Liposomal Nanoparticulate Drug Delivery Systems: Strategies to Destabilize Biological Membranes at the Target Tissue

Abstract

Liposomal nanoparticulate drug delivery systems (LNDDSs) are clinically validated nanomedicine platforms seeing regular use in oncology and infectious disease. Their applications have rapidly expanded with several tissue targeting formulations in early-phase clinical trials. Beyond small molecular drugs, LNDDSs are increasingly employed for delivery of nucleic acid therapeutics, such as ribonucleic acid (RNA) based vaccines and immunomodulators. Recent advances in nanomaterials have enabled LNDDSs not only to transport therapeutic agents across systemic biological barriers but also to selectively destabilize plasma and organelle membranes, such as endosomes and mitochondria, addressing a wide range of diseases. This review systematically examines design strategies for LNDDSs that traverse key biological barriers focusing on the blood-tumor barrier, blood-brain barrier, and lymphatic transport barriers. We further explore approaches including fusogenic, pH-, redox- and, enzyme-sensitive and externally (ultrasound and thermal) triggered LNDDSs to facilitate internalization and membrane destabilization for specific organelle-targeting. The mechanisms and representative formulations and of membrane interactions, and clinical progress are discussed. Finally, the translational opportunities and challenges, and future perspectives for rational design of next-generation LNDDSs are addressed.

Introduction

Liposomal nanoparticulate drug delivery systems (LNDDSs) trace their origins to the pioneering work by Alec Bangham in 1960s, who first described liposomes during his studies on phospholipid bilayers.¹ This discovery marked the beginning of modern lipid-based nanocarrier research in drug delivery. Liposomes are self-assembled phospholipid-based bilayer vesicles that mimic the architecture of biological membranes with the ability to interact with and, in some cases, traverse physiological barriers. These biocompatible vesicles are capable of encapsulating hydrophilic and hydrophobic therapeutics and protecting them from premature metabolism. These unique properties immediately positioned liposomes, particularly nanosized liposomes, as versatile drug delivery platforms, driving the rapid evolution of LNDDSs into a cornerstone of modern nanomedicine.

Over subsequent decades, advances in lipid chemistry, vesicle preparation and surface engineering transformed LNDDSs into clinically validated nanomedicines following clinical trials began in the early 1980s.² The clinical approval of Doxil^® (liposomal doxorubicin) in 1995 represented a major milestone in oncology, demonstrating improved pharmacokinetics, and reduced systemic toxicity thus the translational potential of LNDDSs. The translation catalyzed continued innovation in nanoparticle design for tumor-targeted drug delivery.³ Meanwhile, their application expanded to infectious diseases, following the approval of AmBisome^® in 1997, a liposomal amphotericin B for systemic and invasive fungal infections. LNDDSs have recently evolved beyond primarily delivering small molecular drugs, to ribonucleic acid (RNA)-based vaccines and immunotherapies, marking a significant expansion in their clinical success.⁴ Currently approved LNDDSs based products are summarized in Table 1.

Table 1.

Examples of FDA-Approved LNDDSs-Based Nanomedicines for Various Clinical Applications

Product Name	Active Agent	Year of First Approval	Admin Route	Liposome Type/Particle Size
Cancer Therapy
Doxil/Caelyx	Doxorubicin	1995	IV	PEGylated (stealth liposome), 100 nm
DaunoXome	Daunorubicin	1996	IV	Small unilamellar liposome, 45–80 nm
Myocet	Doxorubicin	2000	IV	Non-PEGylated liposome
Marqibo	Vincristine sulfate	2012	IV	Sphingomyelin/cholesterol liposome, 100 nm
Onivyde (MM-398)	Irinotecan	2015	IV	PEGylated liposome, 80–140 nm
Vyxeos	Daunorubicin + Cytarabine (5:1)	2017	IV	Non-PEGylated liposome, 100 nm
THE 001	Doxorubicin	2025	IV	Thermosensitive liposome
Antifungal & Anti-infective products
AmBisome	Amphotericin B	1997	IV	Small unilamellar liposome
Arikayce	Amikacin	2018	Inhalation	Liposomal inhalation suspension
Pain Management
DepoDur	Morphine sulfate	2004	Epidural	Multivesicular liposome (DepoFoam)
Exparel	Bupivacaine	2011	Local injection	Multivesicular liposome (DepoFoam)
Vaccines
Epaxal	Hepatitis A virus	1993	IM	Virosome (liposome-based viral envelope mimic)
Inflexal V	Influenza virus antigens	1997	IM or deep SC	Virosome (cephalin-containing liposome with viral envelope glycoproteins anchored in the bilayer), 150 nm
Shingrix	Glycoprotein E + AS01 adjuvant	2017	IM	Liposome-based adjuvant system, 50–100 nm, unilamellar
Mosquirix	RTS, S/AS01 (antigens)	2021 (WHO recommendation)	IM	Liposome-based adjuvant for lymph node targeting,100–150 nm, unilamellar
Comirnaty®	COVID-19 S-Protein encoding mRNA	2021	IM	Lipid nanoparticle, evolved from liposomes, 60–100 nm

Abbreviations: IV, intravenous injection; IM, intramuscular injection; SC, subcutaneous injection.

As the field has evolved, the ability of LNDDSs to traverse or modulate biological membranes has emerged as a central determinant of targeted delivery of therapeutics to various tissues. Through rational LNDDS design, major biological barriers such as the blood-tumor barrier (BTB), blood-brain barrier (BBB), and lymphatic transport barriers (LTB) can be traversed or modulated, through either passive or active targeting approaches.^5,6 Several parameters of LNDDSs contribute to increased targeting efficiency: (i) particle size dictates vascular permeability and interstitial mobility;⁷ (ii) surface charge influences interactions with serum components and the endothelial luminal surface;⁸ and (iii) surface modification of LNDDSs, such as PEGylation, ligand conjugation, and incorporation of membrane-destabilizing moieties such as cell-penetrating peptides can further improve delivery across key physiological barriers.^9–11

Research has also advanced toward overcoming subsequent cellular and subcellular membrane barriers, which otherwise substantially restrict cellular drug penetration. Advanced LNDDSs are capable of traversing cell membranes via multiple endocytic and transcytotic pathways, followed by escape from endo-lysosomal compartments to enable intracellular drug release.^12,13 This is particularly significant for macromolecules such as nucleic acids, which rely on endocytic pathways to function,¹⁴ and have emerged as promising treatments for numerous diseases.¹⁵ Moreover, modification of LNDDSs with active-targeting moieties (eg, RGD peptides, folic acid) has proven as a promising strategy to induce cell membrane destabilization through receptor-mediated endocytosis (RME), thereby enhancing cellular uptake of the payloads.^16–18 Additionally, exploiting cationic ionizable lipids, such as 1,2-dioleoyl-3-dimethylammonium-propane (DODAP),¹⁹ and heptatriaconta-6,9,28,31-tetraen-19-yl 4-(dimethylamino)butanoate (DLin-MC3-DMA),²⁰ in LNDDSs has demonstrated the ability to facilitate endo-lysosome escape. These lipids remain stable at physiological pH but become protonated in the acidic endo-lysosome compartments, which induces membrane destabilization and promotes lipid mixing and membrane fusion, enabling cargo release to the cytosol while avoiding lysosomal degradation.^21,22 Other strategies include targeting specific organelles, such as mitochondria through the use of ligands including triphenylphosphonium (TPP⁺) and dequalinium (DQA).²³

