There are limitations to the methods used to overcome the BBB for drug delivery. In some clinical trials, investigators have attempted to directly inject the drug into the targeted brain region using MRI and neuronavigation systems to position needles and pumps (Gross et al. 2011; White et al. 2010). Although this method of drug delivery is effective, the potential risks and undesirable surgical procedures prevent patients from receiving these treatments. Other methods to bypass the BBB include pharmacological modification of the drug targets in order to increase uptake through the barrier. One successful example of this technique is the addition of cysteine, a small neutral amino acid, to a non-transportable drug. This makes the drug appear to be a large amino acid, thereby allowing it to be taken up into the BBB (Killian et al. 2007). Although drug modification can be effective for bypassing the BBB, there are challenges in achieving therapeutic concentrations in the target region.

As an effort to address the limitations regarding drug delivery methods, researchers have attempted to find new ‘non-invasive’ methods to circumvent the BBB. Recent successes with intranasal delivery of therapeutic agents have gained attention. The highly vascularized nasal mucosa is adjacent to the cerebrospinal fluid (CSF), thus rapid blood concentrations can lead to therapeutic CSF drug concentrations. It has been suggested that drugs travel to the CSF via the olfactory neuroepithelium or via neuronal transport (Pires et al. 2009). Recent studies have shown that fluorescent tracers delivered intranasally may move through the perivascular spaces to reach the brain (Lochhead and Thorne 2012). These experiments have been promising in rodent models, but may be difficult to validate in humans where drug penetration throughout large brain regions is limited (Pardridge 2012).

Focused ultrasound (FUS) has been used to non-invasively and transiently open the BBB to allow drugs to pass from the blood stream into the brain (Hynynen et al. 2001). FUS has been effective as a method for delivering small and large therapeutic agents in preclinical disease models. The first clinical trials are rapidly approaching, and FUS has the potential to revolutionize drug delivery methods and be the first step to effective treatment of brain disorders.

The use of focused ultrasound in the brain has been investigated since the 1950s (Fry and Fry 1960). It was first observed that heating of brain tissue through application of continuous ultrasound could result in BBB opening, but this was often accompanied by hemorrhage (Bakay et al. 1956; Shealy and Crafts 1965) or tissue coagulation (Patrick et al. 1990). At the time, both thermal- and cavitation-related mechanisms were thought to contribute to BBB opening. Further studies have confirmed that thermal mechanisms can lead to increased drug uptake by endothelial cells (Cho et al. 2002), but that thermally-induced BBB opening has so far been accompanied by tissue damage (McDannold et al. 2004). As an alternative, cavitation-related BBB opening has been explored using pulsed ultrasound application. Cavitation refers to ultrasound-induced generation, oscillation, and in its extreme form, subsequent collapse of bubbles within a tissue. Similar to thermal mechanisms, in the 1980s and 1990s, cavitation-mediated ultrasound-induced BBB opening was unpredictable and often related to damage of the surrounding blood vessels and brain tissues (Vykhodtseva et al. 1995). Several parameters were investigated, and in general it was observed that tissue damage occurred more often as pulse duration, number of pulses and repetition frequency increased (Vykhodtseva et al. 1995; Mesiwala et al. 2002).

In 2001 it was demonstrated for the first time that preformed microbubbles, delivered intravenously, could be used as cavitation nuclei, thereby reducing the ultrasound energy required to induce BBB opening, producing much more consistent results (Hynynen et al. 2001). When the circulating microbubbles pass through the focal area of the ultrasound, they concentrate the acoustic energy inside the blood vessel. The ultrasound energy causes the microbubbles to expand and contract at the frequency of the ultrasound, a process known as stable cavitation (Fig. 16.2).

It is important to note the considerable differences between the high pressure generation and collapse of bubbles observed during the initial attempts at ultrasound-mediated BBB opening (Vykhodtseva et al. 1995), and then the stable oscillation of preformed bubbles that was introduced in 2001 (Hynynen et al. 2001). The addition of microbubbles has reduced the amount of ultrasound energy required to open the BBB by 100-fold. Since less energy is required, the risk of skull heating by the ultrasound is reduced, thereby making clinical treatments feasible. Also, at lower ultrasound pressures, the risk of brain tissue hemorrhage and damage is significantly reduced. Instead, the gently oscillating microbubbles interact with the endothelial cells and stimulate the opening of the tight junctions and increase transcellular molecular transport, leading to safe and effective drug delivery to the brain (Hynynen et al. 2001, 2006).

