Qura Manual of Fetal Imaging

QURA MANUAL OF FETAL IMAGING

QURA MANUAL OF FETAL IMAGING
Author
Dr. Somya Dwivedi
Co- Authors
Dr. Abhishek Dwivedi Dr. Aman Gupta
Qura Research and Diagnostic Centre
Plot No 4, Kolar Rd, Janki Nagar, Chuna Bhatti, Bhopal, Madhya Pradesh 462016

Lakecity Publishing
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© Lakecity Publishing 2023
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First Published in 2023
Designed By : Lakecity Publishing

Preface
Fetal Medicine and imaging is an ever-growing branch with vast knowledge and potential. Obstetric imaging is practised all over the world and is considered as one of the most essential and common scans done in day to day life, improving the perinatal outcome, prognosis and postnatal life expectancy. In this view, specially in Indian scenario where scans are done both in urban normal and high risk clinics as well as rural on sometimes low resolution machines, it is very crucial to emphasise on the correct guidelines to diagnose basic and complex anomalies antenatally.
Evaluation of a fetus is not only an important aspect for the parents to be but also for the fetus who is going to be born. Needless to say, many abnormalities are missed by just taking improper views and lesser known techniques.
In this manual we impart not only knowledge about the important topics of fetal imaging, but also we include the standard guidelines given by renowned accredited societies which are evolving over the period of time that should be followed worldwide for scanning of fetuses, in all trimesters. In this manual, emphasis has been laid on the techniques for taking proper planes and protocols to clear the concepts for our basic routines scans to doing high-risk obstetric scans as well. Pattern recognition and anatomical landmarks are stressed upon which enhances the detection rate of antenatal abnormalities.
In the final section the guidelines provide us a baseline which is followed all over the world, to scan the fetus with the best perinatal outcome given by a sonologist, which we know will keep evolving like the editions of our book in future.
The reader will not only gain knowledge about the embryology, ultrasound features, counselling but also the management of the complicated pregnancies. We encourage the reader to further refer and read the complexities and refer the patient to appropriate specialist at high risk centers if need be. This manual will help the reader to perform and document features, images systematically in a standardised fashion in all types of pregnancies. It will also provide a verifiable and a standardised method which can be implemented in day to-day practice for improving the perinatal mortality, outcome and morbidity.

Acknowledgement
We are immensely thankful to all the contributors Dr. Abhishek Dwivedi and Dr. Aman Gupta for their continuous support and guidance to give the first edition its final look and share there vast knowledge in the given chapters. I would also like to thank my family specially my parents, my husband and daughters for providing unending support and encouragement to share my work and empower the younger and enthusiastic generation about obstetric imaging and also advise to work towards their goals and passion, like the love for imaging the beauty of life that is the ‘FETUS’. My book won’t be complete if I don’t thank my biggest supporter and stimuler my brother Mr. Abhishek Mohan Gupta, who provided me the platform and the thought process to write this book and give its beautiful shape, which you have in your hands.We would like to express gratitude to Lakecity Publishing for there untiring efforts, support and professional help in formulating and editing the book to give it’s final form.
Last but not the least, I would like to thank our dear readers, with sincere hope that your love for obstetric imaging will increase and you will have more clearity towards this fascinating branch.
Regards,
The Author

About the Authors
DR. SOMYA DWIVEDI
Present Designation
Director & chief radiologist of QURA DIAGNOSTICS ,BHOPAL; MBBS,MD Radiodiagnosis;Fellow in fetal imaging (MEDISCANS), FMF Certified for NT, Pre-eclampsia ,cervical assessment and fetal abnormalities,IOTA Certified for women imaging,I.S.U.O.G. Certified for Advanced obstetric Imaging
Present Affiliation
National Council member IRIA Yuwa, Social and Utsav wing, Member IRIA,IFUM,SOCIETY OF FETAL MEDICINE,ISUOG,I.O.T.A. MEMBER AFFLIATION, FMF-INDIA, Indian Medical association,Teaching faculty with ICRI & IRIA
Areas of Interest: Fetal & Women Imaging And Interventions
Major Achievements (Honours) (Awards) (Publications):
• 5 Gold Medals for paper presentations in USCON,IRIA NATIONAL AND STATE PLATFORMS and IMA
• Multiple paper publications in Indexed national journals
• Qura Diagnostics has received ‘Best diagnostic and imaging centre of
Madhya Pradesh’ for three consequtive years
• She runs a social initiative “INNER SHE” started in 2018 January to empower
and educate women and girls about their health and well being. Till now 8000 women have been educated for free in this mission.

Past work profile:
DR. ABHISHEK DWIVEDI
Assistant Professor:- M.Y.H.& M.G.M.M.C.,Indore from Aug.1999 to June.2005
Associate Professor: M.Y.H. & M.G.M.M.C.,Indore from June.2005 to Sept.2009
Professor & Head: S.A.I.M.S. Medical College & P.G.Institute,Indore from Oct.2009 to March.2020.
Academic Post held:-
• Secretary, IRIA Indore Br.2008-09.
• Member, Scientific Committee, 66th IRIA National conference 2013, Indore. • Chairman, Scientific Paper Committee, 66th IRIA National conference 2013,
Indore
• Chairman, Scientific Committee, 29th Annual M.P. State IRIA Conference. • Joint Organizing Secretary, USCON 2018,Indore.
• Academic Secretary-ISUMB-2017-2019
• Treasurer ISUMB-2019-2021.
• Member EC-IFUMB-2020.
• Vice President, IRIA Indore Br.2021-22
• Organizing Chairperson, 36th Annual M.P. State IRIA Conference.
Paper Publications:- More than 28 publications in National and International

DR. AMAN GUPTA
Profesional Qualifications:-
M.B.B.S.:-M.G.M. Medical College, Indore. 4th position in D.A.V.V. with Gold Medal in Surgery.
M.D.(Radiodiagnosis):-M.G.M.Medical College, Indore.
AWARDS:-
• Sonar Films Bombay Gold Medal for 1st position in
• Surgery in MBBS by DAVV, Indore.
• Gold Medal and Certificate of Merit by MGM Medical college, Indore for 1st
position in Surgery.
• Dr. Khanbhadur Ahmadbux Award by IMA, Indore branch. • Dr. Anil Gupta Memorial Award by IMA, Indore branch.
Present work profile:
• Professor & Head, Sevakunj Hospital and LNCT Medical College, Indore • Consultant Radiologist at St. Francis Hosp. & Research Centre, Indore
• Consultant Radiologist-Agrawal Diagnostic Center, Indore

CONTENTS
1. Diagnostic Ultrasound-Basic Physics and Knobology
2. DopplerUS
3. Basics of Genetics
4. PEDEGREE Module
5. First Trimester Ultrasound
6. Early OBS Complication Cases
1 31 65 91
99 123
7. USG of Obstetric Emergencies 157
8. Unusual Ectopic Pregnancies
9. Management of Ectopic Pregnancy
10. CAKUT
11. Screening and Follow Up in Pre-Eclampsia 221
12. Opthalmic Artery Doppler in Preeclampsia
13. Duodenal Atresia
14. Gestational Diabetes Mellitus
15. IOTA Adnex Model 305
16. Obstetric Doppler
17. Rare Cases
18. Guidelines 353
177
245 255 267
201 211
331 345

Diagnostic Ultrasound- 1 Basic Physics and Knobology
Author: Dr. Aman Gupta
DIAGNOSTIC ULTRASOUND
Many of the differences between x-ray and ultrasound imaging stem from the nature of the energy being used, and how it propagates. Whereas x-rays can be considered either as discrete photons or as waves, ultrasound is solely a wave phenomenon. It behaves like light waves in some ways, although it carries a different form of energy. For example, it travels in straight lines unless it is caused by some material to change its direction; it undergoes partial reflection at the boundary between two media, and is subject to refraction, diffraction, scattering and absorption, all of which are relevant to diagnostic imaging. However, sound waves are fundamentally different from light, x-rays and all other electromagnetic waves, in that they cannot travel through a vacuum. They exist only as vibrations of the particles of a medium and their propagation depends on the density and stiffness of the medium. The partial reflection and scattering, which are crucial to imaging, occur whenever the travelling ultrasonic wave meets a change in the density or stiffness of tissue. The result in many instances is that an echo is sent back to the transducer, and diagnostic ultrasound relies on these echoes to form its image. The pulsed ultrasound waves emitted from the transducer travel at a fairly constant speed in their given direction, so that the range and direction of the features that produce the echoes can be measured and plotted. The image is therefore an ‘echo-map’. The reflected ultrasound waves offer another diagnostic possibility: if scattered from moving blood cells within the body, they will undergo Doppler shift (a change in frequency). Measurement of the shift allows the blood flow in specific parts of the body to be examined, thus providing functional information.
