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Theranostics 2014; 4(4):432-444. doi:10.7150/thno.8074 This issue Cite

Review

Combining Microbubbles and Ultrasound for Drug Delivery to Brain Tumors: Current Progress and Overview

Hao-Li Liu¹, Ching-Hsiang Fan², Chien-Yu Ting², Chih-Kuang Yeh^2,

1. Department of Electrical Engineering, Chang-Gung University, 259 Wen-Hwa 1st Road, Kuei-Shan, Tao-Yuan, Taiwan 33302
2. Department of Biomedical Engineering and Environmental Sciences, National Tsing Hua University, No. 101, Section 2, Kuang-Fu Road, Hsinchu, Taiwan 30013

✉ Corresponding author: Department of Biomedical Engineering and Environmental Sciences, National Tsing Hua University, No. 101, Section 2, Kuang-Fu Road, Hsinchu, Taiwan 30013. Tel: +886-3-571-5131 ext. 34240; Fax: +886-3-571-8649. E-mail: ckyehnthu.edu.tw More

Abstract

Malignant glioma is one of the most challenging central nervous system (CNS) diseases, which is typically associated with high rates of recurrence and mortality. Current surgical debulking combined with radiation or chemotherapy has failed to control tumor progression or improve glioma patient survival. Microbubbles (MBs) originally serve as contrast agents in diagnostic ultrasound but have recently attracted considerable attention for therapeutic application in enhancing blood-tissue permeability for drug delivery. MB-facilitated focused ultrasound (FUS) has already been confirmed to enhance CNS-blood permeability by temporally opening the blood-brain barrier (BBB), thus has potential to enhance delivery of various kinds of therapeutic agents into brain tumors. Here we review the current preclinical studies which demonstrate the reports by using FUS with MB-facilitated drug delivery technology in brain tumor treatment. In addition, we review newly developed multifunctional theranostic MBs for FUS-induced BBB opening for brain tumor therapy.

Keywords: Microbubbles, brain tumor, focused ultrasound, blood-brain barrier, chemotherapy

1. Brain Tumors and the Blood-Brain Barrier

1.1 Current status of brain glioma treatment

Nearly 20,000 patients in the U.S. are diagnosed annually with primary malignant brain cancers, among which more than 50% are reported as glioblastoma multiforme (GBM), making this the most common malignant brain cancer in adults, and it is responsible for half of cancer patients' deaths [1]. Gliomas can be divided into astrocytic and oligodendroglial tumor types, and are classified as grades I to IV based on the tumor growth rate. The median survival times are reportedly 5-15 years and 9-12 months in patients with low- and high-grade gliomas, respectively [2, 3].

GBM patients first undergo debulking surgery to remove most of the tumor mass, followed by chemotherapy and/or radiation therapy. A phase-III randomized trial found that the prognosis of GBM patients remains poor after debulking surgery and radiation, with a median survival time of only 12 months [4]. Chemotherapy is considered to be an important treatment modality for malignant brain tumors [5]. The most common adjuvant chemotherapy drugs to be administered systemically are carmustine (also called BCNU), PCV (comprising procarbazine, lomustine, and vincristine), and the first-line chemotherapeutic agent temozolomide (TMZ). However, such chemotherapeutic drugs provide a limited and short-duration response, typically resulting in disease stabilization that lengthens survival by only a few months. In addition, the side effects of chemotherapy negatively impact the already poor quality of life during the patient's remaining life span. Up to date no specific drug or regimen has shown superior efficacy in GBM treatment, and delivery methods of chemotherapeutic agents to treat GBM have met with limited success. The dilemma is that while increasing the concentration at which a chemotherapeutic agent is administered in an attempt to increase the dose delivered to the tumor may improve the treatment outcome, this is likely to also result in substantial systemic toxicity [6]. There is therefore an urgent need to develop techniques for delivering a chemotherapeutic agent into the central nervous system (CNS) so that it reaches a sufficiently high therapeutic dose in the targeted brain tumor region while minimizing its systemic concentration.