Following this first study, several groups have reported effective drug delivery using different ultrasound parameters in the presence of microbubbles. The factor that most significantly affects the amount of BBB opening is the ultrasound pressure. There is a delicate balance between applying enough ultrasound pressure to induce BBB opening for drug delivery, but limiting the ultrasound to avoid damage to the vessels and surrounding brain tissue. As pressure increases, the amount of microbubble expansion and contraction also increases, but too much pressure can lead to microbubble collapse and ensuing tissue damage. The relative amount of BBB opening required is based on the relative size of the therapeutic agent that needs to cross the BBB. Recent studies have demonstrated that greater ultrasound pressure is required to allow a 2000 kDa drug to cross the BBB, compared to the ultrasound pressure required for a 70 kDa drug (Chen and Konofagou 2014). The threshold of ultrasound pressure required for BBB opening is related to the mechanical index, defined as the peak negative pressure in-vivo by the square root of the frequency (McDannold et al. 2008a, b).

In rodents, ultrasound frequencies ranging from 28 kHz (Liu et al. 2010a, b) to 8 MHz (Bing et al. 2009) have been used to open the BBB. The size of the focal spot (or the treatment area) decreases with higher frequency, so conceptually, a higher frequency transducer would be better suited for targeting small brain nuclei. However, higher ultrasound pressures are required to induce opening when higher frequencies are used (McDannold et al. 2008a, b), therefore, the clinically relevant frequency range for therapeutic ultrasound is likely between 0.2 and 1.5 MHz. In addition to frequency, the duration of the ultrasound burst and the frequency of the burst repetitions have both been investigated to better understand their effects on BBB opening. Ultrasound burst length ranging from a few microseconds to a 100 ms have been reported, and in general BBB opening is greater as the burst length is increased, with no real benefit observed after 10 ms (McDannold et al. 2008a, b; Choi et al. 2011, O’Reilly et al. 2010). The repetition frequency of the bursts has not significantly impacted the amount of BBB opening (McDannold et al. 2008b), however it is understood that if the burst repetition frequency is too high, the microbubbles cannot reperfuse the area, thereby limiting the effectiveness of the low ultrasound energy (McDannold et al. 2007; Yang et al. 2008; Goertz et al. 2010; Weng et al. 2010; Choi et al. 2010; Vlachos et al. 2011). The microbubbles themselves can impact BBB opening. As microbubble size and dose are increased, there is greater BBB opening and potential for damage (Samiotaki et al. 2012; Wang et al. 2014).

Until recently, FUS-mediated BBB opening was monitored by post-treatment magnetic resonance imaging (MRI) and histological methods; however these methods were limited in their ability to provide feedback during FUS treatments. Over the past few years, several different groups have developed methods to detect and evaluate microbubble activity in-vivo (McDannold et al. 2006; Tung et al. 2011; O’Reilly and Hynynen 2012, 2013; Arvanitis et al. 2012). It was demonstrated that the frequency content emitted from the microbubbles during the ultrasound treatment can be used for online control of the exposures (O’Reilly and Hynynen 2012).

There are several advantages to FUS-induced BBB opening with microbubbles over other methods used.

FUS has proven to be an effective method for transiently opening the BBB and allowing therapeutic agents to cross into the brain parenchyma. Yet, the physical and cellular mechanisms by which the barrier is opened are still unknown. It has been shown that at low ultrasound pressures, microbubble oscillations put stress on the vascular endothelium (Hosseinkhah and Hynynen 2012), which responds to dynamic shear stress (Krizanac-Bengez et al. 2004). Although both the expansion and the contraction phases of the bubble under ultrasound can impact the blood vessel wall, recent evidence has indicated that the majority of changes occur during the contraction phase (Chen et al. 2011). Bubble contraction exerts forces that pull the vessel wall towards the interior of the lumen. There are also circumferential stresses and radiation forces that are generated by an oscillating microbubble that may activate the mechanosensitive ion channels in the vascular endothelium (Traub et al. 1999).

At higher pressure, microbubble oscillation can turn into violent collapse, a process known as inertial cavitation. It has been suggested that inertial cavitation could be a potential mechanism for BBB opening, but due to the extreme local temperature rise and microjetting, it is suspected that inertial cavitation is linked to red blood cell extravasation and more serious vessel damage (Nyborg 2001; Deng et al. 2004; Vykhodtseva et al. 2008). Recently developed methods, which monitor microbubble acoustic emissions during ultrasound exposure, have determined that BBB opening can occur in the absence of inertial cavitation. Analysis of the acoustic emissions demonstrated rises in subharmonics, harmonics and ultra harmonics due to bubble oscillation. These harmonic changes are correlated to BBB opening (McDannold et al. 2008b; O’Reilly and Hynynen 2012). The acoustic emissions observed during inertial cavitation show increased broadband emission, and these signals are correlated experimentally with red blood cell extravasation and damage, thus confirming that BBB opening is not due to inertial cavitation.

Less is known about the cellular mechanisms of FUS. Initial studies delivered labeled endogenous IgG and tracked its entry through the BBB following FUS in the presence of microbubbles (Sheikov et al. 2004, 2008). The studies suggested that antibodies passed through the BBB using both the paracellular and transcellular delivery routes.

