Musculoskeletal Ultrasound Marnix T van Holsbeeck, Joseph H Introcaso
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Physical Principles of Ultrasound Imaging1

Thomas Gauthier,
Joseph Introcaso,
Marnix van Holsbeeck
 
INTRODUCTION
The human ear can detect sound waves with frequencies in the approximate range of 20 hertz (Hz) to 20 kHz. Ultrasound waves are defined as sound waves having a frequency greater than that which is audible by the human ear, i.e. above approximately 20 kHz. Modern diagnostic ultrasound scanners operate in the frequency range of 1–15 MHz, with the specific choice of an operating frequency range depending on the application, e.g. small parts versus abdominal imaging.
The practical use of ultrasonography has slowly evolved, as new technologies have emerged to eventually enable real-time B-mode imaging (a B-mode image is defined as a cross-sectional image of tissues and organ boundaries within the body). Technological developments, encompassing World War II era invention of the SONAR (SOund Navigation And Ranging) system and, more recently, the introduction of digital computing in modern diagnostic ultrasound scanners, have paved the way to real-time ultrasound imaging as we know it today.
Ultrasonography is now widely used in medicine for both diagnosis and guidance of interventional procedures, such as biopsies and fluid aspirations. It is particularly effective in imaging body soft tissues, from superficial structures, such as muscles and tendons, to deeper structures such as abdominal organs (liver and kidney).
 
FUNDAMENTAL PRINCIPLES
Both sound waves and X-ray photons are forms of energy transmission. However, this is where the similarity between the two ends. Their interactions with matter are quite different. The way in which these two forms of energy interact with matter determines how they can be used in medical imaging.
Unlike radiography where the image is produced by energy transmitted through the body, ultrasound imaging most commonly utilizes energy reflected back to the source to produce an image. This is referred to as pulse–echo imaging. X-rays are best transmitted through a vacuum, whereas sound requires matter for its transmission. The speed at which X-ray photons travel is constant. On the other hand, the speed of sound varies with the type of material in which sound propagates. Table 1.1 lists the speed of sound in a variety of materials. The factors that determine the speed of sound in a material are the material's density and stiffness. Mathematically, the speed of sound in a material is equal to the square root of the ratio of the material's stiffness to the material's density. Hence, low density and high stiffness of a material lead to high speed of sound.
Sound is reflected at interfaces between materials. Two factors influence reflectivity: the acoustic impedance of the two materials and the angle of incidence of the sound wave. The angle of incidence of a sound wave is defined as the angle between its direction of propagation and a line perpendicular to the interface between the two materials. The acoustic impedance of a material is equal to the square root of the product of the material's density and stiffness. Reflectivity is greatest at interfaces between materials with dissimilar acoustic impedance, as exhibited by the following equation:
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Table 1.1   Speed of sound through various substances.
Transmitting substance
Speed of sound (m/sec)
Air
331
Fat
1,450
Water
1,540
Liver
1,549
Blood
1,570
Muscle
1,585
Cortical bone
4,080
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Fig. 1.1: Angle of incidence.
Reflection of sound and angle of incidence (Φ).
Table 1.2   Acoustic impedance of various materials.
Material
Acoustic impedance (g/cm2 sec × 10−5)
Air
0.0004
Fat
1.38
Water
1.54
Blood
1.61
Muscle
1.70
Cortical bone
7.8
where R is referred to as the amplitude reflection coefficient; 100 × R represents the percentage of the incident ave's amplitude which is reflected at the interface between materials of acoustic impedance Z1 and Z2, respectively, assuming an incidence angle of 0° (normal incidence). In clinical practice, the ultrasound wave may not necessarily travel through interfaces under normal incidence condition.
The data in Table 1.2 indicate that interfaces between soft tissue and air should be highly reflective, which is what we observe clinically. Keeping this in mind, it is clear why a coupling gel must be used to ensure contact between the ultrasound transducer and the patient's skin. The previous equation tells us that 99.9% of a sound beam is reflected at any tissue–air interface. In areas of poor contact with the skin, where an air gap exists, essentially no energy is available for imaging.
Reflection of a sound wave away from the transducer varies greatly with the angle of incidence. The least reflection away occurs with the sound wave perpendicular to the reflecting interface, i.e. under normal incidence condition. As the incidence angle increases, the amplitude or intensity of the reflected sound wave increases. Beyond a certain angle, referred to as the critical angle, the entire sound wave is reflected. In addition, the direction of the reflected wave is determined by the angle of incidence (Fig. 1.1). This is important to keep in mind when imaging an object with a curved surface, such as the humeral head, the femoral condyles, or the diaphysis of a long bone. As the angle of incidence increases, the beam will be reflected away from the transducer and will not contribute to the image (Fig. 1.2). Therefore, the optimal scanning pattern for a curved surface is an arc that keeps the beam perpendicular to the surface of the object to be imaged (Fig. 1.3).
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Fig. 1.2: Suboptimal musculoskeletal ultrasound.
Improper imaging of a curved object.
Under “non-normal” incidence conditions, where the angle of incidence is > 0, refraction of an ultrasound wave occurs at the interface between two materials with two different speeds of sound. Refraction is a change in the propagation direction of a sound wave when it crosses the interface between two dissimilar materials provided that the incident wave's propagation direction is not perpendicular to the interface. The amount of “bending” a sound wave experiences at a boundary where the speed of sound changes, is related to how much the speed of sound is changing across the boundary between the two substances. Refraction of ultrasound waves can cause the image of a target to be displaced from its true relative position in the patient. In most circumstances, the error introduced by refraction is not significant. However, under certain conditions, the true location of an object will differ significantly from its imaged position.
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Fig. 1.3: Optimal musculoskeletal ultrasound.
Optimal imaging of a curved object.
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As a sound wave travels in a material, a portion of its energy is absorbed by frictional forces. The energy is converted to heat and no longer contributes to the imaging process. Viscosity, relaxation time, temperature of the material, and the frequency of the sound wave, all affect absorption. Of these factors, the one that can be modified in the clinical setting is the operating frequency of the transducer. The degree of absorption of a sound beam in soft tissue is directly proportional to its frequency. If the frequency of the sound wave is doubled, absorption will double, therefore, decreasing the depth of penetration of the ultrasound waves. On the other hand, spatial resolution increases with frequency. The optimal transducer for a specific examination is the one operating at the highest possible frequency while penetrating to the desired depth within the soft tissues being examined. Appropriate transducer selection will result in images with greatest possible spatial resolution.
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Figs. 1.4A to D: Transducers.
(A) Sector scanning transducer. (B) Annular array transducer. (C) Radial array transducer. (D) Linear array transducer.
 
