Current approaches to cancer treatment encompass surgery[1], radiotherapy [2], [3], chemotherapy [4], immunotherapy [5], hormone therapy [6], targeted therapy [7], or a combination of these strategies [8]-[10]. Chemotherapy is one of the most widely employed methods compared to other techniques, where chemotherapeutic agents are employed to eradicate tumor cells [11]; however, conventional chemotherapy encounters challenges in achieving precise drug delivery to tumors, with approximately only 1% of the injected dose reaching the tumor site after systemic administration [12], [13]. Furthermore, the heterogeneity of metastatic tumors presents varying genetic characteristics from primary tumors, making it challenging to devise a unified approach [14], [15]. Breast cancer has emerged as the most prevalent cancer globally [16]-[18]. It Is the second leading cause of death globally and comprises approximately 12.5% of diagnosed cancer cases [17], [19].
Surgery is deployed as the primary treatment for the removal of malignant cells. However, the choice of treatment depends on the stage, type of tumor, size, grade, proliferation rate, and involvement of lymph nodes. Surgery includes lumpectomy (partial breast tissue removal) and mastectomy (complete breast tissue removal) [20], [21]. Yet surgery may not yield the most effective results for aggressive and metastatic tumors, and adjuvant therapies like chemotherapy, radiation, hormonal therapy, and targeted therapies come into play. These treatment methods enable clinicians to tailor therapeutic approaches depending on the tumor’s behavior while monitoring, evaluating, and adjusting the tumor’s response to chemotherapy or hormonal therapy. This helps conserve breast tissue and enhance overall patient well-being.
Roughly 25% of breast cancer cases exhibit overexpression of the Human Epidermal Growth Factor Receptor (HER2), a proto-oncogene, that correlates with malignant transformation and notably lower survival rates among breast cancer cases that have undergone lymph node metastasis. HER2 overexpression, also known as HER2-positive breast cancer, is widely used as a significant biomarker for breast cancer treatment and helps tailor personalized treatment strategies [17], [22]. Monoclonal antibody (mAb)-based breast cancer treatment strategies include Trastuzumab (Herceptin) [23]-[25], Pertuzumab (Perjeta) [26], [27], Margetuximab (Margenza), [28], [29], Neratinib (Nerlynx) [30], [31], Tucatinib (Tukysa) [32]-[34], DS-8201 (Enhertu) [35], [36], and Ado-trastuzumab emtansine (Kadcyla) [37], [38]. Trastuzumab (Herceptin) is a monoclonal antibody that received FDA approval in 1998 for the treatment of HER2 + breast cancers [39]-[43]. It specifically targets the HER2 protein overexpressed on the surface of cancer cells and blocks the cell signaling pathways for growth and division [22], [44], [45].
To address the concerns related to drug release and negative side effects of chemotherapeutic drugs, researchers are exploring the nano-sized cavities of nanoparticles that act as a vehicle to effectively encapsulate and transport the drug to the target tumor site [46], [47]. A recommended size of nanocarriers < 200 nm is crucial to attain passive targeting, which enables drug carriers’ accumulation at the tumor site due to their extravasation through the leaky tumor vasculature and poor lymphatic drainage. This is called the enhanced permeability and retention (EPR) effect [48]-[50]. In contrast, targeting ligands/ moieties conjugated to the surface of nanocarriers facilitates active targeting by binding to the overexpressed receptors on the cancer cell surfaces. A schematic representation of passive and active targeting mechanisms is illustrated in Fig. 1. A study conducted by Ayub and Wettig demonstrated promising outcomes by nanoparticles, particularly in treating various types of cancers, especially brain cancer [51]. Liposomes have been extensively explored as carriers for imaging agents, active drugs, nucleic acids, and proteins [52]-[57].
Liposomes were first discovered as biological models by Bangham in 1961 [58]. Gregory Gregoriadis was the first to explore the concept of drug encapsulation within liposomes in 1973 [59]. Liposomes resemble cell membranes in composition. They are biocompatible, biodegradable, non-toxic, stable in the physiological environment, and reduce the toxicity of the encapsulated drug [60]-[63]. Most widely used liposomal nano-drug delivery systems with encapsulated active hydrophobic and hydrophilic drugs have paved the way for novel approaches that adapt to the tumor environment, utilizing the pathophysiology of tumors [60], [64]-[67]. Additionally, they offer a controlled distribution and release of therapeutic agents at the target site, thus reducing the necessity for frequent dosing, potentially improving patient compliance, and minimizing some side effects [65], [68]-[70].
