Nano-enabled delivery of anesthetics: mechanistic insights, technological advances and translational challenges

Abstract

The application of anesthetics constitutes a cornerstone of modern medicine, yet conventional agents are plagued by suboptimal pharmacokinetics and potential toxicity. Nanotechnology has emerged as a transformative paradigm, facilitating the design of sophisticated carriers that enable unprecedented precision in anesthetic delivery. This narrative review provides a critical synthesis of the field, advancing beyond conventional material cataloguing to deliver a comprehensive “bench-to-bedside” analysis. We systematically elucidate how the fundamental physicochemical properties of lipid-based, polymeric, and inorganic nanomaterials can be engineered to develop smart stimuli-responsive systems capable of on-demand analgesia and targeted central nervous system delivery. More importantly, this review rigorously examines three pivotal translational challenges, including long-term toxicology, immunogenicity modulation, and the navigation of complex regulatory pathways, that must be overcome. By translating clinical needs into well-defined material design criteria, this work establishes a multidisciplinary roadmap for researchers, encompassing materials chemists, pharmacologists, and clinicians, to collaboratively develop the next generation of nano-anesthetics with enhanced safety profiles and therapeutic efficacy.

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1. Introduction

Modern anesthesia is a cornerstone of surgical medicine, yet both local and general anesthetics are plagued by significant limitations in their effectiveness and safety profiles.¹ Fundamentally, these include a short duration of action that necessitates frequent redosing, poor tissue specificity leading to systemic toxicity, and narrow therapeutic indices that create significant safety risks.²

Local anesthetics (LAs), such as lidocaine and bupivacaine, provide rapid nerve conduction blockade but offer a short duration of action, typically lasting only minutes to a few hours. For instance, a conventional bupivacaine nerve block rarely exceeds eight hours, often necessitating repeated dosing or continuous catheter infusions that increase the risk of infection and local tissue damage.^3,4 A more critical issue is their poor tissue targeting. LAs rapidly diffuse from the injection site into systemic circulation, which not only shortens their therapeutic window but also leads to systemic absorption. Once in circulation, high LA concentrations can induce severe central nervous system (CNS) toxicity (e.g., seizures, loss of consciousness) and cardiotoxicity, such as myocardial depression and refractory arrhythmias.^5,6 Additionally, some LAs have been reported to exert direct cytotoxic effects on cartilage, tendons, and neuronal cells. General anesthetics (GAs), including both intravenous and inhalational agents, induce reversible unconsciousness by acting on the CNS but are associated with poor tissue selectivity.⁷ As small, lipophilic molecules, they must cross the blood–brain barrier (BBB), leading to their widespread distribution throughout the body. This nonspecific action suppresses vital physiological functions, causing common side effects such as respiratory depression and hypotension.⁸ Owing to their narrow therapeutic indices, slight dosage increments can be life-threatening, leading to severe complications such as arrhythmias and malignant hyperthermia.^9,10 Lipophilicity also causes GAs to accumulate in nontarget tissues, prolonging the recovery time. Furthermore, emerging evidence from animal studies indicates that exposure to GAs in early life may lead to neuronal apoptosis and long-term cognitive impairment,¹¹ whereas agents such as ketamine present risks of tolerance and abuse. Overall, the fundamental trade-offs between the efficacy and safety of traditional anesthetics result from poor pharmacokinetics and a lack of targeting. These trade-offs highlight the urgent need for innovative drug delivery strategies.

Nanotechnology offers a powerful strategic shift in anesthetic drug development, directly addressing the above-mentioned limitations.¹² By engineering nanocarriers with precise physicochemical properties, it is possible to control drug loading and release with unprecedented accuracy, offering a clear path to overcoming the deficiencies of traditional anesthetics.¹³ Nanoparticles (NPs) enable sustained and targeted drug delivery, thus reducing the dosing frequency and enhancing drug accumulation at the intended sites. Additionally, surface-functionalized NPs can cross biological barriers, including the BBB, facilitating targeted delivery to the CNS, improving efficacy, and minimizing systemic toxicity.¹⁴ The field has rapidly evolved to exploit a diverse palette of nanomaterials. Lipid-based nanoparticles, which leverage their high biocompatibility, represent the most clinically advanced platforms.¹⁵ In parallel, polymeric nanoparticles offer unmatched versatility and customizable biodegradability, allowing for precisely tailored drug release profiles.¹⁶ More recently, inorganic nanomaterials have introduced novel functionalities, such as magnetic guidance and photothermal conversion, enabling externally controlled, on-demand drug release that was previously unattainable.¹⁷ Furthermore, smart stimuli-responsive nano-delivery systems are advancing precise, intelligent drug delivery modalities, offering potential not only to enhance anesthetic efficacy but also to pave the way toward personalized anesthetic care.¹⁸ In essence, this diverse array of nanocarriers represents a clear therapeutic evolution: from clinically-established lipid vehicles, to highly adaptable polymeric systems, and ultimately, to smart inorganic platforms capable of on-demand, externally-controlled action.

Despite this promise, significant hurdles in biological safety, manufacturing, and regulatory navigation impede clinical translation.¹⁹ To bridge the gap between preclinical promise and clinical reality, this narrative review provides a comprehensive analysis of nano-anesthetics, including both local and general anesthetics, systematically summarizing recent technological advances, critically evaluating the key translational hurdles, and outlining future directions for development and clinical integration (Fig. 1).

Fig. 1 A graphical overview of the development and application of nanomedicine in anesthesia. This schematic illustrates the bench-to-bedside pathway for nano-anesthetics, from fundamental nanocarrier design to clinical implementation. The primary classes of nanocarriers, including lipid-based, polymer-based, and inorganic platforms, are engineered with specific mechanisms such as passive targeting in inflamed tissues, active targeting via surface ligands, and stimuli-responsiveness to triggers like light or magnetic fields. Following rigorous preclinical (in vitro and in vivo) and clinical evaluation, these nanomedicines are applied to enhance local anesthesia through sustained release, improve general anesthesia by traversing the blood–brain barrier, and revolutionize pain management with on-demand, controlled analgesia.

2. Mechanisms and advantages of nano-delivery systems

Carriers of this type act as platform technologies. The fundamental mechanisms that govern the behavior of these carriers include passive targeting, active targeting, stimuli responsiveness, and surface functionalization, and they are often universal across different therapeutic fields. While research into the specific targeting mechanisms for nano-anesthetics is still an emerging area, the foundational principles have been extensively elucidated in more mature fields like cancer nanomedicine and the delivery of genetic drugs.²⁰ The core strategies for navigating biological barriers and achieving organ- or cell-specific delivery are largely conserved, as elegantly summarized in Fig. 2.²¹ These include passive targeting driven by physicochemical properties, active targeting via surface ligand–receptor interactions, and leveraging endogenous pathways for cellular uptake. This mechanistic framework, although often discussed in the context of other diseases, provides a crucial conceptual roadmap for the rational design of next-generation targeted nano-anesthetics.²²

Fig. 2 Fundamental mechanisms of nanoparticle targeting for potential anesthetic delivery. (a) Passive targeting based on tuning the physicochemical properties (e.g., size, surface charge) of nanoparticles to promote diffusion across the endothelium and preferential accumulation in target tissues with fenestrated vasculature or leaky barriers, a principle that could be exploited for localized anesthetic delivery in inflamed or surgical sites. (b) Active targeting achieved through surface conjugation of targeting ligands (e.g., antibodies, peptides, aptamers) that exhibit high affinity and specificity for cognate receptors overexpressed on target cells, such as pain-sensing neurons or inflammatory cells, enabling precise delivery of encapsulated anesthetics or modulators of pain pathways. (c) Leveraging protein binding in the serum to indirectly enhance targeting, where nanoparticles interact with specific serum proteins that subsequently bind to receptors expressed in the target organ, potentially offering a strategy to direct anesthetic-loaded nanoparticles to specific tissues or cells involved in pain processing. Adapted with permission from ref. 21 (Dilliard SA, Siegwart DJ. Passive, active and endogenous organ-targeted lipid and polymer nanoparticles for delivery of genetic drugs. Nat. Rev. Mater. 2023;8(4):282–300. doi: 10.1038/s41578-022-00529-7. Copyright 2023, Springer Nature).

2.1. Passive targeting

Passive targeting refers to the natural accumulation of drug-loaded NPs at specific sites on the basis of physiological or anatomical characteristics. A well-known example is the enhanced permeability and retention (EPR) effect, which was first described in tumors, where NPs preferentially accumulate due to increased vascular permeability and impaired lymphatic drainage.^23,24 Passive targeting requires no special ligand modifications and instead relies on optimized carrier sizes and properties. Generally, particles smaller than 200 nm are suitable for enhanced vascular extravasation and retention in targeted tissues.²⁵ In anesthesia applications, passive targeting mainly enhances local anesthetic drug retention in targeted tissues. For example, when LAs are encapsulated within NPs or gels, their release kinetics are significantly slowed, preventing rapid diffusion away from the injection site.²⁶ One study demonstrated that bupivacaine-loaded poly(lactic-co-glycolic acid) (PLGA) nanoparticles extended the duration of analgesia from several hours, which is typical for free drug solutions, to approximately 48 hours during nerve block procedures.²⁶ This prolonged effect was attributed to the sustained release of the drug from the nanoparticles passively accumulating near the nerve. Similarly, postoperative inflammatory microenvironments, characterized by increased capillary permeability, may also enhance nanoparticle retention at surgical sites, thereby increasing the local concentration of analgesics.²⁷ However, it is crucial to note that the clinical translatability of the EPR effect is a subject of significant debate. As highlighted in recent critical reviews, the inherent heterogeneity of human tumors often leads to an unreliable or absent EPR effect, a factor implicated in the clinical failure of many nanomedicines that rely solely on passive targeting.²⁸ This translational challenge underscores the imperative for developing more robust active targeting and stimuli-responsive strategies to ensure effective drug accumulation, moving beyond a dependence on the often-unpredictable nature of passive targeting.