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One of the principal attractions of ultrasonic imaging is its freedom from the harmful effects associated with ionizing radiations. Its growth to its present dominant role is also due to it being relatively cheap and mobile, to its ‘real- time’ nature, and to an ever-increasing appreciation that it can provide unique diagnostic information about the structure, properties and mechanical behavior of tissue and blood flow. The major limitations to ultrasonic investigations are the barriers presented by gas and bone, and the dependence on a relatively high degree of operator skill and experience. In addition, mapping distortions and artefacts occur, so some understanding of the basic physics of sound propagation and of the techniques used in scanning equipment is necessary if high quality scans are to be produced and their limitations appreciated.
Successful medical applications of ultrasound began in the late 1940s and early 1950s. Since then the application of ultrasound to medical diagnosis has seen continuous development and growth over several decades. Early, primitive display modes, such as A-mode and static B-mode, borrowed from metallurgical testing and radar technologies of the time, have given way to high-performance, real- time imaging. Improvements in technology have been followed by widespread acceptance and use of ultrasound in medical diagnosis. Applications have progressed from simple measurements of anatomical dimensions, to detailed screening for fetal abnormalities, detection of subtle changes in tissue texture and detailed study of blood flow in arteries, visualize moving structures in 3D, and make measurements related to the stiffness of tissues.Modern ultrasound
equipment uses a pulse-echo approach with a brightness-mode (B-mode) display. Fundamental aspects of the B-mode imaging process include basic ultrasound physics, interactions of ultrasound with tissue, ultrasound pulse formation, scanning the ultrasound beam, and echo detection and signal processing.
BASIC ULTRASOUND PHYSICS
Ultrasound is a high-frequency sound wave with frequencies above the upper auditory limit of 20 kHz, which can be used to form images of internal body organs. Ultrasound travels through the tissues of the body in a way which makes it possible to form useful images however, the image-formation process includes some approximations, which give rise to imperfections and limitations in the imaging system. In order to be able to use diagnostic ultrasound systems effectively and to be able to distinguish imperfections in the image from genuine diagnostic information, the user must have an appreciation of the basic principles of ultrasound propagation in tissue.
The sound waves used to form medical images are longitudinal waves, which propagate (travel) through a physical medium (usually tissue or liquid). Here, the particles of the medium oscillate backwards and forwards along the direction of propagation of the wave (Fig.-1). Where particles in adjacent regions have moved towards each other, a region of compression (increased pressure) results,
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but where particles have moved apart, a region of rarefaction (reduced pressure) results. Each repetition of this back-and-forth motion is called a cycle, and each cycle produces a new wave. The wave length is the distance between two bands of compression, or rarefaction, and is represented by the symbol ‘λ’ and has units of meters or millimeters.
Fig.1:-Propogation of a longitudinal wave, particle motion is aligned with the direction of travel, resulting in bands of high and low pressure
Once the sound wave has been generated, it continues in its original direction until it is either reflected, refracted, or absorbed. The frequency of the wave is the number of oscillations or wave crests passing a stationary observer per second and is determined by the source of the sound wave (Fig.2). Frequency is normally given the symbol ‘f ’ and has units of hertz (1 Hz = 1cycle per second), in honor of famous German physicist Heinrich R. Hertz; sound waves with frequencies in the approximate range 20 Hz to 20 kHz can be detected by the human ear. Sound waves with frequencies above approximately 20 kHz cannot be heard and are referred to as ultrasound waves.
Fig.2:- The frequency ‘f ’ of a wave is the number of wave crests passing a given point per second. The wavelength ‘λ’ is the distance between wave crests.
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SPEED OF SOUND
The speed or velocity of propagation of a sound wave for body tissues in the medical ultrasound range, is independent of frequency, and depends primarily on the physical makeup of the medium it is travelling in. In gases the speed of sound is relatively low in relation to values in liquids, which in turn tend to be lower than values in solids. The material properties which determine the speed of sound are (1) density and (2) stiffness or compressibility. Density is a measure of the weight of a material for a given volume and is normally given the symbol
‘ρ’. Stiffness is a measure of how well a material resists being deformed when it is squeezed. This is given by the pressure required to change its thickness by a given fraction. Stiffness is usually denoted by the symbol ‘k’. Dense materials tend to be composed of massive molecules, and these molecules have a great deal of inertia. They are difficult to move or to stop once they are moving. Because the propagation of sound involves the rhythmic starting and stopping of particulate motion, dense material made up of large molecules (i.e., large in mass) transmit sound at less velocity. The velocity of sound is inversely related to the compressibility of the conducting material; that is, the less compressible (more stiffer) a material, the more rapidly it transmits sound. Hence, low density and high stiffness lead to high speed of sound whereas high density and low stiffness lead to low speed of sound. Mathematically this is expressed in the following equation:
Speed (velocity) of sound = v = √k/ρ.
Although gases have low density, they have very low stiffness (high compressibility), leading to relatively low speed of sound compared to liquids and solids. Sound travels slowest in gases, at intermediate velocity in liquids, and most rapidly in solids. All body tissues, except bone, behave like liquids, and therefore they all transmit sound at about the same velocity. A velocity of 1540 m per second is used as an average for body tissue. Table. -1 shows typical values for the speed of sound in various materials, including a number of different kinds of human tissue.
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Frequencies and wavelengths used in diagnosis
The ultrasound frequencies used most commonly in medical diagnosis are in the range 2–15 MHz, although frequencies up to 40 MHz may be used in special applications and in research. The wavelengths in tissue which result from these frequencies can be calculated using the equation: which relates wavelength ‘λ’ to the frequency ‘f ’ and velocity ‘v’ of a wave:
λ = v/f
Assuming the average speed of sound in soft tissues of 1540 ms−1, values of λ at diagnostic frequencies are within the range 0.1–1 mm (Table 2). The wavelength of the ultrasound wave has an important influence on the ability of the imaging system to resolve fine anatomical detail. Short wavelengths give rise to improved resolution, i.e., the ability to show closely spaced targets separately in the image.
INTERACTIONS BETWEEN ULTRASOUND AND MATTER
The types of interactions between sound and matter are similar to those of light, and include the following: (1) reflection, (2) refraction, and (3) absorption.
Reflection
When a sound wave travelling through one medium meets an interface with a second medium of different acoustic impedance, some of the wave is transmitted into the second medium and some is reflected back into the first medium. The amplitudes of the transmitted and reflected waves depend on the change in acoustic impedance. Acoustic impedance is a fundamental property of matter. The impedance of a material is the product of its density and the velocity of sound in the material:
z =p ×v
where, z = acoustic impedance (Rayls) p = density (g/cm3)
v = velocity of sound (cm/sec)
Acoustic impedance ‘z’ has units of kg m−2 s−1 , but the term ‘rayl’ (after Lord Rayleigh) is often used to express this unit. Table.3 gives values of z for some common types of human tissue, air and water. Values for z in most human soft
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tissues are very similar. For air, which has low density and low stiffness, z is very small. For bone, which has high density and high stiffness, z is much higher.
The ratio of reflected to incident pressure is commonly referred to as the amplitude reflection coefficient ‘RA’. It is very important to ultrasound image formation as it determines the amplitude of echoes produced at boundaries between different types of tissue. For most soft tissue to soft tissue interfaces, the amplitude reflection coefficient is less than 0.01 (1%). This is another important characteristic for ultrasound imaging as it means that most of the pulse energy at soft tissue interfaces is transmitted on to produce further echoes at deeper interfaces. The amplitude reflection coefficient at a tissue–fat interface is about 10% due to the low speed of sound in fat. At an interface between soft tissue and air, as might be encountered within the lungs or gas pockets in the gut, the reflection coefficient is 0.999 (99.9%), so no further useful echoes can be obtained from beyond such an interface. For this reason, it is important to exclude air from between the ultrasound source (the transducer) and the patient’s skin to ensure effective transmission of ultrasound. At an interface between soft tissue and bone, the amplitude reflection coefficient is approximately 0.5 (50%), making it difficult also to obtain echoes from beyond structures such as ribs. Reflection at strongly reflecting interfaces can lead to a number of image artefacts.
The amount of reflection is determined by the angle of incidence between the sound beam and the reflecting surface. The higher the angle of incidence (i.e., the closer it is to a right angle), the less the amount of reflected sound as the angles of incidence and reflection are equal. In medical ultrasound, in which the same transducer both transmits and receives ultrasound, almost no reflected sound will be detected if the ultrasound strikes the patient’s surface at an angle of more than 3° from perpendicular.
Refraction
When sound passes from one medium to another its frequency remains constant but its wavelength changes to accommodate a new velocity in the second medium.
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The change in wavelength necessitates a change in direction. This bending of waves as they pass from one medium to another is called refraction. When the wave crosses an interface where the velocity (speed) of sound increases, the angle to the normal also increases. Conversely, when the wave experiences a reduction in the speed of sound as it crosses the interface, the angle to the normal also decreases. The relationship between the angles θi , θt , v1and v2 is described by Snell’s law:
(sin θi )/(sin θt) = v1/v2
where, sin θi = incidence angle (see Fig.3) sin θt = transmitted angle
v1 = velocity of sound for incident medium
v2 = velocity of sound for transmitting medium.
