Understanding Ultrasound - The Basics - SD

Ultrasound Fundamentals

Further topics will include discussion of attenuation, artifacts, and harmonic imaging.

Ultrasound is a form of mechanical energy. It is a series of pressure waves, which are propagated through matter. Sound is an example of such a mechanical, a pressure wave, which is detected by the ear. Ultrasound uses much higher frequencies of pressure waves in order to produce the pressure waves which are propagated in tissue.

Piezoelectric Effect

Certain materials are used which respond to electrical stimulation by changing size. This property of materials to respond to an electrical potential by changing their size is called the piezoelectric effect, and was first described in the 1890s by Pierre Curie. Crystalline materials such as quartz and a variety of artificial materials will change their crystal size when stimulated by small electric potential. This change in size then produces pressure, which can be transmitted or propagated into tissues. The same materials which produce the pressure, which forms the ultrasound waves, which are propagated, also serve as the detectors of the returning pressure waves. The echoes of the back scattered ultrasound, energy.

When these pressure waves strike the pazo electric material, they induce a tiny, voltage change, which can be amplified and used ultimately to generate the two dimensional ultrasound image.

Ultrasound Transducer

The typical ultrasound transducer consists of multiple piezoelectric elements, often several hundred, and this image, the small linear structures are individual transducer elements, each of which is provided with wires which conduct the small voltage to stimulate the transducer or to respond to pressure changes to amplifiers, which then, convey the information to the ultrasound machine where it is processed to produce the 2D ultrasound image.

Ultrasound Wave Propagation

The mechanical pressure waves of ultrasound are propagated in tissues at the molecular level. When pressure is applied to tissues, the molecules are slightly displaced, and as a series of pressure waves are passed, this pressure propagates as a series of alternating areas of high pressure and low pressure.

We can define these changes in pressures that occur with ultrasound in several ways. First, we can measure the spatial distance between corresponding portions of the pressure wave. This is typically measured in millimeters and constitutes the wavelength of the ultrasound. We can also measure the length of time it takes to pass an ultrasound pressure wave through a given point. This is measured in seconds or fractions of a second, and here we've illustrated a pressure wave, which has a duration of one 10 million of a second. This interval is called the period of the ultrasound wave. The period is not commonly used because the numbers are difficult to manage, but instead we use the reciprocal of the period, which is called the frequency. In this example, a period of one 10 million of a second is equal to 10 million cycles per second or 10 megahertz.

The frequencies used in diagnostic ultrasound are quite high, Generally ranging between one and 20 megahertz. This is compared to the frequencies of audible sound, which are in the range of several hundred to several thousand hertz. Each ultrasound frequency has a specific wavelength. In the case of diagnostic ultrasound, a frequency of one megahertz has a wavelength of approximately 1.5 millimeters, and a frequency of 15 megahertz has a wavelength approximately 0.15 millimeters.

The mechanical energy of pressure waste, produced by ultrasound propagates through the body at a constant velocity. This velocity is equal to the product of the frequency and the wavelength, so that if the velocity is constant, an increase in frequency will be accompanied by a decrease in wavelength. And conversely, a decrease in frequency will be accompanied by an increase in wavelength as we'll see later. The wavelength is one of the important determinants of resolution of an ultrasound device.

Propagation Velocity in Tissues

Sound travels at different speeds and different materials and air. The propagation velocity of sound is relatively slow in the range of 330 meters per second in the body. Different tissues conduct sound at considerably higher velocities fat at a velocity of 1,450 meters per second. Fluid or cyst content at about 1,480 meters per second. With an average propagation velocity of sound and soft tissues of 1,540 meters per second, Blood propagates sound at a slightly faster velocity. The ultrasound machine must assume a specific velocity for the propagation of sound in order to properly place echoes in their respective locations within an image. And for this purpose, the machine assumes the sound will travel 1,540 meters per second.

