Introduction to Diagnostic Ultrasound Technology PDF

Introduction to Diagnostic Ultrasound Technology Course purpose: The aim of this course is to give you a good start in understanding the physics of diagnostic ultrasound. It introduces the main concepts and explains why they are important clinically. Course structure: The course parallels my textboo...

Introduction to Diagnostic Ultrasound Technology Course purpose: The aim of this course is to give you a good start in understanding the physics of diagnostic ultrasound. It introduces the main concepts and explains why they are important clinically. Course structure: The course parallels my textbook The Physics and Technology of Diagnostic Ultrasound: A Practitioner's Guide (Second edition).* It's important to realise that the textbook provides significantly more detail. So, I strongly recommend that you regard this online course as an introduction and use the textbook to fill in the details you need to understand to be an ultrasound expert. This module: This module discusses the nature of ultrasound waves, some of their properties and the interaction of ultrasound with human tissues. It parallels Chapter 2 of the textbook. Chapter 1: Ultrasound Interaction with Tissue Lesson 1: Ultrasound Waves and Propagation Ultrasound is simply a very high-frequency sound. Diagnostic ultrasound frequencies range from 2 MHz to 20 MHz*. This is around 1000 times higher than the frequency range of audible sound (20 Hz to 20 kHz). As we will see soon, these high frequencies are used to achieve the best possible image resolution**. *Hz = cycles per second; kHz = 1000 Hz; MHz = 1,000,000 Hz **ability to image small structures The ultrasound wave consists of oscillating pressures. It is generated when the ultrasound transducer vibrates against the skin surface (at the ultrasound frequency). When the transducer is moving towards the body it compresses the tissues, causing an increase in pressure (compression). When it moves away from the body it decompresses the tissues, causing a decrease in pressure (rarefaction). In this way, oscillating changes of pressure are created. These constitute an ultrasound wave that travels through the tissues at a constant speed. If we could focus on a single point in the tissues, we would see the pressure oscillating as a function of time, increasing and decreasing equally above and below the mean pressure in the tissue. The amplitude (A) is the maximum change of pressure from its mean value in the tissues. It determines the amount of energy in the ultrasound wave and therefore the level of exposure of the tissues. The period (T) is the time occupied by one cycle. It is related to the ultrasound frequency (f) (the number of cycles per second) as follows: T = 1/f We can also imagine a situation where we could look at the variation of pressure as a function of depth into the tissue at one instant in time. (This is like taking a photograph of ocean waves at one instant.) We would find a similar pattern of pressure variations with depth. The wavelength (λ) is the physical length of a single cycle. (This is like the distance from the top of one ocean wave to the top of the next one.) It is very closely related to image resolution; the smaller the wavelength the better the resolution will be. The speed at which the ultrasound wave travels is referred to as the propagation speed. It is conventionally represented by the symbol c. The average value of c in normal soft tissue is 1540 m/sec. The equation below describes the relationship between propagation speed, frequency and wavelength. c=f⨉λ This can be rearranged to allow the wavelength to be calculated for a given frequency. λ = c/f This table lists a number of typical ultrasound frequencies and wavelengths in soft tissue (assuming c = 1540 m/sec). Notice that for all frequencies the wavelength is less than one millimetre. Also notice that the wavelength becomes smaller as the frequency increases. Since the resolution of an ultrasound image is directly related to the wavelength it is clear that: The higher the frequency the better the resolution. Unfortunately, we will see shortly that it is not possible to keep increasing the frequency to achieve ever better resolution. The physics of ultrasound places a limit on the frequency that can be used in a given clinical situation. Lesson 2: Attenuation Diagnostic ultrasound gives useful information precisely because it interacts strongly with soft tissue. In this and the following two sections, the major types of interaction will be discussed. These are: attenuation. reflection. scattering. refraction. Attenuation As an ultrasound wave travels through tissue, it becomes progressively weaker. This is referred to as attenuation. In soft tissue, the primary cause of attenuation is the absorption of some of the wave’s energy by the tissue. This happens because tissue is not perfectly elastic and so there is some friction as it moves back and forth in response to the pressure variations in the wave. This friction causes heating of the tissues and so the temperature increases. Since energy cannot be created or destroyed, this process removes energy from the ultrasound wave, making it weaker. Other factors can contribute to attenuation. As discussed in the next section, reflection and scattering from structures in the body cause some of the ultrasound energy to be deflected. This energy is lost from the wave, resulting in weakening (i.e. attenuation) of the ongoing ultrasound. If the ultrasound beam diverges (because of defocusing), the energy in the beam is spread over a greater area and so the intensity decreases. (Think of a torch being defocussed.) Referring to the figure on the left, the amount of attenuation is calculated as the ratio of the initial ultrasound intensity (I1) to the final intensity (I2). Since it is a ratio, it is usually measured in decibels (dB). So, the definition of attenuation is: attenuation = 10 ⨉ log (I1/I2) dB Using decibels makes it particularly easy to calculate tissue attenuation for a given situation: attenuation = (α ⨉ L ⨉ f) dB Here α is the attenuation coefficient of the specific tissue involved (in dB/cm/MHz), L is the total distance travelled by the ultrasound (in cm) and f is the ultrasound frequency (in MHz). For typical soft tissue, the attenuation coefficient α is approximately 0.5 dB/cm/MHz Example Consider a situation where the machine is operating at 3 MHz If the transmitted ultrasound travels to a depth of 20 cm in soft tissue, it will be attenuated by 30 dB, as shown below. attenuation of transmitted ultrasound = (0.5 ⨉ 20 ⨉ 3) = 30 dB This means the transmitted intensity is reduced by a factor of 30dB = 1,000*. Similarly, the echo returning from a reflector at this depth will be attenuated by 30 dB as it travels back to the transducer. Thus, the total round path (there and back) attenuation is 60 dB. This means that the echo coming from 20 cm depth will be 60 dB weaker (i.e. 1,000,000 times lower in intensity) than an echo from a similar reflector at the skin surface. This example highlights how substantial the attenuation of ultrasound is. *See Chapter 1 of my textbook for a discussion of decibels. A second example Suppose we now decide to scan using 6 MHz instead of 3 MHz in an attempt to improve image resolution. The round path attenuation will then be 120 dB, corresponding to a total reduction in intensity by a factor of 1,000,000,000,000! Echoes that have been attenuated this much are so small that they cannot be detected by the machine so they will not appear in the image*. *All electronic devices, including ultrasound machines, are affected by electronic noise due to the random motion of electrons. Ultrasound echoes whose amplitude is comparable to the machine’s electronic noise or weaker are not detected. Penetration (P) When the round path attenuation exceeds the maximum that the machine can tolerate, the echoes are too small to detect and they are not displayed. The depth (P) at which this happens (i.e. the depth beyond which the echoes are not detectable) is referred to as the depth of penetration (or simply the penetration). For a given machine and clinical setting the round path attenuation for echoes coming from the penetration depth is: attenuation = α ⨉ (2P) ⨉ f = 2α(P ⨉ f) This is the maximum attenuation the machine can tolerate without the echoes becoming undetectable. For a given machine this value will be constant. The attenuation coefficient of the tissue (α) is also constant for a given clinical situation. Therefore, the bracketed term on the right-hand side of this equation (P × f) must be constant. As the frequency increases the penetration must decrease to keep the product (P × f) constant. As the frequency increases the penetration decreases. Lesson 3: Reflection and Scattering Reflection and scattering are the two mechanisms that produce echoes. Reflection: the interaction of ultrasound with relatively large and smooth surfaces (think of light reflecting from glass.) Scattering: the interaction of ultrasound with small structures (red blood cells, capillaries, etc.) within the tissues (think of light scattering from the tiny water droplets in a fog.) Reflection, acoustic impedance When ultrasound strikes an interface between two tissues, some of the ultrasound energy is reflected and the rest continues into the deeper tissues. The more different the two tissues are the more of the energy is reflected. In this diagram, z1 and z2 are the acoustic impedance values of the first and second tissue respectively. The acoustic impedance of a tissue is z=ρ⨉c where ρ is the density of the tissue (its weight per unit volume) and c is the propagation speed. The units are Rayls. Reflection coefficient The reflection coefficient R is a measure of the fraction of the ultrasound energy that is reflected. It can be calculated as follows: R = (z1 - z2)2 / (z1 + z2)2 A reflection coefficient of 0.01, for example, would mean that 1% of the ultrasound energy is reflected and 99% passes through the interface. It is easy to show that when the two tissues have the same acoustic impedance (z 1 = z2) the reflection coefficient is zero and no energy is reflected. It can also be shown that when z1 and z2 are very different the value of R is close to one, which means that almost all the energy is reflected and very little passes through the interface. Reflection geometry So far, we have focussed on the special situation where the ultrasound is incident on the interface at right angles (i.e. at an angle of 90° to the interface); this is called perpendicular incidence. What about the more general situation? When the incident angle is not 90°, the reflected ultrasound does not travel back to the transducer. As a result, the echo from the interface is not detected and so it is not seen in the image. Echoes caused by reflection will only be detected by the machine and displayed when the ultrasound strikes the interface at perpendicular incidence. Scattering The word scattering describes the interaction of ultrasound with small structures such as red blood cells and capillaries. Scattering differs from reflection in two important ways. 1. Scattered energy is distributed in all directions. Reflected ultrasound goes in a single direction. 