Free ARDMS SPI Exam Flashcards

c = f × λ (propagation speed = frequency × wavelength). Frequency is fixed by the transducer and never changes with medium, so wavelength is forced to change whenever propagation speed changes — i.e., whenever sound enters a different tissue.

1540 m/s (1.54 mm/µs) — the average speed every ultrasound system assumes when calculating depth. Speed ranking from slowest to fastest: air (~330 m/s) < fat (~1450 m/s) < soft tissue (1540 m/s) < bone (~4080 m/s). Stiffness, not density, is the dominant driver of speed.

λ (mm) = 1.54 / f (MHz). Example: a 5 MHz transducer produces a wavelength of about 0.308 mm in soft tissue. This shortcut only applies at the assumed 1.54 mm/µs soft-tissue speed.

Period = 1/Frequency — the time required for one complete cycle. Period and frequency are reciprocals fixed by the source (transducer) and do not change when sound enters a new medium. Example: a 5 MHz transducer produces a 0.2 µs period.

dB = 10 log(I2/I1) — the logarithmic ratio comparing two intensities. A −3 dB change corresponds to a 50% intensity loss. Decibels compress the enormous linear range of echo amplitudes into a workable scale for gain, TGC, attenuation, and dynamic-range calculations.

≈0.5 dB/cm/MHz, one-way. Total attenuation (dB) = 0.5 × frequency (MHz) × path length (cm). Example: a 5 MHz beam traveling 4 cm one way loses 0.5 × 5 × 4 = 10 dB. Attenuation rises with both higher frequency and greater depth.

The depth at which beam intensity is reduced to exactly half its starting value (a −3 dB drop), also called the depth of penetration. Higher frequency → shorter HVL → less penetration; lower frequency → longer HVL → greater penetration.

Z = ρ × c (density × propagation speed), measured in rayls. Impedance depends on the product of density and speed together — the greater the impedance mismatch between two tissues at a boundary, the stronger the reflection produced there.

IRC = ((Z2 − Z1) / (Z2 + Z1))², for a beam striking a boundary perpendicular (90°). The remaining, non-reflected intensity transmits forward as ITC = 1 − IRC. A soft-tissue/air or soft-tissue/bone interface has a huge impedance mismatch and reflects nearly 100% of the beam.

sinθt / sinθi = c2 / c1. Refraction (beam bending) requires BOTH oblique incidence (angle other than 90°) AND a propagation-speed difference between the two media — if either condition is missing, the beam continues straight through.

Range = (c × t) / 2, where c = 1540 m/s (assumed) and t = total round-trip echo travel time. Dividing by 2 removes the double-counting of the outbound and return trip. Clinical shortcut: round-trip time ≈ 13 µs per cm of depth.

Pulse repetition period PRP = 1/PRF (typical imaging PRF ~1–10 kHz, set automatically from the selected depth). Duty factor DF = PD × PRF (= PD/PRP); imaging duty factor is very low (~0.1–1%), while continuous-wave Doppler has a duty factor of 100%.

Certain ceramics (lead zirconate titanate, PZT) generate a voltage when compressed and physically deform when a voltage is applied — this reciprocal property lets one element both transmit and receive. PZT is polarized above its Curie point during manufacture; reheating a finished element above the Curie point (e.g., autoclaving) permanently destroys its piezoelectric properties.

Resonant frequency = c(PZT) / (2 × element thickness), where c(PZT) is the speed of sound within the PZT ceramic itself. A thinner piezoelectric element produces a higher resonant frequency; a thicker element produces a lower frequency. This dimension is fixed at manufacture and cannot be reprogrammed.

NZL = D²/(4λ) = r²/λ, where D is the transducer element diameter, r is its radius, and λ is wavelength. A larger diameter or higher frequency (shorter λ) both lengthen the near zone. The natural beam waist (best unfocused lateral resolution) sits at the boundary between the near and far zones.

Axial resolution = SPL/2 = (# cycles × λ) / 2. Also called longitudinal, range, radial, or depth resolution (mnemonic LARRD). A smaller number is better; improves with higher transducer frequency or more damping (shorter pulse).

Lateral resolution = beam width (also called angular, transverse, or azimuthal resolution — mnemonic LATA). Best (narrowest) at the focal zone; degrades in the near and far field. Improves with focusing, higher frequency, and a larger active aperture.

Resolution in the dimension perpendicular to the scan plane, set by element height and the (usually fixed) elevational lens focus. It is typically the worst of the three resolution types; 1.5D arrays add electronic elevational focusing to improve it and reduce slice-thickness (partial-volume) artifact.

Frame rate falls as any of these increase: imaging depth, sector width, line density, or number of focal zones — each requires more pulses or more wait time per frame. Temporal resolution (frame rate) and spatial resolution (image detail) trade off directly against each other under a fixed PRF ceiling.

Output power (pulser/transmit control) sets the amplitude of the pulse sent INTO the patient — it raises intensity and bioeffect risk (TI/MI). Overall gain (receiver control) only amplifies echoes already received — it has zero effect on patient acoustic exposure. ALARA favors raising gain over output power to brighten an image.

A receiver function that applies progressively more amplification to echoes returning from greater depth, compensating for depth-dependent attenuation so identical reflectors at different depths appear equally bright. Unlike overall gain, TGC amplification is depth-specific, not uniform.

The ratio (in dB) of the largest to the smallest signal amplitude the system processes. Narrow/low dynamic range = high contrast (fewer gray shades, sharper black/white transitions); wide/high dynamic range = low contrast (more gray shades, smoother look).

Number of gray shades = 2ⁿ, where n = bits per pixel. Most clinical grayscale ultrasound systems use 8 bits per pixel, yielding 256 (2⁸) shades of gray — the standard clinical benchmark.

A-mode: 1D amplitude spikes plotted against depth, no gray scale (used today mainly in ophthalmic biometry). B-mode: brightness-encoded dots swept into the real-time 2D grayscale image. M-mode: a single scan line plotted against time — gives the best temporal resolution of any mode, used for rapid repetitive motion such as valve leaflets or fetal heart activity.

Transmits at a fundamental frequency but forms the image from echoes returning at the second harmonic (2× the fundamental), which is generated within tissue by nonlinear propagation — not at the reflector. Benefits: reduced near-field clutter and reverberation, reduced side-lobe/grating-lobe artifact, and improved lateral resolution.

Reverberation: sound bounces repeatedly between two strong parallel reflectors, producing equally spaced parallel bands of decreasing brightness. Comet tail: reverberation from tiny, closely spaced strong reflectors (metal, crystals) that merges into a solid tapering tail. Mirror image: a real structure near a strong curved reflector (classically the diaphragm) is duplicated on the far side of that reflector.

Acoustic shadowing: anechoic band deep to a structure that attenuates far more than TGC expects (calculus, bone, gas). Posterior enhancement: bright band deep to a low-attenuation structure (cyst, fluid). Refraction/edge shadowing: thin shadow at a curved margin where the beam bends per Snell's law. Side lobes and grating lobes: off-axis beam energy striking a strong reflector produces a spurious echo inside an otherwise anechoic structure.

3D = a single acquired volume, viewable from any plane after acquisition (via mechanical wobbler or 2D matrix array). 4D = real-time 3D — successive volumes update live, showing motion. Ultrasound contrast agents are intravascular gas microbubbles, highly nonlinear scatterers exploited by contrast-specific harmonic imaging for blood-pool/perfusion enhancement.

f_D = (2 × f_t × v × cosθ) / c, where f_t = transmit frequency, v = reflector (blood) velocity, θ = Doppler angle, c = 1540 m/s. The factor of 2 accounts for the round trip — relative motion affects the wave once when the moving reflector receives it, again when the transducer receives the returning echo.

cos 0° = 1 (maximum Doppler shift); cos 60° = 0.5 (shift reduced to 50% of maximum); cos 90° = 0 (no Doppler shift detected even with real flow present). Keep the Doppler angle at 60° or less for velocity measurements — cosine changes steeply above 60°, magnifying any angle-estimation error.

Flow toward the transducer compresses returning wave cycles → f_r > f_t → positive shift (displayed above the spectral baseline; red by the BART color convention). Flow away from the transducer stretches wave cycles → f_r < f_t → negative shift (below baseline; blue). No motion along the beam axis (90° angle) → zero shift, even with real flow present.

Uses two crystals — one continuously transmits, one continuously receives. No PRF means no Nyquist limit, so CW never aliases and has an unlimited maximum measurable velocity. Trade-off: range ambiguity — it cannot localize the depth a signal came from. Used for very high-velocity jets, such as severe or prosthetic valve stenosis.

Uses a single crystal that pulses like B-mode, timing returning echoes to place a sample volume (sample gate) at one selected depth — giving true range resolution that CW lacks. Because it is a sampled system, PW Doppler is subject to the Nyquist limit and can alias.

Nyquist limit = PRF/2 — the highest Doppler shift a pulsed system can measure without ambiguity. Deeper sample volumes require a longer round-trip time and therefore a lower achievable PRF, which lowers the Nyquist limit and makes aliasing more likely at the same true velocity.