Pulsed Ultrasound and Imaging

The pressure produced by an ultrasound transducer is propagated in a series of pulses at high velocity through tissue with relatively long intervals separating the pulses of ultrasound energy. And it is actually possible to produce ultrasound in two fashions, one in a continuous fashion in which the transducer has emitted a continuous stream of, ultrasound pressure waves or, and a transducer that emits short brief pulses of ultrasound energy separated by relatively long intervals. For imaging, very short pulses are used generally consisting of two to three cycles of ultrasound separated by very long intervals in which no sound is transmitted. The use of pulse ultrasound is essential for imaging applications.

In imaging, the transducers produces a small pulse of energy, which travels into the tissue until it encounters a reflecting interface, causing a small amount of the energy to return to the transducer where it is detected and displayed In order to produce the image.

Echo Ranging Principle

The ultrasound machine relies on a very simple principle called echo ranging. In a sense, the ultrasound machine functions as no more than a sophisticated, stopwatch. We consider a transducer at the surface of a piece of tissue with an object at some unknown depth, beneath the surface and fire a pulse of ultrasound from the transducer. It will take a certain finite length of time for the pulse to travel to the area of interest. It'll take a similar length of time for the sound to travel back to the transducer where it's detected. In this example, we've indicated that it takes about 0.145 milliseconds from the production of the ultrasound pulse by the transducer until the echo returns and is detected. If we can measure this time very accurately, then it becomes a simple matter of arithmetic to calculate how far the sound pulse has traveled. In this case, the total distance travel is about 22.4 centimeters. Since this represents the total distance from the transducer to the object and back to the transducer, if we divide by two, we determine the precise depth of the object from the transducer. By making hundreds of thousands of these measurements throughout the field of view, we can generate a precise position of each of the reflectors within the object and use the position information obtained in this way to generate the two dimensional ultrasound image.

Ultrasound Display Modes

The simplest mode of ultrasound was utilized, over 50 years ago. In this case, the ultrasound transducer produced a pulse of energy accompanied by a voltage spike, which is shown in the left of the oscilloscope, diagram. The pulse then travel to the object, encountered an interface from which the sound was reflected. The reflected energy then produced a small voltage change in the transducer, which is displayed on the oscilloscope, as a vertical deflection. Additional sound traveled deeper into the object until another interface is encountered producing a second voltage change. Somewhat weaker because the interface is weaker and positioned, somewhat delayed in time because the interface is further from the transducer. Notice that time and distance are equivalent as far as the ultrasound machine is concerned because the machine interprets a delay in time as indicative of a greater distance of the reflecting interface from the transducer. And this example, the process is, repeated again until the third interface is detected, and the final display shows three voltage spikes on the oscilloscope screen. The height of the spikes indicating the strength of the reflector, the amplitude, and therefore this is called amplitude mode or a mode ultrasound.

Now by replacing the vertical deflection of voltage on the zillo scope screen with small dots which vary in brightness according to amplitude, we create the DB mode or brightness mode of display. In this case, each of the interface is represented by a bright dot, which varies in proportion to the strength of the interface. For example, the small low amplitude spike in the middle of the object appears darker and brightness than the stronger interfaces on either side. If we take multiple measurements using B mode display by moving the transducer around the object, we can draw in a two dimensional image of the interfaces, formed by the object. And in fact, this is how early B mode ultrasound, was first used by manually moving a transducer over the body, drawing in and scanning in each of the reflectors to produce the two dimensional image. Today, beam mode is used in an automated fashion by transducers that steer the beam to accomplish the same task previously accompanied manually.

M mode ultrasound is just a variation of B mode ultrasound in which the position of reflectors is monitored over time so that if an object moves As the scan is being performed and this motion is viewed over a interval of time, this can be displayed as a BM mode image, showing movement of structures along a single line of sight over an interval of time. In this case, the motion of the, intraventricular septum of the fetal heart is shown clearly moving.