Fig.3:- Refraction and reflection of sound waves
Snell’s law shows that the angles of the incident and transmitted waves are the same when the speeds of sound in the two media are the same. The change in the direction of propagation of the wave as it crosses the boundary increases with increasing change in the speed of sound, i.e., larger changes in the speed of sound give rise to stronger refraction effects. Refraction can cause artifacts. Refraction artifacts cause spatial distortion (real structures are imaged in the wrong location) and loss of resolution in the image.
Attenuation
When an ultrasound wave propagates through soft tissue, the energy associated with the wave is gradually lost so that its intensity reduces with distance travelled, an effect known as attenuation. The pattern is that of an exponential decay curve, where the rate of decrease is rapid at first but becomes more gradual as the intensity reduces. The rate at which the intensity of the wave is attenuated, in dB cm−1, is referred to as the attenuation coefficient. An ultrasound wave
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can be attenuated by several mechanisms as it travels through tissue. The most important mechanism is absorption, in which ultrasound energy is converted into heat. In addition the intensity in the beam is reduced due to scattering of ultrasound, reflection and divergence or spreading of the beam with distance.
Absorption
Absorption of ultrasound in fluids is a result of frictional forces that oppose the motion of the particles in the medium. The energy removed from the ultrasound beam is converted into heat. To be precise, the term “absorption” refers to the conversion of ultrasonic to thermal energy. Three factors determine the amount of absorption: (1) the frequency of the sound, (2) the viscosity of the conducting medium, and (3) the “relaxation time” of the medium.
Particle freedom decreases and internal friction increases with increasing viscosity. This internal friction absorbs the beam, or decreases its intensity, by converting sound into heat. In liquids, such as urine, amniotic fluid and blood which have low viscosity, very little absorption takes place. In soft tissues with high collagen content such as tendons and cartilage viscosity is higher and a medium amount of absorption occurs, whereas bone shows high absorption of ultrasound. The relaxation time is the time that it takes for a molecule to return to its original position after it has been displaced. It refers to the resilience of a material. Two substances with the same viscosity may have different relaxation times. The relaxation time is a constant for any particular material. When a molecule with a short relaxation time is pushed by a longitudinal compression wave, the molecule
has time to return to its resting state before the next compression wave arrives. A molecule with a longer relaxation time may not be able to return completely before a second compression wave arrives. When this happens, the compression wave is moving in one direction and the molecule in the opposite direction. More energy is required to reverse the direction of the molecule than was needed to
move it originally. The additional energy is converted to heat.
Attenuation of ultrasound by biological tissues increases with frequency. The attenuation coefficient of most tissues, increases approximately linearly with frequency, i.e., doubling the frequency will increase the attenuation by approximately a factor of 2. Hence, for most tissues, it is possible to measure ultrasound attenuation in dB cm−1MHz−1.This allows the total attenuation of an ultrasound pulse to be calculated easily from the frequency and the distance travelled. For example, if a tissue attenuates by 0.7 dBcm−1MHz−1, then a 5 MHz ultrasound wave, after travelling a distance of 10 cm, will be attenuated by 5 MHz × 10 cm × 0.7 dB cm−1 MHz−1 = 35 dB. Table.4 summarizes measured values of attenuation for some human tissues expressed in units of dBcm−1MHz−1.Values are typically in the range 0.3–0.6 dBcm−1 MHz−1 for soft tissues but much lower for water (and watery body fluids). Attenuation in bone is very high and does not increase linearly with frequency as in the case of soft tissues. When attenuation is
large, the echoes returned from deeper targets may be too weak to detect. Hence, for imaging large or deep organs, a low frequency (3–5 MHz) must be used. High
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frequencies (10–15 MHz) can only be used to image relatively small, superficial targets (e.g., thyroid) as they are attenuated more rapidly. The short wavelengths associated with high-frequency ultrasound leads to improved resolution of image detail. The operator must choose the optimum frequency for each particular application. This is a compromise that ensures that the best image resolution is obtained, while allowing echoes to be received from the required depth.
BASIC PRINCIPLES OF ULTRASOUND IMAGE FORMATION
A B-mode image is a cross-sectional image representing tissues and organ boundaries within the body. It is constructed from echoes, which are generated by reflection of ultrasound waves at tissue boundaries, and scattering from small irregularities within tissues. Each echo is displayed at a point in the image, which corresponds to the relative position of its origin within the body cross section, resulting in a scaled map of echo-producing features. The brightness of the image at each point is related to the strength or amplitude of the echo, giving rise to the term B-mode (brightness mode). Usually, the B-mode image bears a close resemblance to the anatomy, which might be seen by eye, if the body could be cut through in the same plane. Abnormal anatomical boundaries and alterations in the scattering behavior of tissues can be used to indicate pathology. To form a B-mode image, a source of ultrasound, the transducer, is placed in contact with the skin and short bursts or pulses of ultrasound are sent into the patient. These are directed along narrow beam-shaped paths. As the pulses travel into the tissues of the body, they are reflected and scattered, generating echoes, some of which travel back to the transducer, where they are detected. These echoes are used to form the image. To display each echo in a position corresponding to that of the interface or feature (known as a target) that caused it, the B-mode system needs two pieces of information. These are (1) the range (distance) of the target from the transducer-echo ranging- and (2) the direction of the target from the active part of the transducer, i.e., the position and orientation of the ultrasound beam.
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Echo ranging
The range of the target from the transducer is measured using the pulse–echo principle. After transmission of the short pulse, the transducer operates in receive mode, effectively listening for echoes. These begin to return immediately from targets close to the transducer, followed by echoes from greater and greater depths, in a continuous series, to the maximum depth of interest. This is known as the pulse–echo sequence. The 2D B-mode image is formed from a large number of B-mode lines, where each line in the image is produced by a pulse–echo sequence. As the transducer transmits the pulse, a display spot begins to travel down the screen from a point corresponding to the position of the transducer, in a direction corresponding to the path of the pulse (the ultrasound beam). Echoes from targets near the transducer return first and increase the brightness of the spot. Further echoes, from increasing depths, return at increasing times after transmission as the spot travels down the screen. Hence, the distance down the display at which each echo is displayed is related to its depth below the transducer. The rate at which the display spot travels down the screen determines the scale of the image. The different reflectivity’s of various structures encountered by the pulse cause a corresponding variation of the detected echo strength. The detected echo signals are processed and translated into luminance, resulting in a
“brightness-mode” or B-mode image display. In B-mode images, more reflective structures appear brighter than less reflective structures. A complete image is obtained by repeating this pulse-echo cycle.
Let us consider a linear array probe (Fig.4), during the first pulse–echo sequence, an image line is formed, say on the left of the display. The active area of the transducer, and hence the beam, is then moved along the array to the adjacent beam position. Here a new pulse–echo sequence produces a new image line of echoes, with a position on the display corresponding to that of the new beam. The beam is progressively stepped along the array with a new pulse–echo sequence generating a new image line at each position. One complete sweep may take perhaps 1/30th of a second. This would mean that 30 complete images could be formed in 1s, allowing real-time display of the B-mode image. That is, the image is displayed with negligible delay as the information is acquired.
Fig.4:- Formation of a 2D B-mode image. The image is built up line by line as the beam is stepped along the transducer array.
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B-mode formats
Various scan formats are used for different applications for B-mode imaging. In a linear transducer array, a large number of small transducer elements arranged in a straight line. The ultrasound beams, and hence the B-mode lines, are all perpendicular to the line of transducer elements, and hence parallel to each other. The resulting rectangular field of view is useful in applications, where there is a need to image superficial areas of the body at the same time as organs at a deeper level. A curvilinear transducer gives a wide field of view near the transducer and an even wider field at deeper levels. This is also achieved by the trapezoidal field of view. Curvilinear and trapezoidal fields of view are widely used in obstetric scanning to allow imaging of more superficial targets, such as the placenta, while giving the greatest coverage at the depth of the baby. The sector field of view is preferred for imaging of the heart, where access is normally through a narrow acoustic window between the ribs. In the sector format, all the B-mode lines are close together near the transducer and pass through the narrow gap, but diverge after that to give a wide field of view at the depth of the heart. Transducers designed to be used internally, such as intravascular or rectal probes, may use the radial format as well as sector and linear fields of view. This format may be obtained by rotating a single element transducer on the end of a catheter or rigid tube, hence, the B-mode lines all radiate out from the center of the field of view (Fig.5).
Fig.5:- Scan line arrangements for the most common B-mode formats. These are (a) linear, (b)curvilinear, (c) trapezoidal, (d) sector and (e) radial.
Transducer
The transducer is the device that converts electrical pulses into ultrasonic pulses that can be transmitted into tissues, and, conversely, ultrasonic echo pulses reflected back from the tissues into electrical signals. The simplest way to interrogate all the scan lines that make up a B-mode image is to physically move the transducer so that the beam is swept through the tissues as the pulse–echo cycle is repeated. This was the original method used but it has been superseded by electronic scanning methods which use multi-element array transducers with no
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moving parts. Array transducers allow the beam to be moved instantly between positions, and give the additional benefit of allowing the shape and size of the beam to be changed to suit the needs of each examination. The beam-former is the part of the scanner that determines the shape, size and position of the interrogating beams by controlling electrical signals to and from the transducer array elements. In transmission, it generates the electrical signals that drive each individual transducer element, and in reception it combines the individual echo sequences received by all the transducer elements into a single echo sequence. The echo sequence produced by the beam-former for each scan line is then amplified and processed in various ways before being used in the formation of the B-mode image.