EQUIPMENT
Different types of ultrasound transducers are currently available, including (but not limited to) mechanical sector scanners, mechanical linear scanners, linear arrays, curved arrays, sector or phased arrays, and matrix arrays (Figs. 1.4A to D). Linear and matrix arrays are currently preferred for musculoskeletal ultrasound imaging. A matrix array is a rectangular transducer with the crystal divided in a matrix of rows and columns of transducer elements. Due to the fact that matrix arrays allow beam forming in three dimensions, and allow focus in elevation (beam width) and lateral planes, images can be obtained with higher spatial and contrast resolution. These matrix transducers will in all likelihood soon overtake other transducers in imaging of the musculoskeletal system (Fig. 1.5). Even though all of the above-mentioned transducers differ in shape, size, field of view, and clinical applications which they are most appropriate for, they all are made of common building blocks which include a piezoelectric crystal.
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Fig. 1.5: A conventional linear transducer compared to a matrix area transducer.
The footprint of a linear area transducer (top) is made up of a row of individual crystals. The transducer has a narrow zone of focus. The footprint of a matrix area transducer (bottom) in contrast contains rows and columns of crystals. Beam forming is now possible in three dimensions and beam thickness is more uniform. These technological changes make the image more consistent from the near field to the far field.
Piezoelectric materials, discovered by Pierre and Jacques Curie in 1880, have unique mechanical and electrical properties. When a positive or negative voltage is applied across a piezoelectric plate, it expands or contracts. Conversely, when compressed or stretched by an external force, this plate generates a positive or a negative voltage. This unique property makes piezoelectric materials suitable for both transmitting and receiving acoustic waves. The most common piezoelectric material used in modern ultrasound transducers is a synthetic ceramic material, PZT (lead zirconate titanate). Both sides of the piezoelectric plate are then coated with conductive paint to serve as electrodes. These electrodes are connected to electrical leads which carry an oscillating voltage, either sent from a signal generator to the piezoelectric material to create an acoustic wave, or received from the piezoelectric material as a result of the backscattered acoustic wave received by the same. Ultrasound transducers also typically contain an acoustic lens, whose role is to complement beam focusing achieved electronically by controlling the phase of voltages fed into individual piezoelectric transducer elements. In a linear array transducer, 5acoustic beams can only be controlled electronically in the scan plane. An acoustic lens ensures adequate, fixed mechanical focusing in the elevation plane, which has to be carefully chosen for the anticipated use of the transducer given that it cannot be dynamically controlled.
The operating characteristics of an ultrasound transducer are described by its center frequency (f0) and bandwidth (BW). The −6-dB bandwidth of an ultrasound transducer is the range of frequencies over which the output peak-to-peak acoustic pressure for a given applied peak-to-peak voltage is greater than half of the maximum output peak-to-peak acoustic pressure. Nowadays, most commercial ultrasound transducers have a high enough bandwidth to accommodate transmitting (and receiving) acoustic waves of multiple center frequencies. They are referred to as multifrequency ultrasound transducers. An ultrasound transducer center frequency and bandwidth may be combined in the Q factor, which is defined as f0/BW. Two properties of a transducer are described by the Q factor: the purity and the persistence (a.k.a. ring down time) of the sound produced by the transducer. Ring down time refers to the time it takes for the piezoelectric material to stop vibrating. Piezoelectric materials with a high Q factor produce a pure (i.e. narrow bandwidth) sound but have a long ring down time. On the other hand, piezoelectric materials with a low Q factor produce a broad bandwidth sound which contains a wide variety of frequencies, but have a short ring down time. Since broad bandwidth equals short transmitted acoustic wave, which, in turn, equals best image axial resolution, low Q factor piezoelectric materials are most desirable for use in medical ultrasound imaging.
The sound beam emitted by the transducer has a shape which changes with its distance from the transducer (Fig. 1.6). In the near field, the borders of the sound beam are almost parallel, referred to as the Fresnel zone. At a certain distance from the transducer, the beam diverges, referred to as the Fraunhofer zone. The point of transition varies with the frequency and width of the sound beam. As the frequency of the acoustic wave and the width of the acoustic beam increase, the Fresnel zone extends deeper within the imaged medium. Another feature of a typical sound beam profile is side lobes, which are almost parallel to the main beam. The intensity of these side lobes is usually <1% of the intensity of the main beam and, thus, side lobes are rarely clinically significant. When they do contribute to an image, they have the same effect as beamwidth artifact, described in Chapter 2.
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Fig. 1.6: Shape of the sound beam for a conventional transducer.
Fresnel and Fraunhofer zones of a sound beam.
 