2. Scattered energy is generally much weaker than reflected energy. Echoes caused by scattering are generally displayed as low- to mid-level grey tones in the image. Most of the echo information in a typical ultrasound image comes from scattering from soft tissue, not reflection from interfaces between different tissues. So, an understanding of the character of scattered echoes and the way they look in the image is important. Speckle As the diagram to the left shows, the echo signal at any moment is the sum of many echoes, each caused by an individual scatterer within the volume of tissue occupied by the ultrasound pulse. These scatterers are randomly positioned relative to each other, so their echoes add together randomly. This causes random variation in the amplitude of the echo signal received by the transducer. The result is the random granular echo texture in the image that we call speckle. Example Notice how the speckle (the echo texture) in this image varies with depth. Close to the transducer the texture is quite fine-grained whereas at greater depths it is coarser. The phantom material is the same at all depths, highlighting the fact that speckle does not directly reflect a tissue property. It is important to realise, however, that the average brightness of the speckle is meaningful, and it will vary as tissue properties vary. Lesson 4: Refraction Refraction of light You are probably familiar with refraction of light – the bending of the light’s path as it passes through different materials. Examples include: a prism – glass with a triangular cross-section (see diagram); the bending of light as it passes from water to air. Light is refracted whenever it travels from one medium (e.g. air) into another medium that has a different propagation speed (e.g. glass or water). Refraction of ultrasound Similarly, ultrasound is refracted whenever it passes through an interface between tissues with different ultrasound propagation speeds (for example, from liver tissue to fat). The geometry of refraction is determined by measuring the direction of travel of the ultrasound relative to a line drawn at right angles to the interface, as shown in the diagram. The amount of refraction can then be calculated using Snell’s Law: (sin θi)/c1 = (sin θt)/c2 where θi is the incident angle and θt is the transmitted angle. This equation explains a number of features of refraction, including: When two tissues have the same propagation speed (c1 = c2), θi = θt which means there is no refraction, and the beam direction is unaltered. A special case occurs when the beam is at perpendicular incidence to the interface. In this case, θi is 0° so sin θi = 0. According to Snell's Law, this means that θt is also  0°. So there is no change of beam direction (refraction) when the beam is at perpendicular incidence, regardless of the propagation speeds. When the propagation speed in the second tissue is lower than in the first (c2 < c1), θt is smaller than θi as shown in the diagram above. When c2 is greater than c1, θt is larger than θi as shown below. View the four images below. They show that the amount of refraction increases as the incidence angle increases. They also reveal a problem that occurs with large incidence angles when the second tissue has a propagation speed higher than the first. 1. 2.  3. 4. Critical angle As shown above, ultrasound cannot penetrate into the second tissue if (a) the propagation speed is higher in the second tissue, and (b) the incidence angle is equal to or larger than the critical angle. The value of the critical angle (θc) for a given pair of tissues is found by solving Snell's Law for the situation where θt = 90°. This gives us θc = sin-1 (c1/c2) Summary When ultrasound passes through an interface between two tissues with different propagation speeds: the beam path is bent, except for the special case of perpendicular incidence; the amount of bending increases for large differences in propagation speed and large incident angles; in the situation where the propagation speed in the second tissue is higher than in the first, a critical angle exists. When the incident angle is larger than this critical angle, total reflection occurs and no energy is transmitted into the second tissue. Introduction to Diagnostic Ultrasound - Pulsed Ultrasound and Imaging This module deals with the ultrasound pulse and the basic principles of ultrasound imaging. It parallels Chapter 3 of the textbook. Lesson 1 – Transmit Pulse The term pulse refers to ultrasound energy that starts and then stops again shortly afterwards, that is a short burst of ultrasound energy. The diagram below shows a typical ultrasound pulse. The opposite of a pulse is continuous wave ultrasound, which continues indefinitely. Continuous wave Doppler is the only diagnostic ultrasound modality that uses continuous waves. As the figure shows, the pulse used for ultrasound imaging is typically three cycles long. The pulse duration is therefore approximately three times the period of the ultrasound wave. Remembering that the period (the duration of a single cycle) is equal to (1/f) where f is the frequency, we can see that the pulse duration is approximately (3/f). So the pulse duration becomes smaller as the frequency increases. For example, we can see that the pulse duration at 3 MHz is approximately 1 μs*, while at 10 MHz it is 0.3 μs. We will see later that the depth resolution (or axial resolution) of the ultrasound image is proportional to the pulse duration. * 1 μs = 1 microsecond, or 1 millionth of a second Increasing the ultrasound frequency shortens the pulse duration and improves the axial resolution of the image. Lesson 2 – Pulse Repetition Frequency (PRF) To build up an image, the machine repeats the process of transmitting a pulse and receiving echoes from that pulse at regular intervals (see diagram). For each transmit pulse, the ultrasound beam is moved to a new position, so it passes through a different area of tissue. (The movement of the beam is referred to as scanning; it will be discussed later.) So, the machine must generate a number of transmit pulses – usually 100 or more – to produce each image. The machine must create images at a sufficient rate that they are seen as "real-time" (movie-like). Generally, this means there should be at least 15 or 20 images each second. So, it needs to send out transmit pulses as rapidly as possible. For example, if each image requires 100 transmit pulses to produce and we require an imaging rate of 20 frames per second (20 images per second), it is easy to see that the machine must transmit a total of 2000 pulses each second. The number of pulses transmitted each second is called the Pulse Repetition Frequency (abbreviated PRF). For the machine’s electronics, transmitting 2000 or more pulses per second is no problem. However, we will see shortly that there is an important practical consideration that limits the PRF. Lesson 3 – Pulse-Echo Principle We now come to the fundamental concept that underlies diagnostic ultrasound – the pulse- echo principle. Simply put, by accurately measuring the interval between the time when each pulse is transmitted and the time when each echo returns, the ultrasound machine can calculate the distance between the probe and the structure that caused the echo. Consider the situation shown in this figure. The total round path distance travelled by the ultrasound (from the probe to the reflector, then from the reflector back to the probe) is (2 ⨉ d). The time taken to travel this distance, and therefore the time between the transmit pulse and the echo, is t = (distance travelled)/(speed) = (2 ⨉ d)/c Rearranging this equation, we can see how the machine calculates the depth of each reflector or scatterer from the echo arrival time d = (c ⨉ t)/2 The machine must know the value of the propagation speed (c) to use this relationship. Since it has no way to determine the actual value for the tissues under examination, it assumes the average value in soft tissue (that is 1540 m/sec). As an example, consider an echo that arrives 130 μsec after the transmit pulse. Using the above equation the machine calculates the depth as 10 cm. (It may be useful to remember that for every centimetre of depth, the echo delay is 13 μsec.) Lesson 4 – PRF Limitations We have just seen that the time between the transmitting of a pulse and receiving an echo is proportional to the depth of the structure that causes the echo. We have also seen that real- time scanning requires the ultrasound machine to create images as quickly as possible. We will now see how the time taken for ultrasound to travel in the tissues limits the rate of imaging. This occurs because of a fundamental rule. The machine must not transmit again until all detectable echoes from the previous transmit pulse have been received If the machine violates this rule, echoes from the new transmit pulse mix with echoes from the previous pulse and an artifact known as range ambiguity occurs. We will consider this in more detail later. Consider a scan where the penetration depth is P. (Reminder: this means that echoes from tissues at depths greater than P are not detectable). Using the relationship between depth and echo arrival time, we can calculate when the last detectable echo will be received. Its arrival time is tp = 2P/c This is the shortest time allowed between one transmit pulse and the next. So it follows that the maximum number of transmit pulses per second (the maximum PRF) is one second divided by this time PRFmax = c/2P The machine is designed to create as many images each second as possible. So it estimates the depth of penetration P (taking into account the ultrasound frequency being used) then calculates the maximum PRF consistent with that depth of penetration. Notice that the relationship above between the maximum allowable PRF and the depth of penetration is an inverse one. When the depth of penetration is small (e.g. in a musculoskeletal or carotid artery scan), a high PRF can be used and the frame rate is also high. Conversely, having a large depth of penetration (as generally required in an abdominal or obstetric scan, for example) forces the machine to use a relatively low PRF and so decreases the rate of imaging. Lesson 5 – Frame Rate Limitations Now that we know how to calculate the maximum PRF that the machine can use, it is relatively easy to calculate the maximum possible imaging rate. (Note that the number of images created each second is called the frame rate, abbreviated FR.) Suppose the machine requires N transmit pulses to create each image. It is easy to see that the total number of pulses required each second is simply (N × FR). Since the total number of transmit pulses available each second is (by definition) the PRF, we have N ⨉ FR = PRF So, the frame rate is given by FR = PRF/N Using the equation in the previous section relating the maximum allowable PRF to the depth of penetration (P), we get the following relationship between the maximum possible frame rate, the depth of penetration and the number of transmit pulses required to create each image FRmax = c/(2 ⨉ P ⨉ N) and so (FRmax ⨉ P ⨉ N) = c/2 So, the product of the three quantities on the left of this equation is limited by a physical constant that we cannot alter (the propagation speed divided by two). If we want to increase any one of these three quantities (e.g. the frame rate), one or both of the other two must be decreased to compensate. For example, consider an abdominal scan where we require a depth of penetration of 25 cm. Suppose that the ultrasound machine needs a total of 100 transmit pulses to create each image. Then the above equation tells us that the maximum possible frame rate is approximately 30 frames per second. This is adequate for many applications, but there are some situations (such as echocardiography) where we want a higher frame rate. Suppose we wanted a frame rate of 60 frames per second, but still required 25 cm depth of penetration. Then the equation above tells us that the number of pulses for each image would need to be decreased to 50. It is sometimes possible to achieve this by reducing the field of view (the width of the image) so that fewer pulses are needed to create each image. Lesson 6 – Principles of Image Formation The previous sections have described the basics of ultrasound imaging. The machine’s probe transmits a short pulse of ultrasound into the patient’s body. This pulse travels along a defined pathway called the ultrasound beam. Structures within the beam cause some of the transmitted energy to be reflected and scattered. Some of the reflected and scattered energy travels back along the beam to the probe where it is detected and converted into an electrical signal. The signal due to an individual structure is referred to as an echo. As we have seen, by measuring the exact arrival time of an echo the machine can calculate the distance from the probe to the structure that caused it. The machine also knows the position of the beam relative to the probe, and so it can define relatively accurately where the structure is within the ultrasound image. This allows it to place a dot in the image to represent the structure, as shown in this diagram. The brightness of the dot