Nanocarriers were first discovered in the early 1960s, when scientists proposed the application of liposomes for drug delivery [1]. Since then, many nanocarrier systems have been developed and approved for marketing by the U.S. Food and Drug Administration (FDA) and the European Medicines Agency (EMA). Nanocarriers have been extensively used in drug delivery owing to their exceptional physicochemical properties, such as nanometric particle size, surface charge, high entrapment efficiency, and drug loading capacity [2]. Nanocarriers used in cancer treatment have received particular attention from researchers worldwide, since most conventional chemotherapeutic drugs cause systemic toxicity resulting from their poor stability in biological systems, non-selectivity, and non-specificity toward cells expressing the targeted receptors [3]. The first evidence on the feasibility and effective use of nanocarriers in cancer treatment was reported in 1976 by Langer et al. [4], who prepared the first sustained-release, long-circulating poly(ethylene glycol)-poly(lactic acid-ethanolic acid) (PEG-PLEA) nanoparticles, which were later approved by the FDA as the first nanomedicine for therapeutic use in cancer treatment. Generally, nanocarriers are categorized into two main classes: organic and inorganic nanocarriers ^[5][6][5,6]. Organic nanocarriers include nanoemulsions (10–1000 nm), nanosuspensions (<1 µm), nanoliposomes (50–450 nm), polymeric nanoparticles (10 nm to 1 µm), solid-lipid nanoparticles (10–1000 nm), and nanodendrimers (15–200 nm), while inorganic nanocarriers include gold nanoparticles (5–400 nm), silver nanoparticles (1–100 nm), mesoporous silica nanoparticles (30–300 nm), and superparamagnetic iron oxide nanoparticles (100 nm to 5 µm). These tumor-targeting, nano-sized drug delivery systems were developed primarily to reduce the systemic toxicity of chemotherapeutic drugs through encapsulation into nanocarrier systems, which allowed for site-specific delivery with improved passive and active drug targeting (i.e., disease-specific targeted therapeutics). Of all these nanocarriers, liposomes are very promising drug delivery systems with the advantages of being non-toxic, biocompatible, and biodegradable [2]. Liposomes were first discovered by Bangham et al. [7] in 1964. They discovered how membrane molecules interact with water to form unique structural forms, which were described as swollen phospholipid systems [7]. Briefly, liposomes are defined as vesicular systems consisting of one or more concentric spheres of phospholipid bilayers separated by aqueous or buffer compartments ^[7][8][7,8]. When phospholipids are dispersed in an aqueous medium like water or buffer, the hydration of phospholipid polar heads results in a heterogeneous mixture of spherical structures, generally referred to as vesicles, most of which contain multiple phospholipid bilayers forming concentric spherical shells ^[7][8][7,8]. Those were the liposomes first reported by Bangham et al. ^[1][7][8][1,7,8], nowadays referred to as multilamellar large vesicles (MLVs). The sonication of these lipid dispersions results in the size reduction of these liposomes to vesicles containing only a single bilayer with diameters ranging from 20 to 100 nm, later referred to as small unilamellar vesicles (SUVs) [9]. Large unilamellar vesicles (LUVs, 100–1000 nm) are intermediate in size between MLVs (>700 nm) and SUVs [9]. The main components of liposomes are phospholipids and cholesterol, which are naturally occurring substances ^[9][10][9,10]. Phospholipids are amphiphilic molecules with hydrophobic non-polar tails and hydrophilic polar heads. These amphiphilic molecules spontaneously organize into liposomes in an aqueous or buffer environment, driven by hydrophobic interactions and other intermolecular interactions [11]. The proper choice of phospholipid is important to achieve the desired effects. Table 1 shows the most commonly used phospholipids in the preparation of liposomes (data extracted from the Sigma-Aldrich (Burlington, MA, USA) and Avanti Polar Lipids (Alabaster, AL, USA) databases). Figure 1 shows the classification of liposomes according to their structures, sizes, compositions, and preparation methods.