IMAGING
Ultrasound images are composed of a matrix of picture elements. Grayscale images are produced by representing backscattered echoes returning to the transducer as picture elements (pixels). The pixels' intensity (or brightness) is then calculated on the basis of the intensity of the associated backscattered echo. When imaging a given target, the location of the associated pixel or group of pixels in the ultrasound image is determined by the combined knowledge of speed of sound in the imaged medium and time of flight for that particular target. Time of flight is the time it takes for an ultrasound wave to propagate from the transducer to the target and back.
In ultrasound imaging, there are two types of spatial resolution which affect image quality: axial and lateral. Axial resolution is defined as the shortest distance between two targets aligned on the beam axis for which those two targets can be displayed as two separate objects. Axial resolution is determined by the length of the acoustic pulse, which itself is governed by transducer operating frequency and Q factor. In clinical practice, the operator only has control over the operating frequency. The higher the operating frequency, the better the axial resolution.
Lateral resolution is defined as the shortest distance between two targets located at the same depth for which those two targets can be displayed as two separate objects. Two key settings which control beamwidth are adjustable by the operator. The focal depth and the number of focal zones simultaneously activate. Beamwidth is tightest in the vicinity of the focal depth. When multiple focal zones are activated simultaneously, beamwidth can be kept tight over a greater depth. Activating more or fewer focal zones controls the depth of field, with the depth of field increasing when more focal 6zones are activated. However, the disadvantage of activating multiple focal zones is a resulting decrease in image frame rate during real-time imaging.
Divergence of the sound beam, absorption, and scattering, all play a role in attenuation of the ultrasound beam within the body. A sound beam undergoes an exponential decrease in intensity as it passes through tissue. If uncorrected, this attenuation would result in images that are markedly decreased in diagnostic information with increasing distance from the transducer. Correction for attenuation is made by amplifying echoes returning to the transducer using an exponential function based on the time of flight. This type of correction is called time gain compensation. The examiner may modify the correction function using controls on the ultrasound unit to optimize the information displayed.
Newer ultrasound units utilize spatial compounding as another means of improving ultrasound image quality. A better contrast-to-noise ratio can be obtained by insonating the tissues from different angles and then combining these images into a single spatially compounded ultrasound image. Currently available portable computing power allows compound image acquisition from nine different angles to be incorporated into real-time imaging of the soft tissues. This technique reduces artifact resulting from anisotropy (seeChapter 2); it reduces scatter, improves contrast resolution, and provides better definition of the margins of lesions. A beam steering option separated from spatial compounding is available on some equipment. This beam steering allows us to counteract anisotropy of tendons when scanning over particularly rough or sloping bone surfaces.
 