To mitigate the rapid clearance of liposomes by the reticuloendothelial system (RES), liposomes are coated with highly flexible polyethylene glycol (PEG) polymer [71]. PEG-coated liposomes experience reduced immune system recognition, enhanced stability in vivo, and prolonged circulation lifespan. This provides the protected payloads with a sustained presence in the bloodstream, providing an extended period to reach their target sites [72]-[74]. Furthermore, by introducing targeting moieties to modify liposomal surfaces, researchers have improved its tumor targeting capabilities, delivering therapeutic drugs to tumors and enhancing the efficacy of cancer treatment while limiting their interaction with healthy cells and minimizing the off-target effects [22], [46], [75]-[79]. This method offers a more tailored and personalized therapeutic approach based on individual cancer treatment needs. Liposomes conjugated with Herceptin have been employed for targeted breast cancer therapy and demonstrated a significant enhancement in anticancer drug uptake and cellular toxicity levels [77], [80].
Recent studies have shifted their focus to smart drug delivery systems (SDDS) for strategic drug delivery, including release from nanoparticles at the target microenvironment, minimizing premature release, and controlling the drug release rate upon exposure to a stimulus at the target site. Smart liposomes can release drugs promptly in response to distinct physiological characteristics. These can be classified into internal and external triggers. Internal triggers include variations in pH levels [81], enzyme activity fluctuations [82], redox gradients [83], [84], and hormone levels [85]. In contrast, external stimuli include temperature [86], [87], magnetic field [88], ultrasound (US) [89], and high-energy radiation [22], [90]-[92]. These provide platforms for personalized and targeted medicine, where the controlled release of therapeutic agents is adjusted accurately to align precisely with the patient’s condition [47], [78]. The choice to utilize low-frequency ultrasound (LFUS) as the stimulus in this study stems from its safe, non-invasive characteristics, customizable parameters, and precise targeting capabilities [89], [93]-[96]. Figure 2 represents a schematic illustration of a smart liposome-based drug delivery system.
Crucial ultrasound parameters include frequency, power density, and pulse duration. Drug delivery systems employ high-intensity and low-intensity focused ultrasound to induce synergistic effects in the controlled release of chemotherapeutic drugs. Furthermore, ultrasound tends to induce necrosis. Thus, it is pertinent to utilize pulsed-wave (PW) Doppler US to allow dissipation of heat between successive pulses [98]-[101]. Exposure to high-frequency ultrasound locally heats the body tissues; this helps the accumulation of nanoparticles at the target site and activates temperature-sensitive nanoparticles. However, high hyperthermia > 43oC ceases tissue blood flow, leading to rapid cell death (necrosis). Figure 3 illustrates a visual representation of the thermal effects induced on tissues upon exposure to ultrasound. Moreover, the US propagates through a medium as high-pressure or low-pressure waves, inducing pressure variations within a medium. This imparts energy to particles of the propagating medium and leads to the production of small gas pockets. This phenomenon is called acoustic cavitation. Variations in pressure cause gas bubbles to linearly oscillate, creating strong shear forces that temporarily permeabilize cell membranes and help penetrate the nanocarriers into the tumor tissue. Figure 4 depicts a schematic representation of the mechanical effects of ultrasound, with microbubbles (MBs) undergoing stable cavitation or inertial cavitation. Moreover, when oscillations become non-linear with the increase in US intensity, rapid growth and subsequent implosion of gas bubbles take place; this is called collapse cavitation. The implosion is accompanied by high-pressure shock waves and sometimes the production of a sonic jet that leads to the sonoporation of the cell membrane. [101]-[105]. Figure 5 depicts an illustration of microjets created by microbubbles undergoing collapse cavitation.
Despite their clinical use, liposomes still encounter limitations in swiftly achieving optimal chemotherapeutic drug concentrations at the target site, due to their adequately stable liposomal membranes and inherent lack of responsiveness to ultrasound. This reduces their potential effectiveness against cancer. Researchers have investigated strategies to enhance the responsiveness of liposomes to ultrasound by incorporating MBs, nanobubbles, and phase-changing nanoemulsion within or upon droplets [74], [106]-[112]. Microbubbles are composed of lipid shells filled with perfluorocarbon gas. Olsman et al. investigated the effect of focused ultrasound (FUS) and MBs on the transferrin (Tf) targeted liposomes in enhancing the permeability of the blood-brain barrier (BBB) in rats, which overexpress Tf receptors in the BBB. The study revealed that FUS and microbubbles helped safely increase blood-brain barrier permeability and recorded a 40% increase in the accumulation of Tf-targeted liposomes in the brain hemisphere compared with isotype immunoglobulin G (IgG) liposomes [113].
