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Investigations and research Ultrasound-triggered image-guided therapy C. Moonen UMR 5231 CNRS/University Victor Segalen, Bordeaux, France. N. de Jong Erasmus Medical Centre, Department of Biomedical Engineering, Rotterdam, the Netherlands. O. Steinbach M. Böhmer S. Langereis Philips Research Europe, Biomolecular Engineering, High Tech Campus, Eindhoven, the Netherlands. The local delivery of therapeutic molecules mediated by ultrasound is a novel approach to addressing unmet clinical needs by providing a minimally-invasive platform for targeted delivery of pharmaceuticals [1]. It enables high concentrations of drugs to be deposited under image guidance to specific locations in the body. It can, therefore, be a preferred treatment option for non-systemic diseases. This produces more constant and controlled drug concentration profiles with favorable pharmacokinetics. E Ultrasound mediated delivery enables high concentrations of drugs to be deposited to specific locations in the body. With the advancement in imaging techniques, local drug delivery can now be applied in the human body with an unprecedented spatial and temporal resolution. Apart from delivering drugs to the desired location, another advantage of using ultrasound is the increase of cellular uptake through an effect called sonoporation, which gives the method an additional advantage over needle- or catheter-based procedures. Ultrasound for therapy is preferably highly focused, in order to expose a well-defined area, while steering of the ultrasound beam should be flexible and fast to allow for the controlled exposure of a region of interest with a potentially complex shape. Two types of ultrasound-mediated delivery systems can be distinguished: pressure-mediated delivery and temperature-mediated delivery. E Targeted delivery produces more constant and controlled drug concentration profiles with favorable pharmacokinetics. 40 MEDICAMUNDI 54/1 2010 Pressure-mediated delivery is performed using microbubbles, currently clinically applied as ultrasound contrast agents, that react to short ultrasound pulses by oscillation or bubble destruction [2]. As this is a fast process and the microbubbles and their destruction can be conveniently imaged using ultrasound, ultrasound imaging is the preferred modality for guiding the drug delivery. In the case of temperature-mediated delivery, in which ultrasound is applied for a sufficient amount of time to establish a local temperature increase, magnetic resonance imaging (MRI) is the most appropriate modality for image guidance, as it offers the possibility of therapy planning as well as monitoring of the local temperature distribution during the drug delivery procedure [3]. The first application of high intensity focused ultrasound (HIFU) in combination with MRI was in the ablation of uterine fibroids, which gives an excellent starting point for temperature monitoring during HIFU treatment. In this overview we present recent developments in ultrasound-triggered, image-guided drug delivery with a focus on pressure- and temperature-sensitive drug delivery vehicles. For temperature-induced delivery we present material systems, where drug and imaging agent are both encapsulated in temperature-sensitive carriers that allow the monitoring of temperature-sensitive delivery vehicles using MRI and disintegrate upon temperature increase. For pressure-sensitive microbubble agents we discuss the options of drug delivery with either co-injected drugs or drugs incorporated into the microbubble carrier, and focus on the induced permeability increase of cells. Finally an outlook on potential applications is given, ranging from the improvement in small-molecule drug delivery to gene-silencing using small interference RNA as a new therapeutic drug format. Temperature-triggered, MRIguided drug delivery During ultrasound energy deposition, the actual tissue temperature is difficult to predict. Physiological processes may modify local heat conduction and energy absorption. Blood flow may increase during temperature increase and thus change heat conduction. Unlike ablation treatments, the exposure to an elevated temperature for drug delivery using ultrasound should remain well below the damage threshold. In order to improve the therapeutic efficiency and the safety of the intervention, real-time mapping of temperature and thermal dose with MRI appear to offer the best strategy to optimize such interventions [4]. A particular advantage of MRI for guiding thermal procedures is that MRI not only allows for temperature mapping but can also be used for target definition. Although temperature effects on MR measured parameters (resonance frequency, relaxation, diffusion) had been well described previously, the first report of temperature mapping by MRI appeared in 1983 [5]. This method was based on the longitudinal relaxation time (T1). T1 based methods have been used in the clinic and preclinical applications, in particular at low magnetic field [6, 7]. Since then, alternative MR temperature imaging methods have been proposed. The most widely used MR temperature mapping is based on the temperature dependence of the water proton resonance frequency (PRF) [8-11]. The excellent linearity of the temperature dependency of the proton resonance frequency and its near-independence with respect to tissue type make the PRF-based method the method of choice, in particular at mid to high field strength. The principles of MR thermometry have been described in previous publications [3, 12]. As an indication, a precision of about 1°C can be reached with the PRF method at 1.5T for a spatial resolution of 2 mm, and a temporal resolution of 1s. In future applications, the incorporation of MRI contrast agents into temperature-sensitive drug vehicles will not only offer the possibility of guiding and controlling the temperature in the lesion, but can also be used to visualize and quantify the drug delivery process. However, MRI contrast agents have an effect on the relaxation phenomena, which are also temperature-dependent, and therefore complicate the quantification and reliability for monitoring the drug release. Some compromise must be made between sensitivity to co-release of such contrast agents and sensitivity to temperature change. Promising contrast agents to allow for temperature mapping and simultaneous following of the release are 1H CEST (Chemical Exchange Saturation Transfer) [13] and 19F MR contrast agents. Temperature-sensitive liposomes A drug delivery technology that has advanced to clinical trials for MRI-guided, temperatureinduced drug delivery is based on temperaturesensitive liposomes. Liposomes consist of a lipid bilayer encapsulating an aqueous interior containing the active molecules. The 1 G Figure 1. Temperature-sensitive liposomes containing co-encapsulated hydrophilic drugs and MRI contrast agents that are released at the phase transition temperature (Tm) of the lipid bilayer. 2 G Figure 2. Temperature-sensitive liposomal CEST and 19F MR contrast agents. Upon approaching the Tm of the lipid membrane the CEST effect vanishes, while the 19F NMR signal appears. temperature-sensitive liposomes release the encapsulated molecules at their phase transition temperature (Tm), where the lipid membrane displays a transition from a gel to a liquid crystalline phase. Studies with doxorubicin-loaded temperaturesensitive liposomes in combination with an externally applied regional temperature increase clearly showed the improved efficacy of temperature-induced drug delivery [14, 15]. Instead of relying on liposomal accumulation in the tumor, mild hyperthermia was applied during the first hour after injection of the temperaturesensitive liposomal formulation of doxorubicin. As a result, doxorubicin was rapidly released in the microvasculature of the tumor and subsequently taken up by the tumor cells. The co-encapsulation of hydrophilic drugs and MRI contrast agents is shown diagramatically in Figure 1. Temperature-sensitive liposomal 1 H CEST and 19F MR contrast agents have been explored as a potential carrier system for MRI guided drug delivery in combination with local MEDICAMUNDI 54/1 2010 41 hyperthermia induced by focused ultrasound [16] (Figure 2). In their lumen, these temperature-sensitive liposomes contain both a chemical shift agent (for 1H Chemical Exchange Saturation Transfer (CEST) detection) as well as a highly fluorinated compound (for 19F detection). The encapsulation of a chemical shift agent in the lumen of the temperature-sensitive liposome provides a pool of water protons in their aqueous core that can be selectively saturated using a narrow radio frequency pulse, and subsequently the saturation is transferred to the water of the bulk phase. E Pressure-mediated delivery implies an extension of the current application of microbubbles. E Microbubbles consist of gas encapsulated in a phospholipid, protein (albumin) or polymer shell. At temperatures below the Tm of the temperature-sensitive liposomal contrast agent, the chemical shift agent remains intraliposomal and the CEST effect is obtained, whereas the intensity of the simultaneously measured 19F NMR signal is negligible. Upon reaching the Tm, the CEST effect disappears, due to the release of the chemical shift agent. Simultaneously, the 19F MRI probe is freed from the influence of the paramagnetic shift agent, causing an appearance of a 19F MRI signal. The combined CEST and 19F MR temperaturesensitive liposomal carrier provides CEST MR based contrast enhancement, to localize the liposomes before release, while the 19F MR signal can potentially be used to quantify the local release of drugs with MRI. Pressure-sensitive systems An ultrasound device suitable for ultrasoundtriggered, ultrasound image-guided therapy preferably consists of a dual transducer set-up [17] comprising a focused transducer and an imaging transducer. The focused transducer is used to deliver a preset intensity for a precisely defined time at a well-defined location. Mostly the frequencies used range between 1-2 MHz, with an approximate focal zone of a 1-2 mm in diameter and 6-10 mm in length. The imaging transducer is then used to obtain a wider view around the targeted area and, for instance, to follow the inflow of microbubbles. The combination allows acoustic parameters such as frequency, pressure, and pulse length to be set independently for the therapeutic and imaging ultrasound transducers. Pressure-mediated delivery implies an extension of the current application of microbubbles. These microbubbles consist of gas encapsulated in a phospholipid, protein (albumin) or polymer shell. 42 MEDICAMUNDI 54/1 2010 Microbubbles with lipid shells and albumin shells are already used clinically. They are very sensitive ultrasound contrast agents, as only about 109 bubbles have to be injected, leading to efficient imaging of wall motion abnormalities and perfusion. However, these microbubbles disappear quickly from the blood stream. They are taken up by the liver and spleen and, especially, because they may contain a fraction of bubbles above 4-5 micron they may be trapped in the lung capillaries. The fluorinated gas used in these microbubbles is very hydrophobic, but gas exchange does take place in the lungs. Polymer-shelled microbubbles are easily prepared with a more narrow size distribution. This avoids significant accumulation in the lungs. Polymer shells are more rigid, leading to different acoustic characteristics from those of lipidshelled agents, governed by shell inhomogeneities and higher activation thresholds in terms of the acoustic pressure needed for destruction. If polymers with a hydrophobic moiety are chosen, liquid inflow into the capsules can be prevented and gas exchange in the lungs does not occur for these types of agents because fluorinated gases are not necessary [18]. Because of the typical size of microbubbles, less than 1 micron, liver and spleen accumulation is very difficult to avoid. The microbubbles are especially efficient in permeabilizing the vessel walls when exposed to ultrasound [19, 20]. This allows for delivery of material in tissue where normally no, or only limited, extravasation would take place. Examples can be found in cardiology, where local gene delivery has been established in preclinical models both for marker genes as well as for therapeutic genes [21]. In addition, in tumors, an increase of the delivery of model compounds in the tumor tissue has been demonstrated [22]. As the increased cell permeability caused by sonoporation remains for a limited time, having a high local drug concentration is advantageous. Several concepts have been developed to attach drugs to the bubble shell, to build them into the shell, or to put them in an additional liquid reservoir in the inside of the microbubble [23]. For incorporation in the shell, a polymeric microbubble with a relatively thick shell offers the highest volume for drug incorporation. However, drugs incorporated in the shell do not release quickly when the bubble is destroyed, as they tend to remain associated with the bubble fragments that are formed [24]. For this reason, systems in which drugs are present in an additional liquid reservoir form a more useful delivery vehicle, because half the volume of the microbubble or more can be used. High CEST F Figure 3. Three still frames of a Brandaris movie recorded at a frame rate of 10 Mfs. The bubbles were lying against the cell and there were insonified with 10 cycles of a frequency of 1 MHz. Left: before ultrasound, Middle: maximum expansion. Right: maximum compression 3 4 amounts can be incorporated, and when the bubble breaks the release is instantaneous [25]. However, this method is restricted to hydrophobic drugs. A more flexible method is to attach multilayers [26] or liposomes to the outside of microbubbles [27-29]. In this case, both lipid and polymeric microbubbles can be used and both hydrophilic and hydrophobic drugs can be incorporated. The delivery of doxorubicin, siRNA and pDNA has been established. Ultrafast cameras allow the microbubble behavior to be studied on the microsecond timescale, and their interaction with cells to be observed in detail [30]. A camera developed at Erasmus Medical Centre and Twente University uses the rotating mirror principle to project images onto 128 CCD cameras mounted over an angle of 70˚. This camera provides recording speeds up to 25 million frames per second (25 Mfs) which is well above the ultrasound frequency, so that microbubble behavior during an ultrasound pulse can be studied in detail. It has been found that lipid-shelled microbubbles can show both linear and strong non-linear behavior, which is useful for contrast-specific ultrasound imaging [31, 32]. Polymer-shelled microbubbles with a thin, usually inhomogeneous shell show reversible indentations in the shell before they actually break and release the encapsulated gas [25]. This indentation results in signals visible on an ultrasound scanner. The ultrafast camera has been used to investigate sonoporation as well. Van Wamel et al. [33] presented results that demonstrate that (endothelial) cell membranes follow the oscillations of adjacent microbubbles. The amplitude of the oscillations is on the micrometer scale and the time scale is in microseconds, so the movements in the cell membrane are very fast, in the order of m/s. The displacement of the cell caused by the expansion and compression of the bubbles around the cell has an amplitude of 2 μm, indicated by arrows in Figure 3. G Figure 4. Fluorescence images of propidium iodide uptake of the one cell surrounded with bubbles. The left frame is before US exposure. The right frame is the fluorescence image taken 15 s after US was turned off. The two circles indicate the positions of the microbubbles. This leads to an enhanced uptake of a marker dye as shown in Figure 4 (right frame). This shows the propidium iodide uptake recorded 15 seconds after the ultrasound has been switched off. Only cells with bubbles in close proximity showed uptake. After three minutes, no propidium iodide was present in the cell, indicating that the cell was still viable. This experiment clearly demonstrates that it is not necessary to induce large pores into membranes to change the porosity, and conditions exist at which cell membrane becomes only temporarily porous. In recent work, the importance of changes in calcium channels and the need for ATP in the membrane because of ultrasound and microbubbles has been highlighted [34, 35]. MEDICAMUNDI 54/1 2010 43 E The future of ultrasoundtriggered drug delivery will be determined by the combination of many factors. Therefore, not only the opening but also the closure of cell membranes after ultrasound treatment is a very relevant item in ultrasoundinduced drug delivery research. Summary and conclusions The first application of HIFU/MRI was in the ablation of uterine fibroids, which gives an excellent starting point for temperature monitoring during HIFU treatment. Further clinical trials are running for delivery of doxorubicin using temperature-sensitive liposomes (Thermodox™, Celsion). New options to improve image-guided drug delivery will need new imaging technology, material development and an expansion of the knowledge on mechanisms of uptake of drugs into cells, especially for high molecular weight therapeutics. These complex questions are currently addressed in public-private consortia such as the European Framework 7 project “Sonodrugs”. Specific attention needs to be paid to the choice of drug and its formulation. Small-molecule drugs, such as doxorubicin and paclitaxel, are well- established therapies. By depositing a larger dose at the region of interest, the efficacy of cancer therapy can be improved. Novel therapeutic formats, using DNA-based gene therapy or gene silencing by therapeutic siRNAs, need specific mechanisms to enter the cell, currently often by the use of viral vector. With sonoporation, a different way of entering the cell is introduced, avoiding potential adverse reactions and complex delivery systems such as viruses. 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