Common features of all transducers and transducer elements
All transducers or transducer elements have the same basic components: a piezoelectric plate, a matching layer and a backing layer (Figure.6). Usually, there is also a lens, in array transducers it is usual for one large lens to extend across all the transducer elements. The most important component is a thin piezo-electric crystal element located near the face of the transducer. The front and back faces of the crystal are coated with a thin conducting film to ensure good contact with the two electrodes that will supply the electric field used to strain the crystal. The surfaces of the crystal are plated with gold or silver electrodes. The outside electrode is grounded to protect the patient from electrical shock, and its outside surface is coated with a watertight electrical insulator. The inside electrode abuts against a thick backing block that absorbs sound waves transmitted back into the transducer. The housing is usually a strong plastic. An acoustic insulator of rubber or cork prevents the sound from passing into the housing. The number, size, shape and arrangement of transducer elements vary according to the transducer type and application.
Fig.6:- (A)Schematic diagram (B) Cut section showing basic component elements in an imaging ultrasound transducer.
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1. DIAGNOSTIC ULTRASOUND- BASICS OF PHYSICS AND KNOBOLOGY
From the outside, a linear-array transducer appears as a moulded block designed to fit comfortably into the operator’s hand with a rubber lens along the face that makes contact with the patient. Behind the lens is a matching layer, and behind this is a linear array of typically 128 regularly spaced, narrow, rectangular transducer elements, separated by narrow barriers (kerfs), made of an inert material, usually a polymer or epoxy. Some linear-array transducers have as many as 256 elements (these numbers are chosen, since they are ‘round’ numbers in binary terms, and hence more convenient for digital control and processing). The width of each array element is typically about 1.3 wavelengths, being a compromise that gives a reasonably wide array. Since all transducer dimensions are proportional to wavelength, high-frequency transducers are smaller than low-frequency ones. Thus, assuming 128 elements, a 3MHz (λ = 0.5 mm) transducer might typically have a lens face measuring about 83mm (128 × 1.3 × 0.5) by 15mm, with each element being about 0.65mm wide whereas a 7.5 MHz (λ = 0.2 mm) transducer will have a lens face measuring about 33mm by 6mm, with each element being about 0.25mm wide. The front electrodes of all the elements are usually connected together, so they share a common electrical lead. However, the rear electrode of each element is provided with a separate electrical lead (Figure.7), allowing the signals to and from each element to be individually processed by the beam- former.
Fig.7:-Section through a linear-array transducer.
Piezoelectric plate
The actual sound-generating and detecting component is a thin piezoelectric plate. Certain materials expand or contract when a positive or negative electrical voltage is applied across them and, conversely, generate positive or negative voltages when compressed or stretched by an external force. This is called the “piezoelectric effect”, first described by Pierre and Jacques Curie in 1880. Piezoelectric materials are made up of innumerable dipoles arranged in a geometric pattern. An electric dipole is a distorted molecule that appears to have a positive charge on one end and a negative charge on the other. The positive and negative ends are arranged so that an electric field will cause them to realign, thus changing the dimensions of the crystal. Some piezoelectric materials, such as quartz, occur naturally but
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the piezoelectric material normally used for transducers in medical imaging is a synthetic ceramic material:lead zirconate titanate (PZT).Various types of PZT are available and are chosen according to the properties required, such as high sensitivity, or the ability to cope with large acoustic powers. The thin plate of PZT is coated on both sides with conductive paint, forming electrodes, to which electric connections are bonded. In order to transmit an ultrasonic pulse, a corresponding oscillating voltage is applied to the electrodes, making the PZT element expand and contract (“strain”) at the required frequency. The term “strain” refers to deformity of the crystal caused when a voltage is applied to the crystal. The back and forth movements of the front face send an ultrasonic wave into the patient’s tissues. In reception, the pressure variations of returning echoes cause the PZT plate to contract and expand. The voltage generated at the electrodes is directly proportional to the pressure variations, giving an electrical version (the echo signal) of the ultrasonic echo (Fig.8).
The PZT slab vibrates most strongly at the frequency for which its thickness is half a wavelength, giving rise to the term ‘half-wave resonance’. Resonance occurs because an ultrasound wave propagating across the thickness of the PZT plate is partially reflected at the front and back faces, and thus continues to travel back and forth (reverberate) within the PZT plate. As the plate thickness is equal to half the wavelength, the wave travels a full wavelength in the PZT on each round trip. This means that the reflected wave arrives back in phase with the original wave, and adds constructively to it to produce a greater output. A PZT plate with a thickness equal to half a wavelength at the required centre frequency is therefore used, as this will resonate and produce a large output at this frequency.
Fig.8:- Piezoeletric elements vibrate to generate sound wave when applied with a voltage and conversely, generate a voltages when compressed or stretched by an external force.
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Backing layer
PZT has advantages as a transducer material in that it is efficient at converting electrical energy to mechanical energy, and vice versa, and is relatively easy
to machine or mould to any required shape or size. However, it has one significant disadvantage in that it has a characteristic acoustic impedance
that is about 20-times higher than that of the soft tissue. If the front face of
the PZT plate were to be in direct contact with the patient, a large fraction (approximately 80%) of the ultrasound wave’s power would be reflected at the PZT–tissue interface. This unwanted ringing can be much reduced by having a backing (damping) layer behind the PZT, made of a material with both a high characteristic acoustic impedance and the ability to absorb ultrasound. If the impedance of the backing layer were identical to that of PZT, all the sound energy would cross the boundary between the PZT and the backing layer, and none would be reflected back into the PZT. Once in the backing layer the sound would be completely absorbed and converted into heat. This would eliminate ringing, but would be at the expense of sensitivity, as some of the energy of the electrical driving pulse and also of the returning echo sound pulses would be wasted as heat in the backing layer. In modern practice, a backing layer with
an impedance somewhat lower than that of the PZT is used. This compromise impedance is chosen to give a useful reduction in reflection without lowering sensitivity too much. The remaining ringing is removed by using matching layers.
Matching layer(s)
Apart from the ringing problem, the fact that only about 20% of the wave’s power would be transmitted through the front PZT–patient interface means there is also a potential problem of poor sensitivity. In order to overcome this, at least one ‘impedance matching layer’ is bonded to the front face of the PZT.
A single matching layer can increase the transmission across the front face to 100%, provided that two important conditions are met.
(1) the matching layer should have a thickness equal to a quarter of a wavelength.
(2) it should have an impedance equal to √((Zpzt ×Zt)), where Zpzt is the impedance of PZT and Zt is the impedance of tissue.
The sound reverberates back and forth repeatedly within the matching layer, producing a series of overlapping waves, at the interface to the patient, that are in phase with each other and hence combine to give a large resultant wave. At the same time, another series of waves is sent back into the PZT, which together exactly cancel out the wave originally reflected at the PZT–matching layer interface (Figure 9).
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Fig.9:- Quarter-wave matching layer. Reverberations within the plate produce multiple transmissions into the patient that reinforce each other to give a large amplitude resultant pulse.
One hundred per cent transmission through the matching layer only occurs at one frequency for which its thickness is exactly one-quarter of a wavelength. It is normal to choose the matching layer thickness to be correct for the centre frequency of the pulse, as there is more energy at this frequency than at any other. For a transducer, it is the range of frequencies over which its efficiency, as a converter of electrical energy to sound energy or vice versa, is more than half its maximum. It is evident that, although the matching layer improves sensitivity at the centre frequency, it also acts as a frequency filter, reducing the bandwidth. A large transducer bandwidth is crucial to good axial resolution. The transducer bandwidth can be increased by using two, or more, matching layers, progressively reducing in characteristic impedance from the PZT to the patient’s skin. The reason why this leads to a greater bandwidth is that the impedance change, and hence the reflection coefficient, at the PZT front face is less than that for the case of a single matching layer. This applies at all frequencies, so there is less difference in the performance at the centre frequency relative to that at other frequencies. Improvements in backing layer and multiple-matching-layer technologies have meant that −3dB bandwidths greater than 100% of centre frequency are now available.
A large transducer bandwidth is also needed for harmonic imaging and other modern developments. It is also a prerequisite of a ‘multi-frequency transducer’. This type of transducer allows the operator to select one of a choice of operating frequencies according to the penetration required. Whatever centre frequency is selected, short bursts of oscillating voltage at that frequency are applied across the PZT plate to produce the ultrasound transmission pulses. In order for a single transducer to be able to operate at three frequencies, say 3, 5 and 7 MHz, it would need to have a centre frequency of 5 MHz and a bandwidth of 4 MHz, which is 80% of the centre frequency (Figure.10). Note that the bandwidths of the pulses generated at the upper and lower frequencies must be less than that of the transducer itself; this means the axial resolution to be expected from a probe in
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multi-frequency mode will be less than that which would be possible if the whole transducer bandwidth were used to generate a full 5 MHz bandwidth pulse .
