some other identifier (Figure 1.9). Objects located near the probe marker on the transducer will appear near the probe marker on the screen. Objects opposite the probe marker will appear on the other side of the screen marker.Image not available in HTML version
Figure 1.9 Screen markers are found on the top of the screen, usually on the left for emergency ultrasound applications. Courtesy of Emergency Ultrasound Division, St. Luke's–Roosevelt Hospital Center, New York, New York.
For the most part, bedside ultrasound keeps the screen marker on the left‐hand side of the screen. However, formal echocardiography is performed with the marker on the right‐hand side of the screen, so most machines have a button that lets you flip the screen marker back and forth. This manual describes all images with the marker on the left to keep machine settings constant. It is important to know this fact because echocardiographers will have different probe positions (180 degrees different) based on their different screen settings.
As one grows more comfortable with scanning, the probe and ultrasound beam become an extension of the arm (Figure 1.10). It becomes natural to understand that moving your hand a certain way yields predictable changes in the image orientation. For novice users, it is helpful to review the standard orientation of the probe. Like any object working in three dimensions, the probe (and therefore the ultrasound beam) can be oriented in an x, y, or z axis. A simple analogy would be the orientation of an airplane. An ultrasound transducer is pictured in the figure in three different orientations (short side, long side, and facing out of the page), with its beam colored green to illustrate the concept.Image not available in HTML version
Figure 1.10 Orienting the probe in three dimensions.
Pitch refers to movement up or down. For a transducer in a transverse orientation on the abdomen, this would refer to tilting or “fanning” the probe toward the head or feet. Yaw refers to a side‐to‐side turn. This would correspond to angling the same probe left or right toward the patient's flanks. Finally, roll refers to spinning on a central long axis. If this motion is done with the aforementioned probe, the transverse orientation would become sagittal. At first, focus on moving the probe in one plane at a time, and note the impact on the image. Novice users often become disoriented when they believe that they are moving in one plane but are truly twisting through multiple axes at once.
Probe Positioning When Scanning
When obtaining a longitudinal or sagittal view (Figure 1.11), the transducer is oriented along the long axis of the patient's body (i.e., the probe marker is pointed toward the patient's head). This means that you will see the cephalad structures on the side of the screen with the marker (here, on the left side).
The transverse or axial view (Figure 1.12) is obtained by orienting the transducer 90 degrees from the long axis of the patient's body, producing a cross‐sectional display. For the vast majority of indications, the probe marker should be oriented toward the patient's right. Again, if the marker is pointed to the right, the structures on the right side of the body will appear on the side of the screen with the marker.
The coronal view (Figure 1.13) is obtained by positioning the transducer laterally. The probe marker is still pointed to the patient's head so the cephalad structures are on the left side of the screen (marker side). In this view, the structures closest to the probe are shown on the top of the screen, and as the beam penetrates, the tissues furthest from the probe are on the bottom of the screen.Image not available in HTML version
Figure 1.11 Longitudinal probe position.Image not available in HTML version
Figure 1.12 Transverse probe position.Image not available in HTML version
Figure 1.13 Coronal probe position.Image not available in HTML version
Figure 1.14 Focal zone. Courtesy of Emergency Ultrasound Division, St. Luke's–Roosevelt Hospital Center, New York, New York.
Understanding the Formed Image
To review, a number of conventions have been almost universally adopted for translating the electrical information generated by the transducer into an image on a display screen. We say “almost” because, as mentioned previously, cardiologists have reversed their screen marker ; instead of placing it on the left side of the screen, they place it on the right. Because bedside ultrasound includes abdominal and other imaging, we leave the marker on the left side and teach you to hold the probe 180 degrees reversed from the cardiology standard when doing bedside cardiac imaging. By doing this, the images you create will appear the same as the cardiologists' on the screen.
Again, to obtain these conventional views, you must know the orientation of the transducer's beam. The convention is that the probe indicator or marker should be to the patient's right or the patient's head. The screen marker should be on the left of the screen (see figures in previous section).
Adjusting the Image
Some ultrasound machines allow the operator to choose where to focus the narrowest part of the ultrasound beam. By adjusting the focal zone (Figure 1.14), you can optimize lateral resolution. Focus is usually adjusted by means of a knob or an up/down button on the control panel.