1.2 The blood-brain barrier in brain tumors

The blood-brain barrier (BBB) is a highly specialized structure in CNS blood vessels and capillaries that comprises arachnoid membranes, cerebral capillary endothelial cells, and the choroid plexus epithelium. These layered cell structures construct the so-called tight junctions also known as the zonula occludens (Fig. 1) [7]. The tight junctions of the cerebrovascular endothelium contain membrane-associated guanylate kinases such as ZO-1 and ZO-2, cadherins (single-pass membrane-spanning molecules), and cingulin [8]. In addition to tight junctions, the low endocytic activity and absence of fenestrations also limit transcellular transport. By forming an almost impermeable barrier to the diffusion of large (>200 kDa) and hydrophobic molecules, the BBB not only protects the normal brain parenchyma from foreign toxic substances, but also blocks the delivery of many potentially effective diagnostic or therapeutic agents in cases of CNS disease [9]. The normal physiological function of the BBB therefore needs to be temporarily disrupted to allow the diffusion or local delivery of macromolecular therapeutic or diagnostic agents into the brain.

The integrity of the BBB is typically highly heterogeneous within tumor tissue, resulting in highly variable BBB permeability within different areas of the same tumor. Brain tumors are usually most permeable in the tumor core whereas the BBB remains relatively intact at the tumor peripheral regions [10]. The BBB of the peripheral glioma has been shown to remain highly functional [11-13], and previous clinical studies have demonstrated that brain tumor cells can migrate great distances from the enhancing regions of the tumors [14, 15]. The intact BBB of tumor-infiltrating regions (mostly the tumor periphery) severely restricts treatment efficacy and is a critical factor in the high rate of GBM recurrence. For this reason, enhancing the BBB permeability of the tumor periphery represents an important potential strategy for improving treatment efficacy.

1.3 BBB-disrupting strategies for enhanced drug delivery to brain tumors

Several clinical and preclinical methods exist for delivering chemotherapeutic agents for GBM treatment. Intravenously (IV) administered agents reach the brain tissues from either the blood or cerebrospinal fluid (CSF) after penetrating the BBB or blood-CSF barrier.

Interstitial delivery involves either IV injection [16] or the implantation of biodegradable wafer-containing drugs attached inside surgically removed brain-tumor cavities [17, 18] to bypass the impermeable BBB. In a 240-patient clinical trial the median survival time was 14 months for implantation of a BCNU wafer compared to about 12 months for placebo [18]. Similar results were obtained for Gliadel^®, but its use was associated with various adverse effects, including intracranial hypertension, CSF leakage, brain edema, seizures, intracranial infection, and abnormal healing [19].

 Fig 1 

The blood-brain barrier prevents diffusion of harmful molecules as well as the delivery of therapeutic drugs to the brain.

High-concentration chemotherapeutic agent administration obtained by interstitial infusion (or so-called convection-enhanced delivery) has also been reported, where a pressure gradient is used to generate continuous bulk fluid flow through the brain interstitium [20]. The distribution of the chemotherapeutic agent depends on its molecular weight, total concentration, and polarity, as well as the total volume infused and the rate of infusion. Current challenges include heterogeneous drug distributions and high or varying tumor interstitial fluid pressures, which can lead to faster drug efflux out of the injection site and thereby significantly degrade the efficacy.

BBB permeability can also be increased by intra-arterial osmotic agents or hypertonic solutions [11, 21, 22]. Phase-II patient trials have demonstrated that therapeutic outcomes are enhanced when the BBB permeability is higher (expected survival of 17.5 vs. 11.4 months). However, a major obstacle of using osmotic pressure to increase BBB permeability is the lack of specific targeting. Complications include neurological deficits, syndromes similar to stroke, possible migration of tumor cells, temporal seizures, and new tumor-nodule formation at distant brain locations [23]. Current strategies for enhanced drug delivery have therefore been limited by their invasiveness and/or lack of specific targeting.

2. Current Status of Microbubbles and their use to Enhance Drug Delivery to the Brain

2.1 Microbubbles as ultrasound contrast agents

The use of microbubbles (MBs) in echocardiography was first reported in 1968 [24]. Due to the absence of shell structures, these MBs had short half-lives (within a few seconds) that limited their clinical applications. Updated MB designs have led to higher stability through increased molecular weight, low solubility, incorporation of a low-diffusivity gas such as nitrogen or perfluorocarbon, and use of a biodegradable shell material such as albumin, phospholipids, or polymers [25, 26]. MBs are highly echogenic in vivo due to the mismatch in acoustic impedance (i.e., the product of density and speed of sound) between their gas cores and surrounding tissues [27]. IV administered MBs are capable of increasing the intensity of backscattered ultrasound by up to 20-30 dB [28], therefore serving as excellent ultrasound imaging contrast agents. MBs are currently applied in routine clinical examinations including organ perfusion and enhanced diagnosis in highly vascularized tumor structures [25, 29], and in diagnoses of cardiovascular and renal diseases [30, 31].