Fig.10:- The −3 dB bandwidth of a transducer is the range of frequencies over which the output power for a given applied peak-to-peak voltage is within a factor of 2 of the maximum. A multi-frequency probe have a large bandwidth, so that it can transmit and
receive pulses with several different centre frequencies.
Developments in transducer technology
The performance of the transducer in terms of its efficiency and bandwidth is critical to the overall performance of the ultrasound system. Transducer performance has been improved through the development of new materials and fabrication techniques and through the introduction of new transducer technologies. Traditional PZT materials are piezoceramics formed from small polycrystalline grains of the base materials, moulded to shape and fired at high temperatures. During transducer manufacture, the PZT must be polarized to activate its piezoelectric properties, by heating while a high voltage is applied to the electrodes. The granular structure limits the alignment of piezoelectric domains during this process and hence the piezoelectric efficiency of the device. Transducers are now commercially available which use alternative piezo electric materials grown as single crystals. These include lead titanate (PT) doped with various other elements, such as lead, magnesium and niobium (PMN–PT) or lead, zinc and niobium (PZN–PT). When cut into wafers for transducer manufacture, the single-crystal structure allows complete polarization of the material and hence a stronger response in transmission and reception (greater sensitivity and penetration). The improved performance of these materials allows manufacturers to design transducers with increased bandwidth without loss of sensitivity, which is used to advantage in multi-frequency operation and harmonic imaging. However, the increased electrical and mechanical fragility of these materials at present limits their application to lower frequencies, which make use of greater element width and thickness. Prototype transducers have
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been demonstrated using capacitive micro-machined silicon transducers. The use of silicon technology gives the possibility of fabricating the transducer and its associated electronics on the same piece of silicon.
Focusing of an ultrasonic beam
For imaging purposes, a narrow ultrasound beam is desirable as it allows closely spaced targets to be shown separately in the image (lateral resolution). Narrow ultrasound pulses are produced by focusing the transducer. This may be accomplished for transducers with single piezoelectric elements by curving the element or by using an acoustic lens (Fig.11). These approaches are referred to as mechanical focusing. All ultrasound beams (focused or not) have widths (and lateral resolutions) that vary with distance from the transducer. The point of narrowest width occurs at the focal distance. For unfocused transducers, this point is also called the near field length. The region closer than this to the transducer is called the near field or Fresnel zone, and the more distant region is called the far field or Fraunhofer zone.
Fig.11:- Diagrams show the ultrasound beam profiles from (a) an unfocused transducer,(b) a mechanically focused transducer with a curved piezoelectric element, and (c) a mechanically focused transducer with an acoustic lens.
Focusing can be accomplished only in the near field of the transducer. For any particular frequency, the mechanical focusing characteristics are fixed as a function of the element curvature or the shape and composition of the lens. The near zone increases in length with increasing frequency. High-frequency beams have two advantages over low-frequency beams: depth resolution is superior, and the Fresnel zone is longer. It would seem logical to use high frequencies for all imaging. High frequencies, however, have a major drawback related to penetration. Tissue absorption increases with increasing frequency, so a relatively low-frequency beam is required to penetrate thick parts.
Multiple-zone focusing
Further improvement in lateral resolution, is possible by sub-dividing each scan line into two or more depth zones and interrogating each zone with a separate transmission pulse, focused at its centre. The operator might select transmission foci at two different depths – F1 and F2. These would be indicated by two
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arrowheads or other focus indicators down the side of the image. One pulse would be transmitted with a focus at F1 and echoes from depths up to about half-way between F1 and F2 would be captured. Then a second pulse would be transmitted with a focus at F2 and echoes from all greater depths would be captured. The greater the number of transmission focal zones, the greater the depth range over which the ‘effective transmission beam’ is narrow. Unfortunately, the greater the number of focal zones, the longer is spent on each scan line, and so the lower the frame rate.
3D/4D Transducers
Features in the individual 2D image data sets can be used to register the various image planes with one another allowing 3D rendered displays of tissue structures. The resulting 3D volume data set can then be processed to create a number of alternative modes of display. Repetition of the 3D acquisition and display at a few hertz (refresh rates up to 20Hz are possible) results in a moving 3Ddisplay, referred to as 4D.There are two commonly used approaches to the design of 3D / 4D transducers. The mechanical 3D transducer (Fig.12A), incorporates an array transducer, mounted on a swivel inside an enclosed bath of acoustic coupling fluid.
The transmit pulse and returning echoes pass through a thin acoustic window to acquire a 2D section from the adjacent tissues. A motor, coupled to the swivel by pulleys and a drive belt, rotates the transducer array so that the 2D scan plane is swept at 90° to the imaged plane, to interrogate and acquire echoes from a pyramidal volume of tissue. The orientation of each scan plane is measured by an angle encoder attached to the motor. A sequence of closely spaced, adjacent 2D sections is acquired, which constitutes a 3D volume data set. By repeating the sweeping motion of the array from side to side, a 4D image can be created. Due to the mechanical nature of the transducer movement, 4D acquisition rates are limited to a few volumes per second. The spatial resolution of the acquired 3D data, have reduced lateral resolution in the elevation plane, due to the lack of electronic focusing in the elevation direction.
The alternative, commonly used 3D transducer is the 2D matrix array (Fig.12B). This transducer contains several thousand square elements arranged in a two- dimensional matrix. Using beam-steering techniques, the matrix array can be used to create a 2D sector scan in a single plane. By applying beam-steering techniques also in the orthogonal direction (elevation), the 2D sector can be swept sideways to describe a pyramidal volume. As the matrix of elements is square, it is possible to achieve equivalent resolution in the lateral (scan plane) and elevation directions by applying the same degree of electronic focusing and aperture control to both planes. As the matrix array has no moving parts, the 3D acquisition rate is limited only by speed-of sound considerations. If the imaged depth and angle of sweep are restricted, volume rates of up to about 20Hz are possible. The stored volume data can be interrogated and displayed as a sequence
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of adjacent 2D sector scans in any B-mode plane or as a C-scan, where the imaged section is parallel to the transducer face.
Fig.12:-(A) The mechanical 3D transducer. (B) The 2D matrix-array transducer, volume data can be displayed as a series of B-scan planes or as a C-scan.
Endo-cavity transducers
Endo-cavity transducers are intended for insertion into a natural body cavity or through a surgical opening. The ability to place the transducer close to a target organ or mass means that there is less attenuation from intervening tissue, which in turn means a higher frequency may be used and superior lateral and axial resolution obtained. The image distortions and artefacts due to any tissue heterogeneity or strongly reflecting or refracting interfaces between the transducer and the target are also reduced. The choice of beam-forming techniques in endo- cavity transducers being determined primarily by the anatomical features and constraints of the particular application. Thus, a curvilinear array offers a field of view of an appropriate shape for trans-vaginal scanning. The wide field of view close to a linear array is suitable for trans-rectal imaging of the prostate. Phased arrays give a wide field of view for visualizing the left side of the heart from a transoesophageal probe (Fig.13).
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Fig.13:-Endo-cavitatory transducers. (a) Curvilinear transducer for trans-vaginal scanning. (b) ‘End-fire’ curvilinear-array transducer for trans-rectal or trans-vaginal scanning. (c) ‘Bi-plane’ trans-rectal transducer with both a linear array and a curvilinear array,allowing both transverse and longitudinal scans of the prostate. (d) Transoesophageal transducer with two phased arrays set at right angles, giving two orthogonal cross sections of the heart.
DIAGNOSTIC ULTRASOUND MODES
All modes are based on the principles of pulse-echo mapping. Real-time capability is achieved if the rate of transmission of pulses, called the pulse repetition frequency (PRF), is sufficient to allow the scanner to keep up with any changes in the anatomical region of interest, or with movement of the probe. However, an upper limit to the PRF is set by the requirement that there should be time for the echoes from each pulse to return from the deepest reflecting or scattering structures before the transmission of the following pulse. Otherwise late echoes from one transmitted pulse will be confused with early echoes from the following pulse.
A-Mode
In an A-mode scan, the amplitudes of echoes along a single scan line are represented as positive vertical deflections of a real-time graph on a display screen (Fig.14A). A-mode scans provide the most accurate way of measuring the amplitudes of echoes and the distance between two targets. A-mode is specially provided as an additional feature on some B-scan equipment, for example, in eye scanners, for measuring various axial dimensions.
M-Mode
M stands for motion, and M-mode shows how the positions of reflecting surfaces along a single scan line vary with ‘physiological time’ (Fig.14B). It is used primarily in cardiac work, in conjunction with B-mode, via which the direction of the M-line can be monitored and adjusted to run through the moving interfaces of interest (e.g. valve leaflets, chamber walls). On activating M-mode, the B-mode scanning
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action stops and all further transmitted pulses are sent down the selected scan line. The echoes resulting from each transmission pulse are presented as brightness modulations along a corresponding vertical line of the M-mode display, with the brightness (grey level) indicating echo amplitude. Each transmission-reception sequence results in a new M-mode line, displayed alongside the previous one, the data being continuously scrolled across the display. Most machines save the most recent say 10s-worth of the data that has been scrolled off the display, allowing it to be replayed if the acquisition of new information is frozen. There is often provision to show other physiological waveforms, such as an echocardiogram (ECG) or intra-cardiac pressure waveform, on the same timescale as the M-mode display.
