Ultrasound-Guided Regional Anesthesia

peripheral nerve blocks are frequently performed in children to provide anesthesia or analgesia during the perioperative period. Success depends on the ability to accurately place the needle—and thereby the local anesthetic—close to the target nerve without causing injury to the nerve or adjacent structures. Peripheral nerve blocks are not without risk and can pose a serious challenge even to the experienced anesthesiologist because they are usually performed after the child is anesthetized. In the past, clinicians relied on anatomic landmarks, fascial clicks, loss of resistance, or nerve stimulation to position the needle in the vicinity of the nerve. Anatomic landmarks provide valuable clues to the position of the nerve, but they are surrogate markers, lack precision, vary among children of different ages, and may be difficult to locate in obese children. Even nerve stimulation, which has been recommended as the gold standard for nerve localization, may not always elicit a motor response and its use does not guarantee success or preclude complications. Moreover, the accuracy of needle placement cannot be predicted with any of these methods, which may lead to multiple attempts to place the needle that may result in pain and possibly an incomplete or failed nerve block.

Various imaging modalities, such as fluoroscopy, computed tomography (CT), and magnetic resonance imaging (MRI), improve the accuracy of block placement in adults. However, these adjuncts are rarely used in children and are not practical in the operating room. The use of ultrasound (US) to guide peripheral and central neuraxial blocks has improved both accuracy and safety in both adults and children. In this chapter, the basic principles of US imaging and US-guided regional anesthesia (USGRA) in children are described. It is assumed that the reader has a basic understanding of common landmark-based nerve block techniques in children (see Chapter 42 ).

The use of US for regional anesthesia dates back to 1978, when La Grange et al. used a Doppler flow detector to locate the subclavian artery and guide supraclavicular brachial plexus blocks. In 1994, Kapral and colleagues published the first report on direct sonographic visualization in regional anesthesia. They used US to directly visualize the brachial plexus and observe the spread of the local anesthetic in real time during supraclavicular brachial plexus block. Today, US is commonly used to guide regional anesthesia techniques in both adults and children. Improvements in US technology and the availability of portable US machines with high-resolution imaging capabilities allow direct visualization of the peripheral nerves and central neuraxial structures in children; this can be compared to removing a blindfold from the anesthesiologist performing regional anesthesia. Currently, outcome data that prove US increases the safety and efficacy of regional anesthesia in children are rapidly accumulating.

Principles of Ultrasound

Sound is a form of mechanical energy that propagates through a medium as a wave of alternating pressure, causing local regions of compression and rarefaction ( Fig. 43.1 ). The frequency (f) of sound is the number of cycles of oscillations per second made by the sound source and the particles in the medium through which it moves. It is expressed in hertz (Hz, cycles per second). Sound waves propagate symmetrically away from the source at a constant velocity (v), which is the speed of sound in the medium. Distance between the wavefronts is the wavelength (λ) of the sound. The speed of sound through a medium can thus be represented as:

v = f \times λ

$v = f \times λ$

Amplitude is the strength of a sound wave, and the unit used to describe it is decibels (dB). The velocity of transmission of sound through a medium depends on its acoustic impedance and is determined by factors such as the stiffness, elasticity, and density of the medium. This accounts for the varying velocity of sound transmission through different tissues in the human body ( Table 43.1 ). The average velocity of sound transmission through biologic tissue is 1540 m/second. If the time taken by the US signal to return to the transducer is known, the distance of the target from the transducer (depth) can be computed.

TABLE 43.1

Propagation Velocity of Sound in Body Tissues

Tissue	Propagation Velocity of Sound (m/second) ^a
Bone	4080
Muscle	1580
Blood	1570
Kidney	1560
Liver	1550
Soft tissue (average)	1540
Water	1480
Fat	1450
Lung	600
Air	330

a Medical ultrasound device measurements are based on an assumed average propagation velocity of 1540 m/second.