Numerous liposomal formulations have been clinically approved for human use to treat cancer and other chronic diseases, and are currently available on the global pharmaceutical market. Table 2 shows FDA- and EMA-approved liposomal drug formulations ^[20][21][20,21]. The Orange Book identifies drug products approved on the basis of safety and effectiveness by the FDA [20]. The electronic medicines compendium (EMC) identifies drug products approved for human use in the UK and Europe [21].

Figure 3) [22]. Since targeting the overexpressed surface receptors of cancer cells in order to enhance cellular uptake and intracellular activity is a promising approach, several cell surface strategies have been developed so far, aiming towards achieving the targeted inhibition of these receptors ^[22][24][22,24]. The efficiency of active targeting and ligand receptor interaction is dependent on certain factors [25], such as (1) the extent of receptor expression level on tumor cells relative to non-tumor cells, (2) the availability of surface receptors on tumor cells, (3) the internalization rate, and (4) the heterogeneity of receptor expression in tumor cells. Active targeting can be achieved through the surface engineering of liposomes via decoration with aptamers (oligonucleotides), carbohydrates, glycoproteins, monoclonal antibodies (mAbs) and their fragments, peptides, proteins, or other small molecules adsorbed onto the liposomal surface ^[22][24][22,24]. Figure 3 shows the surface modification of liposomes for active targeting. Figure 4 shows the distinction between passive and active targeting. Table 3 shows some examples of liposomes and their ligands used for active targeting.

. However, despite all these advantages, PEG-functionalized liposomes showed some serious drawbacks. For instance, PEG demonstrates potential immunogenicity owing to the activation of complement in response to antibodies ^[42][43][42,43]. In fact, PEG-based delivery systems support the phenomenon of accelerated blood clearance (ABC), owing to the formation of anti-PEG immunoglobulin M (IgM) antibodies by the spleen after initial administration. Following the second administration, the anti-PEG IgM binds to PEG groups on the surface of liposomes, resulting in the activation of the complement system, which subsequently leads to opsonization of the liposomes by C3 fragments and consequently enhances the cellular uptake of liposomes by the Kupffer cells in the liver, which in turn greatly affects the drug’s bioavailability [43]. Even though the ABC phenomenon poses a significant challenge for certain drugs, it does not pose a critical problem for PEGylated liposomes used in cancer therapy, owing to the high lipid content of liposomes encapsulating anticancer cytotoxic agents [43].

The adsorption of active targeting moieties on the surface of liposomes has played a significant role in enhancing liposomal accumulation in cancer cells since this structural characteristic helps to increase the therapeutic index of the encapsulated drug, maximize on-target effects, and minimize off-target effects. Briefly, active targeting is a surface modification process where active targeting moieties are adsorbed onto the liposomal surfaces, which substantially help to recognize and bind specifically to target cells through ligand–receptor interactions [44]. Actively targeting liposomes are made by grafting moieties such as aptamers, carbohydrates, glycoproteins, mAbs and their fragments, peptides, proteins, and small molecules adsorbed onto the liposomal surfaces (Figure 3). The targeting moiety can be either inserted directly into the lipid membrane or attached specifically to the distal end of the polymer ^[38][45][38,45]. Aptamers are short single-stranded oligonucleotides of deoxyribonucleic acid (DNA) or ribonucleic acid (RNA) with typically 20–100 base pairs. They have been widely studied as targeting agents to deliver various drugs and photosensitizers with high specificity to cell receptors and binding affinity at the nanomolar level [46]. Aptamers can be easily synthesized chemically and modified for conjugation. Different aptamers recognize different molecular targets. Cell-type-specific aptamers can be designed and synthesized by the exponential enrichment (SELEX) method, which can generate cell-type-specific aptamers with a high target affinity [47]. Transferrin is a blood plasma glycoprotein that plays a key role in iron metabolism by binding iron and transporting it into cells via TFR-mediated endocytosis [48]. Tumor cells often require excess amounts of iron for fast proliferation; thus, transferrin receptors on tumor cells are overexpressed in order to increase the uptake of iron from the plasma ^[48][49][48,49]. Folate (folic acid) is an essential vitamin for DNA synthesis and cell division. Folate and its derivatives have a high affinity for FRs and are internalized by FR-mediated endocytosis [50]. The receptor–ligand interaction affects the rate of cellular internalization, which in turn influences the accumulation of liposomal drugs in the tumor cells. Folate has the advantages of fast internalization and intracellular recycling rate, which subsequently fastens the delivery of drugs and photosensitizers to tumor cells ^[50][51][50,51]. Antibodies can recognize tumor cells by binding to specific antigens overexpressed by tumor cells with high affinity. Antibody targeting has been extensively studied for the delivery of drugs and photosensitizers to tumor cells; however, the therapeutic uses of mAbs are limited owing to antibody recognition and immunologic response [52]. Antibody fragments, such as single-chain variable fragments (scFv) and antigen-binding fragments (Fab), have been developed to modulate the overexpression of specific membrane antigens. Antibody fragments are smaller in size than mAbs, which allows them to efficiently reach their targets [52]. Although peptides have low molecular weights and a lower immunogenicity than antibodies, which lead to improved tumor penetration, they have a lower binding affinity. Peptides can be easily synthesized chemically and modified to obtain the desired physicochemical and biopharmaceutical properties. However, the presence of PEG polymers on the liposome surface may prevent peptides from interacting with cells. Therefore, peptides should be projected away from the liposome surface to avoid the shielding of the PEG polymer [53]. To conclude, the proper selection of targeting ligands is necessary to achieve high binding affinity and specificity. A recent trend in liposome surface functionalization includes the decoration of the liposomal surface with two ligands (i.e., dual-targeting). Dual-targeted liposomes offer numerous advantages, such as targeting multiple receptors, delivering more than one drug to target sites, enabling the encapsulated drugs to exert enhanced therapeutic effects, and reducing normal tissue toxicity and damage ^[54][55][54,55]. Two strategies are commonly used for dual-targeting based on different ligand combinations. The first strategy is based on the combination of two targeting ligands that leads to improved tumor selectivity and the cellular uptake of liposomes by specific tumor cells. The second strategy is based on the combination of targeting ligands with cell-penetrating peptides (CPPs) to achieve selective cellular uptake and enhanced anticancer efficacy toward specific tumor cells ^[55][56][55,56]. A dual-ligand combination involves the use of two ligands to target different receptors, which can be expressed only on one cell or on different cells. These ligands can be combined into one molecule or even placed separately on the surface of liposomes [56]. Common examples of dual-ligand combinations are those in which the second ligand is combined with either RGD, HA, or transferrin, since those ligands are overexpressed on different tumor cells [56]. In general, dual-ligand combinations can be classified into three different classes based on the types of target cells and the target sites of action. In the first class, two ligands are used to target one cell type, which is overexpressed on two receptors. The second class involves the use of two ligands to target two cell types, while the third class combines cell membrane targeting with intracellular organelle targeting (such as nuclear targeting or mitochondrial targeting) ^[55][56][55,56]. In fact, surface-modified liposomes with one or more targeting ligands exhibit improved cellular uptake compared with conventional liposomes due to receptor-mediated endocytosis. However, the cellular uptake efficiency of liposomes by tumor cells is still limited since the receptors on the surface of tumor cells dynamically change with tumor progression, leading to the saturation of receptor–ligand binding ^[56][57][56,57]. The discovery of CPPs helped to overcome these barriers, especially the cell membrane barrier, and subsequently enhanced cellular uptake ^[57][58][57,58]. However, in vivo applications of surface-functionalized liposomes with CCPs are limited since the non-specificity of CPPs may penetrate normal cells and lead to systemic toxicity ^[59][60][59,60]. Furthermore, the surface-functionalized liposomes with CCPs may be easily recognized by the RES owing to the high density of surface positive charge [61].