2.2. Active targeting

Active targeting involves modifying nanoparticle surfaces with specific ligands or recognition molecules that selectively bind receptors on targeted cells or tissues, effectively “guiding” the drug to precise sites.²⁹ Common targeting ligands include antibodies or their fragments, peptides, carbohydrates, and folic acid, all of which selectively interact with biomarkers overexpressed on target cells. For example, NPs conjugated with a transcriptional transactivator peptide (TAT) significantly enhance lidocaine penetration across nerve membranes and skin barriers, substantially increasing sensory nerve blockade efficiency and duration.³⁰ Similarly, NPs decorated with peptides containing the arginine–glycine–aspartic acid (RGD) sequence bind specifically to integrin receptors that are overexpressed on endothelial cells at inflammatory pain sites, increasing the local drug concentration.³¹ The primary advantage of active targeting is its ability to achieve a tissue-selective drug concentration following systemic administration, reducing off-target effects and the overall drug dosage. For instance, NPs designed to cross the BBB can preferentially bind neuronal or glial cells, enhancing CNS drug availability.³² Notably, targeting efficacy is governed by multiple parameters including target selection, ligand density, and binding affinity. With ongoing research into molecular pain pathway biomarkers, active targeting could facilitate precise drug delivery to critical pain-processing areas such as the spinal dorsal horn or peripheral nerves. A prime example of mechanism-driven active targeting is the treatment of bone metastasis, where the goal is to disrupt the vicious cycle between tumor cells and the bone microenvironment.³³ To achieve this goal, nanocarriers are often functionalized with bone-homing moieties such as bisphosphonates, which exhibit a high binding affinity for hydroxyapatite in areas of high bone turnover, thereby concentrating the therapeutic payload at the site of pathology.³³

2.3. Surface functionalization and prolonged circulation

Surface properties significantly influence the in vivo fate of nanocarriers, and functionalization techniques provide additional therapeutic advantages. Among these methods, PEGylation involves attaching polyethylene glycol (PEG) chains to the surface of nanoparticles. PEG, a hydrophilic and inert polymer, forms a protective hydration layer around NPs, effectively reducing protein adsorption and recognition by macrophages, thus prolonging systemic circulation.³⁴ For anesthetic agents delivered intravenously, such as GAs or opioid analgesics, longer blood circulation allows a higher concentration of the drug to target tissues before clearance by the reticuloendothelial system. Moreover, PEGylation also reduces immunogenicity and the risk of allergic reactions.³⁵ Another surface-modification approach involves the introduction of cationic groups, such as quaternary ammonium or amines, which make the resulting nanoparticles positively charged. Cell membranes and vascular endothelia are typically negatively charged; thus, positively charged NPs demonstrate enhanced adherence and cellular uptake via electrostatic attraction.³⁶ Cationic modifications can also improve penetration across biological barriers, including facilitating adsorptive-mediated transcytosis across the BBB, thus enhancing drug delivery into the CNS.³⁷ However, excessive cationic charge densities can increase serum protein adsorption or activate coagulation pathways, necessitating careful optimization to achieve a balance.

Other promising functionalization strategies include coating NPs with biomimetic materials, such as erythrocyte membranes or albumin, to improve biocompatibility and extend circulation time, or conjugating specific antibody fragments to enable active targeting. A particularly sophisticated biomimetic strategy is ‘homologous targeting,’ where nanocarriers are camouflaged with membranes derived from the specific cell type they are intended to target.³⁸ This approach allows the nanocarrier to inherit a full complement of native surface proteins, which can act as crucial signals to regulate immune cell interactions and mitigate clearance by phagocytes.³⁸ Concurrently, other cell-specific adhesion molecules on the membrane enable highly specific binding to the tissue of origin. The implications of this principle for nano-anesthetics are profound, suggesting a future where nanocarriers coated in neuronal or macrophage membranes could be engineered for unprecedented, specific targeting of nociceptive pathways or sites of inflammation.

2.4. Stimuli-responsive release

Stimuli-responsive nanodelivery systems are designed to release drugs upon specific internal or external triggers, enabling precise control over the timing and location of anesthetic drug delivery.³⁹ Different stages of anesthesia often require varying intensities and durations of drug effects; thus, dynamically adjustable release mechanisms can significantly increase precision in anesthetic management.

Stimuli-responsive systems can be categorized into two types: endogenous and exogenous. Endogenous stimuli leverage physiological or pathological factors, such as pH, enzyme concentration, or temperature. For example, inflammatory or injured tissues typically present mildly acidic environments rich in specific enzymes.⁴⁰ Acid-sensitive NPs remain stable in neutral blood conditions but rapidly disintegrate to release anesthetics when they encounter the acidic inflammatory microenvironment at surgical sites. For instance, the release of zinc ferrite NPs with loaded lidocaine is relatively slow at physiological pH but more rapid and complete when the pH is acidic.⁴¹ The chemical basis for this pH-responsiveness is typically achieved through two primary materials science strategies.⁴² The first involves the use of a polymer matrix containing ionizable groups. A classic example is chitosan, a polymer rich in amino groups. In the acidic microenvironment of inflamed tissue, these amino groups become protonated, leading to electrostatic repulsion that causes the polymer network to swell and release its encapsulated drug payload.⁴³ A more precise method involves covalently conjugating the drug to the nanocarrier via an acid–labile linker, such as a hydrazone bond, which is stable at physiological pH but hydrolyzes rapidly under acidic conditions to release the free drug directly at the site of action.⁴⁴ Thermosensitive hydrogels represent another endogenous-responsive system. They often employ polymers (e.g., hydrogels) that exhibit a specific temperature. Below this temperature, the polymer chains are hydrated and present as a solution, but above this temperature (usually the body temperature), they undergo a phase transition, becoming hydrophobic and self-assembling into a viscous gel, creating a depot that provides highly effective sustained drug release.⁴⁵ Enzyme-responsive systems utilize elevated concentrations of enzymes, such as lysozymes or matrix metalloproteinases, present in the tumor microenvironment to trigger rapid degradation of carrier matrices and subsequent drug release. While enzyme-responsive systems are well-established in other fields such as oncology, their application for anesthetic delivery remains largely unexplored but promising for future research.

Exogenous stimuli including light, ultrasound, or magnetic fields, allow clinicians precise, on-demand control over anesthetic release. The mechanism for these systems is rooted in the ability of specific functional materials to act as energy transducers. Photo-responsive nanocarriers incorporate photosensitive groups or photothermal agents, rapidly releasing drugs upon exposure to specific wavelengths of light.⁴⁶

The photothermal mechanism relies on materials such as gold, which converts near-infrared (NIR) light into localized heat via surface plasmon resonance. This heating can then trigger a phase transition in a co-loaded thermosensitive polymer, causing drug release.⁴⁷ More advanced systems combine this with a photochemical effect, where a photosensitizer generates reactive oxygen species to disrupt the carrier, or utilizes photocleavable linkers that break upon irradiation to directly release a conjugated prodrug.^46,48

Ultrasound-responsive systems exploit mechanical vibrations or heating induced by ultrasonic waves. The primary mechanism is acoustic cavitation, where gas-filled or phase-change liquid nanocarriers, such as perfluorocarbons, rapidly oscillate or vaporize in the ultrasound field, creating microshock waves that can physically disrupt the carrier matrix or transiently open biological barriers.^49,50 Magnetically responsive systems utilize magnetic NPs heated by alternating magnetic fields (AMFs). The physical mechanism involves the rapid relaxation of the nanoparticle's magnetic moment through Néel and Brownian relaxation processes, which dissipate energy as heat. This localized hyperthermia can then be used to melt a thermosensitive coating, such as gelatin, to release the encapsulated drug on-demand.⁴¹

3. Common nanocarriers in anesthetic drug delivery

The design and optimization of nanocarriers play essential roles in anesthetic drug delivery. Different nanocarrier types meet specific delivery requirements, such as improved solubility, targeted delivery, sustained release, and stimuli-responsive properties. In terms of the materials and mechanisms, the nanocarriers used in anesthesiology mainly include lipid-based, polymer-based, and inorganic nanomaterials. These systems have been extensively applied in local anesthesia, general anesthesia, and postoperative pain management.

3.1. Lipid-based nanocarriers

Lipid-based nanocarriers, which have excellent biocompatibility, controlled drug release, and low immunogenicity, have been extensively researched and widely applied in anesthetic drug delivery.¹⁵ Lipid-based systems (Table 1) mainly include liposomes, solid lipid nanoparticles (SLNs), and nanostructured lipid carriers (NLCs).