is determined by the intensity of the echo. A weak echo is shown as a dark grey, a moderate echo as a mid-grey and a strong echo as white. Black is shown when no echoes are detected. Three standard types of probes are used with ultrasound machines. Linear array probes. Curved array probes (sometimes called curvilinear or convex probes). Phased array (or sector) probes*. *The word array is used because the transducer (the active element in the probe) is sliced into a large number of identical elements. We will see that this allows the machine to use electronic focussing and scanning of the ultrasound beam. The diagrams below show the scan patterns and image shape for these three probe types. Linear Array. Curved Array. Phased Array. Why are different types of probes used in different clinical applications? The answer lies in their relative advantages and limitations. Considerations include the contact area of the probe (often called its footprint), the available acoustic window into the patient (i.e. the area in the superficial tissues through which ultrasound can travel) and the shape and elasticity of the skin where the probe is to be placed. In some areas, the acoustic window is extremely limited (for example, transthoracic echocardiography, where the ultrasound must travel between ribs and must avoid the lungs). In other areas, the acoustic window is large and therefore not such a concern (for example late pregnancy scans). In brief: Linear array: ideal for relatively flat surfaces such as the neck and limbs; however, the field of view is limited by the length of the probe. Curved array: widely useful; the field of view diverges with depth, so it is not so limited by probe size or the acoustic window; the degree of divergence depends on the degree of curvature of the probe and so the probe can be tailored to suit a given clinical application. Phased array: used in echocardiography; compact, easily manipulated and able to make use of a very small acoustic window; major limitation is the lack of detail close to the probe. Endocavity and laparoscopic probes Endocavity probes include transvaginal (see figure), transrectal, transoesophageal and endoscopic probes. As with standard probes, they use array technology to focus and scan the beam. These probes are used for two reasons. First, some areas are virtually impossible to reach with ultrasound from the skin surface because the ultrasound is obstructed by gas and bone. Endoscopic and laparoscopic probes can get much closer and so they eliminate problems relating to overlying tissues. Second, they use higher ultrasound frequencies than transcutaneous probes, so image resolution is improved. The frame rate is also increased because of the smaller depth of penetration required. 3D Probes As will be discussed in a later module, 3D imaging places an extra requirement on the probe. The scan plane must be swept through the patient’s tissues to acquire echo information from a three-dimensional volume. Historically this has been achieved by moving the probe, either manually or mechanically. More recently, however, it has become possible to achieve the same thing electronically using matrix transducers, which will be discussed briefly in the Module 3. Transducers and Focussing This module describes how the ultrasound transducer generates the transmit pulse and receives echoes. It also covers how the machine focuses and scans the ultrasound beam. It parallels Chapter 4 in the textbook. Lesson 1 – Transducer Principles The active component in an ultrasound probe is the transducer*. It generates the transmit pulse and receives the returning echoes. Transducers are made from piezoelectric** materials. A few naturally occurring piezoelectric materials exist (for example quartz) but they have poor efficiency. So specialised materials have been developed for ultrasound transducers. These may be piezoelectric ceramics or "composite" transducer materials which can operate over a wider range of frequencies. *Transducer: any device that can convert energy from one form to another. For example, a microphone is a transducer since it converts sound to electrical signals. **Piezoelectric: a material that generates an electrical voltage when it is compressed or expanded. Piezoelectric materials also show the inverse piezoelectric effect; they expand and contract when an electrical voltage is applied to them The transducer is grey in this diagram. Immediately behind it is a backing layer. This absorbs ultrasound energy that would otherwise travel from the transducer into the probe, causing unwanted echoes. The acoustic impedance of the ultrasound transducer is substantially different from the impedance of human tissue. This ‘mismatch’ would cause some of the transmitted energy to be reflected back into the transducer and some of the returning echo energy to be reflected back into the tissues. To address this, a thin matching layer is placed between the transducer and the probe face. This makes it seem as if the transducer and tissues have identical acoustic impedances. Coupling gel is placed on the skin to exclude air from the space between the probe and skin. Lesson 2 - Focussing In the previous module, we saw that a dot is placed in the image for each echo. The position of the dot in the image is determined by: the echo arrival time (used by the machine to calculate the depth of the point); the position of the beam. Unfortunately, the beam often has a substantial width. Since the machine doesn't know whether the structure causing each echo is on the midline of the beam or to one side or the other of the beam, it has to make an assumption. It displays every echo as if it had come from the midline of the beam, as shown in the diagram below. This significantly degrades the image, as we will see in a later module. We want to get as close as possible to the ideal of having a very narrow beam. This is why focusing the beam is such an important aspect of diagnostic ultrasound. This diagram shows a typical focused ultrasound beam. The diffraction limit (defined by the broken black lines) determines how well the machine can focus the ultrasound beam. The beam can never be narrower than the diffraction limit. In the focal zone, the width of the beam is equal to the diffraction limit (that is, the blue lines touch the black lines). At all other depths the beam is wider than the diffraction limit. This diagram shows a magnified view of the diffraction limit. The divergence angle θ can be calculated from the transducer aperture (A) and the ultrasound wavelength (λ) sin θ = (1.22 ⨉ λ)/A θ = sin-1(1.22 ⨉ λ)/A Note the aperture is the effective size of the transducer. We will see later that ultrasound machines use array transducers which are divided into a large number of small transducer elements. For linear and curved array probes, the machines use a sub-group of these elements to form each beam, not the entire length of the transducer, so the aperture is smaller than the overall length of the transducer. Clearly, focussing is good (the beam is narrow) when the diffraction limit is narrow. Remembering the inverse relationship between frequency and wavelength, the equation above tells us how to make the beam as narrow as possible. The best focusing is achieved when the aperture is large, and the frequency is high. Since the beam width at the focal depth is equal to the diffraction limit, we can calculate it using the equation above. beam width at focus = (2.44 ⨉ λ ⨉ F)/A where F is the depth of focus. Example For a 4 MHz probe with an aperture of 2.5 cm and a focal depth of 10 cm, the above equation tells us that the beam width at focus is 3.8 mm. This means that at all other depths, the beam width will be greater than this. Also, note that the beam width at focus is directly proportional to the focal depth. So, doubling the focal depth to 20 cm would change the beam width at focus to 7.5 mm; halving the focal depth would reduce the beam width at focus to 1.9 mm. Lesson 3 – Electronic Focussing and Scanning The intensity at any point in the ultrasound beam is defined as the amount of transmitted energy passing through that point per second per square centimetre. Intensity is a useful concept because it determines the strength of any echoes coming from the tissues at that point, and the likelihood of any harmful bioeffects due to ultrasound exposure. Consider the ultrasound beam. Imagine a certain amount of ultrasound energy travelling into the tissues within this beam. (Ignore tissue attenuation for simplicity.) Where the beam is wide, the energy is spread out over a large area and so the intensity is relatively low. At the focal point, the area of the beam is small and so the energy is highly concentrated, and the intensity is high. This gives us a clue about how focussing works. If we want to make the beam focus at a certain depth, we must ensure that the intensity in the middle of the beam is as high as possible at that depth. Now we'll look at how ultrasound machines achieve this. Array transducers As mentioned earlier, ultrasound machines use array transducers. This term refers to the large number of fine parallel cuts made in the transducer material (see diagrams below). These cuts transform the transducer into an array of identical transducer elements, generally just a fraction of a millimetre wide. Each element can act as a separate transducer, transmitting and receiving ultrasound independently of its neighbours. This enables the machine to focus and steer the ultrasound beam electronically and scan it through the tissues to acquire a two-dimensional image. Linear and curved arrays use a sub-group of the elements at any one time to transmit and receive ultrasound, not the entire length of the transducer. Phased arrays are different; they use all the elements. To create a beam that focuses at the point P, the outermost elements of the aperture transmit first, followed by the elements next to them and so on. The central elements transmit last. For simplicity, only pulses transmitted by the outermost and central elements are shown. The timing of the transmit pulses is arranged so that they lie on a circle (broken line) centred at P. (Right) As a result, the pulses arrive simultaneously at the point P. Transmit focussing It is now relatively easy to understand how an array transducer creates a focussed transmit beam. The machine transmits using all the active transducer elements (that is, all the elements in the aperture). The transmit time varies from element to element to ensure all the transmit pulses arrive at the focal point simultaneously. This guarantees that the intensity is maximum at that point and therefore the beam is focussed at that point. The electrical pulse from the transmitter is delivered to each transducer element via an electronic delay. The elements with the shortest delays transmit first. Those with longer delays transmit later. Just five delays are shown in this diagram for simplicity. In reality, there would be the same number as the number of elements in the transducer aperture (for example, 128). Transmit beam-former This diagram shows how the ultrasound machine's electronics focus the beam. In this example, the beam is focussed, and it is being steered to the left. The machine has calculated the required delays and set up the electronics accordingly. When the user alters the depth of focus or the beam direction changes, the machine quickly recalculates and adjusts the delays. The user sets the depth at which the image will be focussed and this is displayed on the screen. The machine then calculates the necessary delays. It is also possible to set multiple focal depths for improved uniformity of focus as shown below. However, this reduces the frame rate substantially. Echoes returning from the point P arrive first at the centre element of the aperture and last at the outermost elements. Receive focussing How do array transducers focus when they are receiving echoes? The principle is very similar. Again, electronic delays are used, in this case to delay the echo received by each element by an appropriate amount. Consider an echo returning to a phased array transducer from a point P in the tissues. As the diagram shows, the echo arrives at a slightly different