Image Resolution

This is the second in a series of, presentations dealing with basic ultrasound principles. The emphasis in this presentation will be image resolution. High quality imaging requires several attributes. One of the more important of these is spatial resolution as well as contrast resolution, temporal resolution and freedom from artifacts.

An ultrasound, the consideration of resolution must take into account resolution in three planes along the axis of the ultrasound beam, so-called axial resolution, as well as in the slice thickness plane and in the plane perpendicular to the axis of the beam.

Axial Resolution

Axial resolution refers to the ability to identify as separate two objects that lie, one behind the other, along the axis of the ultrasound beam. In this case, two small bright echoes are seen in each of the images, one behind the other representing small microcalcifications in the carcinoma of the breast. It is the axial resolution of the ultrasound system that provides the ability to identify these as two separate calcifications rather than seeing them merge together as one single echo. Axial resolution is primarily determined by the length of the ultrasound pulse. The ultrasound pulse consists of two or three cycles of ultrasound, and the overall length therefore is determined by the wavelength and the number of pulses. If the ultrasound pulse is relatively long compared to the distance that we are trying to resolve, the objects will blend together and produce a single, echo on the final display. If the ultrasound pulse is shorter than the distance between the objects, it then becomes possible to identify each of the objects separately and have them displayed as separate in the image.

Here we can compare the actual resolution at five and 15 megahertz. The line indicates three objects, one located at the origin, another at four millimeters, or four tenths of a millimeter and a third at one millimeter at five megahertz. The wavelength is approximately 0.3 millimeters or 300 microns, and the ultrasound pulse is a little less than a millimeter. In this case, it would be possible to resolve the objects that lie a millimeter apart, but not to resolve objects that lie much less than a millimeter from one another. At 15 megahertz, the wavelength is only 0.1 millimeters and the pulse is approximately 0.3 millimeters. Here it becomes possible to resolve not only the objects that lie a millimeter apart, but objects that may lie as close together as three or four tenths of a millimeter. This again, illustrates the high resolution, capabilities of a high frequency transducer. In this case, a 15 megahertz transducer evaluating very small and closely placed calcifications within a breast cancer.

Lateral Resolution

Now, in addition to axial resolution, the quality of the ultrasound image is also affected by the width of the ultrasound beam, and as we can see in this photograph, the ultrasound beam, which is represented by the pattern emerging from the transducer on the left, changes in shape. As we move away from the transducer, it becomes very narrow, and then the, field expands again. This point of maximal narrowing of the ultrasound field is called the focal zone. It determines the lateral resolution of the ultrasound system. The lateral resolution in turn determines the ability to resolve a separate two objects that lie side by side and the plane perpendicular to the direction of the ultrasound beam. Here again, looking at examples of small calcifications when they're breast cancer, we can see that these tiny calcifications in each example lying side by side are clearly resolved as two discrete objects rather than merging together as one. This illustrates the lateral resolution capabilities of this system. Lateral resolution is extremely important because it is controlled by the user by precisely timing the firing of the ultrasound elements. The point of greatest narrowing of the ultrasound beam, can regulated. It can be placed closer or further to the transducer. Notice that outside of the focal zone point B in this image, the ultrasound film field becomes wider and therefore the lateral resolution diminishes in these locations. In addition, the transducer typically has a dead zone very near the transducer in which very little useful information can be obtained due to interference of the, ultrasound energy at this location,

The placement of the focal zone is one of the most important user controls affecting overall image quality. Normally, focal zone placement is displayed as an arrow or marker along the depth axis of the ultrasound device and the image on the left. The focal zone is placed, near the bottom of the image, not in an appropriate location to provide optimal resolution of the, soft tissue mass. In the middle of the screen on the right, the focal zones are placed, at the level of the, area of interest and provide the optimal image quality at this depth.

