Targeted delivery of microbubbles and nanobubbles for image-guided thermal ablation therapy of tumors

Thermal ablation treats tumors by localized deposition of thermal energy that causes protein denaturation and coagulation necrosis. Commonly used thermal ablation techniques include laser ablation, radiofrequency ablation, microwave ablation and high-intensity focused ultrasound (US) ablation [1–4]. These techniques have been explored for minimally invasive management of tumors in various locations, including liver, kidney, lung, prostate, bone and breast. The potential major advantages of thermal ablation processes include minimal patient trauma, a low incidence of complications, an outpatient treatment approach, rapid post-procedural recovery and a more rapid return to normal activities. However, the widespread application of these thermal ablation techniques remains controversial due to ongoing concerns with regards to its overall clinical efficacy and its impact on disease-free survival and local recurrence rates as compared with surgical resection. The primary reasons for these concerns are poor process control, incomplete destruction of residual cancer cells and the potential for incidental thermal injury to surrounding normal tissue structures.

Efforts have been made to improve ablation process control and to maximize therapeutic efficacy. Both in vivo and ex vivo thermal ablation experiments have previously evaluated the factors contributing to the size and shape of the resultant zone of ablation, such as ablation power, duration, heating mode, use of perfusate, supplement of intratumoral drug injection, cooling of the thermal ablation device and vascular blood flow [5–9]. Numerical models have also been developed to simulate the ablation process and optimize the ablation parameters [10–12]. However, the actual implementation of these experimental and numerical advances into the clinical setting faces many potential challenges. One major challenge is the lack of real-time, intraprocedural imaging tools for quantitative assessment of the tumor boundaries and of the margins of the ablation zone. Clinical imaging systems that have been explored for tumor ablation applications include US, computed tomography, PET and MRI [13–16]. However, these systems are not optimized for real-time imaging owing to various related factors, such as low sensitivity to tissue thermal damage, high susceptibility to ablation-induced tissue gas bubbles, potential for concomitant radiation exposure and cost and complexity of such imaging systems.

MRI-guided focused US surgery (MRgFUS) is a relatively recent approach to thermal ablation therapy, which integrates imaging feedback control [17]. The clinical feasibility of MRgFUS has been demonstrated on various tissues, including breast, liver, prostate, bone and brain. MR thermometry detects tissue temperature changes in order to estimate the accumulative thermal dose. Since the assessment of the margins of the ablation zone involves the Arrhenius integral of the tissue temperature history [18,19], MR thermometry is vulnerable to the accumulation of errors associated with thermal damage modeling, tissue displacement, temperature accuracy, tissue property variations and other transient changes during the thermal ablation process. Additionally, further technology development is necessary to enhance the imaging sensitivity in adipose tissue and to reduce the size and the cost of the MRgFUS imaging system.

In order to take full advantage of thermal ablation technology and maximize its clinical potential in a minimally invasive approach to malignancies, it is important to develop and optimize imaging systems and contrast agents for real-time quantitative assessment of the tumor boundaries and the zone of ablation. In our opinion, the ideal imaging system to guide the thermal ablation process of tumors should possess the following features:

• • The imaging system should not interfere with the thermal ablation procedure;

• • The accuracy for assessing the zone of ablation should not be affected by motion artifacts and other temporally accumulative errors during the thermal ablation process;

• • The imaging system should be able to detect both the tumor boundaries and the margins of the ablation zone for close-loop feedback control of the thermal ablation process;

• • In order to maximize its clinical usability and flexibility, the imaging system should ideally consist of a low-cost, portable, handheld imaging device that integrates multiple imaging modalities, rather than using a single imaging modality;

• • The contrast agents used for imaging of the tumor boundaries and the zone of ablation should be nontoxic, biocompatible and biodegradable.

At the current time, imaging systems for thermal ablation therapy that satisfy all the aforementioned design requirements are not commercially available. However, several recent technical could be incorporated into a thermal ablation imaging strategy that employs a handheld imaging device for the assessment of tumor boundaries and margins of the ablation zone. One such technical advance is the potential utilization of heat-sensitive microbubbles (MBs) and nanobubbles (NBs) to specifically target molecular markers that are overexpressed by various tumors. A wide variety of imaging agents can be encapsulated in MBs and NBs for concurrent utilization in conjunction with US, fluorescence (FL) and photoacoustic (PA) imaging. The propagated heat in the thermal ablation process can be used to activate MBs and NBs at a tissue-lethal thermal dose and introduce significant contrast changes within the tissues that can be utilized for the assessment of the zone of ablation. Herein, we will discuss the rationale and methodology for such an innovative imaging system for thermal ablation therapy.

Ultrasound, fluorescence & photoacoustic imaging

Diagnostic US uses ultrasonic scattering as the primary imaging contrast to visualize internal organs and body structures, with an imaging depth of up to approximately 60 mm and with a submillimeter spatial resolution. US can be used to localize the tumor and to guide the placement and positioning of the thermal ablation device [14]. However, clinical US alone is not ideal for real-time thermal ablation guidance. This is due to the image interference and image artifact seen on diagnostic US that is caused by thermal ablation-induced gas bubble formation within the tissues, as well as the intrinsic inability of diagnostic US alone to provide any molecular functional information during the thermal ablation process.