Table 1 Representative lipid-based nanocarriers described in the literature for anesthetic delivery

Nanocarrier type/key innovation	Loaded drug(s)	Preclinical tests	Key quantitative outcome(s)	Clinical stage	Ref.
Conventional liposomes (multilamellar vesicles)	Lidocaine	In vivo: Canine epidural nerve block	Prolonged epidural nerve blockade ∼3-fold vs. free drug (170 vs. 61 min) in a canine model via slow, diffusion-controlled release	Preclinical	53
Ionic-gradient liposomes (IGLs) (active loading)	Bupivacaine	In vitro: Trapping efficiency. Release and clearance tests	Increased trapping efficiency to 64–82%; extended nerve block >3-fold (∼6.5 h vs. ∼2 h) by trapping protonated drug via a pH gradient	Preclinical	57
		In vivo: Guinea pig sciatic nerve block (pin-prick test)
Multivesicular liposomes (MVLs) (structural innovation)	Bupivacaine	In vivo: Murine subcutaneous infiltration (hot plate test); Pharmacokinetics (drug retention at injection site)	Achieved a >15-fold increase in analgesic duration (26 h vs. 1.58 h) in mice, leading to the FDA-approved Exparel®	Approved	58
Hybrid liposomes (drug-in-cyclodextrin-in-liposome)	Ropivacaine	In vitro: Encapsulation efficiency; release kinetics; cytotoxicity (MTT assay)	Prolonged in vivo sensory block by 1.7-fold (from 180 to 300 min) with reduced cytotoxicity due to synergistic drug hosting	Preclinical	59
		In vivo: Murine plantar test (von Frey test)
Ultra-long-acting suspension (hydrophobic ion pairing)	Ropivacaine	In vitro: Drug release kinetics	Provided analgesia for >10 days in vivo by forming a poorly soluble drug-surfactant complex with a >5-day release profile	Preclinical	60
		In vivo: Murine sciatic nerve block (von Frey & hot plate tests)
Surface-modified liposomes (brain targeting)	Etomidate (GA)	In vitro: Release kinetics; BBB permeation model	Lactoferrin modification enhanced brain delivery of the GA, resulting in a faster onset of anesthesia and more rapid recovery	Preclinical	61
		In vivo: Rat general anesthesia (loss of righting reflex)
Solid lipid nanoparticles (SLNs) (solid matrix)	Articaine	In vitro: Encapsulation efficiency; Release kinetics; Cytotoxicity	Achieved an encapsulation efficiency of 65.7% and provided a controlled release profile with reduced cytotoxicity	Preclinical	65
Nanostructured lipid carriers (NLCs) (vs. SLNs)	Lidocaine + prilocaine	In vitro: Encapsulation efficiency; drug release behavior	NLC formulation demonstrated a significantly stronger and more prolonged in vivo anesthetic effect compared to the SLN formulation	Preclinical	78
		In vivo: Murine tail-flick test
Functionalized NLCs (transdermal delivery)	Bupivacaine	In vitro: Skin permeation (Franz cells)	Hyaluronic acid modification led to a 2.5-fold increase in cumulative drug permeation through the skin	Preclinical	77
		In vivo: Rat tail-flick test
NLC-in-film hybrid (dental anesthesia)	Lidocaine + Prilocaine	In vitro: Release kinetics; mucoadhesion	Hybrid nanofilm prolonged anesthesia for >7 hours in mice, a 3-fold improvement over the commercial EMLA cream	Preclinical	79 and 188
		In vivo: Murine tail-flick test

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3.1.1. Liposomes. Liposomes are classic lipid-based carriers consisting of one or more concentric phospholipid bilayers, forming uni- or multi-lamellar vesicles.⁵¹ The excellent biocompatibility and ability of these materials to serve as sustained-release depots have made them a cornerstone of nano-anesthetic research since their inception. The pioneering work of Gesztes and Mezei in 1988 demonstrated this potential, showing that a 0.5% liposomal formulation of tetracaine could provide effective topical anesthesia on intact skin for at least 4 hours, whereas a conventional cream was ineffective.⁵² This principle was further quantified by Mashimo et al., who reported that the epidural administration of 2% liposomal lidocaine in a canine model prolonged the duration of nerve blockade by approximately 3-fold compared with that of the free drug solution (170 ± 49.5 min vs. 61 ± 18.1 min, respectively).⁵³ The authors attributed this significant prolongation to the slow, diffusion-controlled release of lidocaine from the multilamellar lipid bilayers. However, these early successes also highlighted a key material science challenge: the performance of simple liposomes is highly drug-dependent. For example, de Araujo et al. reported that while encapsulating mepivacaine in large unilamellar vesicles prolonged sensory blockade by 1.3- to 1.7-fold, the same formulation provided no significant advantage for the more lipophilic bupivacaine.⁵⁴

Early liposomal formulations also suffer from significant limitations, including poor drug loading and rapid drug leakage during storage.^55,56 A major breakthrough came with the development of active loading techniques. Mowat et al. pioneered the use of a transmembrane pH gradient, creating so-called ionic-gradient liposomes (IGLs) that can actively accumulate weak-base anesthetics such as bupivacaine.⁵⁷ By establishing an acidic interior (pH 4.0) relative to the external physiological environment (pH 7.4), the uncharged form of bupivacaine diffuses across the lipid bilayer and becomes protonated and trapped within the liposome's aqueous core. This method increased trapping efficiency to a range of 64% to 82%. It also extended the duration of nerve blockade by more than three times in a guinea pig model. The half-maximal response recovery time for 0.75% bupivacaine increased from about 2 hours to about 6.5 hours. This work established that engineering the internal chemistry of liposomes was a key strategy to overcome their inherent limitations.

The clinical translation of liposomal anesthetics was ultimately realized through a structural innovation: multivesicular liposomes (MVLs). This technology, which forms the basis of the approved formulation Exparel®, consists of large, non-concentric lipid chambers that can encapsulate high concentrations of drugs with high stability. Preclinical studies by Grant et al. on an early MVL formulation of 2% bupivacaine demonstrated a dramatic prolongation of analgesic duration in mice, with a median duration of 26 hours compared with only 1.58 hours for the standard bupivacaine solution.⁵⁸ This greater than 15-fold increase in efficacy was directly correlated with prolonged drug retention at the injection site, where the drug was detectable for more than 96 hours, compared to less than 8 hours for the standard solution.

Recent innovations have focused on creating even more sophisticated hybrid liposomal systems to further refine release kinetics and enhance drug loading. Vieira et al. developed a ternary “drug-in-cyclodextrin-in-liposome” platform for ropivacaine, which leverages the synergistic effects of the two carriers to reduce cytotoxicity and prolong the in vivo sensory block by approximately 1.7-fold (from 180 min to 300 min).⁵⁹ More recently, Tan et al. utilized hydrophobic ion pairing (HIP) to create a micron-sized ropivacaine lipid suspension. By forming a stable, poorly soluble complex between ropivacaine and docusate sodium, this formulation achieved an in vitro release profile of over 5 days and provided analgesia for more than 10 days in vivo, revealing a state-of-the-art strategy for weekly, long-acting pain relief.⁶⁰ Furthermore, the application of liposomes has expanded beyond local anesthetics to include general anesthetics and brain targeting. Wu et al. demonstrated that modifying the surface of liposomes with lactoferrin, which is a ligand for receptors on the blood–brain barrier, notably enhanced the delivery of GA etomidate to the brain in rats, leading to a faster onset of anesthesia and quicker recovery.⁶¹

In conclusion, while liposomes offer proven biocompatibility and an established ability to prolong analgesia, their inherent limitations in stability and drug retention have driven a clear evolutionary path: transforming these classic passive carriers into precisely engineered hybrid and surface-functionalized systems for a new generation of targeted and controlled anesthetic therapies.

3.1.2. Solid lipid nanoparticles (SLNs). Introduced in the early 1990s as a second-generation lipid carriers, solid lipid nanoparticles (SLNs) were developed to overcome some of the stability and scalability issues associated with liposomes.⁶² SLNs are produced by replacing the liquid oil of a conventional o/w emulsion with a solid lipid, creating a solid particle matrix that is stabilized by surfactants.⁶³ This solid matrix offers superior physical stability and better protection for encapsulated labile drugs. Furthermore, their production is readily scalable via established industrial techniques such as high-pressure homogenization, which is a significant advantage over more complex liposomal preparation methods.⁶⁴

The solid matrix of SLNs has been shown to be effective for encapsulating various anesthetics. For example, de Melo et al. reported an encapsulation efficiency of 65.7% for articaine in SLNs, which, when incorporated into a hydrogel, provided a controlled release profile and reduced cytotoxicity compared with the free drug.⁶⁵ More strategically, SLNs offer a solution for delivering potent but highly toxic anesthetics that are difficult to formulate in conventional carriers. Dibucaine, for example, is a highly potent local anesthetic but carries a significant risk of systemic toxicity, with lethal toxicosis reported in both humans and animals.⁶⁶ Its bulky molecular structure and strong interaction with lipid bilayers also complicate its incorporation into traditional liposomal structures.^67,68 By entrapping dibucaine within a solid lipid matrix, SLN formulations were shown to reduce its systemic toxicity and enhance its bioavailability.⁶⁹

Despite these advantages, SLNs possess a critical, inherent flaw rooted in their materials science: instability due to lipid polymorphism. The solid lipids used to form SLNs can exist in different crystalline structures (polymorphs), such as the less-ordered, higher-energy α and β′ forms, or the highly ordered, low-energy β form.⁷⁰ During production via melt-homogenization, the lipid matrix often solidifies into less-ordered states, which have imperfections in their crystal lattice that can accommodate drug molecules. However, during storage, a slow polymorphic transition to the most stable β form can occur. This highly ordered, “perfect” crystal structure has minimal space for foreign molecules, leading to the gradual expulsion of the encapsulated drug from the particle matrix over time.⁷¹ This phenomenon not only compromises the long-term stability of the formulation but also leads to an uncontrolled burst release of the expelled drug. This fundamental material science bottleneck directly motivated the development of the next generation of lipid carriers designed to resolve this issue.