At present, three commercial MB agents—Optison^TM (GE Healthcare, WI, USA), Definity^® (Lantheus Medical Imaging, MA, USA), and SonoVue^® (Bracco, Milano, Italy)—are licensed for clinical diagnostic applications (Table 1). The commercial MBs in these agents are typically larger than 1 μm and have imaging durations within the range of 5-10 min.

2.2 Therapeutic applications of MBs

In addition to their contrast-enhancing ability for diagnostic applications, MBs also possess unique properties for therapeutic applications [32-34]. MBs excited by ultrasound are capable of physically interacting with surrounding media through stable or inertial cavitation (Fig. 2). Stable cavitation generated by ultrasound excitation causes repetitive MB contraction and expansions that induce the flow of liquid around the MBs (called microstreaming) that in turn applies shear stress to cell membranes, leading to ion channel/receptor modulation and therefore affects on cell permeability [35]. On the other hand, in inertial cavitation the excessive ultrasound pressure causes abrupt MBs destruction or collapse to produce strong mechanical stress (i.e., shock waves), microstreaming, microjets [36], or even free radical production [37] in the surrounding media. Shock waves and micro-jets are strong forces that cause cell membrane perforation and even blood vessel permeabilization [38, 39]. Recently, the use of MBs in conjunction with non-thermal pulsed-mode ultrasound has been investigated to enhance blood-tissue drug permeability for therapeutic applications [40-43].

 Fig 2 

Physical mechanisms underlying the biological effects induced when microbubbles are excited by ultrasound energy.

3. MB-facilitated Focused-Ultrasound-Induced BBB Opening

3.1 Concepts of MB-facilitated focused-ultrasound-induced BBB opening for CNS drug delivery

Treatment with non-thermal burst-mode ultrasound in the presence of MBs has been confirmed to be able to induce local and reversible BBB opening due to the disruption of tight junctions in CNS capillaries (Fig. 2) [44-46]. Noteworthy, the size and resonance frequency of MBs (in general, the smaller the size of MBs, and the higher the resonance frequency of MBs) were reported to be critical in MB-facilitated focused-ultrasound (FUS)-induced (MB-FUS) BBB opening. Larger MBs (such as SonoVue^TM) allow sufficient size expansion to stimulate vessel walls, and thereby BBB opening can be achieved with a lower acoustic pressure. When MBs collapse inside a blood vessel, smaller ones (< 2 µm) would be likely to be fragmented at some distance from the endothelial wall, whereas larger MBs would expand and fragment while in contacting with endothelial walls, and may increase the occurrence of endothelial cell destruction, capillary lumen damage, or erythrocyte extravasations due to inertial cavitation [47, 48]

The abilities of several types of molecules to penetrate the intact BBB upon MB-FUS BBB opening in preclinical settings have been investigated. Trypan blue (872 Da) and Evans blue (960 Da) are dyes that are widely used to identify the BBB opening region [49, 50]. In addition, gadolinium-based contrast agents such as Gd-DTPA (573-928 Da) used in magnetic resonance imaging (MRI) imaging are commonly administered to monitor the location and efficacy of BBB opening [44-46].

Various other imaging tracers have been delivered across the BBB, including horseradish peroxidase (40 kDa) [45, 51], monocrystalline iron oxide nanoparticles (20 nm, 10 kDa) [51], Alexa Fluor 488 (10 kDa) [52], Texas-Red-tagged dextran (3-70 kDa) [47, 54-56], lanthanum chloride (139 Da) [57], 99mTc diethylenetriamine pentaacetate (492 Da) [58, 59], superparamagnetic iron oxide (SPIO, 60 nm) [60, 61], ionic manganese [62], and gold nanorods [63].

Compared to alternative brain drug delivery strategies such as hypertonic infusion [23] and the use of modified lipophilic chemicals, MB-FUS is an entirely noninvasive procedure capable of local rather than systemic BBB disruption, thus minimizing off-target effects. The reversibility of this BBB-disrupting method provides a time window of several hours for drug release, allowing enhanced permeability and retention of the drug specifically in the tumor region [64-66]. MB-FUS thus provides an attractive alternative for elevating the local concentration of chemotherapeutic agents for the treatment of CNS diseases.

