Polymeric nanocarriers for nose-to-brain drug delivery in neurodegenerative diseases and neurodevelopmental disorders

Abstract

Neurodegenerative diseases are progressive conditions that affect the neurons of the central nervous system (CNS) and result in their damage and death. Neurodevelopmental disorders include intellectual disability, autism spectrum disorder, and attention-deficit/hyperactivity disorder and stem from the disruption of essential neurodevelopmental processes. The treatment of neurodegenerative and neurodevelopmental conditions, together affecting ∼120 million people worldwide, is challenged by the blood–brain barrier (BBB) and the blood–cerebrospinal fluid barrier that prevent the crossing of drugs from the systemic circulation into the CNS. The nose-to-brain pathway that bypasses the BBB and increases the brain bioavailability of intranasally administered drugs is promising to improve the treatment of CNS conditions. This pathway is more efficient for nanoparticles than for solutions, hence, the research on intranasal nano-drug delivery systems has grown exponentially over the last decade. