Piezoelectric Effect and Reverse Piezoelectric Effect in Ultrasound

This section explains how acoustic impedance differences determine ultrasound behavior at tissue boundaries. Key principles include: (1) Rayls serve as the standard unit for measuring acoustic impedance; (2) Large impedance differences between tissues cause strong reflection back toward the ultrasound machine, while similar impedances allow most energy transmission; (3) The magnitude of impedance difference, not direction, determines reflection strength; (4) High acoustic impedance indicates stiff, less compressible tissue; (5) When ultrasound encounters a highly compressible tissue with nothing to transfer energy onto, energy returns as the medium returns to rest. These principles form the foundation for understanding echo generation and image formation in diagnostic ultrasound.

This section covers the basic principles of how ultrasound waves interact with different tissues in the human body. Key concepts include: (1) Wave reflection occurs when ultrasound encounters boundaries between different media, producing both transmitted and reflected waves; (2) The intensity relationship states that incident intensity equals reflected plus transmitted intensity (I = Ir + It); (3) Ultrasound uses pulse transmission instead of continuous waves because transducers cannot emit and detect simultaneously, and pulses allow clear separation of emitted and reflected signals; (4) Acoustic impedance (Z = ρ × c) combines density and wave speed to characterize how waves travel through different materials.

When sound encounters a boundary between media with different acoustic impedances, part reflects and part transmits. The pressure reflection coefficient is: R = (Z₂ - Z₁)/(Z₂ + Z₁). When Z₂ < Z₁, reflection occurs with 180° phase inversion. The pressure transmission coefficient is: T = 2Z₂/(Z₁ + Z₂). Intensity coefficients are: R_I = R² and T_I = (4Z₁Z₂)/(Z₁ + Z₂)² × (Z₂/Z₁). At air-tissue interfaces, ~99% of energy reflects due to large impedance mismatch (Z_air ≈ 430 kg/m·s, Z_tissue ≈ 1.46 × 10⁶ kg/m·s). Ultrasound gel eliminates this mismatch for medical imaging. Sound refracts at interfaces according to Snell's Law: sin(θ₁)/v₁ = sin(θ₂)/v₂, enabling acoustic lenses for focusing sound. Sound attenuates exponentially in tissue: A = A₀ × e^(-αx), where α is the attenuation coefficient. The half-value layer (HVL) is the thickness reducing intensity by half: HVL = ln(2)/α. Different tissues have different HVLs, and higher frequencies experience greater attenuation.

Ultrasound imaging relies on wave physics principles. Specific acoustic impedance Z = ρ × c determines how waves interact with tissues. At boundaries with similar Z values, minimal reflection occurs; dissimilar Z values cause strong reflections. The reflection coefficient α = [(Z₁-Z₂)/(Z₁+Z₂)]² quantifies this effect. As waves penetrate tissues, intensity decays exponentially: I = I₀e^(-μx), where μ is the linear attenuation coefficient. Different tissues absorb ultrasound at different rates, affecting image penetration and quality. These principles explain why gel is used to match impedances and why certain tissues appear brighter or darker in images.

Acoustic impedance (Z) measures a material's resistance to sound wave movement, calculated as Z = ρ × v (density multiplied by velocity), typically expressed in units of 10⁶ kg/m·s. When sound waves encounter boundaries between different media, they partially reflect and partially transmit according to conservation of energy. The reflection coefficient R = [(Z₂ - Z₁)/(Z₂ + Z₁)]² determines how much energy reflects versus transmits. Large impedance differences cause high reflection, while similar impedances allow greater transmission. This fundamental principle governs how sound interacts with different tissues and materials in any acoustic system.

Ultrasound systems use three primary imaging modes: A-mode displays signals as amplitude spikes on a single line, useful for measuring cardiac chamber dimensions; B-mode uses multiple scan lines with amplitudes translated to gray shades, creating 2D cross-sectional images; M-mode combines B-mode with time-based display to track structure movement over time, essential for cardiac function assessment. Time Gain Compensation (TGC) corrects for beam attenuation by applying increasing amplification to signals from deeper regions, ensuring structures at different depths appear with similar brightness regardless of their actual echogenicity.

Ultrasound imaging uses three fundamental display modes: A-Mode (Amplitude Mode) is the oldest and simplest technique that displays ultrasound pulses as spikes on a one-dimensional graph, showing signal amplitude versus depth; B-Mode (Brightness Mode) converts echo amplitudes into brightness levels to create two-dimensional cross-sectional images with horizontal and vertical lines indicating tissue distance and structure; M-Mode (Motion Mode) is used for imaging moving structures like heart valves, providing high temporal resolution by displaying motion along a single ultrasound line over time.