The human ear can detect sound between 20 and 20,000 Hz. US is sound with a frequency beyond 20,000 Hz (20 KHz). For medical imaging, US typically uses a frequency between 1 million and 15 million Hz (megahertz [MHz]) and is produced by a piezoelectric crystal (element) within the transducer. When an electrical field is applied to the surface of an element in a transducer, it undergoes dimensional changes that cause it to vibrate and produce sound. The element is typically driven by a pulsed alternating voltage. This results in the generation of short pulses of US that are emitted into body tissues. Between successive short pulses of US generation, the transducer does not transmit but rather functions as a receiver of the reflected US energy (i.e., the echoes). The percentage of time that a transducer is transmitting is termed the duty factor and is typically less than 1%. The US transducer thus has a dual function—it functions as both a transmitter and a receiver.

The emitted US signal travels through the tissue medium, and when it encounters a tissue interface it is reflected back. The degree of reflection of US from tissues is related to the changes in acoustic impedance (Z) between two tissue interfaces. The reflected echoes are detected by the transducer and converted into electrical energy; they are then processed by the US machine according to their strength and displayed as dots on the monitor. The brightness of each dot corresponds to the strength of the echo signal. Strong echoes produce bright white dots, weak echoes produce gray dots, and anatomic structures that do not reflect US appear as black dots. The position of the dot on the monitor represents the depth from which the echo is received. When all of these dots are combined, they produce a complete image of the area scanned.

Modes of Ultrasound

B-Mode (Brightness) or Two-Dimension Mode

B-mode (brightness) or two-dimensional (2D) mode, is the most commonly used US mode. In this mode, the echo is converted to a dot and the brightness of the dot represents the amplitude of the returning signal. The position of the dot on the display represents the depth from which the signal is returning and depends on the round-trip time of the US signal. Multiple scan lines across a plane are combined to produce a single 2D gray-scale image. A series of frames are then displayed in rapid succession to give the impression of constant motion, the quality of which depends on the number of images displayed per second, that is, the frame rate .

M-Mode (Motion)

Motion, or M-mode, US is directed along a single scan line (sample line), and reflected signals along this scan line are converted to a brightness scale and displayed against a time axis. Because M-mode is produced from US signals along a single scan line, the 2D anatomy of the underlying body tissues should be studied first using the 2D mode. M-mode is of particular interest when time resolution is necessary, such as when examining a target with rapid movement (e.g., the mitral valve during echocardiography).

Doppler Ultrasound

Doppler US (based on the Doppler principle) detects a shift in frequency between the emitted US waves and their echoes. It is used to detect and measure blood flow, and the major reflector for this purpose are red blood cells. Several modes are available:

Color Doppler measures and color codes the direction and magnitude of the mean Doppler frequency shifts that occur in moving red blood cells and superimposes a color depiction of these data on the gray-scale image ( Fig. 43.2A ).
Power color Doppler depicts the amplitude, or power, of the Doppler signals ( Fig. 43.2B ). This allows better sensitivity for visualization of small vessels, but at the expense of directional information.
Pulsed Doppler allows a sampling volume (or gate) to be positioned in a vessel visualized on the gray-scale image and displays a spectrum of the full range of blood velocities within the gate plotted as a function of time ( Fig. 43.2C ).

The Ultrasound Machine

US machines are made up of the following components: a monitor (where the clinical images are displayed), the US unit (where the signals are processed), the control panel (with the knobs and controls), one or more transducers , and a data storage device . For an anesthesiologist's first encounter with a US machine the wide array of knobs and controls available may be confusing. However, several controls are common in most US machines, and a clear understanding of their function—that is, knobology —is essential for optimal imaging.

Presets

Most US machines have a number of presets, which are factory set, to allow optimal US imaging of a specific area of the body or type of examination. Some of the categories of presets that are available include small part, vascular, breast, nerve, musculoskeletal, abdominal, and so on. For example, if a small part preset is chosen, the US machine assumes that the operator is scanning for small, relatively superficial structures and automatically adjusts the depth, power, focus, gain, and time-gain compensation (TGC) to allow optimal imaging of superficial structures. Some US machines also allow customized presets according to clinical requirements. Power output is the amount of energy transmitted from the US transducer. In most machines, the power cannot be adjusted by the operator but is automatically set when a particular preset is chosen.