time at each of the elements. Receive focussing. The echo arrives at the centre element first, so it requires the longest delay. Conversely, the echo arrives at the outermost elements last and so they require the shortest delay. This diagram shows how the machine uses electronic delays to compensate for these different arrival times so that the echoes are simultaneous and in phase when they are added together electronically. This ensures that the echo strength is maximised and so the receive beam width is minimised. Dynamic receives focussing There is one important difference between transmit focussing and receive focussing. Whereas the transmitted beam has a single depth of focus (set by the user and displayed on the screen), receive focussing is done automatically without user intervention and it is dynamic. This refers to the fact that at each instant the machine knows what depth the echoes are coming from (for example, 13 μsec after the transmit pulse the echo is coming from a depth of 1 cm). It therefore dynamically adjusts the receive focussing delays so that at each instant the beam is correctly focussed at the depth from which the echoes are coming. This is known as dynamic focussing. Scanning the beam A linear array typically contains 256 (or more) elements. As discussed above, each beam is created using a subset of these elements (generally a group of 128). To create a complete image, the machine steps the aperture (and hence the beam) along the face of the transducer (see diagrams below). The beams are at 90° to the transducer face and hence the image is rectangular. A curved array is similar in operation to a linear array, with the beams again being at 90° to the transducer face. The curvature of the transducer causes the beams to change direction as they step along the transducer face, and hence there is a radial component to the scan pattern. A phased array probe scans the beam by steering it sequentially in a number of different directions. A typical phased array sweeps the beam through approximately 90°, with each step being approximately 1°. Matrix arrays As mentioned earlier, standard linear, curved and phased array transducers are one- dimensional (1D), since they are divided into a large number of elements along their length. In recent years manufacturers have introduced a limited number of two-dimensional (matrix) arrays (see diagram). Matrix arrays have two major advantages. First, they can focus the beam in the direction across the width of the transducer (the elevation plane or slice thickness direction) using electronic beam-forming. This improves the slice thickness compared with standard probes and so improves image quality. Second, matrix arrays can steer the entire scan plane in a number of different directions, thus sweeping it through a volume of tissue. They can therefore be used to acquire 3D images, eliminating the need for manual or mechanical movement of the probe. Also, since electronic scanning and focussing can be extremely fast, 2D arrays have made it possible for the first time to produce 3D images in real time. As technology and manufacturing techniques improve, the cost of these probes is likely to fall and they may be more widely used than at present. Lesson 4 – The Ultrasound Beam and Image Quality Sidelobes The diagrams below show how the ultrasound intensity varies across the beam. In particular, they show that the intensity is maximum at the centre of the beam and weakens progressively towards the edges of the beam. A low-level sidelobe (an unwanted additional beam) is seen either side of the main beam. The vertical scale of the second diagram is logarithmic. This emphasises the weaker portions of the beam pattern so that the sidelobes are more obvious. You will notice that the sidelobes are weak compared with the main beam. The first sidelobe is approximately 20 dB (i.e. 100 times) weaker than the main beam, and the other sidelobes are even weaker. Even so, in some circumstances they can degrade the image by causing sidelobe artifacts (this will be discussed in a later module in more detail). Slice thickness So far, we have concentrated on how the machine focusses and scans the ultrasound beam within the scan plane (that is, a plane running along the length of the transducer, see below). We must recognise, however, that the beam is three-dimensional, so we also need to consider the beam profile in the elevation plane (that is, a plane running across the width of the transducer, as shown in the diagram). For standard linear and curved array probes (that is, non-matrix probes) the aperture (the active area on the transducer used to transmit and receive at any given time) is significantly smaller in the elevation plane than in the scan plane. This means that the beam cannot be focussed as well in the elevation plane as it is in the scan plane. Consequently, slice thickness artifact can be a significant problem in ultrasound imaging. This will be discussed in more detail in a later module. Example For example, consider a 5 MHz transducer with an aperture A of 20 mm. The beam width at focus (in the scan plane) can be calculated as usual beam width = (2.44 ⨉ λ ⨉ F)/A If the focal depth F is set at 5 cm, for example, we have λ = 0.31 mm (since f = 5 MHz), F = 50 mm, A =20 mm. The beam width is 1.8 mm. The beam width in the elevation plane can be calculated using the same equation. For example, if the transducer is 7 mm wide and the elevation focus is also set at 5 cm we have λ = 0.31 mm, F = 50 mm, A = 7 mm and so the elevation plane beam width at focus is 5.4 mm. So the beam width at focus is three times worse (larger) in the elevation plane. How does the machine focus the beam in the elevation plane? For standard (non-matrix) probes, it uses a lens. This means that the focal depth is fixed and cannot be varied by the user. For matrix probes the machine uses electronic focussing. In addition, matrix probes are wider than standard probes. The use of a wider transducer and electronic focussing means that matrix probes achieve significantly better focussing in the elevation plane than standard probes. Slice thickness artifacts are therefore significantly reduced. Ultrasound Instrumentation This module describes how the machine processes ultrasound echoes to produce an image. It also reviews the main controls available to the user and how they can be used to optimise image quality. It parallels Chapter 5 of the textbook. Lesson 1 – Introduction Ultrasound machines vary widely in how they achieve the functions they perform. Look at the variety of equipment available today! It ranges from large full-featured machines to pocket- sized "machines" based on smartphones (see below). Naturally, the electronics and processing varies substantially from one machine to another. In this module, we will be looking at the basic elements that are common to virtually all machines. An ultrasound machine contains some analogue electronics. This is needed to deal with time-varying electrical voltages, like the transmit pulse and the echo signal. However, most of the machine’s functions are carried out by digital electronics. The echo signal is converted into digital data (a series of 1s and 0s) which can then be processed digitally under software control. This has a number of advantages. Digital processing is stable and reliable. From the manufacturer’s point of view, digital processing is also very flexible. Software can be changed easily to add new functions or improve existing functions. One final comment. Much of the terminology relating to diagnostic ultrasound has become standardised, but you will still find variations between different authors and different parts of the world. Furthermore, equipment manufacturers often coin their own terms for a given function as a way of differentiating themselves in the market. In this book, we use generic terms as far as possible. This block diagram is probably intimidating at first glance. However, we will discuss the functions shown in it one by one, so you should be able to build up your understanding of the machine in relatively easy steps. For convenience, this module deals only with standard grey scale imaging. Doppler instrumentation and newer modes of operation, such as spatial compound imaging, harmonics and plane wave imaging, will be discussed in later modules. Lesson 2 – Probe, Transmitter, Beam-former Probe The construction of the probe was covered in the previous module, but a few additional comments are needed. A typical array transducer uses 128 transducer elements to transmit and receive at any one time. Generally, this means that there are at least 128 wires within the transducer cable connecting the probe to the machine. This explains why the cable is quite large. The weight of the cable can become an issue when you are scanning for extended periods of time, and some kind of support may be needed to reduce its impact. Controls With most machines, the user can choose which probe to use in each examination. As explained earlier, the choice is based on the depth to be scanned and the acoustic window. Transmitter and transmit beam-former The transmitter generates the electrical transmit pulse that is applied to the transducer. The machine’s software controls the frequency, amplitude, pulse duration and pulse repetition frequency of this pulse. As explained in the previous module, the transmit beam-former uses electronic delays to form an appropriately focussed (and, if necessary, steered) ultrasound beam. Controls Like all the machine’s settings, the transmit pulse parameters are set to default values by the machine at the start of the examination. You can change some of these settings if needed. The transmit pulse amplitude can be increased or decreased using the power control. Since increased power means increased patient exposure to ultrasound energy, the power should be increased only if other measures (e.g. increasing the gain and using a lower frequency) have been tried but they have not been sufficient to get the required information. In most machines it is also possible to change the ultrasound frequency (within the limits of the probe’s overall frequency range). Often this is done by selecting a ‘general’, ‘penetration’ or ‘resolution’ frequency setting. When the penetration option is chosen, the transmit frequency is reduced to improve penetration. Conversely, when the resolution option is selected, the frequency is increased so the resolution is improved. Although the machine sets the depth of focus to a suitable value for the ‘typical’ patient, you should always check the focal depth and adjust it if necessary. Receive beam-former This was also discussed in the previous module. Electronic delays are used to create the receive beam. The delays are adjusted continuously as the echoes return from increasing depth, so the beam is always focussed at the correct depth. Controls You have no control over receive focussing. Lesson 3 – Amplifier, TCG Amplifier Generally, the echo signal from the transducer is weak, its amplitude ranging from microvolts to millivolts. It therefore needs to be amplified to a higher level to enable subsequent processing. This also reduces the degrading effect of electronic noise, which is present in all electronic systems. Controls The amplifier’s gain is preset by the machine. However the user can adjust it to vary the overall brightness of the image using the machine’s gain control (sometimes called the B gain or overall gain control). (Top) Image without adequate TGC, showing loss of echoes with depth. (Bottom) With the TGC adjusted correctly, the brightness is similar throughout the display. TGC As discussed earlier, both the transmitted ultrasound pulse and the returning echoes are attenuated as they pass through tissue. For example, at a frequency of 6 MHz in normal soft tissue the intensity is reduced by 3 dB (i.e. it is halved) for each centimetre of travel. This means the received echo is reduced by a total of 6 dB (i.e. by a factor of four) for each centimetre of depth into the patient. If the machine did not compensate for this, echo information would be visible for superficial tissues but it would quickly fade with depth and deeper tissues would not be seen. The ultrasound