3.1.3. Nanostructured lipid carriers (NLCs). To address the fundamental stability bottleneck of SLNs, a second-generation solid lipid nanoparticle, the nanostructured lipid carrier (NLC), was developed in 2000.⁷² The key material science innovation of NLCs is the creation of a less-ordered, “imperfect” solid core. This is achieved by blending a solid lipid with a structurally incompatible liquid (oil), which disrupts the formation of a perfect crystal lattice upon cooling.⁷³ This amorphous, nanostructured core creates numerous defects and voids that can accommodate drug molecules, thereby significantly increasing the drug loading capacity and, most importantly, preventing drug expulsion during storage.⁶²

This refined design translates into demonstrably superior performance and versatility. NLCs have been successfully used to encapsulate a wide range of anesthetics, including benzocaine,⁷⁴ procaine,⁷⁵ lidocaine,⁷⁶ and bupivacaine.⁷⁷ Direct comparative studies consistently show that NLCs outperform SLNs. For example, in a study co-loading lidocaine and prilocaine, while both carriers achieved high encapsulation (>90%), the NLC formulation produced a significantly stronger and more prolonged in vivo anesthetic effect.⁷⁸ The robust NLC matrix also serves as an excellent platform for complex formulations. Ribeiro et al. developed NLCs for dental applications that sustained drug release for over 8 hours and subsequently incorporated them into a pectin-based hybrid nanofilm.⁷⁹ This NLC-in-film system prolonged anesthesia for more than 7 hours in mice, representing a 3-fold improvement compared with both the control film and the commercial EMLA cream, and also exhibited superior mucoadhesive and mechanical properties.⁷⁹

A major focus of NLC research is leveraging them for problem-driven design to address specific clinical challenges. A classic example is anesthesia failure in inflamed tissues, where local acidosis neutralizes traditional amine-based LAs. To circumvent this, Rodrigues da Silva et al. formulated the non-ionizable anesthetic butamben into NLCs. Compared with the standard anesthetic articaine, the resulting formulation not only has poor aqueous solubility but also has significantly greater and more prolonged analgesic effects in an inflammatory hyperalgesia model, demonstrating how NLCs can be used to engineer solutions for specific pathological microenvironments.⁸⁰

Furthermore, the stable NLC scaffold is an ideal platform for advanced surface functionalization to overcome biological barriers, particularly for transdermal delivery. Surface modification with hyaluronic acid (HA), for example, has been shown to leverage interactions with CD44 receptors on keratinocytes to enhance skin penetration.⁷⁷ In a bupivacaine-NLC system, HA modification led to a 2.5-fold increase in cumulative drug permeation through the skin compared with that of the free drug.⁷⁷ More advanced strategies employ cell-penetrating peptides such as the TAT peptide, which can directly facilitate translocation across the stratum corneum.⁸¹ In addition, multidecoration with both TAT and a permeation enhancer such as pyrene butyrate has been shown to further improve the anesthetic efficiency of NLC-loaded lidocaine.⁸² The synergy between material and physical enhancement was demonstrated by Hasanpour et al., who used 3D-printed microneedles to create microchannels in the skin, dramatically increasing the transdermal flux of NLC-encapsulated lidocaine applied subsequently.⁸³ The research focus has now decisively shifted from fixing inherent stability issues to leveraging NLCs as versatile scaffolds for advanced functionalization and application-specific design. A prime example is the engineering of multifunctional nanocarriers for enhanced topical anesthesia, as depicted in Fig. 3.⁸⁴ These state-of-the-art systems employ a multi-pronged strategy: co-loading a local anesthetic, levobupivacaine, with a synergistic adjuvant, dexmedetomidine, while functionalizing the surface with cell-penetrating peptides to overcome the stratum corneum barrier. This approach, which integrates structural optimization, synergistic co-delivery, and active surface targeting, exemplifies the current frontier in the development of high-performance topical nano-anesthetics.

Fig. 3 Schematic representation of advanced lipid-based nanocarriers for enhanced topical anesthesia. (A) TAT peptide-modified nanostructured lipid carriers (NLCs) and (B) lipid–polymer hybrid nanoparticles are engineered to co-load the local anesthetic levobupivacaine (LBVC) and the adjuvant dexmedetomidine (DMED). The surface functionalization with TAT is designed to facilitate penetration across the stratum corneum for transdermal delivery. Adapted with permission from ref. 84 (Li M, Feng S, Xing H, Sun Y. Dexmedetomidine and levobupivacaine co-loaded, transcriptional transactivator peptide modified nanostructured lipid carriers or lipid-polymer hybrid nanoparticles, which performed better for local anesthetic therapy? Drug Delivery. 2020 Dec;27(1):1452–1460. doi: 10.1080/10717544.2020.1831105. Copyright 2020, Taylor & Francis).

3.2. Polymer-based nanocarriers

Owing to their exceptional tunability, polymer-based nanocarriers, which are composed of natural, semisynthetic, or synthetic polymers, are arguably the most versatile class of delivery systems.⁸⁵ A wide array of polymeric nanomaterials (Table 2) have been developed, including micelles, nanogels, and solid nanoparticles, which are often constructed from benchmark biodegradable polyesters such as poly(lactic-co-glycolic acid), polylactic acid (PLA), and poly(ε-caprolactone) (PCL).^86–89 Drug loading into these carriers is achieved through two primary strategies. The most common method is physical encapsulation, where drug molecules are entrapped within the polymer matrix, and release is governed by mechanisms such as diffusion or polymer degradation.⁹⁰ A more advanced strategy is chemical conjugation, where the drug is covalently bonded to the polymer backbone via a cleavable linker. This approach was powerfully demonstrated by Zhao et al., who conjugated the potent neurotoxin tetrodotoxin (TTX) to a biodegradable polymer via hydrolyzable ester linkages. This polymer-TTX conjugate enabled sustained release over 3 days from a single injection, achieving a long-lasting nerve blockade that was previously unattainable without significant systemic toxicity.⁹¹ By leveraging these diverse design strategies, polymeric systems can be precisely engineered to meet a wide range of anesthetic delivery challenges.

Table 2 Representative polymer-based nanocarriers described in the literature for anesthetic delivery

Nanocarrier type/key innovation	Loaded drug(s)	Preclinical tests	Key quantitative outcome(s)	Clinical stage	Ref.
PLGA nanoparticles (benchmark system)	Bupivacaine	In vivo: Sheep intercostal nerve block (motor & sensory function tests)	Provided dose-dependent nerve blockade lasting up to 48–72 hours in a large animal (sheep) model via slow polymer degradation	Preclinical	26
Polymer–drug conjugate (chemical conjugation)	Tetrodotoxin (TTX)	In vitro: Hydrolysis/release kinetics	Covalent conjugation via a hydrolyzable linker enabled ultralong-acting nerve blockade for over 3 days	Preclinical	91
		In vivo: Rat sciatic nerve block (hot plate & motor function tests).
Mixed polymeric micelles (for propofol)	Propofol (GA)	In vitro: Drug loading & solubility studies.	Increased propofol's aqueous solubility by 72% (to 4.69 μg mL^−1) and achieved a high drug loading of 32.4% (compared with earlier report of 12%)	Preclinical	105
		In vivo: Rat Sleep/recovery studies
Injectable micellar hydrogel (“flower-type” micelles)	Lidocaine	In vitro: Gelation behavior	In situ crosslinking of micelles formed a gel depot that prolonged analgesia and mitigated local inflammation from acidic byproducts	Preclinical	107
		In vivo: Rat local infiltration (von Frey test); Histology (inflammation).
Thermosensitive nanogel (in situ depot)	Lidocaine	In vitro: Drug release kinetics; Rheology	PCL–PEG–PCL hydrogel formed a gel depot at body temperature, providing sustained release for over 350 hours in vitro and increasing in tail-flick response latency as compared to the normal saline/lidocaine group (1.67, 45.00 and 190.00 min)	Preclinical	110
		In vivo: Murine local infiltration (pin-prick test)
Hybrid nanogel (“nanocrystal-in-nanogel”)	Bupivacaine	In vitro: Drug loading; release kinetics; cytotoxicity	Achieved an ultra-high drug loading of 84.8 wt% by encapsulating drug nanocrystals, providing nerve block for >8 hours	Preclinical	112
		In vivo: Rat sciatic nerve block (hot plate test)
Multifunctional nanofibers (co-delivery)	Lidocaine + fusidic acid	In vitro: Drug release; antibacterial activity; cell viability	Electrospun PLGA fibers simultaneously prevented bacterial infection and provided prolonged pain relief in an infected wound model	Preclinical	98
		In vivo: Rat infected wound model (healing & pain scores)

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3.2.1. Polymeric nanoparticles. Among polymeric platforms, nanoparticles fabricated from the approved copolymer PLGA have emerged as the most extensively studied “workhorse” system, owing to their excellent biocompatibility, biodegradability, and highly tunable drug release properties.⁹² The potential of PLGA for sustained anesthetic release was recognized as early as 1981 by Wakiyama et al., who reported that PLGA-encapsulated tetracaine exhibited a prolonged release profile in vitro.⁹³

A landmark preclinical study by Dräger et al. established the clinical relevance of this platform. They applied a PLGA-based bupivacaine formulation to intercostal nerve blocks in sheep, a model chosen for its anatomical similarity to adult humans. The results were compelling: the formulation provided a dose-dependent nerve blockade lasting up to 48–72 hours, a dramatic increase from the 12-hour limit of a single bupivacaine injection, laying crucial groundwork for clinical translation.²⁶ Subsequent innovations have focused on refining the formulation and enhancing safety. For example, Moraes et al. demonstrated that PLGA nanospheres could effectively encapsulate a novel bupivacaine enantiomeric mixture (S75:R25), which not only extended the anesthetic duration but also reduced in vitro cytotoxicity by approximately 20% compared with the free drug.⁹⁴ The versatility of PLGA has been shown for various other LAs, including ropivacaine.⁹⁵ A particularly novel application was demonstrated by Horie et al., who applied lidocaine-loaded PLGA particles to the round window membrane of the cochlea in mice, achieving sustained perilymph drug concentrations for up to three days for the potential treatment of tinnitus.⁹⁶

However, a key material science challenge of PLGA is that its degradation produces acidic byproducts (lactic and glycolic acid), which can cause local inflammation and may degrade acid-labile drugs. Addressing this and other physical limitations (e.g., poor tissue retention of simple suspensions) has driven the development of advanced hybrid and multifunctional systems.

Campos et al. developed a composite system by incorporating chitosan-coated PLGA nanoparticles into a thermosensitive hydrogel. This hybrid formulation, loaded with bupivacaine and the natural compound limonene, demonstrated improved skin permeation and a sustained anesthetic effect lasting over 8 hours.⁹⁷ Pushing the boundaries of multifunctionality, Alsulami et al. recently developed an electrospun PLGA-based nanofiber system co-loaded with lidocaine and the antibiotic fusidic acid. This dual-action wound dressing was shown to prevent bacterial infection while simultaneously providing prolonged pain relief, thereby accelerating the healing process for infected wounds.⁹⁸ These examples highlight the ongoing evolution of PLGA from a simple nanosphere into a sophisticated and integral component of next-generation, multifunctional therapeutic systems.

Furthermore, researchers have engineered sophisticated composite platforms in which PLGA nanoparticles act as critical components for achieving multistage, sequential drug release. A state-of-the-art example is an injectable electrospun PLGA fiber-hydrogel composite designed for prolonged regional analgesia, as illustrated in Fig. 4.⁹⁹ This system achieves a precisely timed, synergistic effect: the hydrogel provides rapid release of clonidine for immediate pain control, while the embedded PLGA nanofibers subsequently ensure the long-term, sustained release of ropivacaine. This elegant design, which leverages multiple polymeric components to orchestrate a sequential release cascade, represents a significant step towards more advanced, patient-tailored pain management solutions.

Fig. 4 Schematic illustration of an injectable, electrospun fiber-hydrogel composite for sequential and prolonged regional analgesia. (A) Fabrication of Fiber–Rop/Gel–Clo by combining electrospun PCL fibers loaded with ropivacaine (Rop) and thermosensitive Pluronic F127 hydrogel encapsulating clonidine (Clo). (B) The rat sciatic nerve blockade effect of the Fiber–Rop/Gel–Clo composite. (C) The mechanism of long-acting and walking regional analgesia in vivo. Adapted with permission from ref. 99 (Chen S, Yao W, Wang H, Wang T, Xiao X, Sun G, Yang J, Guan Y, Zhang Z, Xia Z, Li M, Tao Y, Hei Z. Injectable electrospun fiber-hydrogel composite sequentially releasing clonidine and ropivacaine for prolonged and walking regional analgesia. Theranostics. 2022 Jun 21;12(11):4904–4921. doi: 10.7150/thno.74845. Copyright 2022, Ivyspring International Publisher).

3.2.2. Polymeric micelles. Polymeric micelles are nanoscale core–shell structures (typically <100 nm) formed by the self-assembly of amphiphilic block copolymers in aqueous solution. The hydrophobic core serves as an ideal reservoir for poorly water-soluble drugs, whereas the hydrophilic shell (often PEG) ensures colloidal stability and prolongs systemic circulation.^100,101 A primary application in anesthesia has been to improve the formulation of propofol, a vital but notoriously insoluble intravenous GA. Traditional lipid emulsions containing propofol carry risks of hyperlipidemia and microbial contamination. To address this, Boscan et al. developed a micellar formulation that not only eliminated the need for lipids but also demonstrated a superior anesthetic profile in horses, with faster, smoother recoveries and a 20% reduction in the required induction dose compared with the commercial emulsion.¹⁰² However, a key material science challenge for polymeric micelles is achieving high drug loading. Early propofol micelles based on PVP-PLA copolymers showed limited loading capacity.^103,104 A significant advance was made by Li et al., who engineered mixed micelles composed of mPEG-PLA and the surfactant Solutol HS 15. This strategy dramatically increased the aqueous solubility of propofol by 72% (to 4.69 μg mL⁻¹) and achieved a high drug loading of 32.4%, creating a stable, clinically viable formulation.¹⁰⁵ In the context of local anesthesia, Lalatsa et al. demonstrated that lidocaine-loaded micelles achieved a 2-fold greater transcorneal flux than EMLA cream did, highlighting their potential for ophthalmic applications.¹⁰⁶ More recently, Nagasaki et al. engineered “flower-type” redox-active micelles that, when injected, form an in situ crosslinked hydrogel. This innovative design, used to deliver lidocaine, not only prolonged analgesia but also mitigated the local inflammatory response often caused by the acidic degradation products of PLGA-based systems.¹⁰⁷

3.2.3. Nanogels (polymeric nanohydrogels). Nanogels are 3-dimensional, crosslinked polymer networks that are hydrophilic and swell in aqueous environments, combining the advantages of both hydrogels and nanoparticles.¹⁰⁸ Inherent smart behavior is a key reason why hydrogel platforms are being extensively explored for advanced applications requiring on-demand action, from regenerative medicine to controlled drug delivery.¹⁰⁹ Thermosensitive nanogels, for example, leverage polymers that undergo a phase transition at physiological temperatures. Gong et al. developed an injectable hydrogel based on thermosensitive PCL–PEG–PCL copolymers that is a liquid at room temperature but forms a gel depot in situ upon injection into the body.¹¹⁰ This system provided sustained release of lidocaine for more than 350 hours in vitro and significantly prolonged local anesthetic effects in mice.¹¹⁰

A primary challenge for nanogels, however, has been their limited capacity for loading hydrophobic drugs.¹¹¹ To overcome this, Alejo et al. pioneered a groundbreaking “nanocrystal-in-nanogel” hybrid strategy. By first formulating bupivacaine into high-payload nanocrystals (drug loading >99%) and then physically entrapping them within a thermosensitive nanogel, they achieved an ultrahigh overall drug loading of 84.8 wt%. This composite system provides an exceptionally long-lasting sciatic nerve block in rats (>8 hours at a low 2 mg dose) with minimal toxicity.¹¹² The versatility of nanogels is further enhanced by creating composites; a particularly compelling future direction lies in the synergistic integration of advanced hydrogel platforms, such as gelatin methacryloyl (GelMA), with bio-derived nanocarriers such as extracellular vesicles (EVs). This strategy leverages GelMA as an engineered, injectable depot that provides structural support and modulates the sustained release of EVs, which in turn function as highly biocompatible delivery shuttles. The profound potential of this synergy has already been demonstrated in regenerative medicine, where GelMA hydrogels functionalized with osteoblast-derived EVs have been shown to significantly promote stem cell osteogenic differentiation and accelerate bone regeneration.¹¹³

In addition to drug delivery, the porous and functional nature of nanogels makes them promising candidates for drug scavenging.¹¹⁴ Tan et al. demonstrated that poly(N-isopropylacrylamide)-based nanogels could effectively bind and “sponge up” local anesthetics such as bupivacaine, suggesting a novel therapeutic strategy to reverse local anesthetic systemic toxicity in cases of accidental overdose.¹¹⁵ More recently, the focus has shifted to complex hybrid systems for managing neuropathic pain, where esketamine-loaded nanoparticles have been incorporated into a hydrogel.¹¹⁶ This system provides sustained pain relief by modulating spinal astrocyte activation and reducing DRG neuron excitability, revealing a sophisticated, mechanism-driven approach to pain management.¹¹⁷

3.2.4. Dendrimers. Dendrimers are highly branched, monodisperse macromolecules with a well-defined, three-dimensional architecture.^118,119 A representative example is the poly(amidoamine) (PAMAM) dendrimer, which offers a high density of surface functional groups and internal cavities for drug encapsulation.¹²⁰ The unique structure of PAMAM dendrimers allows provides for exceptionally high drug payloads; for example, a fifth-generation (G5) PAMAM dendrimer can encapsulate up to ∼100 molecules of analgesics such as morphine or tramadol, releasing them in a sustained, pH-responsive manner.¹²¹ However, a major hurdle for the clinical translation of unmodified PAMAM dendrimers is their inherent cytotoxicity, which is primarily attributed to the high density of cationic primary amine groups on their surface. This challenge has driven extensive research into surface engineering as a strategy to both mitigate toxicity and add functionality. Chandrasekar et al. demonstrated a classic approach by modifying the surface of G4-PAMAM dendrimers with folic acid. This not only neutralizes the surface charge, significantly reducing cytotoxicity but also introduces the ability to actively target folate receptors that are often overexpressed on inflammatory cells, thereby guiding the drug-loaded dendrimer to sites of inflammation.¹²² These works exemplify the critical role of surface chemistry in transforming dendrimers from simple high-capacity carriers into safe and targeted delivery systems.¹²³

3.3. Inorganic nanomaterials

Inorganic nanomaterials (Table 3) have emerged as powerful platforms for anesthetic delivery, offering unique physicochemical properties not found in their organic counterparts.^124,125 These include exceptional drug loading capacities, tunable surface chemistry, and, most notably, the ability to respond to external physical stimuli such as light and magnetic fields.^126,127 However, their primary and most significant translational hurdle is their poor or nonexistent biodegradability, which raises critical concerns about long-term bioaccumulation and chronic toxicity.¹²⁸

Table 3 Representative inorganic nanomaterials described in the literature for anesthetic delivery

Nanocarrier type/key innovation	Loaded drug(s)	Preclinical tests	Key quantitative outcome(s)	Clinical stage	Ref.
Photothermal hybrid system (AuNRs-in-Hydrogel)	QX-314	In vitro: Photothermal conversion; NIR-triggered release	NIR-triggered system prolonged sensory blockade to 90.0 min, significantly longer than lidocaine (37.5 min) via thermal activation of TRPV1	Preclinical	47
		In vivo: Rat sciatic nerve block (foot-flick test)
Synergistic photothermal/photochemical system	Tetrodotoxin (TTX)	In vitro: Release kinetics; singlet oxygen quantum yield	Integrating AuNRs and a photosensitizer in a liposome enabled repeatable, on-demand nerve blockade at lower light irradiance	Preclinical	48
		In vivo: Rat subcutaneous infiltration (sensory test)
Magnetic hyperthermia system (FeAu@Gelatin)	Lidocaine	In vitro: Magnetic heating; cytotoxicity (MTT assay)	AMF-induced heating from the FeAu core melted the gelatin shell, triggering in vivo local anesthesia	Preclinical	135
		In vivo: Rat intramuscular injection (observation of anesthesia)
Magnetic guidance system (“nanoanesthesia”)	Ropivacaine	In vivo: Rat ankle block with external magnet (behavioral tests); pharmacokinetics	IV-administered magnetic NPs were concentrated at the target nerve using an external magnet, achieving a successful nerve block	Preclinical	133
Porous carrier with hybridization (MSN-in-Hydrogel)	Bupivacaine	In vitro: Drug loading; release kinetics	Composite system provided sustained release for over 60 hours and prolonged the in vivo anesthetic effect to 36 hours	Preclinical	141
		In vivo: Rat tail-flick test
Biodegradable hybrid framework (hollow organosilica)	Ropivacaine	In vitro: Drug loading; Ultrasound-triggered release	Ultrasound-triggered on-demand release provided analgesia for over 6 hours (three-fold longer as compared to single free ropivacaine injection) from a biodegradable organosilica framework	Preclinical	49
		In vivo: Murine incision pain model (mechanical & thermal tests)
Advanced theranostics (FUS-Triggered perfluorocarbon NPs)	propofol (GA)	In vivo: Non-human primate deep brain delivery (behavioral modulation & MRI)	Achieved on-demand drug release in deep brain regions of non-human primates, a landmark for CNS delivery	Preclinical	147