Doppler Flow Imaging
Doppler ultrasound imaging is rapidly finding increasing application in the diagnosis of disorders of the musculoskeletal system. Local hyperemia is often associated with focal tendon lesions and synovial reaction in inflammatory arthropathies. Granulation tissue at sites of healing is also quite vascular. These alterations in tissue vascularity can often be observed using Doppler ultrasound imaging.
The basic principle of Doppler ultrasound lies in the observation that the frequency of a sound beam reflected back to the source is altered when it encounters a moving object. The frequency increases if the object is moving toward the source and decreases if it is moving away from the source. The change in frequency is proportional to the velocity of the object, which can be determined by the following equation:
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where ΔF is the change in frequency of the sound beam, FT is the transmitted frequency, v is the velocity of the object, c is the speed of sound within the substance, and θ is the angle of incidence of the sound beam relative to the direction of motion of the object. An important aspect of this equation to note is the effect of the angle of incidence of the sound beam on frequency shift. The greatest frequency shift occurs when the sound beam is traveling along the same trajectory as the object being evaluated. No frequency shift occurs when the angle of incidence of the ultrasound beam is 90° (cos 90° = 0); therefore, no flow will be detected.
In musculoskeletal imaging, the two most common ways of evaluating flow-related information involve either color flow or power Doppler imaging. Color flow Doppler images present the frequency shift data by converting them into a spectrum of color, which encodes both directional and velocity information. The benefit of having both types of information is the principal advantage of color flow Doppler imaging. The disadvantages of color flow are that it is extremely sensitive to the angle of incidence of the sound beam, aliasing can occur, and noise will result in image artifacts. Aliasing is an artifact related to the pulsed Doppler technique when large frequency shifts occur. If the sampling frequency [pulse repetition frequency (PRF)] is too low (less than half of the frequency shift encountered), then aliasing will result. This can be dealt with by either increasing the PRF or increasing the Doppler angle, which will result in a smaller frequency shift.
Power Doppler differs from color flow in that it displays in color information on the amplitude (power) of the Doppler signal, rather than the frequency data directly. This approach has several advantages. Power Doppler is less angle dependant; no aliasing is experienced, and noise results in much less image degradation. In addition, power Doppler is much more sensitive to slow flow. The major disadvantages of this technique are that the direction and velocity information is lost. Fortunately, in musculoskeletal ultrasound, direction and velocity information is usually of little value. Therefore, power Doppler usually proves to be the more valuable technique.
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Extended Field-of-View Imaging
One of the significant limitations of ultrasound imaging has been its limited field of view. However, recent advances in computer hardware and software have made possible extended field-of-view ultrasound imaging without the use of an articulated arm or special position sensors (SieScape, Siemens Medical Systems, Iselin, NJ). High-speed video image processing hardware analyzes sequential frames acquired from a linear array transducer as it is moved slowly over the region being examined. The direction of motion is determined by dividing each image into a group of blocks of equal size and comparing sequential image frames. Each block is examined to determine a motion vector for that block, which is dependent on the degree of change occurring within that block. The individual motion vectors are then analyzed to determine the overall direction of transducer motion. If all vectors are in the same direction and have equal magnitude, then transducer motion must be linear. Transducer motion with a rotational component will yield vectors which vary in direction and magnitude. This type of imaging proves most valuable when evaluating long muscles, tendons, and vessels. Measurement across the extended field of view is accurate and proves useful in selection and preoperative assessment of soft tissue grafts.
 