machine therefore provides a built-in Time Gain Compensation (TGC) function (see diagram). This is occasionally called Depth Gain Compensation. As the name suggests, TGC applies a time-varying gain to the signals, beginning with a relatively low gain for superficial echoes, increasing steadily with time as the echoes come from greater depths. Controls In reality, tissues are variable in their attenuation. For example, liquids attenuate much less than soft tissue. So the machine provides a manual TGC function which operates in addition to the built-in TGC function. The purpose of this is to correct for these variations and allow you to achieve a uniform grey scale appearance at all depths. Generally, the TGC control is a series of sliders. Each slider corresponds to a specific depth in the image. Moving a slider to the right increases the gain at that depth, while moving it to the left decreases the gain. The TGC control often needs adjusting as patients are scanned. Lesson 4 – Dynamic Range Compression Dynamic range The term dynamic range is used to describe the overall range of echo intensities received by the ultrasound machine. It is defined as the ratio of the strongest to the weakest echo and is generally measured in decibels (abbreviated dB, see the Mathematics review in Chapter 1 of the textbook). The dynamic range (DR) is  DR = 10 ⨉ log(Imax/Imin) dB where Imax and Imin are the intensities of the strongest and weakest echoes respectively. As the diagram below shows, the dynamic range is large (60 dB or more) since the image contains both weak echoes (produced by soft tissue scattering) and much stronger echoes (caused by reflection of ultrasound from tissue interfaces). Echo intensity plotted on a logarithmic scale. Both the intensity values (in red) and the equivalent decibel values (in black) are measured relative to the weakest detectable echo at the left-hand end. We have seen that the machine displays each echo as a dot in the image. The brightness of the dot is determined by the echo intensity. Echo-free areas are black, the strongest echoes are displayed as white. Echoes with intermediate intensity are displayed as varying shades of grey. Importantly, the dynamic range of the machine’s display (and our eyes) is much more limited than the dynamic range of the echoes. We can only see variations of brightness on the screen over a range of about 1000 to 1, that is a dynamic range of 30 dB. This means there is a mismatch between the dynamic range of the echoes (60 dB or more) and the display. If echo intensity was translated directly into displayed intensity on the screen, we would only be able to see a 30 dB range of echo strengths. We would not be able to see any of the low-level soft tissue echoes. The machine must compress the dynamic range of the echo signal so it matches the dynamic range of the display. A typical dynamic range compression curve. The machine uses this curve to determine the brightness of the dot it displays for each echo, in other words how echo intensity maps to grey scale brightness values on the display. This diagram shows a typical dynamic range compression curve. The shape of this curve determines two important aspects of the ultrasound image the range of echoes displayed in the image; in this example the range is 60 dB; how variations in echo strength affect the level of grey displayed at each point. For example, the broken line shows that an echo 30 dB above the minimum is displayed as a light grey (almost white) value 25 dB above the display’s black point. An important feature of this compression curve is that the stronger echoes are more compressed; all the strong echoes are close to white in the image. In contrast, variations in the strength of the soft tissue echoes cause noticeable variations of brightness in the image. So, the compression curves has been designed to optimise your ability to differentiate between different types of soft tissue. Controls Ultrasound machines offer a range of compression curves. Usually, the machine chooses an appropriate curve without the need for user intervention. However, you can select different curves using the machine’s controls. (The control may be called compression or dynamic range.) The machine may display the dynamic range setting in dB or as a ‘curve’ setting (C1, C2, etc.). The images below show the effect on an abdominal image of changing the dynamic range setting. Lesson 5 – Scan Converter The main purpose of the scan converter is to place each echo in its correct position in the image memory. To do this, the scan converter must take account of factors such as echo arrival time probe geometry beam position and direction image depth (set by the user) image zoom (if used). An early abdominal ultrasound image showing visual degradation of the image because individual lines of sight (and the gaps between them) are visible. Depending on the line density (that is, the spacing between one beam and the next), echoes may not be stored in every storage location (pixel) in the image memory. If some pixels lack echo information, this can substantially degrade the image (see diagram). The scan converter therefore uses a mathematical technique known as interpolation to fill any empty pixels with appropriate values so there are no gaps in the image. Controls There are no user controls associated with the scan converter. Lesson 6 – Image Memory The image memory is a section of the machine’s digital memory where the ultrasound image is stored. It can be thought of as a two-dimensional array of digital memory locations (called pixels). The number stored in each pixel determines what shade of grey (or colour) is displayed at the corresponding point in the machine’s display. Ultrasound machines store many (typically a hundred or more) images stretching back several seconds from the most recent image. When you freeze the image (which stops the imaging process), you can then scroll back through these previous images to find the optimum one for your purpose. This cineloop function is often very useful. In 3D imaging (this will be discussed later in more detail), things become more complicated. Images are acquired from many different scan planes and these are processed to create a three-dimensional view of the anatomy. Controls As mentioned above, two controls associated with the image memory are the freeze button and the cineloop function. Another useful imaging function is the ability to zoom in on a region of anatomy. The preferred way to do this is while scanning the patient. The region of interest is identified then zoom is activated. The machine enlarges the region so that it occupies the entire image memory (see below). Since it is now scanning a smaller region, the frame rate increases. This can be a significant advantage, for example when imaging a fetal heart. This is known as write zoom or pre-processing zoom since it is applied to the echo data before it is written into the image memory. Most machines also provide a read zoom or post-processing zoom function, where part of a stored image can be magnified on the display. This type of zoom has two disadvantages relative to write zoom. First, the image may become pixellated (jagged) as it is magnified. Second, the frame rate is unchanged, since the image was acquired and stored before it was zoomed. The depth control determines the depth of tissue displayed, that is, the number of centimetres displayed on the screen. Reducing the depth may allow the machine to increase the PRF, which increases the frame rate. Generally, the depth is set so that all the tissues of interest are displayed in the image, and they occupy as much of the image area as possible. Lesson 7 – Pre-processing This term refers to changes made to the echo information immediately before it is stored as an image. Once these changes have been made and the image has been stored, the changes cannot be undone. This is different to post-processing, whose effects are reversible. Examples of pre-processing include depth write zoom frame averaging (persistence) compound imaging extended field of view imaging. The depth and write zoom controls have been described already. Frame averaging (also known as persistence) means that the machine takes the most recently acquired image and averages it with the previous image. This is not desirable when imaging fast-moving tissues such as the heart, but in other clinical areas it has a significant benefit; it reduces soft tissue speckle. As described earlier, speckle results from large numbers of low-level echoes from soft tissue combining to produce the echo displayed in the image. Speckle is often a problem since it distracts the eye and makes it difficult to see variations in echo strength in the image. Frame averaging relies on there being some movement of the tissues (for example because of respiration). Even a small amount of movement significantly alters the speckle pattern. When two (or more) different speckle patterns are combined, the resulting image contains less speckle. Compound imaging will be described in more detail in a later module. Briefly, the machine acquires several images. For each image the beams are steered in a slightly different direction. The machine then combines these images to produce a single compound image. Advantages of compound imaging include smoother speckle (for similar reasons to those described above for frame averaging) and better imaging of tissue interfaces. Extended field of view (or panoramic imaging) will also be discussed later. The probe is moved along the patient’s skin while scanning and the image is extended to create a larger field of view. Controls The user has controls for both the depth and zoom functions. Frame averaging/persistence can also be adjusted, though generally, the machine’s preset value is appropriate. Compound imaging and extended field of view are both optional modes of operation that the user can select when needed. Lesson 8 – Post-processing Post-processing functions are applied to the stored image as it is being read out of the image memory for display. Since these functions do not alter the stored image, their effects are reversible. They include colour mapping post-processing curves post-processing (read) zoom measurements. Some users find that they can see subtle variations in echogenicity better if the image is displayed using colour mapping (or B-colour). Instead of using the standard range of grey scale values from black to white, a colour scale is used. Some examples are shown below. Post-processing curves function in much the same way as the compression curve. They allow the user to fine-tune the way variations in echo strength are converted to shades of grey. An important post-processing function is measurement. Ultrasound machines allow the user to measure distances (e.g. the femur length in a fetal scan), circumferences (e.g. the abdominal circumference of a fetus), areas (e.g. the cross-sectional area of the left ventricle at a specific point in the heart cycle) and volumes (e.g. of a cyst). It is important to recognise that while these measurements may be displayed on the screen with three-figure precision, they are unlikely to be this accurate. Image artifacts (which will be discussed in the next module) cause blurring of tissue interfaces and lead to inaccuracy in the placement of the measurement calipers on the screen. It is not uncommon to find variations in measurements from the same image of up to 5% between individual sonographers. Since there can be significant variation between individuals in how they make measurements, detailed guidelines and protocols have been developed for critical application areas (such as early fetal measurements) to improve accuracy and repeatability. Controls Post-processing controls include read zoom, colour mapping, post-processing curves and measurements. Lesson 9 – Image Display and Storage Display The display used by an ultrasound machine is similar to a computer display. It can be expected to be stable and reliable and should require little or no adjustment or calibration. Image storage Since the ultrasound image is digital, several options are available for storing images. These include short-term storage in the machine’s digital memory; storage on