Frequency

This control is used to select the desired frequency, within certain limits, of a broadband transducer—that is, a transducer that serves a range of frequencies. In some US systems (Edge, SonoSite Inc., Bothell, WA), this is available as an image optimization control. In the “Res” (resolution) setting the highest frequency of the broadband transducer is selected; in the “Pen” (penetration) setting the lowest frequency is selected; and in the “Gen” (general) setting an intermediate frequency is selected.

Gain

The gain control adjusts the amplification of the returning acoustic signals and is used to optimize the US image ( Fig. 43.3A ). Reduced gain produces a dark image ( Fig. 43.3B ) and detail is masked. In contrast, too much gain produces a white image and detail is saturated ( Fig. 43.3C ). Some US machines have separate controls for overall gain and gain for the near and far fields. “Auto-gain,” by which the US machine automatically adjusts the gain, is also available in some machines.

Time-Gain Compensation

US energy is progressively attenuated as it travels through tissue. Therefore signals returning from reflectors at a depth are weaker in strength. By selectively amplifying the echoes from greater depths, using the TGC method or depth-gain compensation (DGC), equal reflectors at unequal depths are displayed as structures of equal brightness on the monitor. TGC is preset to a large degree, and the operator can make fine adjustments if necessary. The TGC control is presented as a series of sliders arranged in a vertical fashion on the control panel. Each of the sliders adjusts the amplification of the returning US signals at a specific image depth.

Depth

Adjustment in the displayed depth may be necessary, depending on the location of the target, the patient's body size, or other anatomic factors. A depth greater than necessary should not be chosen because this reduces the frame rate and resolution of the image.

Focus (Focal Zone)

The focus of the US signal occurs at a point at which the beam is at its narrowest width. It is also the region where lateral resolution is the best. The focus point should therefore be positioned at the depth at which the relevant anatomic structures are located. In some US machines, the operator can select multiple focal zones, but this markedly reduces the frame rate and thus should not be routinely used.

Freeze and Unfreeze

The “freeze” function allows the operator to lock a static image on the monitor. A number of frames (usually 20 or more) are also simultaneously stored in a memory bank. These stored frames can be scrolled back and forth. The selected still image can then be used for annotation, documentation, storage, review, or teaching. Pressing the freeze button once again will unfreeze the image.

Ultrasound Transducers

The transducer functions both as a transmitter and a receiver of the ultrasound signal. Three types of transducers are currently used ( Fig. 43.4 ): (1) in a linear-array transducer, the piezoelectric crystals are arranged in a linear fashion and sequentially fired to produce parallel beams of ultrasound in sequence, creating a field of view that is rectangular and as wide as the footprint of the transducer ( Fig. 43.4A ); (2) a curved linear-array transducer has a curved surface, creating a field of view that is wider than the footprint of the probe ( Fig. 43.4B ), but at the cost of reduced lateral resolution in the far field as the scan lines diverge; (3) a phased-array transducer has a small footprint, but the ultrasound beam is steered electronically to produce a sufficiently wide far field of view. The ultrasound beam diverges from virtually the same point in the transducer ( Fig. 43.4C ). Phased-array transducers are routinely used for transthoracic echocardiography. The footprints of these transducers are small enough to fit between the ribs and still produce a wide far field of view to image the heart. US transducers are typically broadband and can serve a range of frequencies. For example, a transducer with the notation HFL38/13-6 indicates that it is a high-frequency broadband (13-6 MHz) linear transducer with a 38-mm footprint. Note that the nomenclature used for transducers varies among manufacturers of US devices.