Firstly, the paracellular space was increased in the FUS application region. This widening of the tight junctions was accompanied by a reduction in the quantity of tight junction proteins (Sheikov et al. 2004, 2008), suggesting increased paracellular transport following FUS application. Specifically, there was a reduction in key tight junction proteins (occludin, claudin-5 and zonula occludens) between 1 and 2 h after FUS application, but these levels were returned to normal within 4 h (Sheikov et al. 2004). These studies have been supported by research using oxidants to open the BBB, in which there was a consistent and similar reduction in occludin and zonula occludens (Musch et al. 2006). With respect to FUS, it has been hypothesized that FUS induces either protein reorganization in the tight junction which masks the antigens, or that the proteins are downregulated (Sheikov et al. 2008). It is known that FUS can downregulate other proteins, such as the gap junction protein Connexin 43 (Alonso et al. 2011). The change in gap junction proteins was suggested to buffer the hyperexcitability of the brain, which may occur after BBB opening (Alonso et al. 2011).

Secondly, there was labeled antibody present inside vesicles within the endothelial cells. Along with an increase in the number of cytoplasmic channels and vesicles in FUS-treated regions, it was evident that antibody was being taken up by the endothelial cells and passed through the BBB transcellularly (Sheikov et al. 2004, 2008). Related studies have demonstrated both in-vitro and in-vivo that caveolin, a membrane protein involved in receptor independent endocytosis, is upregulated after FUS (Lionetti et al. 2009; Deng et al. 2012). Uptake of the drug into the cells may occur during the deformation of the endothelial cells due to the shear stress generation by the microbubble in the ultrasound field (van Wamel et al. 2006; Meijering et al. 2009).

In-vivo, two-photon microscopy has been used to visualize FUS-mediated BBB opening. Two studies have attempted to fit the leakage observed during imaging sessions to the cellular mechanisms of FUS (Raymond et al. 2008; Cho et al. 2011). In these studies, fluorescent dye (10–70 kDa) was injected intravenously to visualize the vasculature and to observe the leakage kinetics as the dye passed through the BBB. Two different patterns of dye leakage from a single vessel were observed. Fast leakage was characterized by the dye rapidly crossing across the BBB within the first 5 min after FUS application, as well as appearing to arise from a point source along the vessel wall (Cho et al. 2011). Conversely, slow leakage began 5–15 min after the onset of FUS and occurred along the length of the vessel (Cho et al. 2011). The authors speculated that these leakage types corresponded to the different cellular mechanisms of FUS. Fast leakage was hypothesized to occur when tight junction proteins were downregulated, leading to tight junction widening and ‘fast’ leakage of dye through the BBB. Slow leakage was hypothesized to occur when the endothelial cells were stimulated to uptake dye and transfer it through the BBB. A next step will be to correlate two-photon microscopy with acoustic emission analysis to understand the relationship between the dye leakage type and the microbubble emissions.

Over 100 primary research studies have used FUS to deliver tracers or therapeutic agents through the BBB and into the brain. Evaluation of BBB opening and quantification of drug delivery is done by delivery of a contrast agent or tracer molecule. Most commonly, gadolinium-based MRI contrast agents (500–900 Da) are delivered during FUS treatment, and the extravasation of these agents into the brain is used to confirm and evaluate BBB opening using MRI (Hynynen et al. 2001). MRI images are analyzed and the hyperintense regions, indicative of contrast agent delivery, are measured and used as an indicator of the ‘amount’ of BBB opening (Treat et al. 2007). It has been demonstrated that the relative enhancement of the hyperintense regions is correlated to the amount of vessel damage that occurs in the brain (Hynynen et al. 2001; McDannold et al. 2008a, b; O’Reilly and Hynynen 2012). Within the safe range of BBB opening (enhancement <30 %), there is the clear advantage of more drug delivery when the BBB is opened to a greater extent. This is particularly true when delivering large molecules. As an example, Jordao et al. measured the amount of endogenous IgM, an antibody that weighs 6 times more than the more abundant IgG. They demonstrated that the percentage increase in IgM in the brain after FUS-mediated BBB opening is positively correlated to relative amount of BBB opening, quantified by increased enhancement on MR images (Jordão et al. 2013). There is room for further experimentation in this area to precisely optimize the amount of BBB opening during each treatment with respect to the drug that is being delivered. It is understood that larger drugs require more BBB opening to reach therapeutic concentrations in the brain, and balancing optimal drug delivery with the risk of vessel damage will be important (Yang et al. 2011; Nhan et al. 2013; Chen and Konofagou 2014). Many other studies have supported these findings and have validated the use of MRI, not only as a method to choose the target location, but also as a method to evaluate drug delivery and success of the treatment (Kinoshita et al. 2006; Treat et al. 2007; Liu et al. 2010b; Aryal et al. 2013; Fan et al. 2013a). It is also possible to combine the therapeutic agents with MRI contrast agent molecules, and thus directly quantify the amount of drug delivered (Ting et al. 2012; Fan et al. 2013a, b).