This review examines current and emerging strategies by which LNDDSs are engineered to selectively cross and destabilize various biological barriers focusing on BTB, BBB, and LTB (Figure 1A) as well as cellular and intracellular barriers (Figure 1B). We summarize mechanistic insights into membrane interaction, fusion, disruption, and transcytosis of LNDDSs, and evaluate how lipid composition, surface modifications with ligands and endogenous stimuli-responsive functionalities contribute to controlled membrane destabilization. Additionally, we highlight the growing role of externally triggered destabilization, focusing on ultrasound and hyperthermia responsive LNDDSs that enable spatiotemporal control of membrane perturbation to promote localized on-demand drug release. Together, these approaches illustrate the evolving landscape of precision nanomedicine aimed at enhancing tissue-specific delivery, intracellular trafficking, and thus therapeutic efficacy. Other significant advances in LNDDSs, including strategies destabilizing dense extracellular matrix (ECM) in tumors, such as surface functionalization or the co-delivery of hyaluronidase, collagenase or matrix metalloproteinases (MMPs),^24,25 are beyond the scope of this review.

Figure 1.

Schematic illustration of multi-scale transport and membrane destabilization pathways of liposomal nanoparticulate drug delivery systems (LNDDSs). (A) Systemic biological barriers including (i) blood tumor barrier (BTB) where LNDDSs navigate a pathological microenvironment including the leaky vasculature, dense extracellular matrix (ECM), and high interstitial fluid pressure (IFP), (ii) blood brain barrier (BBB), and (iii) lymphatic transport barriers (LTB). (B) Cellular/intracellular organelle membranes destabilization.

Abbreviations: AME, Adsorption-mediated endocytosis; AMT, Adsorption-mediated transcytosis; RME, Receptor-mediated endocytosis; RMT, Receptor-mediated transcytosis.

Mechanisms of LNDDSs to Cross Various Bio-Membrane Barriers

LNDDSs bypass systemic barriers predominantly via paracellular and transcellular transport. Recent advances have substantially refined our understanding of how LNDDSs traverse or destabilize biological membranes. While classical mechanisms including paracellular transport, member fusion, endocytosis, and RME remain important; recent focus has shifted to transcytosis-mediated bio-membrane crossing and particularly incorporating targeting ligands, namely receptor-mediated transcytosis (RMT) which further increases the delivery efficiency (Figure 1).^26–29 This has been driven by insights into how LNDDSs can actively engage and modulate endothelial transport mechanisms. In parallel, membrane fusion, lipid exchange and transient membrane destabilization contribute substantially, especially for liposomes incorporating ionizable or fusogenic lipids.³⁰

Paracellular Transport

The paracellular pathway, via intercellular spaces between adjacent endothelial cells, is a key route for small-molecule transport. In healthy tissues, tight junctions effectively block paracellular transport, limiting the passage of LNDDSs across normal blood vessels. In contrast, the BTB permits paracellular transport of LNDDSs through the Enhanced Permeability and Retention (EPR) effect, enabling tumour accumulation by exploiting the abnormal and leaky tumour vasculature (Figure 1A). Since the effective pore size of the BTB is often smaller, requiring nanoparticles to be optimised within a narrow hydrodynamic diameter range (typically <150 nm) to achieve efficient tumour penetration. For brain targeted delivery, the healthy BBB remains largely impermeable to conventional therapeutics, BTB formation provides a critical opportunity for passive paracellular extravasation, allowing LNDDSs to partially overcome central nervous system (CNS) delivery barriers.

The EPR effects serve as the “gold standard” for designing tumour-targeted nanomedicines and drug delivery systems in the past decades. A major obstacle to consistent EPR-mediated drug delivery is the pronounced heterogeneity of the BTB across tumour types and stages, together with dense extracellular matrix (ECM) and elevated interstitial fluid pressure (IFP), which further hinder nanoparticle penetration and distribution.³¹ As a result, the translation from animal models to humans is one of the most debated “translation traps” in modern oncology.³²

Membrane Fusion

Fusogenic LNDDSs may directly merge their lipid bilayers with the plasma membrane, allowing efficient delivery of payload into the cytosol, while bypassing slower endocytic pathways and the subsequent endosomal trafficking.³³ Membrane fusion (Figure 1B), once thought to be minor, plays an expanded role in LNDDS trafficking. At the subcellular level, low pH-triggered fusion within endosomes is a major pathway for ionizable and polyethylene glycol (PEG)-sheddable LNDDSs, facilitating lipid mixing and efficient endosomal escape.^34–36 At the plasma membrane, liposomes can also undergo hemifusion or partial fusion events, promoted by cone-shaped or non-bilayer-forming lipids such as 1,2-dioleoyl-sn-glycero-3-phosphoethanolamine (DOPE). These lipids convert the lamellar phase to the highly fusogenic inverted hexagonal phase (HII) and are often paired with cholesteryl hemisuccinate (CHEMS) to lower the energetic barrier for membrane deformation and fusion intermediates.³⁷

Of note, these fusion processes are often transient, causing localized membrane destabilization rather than complete rupture, and can be modulated by the protein corona (eg, apolipoproteins, immunoglobulins, and complement proteins), which affects both membrane affinity and uptake specificity.³⁸ Highly charged cationic liposomes can fuse with the cell membrane within seconds.³⁹ The fusion-mediated cargo delivery results in the incorporation of the foreign lipids, induces membrane curvature, reduces surface tension, and generates short-lived pores. However, once the liposomes are removed, cells rapidly return to their original state.³⁹ Examples of the most commonly used cationic lipids in LNDDSs include DOTAP (1,2-dioleoyl-3-trimethylammonium-propane) which forms stable lipoplexes with nucleic acids, and DOTMA (N-[1-(2,3-dioleyloxy)propyl]-N,N,N-trimethylammonium chloride), a first-generation cationic lipid used in Lipofectin^®.

Endocytosis

Endocytosis represents the predominant pathway by which LNDDSs enter cells. Upon reaching the cell surface, LNDDSs are engulfed by the cell membrane and internalized through membrane invagination, resulting in the formation of endocytic vesicles.⁴⁰ Endocytosis occurs via multiple distinct pathways, including adsorptive-mediated endocytosis (AME), RME such as clathrin- and caveolae-mediated, as well as other specific receptor-mechanisms, and macropinocytosis, each with unique cellular uptake characteristics. AME, also referred to as adsorptive-mediated transcytosis (AMT), occurs when cationic LNDDSs electrostatically interact with negatively charged cell-surface structures such as sialic acid residues on the apical surface of brain microvessel endothelial cells. This nonspecific electrostatic interaction promotes cellular uptake and has been widely exploited for transport across the BBB. Compared with RME, AME exhibits lower target specificity but higher uptake capacity.

Particle size plays a critical role in determining the dominant cellular internalization pathway. Nanoparticles in the range of ~100–150 nm are efficiently internalized via clathrin-mediated endocytosis, whereas smaller particles (~50–80 nm) may preferentially engage caveolae-mediated endocytosis. Larger nanostructures >150–200 nm can enter cells through macropinocytosis, a non-selective pathway characterized by membrane ruffling and large vesicle formation.

Clathrin-mediated endocytosis facilitates the delivery of cargoes encapsulated in LNDDSs via transmembrane receptors, such as transferrin and low-density lipoprotein receptors (LDLR).⁴¹ For example, DOTAP⁴² and 1,2-Dioleoyl-sn-glycero-3-phosphocholine (DOPC)-incorporated LNDDSs (~100 nm) are predominantly internalized via the clathrin-mediated endocytosis pathway.⁴³ Caveolae are cholesterol-rich invaginations of the plasma membrane that mediate the intracellular trafficking of LNDDSs. Caveolae-mediated endocytosis facilitates the delivery of smaller, neutrally charged, cholesterol-rich LNDDSs from the vasculature to tumor cells, especially in late-stage tumors characterized by extensive vascularization.⁴⁴ LNDDSs modified with folic acid or albumin are more likely to enter the cells through caveolae-mediated endocytosis by interacting with glycoprotein 60 receptor overexpressed on caveolae.⁴⁵ On the other hand, the cellular uptake of negatively charged LNDDSs also relies on caveolae-mediated endocytosis, as exemplified by the clinically approved liposomal doxorubicin, which contains a high cholesterol content (~38 mol%) and exhibits a zeta potential of approximately −2.6 mV.⁴⁶ As a complementary pathway, macropinocytosis facilitates the cellular internalization of submicron- and micron-sized LNDDSs.

Of note, in many cases, a single LNDDS may simultaneously utilize multiple endocytic pathways, possibly due to the heterogeneity of particle size and lipid composition. For example, clathrin-mediated endocytosis accounted for only 46% of the cellular uptake of PEGylated pH-sensitive liposomes (DOPE/CHEMS) with an average size of 130 nm.⁴⁷

Active tumor-targeting strategies have demonstrated strong potential for enhancing tumor specificity by exploiting RME through ligands (adaptors, RGD peptides, folic acid) that recognize overexpressed receptors on target cells.⁴⁸ As an example, hyaluronic acid (HA) functionalized liposomes bind to CD44 receptors, promoting cancer cell-liposome membrane fusion and enhancing chemotherapeutic outcomes in pancreatic cancer.²⁶ Similarly, anti-CD44 adaptor, the intracellular protein that binds to the cytoplasmic tail of the CD44 transmembrane glycoprotein, was conjugated on PEGylated liposomes with thiol-maleimide click reaction, conferring selective recognition of CD44-overexpressing lung and breast cancer cells.⁴⁹

To target the folate receptors on breast cancer cells, De Oliveira Silva et al²⁷ fabricated folate-modified LNDDSs. This targeting strategy markedly prolonged vehicle systemic circulation and enhanced receptor-mediated endocytosis, ultimately achieving 68% tumor growth inhibition and a substantial reduction in pulmonary metastasis. Similarly, Bilgicer’s group modified liposomes with HER2- or VLA-4-antagonistic peptides to facilitate targeting of HER2- or VLA-4-overexpressing breast cancer and myeloma cells, achieving approximately 9-fold and 100-fold enhancements in cellular uptake, respectively.⁵⁰ The exploitation of biomimetic cell membranes, such as cancer-, macrophage-, and natural killer (NK)-cell membranes,^51–53 represents another successful strategy for membrane fusion-mediated targeting, in which homologous adhesion molecules on the membrane surface enable self-recognition and preferential interaction with cells of the same origin.

Following cellular uptake through membrane fusion and/or endocytosis, LNDDSs are initially trafficked to early endosomes, which act as sorting networks. Partial LNDDSs are transported to late endosomes and subsequently sequestered in lysosomes, where enzymatic degradation significantly reduces the payload bioavailability.⁵⁴ Rationally engineered LNDDSs can exploit the progressive acidification of the endosomal compartment to trigger pH-responsive destabilization of the lipid bilayer, or other pathways, thereby enabling “endosomal escape” before lysosomal sequestration.⁵⁵

Transcytosis

Following endocytosis, LNDDSs may be transported across the cytoplasm and exocytosed on the opposite side (Figure 1A). LNDDSs containing DOPC may be redirected from early endosomes to the Golgi apparatus through recycling endosomes,⁵⁶ thereby being expelled via exocytosis. This “transcytosis” process allows lipid vesicles to cross cellular barriers intact cross polarized cells, such as those in the BBB and intestinal epithelium, transporting their payload from the apical (luminal) to the basolateral (abluminal) side.^57,58 Increasing evidence highlights the importance of these active, dynamic transport mechanisms that navigate the restrictive architectures of pathophysiological barriers particularly the BTB and BBB, positioning transcytosis as the primary gateway governing the interaction of LNDDSs with complex membrane environments.

A major emerging theme is the dominant role of AMT and RMT followed AME and RME, respectively.⁵⁹ AMT is a non-specific process driven by electrostatic interactions,⁶⁰ where mildly cationic or charge-switching liposomes transiently engage the endothelial glycocalyx to facilitate transport.⁶¹ RMT exploits receptors expressed on the luminal surface of endothelial cells or epithelium, such as transferrin, insulin, LDLR, lipoprotein receptor-related protein 1 (LRP1) and transmembrane protein 30A (TMEM30A), and therefore is highly specific. Successful transport requires optimization of ligand density and binding avidity to promote transcytosis while avoiding endosomal sequestration and lysosomal degradation.^29,62,63

In oncology, the recently proposed active transport and retention (ATR) concept,^64,65 based on transcytosis across the BTB, has been demonstrated to complement the EPR effect, which occurs on permeable tumoral vasculature via paracellular pathways, achieving tumor accumulation of LNDDSs. Furthermore, emerging evidence also reveals that LNDDSs can actively stimulate endothelial transcytosis machinery (eg, upregulating caveolin-1 or inducing vesiculo-vacuolar organelle formation),^64,66 suggesting the process is not purely passive. A novel LNDDS co-modified with a cell-penetrating peptide (octaarginine) and a Golgi-targeting moiety containing cysteine (Cys) achieved significant transendothelial transport across the BTB and high tumor accumulation leveraging the ATR mechanism.^57,64

Bio-Membrane Destabilization at Target Tissues by LNDDS Design

Understanding the mechanisms by which LNDDSs traverse diverse systemic biological membranes, such as BTB, BBB and LTB, provides critical guidance on how to design these for enhanced delivery efficiency. The BTB represents a critical physiological barrier to effective LNDDS delivery following intravenous (IV) injection in cancer therapy. For neurological diseases, nearly 98% of small molecule drugs are unable to cross the BBB, a highly selective interface for exchange between blood and brain.⁶⁷ For cancer immunotherapy and vaccines for infectious diseases, where subcutaneous (SC) and intramuscular (IM) administration are more commonly used to target lymph node or spleen, LNDDSs must overcome the LTB to achieve efficient immune activation.⁶⁸ In this section, we comprehensively discuss LNDDSs design strategies (Table 1) to overcome bio-membrane barriers at target tissues, including solid tumors, the brain, and the lymphatic system (Figure 1).

Blood-Tumor Barrier

LNDDSs cross the BTB primarily through two transport pathways: passive extravasation across leaky neo-vasculature via the enhanced EPR effect, and ATR effect (particularly RMT) across the tumor endothelium.^69,70 However, they cannot easily enter and penetrate tumors through the EPR effect due to the excessive extracellular matrix (ECM), interstitial fluid pressure (IFP) and limited vascular extravasation in solid tumors (Figure 1A).⁷¹

In addition, the tumor microenvironment (TME) is a highly heterogeneous and dynamic network of diverse cells (tumor cells, fibroblasts, immune cells, and endothelial cells) embedded within the excessive ECM and nourished by abundant vasculature.⁷² Solid tumors are characterized regionally by acidic pH (6.5–6.9), hypoxia, and aberrant enzymatic activity that restricts LNDDS tumoral accumulation.⁷³ To overcome these obstacles, recent LNDDS designs have focused on selectively destabilizing biological membranes by leveraging stimuli (low pH, redox, or enzyme)-responsive properties^74,75 that trigger controlled drug release within the pathological milieu (Figure 2), allowing free drug to diffuse further across the tumors. Moreover, this review discusses external stimuli, including ultrasound and thermal effects, as strategies to facilitate membrane destabilization. Meanwhile, co-delivery of hyaluronidase, collagenase or matrix metalloproteinases (MMP) destabilizing tumoral ECM has also been explored.²⁵

Figure 2.

Rational LNDDS design that overcomes biological barriers by leveraging stimulus (pH, enzyme, redox, thermal, and ultrasound)-induced membrane destabilization to enhance intracellular drug delivery.

Additionally, pH-sensitive LNDDSs have been established to enhance cancer chemo-immunotherapy by overcoming fibrosis-mediated biological barriers and counteracting the immunosuppressive TME. Zhou et al engineered a doxorubicin-loaded dendritic macromolecule (PAMAM-ss-DOX) encapsulated within pH-sensitive LNDDSs for pH-triggered co-delivery of an immunoadjuvant R848 and losartan, a vasodilating drug to alleviate tumor fibrosis.⁷⁶ While doxorubicin-induced immunogenic cell death substantially reversed the suppressive immune TME, losartan significantly reduced IFP and deactivated tumor-associated fibroblasts, which collectively overcame the stromal barrier in solid tumors. Together, the combination within the LNDDS promoted robust T-cell infiltration to fight cancer.

Aberrant enzyme expression in TME, exemplified by MMP, β-galactosidase, phospholipases and cathepsin, has driven the rise of enzyme-cleavable LNDDSs that leverage enzyme-triggered membrane destabilization.⁷⁷ For instance, Liu et al⁷⁸ developed an MMP- and pH-dual responsive LNDDSs in which an MMP-cleavable linker enabled extracellular release of a PD-L1 inhibitor through MMP2-mediated cleavage, while a pH-responsive polymer stabilized the nanocarrier and facilitated endosome escape. Consequently, low-dose doxorubicin enhanced antitumor immunity by upregulating M6PR on tumor cells and sensitizing them to cytotoxic T lymphocytes and together with PD-1/PD-L1 blockade produced synergistic chemo-immunotherapy in B16F10 melanoma mouse models.

Hypoxia is a unique feature of solid tumors, arising from excessive oxygen consumption during rapid cancer cell proliferation and inadequate blood supply from disorganized vasculature, creating a significant barrier for drug delivery. It particularly compromises reactive oxygen species (ROS)-based cancer therapies, such as chemo- and photo- dynamic therapy, by impairing the cascade reactions required for oxidative damage.⁷⁹ To overcome these barriers, LNDDSs engineered with hypoxia-responsive, ROS-cleavable linkers or thioketal lipids have emerged as a promising approach to enable controlled cargo release. Yu’s group fabricated ROS-activatable liposomes incorporating egg-yolk L-α-phosphatidylcholine as ROS-cleavable linkers and thioketal-conjugated lipids to facilitate the delivery of photosensitizers and Fe₃O₄ nanoparticles.⁸⁰ Upon laser irradiation mediated ROS induction, the liposome membranes destabilize which enhances the Fenton reaction and alleviates tumor hypoxia, amplifying photodynamic therapy against colon cancer.

Blood-Brain Barrier

The BBB is a highly specialized interface between the CNS and the periphery that restricts drug permeability, consisting of endothelial cells and surrounding pericytes, astrocytes, neurons, and microglia within a vascular structure.⁸¹ LNDDSs have increasingly attracted interest to enhance targeting and accumulation within the brain. As illustrated in Figure 1A, well-designed LNDDSs cross the BBB via both passive pathways, such as diffusion and paracellular transport, and active pathways, including RMT and AMT, which are facilitated by surface ligands or cationic charge that promote endothelial uptake and transport into the brain.

To transport LNDDSs across the BBB, surface functionalization with targeting ligands, such as transferrin, lactoferrin, insulin, or apolipoprotein E (ApoE) mimetic peptides, to enable RMT, has been widely exploited.^82,83 For example, Ismail et al⁸⁴ engineered ApoE-peptide conjugated liposomal nanoparticles encapsulating dual-chemotherapeutic agents (temozolomide and artesunate) to enhance BBB traversal via LDLR-mediated transcytosis in temozolomide-resistant brain tumors. This LNDDS achieved deep tumor penetration, efficient drug release, and prolonged survival in a temozolomide-resistant glioblastoma model.

Positively charged LNDDSs (eg, cationic liposomes) can electrostatically interact with negatively charged glycocalyx on the luminal surface of the BBB, thereby enabling efficient AMT. For instance, a cationic liposome modified with transferrin-targeting peptides was designed as a “smart bomb” to cross the BBB and deliver CRISPR-Cas9 for targeted P-glycoprotein (P-gp) knockdown, enabling evaluation of drug pharmaco-resistance.⁸⁵ In brain endothelial cell-containing BBB models, this cationic liposome significantly enhanced cellular uptake and effectively inhibited the P-gp efflux.

Lymphatic Transport Barrier

LNDDSs reach target immune tissues primarily via slow passive diffusion through the dense interstitial matrix into lymphatic vessels, followed by entry into lymph nodes (LNs), a process that can be enhanced through LNDDS design.⁸⁶ The COVID-19 outbreak witnessed the emergence and rapid development of lipid-based mRNA vaccines, such as Moderna’s mRNA-1273 and Pfizer’s BNT162b2, for mRNA delivery. These LNDDSs enabled vaccines to utilize lipid nanoparticles, evolved from liposomes, composed of ionizable lipids (SM-102 in Moderna, ALC-0315 in Pfizer), phospholipids, cholesterol and PEG-lipids to encapsulate and protect the mRNA from systemic degradation.⁸⁷ The mRNA vaccines are internalized into cells via endocytosis, where the endosomal membrane is destabilized, allowing the mRNA to be released into the cytoplasm and subsequently reach ribosomes to encode viral spike protein.⁸⁸ This S-protein is presented to antigen-presenting cells in lymphoid tissues, triggering both humoral and cellular immune responses.

As summarized in Table 2, various ligand-functionalized LNDDSs have demonstrated significant potential in augmenting cancer immunotherapy by actively directing vaccines to LNs and the spleen, where they facilitate immune cell uptake through membrane destabilization/fusion.^89,90 As an example, tecemotide (formerly Stimuvax) is a liposomal formulation encapsulating a mucin 1 (MUC1) peptide that advanced to Phase III clinical evaluation as a therapeutic cancer vaccine. It can trigger a cellular immune response against cells overexpressing MUC1 glycoprotein, such as in non-small cell lung cancer.⁹¹ Lipovaxin-MM is a liposomal vaccine that incorporates the dendritic cell-targeting domain DMS5000 to deliver tumor-associated antigens from MM200 patient-derived cell membrane fragments and has progressed to Phase I clinical trials.⁹²

Table 2.

Examples of Biological Membrane Destabilization at Target Tissues by Formulation Design

Biological Barriers	Hallmarks	Formulation Design	Payloads	Injection Route	References
BTB	Acidic TME; Dense ECM; high IFP	pH-sensitive liposomes; Cell penetrating peptides	Docetaxel/siRNA; Doxorubicin/R848/losartan	IV	[76,93]
	Enzyme dysregulation (eg, Cathepsin B, MMP, hyaluronidase)	Cathepsin B- cleavable peptide; MMP- cleavable substrate; pH-sensitive polymer	Gemcitabine; Doxorubicin/PD-L1 inhibitor	IV	[77,78]
	Hypoxia	ROS- cleavable linker/thioketal lipids	Oxaliplatin/Cypate/Fe3O4	IV	[80]
BBB	Tight junctions; Limited transcytosis; Efflux pumps (P-gp)	Transferrin targeting peptides; ApoE-peptide; Cationic liposome	SynO4 mAb; Temozolomide/artesunate; CRISPR-Cas9	IV	[83–85]
LTB	Limited immune cell targeting; Dense interstitial matrix; Slow lymph flow	Dendritic cells targeting ligands (Mannose; HA, LyP-1); HER2/neu-derived peptide; MUC1 peptide	TLR4 adjuvant; Tecemotide/cisplatin; Ce6	SC/IP/IV	[91,94,95]

Other targeting ligands, such as folic acid⁹⁴ and mannose,²⁸ have been employed to engineer LNDDSs, enabling interactions with specific receptors on immune cells, such as dendritic cells, macrophages, and T cells, enhancing robust immune responses across various cancer types. Zamani et al⁹⁵ reported a liposomal formulation composed of a HER2/neu-derived peptide, a toll-like receptor agonist, and a CD4⁺ T helper cell epitope, to target the spleen and activate cytotoxic T cells, ultimately inducing robust anti-tumoral immunity in HER2-positive breast cancer.

Crossing Cellular Membrane and Intracellular Organelle Targeting

When LNDDSs reach the target tissues, additional challenges occur in crossing cellular membranes and achieving precise intracellular organelle targeting. Strategies to enhance receptor-mediated cellular uptake, promote endo-lysosome escape, and selectively destabilize specific organelle membranes, have made significant progress.

Endo-Lysosome Escape

Many internalized liposomes are trapped in endo-lysosomes, requiring them to have efficient “escape” mechanisms to reach the cytosol before reaching a specific organelle. Efforts have increasingly focused on engineering endo-lysosome escaping LNDDSs that promote membrane fusion/destabilization, allowing them to bypass enzymatic degradation in lysosomes.

Inspired by the acidic environment within endo-lysosomes (pH 5.0–6.5), pH-responsive liposomes (PSLs) have been widely explored as an effective intracellular delivery strategy. A detailed discussion of the underlying mechanisms and material chemistries, including comprehensive coverage of pH-responsive designs such as acidic PEG-shedding and charge-reversal strategies, is provided in the review by Kanamala et al.⁷⁴

For example, a multistage PSL was fabricated using a dual pH-responsive peptide that remains charge-shielded and stable in the bloodstream (pH 7.4) but undergoes imine linker cleavage and histidine protonation in the acidic TME, reactivating the cell penetrating peptide for deep tumor penetration. Lower pH in endosomes triggered SA-H8-mediated endosomal escape and dissociation, ensuring efficient cytoplasmic siRNA release and achieving a superior tumor suppression.⁹³

In general, pH-sensitive LNDDSs achieve endosomal escape through three widely proposed mechanisms: (i) destabilization of the liposomal bilayer in the acidic endosomal lumen, leading to drug release; (ii) fusion between liposomal and endosomal membranes, enabling cytosolic delivery; and (iii) disruption or rupture of the endosomal membrane, resulting in payload release into the cytosol.⁹⁶

Fusogenic PSLs composed of DOPE/CHEMS destabilize the endosomal membrane upon internalization and release their contents into the cytoplasm, where the payload can then move towards the nucleus, or other organelles.⁴⁷ Protonation of lipids in backbone at the acidic lumen of endo-lysosomes is a widely used strategy in designing pH-sensitive LNDDSs and enabling endo-lysosome escape under an acidic TME. Incorporation of pH-sensitive moieties, such as the 3β-[N-(N′,N′-dimethylaminoethane)-carbamoyl]cholesterol hydrochloride, imidazole and poly(2-ethyl-2-oxazoline), enables protonation under acidic conditions, turning on their cationic charge.^34–36 This charge shift drives electrostatic fusion with the anionic endosomal membrane, inducing membrane destabilization and endosomal escape.

Compared to protonation-induced membrane fusion, the “proton sponge effect” is a buffering-driven mechanism mainly associated with polycationic polymers and amine-rich LNDDSs, leading to endosomal swelling and membrane disruption.⁹⁷ For instance, zwitterionic oligopeptide lipid-based multistage PSLs were engineered to undergo charge conversion in response to environmental pH. These liposomes exhibited a negative surface charge under physiological conditions and converted to a positive charge in acidic environments, thereby enhancing cellular uptake via electrostatic adsorptive endocytosis. Subsequent endosomal acidification induced a proton sponge-like effect, promoting endo-lysosome escape and cytoplasmic drug release, ultimately resulting in robust antitumor efficacy in xenograft renal carcinoma models.⁹⁸

Calcium acetate has been used as remote loading agent for weakly acidic drugs into liposomes. Additionally, in the work by Wu et al,⁹⁶ Ca²⁺ additionally enabled cytoplasmic delivery of a DNA-toxic agent SN25860 (acidic pKa 4.1) by facilitating intracellular release (Figure 3). When loaded in DOPE/CHEMS based PSLs, PSL-Ca²⁺ was accomplished by two sequent steps sequentially: 1) destabilization of PSL backbone in the acidic endo-lysosomal lumen, allowing release of Ca²⁺; and 2) free Ca²⁺ induced endosomal swelling, or even rupture via the “proton sponge effect”. Enhanced intracellular delivery of SN25860 by PSL-Ca²⁺ translated into 21- and 141-fold lower IC₅₀ values compared with non pH-sensitive liposomes (NPSL) and free drug, respectively. In a breast cancer mouse model, intratumoral drug concentrations 48 h after treatment with PSL-Ca²⁺ reached 56.3 µg/g, approximately 9-fold higher than those achieved with NPSL; 6.76 µg/g; p = 0.001). This corresponded to 0.47% and 0.06% of the total injected dose, respectively. This corresponded to the ex vivo clonogenic assay which demonstrated PSL-Ca²⁺ significantly suppressed proliferation and colony formation. Of note, there was a lack of endosomal “rupture” by NPSL-Ca²⁺ or PSL containing no Ca²⁺.

Figure 3.

Schematic illustration of a Ca²⁺-in-PSL system mediated endo-lysosomal escape via three-step process including Ca²⁺-facilitated proton sponge effect, and confocal microscopic evidence of endo-lysosomal rupture at 45 to 50 min. Enhanced intracellular delivery translated into 21- and 141-fold lower IC₅₀ values compared with NPSL and free drug, respectively, and significantly suppressed tumour cell proliferation and colony formation in an ex vivo clonogenic assay. Reproduced from ref96 with permission.

Mitochondrial Membranes

The mitochondrion is well established as the central regulator of numerous cellular processes, including energy generation, stabilization of ROS, and the signaling of cell survival and death pathways. It has emerged as a critical therapeutic target, particularly in cancer and increasingly in neurodegenerative diseases, where mitochondrial induction of apoptosis or modulation of metabolic function can provide significant clinical benefits.^99,100 However, targeting mitochondria is challenging due to the organelle’s double membrane and highly negative membrane potential (approximately −180 mV).

Selective targeting of therapeutic agents to mitochondria can be achieved through several approaches, with one of the most widely employed strategies involving surface conjugation of LNDDSs with triphenylphosphonium (TPP⁺, named as MITO-Porter)¹⁰¹ or dequalinium (DQA, named as DQAsomes) as a targeting moiety.²³ This delocalized cationic lipid, characterized by a positively charged phosphorus atom bonded to three phenyl groups, functions as a mitochondriotropic molecule, facilitating the penetration of the negatively charged mitochondrial membrane by creating a potential gradient across the mitochondrial surface, allowing it to accumulate inside mitochondria.¹⁰² Potential applications of DQAsomes, first described in 1998, involve mitochondrial gene therapy, antioxidant therapy, and apoptosis-based anticancer chemotherapy.¹⁰³ Building on this, TPP⁺-functionalized liposomes represent a versatile and increasingly sophisticated toolkit for mitochondrial nanomedicine, with growing relevance in oncology, neurodegeneration, and metabolic diseases.¹⁰⁴

Bio-Membranes Destabilization with External Stimuli

The controlled destabilization of biological membranes using external stimuli is a critical strategy for enhancing the delivery of LNDDSs to target tissues as it provides a spatiotemporal mechanism to disrupt both liposomal and cellular membranes.¹⁰⁵ The most commonly employed strategies here use either heat¹⁰⁶ or mechanical^107,108 effects to perturb membranes and enable tissue-specific and intracellular delivery (Table 3). Ultrasound is a commonly used technique to achieve either effect, although various stimuli, including light, magnetic fields and electricity, have also been explored.¹⁰⁵

Table 3.

External Stimulus-Responsive LNDDSs and Interaction Cellular Membrane Effect

External Stimulus	Working Mechanisms	Biomembrane Effects	Outcomes
Thermal trigger	Lipid phase transition increases membrane permeability	Facilitates endosomal escape	Controlled, site-specific drug release
Mechanical trigger	Cavitation induces lipid bilayer disruption; phase transition in thermosensitive lipids	Sonoporation, membrane ruffling, enhanced endocytosis	Rapid drug release, increased cytosolic delivery, enhanced uptake
	Energy increases liposome movement toward cells	Increased liposome-cell contact	Higher local concentration, improved targeting

Temperature-Based Destabilization

Thermosensitive LNDDSs are engineered using phospholipids with well-defined gel-liquid crystalline phase transition temperatures (Tc, typically 40–42°C), such as dipalmitoyl phosphatidylcholine (DPPC). When these systems experience temperatures above their Tc, they deform to release their cargo. The lysolipid component 1-Myristoyl-2-stearoyl-sn-glycero-3-phosphocholine (MSPC) has sometimes been incorporated into bilayers to increase their heat sensitivity. Liposomes prepared using this constituent released ~45% of their loaded doxorubicin within 20 s under mild hyperthermia.¹⁰⁹ Similarly, lipid-peptide conjugates such as leucine zipper peptides have been explored and have shown similar efficacies to other thermal-sensitive LNDDSs in in vivo tumor models.¹¹⁰

Several studies have demonstrated enhanced tumor targeting and therapeutic efficacy of chemotherapeutics delivered via thermosensitive liposomes. Regenold et al¹¹¹ reported that thermosensitive liposomal vinorelbine (ThermoVRL) combined with mild hyperthermia (39–43°C), exhibited rapid drug release within 1 min at temperatures ≥40°C (Figure 4A) and significantly higher tumor drug accumulation than free vinorelbine or non-thermosensitive liposomes (NTSL-VRL) (Figure 4B). Importantly, ThermoVRL treatment induced complete tumour remission in three out of five mice, which was maintained until day 150 post-treatment (Figure 4C).¹¹²

Figure 4.