3.2 MB-FUS BBB opening for delivery of therapeutic substances to the brain

Therapeutic molecules have been shown to penetrate the intact BBB of the normal brain upon MB-FUS BBB opening in preclinical settings. Both Herceptin (150 kDa) and D4 receptor antibodies (150 kDa) have been successfully delivered to mouse brains [67]. Doxorubicin (DOX; 543 Da) has been delivered into the normal rat brain via MB-FUS BBB opening [64], and the amount of methotrexate (545 Da) delivered to the brain was significantly higher for FUS-BBB disruption than for an intra-carotid injection [68]. Therapeutic anti-amyloid-β antibodies (150 kDa) which were applied to remove Aβ plagues for immunotherapy in Alzheimer's disease treatment have been successfully delivered to the brain [69, 70]. Macromolecules such as magnetic nanoparticles (comprising 60-nm SPIO) have also been successfully delivered into the brains of both small and large animals via BBB opening [60, 71, 72]. Polyethylene-glycol-coated gold nanoparticles (10-20 nm) have recently been delivered into the normal rat brain, which could have potential photothermal/photoacoustic imaging applications for combined cancer treatment and diagnosis [63]. Enhanced delivery of boronophenylalanine which has high thermal neutron capture cross-section for boron neutron-capture therapy (BNCT) has been achieved via MB-FUS BBB opening, indicating that this technique has potential in increasing the treatment efficiency of BNCT [73, 74]. Small interfering RNA (13 kDa) has been non-invasively delivered into the striatum to modulate the expression of mutant Huntingtin protein [61, 75]. More recently, stem cells were also delivered to brain tissues using a BBB opening technique [76, 77].

The above studies have demonstrated a wide variety of possible applications of MB-FUS BBB opening for CNS disease treatment; however, the application to brain tumor treatments constitutes the largest proportion (>25%) of preclinical investigations using this approach. The following section provides a comprehensive overview of the current progress of MB-FUS BBB opening applied in brain tumor animal models to confirm the therapeutic efficacy associated with the enhanced delivery of chemotherapeutic agents.

4. MB-FUS BBB Opening for Brain Tumor Therapy

The efficacy of brain tumor treatment with enhanced delivery of chemotherapeutic drugs by MB-FUS BBB opening has been evaluated in a preclinical setting with liposomal DOX, BCNU, and TMZ, as described in Sections 4.1-4.3, respectively.

4.1 MB-FUS BBB opening to enhance liposomal DOX delivery

The earliest attempt at FUS-enhanced delivery for glioma treatment involved DOX in the form of Doxil^® (Ben Venue Laboratories, OH, USA), which was encapsulated in long-circulating pegylated liposomes [64, 78, 79] (frequency = 1.7 MHz, acoustic pressure = 1.2 MPa, burst length = 10 ms, pulse repetition frequency (PRF) = 1 Hz, and sonication duration = 60-120 s). Therapeutic efficacy was evaluated in rats implanted with 9L-glioma-cell tumors by comparing population survival curves among different treatment groups. Longitudinal MRI used to follow the tumor progression revealed that tumor growth in rats with implanted glioma after treatment with MB-FUS+Doxil was delayed compared to control-tumor and Doxil-alone tumor animal groups; the last group showed continued tumor growth after treatment. Median survival was improved 16% relative to control animals by Doxil-alone treatment, and was further improved to 24% by additional MB-FUS treatment (control, 25 days; FUS alone, 25 days; Doxil alone, 29 days; MB-FUS+Doxil, 31 days). The average tumor volume doubling times (T_1/2) were 3.7, 2.2, 2.3, and 2.7 days in the MB-FUS+Doxil, FUS-alone, control, and Doxil-alone groups, respectively. One recent study adopting multiple treatments involving combined MB-FUS BBB opening and Doxil delivery has also demonstrated significant improvements in treatment outcomes [80].

4.2 MB-FUS BBB opening to enhance BCNU delivery

MB-FUS-enhanced BCNU delivery was also investigated in the present study [81]. BCNU has been used clinically for many years as a chemotherapeutic agent for the treatment of gliomas [82]. Although BCNU is lipophilic, which allows it to penetrate the BBB structure, its substantial toxicity limits the overall dose and thus the concentration in the tumor. Only a modest benefit in survival has been shown in clinical trials compared to radiation therapy alone; the median survival improvement was shown to be limited, which improved from 9.4 months for radiation alone to 12 months for combined radiation and BCNU treatment [83].