Fig.14(A):- A-mode display from eye US. Echo amplitude is plotted against axial distance of the echoing feature. (B) M-mode display, with associated B-mode image across mitral valve.M-mode display allows dynamic study of its functioning. The vertical coordinate of the M-mode display represents range (depth) while the horizontal coordinate represents
time.
B-Mode
In a B scan, a probe held in the operator’s hand produces a sequence of beams which automatically sweep across a chosen scan plane in the patient’s body, defining a sectional (tomographic) field of view. The varying angle and position of the scan lines are electronically controlled so that, as echoes are received, the targets’ directions and ranges can be established. Echoes obtained from all scan lines are displayed in their correct respective positions, to form a 2D image of the scan plane. The amplitude of each echo determines the brightness (grey level) of the corresponding point on the display. Hence the name B-mode. (Brightness mode).This is the most widely used type of ultrasound imaging, and is often used in conjunction with other modes, such as M-mode and Doppler.
Doppler Modes
Color flow mapping: A 2D Doppler image of blood flow is color-coded to show the direction of flow to and away from the transducer.
Power Doppler (mode): This color-coded image of blood flow is based on intensity rather than on direction of flow, with a paler color representing higher intensity.
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It is also known as ‘angio’ mode.
Spectral or Pulsed wave Doppler: This Doppler mode uses pulses to measure flow in a region of interest. and displays how the blood velocity at that point is varying with physiological time. It provides functional information about the heart or vascular system.
Duplex: Presentation of two modes simultaneously: usually 2D and pulsed (wave).
Triplex: Presentation of three modes simultaneously: usually 2D, color flow, and pulsed Doppler.
3D: This is a image representation of a volume or 3D object, such as the heart or foetus. Surface rendering can be used to visualize surfaces. Another image presentation is volume rendering, in which surfaces can be semi-transparent or 2D slice planes through the object. Alternatively, there is simultaneous viewing of different 2D slice planes (side by side).
4D: A 3D image moving in time.
B-MODE INSTRUMENTATION (KNOBOLOGY)
The modern diagnostic imaging system is continuing to evolve and, as a result, is becoming more complicated with new modes and features. A typical ultrasound imaging system is with many controls designed to improve conventional gray- scale image quality (Table-5).
The number of actual controls on an ultrasound imaging system may seem bewildering at first, most modern systems start in a default set of control settings or ‘‘presets’’ that are optimized for a particular clinical application or transducer, so that with clinical training and moderate effort, such as the adjustment of the time gain compensation (TGC) controls, a reasonably good image can be obtained quickly. The main controls identified in the system control panel are
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the following:
Power Button
Switching on the ultrasound machine is the first knob to press if the machine is switched off. By switching on the machine, the ultrasound system is given access to a source of electricity. The ‘on’ and ‘off’ button are from a binary numbering system where “0” was for ‘off’ and “1” was for ‘on’,so as to create the universal symbol for Power.
Probe or transducer selection
Typically, most modern ultrasound machines provide three to five active ports/ connectors to which transducers can be plugged into, and probe selection switch allows the user to activate one of the arrays at a time.
Application preset
Modern machines allow the operator to preset an application setting for a certain examination type. Ultrasound imaging is used for a wide range of medical applications. Aside its use in assessing the abdomen, it is also used in obstetrics and gynecology, cardiac and vascular examinations, and other small-part examinations such as breast, thyroid, and musculoskeletal imaging. Different sonographic settings are needed for the various examinations, due to their differences in terms of the depth of region of interest, tissue-type, and the size of organs and structures in that region. Because of the uniqueness of these examinations, adjusting the settings between patients for a different examination can be time consuming, and may compromise the adherence to ALARA principles. In addressing this limitation, the manufacturer makes it easier by allowing the operator to select the type of examination which will activate the pre-defined factory settings for the specific type of examination. By selecting the appropriate ‘Application Preset’ for abdominal examination, the pre-defined factory settings are activated for abdominal sonography.This automatically adjusts the basic settings for the selected examination, which include an adjustment of the transducer frequency, acoustic Output Power, Overall Gain, Dynamic Range, Depth and other related settings. This givesyou a great starting point to further fine-tune your image.
Mode selection
Ultrasound machine will always start in B-mode or“greyscale” mode by default. Mode selection knobs,provides the means for selecting a mode of operation, such as B-mode, color flow, M-mode, or Doppler, individually or in combination
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(duplex or triplex operation).
Output power
The acoustic output power of the machine must be considered at all times by the operator.Selecting the appropriate preset will automatically adjust the output power to the recommended level.However, while it is important to observe the ALARA principle by using the minimum output power possible, the operator must not compromise image quality for output power reduction which may lead to misdiagnosis. In essence, there should be a balance between maximizing image quality with the minimum output power possible as a measure for reducing the risk of biological effect. Usually, the ultrasound machine will display the output power on the screen at all times, allowing the operator to be constantly informed.
Depth of scan control
The Depth knob adjusts the field of view (scan depth in centimeters).Structures within the field of view can be moved far or closer by adjusting the Depth.This is to ensure that the region of interest is closer enough for optimum visualization.It is also to avoid showing regions that are not relevant to the area of interest. The structure of interest should always take the center stage by occupying about two- thirds of the field view.
Gain (Overall)
After optimizing your depth, the next ultrasound setting you should adjust is your gain.Ultrasound gain simply means how bright or dark you want your image to appear. It achieves this by amplifying the echo-signals returning from the body after transmitting the sound waves.All ultrasound machines will have an “Overall” Gain setting that, when increased or decreased, will make the entire ultrasound image brighter or darker. This is good to use when your entire imaged is too dark (under-gained) or too bright (over-gained).
Time Gain Compensation (TGC)
Most other ultrasound machines will allow you to further adjust the gain in even
more specific areas of your ultrasound screen. This ultrasound setting is called “Time Gain Compensation” or TGC. Adjusting the Time Gain Compensation (TGC) allows you to adjust the gain at almost any depth of your ultrasound image, not just the near and far-fields. The top rows of the Time Gain Compensation control the near field gain and the bottom rows control the far-field gain. The TGC creates uniformity in the brightness of the echoes when used in conjunction with the overall gain. The best approach is to center all the TGC settings before adjusting the overall gain. After adjusting the overall gain, the TGC can then be
adjusted to compensate for attenuation at specific depth.
Focus or transmit focal length selection
This allows the location of the transmit focal length to be moved into a region of interest.Focal zone(s) allows the operator to improve lateral resolution in a region of interest.This is an additional measure to minimize the effect of attenuation.
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However, while other controls such as the overall gain and TGC are effective for improving axial resolution, adjusting the focal zone is much more effective for improving lateral resolution.The focal zone normally appears at the lateral side of B-mode as a triangular or hourglass shaped structure or a dot. It can be moved up or down by the operator and should be placed at the region of interest or posterior to that region. To improve lateral resolution in a wider region, more than one focal zones may be selected by the operator. However, increasing the number of focal zones also decreases the frame rate which has the tendency of slowing down the image production time to the detriment of temporal resolution.
Dynamic range control
When an ultrasound pulse is incident on an interface or a scatterer, some of the incident intensity is usually reflected or scattered back to the transducer. For reflection at a large interface, as might be encountered at an organ boundary, the reflected intensity ranges from less than 1% of the incident intensity for a tissue–tissue interface to almost 100% for a tissue–air interface. The intensities of echoes received from small scatterers depend strongly on the size of the scatterer and the ultrasound wavelength, but are usually much smaller than echoes from large interfaces. Hence, the range of echo amplitudes detected from different targets is very large. The dynamic range of signals at the transducer is defined as the ratio of the largest echo amplitude which does not cause distortion of the signal, to the smallest that can be distinguished from noise. The dynamic range is expressed in decibels. Weak echoes produced by scattering within tissue give information about organ parenchyma, while strong echoes from large interfaces give information about the size and shape of organs and other tissue features. To be useful diagnostically, a B-mode image should contain both types of echo. The range of echo amplitudes displayed needs to include echoes from a typical tissue to tissue interface (e.g. liver–fat) as well as echoes due to scattering within tissue. However, this range of echo amplitudes cannot be displayed without further processing. In practice, the range of echo amplitudes which need to be compressed into the display range depends on the application. For example, when trying to identify local tissue changes in the liver, a wide dynamic range of echoes needs to be displayed, so that these weak echoes are made to appear relatively bright in the image. For obstetric work, interpretation of the shapes in the image may be aided by reducing the displayed dynamic range, so that weak echoes are suppressed and areas of amniotic fluid are clearly identified.
Freeze
Freeze control stops transmission of acoustic output. In freeze mode, the image acquisition process (writing) is stopped, so that data in the image memory remain unchanged, but are still read out repeatedly to the display. Freeze mode is used while measurements are made on features in the image and hard-copy records made.