Representative studies illustrating the mechanisms and therapeutic performance of thermosensitive liposomes. (A) Temperature-dependent vinorelbine (VRL) release from thermosensitive liposomes. (B) Blood and tumor accumulation of free VRL, thermosensitive liposomal VRL (ThermoVRL), and non-thermosensitive liposomal VRL (NTSL-VRL) in Rh30 tumor-bearing SCID mice following mild hyperthermia (HT), expressed as percent injected dose per mL blood (%ID/ml) or per g of tumors (%ID/g). (C) Kaplan-Meier survival analysis of Rh30 tumor-bearing SCID mice treated with saline, free VRL, NTSL-VRL, or ThermoVRL with or without HT. * Indicates the significant increase in median survival times when HT is combined (p < 0.05; n = 5). Reproduced from refs^111,112 with permission from Elsevier.

ThermoDox^®, a doxorubicin-loaded thermosensitive liposome, is the most clinically advanced system in this field. In Phase III HEAT and OPTIMA trials for hepatocellular carcinoma, treatment with the thermosensitive liposomes was combined with radiofrequency ablation although it was seen to fall short in both trials as it was unable to extend progression free survival in patients.¹¹³

Mechanical Destabilization

Phospholipid bilayers may be either transiently or permanently disrupted via application of mechanical stimuli such as ultrasound. The technique has long been explored as a means to enhance tissue permeation, where the applied vibrational energy both disturbs bilayers and provides a radiation force to push drugs or other molecules into the cytoplasm.^107,114 This disruption is also known to enhance uptake pathways including endocytosis.¹¹⁵ With ultrasound being the predominant mechanical stimulus explored, this section will focus on phenomena observed with ultrasound energy.

Mechanistically, ultrasound exerts three notable effects that have been noted to improve liposomal drug delivery: heating, streaming and cavitation. Heating has already been discussed and is the result of the mechanical energy being absorbed by and lost into tissue. The heating can also be localized onto LNDDSs and if the generated heat exceeds the Tc of an LNDDS component, this may disassemble the vesicle and trigger drug release.^116,117

Acoustic streaming can be used to move liposomes within the fluid towards cells and has been explored to encourage LNDDS interactions with cellular surfaces. A recent study demonstrated that doxorubicin-loaded liposomes could be internalized into HeLa cells intact via Gigahertz frequency acoustic streaming while bypassing the endosomal pathway.¹¹⁸ The authors argue that this technology could enable the targeting of liposomes into the cell nucleus. This study provides a rare instance where the streaming effect has been used in isolation to enhance cellular uptake of LNDDSs.

Finally, by far the most well explored mode of mechanical bilayer destabilization is cavitation, which is the spontaneous formation and oscillation of gas bubbles in a medium in response to oscillating pressure waves. Cavitation has been used to transiently create pores in cell membranes to promote internalization.^107,115 Similarly, cavitation may be used to open liposomal bilayers and encourage the release of encapsulated drug. Although ultrasound-triggered drug release from intact LNDDS has been reported,^117,119 complete LNDDS disruption may be more favorable to ensure maximal drug release and availability.

The specific phospholipid used impacts ultrasound responsiveness, for example LNDDSs made with unsaturated DOPC have shown more ready ultrasound-triggered release than those made with 1,2-Distearoyl-sn-glycero-3-phosphocholine (DSPC).¹¹⁹ The same study found that incorporating cholesterol in the bilayer reduces ultrasound responsiveness. This is in line with several studies that note more stabilized bilayers are less responsive to ultrasound.^108,120 PEGylated lipids, which are typically included to evade opsonization, have also previously increased the ultrasound sensitivity of LNDDSs.^121,122

The core of LNDDSs may also be modified to increase mechanical responsiveness. Approaches here include encapsulation of either a gas to form a micro/nanobubble or a volatile liquid to form a nanodroplet,¹⁰⁸ although it is debated whether these systems possess a liposomal or micellar structure. If gas is already present inside the LNDDS, cavitation and therefore drug release can occur more readily. For example, when LNDDSs were prepared with and without encapsulated argon, US exposure released over twice as much 5-fluorouracil from the gas containing formulations than their non-gas containing counterparts.¹²³

Similarly, volatile liquids convert into gas following administration of ultrasound, with the process termed acoustic droplet vaporization.¹²⁴ The gas rapidly expands and disrupts the LNDDS to trigger release. In many cases, LNDDSs tend to be tethered to the micro/nanobubbles¹²⁵ or nanodroplets¹²⁶ to mitigate any interactions between the ultrasound-responsive excipients and the drug to be delivered.

Dynamics of LNDDSs Following Different Administration Routes

Drug administration routes may significantly affect the biodistribution and trafficking fate of LNDDSs within the patient’s body, where they encounter specific biological barriers before reaching the target tissues. In this section, we comprehensively compare the transport pathways of LNDDSs to target tissues, as well as the strategies to overcome these biological barriers.

Intravenous (IV) Administration

IV delivery of LNDDSs is the most commonly used route for achieving fast systemic drug exposure, while maintaining the nanostructures to overcome the cellular barriers.¹²⁷ Once injected, LNDDSs encounter a dynamic layer of plasma proteins, which adsorb onto their surface, creating a protein corona (Figure 5).¹²⁸ The physicochemical properties and targeting capabilities of the LNDDSs may shift and ultimately diminish therapeutic efficacy.¹²⁹ When opsonins are adsorbed onto the LNDDSs, they are rapidly recognized and phagocytosed by mononuclear phagocyte system cells (MPS), leading to rapid drug clearance.¹³⁰

Figure 5.

Challenges for LNDDSs upon intravenous (IV) injection. After IV, plasma proteins form a Corona on LNDDSs, triggering opsonization and clearance by the mononuclear phagocyte system (MPS). PEGylated liposomes prolong circulation by reducing protein adsorption but can trigger anti-PEG antibodies, limiting therapeutic efficacy.

To alleviate these systemic clearances, coating LNDDSs with hydrophilic polymers, such as PEG, has been employed to provide liposomes with “stealth” properties, thereby extending their circulation time in the bloodstream.¹³¹ The stealth effect induces immune evasion through decreasing protein adsorption and thus reducing the uptake and clearance by the MPS.¹³² Unfortunately, anti-PEG antibodies are induced following repeated administration of PEGylated LNDDSs (Figure 5), leading to hypersensitivity reactions.¹³³

To mitigate PEGylation-based challenges, researchers are actively exploring innovative strategies, including optimization of PEG chain length, surface density, and the development of cleavable or exchangeable PEG moieties in the LNDDSs.^134,135 For instance, lipid nanoparticles modified with fast-shedding PEG-lipid (short acyl chain) enhance cellular uptake while also reducing the generation of anti-PEG antibodies compared to those modified with a slow-shedding PEG-lipid (long acyl chain).¹³⁶

Subcutaneous (SC) and Intramuscular (IM) Administration

SC injection is the preferred route for delivering vaccines and immune stimulatory agents to elicit long-lasting immune responses.^137,138 LNDDSs are deposited within the interstitial space, where they gradually diffuse through the dense interstitial matrix before draining into LNs.¹³⁹ By contrast, IM injection is routinely used in prophylactic and therapeutic vaccines. Muscle tissue acts as a local depot, allowing slow diffusion and sustained release of the LNDDSs. The success of the COVID-19 mRNA vaccines highlights the suitability of IM administration for LNDDS-mediated mRNA delivery.¹³⁷

Owing to limited access to the bloodstream and high permeability of initial lymphatic capillaries, LNDDSs preferentially enter lymphatic vessels and subsequently traffic to draining LNs, where they are captured by immune cells for activating the immune response.¹⁴⁰ This process typically requires approximately 12 h, yet unfortunately less than 2% of the LNDDSs ultimately reach the LNs, with the vast majority remaining sequestered at the injection site.⁸⁶ A small fraction of LNDDSs that are not captured by immune cells can migrate to the bloodstream, enabling their redistribution to the spleen and other secondary immune organs.