Elevation Resolution

In addition to axial and lateral resolution, the thickness of the ultrasound slice also should be considered. This is called the elevation, plane or elevation resolution. And as with the CT and magnetic resonance imaging slice thickness can have an important impact on the overall quality of the, final image and an ultrasound machine. The elevation focus or slice thickness is designed into the ultrasound transducer and cannot be modified or altered by the user depending on the expected application of the transducer. The elevation focus may be relatively shallow in this example, at a depth of about 1.5 centimeters. Typical depth for a small parts transducer used to evaluate relatively superficial structures, whereas a transducer use for abdominal or pelvic ultrasound will have the elevation focus at a considerable deeper depth.

Ultrasound Artifacts

The third portion of this discussion of, basic ultrasound imaging will deal with important ultrasound artifacts. The ultrasound machine makes a number of assumptions when it, creates an image, and when these assumptions are not correct, artifacts may arise.

Beam Width and Side Lobes

One assumption that the machine makes is that the ultrasound beam is narrow and uniform in width. As we've seen in the discussion of lateral and elevation focus, this is not the case. The width of the ultrasound beam varies depending on the placement of the focal zone and the inherent, elevation focus incorporated into the design of the transducer. As we saw in an earlier example, the ultrasound field is quite variable in width affected by the focal zone, and also as shown in the image on the right, the ultrasound field produced by a transducer may be extremely complex with multiple areas of, acoustical energy extending from the transducer in addition to the primary beam used for imaging.

Here is a simplified example of the energy pattern produced by a transducer in which the central beam of the ultrasound indicated by A, represents the assumed path of the ultrasound energy by the machine where b and c represent additional areas of relatively high acoustical energy called side lobes. These vary from one transducer to another and are important because if the side lobe interacts with a strong interface, the echo from that interface returns to the deucer And is displayed in the image. Since the ultrasound machine assumes that that echo is coming along the path of the primary beam indicated by a, the echo is displayed in the image at an incorrect location. An example of this type of artifact is seen in this view of the gallbladder. A bright echo is present in bowel adjacent to the gallbladder indicated by arrow B. When this interacts with a side lobe, it produces an artifactual echo labeled a, which extends into the lumen of the gallbladder. Under certain circumstances, such an artifact might be misinterpreted, as indicating a complex structure when in fact the structure is completely normal. Fortunately, this kind of artifact can easily be recognized by slightly changing the position or orientation of the transducer, in which case the artifact should disappear.

Slice Thickness Artifact

Another artifact relates to the fact that the ultrasound machine, and assuming that the ultrasound beam is uniform and very thin, is reflected in this series of images. Here we can see the scan plane of an ultrasound field passing through a cystic structure such as the gallbladder, and when the plane passes directly through the center of the cyst, the wall is detected in a accurate fashion. If, however, the scan plane is near the edge of the cyst or gallbladder averaging due to slice thickness artifact results in a thicker appearance of the wall than is actually the case and may lead to overestimation of wall thickness. Finally, if the focal zone is not properly positioned, the divergence of the ultrasound being and both the elevation and lateral planes may result in further volume averaging creating the appearance of low level echoes in certain portions of the cyst that you've shown here. And in fact, the area C shown in this view of the gallbladder represents artifactual echoes produced by volume averaging in the dependent part of the gallbladder. These should not then be mistaken for sludge or small stones.

Specular Reflectors and Angle Dependence

Another assumption that the ultrasound machine makes is that the beam travels directly to and from the echo producing interface. This involves a little consideration of what kinds of interfaces, scatter ultrasound. In general, there are two types of reflectors. One are called diffuse scatterers. These are very tiny interfaces within tissues and organs that are small relative to the wavelength of the ultrasound being used. The other reflectors are very large flat surfaces, so-called specular or mirror-like reflectors, and most solid organs such as the kidney, liver, or in this case the breast. Much of the tissue echo pattern is produced by diffuse scatterers. In this case, the ultrasound interacts with very small tissue interfaces and it's scattered in all directions. Some of the scattered energy returns to the transducer where it's detected. If you look very closely at the echo pattern within a solid organ, you'll see that it's made up a number of small dots of varying size and brightness. This has a speckled appearance and in fact is called ultrasound speckle.

