## 1 Introduction  
In recent years, novel magnetic resonance imaging (MRI) techniques using hyperpolarized (HP) gases as imaging agents for in-vivo imaging of human lungs have been developed [1]. They successfully complement the standard proton imaging and computed tomography (CT) in providing high resolution structural information, on which very sensitive diagnostic methods for lung diseases such as asthma, COPD (chronic obstructive pulmonary disease), and cystic fibrosis are based [2]. Moreover, dynamic ventilation studies that are impractical using other techniques can be performed [3]. Contrary to the standard MRI that utilizes the thermal polarization of water protons in high magnetic field, the hyperpolarization of noble gases is obtained by optical methods. The techniques are described in detail both for \({}^{3}\)He [4] and \({}^{129}\)Xe [5]. There are several specific requirements for their application in the MR imaging that take into account the relaxation of unrecoverable polarization during storage and delivery [6]. A careful choice of construction materials for the gas container with all connecting tubes is necessary to minimize relaxation losses. Moreover, the magnetic field strength and homogeneity along the transferring route from the gas container to the patient has to be optimized. Additionally, due to recent dramatic rise of \({}^{3}\)He cost, it is important to collect the exhaled gas for recycling [7]. For the method to be reliable and reproducible, it is necessary to employ a dedicated system for monitoring and controlling the patient's breath cycle and dosing hyperpolarized \({}^{3}\)He. It should be compatible with the MRI scanner, which utilizes a high field magnet that is located in the Faraday cage to eliminate any RF interference. Therefore, the components of the ventilator that must be inside the Faraday cage have to be made of non-magnetic material and should not be controlled by the electrical AC power. Additionally, the system should provide a triggering circuitry to synchronize its operation with the image acquisition sequence. The first MRI compatible ventilators were developed for animal studies applications [8; 9]. Since the cooperation of an animal is impossible, the ventilators were equipped with the gaseous anesthesia unit, as well as the breath and heart rate measuring devices for triggering the MR imaging sequence [10]. In the case of human lung studies, full cooperation of the patient with the medical scanner operator is usually feasible, so that a fast MRI sequence can be applied during the breath hold [11]. However, in the case of children, critical care patients, or those requiring permanent mechanical ventilation, the breath control and the sequence triggering may be still necessary to obtain good image quality. For such applications, the standard ventilators were modified by simply extending the gas delivery tubes that connect the main unit located outside the Faraday cage to the patient [12]. More advanced devices were also built, in which only non-magnetic components were used [13; 14]. However, all above systems could not handle the hyperpolarized gases. The first in-vivo MR images of the rat lungs were obtained in 1994 [15], using hyperpolarized \({}^{129}\)Xe as a gaseous imaging agent inhaled by the animal. Subsequently, in-vivo MR images of human lungs were acquired in 1996, with the application of hyperpolarized \({}^{3}\)He [16; 17]. These results opened new possibilities for the application of the technique in medical diagnostics, and stimulated the development of dedicated ventilators with the noble gas administration capability. A detailed description of a device modified for small animal studies was given in [18; 19; 20]. It was followed by a fully scalable system that could be used for humans [21], and a sophisticated ventilation system dedicated to MRI of human lungs with the use hyperpolarized noble gases [22].