a network such as a PACS (Picture Archiving and Communications System); removable media such as a thumb drive, CD or DVD; hard-copy record on film or paper; video (i.e. TV) format recording on DVD. Introduction to Diagnostic Ultrasound Technology Image Artifacts Image artifacts are important in diagnostic ultrasound. They can degrade the image, mislead you at times, and even cause misdiagnosis. This module describes why and how image artifacts occur and how they can be minimised. It also identifies those artifacts that may cause problems and those that can be useful diagnostic signs. It also includes a brief review of the factors that limit overall imaging performance. The module parallels Chapter 6 of the textbook. Lesson 1: Introduction An artifact is any appearance in an image that does not accurately represent the anatomy that is being scanned. Tissues may be missing from the image or distorted or misplaced, or they may be misrepresented (e.g. too bright or too dark). There may also be appearances in the image that do not correspond to real tissues. Artifacts are common in ultrasound images, and it is important to be able to recognise them. Once an artifact is recognised, it can often be ignored because it is unlikely to cause misdiagnosis or have a negative impact on the examination. Sometimes, artifacts are themselves useful diagnostic signs, providing additional information about tissues and helping to highlight tissues of interest. However, sometimes artifacts can be misleading, in which case it is particularly important that they be recognised and minimised (or eliminated, if possible). This breast lesion has produced an artifact called enhancement (arrows). The tissues deep to the lesion are brighter than similar tissues either side of them. This indicates that the lesion has lower attenuation than normal tissues. Enhancement is often diagnostically useful, since it provides additional information about the tissues causing the artifact. Why do artifacts occur? Broadly speaking there are four possible causes: the ultrasound is not behaving as the machine assumes (causing acoustic artifacts); one or more equipment settings are inappropriate. the equipment is faulty. electrical interference. We will concentrate mainly on acoustic artifacts. Imaging assumptions What assumptions does the ultrasound machine make when it converts echo information into an image? 1. The transmitted ultrasound beam is a narrow straight line. This is implied by the fact that the machine places all echoes on a line of sight in the image that corresponds to the centre line of the ultrasound beam. 2. Ultrasound travels along this straight line and does not deviate from it. We have seen that this is not true if, for example, the beam is reflected or refracted. 3. The ultrasound pulse always travels directly from the transducer to a given reflector (or scatterer) and the echo from that object travels directly back to the transducer. You can see that this assumption is required for the machine to calculate the depth of each object from the arrival time of its echo. 4. The propagation speed of the ultrasound is 1540 m/sec in all tissues. This is not true, 1540m/sec is the average value for typical soft tissue. 5. The attenuation is the same in all the tissues being imaged. Again, this is not true. 6. All echoes detected by the transducer are caused by the most recent transmit pulse. We have seen that the machine attempts to ensure this is true by using a PRF that considers the estimated penetration depth. Occasionally, however, it fails and transmits while echoes from the previous pulse are still returning. Lesson 2: Attenuation Artifacts This table shows that attenuation values vary substantially between different tissues. Liquids such as blood and urine have very low attenuation. Many abnormal tissues (for example fibrous tissues) have much higher attenuation. As explained in the previous module, the TGC function corrects for the average attenuation of normal soft tissue. However, when there are relatively small, localised regions with different attenuation to the surrounding tissues, the TGC cannot correct this, and shadowing and enhancement artifacts can occur. These are relatively common. With experience, you will learn to recognise them and to ignore them except when they are useful. Shadowing Shadowing is the term used to describe a darkened line cast behind (deep to) a region whose tissues cause more attenuation than the surrounding tissues. The transmitted intensity deep to the region is reduced compared to the intensity at the same depth in other parts of the image. So, echoes from this region are reduced in amplitude. These echoes then pass through the highly attenuating tissue on their way back to the transducer, which further weakens them. As a result, there is a darkened area in the region behind the highly attenuating tissue. Notice that the shadow is always in the direction of the ultrasound beam. Depending on the amount of attenuation, the shadowed region may be echo-free (a clean shadow) or it may still show the deeper tissues with reduced amplitude. Another possibility is that artifact echoes may be seen in the shadow (dirty shadowing). Shadowing is often useful, since it highlights tissues with unusually high attenuation such as gallstones. (Left) If the beam is wide relative to an attenuating object, only a small fraction of its energy is removed and the shadowing effect is minimal. (Right) If the object is at focus, it blocks more of the energy in the beam and the shadowing is more clearly seen. There is an important practical consideration when you are using shadowing as a diagnostic sign. Consider a small, highly attenuating object. It should be clear from this diagram that the beam should be made as narrow as possible at the depth where the object is located. This maximises the impact that it has on the intensity of the ultrasound and so it maximises the shadowing effect. Enhancement Enhancement is closely related to shadowing. It occurs when a region has lower attenuation than the surrounding tissues. The result is a brightened area deep to that region, the opposite of a shadow. Like shadowing, it is parallel to the beam direction and maximised when the beam is focussed at the depth of the object or region causing the enhancement. Like shadowing, enhancement can be a useful diagnostic sign. Enhancement (arrows) due to the low attenuation of ultrasound by the fluid in the gallbladder. Enhancement (arrows) caused by an abscess in the breast, showing that enhancement is not only caused by echo-free regions. Edge shadowing This is another type of shadowing. The cause is quite different to that described above. Edge shadowing occurs when the ultrasound beam strikes the edge of a curved structure (e.g. a cyst or a blood vessel in cross-section). A combination of reflection and refraction occurs (the exact details depend on the acoustic properties of the tissues). This causes the ultrasound beam to be deflected and broadened (i.e. defocussed), as the figure below shows. You may remember there is an inverse relationship between the width of the beam and the intensity within the beam. Broadening of the beam due to its interaction with the circular structure therefore reduces the beam intensity. As a result, echoes from this region are lower in amplitude than expected, causing a shadow. The ultrasound beam is deflected from the edge of a circular object. This spreads the beam out (i.e. defocusses it) which reduces the beam intensity. The weak echoes coming from the defocussed beam are displayed as if the beam had continued in a straight line, causing edge shadowing. Notice that it is the defocussing that causes the shadow, not the fact that the beam is deflected. Edge shadowing is very common in ultrasound images. Generally, it is of little value diagnostically, and so it is largely ignored. Edge shadowing (arrows) caused by a liquid-filled gallbladder. Edge shadowing is caused by a testis. These two examples show how a variety of curved structures can cause edge shadowing and that they need not be perfectly circular. Lesson 3: Depth Artifacts A number of artifacts lead to incorrect display of the depth of tissues. These include propagation speed, reverberation, ring-down, comet tail and range ambiguity artifacts. Propagation speed artifact When the machine calculates the depth at which to display each echo in the image it assumes that the propagation speed of sound is 1540m/sec in all tissues. In reality the propagation speed varies somewhat from one tissue type to another. Fat, for example, has a significantly lower propagation speed. When ultrasound travels through tissues with propagation speed different to 1540 m/sec, the echoes are not displayed at the correct depth. Consider the diagram below, where ultrasound travels through a fatty region in the liver. Both the transmitted ultrasound and the returning echoes are slowed down while they are travelling through it. Echoes from tissues deep to the fatty region therefore arrive at the transducer later than expected and so they are displayed deeper than they should be. Conversely, when the ultrasound passes through a region with propagation speed higher than 1540 m/sec, echoes from deeper tissues arrive more quickly than expected and so the tissues are displayed more superficially than their true position. A typical appearance is therefore as shown below, where a linear structure (e.g. the diaphragm as imaged through the liver) will appear to be displaced locally in the area behind the region with different propagation speed. A region with lower propagation speed (e.g. a fatty region within the liver) slows down both the transmitted ultrasound and returning echoes, causing the machine to display structures which are deep to the region as being farther away than their true position. The slightly higher propagation speed in these varicose veins makes the tibial plateau (arrows) appear falsely wavy in this image. Reverberation In acoustics, the word reverberation is used to describe the reflection of sound between parallel surfaces, for example the walls of a room. In ultrasound, reverberation means the same thing. It describes the reflection of ultrasound repeatedly back and forth between two strongly reflective parallel tissue interfaces, as shown in the diagrams below. Examples include the interface between the transducer and skin surface when there is not sufficient coupling gel, tissue-bone interfaces, calcified structures, fibrous and air-filled structures. The other requirement is that the ultrasound must strike the tissue interfaces at 90°. When ultrasound is reflected back and forth between a tissue interface and the transducer face, a series of equally spaced echoes are seen, progressively weakening in brightness with depth. A similar appearance is seen when the ultrasound reverberates between two tissue layers. As the first of the two images below shows, when reverberation occurs a series of equally spaced echoes return to the transducer. The reflecting surfaces are displayed in their true positions then replicated a number of times. If soft tissue lies between the reflecting interfaces, it too is replicated, as shown in the second image. Reverberation has caused a series of parallel echoes to appear in the superficial portion of a haemorrhagic renal cyst (arrows). In this image, Reverberation has caused soft tissue echoes to be replicated and appear inside the urinary bladder (arrows). It is relatively easy to recognise that the extra echoes in these images are caused by reverberation. This is because they appear in liquid filled regions which should be echo free. Commonly, however, echoes caused by reverberation are displayed in soft tissue regions (in the liver, for example). It can be difficult or impossible to distinguish between artifact echoes and genuine echoes. So, reverberation is a common cause of image degradation. With experience, sonographers learn to minimise reverberation by manipulating the ultrasound probe (tilting it and modifying the applied pressure) and finding the best window into the anatomy. Ring-down artifact Ultrasound reflects between multiple gas bubbles, creating a continuous series of echoes returning to the