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3.3.1. Photothermally-responsive systems. A leading strategy for on-demand anesthesia involves the use of inorganic nanoparticles as potent photothermal agents. Gold nanoparticles (AuNPs), particularly nanorods, are exemplary in this regard. Their efficacy stems from surface plasmon resonance, a phenomenon where NIR light excites collective oscillations of electrons on the nanoparticle surface, leading to highly efficient and localized heat generation.^47,129 Zhou et al. harnessed this by embedding gold nanorods (AuNRs) within a thermosensitive hydrogel loaded with the lidocaine derivative QX-314. Upon 808 nm laser irradiation, the AuNRs generated sufficient local heat to trigger a phase transition in the hydrogel, releasing the drug and thermally activating TRPV1 channels on nociceptors. This dual-action system prolonged sensory blockade in rats to 90.0 ± 12.2 min, which was significantly longer than that of 1% lidocaine (37.5 ± 12.5 min).⁴⁷ To enhance photosensitivity further, Rwei et al. engineered a sophisticated hybrid system by integrating both AuNRs and a photosensitizer into a single liposome.⁴⁸ This created a synergistic effect, where both photothermal and photochemical mechanisms contributed to drug release, enabling repeatable, on-demand nerve blockade at a lower light irradiance.⁴⁸

While noble metals such as gold are highly effective, the field is actively exploring alternative photothermal transducers to enhance biocompatibility and diversify applications. Iron oxide nanoparticles (Fe₃O₄), for instance, have been integrated into implantable microneedle systems to achieve externally controlled transdermal release of lidocaine.¹³⁰ Further advancing this paradigm, a shift towards more biocompatible and potentially biodegradable inorganic materials represents a crucial and insightful strategy. Copper sulfide nanoparticles (CuS NPs) have emerged as a particularly promising alternative, offering both strong NIR absorbance and an improved safety profile. A state-of-the-art example is the development of a hybrid platform where NIR-absorbing CuS NPs are integrated into a thermosensitive nanogel containing high-payload bupivacaine nanocrystals, as illustrated in Fig. 5.¹³¹ The elegance of this design lies in its use of CuS NPs as highly efficient nano-transducers that convert light into localized heat, triggering a pulsatile, on-demand release of the anesthetic cargo. This approach, which retains the advantage of precise external control while mitigating long-term safety concerns, exemplifies a significant advancement towards clinically viable, on-demand anesthetic systems.^131,132

Fig. 5 Synthesis and NIR-light-triggered release mechanism of a hybrid thermoresponsive nanogel system. (A) Fabrication process of the hybrid nanocarrier. Bupivacaine nanocrystals (BNCs) are first encapsulated within a nanogel shell via precipitation polymerization. In a parallel process, copper sulfide nanoparticles (CuS NPs) are surface-modified using a layer-by-layer technique with polyelectrolytes to create a positive surface charge. The two components are then assembled through electrostatic coupling to form the final CuS@BNC-nanogel hybrid. (B) Mechanism of on-demand drug release. The hybrid nanogels are stable at physiological temperature (T < LCST), effectively retaining the drug. Upon irradiation with an 808 nm near-infrared (NIR) laser, the embedded CuS NPs act as potent photothermal agents, generating localized heat. This increases the temperature above the nanogel's lower critical solution temperature (T > LCST), triggering a rapid phase transition (deswelling) of the gel matrix that results in the controlled, on-demand release of the encapsulated bupivacaine.

3.3.2. Magnetically-responsive systems. Alternating magnetic fields offer another powerful external trigger with excellent tissue penetration. Two main strategies are employed: magnetic guidance and magnetic hyperthermia. Magnetic guidance utilizes a static magnetic field gradient to attract and concentrate drug-loaded magnetic nanoparticles physically at a target site. This “nanoanesthesia” concept was powerfully demonstrated by Mantha et al., who showed that intravenously administered ropivacaine-associated magnetic nanoparticles could be concentrated at the rat ankle via an external magnet, resulting in successful and targeted nerve block.¹³³ The second strategy, magnetic hyperthermia, uses a high-frequency AMF to induce heat in superparamagnetic nanoparticles through Néel and Brownian relaxation mechanisms.¹³⁴ This localized heating can then trigger drug release from a thermosensitive carrier. Ting et al. demonstrated this by developing FeAu@gelatin-lidocaine nanocomposites in which AMF-induced heating melted the gelatin shell, releasing lidocaine and inducing local anesthesia in vivo.¹³⁵

In addition to serving as standalone carriers, a highly promising future direction involves the integration of inorganic nanomaterials as functional cores within biocompatible organic matrices such as polymers. The magnetic polymeric micelles developed by Karami et al.¹³⁶ serve as a quintessential example of this strategy. In that system, the polymer forms the primary carrier structure. They constructed these micelles using mPEG-PCL and superparamagnetic iron oxide nanoparticles (SPIONs) to deliver the analgesic naproxen. It is the inorganic SPION that bestows the critical magnetic responsiveness, acting as the 'engine' for targeted delivery. This paradigm uses inorganic materials rather than a bulk chassis as the functional trigger. This is a key strategy for harnessing their unique physical properties and reducing the potential long-term toxicity of purely inorganic systems.

3.3.3. Porous inorganic carriers. Mesoporous silica nanoparticles (MSNs) have emerged as highly versatile carriers because of their large surface area, tunable pore sizes, and well-defined surface chemistry.^49,137–139 Drug loading is governed by specific interfacial interactions. As Sato et al. demonstrated, the loading of lidocaine onto MSNs is driven by electrostatic interactions between the cationic amine group of the drug (lidocaine-NH⁺) and the deprotonated surface silanol groups (Si–O⁻) of the silica.¹⁴⁰ This strong interaction allows for high drug loading and provides a robust mechanism for controlling release, which was shown to be accelerated by cation exchange in physiological buffers.¹⁴⁰ To create more practical formulations, bupivacaine-loaded MSNs have been embedded within an injectable alginate hydrogel, creating a composite system with dual-level release control and enhanced tissue retention.¹⁴¹

Given the critical challenge of the poor biodegradability of silica, recent innovations have focused on creating biodegradable hybrid frameworks. Gao et al. developed hollow mesoporous organosilica nanoparticles with a thioether-bridged framework. These nanoparticles not only offered high loading for ropivacaine but also demonstrated the ability for on-demand, repeated release triggered by ultrasound, providing analgesia for over 6 hours in a mouse model with excellent biocompatibility.⁴⁹ In another innovative approach, Yin et al. encapsulated a 2D silicene core within MSNs. This system leveraged the unique photothermal properties of the silicene core to achieve NIR-triggered release of ropivacaine, combining the high loading capacity of MSNs with the on-demand functionality of a photothermal agent.¹⁴²

These hybrid strategies represent the most promising path forward for inorganic carriers, aiming to retain their functional advantages while engineering solutions to their inherent safety concerns. To synthesize the extensive information presented on these diverse nanocarrier platforms, a strategic comparison of their core principles, advantages, and fundamental limitations is provided in Table 4.

Table 4 A strategic comparison of nanocarrier platforms for anesthetic delivery

Parameter	Lipid-based nanocarriers	Polymer-based nanocarriers	Inorganic nanomaterials
Core design principle	Biomimetic self-assembly of amphiphilic lipids into vesicular or matrix structures^15,51	Precisely engineered, tunable macromolecules via controlled polymerization^85,92	Harnessing unique, intrinsic physicochemical properties of inorganic elements^49,129,134
Primary advantages	High biocompatibility & clinical translation. Proven safety profile and FDA-approved precedents (e.g., Exparel®)^58,189	Exceptional tunability & versatility. unmatched control over degradation, release, and surface chemistry for diverse applications^85,92	External control & functionality. enables on-demand release (light, magnet, ultrasound) and theranostic capabilities^49,129,135,147
Fundamental limitations	Formulation instability. Prone to drug leakage (liposomes) or expulsion via recrystallization (SLNs)^55,56,71	Biological complexity. potential for local inflammation from acidic byproducts (e.g., PLGA) and immunogenicity^152,190	Non-biodegradability. Poses significant risks of long-term bioaccumulation and chronic toxicity, a major translational hurdle^128,148
Drug-loading efficiency	Moderate. Typically physical entrapment. Can be enhanced by active loading (e.g., IGLs with 64–82% trapping efficiency)^57	High to very high. Supports both physical encapsulation and covalent conjugation. Can reach >80 wt% with nanocrystal-in-nanogel strategies^112	High. Primarily physical adsorption onto high-surface-area structures like MSNs (e.g., 19.7% loading for bupivacaine)^141
Release kinetics	Primarily sustained/passive. Can be extended significantly (e.g., 72 h for MVLs;^189 >10 days for HIP formulations^60)	Highly tunable & sustained. Can be engineered for release over hours, days (e.g., PLGA for 48–72 h), or months^26	Primarily on-demand/triggered. Offers highest temporal control. Sustained release also possible (e.g., >60 h from MSN-hydrogels)^141
Toxicity profile	Generally high safety. Composed of biocompatible lipids, with low local tissue reaction reported for clinical formulations^149,191,192	Variable & material-dependent. Key concern is local inflammation from acidic degradation byproducts of polyesters^190	Major concern. non-biodegradability is the primary issue. Potential for long-term organ accumulation and cytotoxicity^128,148

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3.4. Advanced delivery systems for crossing the blood–brain barrier

A primary challenge in general anesthesia is achieving targeted drug delivery across the blood–brain barrier while minimizing systemic exposure. The BBB is not a simple physical wall but rather a highly complex, dynamic biological interface. Its formidable nature stems from the endothelial cells of brain capillaries, which are sealed by tight junctions that severely restrict paracellular transport and that also express many efflux pumps, such as P-glycoprotein, that actively expel many xenobiotics back into the bloodstream.¹⁴³ Overcoming this barrier requires sophisticated strategies that can transiently and safely modulate its integrity or hijack its natural transport mechanisms.