Tissue Harmonic Imaging
In conventional B-mode imaging, an acoustic wave is transmitted at a fundamental frequency f0 and an image is formed using the same frequency component f0 of the backscattered echoes. On the other hand, tissue harmonic imaging relies on the nonlinear propagation in tissue of an acoustic wave and an image is formed using the second harmonic component 2f0 of the backscattered echoes. This is achieved by filtering the backscattered echoes and removing all frequency components (including fundamental) except for the second harmonic component. In order to achieve complete separation of the fundamental and second harmonic components, the bandwidth of the transmitted acoustic wave must be narrow enough to prevent these components from overlapping. A narrow bandwidth translates into a long transmitted acoustic wave, which, in turn, leads to a poor image axial resolution. The most common pulse sequence used in tissue harmonic imaging is called pulse (or phase) inversion (Fig. 1.7). Pulse inversion tissue harmonic imaging (PI) is a two-pulse sequence, which consists of subtracting rather than filtering out the fundamental component of backscattered echoes. It enables the use of shorter pulses, i.e. broader transmit and receive bandwidths, which, in turn, leads to improved image axial resolution.
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Fig. 1.7: A type of tissue harmonic imaging.
Pulse inversion (PI) harmonic imaging signal processing. Pulse inversion cancels fundamental (linear) echoes while preserving harmonic (nonlinear) echoes. In PI, the second transmitted pulse is an exact replica of the first transmitted pulse, except for a 180°-phase shift. This is illustrated here with a contrast bubble as a reflector of ultrasound. However, this principle would apply to any reflector in the ultrasound beam.
In human tissues, the transmitted ultrasound waves are transformed by interaction with the tissues. The high-pressure components of sound waves move faster through human tissues and these waves then generate higher- frequency waves (harmonics). By listening in on the second harmonic component with the technique described in Figure 1.7, artifacts are reduced. Many artifacts in musculoskeletal imaging relate to interactions of the sound beam with septa in the subcutaneous lipomatous tissues. Harmonic generation only becomes a factor when sound has traveled a distance through tissues. Artifacts of the hypodermis in obese patients are therefore filtered. Harmonic ultrasound beams are also narrower than the beam that exits the transducer footprint and spatial resolution is improved in harmonic imaging (Fig. 1.8).
 
TRANSMISSION ULTRASOUND
Recent developments in equipment and image-processing techniques have made the production of high-quality images using transmission ultrasound possible in much the same way that computed tomography images are produced. Two transducers are utilized, positioned to face each other.
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Fig. 1.8: Tissue harmonic imaging.
This split screen image shows an almost identical transverse section through the same diseased posterior tibial tendon (PTT) at the level of the medial malleolus. The regular grayscale image is displayed on the left while the image obtained with tissue harmonics shows at the right. Notice how much easier it is to visualize the partial-thickness tear (arrow) in the PTT with tissue harmonics. The new technique clearly shows extension of the PTT tear to the surface; this significant finding may change a conservative approach into surgical treatment. Tissue harmonics improves the visualization of the surfaces of a tear and it clears fluid from reverberation artifact—two important features in musculoskeletal ultrasound.
One serves as the acoustic source, and the other as the receiver. The most recent examples have demonstrated exquisite contrast resolution and anatomical detail. Imaging is currently limited to the extremities due to the requirement of a water path. It is hoped that further development will continue, resulting in a clinical transmission ultrasound imaging system. Transmission ultrasound has proven valuable for bone densitometry. A clinical unit is now available for measurement of bone density, using the calcaneus as the standard for evaluation.
 