Liposomes can respond to different internal and external stimuli and thus trigger the release of encapsulated drugs in a controlled manner to specifically target cancer cells. Internal stimuli include enzymes (such as cathepsin B enzyme, a lysosomal protease of the papain family, which is overexpressed in several pathological conditions and malignancies, including the brain, breast, colon, lung, and prostate cancers), pH (such as DOPE that adopts an inverted hexagonal phase II (H_II phase) at low pH and a bilayer structure (L_α phase) at neutral pH to promote the liposomal membrane permeabilization), redox (e.g., disulfide-containing liposomes that can be easily broken down by reducing glutathione at the tumor site, leading to the complete disintegration of the liposomal membrane), and temperature (i.e., temperature can act either as an external stimulus when heat source is applied from outside, or can be internal when cancer pathological lesions have a naturally elevated temperature) ^[62][63][62,63], while external stimuli include light, electrical fields, magnetic fields, and ultrasound waves ^[62][63][62,63]. Light, in the UV-visible-IR region, is a very promising tool for biological and medical applications due to its non-invasive nature, high spatial resolution and temporal control, tuneability over a wide range of wavelengths, convenience and ease of application, and robustness [64]. In comparison with other stimuli, light provides an unparalleled spatiotemporal modulation of molecular processes [65], making it highly suitable for clinical and therapeutic applications ^[64][65][64,65]. Table 4 summarizes the advantages and limitations of different types of stimuli. Light-responsive liposomes have been recently introduced as smart, intelligent drug targeting delivery systems to target drugs for specific sites with high spatial and temporal control over drug release [66]. These systems utilize nonionizing radiation and are mainly composed of biocompatible, biodegradable materials that can be straightforwardly tailored to the target sites for clinical and therapeutic applications [66]. Although most light-responsive liposomes respond to UV irradiation that has a poor tissue penetration and high phototoxicity, optical technologies like laparoscopic tools are now commonly used for reaching deeper located tissues ^[64][65][64,65]. On the other hand, NIR is safer for use, causes less cell damage, and has good tissue penetration. However, the lower energy of NIR may not be efficient enough to induce the desired drug release response from liposomes ^[67][68][67,68].

The tumor-targeted delivery of liposomes can be achieved by two main targeting mechanisms: passive and active targeting.

In the passive targeting mechanism, liposomes are transported through the tumor interstitium to the target cells through capillary fenestrations and channels through passive diffusion or convection [22]. The tumor angiogenesis induces irregularities in endothelial cells with different pore sizes, ranging from 100 nm to 2 μm [23]. The differences in pore sizes and size distributions between the tumor microvasculature of endothelial cells and the tighter structures of normal cells make liposomes more easily accessible to the cancerous sites. Additionally, liposomes exploit the enhanced permeability and retention (EPR) effect for tumor targeting by improving the amounts of drugs delivered to tumor sites [23]. In order to passively target liposomes to tumor cells, liposomes should possess some physical and structural characteristics, such as the following: (1) the size of the liposomes should be in the range of 10–100 nm; (2) they should carry a neutral or anionic charge in order to avoid renal elimination; and (3) they should be protected from the RES ^[22][24][22,24].

Site-specific drug delivery is a method of targeting drugs to specific sites in a manner that increases their therapeutic indexes and reduces their possible side effects and toxicities ^[22][24][22,24]. Liposomes can reach tumor sites passively through the EPR effect [23], while surface-modified (or surface-engineered) liposomes act by binding to specific receptors overexpressed by cancer cells, such as epidermal growth factor receptor (EGFR), folate receptor (FR), transferrin receptor (TFR), and other receptors (