Broadly, nanocarrier strategies to cross the BBB can be categorized as biological or physical. Biological strategies typically involve functionalizing the nanoparticle surface with ligands such as antibodies against the transferrin receptor, and the antibodies bind to specific receptors on endothelial cells and trigger receptor-mediated transcytosis.¹⁴⁴ While elegant, this approach is often target-specific and can be limited by receptor saturation. In contrast, physical strategies employ external energy sources such as ultrasound or magnetic fields to noninvasively and locally increase the permeability of the barrier, offering a more universal and on-demand approach for drug delivery.¹⁴⁵

Focused ultrasound (FUS), often used in conjunction with intravenously administered microbubbles, has emerged as a revolutionary tool for this purpose. When the microbubbles reach the targeted brain vasculature, a low-intensity FUS beam causes them to oscillate and expand (a phenomenon known as acoustic cavitation). This mechanical stimulation transiently and reversibly opens the tight junctions between endothelial cells, creating a temporary window for co-administered nanocarriers to enter the brain parenchyma.^50,146,147 This technique was powerfully leveraged by Wang et al. to “uncage” propofol from perfluorocarbon nanocarriers with millimeter-level spatial and millisecond-level temporal precision in rodents, allowing them to map whole-brain functional networks.¹⁴⁸ In a landmark study pushing this technology closer to clinical reality, Wilson et al. extended this success to non-human primates. They demonstrated that perfluorocarbon nanoparticles loaded with propofol could be safely administered and that their payload could be released on-demand in deep brain regions, achieving a sufficient local drug concentration to reversibly modulate behavior while confirming an intact BBB post-procedure via MRI.¹⁴⁹ The validation of this technique in primates represents a critical step toward the clinical translation of on-demand, targeted general anesthesia (Fig. 6A and B).

Fig. 6 Advanced physical strategies for overcoming the blood–brain barrier (BBB). (A) Focused ultrasound (FUS) mediated delivery, where intravenously administered perfluorocarbon nanoparticles are activated by an external FUS transducer at a precise location in the brain, triggering localized drug release. (B) Overall therapeutic strategy shown in a rodent model, highlighting intravenous (IV) administration followed by targeted FUS application to achieve precise drug delivery across biological barriers like the blood–brain barrier. (C) Magnetically-guided delivery, illustrating the self-assembly of magnetic polymeric micelles (MPMs) co-encapsulating naproxen and superparamagnetic iron oxide nanoparticles (SPIONs), and their subsequent concentration at the BBB in a rodent model under the influence of an external magnetic field.

Magnetic guidance represents another powerful, noninvasive physical strategy. This approach was demonstrated in a key proof-of-concept by Karami et al., who developed hybrid magnetic polymeric micelles previously referenced for their material design.¹³⁶ By applying an external magnet (0.4 Tesla) to the skulls of anesthetized rats, they were able to actively guide the circulating micelles to the brain. The study critically revealed that smaller nanoparticles (137 nm) resulted in significantly greater brain accumulation after 8 hours than did larger particles (242 nm) or the free drug.¹³⁶ This work provides a compelling blueprint for how magnetic gradients can be used to actively steer nanocarriers across the BBB, suggesting significant potential for developing magnetically-guided carriers for general anesthetics (Fig. 6C).

4. Safety and toxicological hurdles: a materials science perspective

The clinical translation of nano-anesthetics, despite their therapeutic promise, is fundamentally constrained by a critical translational gap between preclinical performance and clinical outcomes.¹⁵⁰ This gap originates from safety and toxicological challenges that are inextricably linked to the core physicochemical properties of the nanomaterials themselves. The attributes that enable advanced functionality, such as high surface-to-volume ratios and tailored surface chemistry, also determine the complex interactions of materials with physiological systems.¹⁵¹

4.1. Immunogenicity and inflammatory risks

The nano-bio interface is the crucible where therapeutic efficacy and potential immunotoxicity are determined. An immediate concern for intravenously administered nanomedicines is complement activation-related pseudoallergy (CARPA), an acute, non-IgE-mediated hypersensitivity reaction triggered by nanoparticle opsonization and subsequent activation of the complement cascade.¹⁵² This risk is highly dependent on material properties, particularly for cationic nanoparticles which can provoke strong immune responses.¹⁵³ A more prominent and contemporary challenge arises from the use of PEG, a polymer once considered the gold-standard “stealth” coating. The widespread use of PEG has led to a high prevalence of pre-existing anti-PEG antibodies in the population, a factor now known to cause accelerated blood clearance and hypersensitivity reactions.^34,151 This ‘PEG dilemma’, highlighted by allergic responses linked to PEG–lipid nanoparticles used in certain mRNA vaccines, transforms a former solution into a new problem.³⁴ This reality mandates that materials chemists develop alternative stealth strategies, with promising options including zwitterionic polymers and biomimetic coatings, such as cell membrane cloaking, which can more effectively evade immune surveillance.³⁸

4.2. Biodegradation, bioaccumulation, and long-term toxicity

The long-term safety of a nanocarrier is almost entirely dictated by its in vivo fate, which presents a fundamental material science bottleneck. The challenge of assessing long-term safety is not unique to nano-anesthetics but is a central concern across the entire field of biomedical nanocomposites, where understanding the long-term fate and potential toxicity of these novel materials is a major barrier to widespread clinical translation.¹⁵⁴ While biodegradable carriers such as PLGA are generally preferred, their degradation is not always harmless. The acidic byproducts of PLGA hydrolysis can cause local inflammation. In rat model, PLGA microspheres caused muscle tissue damage near the sciatic nerve.¹⁵⁵

The long-term safety issue is most evident in non-biodegradable inorganic nanomaterials. They serve as a quintessential case study for translational failure. A fundamental conflict exists between their tunable physical properties, such as plasmon resonance and superparamagnetism, and their intractable biological fate.^150,151 The inability of these bacteria to degrade leads to long-term retention and bioaccumulation in the mononuclear phagocyte system. Crucially, studies have shown that silica and silver nanoparticles can persist in neural tissues, raising concerns of potential neurotoxicity.¹²⁸ This uncertainty surrounding the chronic effects of bioaccumulation, cited as a key translational challenge of “incomplete biodegradation and elimination”,¹⁵⁶ is a primary reason why no inorganic nano-anesthetics have advanced to the clinic for systemic applications.

4.3. The imperative for predictive and standardized toxicological frameworks

The high clinical attrition rate of nanomedicines, with success rates for cancer nanomedicines decreasing from 94% in Phase I to 14% in Phase III, clearly illustrates the predictive failure of current preclinical models.¹⁵⁰ This underscores a critical paucity of chronic toxicity data, as most animal studies are short-term and fail to capture delayed effects.¹⁵⁷

To bridge this translational gap, a paradigm shift towards “safe-by-design” is essential.¹⁵¹ This involves two key steps. First, the adoption of advanced in vitro models, such as patient-derived organoids and organ-on-a-chip systems, is necessary to generate more biologically relevant data early in development.¹⁵⁰ Second, the implementation of standardized testing protocols, such as those detailed in the OECD Guidelines for the Testing of Manufactured Nanomaterials, is crucial. A comprehensive toxicological evaluation under such frameworks would require sponsors to provide data on a battery of tests including genotoxicity, carcinogenicity, and developmental and reproductive toxicology (DART).¹⁵⁰ Adhering to such rigorous frameworks, alongside transparent reporting according to standards such as the Minimum Information Reporting in Bio-nano Experimental Literature (MIRIBEL),¹⁵¹ is vital for building the robust safety dossiers required by regulatory agencies.

5. Navigating the translational gauntlet: clinical, manufacturing, and regulatory hurdles

Beyond the intrinsic safety of the materials, the journey of a nano-anesthetic from a laboratory concept to a clinical product is governed by a formidable set of interconnected challenges at the interface of science, engineering, and policy. Successful navigation requires bridging the “traditional gap between scientists and clinicians”¹⁵⁸ and addressing the distinct hurdles of clinical trial design, industrial-scale manufacturing, and an evolving global regulatory landscape.

5.1. Clinical trial challenges and the nuances of success

The landmark approval of Exparel®, a multivesicular liposomal bupivacaine, was based on pivotal phase 4 trials demonstrating prolonged analgesia for up to 72 hours in models of soft tissue and orthopedic surgery (e.g., NCT04644796 for spine surgery).^159,160 However, the clinical superiority and cost-effectiveness of this formulation remain subjects of significant debate, highlighting the challenges of translating initial trial data into broad clinical practice. Post-marketing safety surveillance during its first two years on the market revealed a distinct adverse event profile compared with that of standard bupivacaine HCl, with hypotension and bradycardia being the most frequently reported events.¹⁶¹ More recently, some randomized clinical trials concluded that, for longer analgesia, liposomal bupivacaine showed no significant difference in efficacy or cost-effectiveness compared with standard bupivacaine hydrochloride.^162–164 These cases strongly illustrate that even after achieving regulatory approval, a nanomedicine's value proposition must be rigorously validated against the standard of care in real-world settings. This aligns with the broader challenge that preclinical promise often fails to translate, exemplified by cases such as CPC634 (nanoparticulate docetaxel), which failed in phase II despite excellent phase I data and high tumor accumulation¹⁶⁵ and Opaxio (paclitaxel/poly(L-glutamic acid) nanoconjugate) at the regulatory stage owing to trial design misalignment.¹⁶⁶

5.2. The manufacturing bottleneck: scalability and quality control

Underpinning many clinical challenges is the foundational manufacturing bottleneck: the transition from laboratory-scale synthesis to industrial-scale, good manufacturing practice (GMP)-compliant production.^150,158 Traditional batch methods often show poor reproducibility of critical quality attributes (CQAs), including particle size and drug loading, which are fundamental to product safety and efficacy.¹⁶⁷ To surmount these engineering hurdles, the field is decisively moving towards continuous manufacturing processes such as microfluidics, which offer superior control and consistency.¹⁶⁸ The successful industrial-scale production of COVID-19 mRNA vaccines serves as a landmark proof-of-concept for these advanced manufacturing platforms.¹⁶⁹ This technological shift must be guided by a quality-by-design (QbD) philosophy, a systematic approach endorsed by regulators that ensures that quality is engineered into the manufacturing process from the outset.¹⁷⁰ Ultimately, all the data must satisfy a complex and non-harmonized global regulatory landscape. Our view is that materials scientists can no longer consider regulating a downstream step.