Using the in-house-developed integrated MRI-monitored FUS platform to investigate enhanced delivery of BCNU in mice (frequency = 400 kHz, acoustic pressure = 0.62 MPa, burst length = 10 ms, PRF = 1 Hz, and sonication duration = 30 s), we found a nearly twofold FUS-dependent increase (from 170 to 344 μg) in the dose of BCNU at the tumor. Without pre-sonication, the BCNU concentration was similar in healthy and tumor-bearing brain sites (170 vs. 150 μg). Tumors in the FUS-alone group grew about threefold during the observation period (with the progression period of 20 days), which was similar to that in the control group (about 2.5-fold). Treatment with BCNU alone also resulted in a similar tumor progression. Combining MB-FUS and BCNU was reported to exert the largest tumor-suppressing effect. Moreover, MB-FUS+BCNU also increased the median survival time to over 50 days, compared to 29, 26, and 32 days in the control, MB-FUS-alone, and BCNU-alone groups, respectively. The application of MB-FUS to enhance BCNU delivery to brain tumors therefore appears to suppress tumor growth as well as prolong animal survival relative to the use of either treatment alone.

4.3 MB-FUS BBB opening to enhance TMZ delivery

TMZ is an alkylating agent that was approved for the treatment of newly diagnosed or recurrent brain tumors in 2005 [84, 85]. TMZ is an imidazotetrazine derivative that spontaneously converts to monoethyl triazenoimidazole carboxamide, which is the active metabolic product of both TMZ and dacarbazine. TMZ has the potential advantage of being absorbed orally and then entering the CNS [86]. Recent TMZ phase-III trials showed more promising results, with a clinically significant improvement in the 2-year survival rate of 26.5% (median survival time of 15 months) for radiotherapy plus TMZ compared to radiation alone [87].

We also recently demonstrated enhanced TMZ delivery by MB-FUS BBB opening [88]. In that study, liquid chromatography-tandem mass spectrometry (MS) was used to measure the TMZ levels in both CSF and plasma. We showed that combining MB-FUS significantly increased the TMZ CSF/plasma ratio from 22% to 39% compared to TMZ alone (frequency = 500 kHz, acoustic pressure = 0.6 MPa, burst length = 10 ms, PRF = 1 Hz, and sonication duration = 60 s). The tumor progression over 7 days showed a significant control effect, with the tumor progression ratio reducing from 24 to only 5. Animals receiving high-dose TMZ showed controlled tumor progress, however, their survival was not extended. In contrast, combined delivery of a reduced dose of TMZ (75 mg/ kg per day, 5 days total) with MB-FUS was shown to extend the animal survival significantly compared to control (100 mg/ kg per day, 5 days total) (improvement in the median survival time of 15%; Table 2).

5. Novel Multifunction MBs Facilitate MB-FUS BBB Opening for Brain Tumor Therapy

MB-FUS BBB opening for brain tumor therapy has employed not only commercial agents but also newly designed MBs. Here we review our recent work related to the concept and design of three different types of multifunction MB for FUS-enhanced brain tumor drug delivery (Figs. 3, 4).

5.1 Concept of novel multifunction MBs

The concept of therapeutic agents being encapsulated in or conjugated with MBs has been developed over the past few years. In addition to the synergistic effects of ultrasound and MBs to enhance the permeability of biological barriers such as cell membranes, small blood vessels, and the BBB, as discussed above, MBs can serve as protective drug carriers. Encapsulating unstable agents protects them from degradation in blood, thus prolonging their half-lives in vivo, improving treatment efficacy, and lowering the required dose [89]. Another advantage is that the encapsulated agents can be released during the ultrasound-triggered MB destruction process. Chemotherapeutic drugs can thus be directly and specifically delivered to target tissues via ultrasound-mediated perforations, whereas the uptake of the drugs by non-target tissues is reduced. The encapsulated agents are conjugated close to the shell of MBs, increasing the opportunity for microstreams, shock waves, and microjets to drive them toward the tissues and enhance their uptake in the ultrasound-treated region [26]. Since MBs act as ultrasound contrast agents, the drug delivery process can also be concurrently monitored by detecting the drug-loaded MBs [90].

Several strategies have been proposed for incorporating therapeutic agents in MB carriers [91, 92], including attachment to the outer shell surface, embedding within the shell, dissolving hydrophobic drugs in the oily layer between the gas core and shell, and linking them to the shell, as for example via streptavidin-biotin interactions. An example of the attachment of agents to the outer shell is non-covalent binding of negatively charged nucleic acids to the outer positively charged lipid shell of MBs for ultrasound-mediated gene delivery [93]. Although the payload of nucleic acids could be improved by increasing the amount of positively charged lipids, the presence of excessive charged lipids would disrupt lipid packing, resulting in higher surface tension and subsequently lower MB stability. Instead, a structure consisting of multilayered nucleic acids and positively charged poly-L-lysine was constructed to increase the nucleic acid loading capacity [94].