However, the size of MBs, (diameter greater than about 1 µm) limits transport within the tumor vasculature and precludes MBs from benefiting from the EPR effect. Nevertheless, MBs have been employed as intra-vascular agents to actively target endothelial markers such as VEGFR2 and αvβ3 integrin [114], [115]. The large size of microbubbles (Figure 5) incentivized the development of nano-scale-sized nanobubbles and nanoemulsions that would easily extravasate into the tumor tissues and become endocytosed into the tumor cells. Upon exposure to ultrasound, nanoemulsion droplets of perfluorocarbon liquids transform from liquid to gas, resulting in an increased volume within the liposomal vesicle, subsequently leading to rupture and prompt release of the enclosed drug [97], [112], [116]-[120]. This is known as acoustic droplet vaporization (ADV). It is important to note that lipid bilayers can tolerate only a 3% increase in volume before reaching the rupture point. This substantial increase in volume upon phase change is sufficient to rupture both eLiposomes and the endosome [120]-[122]. This phenomenon helps attain the desirable therapeutic dose while regulating and controlling the release of anticancer drugs at the target site.
Lattin and Pitt designed experiments to investigate the performance of eLiposomes and liposomes at physiological temperatures (37°C). These experiments revealed the stability and capability of eLiposomes to sequester drugs at physiological temperatures. Experiments employed a fluorometer that measured fluorescence in a heated water bath at incubation times of 3, 10, 20, and 30 minutes. They repeated the process for both eLiposomes with large (450 nm) and (100 nm) emulsions. No calcein release was observed from the samples mentioned above, signifying that heating to body temperature alone cannot render eLiposomes unstable. Finally, Triton-X was used to lyse the eLiposomes, which released all calcein sequestered in the eLiposomes, thus indicating that eLiposomes are very stable at physiological temperatures. They further compared the ultrasound-induced release of the encapsulated model drug, calcein, from eLiposomes (containing PFC5 and PFC6) with the two negative controls (without the droplets and with droplets outside the liposomes vesicle). The eLiposomes showed significantly higher calcein release than both control groups, which was attributed to the emulsion droplets inside the liposome vesicles disrupting its membrane structure from within the eLiposomes and creating calcein release. The eLiposomes showed 3–4 times more calcein release than the control groups, which increased further upon increasing ultrasound power intensity and time of exposure. However, after a certain amount of time or energy, no further increase was observed upon increasing the exposure. The study also reported that an increase in power density resulted in an increased tissue temperature; however, this increase in temperature was not responsible for the significantly higher release from eLiposomes compared to conventional liposomes. Furthermore, they studied the behavior of PFC5 eLiposomes and control (conventional) liposomes as a function of US frequency (varying from 20 kHz to 525 kHz) and mechanical indices (MI = 0.53 at 5 W/cm2 and MI = 1.41 corresponding to 35 W/cm2). In this study, PFC5 eLiposomes and control liposomes were exposed to PW ultrasound for 2 to 30 seconds with 525 kHz at 20 kHz pulse repetition frequency. The study demonstrated that frequency significantly affects the phase transition of emulsion droplets. They concluded that lower frequency offers a long window of negative pressure, allowing more time for bubble nucleation and gas expansion; thus, it was concluded that increasing the frequency decreases the threshold of acoustic vaporization. PFC5 eLiposomes showed a significant difference in their drug release compared to control liposomes: about 2–3 times and 3–5 times more drug release was demonstrated by PFC5 eLiposomes when exposed to 5 W/cm2 and 35 W/cm2, respectively; however, the study showed no significant release from control liposomes with the changes in intensities [111].
In this present study, pegylated liposomes encapsulating calcein and nanoemulsion droplets were formulated and functionalized with the monoclonal antibody Herceptin (see Fig. 6). The release of calcein from conventional liposomes, eLiposomes, HER-conjugated liposomes, and HER-conjugated eLiposomes was catalyzed by harnessing low-frequency ultrasound as a trigger.