US wave is generated when an electrical current is applied to an array of piezoelectric crystals located on the transducer surface. Electrical stimulation causes mechanical distortion of the crystals (expansion and contraction) resulting in vibration and production of sound waves. The conversion of electrical to mechanical sound energy is called the piezoelectric effect. Each piezoelectric crystal produces an US wave, and their summation forms the US beam. US waves are generated in pulses (intermittent trains of pressure waves).

The propagation speed of sound varies in different biological media and is dependent on the acoustic impedance of the medium which is the resistance of a tissue to the passage of US. In most soft tissue, the speed average value is supposed to be constant at 1,540 m/s, and this is the value assumed by US machines for all biologic tissues [32, 33].

When the US beam crosses tissues, the amplitude of the original signal is attenuated as the depth of penetration increases. This loss of energy or attenuation is due to absorption in homogeneous media or reflection, refraction, and diffraction at tissue interfaces, e.g., between two tissues with different acoustic impedances (Fig. 5).

Absorption of US wave into a homogeneous medium corresponds to the conversion of acoustic energy to heat. The absorption coefficient is proportional to the square of the frequency, then to explore deep regions as abdomen, only low frequencies (3.5 MHz) are used.

Transmission and reflection occur on interfaces. When an incident US wave hit an interface, a part of US energy depending on the degree of impedance mismatch between the two tissues is reflected. The reflected US waves, called echoes, can be detected by the transducer and forms the basis of ultrasound imaging.

Refraction occurs when an incident wave hits the interface at an angle inferior to 90°; the penetrating pulse direction will be shifted by the refraction process. The transmitted wave propagates in a medium where the speed is not the same; the transmitted angle is different to the angle of incidence (Snell’s law).

Diffusion occurs when the incident wave encounters an interface that is not perfectly smooth (e.g., surface of visceral organs) or when the wavelength of the US wave is larger than the dimensions of the reflective structure (e.g., red blood cells). The reflected echoes scatter in many different directions resulting in signals of similar weak amplitudes.

The US pulses emitted from the transducer are transmitted into the body, reflected on tissue interfaces, and returned to the transducer (echoes). They are then converted into an electric signal which is processed and displayed as an image on the screen using a gray scale.

The degree of brightness, e.g., echogenicity, represents the amplitude of echoes which returns to the transducer (Fig. 6). There is a large difference between the acoustic impedance of bone and any soft tissues, leading to a hyperechoic image. Air and any soft tissue interface or steel interface also give a hyperechoic image. It is clinically important to apply a conducting gel (an acoustic coupling medium) on the transducer surface to minimize this difference.

A typical ultrasound machine consists of a transducer (probe used to send and receive the US waves by converting electrical energy to mechanical (ultrasound) energy and vice versa) and a pulse control unit to change the amplitude, the frequency, and the duration of the emitted pulses. The micro-ultrasound systems operate at frequencies of 30–100 MHz compared to frequencies of 2.5–15 MHz for clinical systems to improve spatial resolution.

Conventional US machine displays 2D images or slices of the body (2D B-mode—brightness mode) and 3D reconstruction. Four-dimensional (4D) ultrasound produces a moving picture version of 3D ultrasound (M-mode—motion mode). A Doppler ultrasound study (power Doppler mode to detect and map the microcirculation or color-flow Doppler mode to visualize the speed and direction of blood flow through the blood vessels [34–36]) may be part of an ultrasound examination.

The Doppler effect occurs when there is a moving source (blood flow of red blood cells) and a stationary receiver (US transducer). The relative motion between the sound source and the receiver changes echo frequencies, e.g., if the source is moving toward or away from the transducer, the perceived frequency is higher or lower than the actual one.

UBM is a noninvasive real-time technique that provides accurate and reliable images in embryos to adult mice of almost the organs observed in humans.

Contrast agents as air-filled microbubbles can be injected intravascularly as a bolus into tissue. These microbubbles have a highly hyperechoic appearance resulting in an increased contrast due to the high echogenicity difference between air and soft tissues. This technique is called contrast-enhanced US imaging (for review, see [37]).

X-ray CT is made possible by the relative transparency of X-rays through the body. As an X-ray beam crosses the biologic matter, it can (1) transmit through the matter without any interaction (primary or direct radiation), (2) interact with the matter and be completely absorbed by depositing its energy (photoelectric effect), and (3) interact and be scattered or deflected in a new direction and deposit a part of its energy (secondary radiation, Compton effect) (Fig. 8).

The total attenuation is the sum of the attenuation due to different types of X-ray interactions with the matter. It depends on the sample thickness, its density, its atomic number, and the X-ray energy (Beer’s law).