Speckle is a type of artifact produced by the diffuse scattering of ultrasound energy. Each small scatterer sends ultrasound energy in a variety of directions, and as multiple scatterers are produced, the acoustical fields, interact in some cases in a constructive fashion and in some cases in a destructive fashion. These areas of constructive and destructive interference produce the heterogeneous appearance, known as ultrasound speckle, and forms the basic underlying texture that is associated with various organs. A speckle, however, tends to diminish contrast, and when a large amount of speckle is present, it may make the identification of certain objects difficult to identify because of poor contrast. In this case, if we can reduce speckle, we can see a dramatic improvement in the ability to detect low contrast objects, and therefore a number of techniques have been developed in an attempt to reduce the speckle and improve contrast by various processes. Here, for example, is a breast cancer shown in conventional imaging, and on the right after a speckle reduction technique called spatial compounding has been used. Notice that both the improved contrast and the better delineation of these small calcifications within the cancer.

When ultrasound encounters a large flat surface, which is relatively perpendicular to the ultrasound beam, some of the ultrasound energy passes through and some is reflected returning to the transducer and displayed in the ultrasound image. If, however, the flat surface lies at an angle to the ultrasound beam, much of the incoming energy may be deflected and not returned to the transducer and therefore not appear in the image. Here in the liver, we can see an example of some of these effects. In this case, the hepatic vein that lies nearly perpendicular to the ultrasound beam is displayed, quite clearly. But as we look at a vein which lies at a greater angle, the wall is less clearly defined, and the diaphragm, which in the middle of the image appears quite clear, disappears almost completely from view. In the left hand side of the image, this is a manifestation of the angle dependence of large flat reflectors such as the diaphragm. The portion of the diaphragm, which lies nearly perpendicular to the transducer beam, reflects most of the sound directly back to the transducer and is displayed very clearly, whereas the portion of the diaphragm that lies at an angle reflects sound off away from the transducer so that the echo from the diaphragm is hardly displayed at all. Here's another example of large flat reflectors, which when seen perpendicular to the sound beam show quite clearly as in the upper image of a breast implant, which is partially collapsed, but notice in the bottom image that the implant actually extends further to the left than is shown in the top image. The reason that the top image does not show the entire implant is because the angle changes, and in that view does not return to the ultrasound transducer.

Refraction and Shadowing

Another effect, related to the direction of the ultrasound beam and the interaction with reflectors is illustrated here where shadowing is seen along the edge of the cyst. This is due in part to refraction of the ultrasound beam. When ultrasound passes from one tissue to another in which the propagation velocities differ slightly, the direction of the ultrasound beam rather than being straight may deviate slightly depending on whether the velocity is increasing or decreasing. As the sound passes from one tissue to another, this deviation of, the direction of the ultrasound as it passes through the interface is called a refraction. In the case of, the cystic lesion, the refraction is responsible in part for the shadowing scene along the edges of the cyst. As the ultrasound strikes the interface between the, cyst fluid and the tissue surrounding it, there's a change in velocity which causes a deviation of the direction of the ultrasound beam, as well as some deviation due to the nature of the, specular reflector, which formed part of the cyst wall. Since this sound bounces off at an angle, it does not return to the transducer and therefore appears as a signal void or shadow along the edge of the cyst.

Reverberation Artifact

A common artifact related to the propagation of sound between strong specular reflectors is reverberation sound, striking one large specular reflector will be partially returned to the transducer. Some will penetrate further striking the second strong reflector producing a second echo. What may then happen if the interfaces are strong and parallel is that the sound may move back and forth between these two reflectors before it returns to the transducer. This requires additional time, and since the ultrasound device assumes increased time to indicate increased distance, the echo is displayed at an incorrect location. This pattern of repeating echoes, progressively deeper within the object is the typical appearance of reverberation and is illustrated here in which the double walls of a breast implant being quite strong. Specular reflectors produce a series of reverberation artifacts,