Ultrasound imaging employs multiple modes: A-Mode (Amplitude Mode) provides 1D distance measurements using pulse-echo principle; B-Mode (Brightness Mode) creates 2D cross-sectional images for cardiac and organ assessment; M-Mode (Motion Mode) displays movement over time for cardiac chamber measurements. Doppler techniques visualize blood flow velocity: Continuous Wave Doppler uses separate transmit/receive crystals for continuous monitoring; Pulsed Wave Doppler uses the same crystal for depth-resolved measurements; Color Doppler overlays flow direction (red/blue) and velocity on 2D images; Power Doppler offers greater sensitivity for detecting low-velocity flow in solid organs.

Ultrasound imaging evolved from simple to sophisticated modes. A-mode displays echo intensity as a function of depth, showing brighter signals for more reflective structures. B-mode converts amplitude information into 2D brightness images, distinguishing calcium (bright) from soft tissue (dark). M-mode provides time-motion displays along a single scan line, showing how structures move during the cardiac cycle—stationary structures appear as horizontal lines while moving structures show characteristic patterns. These foundational modes established the basis for all subsequent ultrasound imaging techniques.

Three primary imaging modalities exist: B-mode converts wave amplitude to grayscale images (256 shades) for anatomical visualization; M-mode plots motion over time with depth on vertical axis and time horizontally, useful for cardiac and valve studies; Doppler measures blood flow using frequency shift—color Doppler shows flow direction with color coding, while spectral Doppler displays velocity as audible or graphic peaks. These modalities work together to provide comprehensive diagnostic information.

Therapeutic ultrasound uses focused sound waves to treat medical conditions through two primary mechanisms: thermal effects (heating tissue to destroy abnormal cells) and mechanical effects (cavitation-induced tissue disruption). High-intensity focused ultrasound (HIFU) can ablate tumors non-invasively by heating tissue to temperatures around 43°C, with the thermal dose measured in cumulative equivalent minutes (CEM) where 240 CEM typically causes cell death. The technology has been applied to treat kidney stones (shock wave lithotripsy since 1980), liver tumors, prostate cancer, and essential tremor in the brain. Ultrasound can also enhance drug delivery by using thermosensitive liposomes that release chemotherapy drugs when heated to specific temperatures, or by opening the blood-brain barrier using microbubbles. Challenges include acoustic refraction through tissues like fat and bone, which can distort the ultrasound beam and reduce treatment effectiveness.

Ultrasound has two main application categories: diagnostic imaging (ecography) and therapeutic applications. Diathermy applies high-frequency ultrasound (MHz range) to deep tissues for therapeutic heating, using non-focused crystals at moderate power (22W). High Intensity Focused Ultrasound (HIFU) represents advanced therapeutic technology using focused ultrasound energy to achieve temperatures of 55-100°C, enabling precise tissue ablation for treating tumors, metastases, and traumatic lesions. HIFU also enables sonoporation for drug delivery and clot disruption. These techniques demonstrate ultrasound's versatility in medical treatment beyond imaging.

In 1927, Wood and Lumis demonstrated that high-intensity ultrasound causes significant biological effects, including tissue damage and cavitation. This led to the development of therapeutic ultrasound, distinct from diagnostic imaging which uses low intensities. The ultrasound scalpel for neurosurgery (late 1940s) used focused ultrasound to precisely cut brain tissue without external incisions. Modern applications include tumor ablation, kidney stone fragmentation, eye surgery, and physiotherapy for heating internal tissues. Safety thresholds are critical: diagnostic ultrasound operates below heating thresholds (typically <1°C rise over 10 seconds), while therapeutic applications intentionally exceed these thresholds for tissue destruction.

Ultrasound has diverse medical applications including lithotripsy for breaking kidney stones (90% success rate), using focused shock waves to fragment stones for natural passage. Medical ultrasound uses different frequencies for different purposes: 2.5 MHz for abdominal imaging, 3.5 MHz for general obstetrics/gynecology, 5 MHz for vascular/breast/pelvic imaging, 7.5 MHz for breast/thyroid, 10 MHz for thyroid/blood vessels, and 15 MHz for musculoskeletal imaging. Higher frequencies provide better resolution but less penetration depth. Applications also include cleaning technology using cavitation bubbles, crack detection in pipelines and construction, and cardiac imaging.

High Intensity Focused Ultrasound (HIFU) is a non-invasive medical technology used for various biomedical applications including thermal ablation of tumors, transcranial surgery, and shock wave lithotripsy. The technique focuses ultrasound energy on a small focal zone where intensity reaches higher levels, enabling targeted treatment without incisions. Clinically, HIFU is applied for prostate cancer treatment (probe inserted into body), liver cancer treatment (patient positioned on MRI with probe reaching tumor region), and breast cancer treatment (patient placed in water tank with HIFU signals delivered from water tank). The focusing of ultrasonic signals is typically achieved using phase delay or focusing lenses on the transducer side, with vertically focused transducers using concave lenses being commonly employed.