Regulatory approval sets a high bar. The U.S. Food and Drug Administration (FDA) requires a direct clinical benefit beyond simple opioid reduction.¹⁷¹ Agency guidance outlines manufacturing and regulatory factors that material design must include manufacturing and regulatory factors at the earliest stage to create a viable translational pathway.¹⁵⁸

6. Future directions

6.1. Multifunctional and integrated nanoplatforms

The next wave of anesthetic nanomedicine is expected to build multifunctional carriers that deliver several therapeutic benefits simultaneously. Dual action formulations combine analgesia with anti-inflammatory activity. One example is HTX 011, which contains bupivacaine and low-dose meloxicam in a biodegradable matrix. This formulation results in prolonged pain control and reduces postoperative inflammation.^172,173 Comparable co-delivery concepts are being investigated for perioperative scenarios that require concurrent analgesia and anti-infection protection or neuroprotection of injured nerves. An equally promising development is the emergence of theranostic carriers that integrate imaging agents such as MRI probes or fluorescent probes into anesthetic nanoparticles. These carriers allow real-time visualization of drug distribution and enable image-guided nerve blocks for greater procedural precision. In addition, micro- and nanorobotic systems capable of autonomously navigating to peripheral nerves, sensing biochemical pain markers, and releasing LAs on demand have been investigated; magnetically or acoustically actuated prototypes have already achieved site-specific payload delivery.¹⁷⁴ Furthermore, the field is beginning to explore the use of engineered living systems as the ultimate biomimetic carriers. This futuristic approach involves harnessing bacteria, stem cells, or even immune cells as autonomous, “living” delivery vehicles that can navigate complex biological environments, sense pathological cues, and produce therapeutics in situ.²⁸ For anesthetic applications, one could envision engineering macrophages to home in on inflammatory tissues post-surgery and release analgesics, or programming bacteria to colonize specific nerve bundles and secrete local anesthetics in a controlled manner. This represents a new frontier in achieving truly personalized and self-regulated anesthesia.

6.2. Stimuli-responsive smart delivery systems

The design of “smart” nanocarriers that respond to endogenous or external stimuli remains a central goal in the field, promising for the release of anesthetics only when and where they are needed. Harnessing the mildly acidic microenvironment of inflamed tissue via pH-responsive formulations continues to be a key strategy.¹⁷⁵ Acid-labile linkers or pH-sensitive polymers remain stable at physiological pH 7.4 yet accelerate lidocaine release at approximately pH 6.5, providing superior analgesia at inflamed sites.¹⁷⁶ Similarly, the pursuit of thermal control drives innovation in systems employing polymers such as PNIPAM¹⁷⁷ or Poloxamer 407,¹⁷⁸ which convert from a liquid to a gel at body temperature, forming in situ depots that sustain drug release; designs tuned to the elevated temperatures of inflamed tissue can further intensify release during pain flare-ups. Enzyme-responsive matrices integrate protease-cleavable linkages that degrade in protease-rich wounds, offering extended lidocaine action in nerve-injury models.¹⁷⁹ Platforms triggered by external energy sources add the ultimate layer of clinical control; for example, ultrasound-activated nanodroplets, for instance, can be vaporized on demand, allowing clinicians to elicit pulsed lidocaine release with high spatiotemporal precision.¹⁸⁰ Collectively, such stimuli-responsive designs represent a decisive move toward self-regulated, patient-tailored anesthetic therapy.

6.3. AI-assisted nanocarrier design and personalized precision medicine

Artificial intelligence (AI) and machine learning are poised to accelerate every stage of anesthetic nanocarrier development.¹⁸¹ On the diagnostic front, for example, advanced algorithms are already being leveraged to classify anesthesia stages by analyzing real-time physiological signals, such as those from near-infrared spectroscopy, to create more reliable monitoring systems.¹⁸² Building on this wave of data-driven innovation, AI is also being applied to the rational design and development of anesthetic nanocarriers.¹⁸³ Predictive models already correlate particle size, drug loading efficiency, and release kinetics, enabling the in silico selection of formulations that meet predefined pharmacokinetic targets.¹⁸⁴ This shifts the paradigm from iterative experimentation to rational, predictive design, thereby addressing critical translational bottlenecks such as data scarcity and poor reproducibility.¹⁸⁵ Recently, AI has been leveraged to predict complex bio-nano interactions that dictate in vivo fate. A powerful case study is the prediction of the nanoparticle protein corona, where machine learning (ML) models, trained on extensive datasets, can now analyze a nanocarrier's physicochemical properties to accurately forecast the composition of the protein layer that forms in biological fluids.¹⁸⁶ This is a significant breakthrough, as it provides a crucial predictive link between initial material design and its ultimate clinical performance. Furthermore, the cutting-edge concept of self-driving laboratories involves automating the entire research and development workflow. These platforms integrate AI algorithms with robotic systems to create a closed-loop for autonomous discovery: the AI designs an experiment, a robot executes the synthesis and characterization, and the results are fed back to the AI to intelligently inform the next iteration with minimal human intervention.¹⁸⁷ This approach has been proposed as a nanomedicine materials acceleration platform (NanoMAP) to systematically generate high-quality, standardized data and rapidly optimize novel nanocarrier formulations, representing a paradigm shift in accelerating the bench-to-bedside translation of nanomedicines.¹⁸⁷

Early demonstrations of AI-modelled nanogel release under different physiological conditions underscore the potential of this technology to rapidly explore vast formulation spaces, reduce the experimental workload, and ultimately deliver precision-engineered nanomedicines for perioperative pain management.

7. Conclusions and perspectives

The development of nano-anesthetics stands at a pivotal juncture, offering a transformative paradigm to overcome the fundamental pharmacokinetic and safety limitations of conventional anesthetic agents. While the rational design of nanocarriers offers unprecedented control over anesthetic delivery, clinical translation is contingent upon navigating a formidable gauntlet of interconnected hurdles: mastering the complex nano-bio interface to ensure safety and long-term stability, bridging the preclinical-to-clinical gap with more predictive models for personalized adaptation, and overcoming the engineering imperative of scalable, reproducible manufacturing. The path forward, therefore, demands an integrated strategy that embeds a “safety-by-design” philosophy into material conception, leverages AI-driven platforms for predictive design and process automation, and treats early regulatory engagement as a prerequisite for the successful development of advanced smart materials.

List of abbreviations

AI	Artificial intelligence
AMFs	Alternating magnetic fields
BBB	Blood–brain barrier
CARPA	Complement activation-related pseudoallergy
CNS	Central nervous system
CQAs	Critical quality attributes
DART	Developmental and reproductive toxicology
DRG	Dorsal root ganglion
EMA	European Medicines Agency
EMLA	Eutectic mixture of local anesthetics
EPR	Enhanced permeability and retention
EVs	Extracellular vesicles
FDA	U.S. Food and Drug Administration
FUS	Focused ultrasound
GAs	General anesthetics
IGLs	Ionic-gradient liposomes
LAs	Local anesthetics
ML	Machine learning
MRI	Magnetic resonance imaging
MSNs	Mesoporous silica nanoparticles
MVLs	Multivesicular liposomes
NIR	Near-infrared
NLCs	Nanostructured lipid carriers
NPs	Nanoparticles
PAMAM	Poly(amidoamine)
PCL	Poly(ε-caprolactone)
PEG	Polyethylene glycol
PLA	Polylactic acid
PLGA	Poly(lactic-co-glycolic acid)
PNIPAM	Poly(N-isopropylacrylamide)
PVP	Poly(N-vinyl-2-pyrrolidone)
QbD	Quality-by-design
RGD	Arginine–glycine–aspartic acid
SLNs	Solid lipid nanoparticles
SPIONs	Superparamagnetic iron oxide nanoparticles
TAT	Transcriptional transactivator peptide
TTX	Tetrodotoxin
TRPV1	Transient receptor potential vanilloid 1

Author contributions

Xiangyu Hu: sorting out the references, writing the original draft and drawing the tables; Tao He: writing the original draft and drawing the figure; Donghang Zhang, Cheng Zhou and Peng Liang: conceiving and revising the manuscript. All authors have read and agreed to the published version of the manuscript, and that all authors agree to be accountable for all aspects of the work.

Conflicts of interest

The authors have no relevant financial or non-financial interests to disclose.

Data availability

No new data were generated or analyzed in support of this review. The work is based on previously published studies, and all data sources are cited accordingly in the text and listed in the bibliography.

Acknowledgements

This work was jointly supported by No. 82271290 (To C.Z.) from the National Natural Science Foundation of China; No. 2023NSFSC0676 (To P.L.) and 2023ZYD0168 (To C.Z.) from the Sichuan Provincial Department of Science and Technology; and No. 24QNMP086 (To T.H.) from the Medical Science and Technology Program of the Sichuan Provincial Health Commission.

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Footnote

† Xiangyu Hu and Tao He contributed equally to this work.

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