Refraction Artifact in Imaging

another somewhat, more complex artifact as shown here in which the path of the ultrasound beam is not what the ultrasound, machine assumes it to be. Here we see two cystic structures within the uterus in this transabdominal scan through a full urinary bladder of the uterus. The first of these is a true structure. The second is an artifact. The reason for this artifact is, illustrated in this example object. A is produced by the sound path directly from the transducer to the object and back again, object B is due to the interaction of a refraction occurring at the junction or the edge of the rectus muscle. The sound is deviated, interacts with the cystic structure, returns to the transducer along the same path, and because this is a longer path, the object is displayed in the image at a slightly deeper depth and also along the expected path of the ultrasound beam, which is to point B rather than to point a. Again, this type of artifact, although uncommon can be quickly corrected simply by changing the angulation or slightly altering the position of the transducer so that the effects of, refraction seen here are no longer present.

Spatial Compounding

Two dimensional ultrasound images are produced by a series of small scan lines that are generated very rapidly across the entire image field. This entire sequence of events occurs at a rate of 20 to 60 times per second, and by creating these multiple scan lines very rapidly, many, many times a second, the impression of realtime motion is achieved. One of the consequences of producing images in this fashion is that the scan lines are all similar in position and in the case of a linear array illustrated here, they're all parallel. This means that the angle of intonation of each scan line to the target is fixed. Therefore, if we examine a, a spherical object such as shown here with such a transducer, the periphery of the, structure which is relatively perpendicular to the sound will be displayed well, whereas the edges, will not be seen as clearly. In addition, special ar speckle artifact may diminish the contrast of small subtle features within the mass.

One solution to this is to scan at multiple angles. This will draw in the margin more completely, and also because speckle is a random event. By scanning at multiple angles, the true signals are enhanced while the speckle is not, and therefore speckle overall is reduced improving contrast. This process of using real time to scan at multiple scan angles is called spatial compound imaging, and has been an important, adjunct in certain types of applications such as breast and small parts imaging. Here is an example comparing a conventional image and a spatially compounded image of a small lesion within the breast. Notice the improved contrast, of the image on the right, better definition of some of the specular reflectors in the subcutaneous fat and reduction of some of the, spurious echoes within the cyst. Also, notice that the edge refraction artifact on the left is considerably reduced in the compounded image On the right, Here again, is a comparison of conventional and compounded images of a small breast cancer in which the contrast is clearly improved in the compounded image, as well as the visualization of the, subtle microcalcifications within the duct.

One of the disadvantages of spatial compounding is that certain artifacts that may be diagnostically useful such as enhancement and shadowing are reduced by the compounding process. Here we can see the cyst on the left in conventional mode, shows very clearly defined acoustical enhancement, whereas this is considerably diminished in the image on the right with compound imaging.

Ultrasound Attenuation

The next topic in this discussion of basic ultrasound principles focuses largely on the issue of ultrasound attenuation. With ultrasound imaging, the ultrasound device makes a number of assumptions, which may or may not be correct. One of these assumptions is that the attenuation of the ultrasound beam is uniform, and in fact, nothing is further from the truth. When ultrasound passes into tissues, it encounters interfaces where tissues differ slightly in physical properties. The degree of difference between these properties, which largely are the propagation velocity and physical densities of these tissues, determines the proportion of sound that will, will be reflected at that interface As the sound therefore is, reflected and absorbed by the tissues ultimately being converted to heat. The intensity of the ultrasound decreases as the depth increases.

here, for example, at five megahertz on the left, if we pass the ultrasound into a specified depth, we'll discover that at a certain point, only 10% of the initial, intensity is present, and if we pass the sound an additional similar depth and another 90% is removed, resulting in a residual intensity of only 1% of what was initially applied at 10 megahertz. The same phenomenon occurs, but much more rapidly here at the same depth that removes only 90% of the energy from a five megahertz ultrasound field. 99% has been removed at 10 megahertz as ultrasound energy is, removed by reflection and attenuation of tissues.