Drugs can be embedded in MBs by simply adding them during MB preparation. However, the degree of drug loading is rather limited in this approach, and it is influenced by the polarity of the drug, with hydrophobic molecules being preferentially packaged. In contrast, the design of acoustically active liposomes with an oil layer between the gas core and lipid has allowed encapsulation of hydrophobic drugs at a high loading capacity within this oil layer [94]. However, high ultrasound intensities are required to trigger the release of such encapsulated drugs [95].

Lastly, drugs can be pre-incorporated into carriers such as liposomes, micelles, or microspheres, and these structures can then be easily attached to lipid MBs, usually via avidin-biotin interactions [96]. Such advanced MB complexes have extremely high drug loading capacities and the advantage of being able to encapsulate both hydrophilic and hydrophobic drugs.

 Fig 3 

Application of multifunction microbubbles for focused-ultrasound-induced brain tumor drug delivery.

 Fig 4 

Size distribution and microscopy images of microbubbles.

The effectiveness of ultrasound molecular theranostics could be improved by exploiting the nonlinear behaviors (e.g., cavitation dose, resonance frequency, or fragmentation threshold) of MBs. The resonance frequency and magnitude of the radiation force for targeting efficiency enhancement from the ultrasound-excited MBs relate to MBs size [97]. Furthermore, insonation acoustic pressure that can induce MBs fragmentation is also size-dependent [98-99]. Thus, controlling the size distribution of multifunctional MBs is crucial for most ultrasound theranostics applications [100]. Multifunction MBs were designed to have similar but with a smaller and more mono-dispersed dimension with the considerations of: maintaining sufficient nonlinear responses under therapeutic/imaging ultrasound frequency excitation, safety improvement to reduce erythrocyte extravasations, and optimization of payload of drug molecules or targeting ligands [47, 48].

5.2 FUS-enhanced brain tumor drug delivery using BCNU-loaded MBs

While MB-FUS-mediated BBB disruption enhances the delivery of BCNU to brain tumors, the short half-life of BCNU (20-50 min in vitro and less than 15 min in vivo) still intrinsically limits its efficacy after systemic application [101, 102]. A lipid-shell-based and BCNU-loaded MB (BCNU-MB; Fig. 3) was therefore proposed, which could not only serve as a drug-carrying vehicle to protect BCNU from rapid degradation, but could also be activated by FUS to concurrently achieve BBB opening and trigger the local release of BCNU [103].

Quantification in normal rats showed that encapsulation of BCNU in MBs prolonged its circulatory half-life by fivefold (from 13.5 to 67.5 min). Compared to traditional IV BCNU administration, synergistically functioning BCNU-MB+FUS (frequency = 1 MHz, acoustic pressure = 0.5 MPa, burst length = 5 ms, PRF = 5 Hz, and sonication duration = 60 s) significantly enhanced drug delivery by 4.2-fold (from 4.2±0.2 μg to 17.9±1.1 μg, mean±SD). The concentration of BCNU was 3.3-fold higher in the sonicated brain than in the contralateral unsonicated brain (5.4±0.4 μg). The cytotoxic effects of encapsulating BCNU were also studied by examining its accumulation in the liver and evaluating liver function. We found that loading BCNU in MBs lowered drug deposition in the liver by 4.8-fold, from 113.6±3.6 to 23.9±3.6 μg, and also lowered aspartate aminotransferase and alanine aminotransferase levels compared to the BCNU-alone group, demonstrating that BCNU-MBs have the potential to reduce liver toxicity and damage.

Synergistic effects between BCNU-MBs and FUS led to an early improvement in treatments of tumor-implanted rats. Tumors in the control group progressed rapidly from 22.1±17.6 mm³ (day 10) to 202.5±24.2 mm³ (day 31), whereas the BCNU-alone group demonstrated temporary but poor control of tumor progression (12.9±9.1 mm³ on day 10, but 56.8±38.8 mm³ on day 17 and 150.8±11.5 mm³ on day 31). In contrast, the tumor size was only 11.0±1.0 mm³ on day 31 when using the BCNU-MB+FUS treatment.