Numerous physicochemical parameters can potentially influence lymphatic absorption and LNs uptake of the LNDDSs. Particle size and injection site are the principal determinants, whereas surface charge or PEGylation exhibits minimal influence on lymphatic trafficking.¹⁴¹ Compared with larger LNDDSs, smaller particles (<150 nm) traverse the interstitium and lymphatics more efficiently, enabling improved delivery to target tissues.¹⁴² The size of LNDDSs also affects the pattern of immune stimulation: larger vesicles (>250 nm) stimulate Th1 cell-mediated immunity by releasing IgG2a, and IFN-γ, whereas smaller particles preferentially elicit Th2 humoral immunity by promoting IL-5 and IgG1 secretion.^143,144

Owing to variations in tissue architecture and interstitial fluid flow, the injection site influences LNDDS distribution. Footpad injection leads to higher LNDDS uptake in LNs compared with flank injection, attributed to low fat-layer in rats and the elevated interstitial pressure that facilitates lymphatic absorption.⁸⁶ Muscles with rich vascularization, such as the deltoid or vastus lateralis, are typically selected for IM administration since their dense capillary networks support more rapid systemic absorption of LNDDSs.¹⁴⁵ Conversely, muscles with lower blood flow are often avoided, as reduced perfusion can lead to slower drug release and prolonged local retention, ultimately diminishing immunization efficacy.

Current Challenges and Emerging Opportunities

Traversing biological membrane barriers, including BTB, BBB, and LTB, is a primary challenge for the wide application of LNDDSs. The current trend in LNDDS formulation is to integrate receptor-mediated targeting with stimuli-responsive membrane destabilization, representing a promising strategy to overcome biological barriers. However, the delivery efficiency of LNDDSs to target tissues remains limited (<2%) following both IV and SC administration due to enzymatic degradation during systemic circulation.^86,146 Future LNDDS design should integrate protective polymeric architectures with tissue-specific responsive lipid components to enhance structural stability while enabling controlled membrane destabilization at the target sites, thereby improving spatiotemporal drug release and tissue selectivity.

Stability is a critical hurdle in the clinical translation of LNDDSs. Structural instability of LNDDS-based nanomedicines during storage, transportation and blood circulation can lead to premature drug leakage, liposome aggregation, or loss of targeting functionality, particularly for stimuli-responsive LNDDSs,^147–149 compromising therapeutic efficacy. Even the conventional PEGylated liposomal doxorubicin has been reported to exhibit particle aggregation, resulting in increased hydrodynamic size and impaired tumor-targeting.¹⁵⁰ An appropriate balance between dynamic membrane responsiveness and structural integrity is a key consideration in improving LNDDS stability. Rational formulation design by lipid selection and optimized surface engineering is essential to maintain long-term stability while preserving controlled membrane destabilization at target sites, thereby enabling reliable manufacturing and successful clinical translation.

Scale-up and manufacturing are other significant challenges for the industrial and clinical translation of the liposomal nanoparticulate. LNDDSs that demonstrate robust performance at the laboratory scale often exhibit altered physicochemical properties upon scale-up, including changes in particle size, encapsulation efficiency, and batch-to-batch variability.^151–153 Successful scale-up will require the development of reproducible, scalable, and regulatory-compliant manufacturing processes. Integrating LNDDS design with manufacturability considerations at an early stage will be essential to bridge the gap between proof-of-concept studies and clinically viable liposomal therapeutics.

Currently, rapid advances in personalized and precision nanomedicine are reshaping the potential of LNDDSs. Cationic liposomes and lipid nanoparticle-like formulations specifically engineered for RNA therapeutics are enabling highly individualized treatment strategies. Their capacity to encapsulate and deliver mRNA, siRNA or miRNA for various applications, and CRISPR-Cas components for genome editing positions these systems at the forefront of next-generation therapeutics.

Hybrid and biomimetic liposomes represent another promising direction, merging the structural stability of synthetic lipids with the biological sophistication of natural vesicles. Innovations such as cell-membrane-coated liposomes⁵⁷ and liposome-exosome hybrids enhance tissue and cell penetration alongside specificity.¹⁵⁴ These biomimetic LNDDS platforms are particularly well suited for navigating complex biological barriers and delivering diverse therapeutic payloads, including nucleic acids, proteins, and small molecules, opening opportunities for improved efficacy and reduced off-target interactions.

In parallel, artificial intelligence (AI) is poised to accelerate progress across the entire LNDDS development pipeline. AI-guided formulation design enables predictive optimization of lipid ratios, particle size, surface chemistry, and encapsulation efficiency, reducing reliance on empirical trial-and-error approaches.¹⁵⁵ Machine-learning models can also forecast nanoparticle-biological interactions, biodistribution patterns, and patient-specific responses, supporting the rational design of highly customized liposomal systems.¹⁵⁶ As computational tools become more deeply integrated into nanomedicine research, they are expected to streamline discovery, enhance reproducibility, and drive the creation of next-generation liposomal therapeutics with unprecedented precision.

Conclusion

The discovery of LNDDSs has evolved into a sophisticated field centered on precision modulation of membrane interactions to improve targeted drug delivery. Understanding and harnessing these mechanisms is now critical for developing next-generation liposomal nanomedicines capable of achieving selective, efficient, and clinically translational drug delivery systems. In this review, we focused on bio-membrane destabilization strategies of LNDDSs in specific tissues, particularly tumors with the BTB, neurological diseases with BBB, and immune-responsive tissues with the LTB. The mechanisms of liposome-cell membrane and intracellular organelle interactions that mediate membrane destabilization or fusion were explicitly described, thereby overcoming the BTB, BBB, and LTB and enabling endosomal escape for efficient drug release. We further discussed external stimuli, such as thermal and mechanical energy, that trigger destabilization of liposomal and cellular membranes, leading to controlled release of the payloads within the LNDDSs. Future research efforts should focus on prolonging systemic circulation time, achieving controlled release at target tissues, improving the stability and extending the shelf life of LNDDSs to enhance transportability and storage, thereby facilitating clinical and industrial translation.

Funding Statement

This research was funded by Cancer Society of New Zealand’s National Research Grant (CSNZ2326).

Disclosure

The authors declare no conflicts of interest in this work.