A shorthand is used to, describe the relative intensity at various locations within the ultrasound field. Here we want to compare the intensity at point A and point B, and in this example, at point A, we have, 100% of our power and at point B 1%. This, can be expressed as a ratio of one over 100, or we can take the log of this ratio, which is a ratio of a hundred to one. The log of that would be two, and if we multiply that by 10 to achieve a decibel Notation, we can describe this difference in the power at point A and B as differing by 20 db. This becomes then a very useful, shorthand for describing, large differences in power at different locations. 20 DB representing a, a hundred fold difference in power.

Now, most tissues, attenuate sound at very roughly one decibel per centimeter per megahertz, and as you can see from this relationship that higher frequency transducers attenuate sound much more rapidly than low frequency transducers.

Now, we mentioned, that the ultrasound machine assumes that attenuation is uniform. We've just, indicated that in fact, that's not the case, that as we go deeper and deeper into tissue absorption reduces the intensity of the ultrasound, and this is reflected in an uncorrected images as a loss of, intensity. As we proceed to greater and greater depth within the tissue, We have to compensate for this attenuation, and we do this in a very simple fashion. By time, gain compensation or depth gain compensation here, for example, at a depth of one centimeter, we are down 10 db or down to 10% of our original intensity, and by adding 10 DB of amplification to the signal at that depth, we can restore the image to its normal value. At a depth of two centimeters, you can see that we have very little signal, but if we add 20 DB of amplification or gain at that depth, we can again restore the image to normal value. The way this is done with ultrasound machines is somewhat automatic, but user, tuning of this is also important, and this is done using the slider controls on the ultrasound device to increase the amount of amplification as we go deeper and deeper into the tissue.

in this example, we've corrected for the attenuation in the mid portion of the image by adding modest amounts of gain, for that depth, and as we go in deeper into the tissue, we add more progressively, more amplification so that we achieve a uniform level of, intensity throughout the image.

Shadowing and Enhancement

Now, shadowing and enhancement are, excellent examples of cases where the assumption that attenuation is uniform are not met, and both shadowing and enhancement are artifacts which are often clinically useful. In the case of the cyst at the surface, we have no attenuation and we've added no gain to correct for attenuation at a depth of one centimeter. Natural attenuation at this frequency is about 10 db, and we've corrected that by adding 10 DB of gain, but the cyst is attenuating only by about three DB rather than 10 db. So when we correct the attenuation deep to the cyst, we actually end up plus 70 B with much greater intensity than we started with, and as we continue deeper into the tissue, the TGC correction still results in a greater value deep to the cyst than it does elsewhere in the image because the attenuation is not uniform. The correction we apply for attenuation results in the artifact of acoustical enhancement deep to the cyst. The same is true for shadowing here at a depth of, one centimeter, attenuation has removed 10 db. We've restored that with gain, but the mass has attenuated 20 db, we've only corrected 10 db, so we still end up, with less intensity than normal, and this is apparent as an area of shadowing deep to the mass.

Bandwidth and Speed of Sound Assumptions

Another consideration, is that of bandwidth. Early transducers produce frequencies over a very narrow range indicated by the center portion of this graph. Modern transducers produce, intensities over a much larger range of frequencies in this example from about six to about 12 megahertz. This has several implications with broadband, with transducers, the high frequency components can assist in imaging the superficial structures, but because the high frequencies do not penetrate the deeper levels, we require the mid frequencies and the lower frequencies of the broadband width to restore the image at to proper levels and these portions.

Another assumption that the machine makes is that the speed of sound and tissues is constant. The machine is assuming that the sound is propagated at 1,540 meters per second, but as we've seen, it travels a little faster in blood, somewhat slower in water, and even more slowly in fat and air because the machines simply uses, the measurement of time and then calculates based on the assumed propagation velocity, the distance that the object lies from the transducer. If the propagation velocity is not what the machine has assumed, the location of the structure in the image will not be correct.