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Zoom
There are two ways to increase the magnification of part of an image. One is to use the conventional depth control, but this involves losing the deeper parts of the image from the display. Hence, features which are deep within the tissue cannot be magnified by this means. The second method, known as write zoom, presents a full-screen real-time image of a restricted area of the normal field of view, selected by the operator allows the display of a selected region of interest (ROI) remote from the transducer. The user outlines the ROI by adjusting the position and dimensions of an on-screen display box,information from within the ROI then fills the image memory.
Calipers (measure)
Calipers are an important feature of ultrasound machines that allows you to measure the distance of specific structures of interest.
Image/Video Capture
All ultrasound machines will allow you to save an image and/or video clip of your ultrasound scan. This is important if you are trying to archive or use any ultrasound images/videos as teaching files.
Analogue-To-Digital Conversion
The echo signals received by the transducer elements are analogue signals. That is, their amplitudes can vary continuously from the smallest to the largest value. At an early stage in the signal-processing chain, the echo signal are converted from analogue to digital form. In digital form, the signal is quantized i.e., it can have only a limited number of values, unlike the analogue signal, which was continuously variable. The analogue-to-digital converters (ADCs) used in modern B-mode systems convert the signal into a much larger number of more closely spaced values, so that the regenerated signal is a much more faithful recording of the original. Also, since the smallest value that can be digitized is smaller, the dynamic range (the ratio of the largest to smallest value) is greater. The precision to which the ADC can measure the amplitude of the signal is normally described by the number of bits (binary digits) available at the output. The number of bits also determines the dynamic range of signals that can be digitized, since this is equal to the ratio of the largest binary value to the smallest. Some modern
B-mode systems can digitize the echo signal to 12 bits, so that 4096 (212) levels are available, and a dynamic range of approximately 72 dB can be processed and stored. Once a signal has been converted into digital form, it consists simply of a set of numbers, which are essentially immune to noise, interference and distortion. Digital information can be stored conveniently in electronic memory or on a magnetic or optical medium without degrading with time. The retrieved image information is identical to the original. The most important advantage of digitization for B-mode imaging is that it makes digital processing of echo
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information possible. Using built-in, dedicated computing devices, digital echo information can be processed by powerful mathematical techniques to improve the image quality.
Image formation and storage
The B-mode image is assembled in the image memory using the processed amplitude signal and information from the beam-former on the time since transmission (range), and the position and direction of the ultrasound beam, i.e. the scan line. This process is called writing to the image memory. Interpolation is used extensively in writing to the image memory. The image memory divides the image into a 2D array of picture elements or pixels. The binary number representing the amplitude of the echo from this point is written to this memory location. To view the image, it must be read out to a display device in a sequence which is compatible with the display format. The display monitor screen is also composed of a 2D array of elements, each of which emits light when activated. These are addressed one at a time by the monitor in a sequence of horizontal lines or raster scan, starting at the top left of the screen and finishing at the bottom right. The stored image is read out from the image memory by interrogating each pixel in turn in synchronism with the display raster. The stored binary values are converted to an analogue level to control the brightness of the display. The process does not degrade the information stored in the memory and can be repeated indefinitely. Hence, the image memory performs the function of scan conversion from ultrasonic line format to monitor display format.
Real-time display
In most medical imaging processes (e.g. X-ray, CT and MRI), there is a significant time delay between image acquisition and image display,so that diagnosis is made on information which may be many minutes old. Formation of a single ultrasound B-mode image in the image memory takes typically of the order of 1/30s, so that, if repeated continuously, 30 images can be formed each second. This rate is referred to as the frame rate and is measured in hertz. Also, the time delay between image acquisition and display is relatively short (a few tens of milliseconds). This leads to the description ‘real-time’ imaging, i.e., images of events within the patient are displayed virtually as they happen. The real-time nature of the process is an important aspect of ultrasound imaging and, coupled with the use of a hand-held probe in contact with the patient, gives a highly interactive form of investigation. Real-time imaging allows the study of the dynamic behaviour of internal anatomy, such as the heart, and can aid identification of normal anatomy (e.g. gut from peristalsis) and pathology (e.g. movement of a gallstone).The display monitor typically displays a complete image in 1/25s, giving 25 images per second. To avoid conflicts between the reading and writing processes, buffer memories are used to store echo data on a temporary basis, so that the two processes can proceed independently of each other.
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1. DIAGNOSTIC ULTRASOUND- BASICS OF PHYSICS AND KNOBOLOGY
Cine loop
Most current B-mode systems have multiple image memories (up to 2000) and continue to store each frame as it is formed in the next available image memory, eventually cycling around to overwrite memory number 1 when number 2000 is filled. The image is frozen at the end of the dynamic event of interest or examination and the multiple image memories allow the last 2000 frames to be reviewed in real time or individual frames to be selected and printed as hard copies. Sections of the stored data can be stored as movie clips as part of a patient report and can also be used for teaching purposes. For the heart, which has a regular, repetitive movement, the recorded frames can be displayed as a cine loop, repeating the cardiac cycle and allowing review and diagnosis after the examination has been completed.
Display, output, storage and networking
A modern ultrasound system has a flat-screen display, has some capacity locally for storing video clips and images, and is connected to an imaging network, so that any hardcopy is performed using a high-quality communal printer, and all data are archived in a large central datastore. The era of the stand-alone ultrasound system which is connected to its own printer and own video recorder is rapidly becoming a thing of the past.
Display: Some time ago, analogue TV-type display monitors gave way to video graphics array (VGA) monitors and their variants. These display a digital image developed within the monitor from an analogue video signal provided by the scanner. More recently, digital video interface (DVI) monitors have appeared, no longer using a bulky cathode ray tube, but a liquid-crystal display (LCD). The benefit of the digital interface is that the image pixels can be transferred from frame store to display (via the post-processing algorithm) without conversion to analogue video and back again.
Local storage: A modern ultrasound imaging system usually contains an on- board computer. Individual images and image sequences may be stored on the local hard disc, and retrieved and displayed as necessary. Older ultrasound systems tended to store image and video data in proprietary formats, which made it difficult to transfer images to an imaging network or for the operator to incorporate digital images into reports and slide-shows. Modern systems allow storage of images in several widely used formats, such as ‘avi’ and ‘mpeg’ for video, ‘tiff’ and ‘jpeg’ for individual images, and DICOM format for networking. They also have computer compatible recording facilities such as a USB port and a DVD writer.
Integration with an imaging network: The modern radiology department is based on imaging systems, including ultrasound scanners, which are connected to an imaging network, picture archiving and communication system (PACS). A PACS system typically has workstations for review and reporting of image data,
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printers for producing hard copy of images and reports, and a large memory store for archiving patient data. Most reporting for ultrasound is performed in real time, with the operator saving images or video clips which illustrate the lesion (if present) or the measurement. Hard copy of images for inclusion in patient notes is usually done using a high-quality laser printer. Instead, if the referring clinician wants to inspect the images he or she can do so at a local terminal, connected to the hospital network, on the ward or in the clinic. Compared to film-based radiology, PACS has considerable advantages including: rapid access of images throughout the hospital, reduced requirement for hard copy and hard copy storage, reduction in lost images, and hence reduction in repeat examinations due to lost images.
Bibliography:-
• Hangiandreou N J. B-mode US: Basic Concepts and New Technology. RadioGraphics 2003; 23:1019–1033.
• Szabo TL. In:Diagnostic Ultrasound Imaging: Inside Out. Burlington, USA. Elsevier Academic Press.2004.
• Hoskins P, Martin K, Thrush A. In:Diagnostic Ultrasound:Physics and Equipment.Second edition. U.K.,Cambridge University Press 2010.
• Ziskin M C. Fundamental Physics of Ultrasound and Its Propagation in Tissue. RadioGraphics 1993; 13:705-709.
• Curry T S, Dowdey J E, Murry R C. Ultrasound. In: Christensen’s Physics of Diagnostic Radiology. 4th Edition. Philadelphia,USA: Williams & Wilkins;1990.p 323-371.
• Fairhead A C,Whittingham T A. Diagnostic Ultrasound. In: Dendy P P, Heaton B. Physics for Diagnostic Radiology.Third Edition. Boca Raton, Florida:CRC Press;2012,p-489-562.
• Lacefield J C. Physics of Ultrasound. In:Dance D R., Christofides S,Maidment A D A, McLean I D, Ng K-H. Diagnostic Radiology Physics:A Handbook for Teachers and Students. Austria:IAEA;2014.p-291-330.
• E. Forster. Ultrasound.In: Equipment for Diagnostic Radiography. Lancaster, England:MTP Press Limited;1985.p-180-189.
• Roberts P A,Williams J. Imaging with Ultrasound. In:Farr’s Physics for Medical Imaging.Second Edition.Europe:Saunders Elseivier Ltd.2008.p-147-168.
• Goldstein A. Overview of the Physics of US. RadioGraphics 1993; 13:701-704.
• Hill C R, Bamber J C, Haar G R. Physical Principles of Medical Ultrasonics.