The median survival times in the control and BCNU-alone groups were 29 and 29.5 days, respectively, demonstrating the limited capabilities of BCNU for brain tumor treatment. Nevertheless, the median survival time in the BCNU-MB+FUS group was increased to 32.5 days (a 12% increase compared to the control and BCNU-alone groups), and the maximum survival time of animals was significantly extended to 59 days.

5.3 FUS-enhanced brain tumor drug delivery using antiangiogenic-targeting drug-loaded MBs

After successfully enhancing drug delivery by BCNU-MBs, targeted brain tumor delivery was attempted [104]. Vascular endothelial growth factor receptor 2 (VEGF-R2) is recognized to be overexpressed in the endothelial cells of gliomas, causing cell proliferation and migration and resulting in excessive angiogenesis in tumor regions. One type of MBs combining VEGF-R2-ligand conjugation and BCNU encapsulation (designated VEGF-BCNU-MBs) was thus designed to target drug delivery specifically to the sites of the tumor vasculature exhibiting overactive angiogenesis, which are characterized by overexpression of the VEGF-R2 receptor (Fig. 3).

Both immunofluorescence staining and ultrasound imaging revealed that the VEGF-BCNU-MBs exhibited prolonged retention in the blood circulation and displayed higher cumulative concentrations in the tumor region. In vivo drug accumulation in tumor tissues was significantly enhanced (by 1.9-fold) in the VEGF-BCNU-MBs+FUS BBB opening group compared to the BCNU-MB+FUS BBB opening group (frequency = 1 MHz, acoustic pressure = 0.5 MPa, burst length = 5 ms, PRF = 5 Hz, and sonication duration = 60 s). Moreover, the tumor-to-normal-tissue concentration ratio when using the VEGF-BCNU-MBs was found to be 7, compared to 2.7 when using traditional MBs (without targeting ability) and FUS exposure. Besides, liver BCNU deposition was significantly lower in the VEGF-BCNU-MB+FUS BBB opening group than in the BCNU-MB+FUS BBB opening group.

The antitumor efficacy was investigated in tumor-bearing rats. Tumors in the untreated groups grew rapidly. Neither the VEGF-R2 targeting ligand alone nor targeting MBs without BCNU (VEGF-MB) combined with FUS BBB opening provided effective control of tumor progression, although transient suppression of tumor progression was observed in the BCNU-alone and BCNU-MB+FUS groups. It was clear that tumor suppression was greatest when using VEGF-BCNU-MB+FUS BBB opening.

Animal survival was not improved in the BCNU-alone, VEGF-MB+FUS BBB opening, and BCNU-MB+FUS BBB opening groups relative to the control group (median survival times of 23, 22.5, 18.5, and 19 days, respectively). The VEGF-R2 targeting ligand appeared to prolong survival (median survival time of 26 days), while survival was further prolonged in the VEGF-BCNU-MB+FUS BBB opening group (median survival time of 42 days).

5.4 FUS-enhanced brain tumor drug delivery using theranostic MBs

For the treatment of gliomas, drugs not only need to be effectively packaged in microcarriers, but their dynamics, distribution, and accumulation at the target site also need to be monitored in vivo. However, even when they are administered concurrently, therapeutic agents and biological probes can perform diverse pharmacodynamic behaviors, implying that a separate probe may actually not reflect the true drug distribution. A DOX-loaded and SPIO-nanoparticle-conjugated phospholipid-based MB structure (DOX-SPIO-MB) was therefore designed to concurrently achieve BBB opening and delivery of therapeutic agent, while serving as a dual contrast agent in both ultrasound imaging and MRI modalities for direct confirmation of drug quantification/deposition (Fig. 3) [105].