Here's an example in which a fatty structure is illustrated deep within the liver. The sound travels more slowly through this area at about 1,415 meters per second compared to 1,540 meters per second in tissue. Because the sound travels more slowly, it takes longer for the echo to return to the transducer. The machine assumes that the sound is traveled at 1,540 meters per second, so the delay in return of the echo is interpreted as indicating that the structures deep to the fatty mass lie deeper in the body than is actually the case, and this results in the displacement of a small segment of the diaphragm from its expected location.

A more subtle example of this occurs when, ultrasound is used to evaluate large cystic structures. The propagation velocity of sound in cyst fluid is a little slower than in normal tissues. That means that the sound passing from the near wall to the far wall of the cyst will take longer because it's, being propagated more slowly. The machine will then assume that, this indicates that the cyst is slightly larger than is truly the case, and in this example, the true diameter is 5.3 centimeters. While the measured diameter in the image is 5.5 centimeters, this overestimation of the size of the cyst may result in an error in the estimation of the volume of the cyst, and in this case, the error is about 10%.

A more dramatic example of this, error in the assumption of the machine regarding the propagation velocity is illustrated in this, scrotal ultrasound. One testicle has been removed and replaced by silicon prosthesis, and in fact, both the prosthesis and the normal testicle are identical inside. However, the slow propagation velocity of sound in the silicone results in its appearance in the image being much larger than is actually the case.

Harmonic Imaging

Ultrasound has been enhanced in, recent years by a number of advanced techniques which have made major contributions to the quality of, image production. One of these is harmonic imaging. This is largely the result of the, development of broadband with transducers that are capable of producing frequencies over a fairly large range. This allows us then to exploit a, a unique property of, sound as it passes through tissue. This relates to the distortion of the pressure wave form. As the sound passes through tissue, the high pressure components of the pressure wave tend to move a little more quickly than the low pressure components, so that as the sound passes deeper and deeper into the tissue, the pressure wave becomes distorted, and this distortion gives rise to the development of multiples of the fundamental or primary frequency. These multiples are called harmonics and are typically at twice or even three times the, fundamental frequency because these, harmonics take some, distance of travel within tissue before they develop. They do not appear near the transducer, but only appear some distance within the tissue. In addition, these harmonics develop only in the highest, energy portions of the ultrasound beam. Therefore, the profile of the harmonic energy is much smaller than that of the fundamental energy as illustrated in this slide.

Well, the broad bandwidth system, it becomes possible to transmit at a primary frequency and to then tune the receiver to detect only the harmonic or the multiple of that frequency with conventional imaging. Now illustrated in this slide, we transmit, at the primary frequency. The acoustic field contains harmonics, but we process only the primary frequency, using the information illustrated in the bottom slide. With harmonic processing, we transmit the primary frequency, but produce the image only using the harmonic information illustrated in red at the bottom. As you can see, much of the information near the transducer, which consists of a great deal of clutter, is removed from the image, and the profile of the beam generated by the harmonics is considerably tighter than that, from the primary. This can have a dramatic effect on the image, contrast and image quality as illustrated in this view of the gallbladder with conventional processing on the left and harmonic processing on the right, some of the clutter, and noise from the gallbladder lumen is suppressed by this technique, and overall contrast and image quality is significantly enhanced here.

Another example of harmonic imaging, comparing a cystic, mass within the breast is shown with conventional imaging on the left and harmonic on the right. Particularly impressive is the reduction of clutter and noise within the, cystic components.

Summary

In summary, then ultrasound is a unique form of imaging, which, draws upon physical principles and interactions with tissues that are very different from those of other forms of, diagnostic imaging. And because this imaging is derived in a totally different manner from other methods, it is quite unique and often very much complimentary to other methods. The purpose of, this summary of, basic principles then is to, emphasize the importance of not only understanding these principles, but understanding how knowledge of these principles can be utilized in optimizing the quality of diagnostic scans and obtaining the maximal diagnostic benefit from the, ultrasound examination.