Second Edition. West Sussex,England.John Wiley & Sons Ltd.2004.
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Doppler US 2 Author: Dr. Aman Gupta
DOPPLER ULTRASOUND
The ability to detect and quantify blood flow by means of Doppler ultrasonography has made this technique an indispensable adjunct to imaging. In the past, use of Doppler US was restricted to evaluation of a few well-defined indications in cardiac disease and assessment of carotid and peripheral arterial and venous disease. The importance of flow information, however, extends well beyond the identification of stenosis, thrombosis, and occlusion of the major arteries and veins. Doppler techniques provide critical diagnostic information non-invasively about the fluid dynamics of blood circulation and abnormalities. Doppler US is playing a growing role in the diagnosis of abdominal, pelvic, and foetal disorders. Familiarity with the elementary principles of Doppler US is essential for proper use of duplex US and colour Doppler imaging. Doppler detection is a blend of physics and specialized signal processing techniques required to extract, process, and display weak Doppler echoes.
The Doppler Effect
The Doppler effect is a change in the perceived frequency of sound emitted by a moving source. The effect was first described by Christian Doppler in 1843. One experiences the Doppler effect regularly while standing next to a street as traffic passes. By noting the pitch of the sound emanating from vehicles (horns, music, engine sounds) and paying attention to the pitch change as the vehicle direction changes from approaching toward the observer to moving away. The
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shift in frequency is related to the contraction or expansion of wavelengths ahead of or behind the sound-emitting moving object (Fig 1). The source is moving towards the observer,on right of figure,as it transmits the sound wave. This causes the wavefronts travelling towards the observer to be more closely packed, so that the observer witnesses a higher frequency wave than that emitted. The source is moving away from the observer, on left of figure, the wavefronts will be more spread out, and the frequency observed will be lower than that emitted. The resulting change in the observed frequency from that transmitted is known as the Doppler shift , and the magnitude of the Doppler shift frequency is proportional to the relative velocity between the source and the observer.
Fig.1:- Doppler effect: as the source moves towards, the observer detects sound at a higher frequency than that transmitted; as the source moves away, the observer detects sound at lower frequency than that transmitted.
In medical ultrasonography, a Doppler shift occurs when reflectors move relative to the transducer. The frequency of echo signals from moving reflectors is higher or lower than the frequency transmitted by the transducer, depending on whether the motion is toward or away from the transducer. The Doppler shift frequency, or simply the Doppler frequency, is the difference between the received and transmitted frequencies (Fig.2).
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2. DOPPLER ULTRASOUND
Fig.2: Diagrams illustrate the interaction of ultrasound with a reflecting interface (blood), when it is stationary (A),moving toward (B) and moving away (C) from the transducer. If the target is stationary (A), the reflected ultrasound has the same frequency or wavelength as the transmitted sound (fr=ft). If the target moves toward the transducer (B), the difference in reflected and transmitted frequencies is greater than 0; if the target is moving away from the transducer (C),this difference is less than 0. The Doppler equation relates this
change in frequency to the velocity of the moving object.
DOPPLER EQUATION
Ultrasonic Doppler equipment is for detecting and evaluating blood flow. A typical arrangement is illustrated in Fig.3. An ultrasonic transducer is placed in contact with the skin surface; it transmits a beam whose frequency is ft. The received frequency fr will differ from ft when echoes are picked up from moving scatterers, such as the red blood cells. The Doppler frequency (fd) is defined as the difference between the received and transmitted frequencies. The fd is calculated by the following: (the Doppler equation)
fd = fr - ft = (2 ft v cosθ)/c
Thus, the Doppler shift frequency fd depends on the frequency of the transmitted ultrasound, the speed of the ultrasound as it passes through the tissue (c) and the velocity of the blood (v). The detected Doppler shift also depends on the cosine of the angle θ between the path of the ultrasound beam and the direction of the blood flow (Doppler angle). This angle is known as the angle of insonation. The angle of insonation can change as a result of variations in the orientation of the vessel or the probe. Many vessels in the upper and lower limbs run roughly parallel to the skin, although if the vessel is tortuous the angle of insonation will alter. In the abdomen there is considerable variation in the orientation of vessels which will result in many different angles of insonation. The operator is also able to alter the angle of insonation by adjustment of the orientation of the probe on the surface of the skin.
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Fig.3:- Arrangement for detecting Doppler signals from blood. The angle θ is the Doppler angle,which is the angle between the direction of motion and the beam axis, looking toward the transducer.
The Doppler angle strongly influences the detected Doppler frequency for a given reflector velocity. When flow is directly toward the transducer, θ is 0 degrees and cos θ is 1. The Doppler frequency detected for this orientation would be the maximum one could obtain for the flow conditions. More typically, the ultrasound beam will be incident at an angle other than 0 degrees, and the detected Doppler frequency will be reduced according to the cos θ term. For example, at 30 degrees, the Doppler frequency would be 0.87 multiplied by what it is at 0 degree; at 60 degrees, it would be 0.5 multiplied by its 0-degree value. Finally, when the flow is perpendicular to the ultrasound beam direction, θ is 90 degrees and cos θ is 0; there is no detected Doppler shift! In practice, the transducer beam is usually oriented to make a 30- to 60-degree angle with the arterial lumen to receive a reliable Doppler signal. The relationship between an angle and the value of its cosine is shown in Figure.4. Since the cosine of the Doppler angle changes rapidly for angles more than 60°,accurate angle correction requires that Doppler measurements be made at angles of less than 60°. Above 60° relatively small changes in the Doppler angle are associated with large changes in cos θ; therefore, a small error in estimation of the Doppler angle may result in a large error in the estimation of velocity. For this reason, the use of a Doppler angle of less than 60° is recommended to obtain accurate estimates of velocity.
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2. DOPPLER ULTRASOUND
Fig.4:- The cosine function. This has a maximum value of 1.0 when the angle is 0°, and a value of 0 when the angle is 90°. Note the cosine of the Doppler angle changes rapidly for angles more than 60°
BASIC DOPPLER US TECHNOLOGY
Continuous-Wave Doppler Us
Conceptually speaking, the simplest technology for detection of flow with US is continuous-wave (CW) Doppler US. The term continuous wave means that sound is emitted from a transmitting transducer 100% of the time. Sound that echoes back must then be detected by a second receiving transducer. A continuous-wave instrument has an arrangement whereby the two transducers have some overlap of their beams, resulting in a region of interest where the Doppler shifted signals may be detected (Fig.5).
Fig.5:-Schematic of the continuous-wave Doppler instrument. The Doppler signal is obtained by demodulating the amplified echo signals and then applying a low- pass filter.
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The Doppler shifted frequencies are obtained by comparison of the transmitted signal with the received signal. In Figure.5, the continuous sound source is transducer #1. The backscattered sound is detected by transducer #2. In the analysis circuit, the outgoing signal frequency, ft, is compared with the returning signal, fr, to determine the shift in frequency related to moving structures. Echoes returning to the transducer have frequency fr. These signals are amplified in the receiver and then sent to a demodulator to extract the Doppler signal. Here, the signals are multiplied by a reference signal from the transmitter, producing a mixture of signals, part having a frequency equal to (fr + ft) and part having a frequency (fr - ft). The sum frequency (fr + ft) is very high-about twice the ultrasound frequency-and is easily removed by electronic filtering. This leaves signals with frequency (fr - ft) at the output, which is the Doppler signal. Continuous-wave (CW) Doppler is done in a variety of instruments, ranging from simple, inexpensive hand-held Doppler units, to “high-end” duplex scanners in which CW Doppler is one of several operating modes. At frequencies between 2 and 10MHz,the Doppler shift falls in the audible range for most physiologic motions. It can be used to detect blood flow or its absence in both arteries and veins, to identify vascular constrictions with their associated eddy currents and venturi jets, and to detect foetal heart motion earlier than with any other method. The filtered output Doppler signal which is within the audible frequency range can be applied to a loudspeaker or headphones for interpretation by the operator. The signals can also be recorded on audiotape or applied to any of several spectral analysis systems.
Take example of a typical examination of arterial blood flow. Assume blood flow to be at an average velocity of 20 cm/sec (v = 0.2 m/s), the operating frequency of the transducer to be 5 MHz (ft =5,000,000 Hz), and the angle θ that the sound beam makes with respect to the vessel lumen to be 60° (cos 60° = 0.5):
Doppler shift = (2(5,000,000)(0.2)(0.5))/1540 = 650Hz
Thus, for relatively slow-moving structures, the Doppler shift signal falls within the audible range.
• A basic Continuous-Wave Doppler unit usually has following controls:
• Transducer Transmit power: to vary the amplitude of the signal from the
transmitter to the transducer, thus changing the sensitivity to weak echoes.
• Gain, to vary the sensitivity of the unit.
• Audio gain, to vary the loudness of Doppler signals applied to loudspeakers.
• Wall filter: it is possible to eliminate signals of certain frequency ranges
from the output. For example, when studying blood flow, relatively low- frequency Doppler signals originating from movement of vessel walls may be eliminated from the output by applying a high-pass filter.
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