Occurs whenever the true Doppler shift exceeds the Nyquist limit (PRF/2) — the pulsed system cannot sample fast enough and displays the signal wrapped to the opposite side of the baseline (spectral) or as an abrupt color reversal (color Doppler). Aliasing never occurs with CW Doppler, which has no PRF.

(1) Shift the baseline, (2) increase PRF/scale (directly raises the Nyquist limit), (3) decrease transmit frequency, (4) increase the Doppler angle toward 90° (up to the 60° working limit), (5) switch to CW Doppler (no PRF, no Nyquist limit, but loses range resolution).

Physiologic (true) broadening reflects genuine turbulence, such as post-stenotic flow. Technical (artifactual) broadening mimics turbulence from instrument settings — excessive spectral gain, an oversized sample gate, or sampling a curved vessel segment — even though the underlying flow is normal.

Rejects the low-frequency, high-amplitude signal produced by moving vessel walls and tissue that would otherwise clutter the spectral display. Set too high, it also removes real low-velocity flow (e.g., venous or diastolic flow), falsely suggesting absent flow — always use the lowest setting that still eliminates wall thump.

Color Doppler uses autocorrelation, a fast statistical technique, to estimate the mean Doppler frequency shift (not a full spectrum) at every pixel in real time. Convention: BART — Blue Away, Red Toward the transducer. Color scale sets the PRF, and therefore the Nyquist limit, for the color box, exactly as spectral scale does for the spectral gate.

Maps signal amplitude/power, proportional to the number of moving red blood cells, instead of mean frequency shift. Angle-independent and cannot alias; far more sensitive to slow/low flow than color Doppler, but shows no directional or velocity information.

Twinkling: rapid red/blue mosaic posterior to a rough, stationary reflector (classically a renal calculus) despite no true flow. Blooming: color signal overflows the true vessel wall from excessive color gain or transmit power. Flash: sudden color burst filling large field areas from tissue/transducer motion overwhelming the wall filter.

ΔP = 4v², where ΔP is the pressure gradient in mmHg and v is the peak velocity in m/s measured at or just downstream of a stenosis. Example: a 3 m/s peak velocity across a stenotic valve gives ΔP = 4 × 3² = 36 mmHg.

RI = (PSV − EDV) / PSV, using peak systolic velocity (PSV) and end-diastolic velocity (EDV). Unitless, ranges 0–1; a higher RI indicates higher downstream vascular resistance and less diastolic flow. No angle correction is needed — the angle term cancels out of the ratio.

PI = (PSV − EDV) / mean velocity over the cardiac cycle. Unlike RI, PI is not bounded at 1, making it more discriminating in very high-resistance vascular beds where RI would flatten near its ceiling. No angle correction is needed.

Laminar flow is orderly (narrow waveform, clear spectral window); turbulent flow is disorganized (spectral broadening), typically developing distal to a stenosis. Plug flow: nearly uniform velocity across the lumen, seen in large vessels near the heart. Parabolic flow: velocity is zero at the wall and maximal at the center, seen in smaller peripheral vessels.

A unitless, real-time estimate of the ratio of the acoustic power currently in use to the power predicted to raise tissue temperature by 1°C. Three application-specific versions: TIS (soft tissue), TIB (bone in the field, e.g., 2nd/3rd-trimester fetal imaging), and TIC (cranial bone, e.g., transcranial or neonatal head imaging).

MI = peak rarefactional pressure (MPa) / √(center frequency, MHz) — estimates cavitation risk; MI rises with pressure and falls as frequency increases. FDA Track 3 limits apply to nearly all diagnostic applications: I_SPTA ≤ 720 mW/cm² (≤50 mW/cm² for ophthalmic) and MI ≤ 1.9.

As Low As Reasonably Achievable — keep total acoustic energy exposure as low as possible while still obtaining a diagnostically useful image. Practiced by preferring receiver gain over output power to brighten images, limiting Doppler dwell time, and freezing the image whenever active scanning is not required.

QA phantoms are manufactured to propagate sound at 1540 m/s, matching the speed every scanner assumes internally. Core QA tests: depth of penetration (sensitivity), axial resolution, lateral resolution, dead zone, distance accuracy (calipers), and uniformity — performed at defined intervals and documented to catch equipment degradation before it affects patient care.

Sensitivity = TP/(TP+FN) — catches true disease; a negative result rules disease out. Specificity = TN/(TN+FP) — clears true health; a positive result rules disease in. PPV = TP/(TP+FP) and NPV = TN/(TN+FN) are both denominated by the test result and are prevalence-dependent, unlike sensitivity and specificity.