Fluorescence imaging characterizes tissue molecular signatures by irradiating endogenous and exogenous fluorophores and detecting their fluorescence emission [20,21]. Typically, in vivo FL imaging is carried out in the near-infrared (NIR) wavelength range. This is because NIR light can penetrate to a depth of several centimeters within biologic tissues, allowing for relatively deep tissue imaging and tomography. NIR imaging has the advantages of low cost, portability, molecular sensitivity and real-time noninvasive monitoring of multiple tissue parameters. However, it has poor depth-resolved resolution (in centimeters), and therefore is not sufficient to provide spatial guidance for tumor localization or for placement and positioning of the thermal ablation device.

Photoacoustic imaging uses a short-pulsed light beam to irradiate biological tissues, and reconstructs tissue functional properties by detecting a heat-induced acoustic pressure wave [22]. PA imaging combines the molecular sensitivity of optical imaging and the spatial resolution of diagnostic US. It allows for imaging of up to a depth of approximately 50 mm, and spatial resolution of better than 700 µm. However, the sampling rate for PA imaging is relatively slow compared with diagnostic US or FL imaging.

In regards to these three separate imaging modalities, it is technically feasible to integrate diagnostic US, FL and PA into a single, multimodal handheld imaging device in order to take advantage of the technical benefits of each of the individual modalities. In doing so, the limitations of each individual imaging modality can be theoretically overcome, thus providing real-time guidance and information regarding the tumor boundaries and the margins of the ablation zone during the thermal ablation process.

Multifunctional active MBs & NBs

Microbubbles, in the size range of 2–3 µm, have been approved by the US FDA for use as a contrast enhancement agent in clinical applications related to diagnostic US. MB-enhanced US imaging can be used to guide placement and positioning of the thermal ablation device and to assess the treatment results during thermal ablation therapy [23]. Commonly used MB shell materials include liposome, albumin and poly(lactic-co-glycolic acid) (PLGA). PLGA is a FDA-approved biodegradable and biocompatible material for implantation and drug-delivery applications [24]. The use of drug-loaded PLGA microspheres and nanoparticles has been described for targeted tumor imaging and therapy [25,26]. To help prolong the residence time of these particles within the body and to reduce their immunogenicity, PLGA microspheres and nanoparticles are commonly covalently attached with the polymer poly(ethylene glycol) in a process called PEGylation [27]. Since many tumors have leaky vasculature, with a typical pore size of approximately 400–600 nm [28], NBs and other nanoparticles that have been fabricated within a certain size range may penetrate through the tumor vasculature and can then potentially target molecular biomarkers that are overexpressed within the interstitial tumor cell environment [29]. This is in contrast to MBs and microspheres, which are clinically available in the size range of 2–3 µm, and can only circulate within the vasculature and target molecular biomarkers that are overexpressed by the endothelial cells that line those vessels [30]. There is the potential for various imaging agents to be concurrently encapsulated within both MBs and NBs for use in a multimodal fashion with US, FL and PA imaging technology [31,32].

Many MBs are filled with heavy gases of low permeability in order to enhance the circulation profile and improve imaging contrast. Perfluorocarbon (PFC) is a group of nontoxic materials with a low boiling point that have been approved by theFDA as the filling materials for many US-enhancing MBs. Superheated PFC droplets can be cavitated by US bursts for enhanced US imaging and US therapy [33]. PFC-encapsulated microspheres and nanoparticles have also been explored as a targeted drug-delivery device [34]. Drug-loaded PFC NBs have been used for combinational ultrasonography and targeted chemotherapy [35].

Image-guided closed-loop thermal ablation of tumors

The technologic advances in portable imaging devices and the emergence of biodegradable multifunctional active agents provides motivation for the development of a closed-loop image-guided tumor ablation schema that would be suitable for use in a minimally invasive approach to the management of various malignancies. Such an image-guided tumor ablation schema would integrate the repetitive steps of tumor boundary imaging, treatment planning, ablation process control and assessment of the margins of the zone of ablation, in a closed-loop fashion. In the past, we have worked on each individual step of this tumor ablation schema. For example, we have conjugated multifunctional MBs and NBs to target molecular biomarkers for tumor boundary imaging [29,32,36] and have designed handheld imaging devices for multimodal dynamic tumor imaging [31,32,37,38]. We have also developed a simplified finite element model to simulate the propagation of the zone of ablation in a real-time fashion, with the hope of better image guidance for the thermal ablation process [39], and have synthesized active MB agents for real-time imaging of the zone of ablation [40].

To help facilitate the clinical translation of this closed-loop image-guided tumor ablation schema from the benchtop to the realm of clinical medicine, we have paid close attention to issues related to clinical safety and efficacy. The proposed imaging technology is low cost, portable, handheld, noninvasive and poses no radiation risk. Likewise, the materials used for the synthesis of these multifunctional active agents are either FDA-approved biocompatible and biodegradable materials, or materials without reported toxicity. If this closed-loop image-guided tumor ablation schema is successfully implemented, we expect that such a schema could dramatically impact upon the current clinical management strategies for various malignancies.

Financial & competing interests disclosure

The authors have no relevant affiliations or financial involvement with any organization or entity with a financial interest in or financial conflict with the subject matter or materials discussed in the manuscript. This includes employment, consultancies, honoraria, stock ownership or options, expert testimony, grants or patents received or pending, or royalties.

No writing assistance was utilized in the production of this manuscript.