Ultrasound Transducer Selection

Resolution is the ability to distinguish two objects that are close together. Axial resolution is the ability to distinguish two objects that are along the axis of the US beam, and lateral resolution is the ability to distinguish two objects that are side by side. High-frequency US (13-6 MHz) has a higher axial and lateral resolution compared with low-frequency US but cannot penetrate as deeply into body tissue. Therefore high-frequency US is used to image superficial structures such as the brachial plexus in the interscalene groove or supraclavicular fossa. A lower-frequency US transducer (10-5 MHz) is suited for slightly deeper structures, such as the brachial plexus in the infraclavicular fossa, whereas a low-frequency US transducer (5-2 MHz) is used to image deep structures, such as the lumbar plexus or the sciatic nerve. Broadband transducers allow a single transducer to be used for scanning over a wide range of depths. Because regional blocks are performed at relatively shallow depths in neonates, infants, and young children, high-frequency linear transducers are used for most procedures. High-frequency linear transducers with a small footprint (13-6 MHz, 25 mm), in either a hockey stick or linear configuration, are particularly suited for young children.

Essentials of Musculoskeletal Ultrasound Imaging

Axis of Scan

In diagnostic ultrasonography, scans are performed in the transverse, longitudinal (sagittal), oblique, or coronal axis. During a transverse (axial) scan the transducer is oriented at right angles to the target, producing a cross-sectional display of the structures ( Fig. 43.5A ). During a longitudinal scan, the transducer is oriented parallel to and along the long axis of the target (e.g., a blood vessel or nerve) ( Fig. 43.5 B ). During USGRA procedures, US scans are most commonly performed in the transverse axis. In this axis, the nerves, the adjoining structures, and the circumferential spread of the local anesthetic are easily visualized.

Probe and Image Orientation

The US image must be properly oriented to accurately identify the anatomic relations of the various structures on the monitor. To facilitate this, all US probes have an orientation marker, which is usually represented by a groove or a ridge on one side of the transducer and corresponds to a green dot (or a logo) on the monitor. By convention, the orientation marker on the transducer is directed cephalad when performing a longitudinal scan and directed toward the right side of the patient when performing a transverse scan. This way the orientation marker on the left upper corner of the monitor always represents the cephalad end during a longitudinal scan or the right side of the patient during a transverse scan. The top of the display monitor therefore represents superficial structures and the bottom of the monitor the deep structures.

Echogenicity

Certain terms are frequently used to describe the sonographic appearance of musculoskeletal structures ( Fig. 43.6 ):

Echogenic : A bright white structure against a dark background
Reflective : Synonymous with an echogenic structure
Isoechoic : A shade of gray that is of the same brightness or echogenicity as the surrounding tissues
Hyperechoic : A shade of gray that is bright white or brighter than the surrounding tissues
Hypoechoic : A shade of gray that is dark or less bright than the surrounding tissues
Anechoic : An absence of echoes, hence blackness

FIGURE 43.6, Schematic diagram to demonstrate the relative echogenicity of various tissues.

Axis of Intervention

The plane of US imaging is approximately 1 mm thick ( Fig. 43.7 ); for a needle to be visible during US imaging it must lie within this narrow plane of imaging. During USGRA procedures, the block needle is inserted either outside of the plane (out-of-plane approach) ( Fig. 43.8A ) or within the plane of the US beam (in-plane approach) ( Fig. 43.8B ). In the out-of-plane approach, the needle is inserted in the short axis and is initially outside the plane of imaging and therefore not visible. It becomes visible only when the needle crosses the plane of imaging and is seen as an echogenic dot on the monitor ( Fig. 43.8A ). It is important to note that this echogenic dot may be just the cross-sectional image of the shaft of the needle as it passes through the plane of the US beam and thus may not represent the tip of the needle. In the in-plane approach, the needle is inserted along the long axis of the transducer in the plane of imaging and therefore both the shaft and tip of the needle are visible.

FIGURE 43.7, The plane of ultrasound imaging. Note that the ultrasound beam is only 1 mm thick and for a needle to be visible during an ultrasound-guided intervention it must lie within this plane of imaging.