Phospholipids play key roles in the stability of liposomes in the systemic circulation, liposomal encapsulation, drug loading efficiency, and drug release at target sites. For instance, the use of 1,2-distearoyl-sn-glycero-3-phosphocholine (DSPC) in the preparation of liposomes resulted in a higher drug encapsulation efficiency compared to 1-palmitoyl-2-oleoyl-sn-glycero-3-phosphocholine (POPC) and 1,2-dipalmitoyl-sn-glycero-3-phosphocholine (DPPC). This was mainly due to the lengthy fatty acid chain of DSPC as well as the rigidity of the acyl chains of DSPC [12]. Furthermore, the use of n-(methoxypolyethylene glycol 5000 carbamoyl)-1,2-dipalmitoyl-sn-glycero-3-phosphatidylethanolamine, monosodium salt (MPEG-5000-DPPE-Na) prolonged the blood circulation time of liposomes, owing to the additional steric hindrance of MPEG-5000-DPPE-Na, which reduced the liposomal uptake by the reticuloendothelial system (RES) [13]. Additionally, 1,2-dipalmitoyl-sn-glycero-3-phosphoglycerol, sodium salt (DPPG-Na) exhibited fusogenic activity that improved the ability of liposomes to cross the cell membrane [14]. In general, liposomes can enter cells either by endocytosis (i.e., the process of capturing liposomes from outside by engulfing them with the cell membrane) or via exocytosis membrane fusion (i.e., the process where two phospholipid bilayers merge into a single continuous bilayer). Anionic liposomes showed faster endocytosis that enhanced their intracellular uptake, while fusogenic liposomes demonstrated an ability to fuse and penetrate the cell membrane ^[2][9][2,9]. Fusogenic liposomes are a particular type of liposome that are capable of causing fusion with biological membranes, thereby improving cell-type-specific delivery and therapeutic efficacy. They are mainly composed of phospholipids, such as dioleoyl-phosphatidylethanolamine (DOPE) and cholesteryl hemisuccinate (CHEMS) [15]. On the other hand, the use of cholesterol in liposomes aims to provide additional rigidity to the bilayer system in order to enhance liposome stability by increasing the molecular packaging of phospholipid molecules, prompting drug retention inside the bilayer system, and reducing the permeability of phospholipid bilayers [16]. In fact, cholesterol does not form bilayers by itself but will dissolve readily in the phospholipid–water bilayer system [10]. The unique feature of liposomes is their ability to compartmentalize and encapsulate both hydrophilic and hydrophobic drugs. This unique feature, along with biodegradability, biocompatibility, safety, non-toxicity, and targetability, made liposomes very attractive nanocarriers to maximize drug delivery and activity ^[2][9][2,9].

There are several methods for the preparation of liposomes (Figure 1) ^[2][9][2,9], such as thin-film hydration (or thin-layer evaporation), reverse-phase evaporation, double emulsification, ether injection, ethanol injection, and detergent removal. However, all these techniques are used for lab-scale production and mostly require the use of organic solvents in high concentrations and ratios. Moreover, they exhibit difficulty in controlling size and intercalation efficiency. Supercritical carbon dioxide (SC-CO₂) (Figure 2) is a novel technique suitable for the large-scale production of liposomes with the advantages of high encapsulation efficiency and uniform particle size distribution without the necessity of post-formation processes such as sonication or extrusion ^[17][18][17,18]. In SC-CO₂, CO₂ is premixed with lipids and then enters a chamber with atomized water droplets. As a result of the high diffusion ability of CO₂ and the reduced viscosity of the solution, lipids would coat water droplets at higher rates, resulting in inverted micelle-like structures, which are further stabilized by another layer of lipids placed at the bottom of the chamber [19].

Functionalized liposomes include long-circulating PEGylated liposomes, targeting ligand-functionalized liposomes, and stimuli-responsive liposomes (Figure 5).

Effective targeting requires the design of smart drug delivery systems with a long circulatory half-life (i.e., to remain in the bloodstream for longer periods of time), which means that liposomes must evade uptake by RES organs. This unique character can be imparted onto liposomes by coating their surfaces with polymers that suppress opsonization by plasma proteins [38]. These liposomes are commonly known as stealth or long-circulating liposomes, owing to their stealth properties that make them resistant to recognition and degradation by enzymes and immune systems ^[38][39][38,39]. The most commonly used polymer to prevent liposome opsonization is polyethylene glycol (PEG). The process of PEG attachment to liposomes is called PEGylation [39]. Of note, PEG, commercially known as macrogol, is a hydrophilic, biodegradable polymer with a general formula of H(OCH₂CH₂)_nOH, where n is the number of oxyethylene groups + 1 [40]. There are different grades of PEG available on the global pharmaceutical market, such as PEG-400, PEG-1500, PEG-4000, PEG-6000, and PEG-20000, that differ in their molecular weights. PEGs with molecular weights of 1000 and higher are solid grades and range in their consistencies from pastes to waxy flakes, while PEGs with molecular weights below 1000 are viscous liquid grades [40]. The adsorption of PEGs onto the liposomal surfaces leads to enhanced blood circulation time, reduced RES uptake, increased biodistribution and target accumulation, and enhanced formulation stability ^[41][42][41,42]