transducer. The result is a bright echo extending in depth from the gas bubbles along the beam direction. A special case of reverberation occurs when the reflectors are small gas bubbles (as commonly found in the gastrointestinal tract, for example). As this diagram shows, multiple echoes return to the transducer as the ultrasound reverberates back and forth among the gas bubbles. The reflection of ultrasound by gas bubbles is highly efficient (very little energy is lost during reflection), so the echoes remain strong for a considerable time. There is also evidence that gas bubbles can resonate when exposed to ultrasound frequencies. This substantially increases the strength and longevity of the echoes. The result is a bright line of echoes parallel to the probe extending to a considerable depth (see images below). This is known as the ring-down artifact. Note that discrete echoes from individual bubbles cannot be identified, with the artifact instead being a relatively continuous line of bright echoes parallel to the beam direction. A typical example of ring-down artifact caused by gas in the liver. A dramatic example of ring-down caused by gas-filled bowel. Comet-tail artifact Comet tail artifact in a gallbladder filled with ‘sludge’ A closely related artifact is the comet tail artifact. This is like the ring-down artifact but shorter-lived. It is seen as a relatively short ‘tail’ extending into the tissues parallel to the beam, tapering and fading in a manner similar to the tail of a comet. It is generally associated with small calcifications and other crystalline structures. Ultrasound reverberates within these crystalline structures, producing a series of echoes. Energy is lost quite rapidly and so the reverberation is short-lived and the artifact fades quickly with depth. Metal objects such as biopsy needles and guide wires can also cause ring-down and comet tail artifacts. Range ambiguity As discussed in an earlier module, the ultrasound machine must not transmit until all detectable echoes from the previous transmit pulse have returned. If this rule is violated, echoes from the earlier transmit pulse are mixed with echoes from the new transmit pulse. The machine cannot differentiate between them and so there is ambiguity about the depth from which each echo comes. Generally, the machine does a good job of satisfying this rule, but occasionally it fails. This can happen, for example, when scanning through liquid (e.g. a filled urinary bladder), since the attenuation is lower than expected. (Top) Timeline showing the transducer transmit and echo signals. The machine is waiting until there are no more detectable echoes before it transmits again, so there is no ambiguity regarding the range (depth) at which to display each echo. For the echo shown in this diagram, the depth is d = ct/2. (Bottom) The machine has doubled the PRF (Pulse Repetition Frequency = number of transmit pulses per second). Consequently it is transmitting while echoes from the previous transmit pulse are still being received. This causes range ambiguity. The diagram above shows the echo signal received from a reflector in normal operation and when there is range ambiguity. In normal operation the echo arrives at time t after each transmit pulse so the machine calculates the depth as d = ct/2. When the PRF is too high (bottom), the echo still arrives at time t after each transmit pulse. However, the machine doesn't know whether the depth should be calculated from time t or from t' since it doesn't know which transmit pulse generated the echo. As a result the echo is displayed twice, at the correct depth d = ct/2 and more superficially at depth d' = ct'/2. A liquid-filled bladder scanned with normal machine settings. The focusing has been changed dramatically, inducing range ambiguity. The regions labelled X and Y in the posterior wall are duplicated as x and y. Note the bladder wall is scaled down in width (but not depth) due to the scan format of the curved probe. If a linear probe had been used, the true image and the artifact would be identical in size and shape. Lesson 4: Beam Dimension Artifacts Beam width artifact As discussed in an earlier module, the ultrasound machine always displays an object as though it were on the centreline of the beam, no matter where in the beam it actually lies. A time sequence showing three successive positions of the beam as the machine scans the tissues from left to right. A single object produces an echo each time, with each echo being displayed in a slightly different lateral position. In reality the lines of sight are more closely spaced than shown here, so the dots in the image merge to form a line. This diagram shows an important consequence. As the beam scans, it steps from one end of the probe to the other. It moves in small steps, so a number of successive beams ‘see’ a given object. As a result the object is displayed several times in different lateral positions while the depth remains constant. It therefore appears in the image as a line whose width is equal to the beam width. Generally, the beam width is significant (several millimetres), so this causes substantial degradation of the image. Image of an ultrasound phantom showing point reflectors displayed as bright lines. Note that the lines are curved since this is a curved probe. This image shows a scan of a tissue equivalent phantom containing a number of equally spaced point targets (fine nylon lines perpendicular to the scan plane). The lateral smearing of these targets in the image is clearly seen. Every object is ‘smeared out’ laterally in the image by an amount equal to the beam width. This diagram shows the effect of beam width on the image when a tissue interface is scanned. The tissue appears 'fuzzy' in the image and thicker than its true size. An exception to this is when the interface is scanned at perpendicular incidence. Since the lateral smearing exactly coincides with the tissue itself, it does not visibly degrade the image. Image resolution is maximised when tissues are scanned at 90°. The effect of beam width on an image of the portal vein. The resolution is good where the beam is perpendicular to the vessel (top arrow). Where the angle is far from perpendicular (lower arrow) the image is fuzzy and poorly defined. This clinical image shows the effects of beam width. The portal vein walls are well resolved in the area where the beam is perpendicular to them (top arrow). However, in areas where the incidence angle is far from 90°, the walls are fuzzy and less well-defined (lower arrow). Unfortunately, this fuzziness reduces the accuracy of measurements. Deciding just where to position measurement markers requires judgement, and this introduces an error that varies from one practitioner to another. In summary, beam width degrades image resolution and reduces the accuracy of measurements. The larger the beam width the worse these effects are. Beam width effects are minimised when tissues are scanned at 90°(perpendicular incidence). Beam profile. This logarithmic plot of intensity across the beam shows the sidelobes either side of the main beam. Sidelobe artifact A previous module introduced the concept of sidelobes. These are unwanted additional beams symmetrically positioned either side of the main beam. Sidelobes are always present. Fortunately, they are considerably weaker than the main beam. Nevertheless, when a strong reflector is present it may produce a detectable echo when it is seen by a sidelobe. The sidelobes (lighter shade of blue) move with the main beam (darker shade of blue) as it scans along the probe face to create the image. (Left) One of the sidelobes encounters a strong reflector; the machine displays it as though it is in the main beam (i.e. to the left of its true position). (Centre) The main beam displays the object in its true position. (Right) The other sidelobe encounters the object; the machine displays it to the right of its true position. A sidelobe artifact in the urinary bladder. Strong echoes from an area where the beam is at perpendicular incidence (centre arrow) cause sidelobe artifacts to be displayed either side of the true location (left and right arrows). Note they are at a constant depth. As the diagram above shows, the result is lower-level echoes either side of the reflector. Sidelobe artifacts are commonly visible when they occur in liquid-filled areas, as shown in this image. At other times they may be present but unrecognised. Ultrasound machines use a technique known as apodisation to reduce the amplitude of sidelobes. The transmit pulse and receive gain for the outer elements of the aperture are reduced. This reduces the sidelobes at the cost of somewhat increasing the width of the main beam. Slice thickness artifact In an earlier module we also discussed slice thickness and its effect on image quality. The diagrams below illustrate the concept of slice thickness. Recall that slice thickness is determined by how well (or poorly) the beam is focussed in the direction across the width of the transducer. Here the beam is in its initial position at the left-handed end of a linear array transducer. To create the image, the machine steps the beam from left to right in the direction shown by the arrow. Due to the large slice thickness dimension of the beam, it sweeps through a substantial volume of tissue. It is common to think of the ultrasound image as showing a thin slice through the patients anatomy. In reality, the slice has considerable thickness, as explained above. The image therefore displays echoes not only from the tissues in the ideal scan but also from nearby tissues. When the image contains anechoic structures (e.g. blood vessels, urinary bladder) slice thickness artifact echoes are more easily seen. An example is shown below. Notice that rotating the probe 90° will often clarify where the artifactual echoes come from. An example of slice thickness artifact. The echoes seen within the aorta (arrows) come from the tissues beside the vessel at the same depth. Diagram showing the aorta passing through the scanned volume. The echoes highlighted in the previous image come from tissues such as those indicated (arrow). These are adjacent to the vessel but they are displayed as if they come from inside it. The probe has been rotated 90 degrees to produce a short axis view of the aorta (arrow). The artifact has disappeared. Speckle As discussed previously, speckle is the term used to describe the echo texture typically seen in soft tissue. It occurs because scattered echoes from the small structures within the soft tissue add randomly. Speckle makes it more difficult for the user to see small structures and subtle variations in soft tissue echogenicity (brightness) and echo texture. Ultrasound machines have various technologies whose aim is to reduce speckle and make soft tissue look smoother. These include frame averaging (also called persistence) compound imaging image processing. Lesson 5: Beam Path Artifacts Mirror image artifact Reflection of ultrasound was discussed in Module 1. A typical reflector is an interface between two tissues that have significantly different acoustic impedances (e.g. the diaphragm, which is the interface between the soft tissue in the liver and the air-filled lungs). The geometry of reflection is very simple; the incident angle is equal to the reflected angle. Since the machine assumes ultrasound always travels in a straight line, it displays the echoes in the wrong location. As the beam scans, it therefore creates a mirror image of the tissues beyond the interface. This is known as a mirror image artifact. It is quite common, especially when scanning the liver. Examples are shown below. A liver lesion and its mirror image behind the diaphragm. Note the mirror image is always deeper than the reflective tissue that causes it. The mirror image is distorted because of the curvature of the diaphragm. Refraction artifact Refraction was also discussed in Module 1. It refers to the change in the direction of the ultrasound beam when it passes through an interface between two tissues with significantly different propagation speeds. The change of direction is more pronounced when the angle of incidence is large (well away from perpendicular incidence) and