FIGURE 43.8, Axis of intervention. A, Out-of-plane and B, in-plane techniques.

Both approaches are commonly used, and no data have shown that one is better than the other. Proponents of the out-of-plane approach have had great success with this method and claim that it causes less needle-related trauma and pain because the needle is advanced through a shorter distance to the target. However, critics express concerns that the inability to reliably visualize the needle and to use tissue movement as a surrogate marker to locate the needle tip during a procedure can lead to complications. The needle is better visualized in the in-plane approach, but this requires good hand–eye coordination. Moreover, there are claims that the in-plane approach also causes more discomfort in awake patients because longer needle insertion paths are required.

Needle Visibility

The ability to visualize the needle during a US (USG) procedure is critical for precision, safety, and success. However, this is often limited by the dispersion of the reflected US signals away from the transducer. Several factors have been identified that can influence needle visibility. The shaft of the needle is better visualized in the long axis than in the short axis, and its visibility decreases linearly with steep angles of insertion and smaller needle diameters. The needle tip is better visualized when it is inserted in the long axis for shallow angle of insertion (<30 degrees) and in the short axis when the angle of insertion is steep (>60 degrees). To overcome the effect of angle on needle visibility, some US machines allow the operator to steer the US beam toward the needle (“beam steering”) during steep needle insertions. However, this requires experience and decreases in needle visibility can still occur. Manufacturers have incorporated microscopic glass beads or reflectors onto the surfaces of block needles, or have laser-etched the tips of block needles to improve their visibility (echogenic needle).

The anesthesiologist's skill in aligning the needle along the plane of imaging is by far the most important factor influencing needle visibility because minor deviations of even a few millimeters from this plane will result in inability to visualize the needle. Even with experience, needle tip visibility is a problem when performing blocks at a depth in areas that are rich in fatty tissue. Under such circumstances, gently jiggling (rapid in-and-out movement) the needle and observing tissue movement or performing a test injection of saline solution or 5% dextrose (0.5–1 mL) and observing tissue distention can help locate the position of the needle tip. Five percent dextrose is preferred for the latter when nerve stimulation is used because it does not increase the electric current required to elicit a motor response.

Anisotropy

Anisotropy, or angular dependence, is a term used to describe the change in echogenicity of a structure with a change in the angle of insonation of the incident US beam ( Fig. 43.9 ). It is frequently observed during scanning of nerves, muscles, and tendons. This occurs because the amplitude of the echoes returning to the transducer varies with the angle of insonation. Nerves are best visualized when the incident beam is at right angles ( Fig. 43.9A ); small changes in the angle away from the perpendicular can significantly reduce their echogenicity ( Fig. 43.9B ). Therefore during USGRA procedures, the transducer should be tilted, from side to side, to minimize anisotropy and optimize visualization of the nerve.

Identification of Nerves, Tendons, Muscle, Fat, Bone, Fascia, Blood Vessels, and Pleura

Nerves

On a transverse scan, nerves appear round, oval, triangular, lip-shaped, or even flat. Nerves also assume different shapes along their course, depending on the surrounding structures. The echogenicity of nerves also varies and depends on the nerve and area scanned. They are generally hyperechoic and stand out in the background of the hypoechoic muscles ( Fig. 43.10A ), but they can also appear hypoechoic with a hyperechoic rim ( Fig. 43.10B ). They also have been described to have a fascicular or honeycomb appearance (i.e., echogenic structures with internal punctate, echo-poor spaces) ( Fig. 43.10C ). On longitudinal scan, the appearance of peripheral nerves has been likened to a “tram track”; that is, parallel hyperechoic lines are seen against a background of echo-poor space ( Fig. 43.10D ).

Tendons

Tendons appear to have numerous fine, parallel hyperechoic lines separated by fine hypoechoic lines (fibrillar pattern) on long-axis scans. Compared with nerves, tendons have more hyperechoic lines and move more than adjacent nerves when the corresponding muscle is contracted or passively stretched.

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