SPIO nanoparticles have been approved for use in clinical diagnosis as an MRI contrast agent. We previously confirmed that FUS BBB opening can be monitored with the aid of SPIO nanoparticles (frequency = 400 kHz, acoustic pressure = 0.62 MPa, burst length = 10 ms, PRF = 1 Hz, and sonication duration = 120 s). Furthermore, the application of an external magnetic force to SPIO is a potential way to achieve active magnetic targeting (MT) to specific tumor regions [106]. IV administered DOX-SPIO-MBs exhibited excellent contrast capability in vivo, with a 25.4% enhancement for ultrasound and a 40.2% decrease for MRI, respectively. Treating tumor-implanted rats with DOX-SPIO-MBs followed by FUS sonication resulted in a 2.1-fold increase in DOX deposition at the brain tumor relative to normal brain tissue (frequency = 400 kHz, acoustic pressure = 325 kPa, burst length = 2.5 ms, PRF = 1 Hz, and sonication duration = 90 s). In contrast, free DOX alone did not result in effective drug deposition in tumor regions. The delivery of SPIO to brain tumors was also investigated by inductively coupled plasma MS, which revealed that the SPIO accumulation capability was highly dependent on the concentration of administered DOX-SPIO-MBs. Similar to DOX delivery, the DOX-SPIO-MB+FUS BBB opening strategy improved SPIO delivery to brain tumors (by 2.7-fold, 11.9±0.9 mg/g tissue). Furthermore, combined FUS exposure and MT provided the most significant SPIO accumulation enhancement in tumor sites (by 4.0-fold, 15.3±0.7 mg/g tissue).

In vivo drug delivery to the tumor site using DOX-SPIO-MB+FUS BBB opening was monitored by detecting SPIO as hypointense signal-loss regions in contrast-enhanced MRI using a sequence with heavy T2* weighting. A gradual enhancement of SPIO accumulation was observed in tumor areas after performing MT for 40 min following the BBB opening process. The total SPIO deposition was increased significantly, by 22.4% (note that 12% deposition was achieved in groups without MT).

The SPIO nanoparticles can be detected by MRI, which allows the in vivo distribution of therapeutic SPIO nanoparticles to potentially be traced or even quantified under an image-guided brain tumor drug delivery procedure. Future work should be focused on correlating DOX delivery and SPIO nanoparticle quantification for clinical theranostic applications, as well as modifying ligands so that they target DOX-SPIO-MBs for use in multiactive targeting.

6. Conclusion and Perspective

MBs not only serve as a diagnostic contrast agent for ultrasound imaging, but also provide the potential for brain drug delivery when combined with focused ultrasound. Here we have reviewed the application of MBs with FUS to temporarily open the BBB as a new strategy for delivering therapeutic agents to the brain. Lipid-shell MBs respond strongly to ultrasound, allowing MB-FUS BBB opening to be performed using commercial diagnostic MBs. We reviewed the current progress in MB-FUS-enhanced chemotherapeutic agent delivery for brain tumor treatment. The design of multifunctional MBs has further expanded the potential of this treatment approach by integrating BBB opening with concurrent drug release, active targeting, or even theranostic features.

MB-FUS BBB opening serves as a promising method for non-invasively and locally enhancing the targeted delivery of therapeutic agents into CNS tumor regions, providing the potential to improve the treatment efficacy of chemotherapy. Doxil has been approved for the clinical use for ovarian tumors, and TMZ as well as BCNU are already clinically approved chemotherapeutic drugs for brain tumor treatment, suggesting that the enhanced delivery of these drugs achieved by MB-FUS BBB opening is highly clinically relevant. Moreover, the use of multifunction MBs that encapsulate chemotherapeutic drugs, with antiangiogenic targeting, as well as magnetic-sensitive modifications provide many new opportunities to further improve the capability of enhanced drug delivery via MB-FUS BBB opening, and bring new research directions toward realizing noninvasive brain drug delivery for improving brain tumor therapy.

However, notwithstanding the advantages of MB-FUS BBB opening, it has also been suspected that the interaction between MBs and FUS may produce unnecessary side effects, for example erythrocyte extravasations, intracerebral micro-hemorrhages, edema, neuron injury, cell apoptosis, and inflammation, which could be arisen by over-excitation of ultrasound (e.g. excessive acoustic pressure or excessive sonication duration) or over dosages of microbubbles [73, 82, 107-110]. As research on MB-FUS BBB opening in enhanced brain drug delivery progresses, it is expected that useful insights for the control of the above-mentioned parameter for BBB disruption will be further investigated.

Acknowledgements

We thank the National Science Council of Taiwan for grants 101-3011-P-033-003 and 101-2627-M-007-001, and the National Tsing Hua University (grant 102N2046E1) for their financial support.

Competing Interests

The authors have declared that no competing interest exists.

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Author contact

Corresponding author: Department of Biomedical Engineering and Environmental Sciences, National Tsing Hua University, No. 101, Section 2, Kuang-Fu Road, Hsinchu, Taiwan 30013. Tel: +886-3-571-5131 ext. 34240; Fax: +886-3-571-8649. E-mail: ckyehnthu.edu.tw

Received 2013-11-7
Accepted 2014-1-20
Published 2014-2-12

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