An X-ray image or projection represents an image of the sum of all local attenuations of an X-ray beam crossing the biological sample. To create a 2D or 3D tomographic image, a set of measured projection data need to be acquired. These projection views are acquired at different equally spaced angular positions around the rodent. A filtered back-projection algorithm [67, 70] is then used to reconstruct the distribution of X-ray attenuation in the slice plane. The reconstructed image is a spatial map of measured attenuation coefficients, e.g., each pixel in the reconstructed image is expressed as CT numbers in Hounsfield unit (HU) [67]. CT numbers are relative to water attenuation, which is assigned to a value of 0 HU. The highest CT number appears white on the CT image (e.g., in bones), and the lowest CT is black (for air); between these extremes are various shades of gray (Fig. 9).

The basic components of a micro-CT scanner are a micro-focus X-ray tube with focal spot diameter allowing high isotropic spatial resolution down to a range of 10–50 μm [68, 71] and a high-resolution X-ray detector, consisting of a charge-coupled device (CCD) array optically coupled to a phosphor screen via an optical lens or an optical fiber [68, 72, 73] or an X-ray image intensifier (XRII) screen coupled to a video readout [17].

The different projections are acquired by rotating synchronously the X-ray source and the detectors array around the animal using either a rotational bed device and stationary X-ray source and detectors (mainly used for in vitro imaging) or a rotational gantry and stationary sample design [21, 68, 71, 74].

Micro-CT imposes large ionizing radiation doses on the rodents for good image quality. Dose applied is crucial especially for multiple examinations of the same animal [75–78]. Whole-body or single-organ dose limitations are discussed in [67, 68].

Micro-CT can be used for noninvasive imaging of various organs such as skeleton, chest, abdominal organs, and brain [21].

Nuclei, composed of protons and neutrons, can become unstable when the number of neutrons is lower than the number of protons or when these numbers are too important. An unstable nucleus will lose its excess energy by the emission of particles. This phenomenon is called nuclear radiation.

There are three types of nuclear radiations:

The unstable nuclei used in nuclear imaging are obtained artificially with a particle accelerator, called a cyclotron, by hitting a target with accelerated charged elementary particles. These radioisotopes are then chemically fixed to a biological vector and can be injected in tissues to be imaged.

SPECT or scintigraphy consists in collimating with specialized submillimeter pinholes and detecting the γ radiations emitted by radioisotopes after their injection and fixation to the interest organs. For this purpose, γ detectors scan the body and change the number of γ photons detected to an electrical signal.

A γ detector (Fig. 14) is made up of a scintillator which transforms a γ photon into a visible photon and a photomultiplier which transforms a visible photon into an electron and multiply then the number of electrons. The scintillator is composed of a crystal; the detector sensitivity and the imaging system spatial resolution will depend on the type and the size of this crystal. Several research groups have developed high-resolution detectors using new materials like plastics or solid-state detectors [125, 126]. The photomultiplier is first formed by a layer alkaline which has the property of changing a visible photon into an electron. This electron enters into a vacuum chamber and is attracted by a first anode covered with alkali oxide, which produces several secondary electrons. These electrons are again attracted by a second anode and multiplied and so on until an electrical signal can be detected.

Dozens of γ detectors are associated in a network of resistors or capacitors to form the γ camera used in scintigraphy. This γ camera can be static and allows obtaining images of radioisotope concentration in an organ with a submillimeter resolution. It is then possible to acquire successive static images with a minimal temporal resolution of about 1 s to study radioisotope bio-distribution (the image intensity is proportional to the tracer concentration) and then quantify the molecular kinetic processes in which they participate [127]. The γ camera can also move along the body to obtain a larger image or a tomography. The tomography consists in the acquisition of a succession of static images with different angle orientations. The γ camera turns around the body with the imaged organ placed at the center of the rotation to keep the same distance between the γ photon source and the receptors. A matrix computation (analogue to the scanner X) provides axial, sagittal, and coronal slices as well as three-dimension volumes.

A PET scanner consists of one or more γ detector crowns placed around the body. When a β⁺ radioisotope is injected, the two γ photons resulting from the interaction of the positron emitted with the nearest electron are detected in coincidence by two γ detectors facing.

Before the radioisotope injection, an attenuation map is acquired using a γ source moving around the body. On the PET–CT systems, the γ source is replaced by the X-ray source. This map allows correcting the effects of attenuation or diffusion (when a γ photon is respectively stopped or deflected by tissues before achieving a γ detector).

Then, the signal from all the detectors is registered into a big matrix called a sinogram and processed. Firstly, the photons whose energy is different from 511 keV are removed. Secondly, the coincidence detection allows only virtually simultaneous detection (less than 15 ns difference) of two γ photons by two detectors facing. This determines the line of response (LOR) where the emission source should be located. Thirdly, the position of the source on the LOR is calculated by measuring the delay between the two detections (in fraction of nanosecond); this is the time of flight technique. Electronic and computer parts of a micro-PET scanner are adapted to these high-precision timing measurements (Fig. 15). The signal is also corrected taking into account the sensibility and the efficacy of each detector. Images of multiple slices or 3D volumes are then calculated using a mathematical transformation. Scanners now achieve 1 mm resolution and more than 3 % sensitivity [10].