when the difference in propagation speed is large. If the beam is at perpendicular incidence, there is no refraction, regardless of whether there is a difference in propagation speed between the two tissues. The ultrasound machine assumes that ultrasound travels in a straight line, so it displays the object at a slightly different position than its true location, as shown in this diagram. The effect of refraction on the ultrasound image is usually small and generally not noticeable. However, particularly striking refraction artifacts can occur in the abdomen when a transverse scan is done with the transducer located on the patient’s midline, and in transthoracic echocardiography (see the textbook for details). Equipment and electrical artifacts The most likely cause of equipment malfunction is damage to the probe. If a probe is dropped or bumped, one or more transducer elements may fail. This can cause an area of drop-out or other abnormalities in the image. The drop-out will be fixed in position relative to the transducer face. As the probe is moved, the artifact will move with it, making it clear that it is an equipment problem. Electrical interference occasionally occurs, generally when portable equipment is being used in electrically noisy environments such as a Neonatal Intensive Care Unit. It is easy to identify as artifact since the patterns displayed are not related to the tissues being scanned. Lesson 6: Performance Limitations Image quality The quality of an ultrasound image can be considered under three different headings. Spatial resolution. Contrast resolution. Temporal resolution. Spatial resolution This term can simply refer to the subjective sharpness and clarity of the image as seen by the user. However, more objectively, it is defined as follows. “Spatial resolution is the minimum spacing required between two small objects (usually referred to as ‘point targets’) so that they will be seen as separate objects in the image. If the objects are spaced more closely than this, they blur together and appear as a single object. We then say they have not been ‘resolved’ in the image. For ultrasound images, spatial resolution is further divided into two components – axial and lateral resolution. We also need to remember the third dimension of the beam – slice thickness. Axial resolution Axial resolution is the minimum separation in depth that is needed for two point targets lying along the same line of sight to be resolved. To express this mathematically, we need to look at the echoes received from the two point targets (see the diagrams below). The two targets will be resolved in the image (seen as two separate objects) as long as their echoes don’t overlap in time. So axial resolution is the minimum distance between the point targets for which their echoes do not overlap. The points targets lie on the same line of the sight at different depth. The echo signal from the two-point targets. As long as the echoes do not overlap, the two targets will be resolved in the image. They will not overlap as long as their separation in time (t2 – t1) is larger than the pulse duration. We can see that the two echoes will not overlap as long as their arrival times (t 1 and t2 ) differ by more than the transmit pulse duration (τ), that is (t2 - t1) > τ We saw in an earlier module that there is a simple relationship between the depth of an object in the tissues (d) and echo arrival time (t) t = (2d) / c Combining these two equations we see that the two targets will be resolved as long as 2(d2 - d1) / c > τ or (d2 - d1) > (cτ / 2) The axial resolution of the ultrasound image is therefore (cτ/2) where τ is the transmit pulse duration. An obvious conclusion is that the transmit pulse should be as short as possible, since this gives the best axial resolution. This is most easily achieved by making the ultrasound frequency as high as possible, consistent with adequate depth of penetration. We can also see that the axial resolution is independent of depth and is the same throughout the image, since the transmit pulse duration is constant. (Notice this is not true for lateral resolution, which varies with depth.) Lateral resolution We saw earlier that the image of a point target is smeared out laterally by an amount equal to the beam width (see diagram below). So it should be clear that two point targets at the same depth need to be separated by at least the beam width in order to be resolved. Lateral resolution is therefore equal to beam width. The finite width of the ultrasound beam causes a point target to be smeared out laterally in the image by an amount equal to the beam width (W). The separation (X) must therefore be larger than the beam width for the echoes to remain separate in the image. We have seen that beam width, like pulse duration, is minimised when a high frequency is used. So using the highest possible frequency gives the best spatial resolution possible. Now let's see how the different spatial resolution values compare. Consider a 6 MHz probe with a pulse duration of 0.5 μsec, an aperture of 2 cm, transducer width of 0.6 cm and a focal depth 0f 4 cm. The axial resolution is 0.39 mm, the lateral resolution is 1.25 mm and the slice thickness is 4.2 mm. Notice that the axial resolution is around three times smaller (better) than the lateral resolution and the lateral resolution is around three times smaller (better) than slice thickness. These are typical values. They highlight the fact that the axial resolution is better than the resolution in other directions. For this reason: “Perpendicular incidence is always preferred when imaging tissues and making accurate measurements.” Contrast resolution Contrast resolution refers to the user’s ability to detect small differences in soft tissue echogenicity. Every ultrasound image contains artifacts (beam width, slice thickness, sidelobe, reverberation etc.). Often, echoes due to artifacts are not identifiable unless they fall into an area that should be echo-free, but they degrade the image whenever they occur. As mentioned earlier, it can be useful to think of starting with an ideal ultrasound image which is then degraded by the presence of artifacts. The more it is degraded, the worse the image quality will be and the harder it will be for the user to see small variations in soft tissue echo characteristics. So it should be clear that image artifacts worsen contrast resolution. For this reason, new modes of operation which reduce artifacts (like harmonic imaging and spatial compound imaging) have become widely used. These will be described later. Speckle also degrades contrast resolution. Again, modes of operation like frame averaging and compound imaging have been introduced to reduce speckle. To maximise contrast resolution, it is also important that the machine’s settings are optimised, particularly the TGC, gain, dynamic range and ultrasound frequency. Incorrect setting of these degrades contrast resolution. Temporal resolution Temporal resolution refers to the ability of the ultrasound machine to produce good quality images when the patient’s tissues are moving, or when the probe is moving relative to the tissues. Tissue movement is normal. Think of the heart, pulsating arteries, respiratory movement of the abdominal organs and fetal movement. It is also common for the user to sweep the scan plane quickly through a region such as the abdomen to get an overview of the anatomy. If the temporal resolution is not adequate, the image is blurred and there may be a noticeable time lag between movement of the probe and the image. Clearly temporal resolution is important, and ultrasound machines are designed to maximise it as far as possible. The parameter with the most direct effect on temporal resolution is the frame rate, the number of images per second. This is generally displayed on the screen, with the units being Hertz or frames per second. The frame rate depends on the machine's Pulse Repetition Frequency (PRF, the number of transmit pulses transmitted each second), the number of lines in each image (L) and the ensemble length (E, the number of transmit pulses required for each line). E is one for normal grey scale imaging, since a single transmit pulse provides the echo data for a complete line in the image. However, it is higher for other operating modes. For example, it is two for harmonic imaging, two or more for multiple focus markers, and higher for compound imaging and colour Doppler. The frame rate is therefore reduced whenever one or more of these modes are activated. Fortunately, many manufacturers have introduced a technical ‘fix’ for low frame rate. Many machines have the ability to form several beams at the same time (for example four). This increases the frame rate substantially. For example, if the multiple beam factor (M) is four, the frame rate is increased by a factor of four. Now we can pull together all the factors that affect frame rate. As we saw earlier, the machine transmits as many pulses as possible each second, subject to the requirement that it must wait until all detectable echoes have returned to the probe before transmitting again. The PRF can therefore be calculated from the penetration depth (P) as follows: PRF = c / (2P) The number of pulses required to make a single image (taking into account multiple beam forming) is: N = (L × E) / M Multiplying this by the number of images each second (the frame rate, FR) therefore tells us how many transmit pulses are needed each second: pulses per second = N × FR = (L × E × FR) / M This must be equal to the PRF, so we can write: (L × E × FR) / M = PRF = c / (2P) which can be rearranged as: (FR × P × L × E) / M = c/2 = 77,000 cm/s The left-hand side of this equation is equal to a fixed physical quantity (c/2). So it follows that we must reduce one or more of the following if we want to increase the frame rate: P, the penetration depth (e.g. by increasing the ultrasound frequency). L, the number of lines in each image (e.g. by reducing the field of view). E, the number of pulses needed for each line (e.g. by turning off harmonics). Example: Consider a probe with the following operating parameters: penetration P = 10 cm; lines per image L = 200; ensemble length E = 2; multiple beam factor M = 4. Using the equation above, we find that the frame rate is 77 Hz. Frame rate is not the only factor that affects temporal resolution. Frame averaging (or persistence) and compound scanning both work by combining multiple images together. So even if the frame rate is high (the rate at which individual images are acquired) the effective frame rate will be lower. In compound scanning, for example, up to nine images are acquired and merged to form a single compound image, so the effective frame rate will be reduced by a factor of nine. Summary Artifacts are often seen in ultrasound images, particularly in fluid-filled areas and other areas with low echo levels. Some artifacts are useful indicators of tissue properties – shadowing and enhancement indicate regions of abnormally high or low attenuation respectively, ring-down artifact usually indicates the presence of gas bubbles, comet tail artifact is often caused by calcifications, etc. Other artifacts have the potential to mislead and possibly cause misdiagnosis. They may also conceal useful information (think of an image with widespread ring-down artifact, for example). In fact, artifacts are present in every image. Aside from the recognisable artifacts referred to above, what is the impact of this? It can be useful to think of the image as being made up of an ‘ideal’ ultrasound image and an overlay of artifact echoes. Clearly, the fewer artifacts there are the more readily subtle variations in tissue echogenicity and texture can be seen. Some of the newer imaging modes (harmonics and compound imaging, for example) can substantially reduce artifacts, and this explains why they are widely used. Artifacts are often a product of the geometric relationship of the probe to the tissues (think of reverberation, for example). Moving or angling the probe can therefore change (or even eliminate) many artifacts. So probe manipulation is useful both to identify suspicious appearances in an image as artifacts and as a way to reduce or eliminate the