Focal depth is usually indicated on the side of the display screen as a pointer. By moving the pointer to the area of interest, the beam is narrowed at that depth to improve the image quality. Not all machines allow this function to be done manually; however, some perform this function automatically at the midpoint of the screen.
Another parameter that can be adjusted by the ultrasound operator is the depth (Figure 1.15). By adjusting the imaging depth, the operator can ensure that the entire tissue or structure of interest is included on the screen. Depth is usually adjusted by means of a knob or an up/down button on the control panel. A centimeter scale is usually located on the side of the display screen to indicate the depth of the tissue being scanned.Image not available in HTML version
Figure 1.15 Depth. Increasing depth from left to right panels.Image not available in HTML version
Figure 1.16 Gain. Increasing gain from left to right panels.
The gain (Figure 1.16) control offers an additional parameter for adjusting the intensity of returned echoes shown on the display screen. In other words, by increasing the gain, you brighten the entire ultrasound field (i.e., the entire display). When you decrease the gain, the ultrasound field darkens. The gain function is somewhat akin to adjusting the volume on your stereo – it increases the overall volume but does not improve the quality of the sound. In the case of diagnostic imaging, it increases the brightness but does not increase the number of pixels per image.
A knob or up/down button on the control panel allows the operator to adjust gain. The gain function has no effect on the acoustic power.
Time gain compensation (TGC) (Figure 1.17) controls on an ultrasound machine allow the operator to adjust the gain at varying depths. Echoes returning from deeper structures are more attenuated simply because they have to travel through more tissue. Without TGC, the far field (bottom of the screen, deeper tissue) would always appear darker than the near field (top of the screen, tissue closest to probe). TGC boosts the gain on the echoes returning from the far fields. Some machines have one button that allows you to adjust the near field relative to the far field. Other machines have multiple slider levers that allow you to control the gain throughout the entire scanning depth.Image not available in HTML version
Figure 1.17 Time gain compensation (TGC). Ultrasound machines control TGC with either sliders that divide the screen into segments or buttons that allow adjustment on in the near or far field (top panels). The bottom panels show increased far field gain on left and increased near field gain on right with a well‐gained image in the middle.
There are a variety of imaging modalities used in diagnostic ultrasound.
A, or “amplitude,” mode is an imaging modality largely of historical interest, although it is used in ophthalmologic applications today (Figure 1.18). It uses an oscilloscope display for returning amplitude information on the vertical axis and the reflector distance information on the horizontal axis. There is no picture ; distance and amplitude are represented by a graph. In the following figure, the vertical axis A represents the amplitude of the signal returned to the transducer, and the depth D is calculated based on the roundtrip time of the ultrasound beam signal.Image not available in HTML version
Figure 1.18 A‐mode.
B, or “brightness,” mode is the modality we have been reviewing up to this point ; it is what we use for diagnostic imaging. B‐mode scanning converts these amplitude waveforms into an image by using the gray scale converter discussed previously. Most scanners now display images with up to 256 shades of gray, allowing for visualization of subtle differences within tissues/structures. As mentioned, the gray scale assignment of each pixel is based on the signal amplitude or strength of the returning wave from a given point.
M, or “motion,” mode plots a waveform that depicts the motion of the tissue/structure of interest relative to the transducer's image plane (line through the structure) on the vertical axis, and time on the horizontal axis (Figure 1.19). This is often used simultaneously with B‐mode scanning to study the motion of valves or to measure/document fetal cardiac activity. Many new bedside ultrasound machines are capable of performing this function.Image not available in HTML version
Figure 1.19 M‐mode.
D, or “Doppler,” mode is an imaging modality that relies on the principle of Doppler/frequency shift. Consider the example of a moving train : a pedestrian at a crossing will hear an increase in the pitch of the train whistle as it approaches and a decrease in pitch as it moves away. However, the train engineer will not hear this change in pitch – this audible shift in frequency – because he or she is traveling with the sound. Doppler ultrasound can sense the movement of the reflected ultrasound waves toward and away from the probe – this is represented either by color changes (color Doppler) or by audible or graphical peaks (spectral doppler).