CONTRAST-ENHANCED ULTRASOUND
Ultrasound contrast agents are microbubbles (their diameter is typically on the order of 3 μm) filled with a low-solubility, high-molecular-weight gas (e.g. perfluorocarbon), and encapsulated in a stabilizing shell. When exposed to ultrasound, microbubbles behave in a highly nonlinear fashion, i.e. they respond differently to different portions of the incident acoustic wave (Figs. 1.9A to C). Positive acoustic pressure compresses a microbubble, while negative pressure dilates it.
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Figs. 1.9A to C: Contrast-enhanced ultrasound.
(A) Incident acoustic wave. (B) Nonlinear microbubble echoes. (C) Frequency spectrum of microbubble echoes.
When a microbubble expands while exposed to negative pressures, its radius can increase by as much as several hundred percent. On the other hand, as a microbubble contracts in response to positive pressures, the decrease in its radius is limited due to the gas inside the microbubble rapidly stiffening, making the microbubble less compressible. As a result, when a microbubble is exposed to ultrasound, its radius oscillates in a nonsymmetric fashion: the microbubble behaves nonlinearly. Nonsymmetric microbubble oscillations lead to backscattered acoustic waves, which are not sinusoidal in shape anymore, but are rather nonsymmetric. In the Fourier domain, this translates into a frequency spectrum with a fundamental (similar to that of incident acoustic wave) and harmonic components.
Low mechanical index harmonic imaging is the preferred dedicated contrast imaging mode in radiological applications of contrast-enhanced ultrasound. This is due to the fact that typical mechanical indices used for conventional B-mode imaging are high enough to destroy most microbubbles. Low mechanical index harmonic imaging relies on those harmonic components to selectively image microbubbles over surrounding tissue. In Figure 1.9, the “frequency spectrum of microbubble echoes” is shown. The first major hump is the fundamental frequency component and is similar to that observed in the incident acoustic wave's frequency spectrum (not shown here). Higher-frequency humps are the second, third, and fourth harmonic frequency components, and their presence confirms the nonlinear nature of microbubble echoes (Fig. 1.9B).
In musculoskeletal imaging, contrast-enhanced ultrasound has been utilized to show in-growth of rotator cuff tendons after surgery (Figs. 1.10A to C). Pannus and the synovial component of an inflammatory mass can also be distinguished from simple fluid in inflammatory joint swelling (Figs. 1.11A and B).
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Figs. 1.10A to C: Contrast-enhanced ultrasound examination after rotator cuff repair. Courtesy: Ronald Adler, PhD, MD, Hospital for Special Surgery, New York, NY, USA.
This 69-year-old woman underwent supraspinatus tendon repair 3 months ago. Long axis supraspinatus gray-scale image (A) and Power Doppler image (B) show an inhomogeneous tendon reconstruction with an anchor defect in the greater tuberosity (arrow). There is very little color flow on the Power Doppler image. After contrast administration (C), with the region of interest (open arrow) placed at the tendon footprint on the greater tuberosity, the contrast intensity curve shows marked contrast enhancement after intravenous contrast injection. Compare this to the minimal flow observed in panel (B). The mean echo amplitude registers in decibels per millimeter squared on the Y-axis and time elapsed since the injection on the X-axis. The harmonic image at peak enhancement shows on top left and the fundamental image on top right.
 
ULTRASOUND-BASED ELASTOGRAPHY
As discussed earlier in this chapter, ultrasound B-mode imaging relies on acoustic impedance mismatch at the interface between different tissue types to differentiate them. Being a “fluid-like” solid, tissue may also be characterized by its elastic properties. It is possible to take advantage of the variations in the elastic properties of tissue to detect pathology. Manual palpation of tissue has indeed been used for centuries as a qualitative measurement of the elastic properties of tissue. Tumors (benign or malignant) near the body surface feel “stiffer” than surrounding tissue. The external force (stress) applied via palpation will likely not compress a tumor as much as surrounding healthy tissue, thereby providing an indication of tumor presence. The difference in deformation (strain) provides the elasticity information that a physician uses, e.g. when searching for lumps in a breast clinical examination or during surgery when other organs may be palpated directly (e.g. liver and lymph nodes).
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Figs. 1.11A and B: Contrast-enhanced finger ultrasound. Courtesy: Andrea S Klauser MD, Medical University Innsbruck, Department of Radiology, Austria.
A proximal interphalangeal joint (PIP) in rheumatoid arthritis examined in the long axis of the digit with harmonic imaging before (A) and after (B) SonoVue contrast administration. Panel (B) shows marked enhancement—white dots (open arrows)—in the synovial proliferation (oval). In this patient, contrast allows us to distinguish solid articular tissue from fluid.
Over the past two decades, a number of academic and industrial research groups have developed different techniques to estimate tissue elastic properties with ultrasound. While there are a variety of implementation pathways chosen for prototyping and development of commercial solutions, all ultrasound-based elastography techniques need to deform tissue, track tissue motion, and reconstruct maps of tissue elastic properties. Ultrasound-based elastography techniques can further be categorized on the basis of how the three aforementioned steps are implemented.
Applications of elastography in the musculoskeletal system will likely focus on the detection of the solid portions of tumors to improve the accuracy of biopsies and on a more accurate detection of fluid in joints to facilitate aspirations. Research early on has also indicated that tendon disease can be more accurately quantified by elastography as “soft” areas in tendons. This will be discussed in the chapters on tendons, elbow, and foot.
 
CONCLUSION
The purpose of this chapter is not to provide a thorough review of ultrasound physics to enable readers to pass their specialty board examination. Its goal is to present technical information about ultrasound imaging that will allow readers to improve the diagnostic quality of images produced daily in clinical practice. Its success can be determined only by the individual reader, but if you have read it to this point, perhaps that too can be considered a success.
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