The left image in Figure 1.20 shows color Doppler. The blue and red do not identify venous and arterial flow – rather, they describe whether flow is toward or away from the probe and depend on probe orientation. The legend on the left of the screen defines the directional color assignment. In this example, red flow is toward the transducer (toward the top of the screen), and blue flow is away from the transducer (toward the bottom of the screen). The right image in Figure 1.20 is an example of pulsed wave or spectral Doppler. Spectral Doppler waveforms can be helpful in identifying and distinguishing venous from arterial waveforms.Image not available in HTML version
Figure 1.20 Color Doppler (left) and spectral Doppler (right).
Power Doppler is a form of color Doppler that uses a slightly different component of returned signal and seems to be more sensitive in low‐flow states (Figure 1.21). This mode sacrifices the ability to demonstrate the direction of flow to gain sensitivity in detecting lower levels of flow. Again, many of the new bedside ultrasound machines are now capable of performing these functions, and physicians can use these capabilities to augment their diagnostic capabilities.Image not available in HTML version
Figure 1.21 Power Doppler.
We review when D‐ and M‐mode functions are useful in the applications sections.
Effects and Artifacts
Understanding image artifacts and their formation is of the utmost importance. Unrecognized artifacts can lead to misinterpretation and can undermine the utility of the bedside ultrasound exam.
Acoustic shadowing is a characteristic ultrasound effect that can aid in the diagnosis of certain conditions (e.g., cholelithiasis) and act as a hindrance to the visualization of distal structures (e.g., rib shadows) (Figure 1.22). It occurs when a sound beam encounters a highly reflective (high attenuation) surface such as bone or calcium. Shadowing appears as a hypoechoic/anechoic area deep to the reflecting structure because so few sound waves can get around or behind the highly reflective structure. Air can also cause shadowing because the ultrasound energy is scattered in all directions at the interface between tissue and air.
Reverberation occurs when the sound beam “bounces” between two highly reflective structures (Figure 1.23). It appears as recurrent bright arcs, called A lines, are displayed at equidistant intervals from the transducer. One clinically important variation on this is when sound gets trapped between two highly reflective structures that are closely opposed, such as visceral and parietal pleura. The fibrous tissue traps the sound beam, and it “bounces” infinitely back and forth such that the reflected echo is interpreted as a straight bright white echo also known as a comet tail or B line. This concept is reviewed again in subsequent chapters because a “comet tail” artifact is a normal finding in a typical lung exam. The reverberation artifacts are clinically important in Chapter .
Refraction occurs when a sound beam obliquely crosses a boundary of tissue with different propagation speeds (Figure 1.24). It appears as an acoustic shadow, originating from the point where the sound beam changes direction.
Mirror images occur when an ultrasound beam undergoes multiple reflections and an incorrect interpretation results. When the beam encounters a bright reflector (R), some of the acoustic energy is reflected backward. When this beam path encounters an object (A), information about its relative brightness is relayed back to the transducer. However, its depth is miscalculated because the machine assumes the ultrasound beam took a straight path toward the target object. Because the reflected path (solid arrows) has a longer roundtrip time than a path directly to and from the target, the machine calculates that the structure is deeper than it is. This yields a false object (B), calculated by the machine to lie along a linear path from the initial ultrasound beam. Mirroring appears as a duplication of structures, with the mirror image always appearing deeper than the real structure (Figure 1.25). The mirror image will disappear with subtle changes in position of the transducer, whereas the real image should be visible in multiple planes.Image not available in HTML version
Figure 1.22 Shadowing.Image not available in HTML version
Figure 1.23 Reverberation and comet tail artifacts.Image not available in HTML version
Figure 1.24 Refraction artifact (see arrow).Image not available in HTML version
Figure 1.25 Mirror image artifact. Block arrow shows mirror image of liver tissue superior to diaphragm.
Enhancement (or posterior acoustic enhancement) is artifactual brightness deep to an anechoic structure (commonly a cystic structure or blood vessel) (Figure 1.26). It occurs when sound crosses an area of low signal attenuation. There is an increase in echogenicity posterior to the low attenuation structures because the sound returns to the transducer with greater intensity than adjacent areas. For example, the beams on the right are uniformly attenuated as they pass through the body. They return to the transducer with far less energy (thinner arrow) than they started with. The beam in the center loses no energy as it passes through the cyst, and thus it has much more energy to return to the transducer.Image not available in HTML version
Figure 1.26 Posterior acoustic enhancement.
© Cambridge University Press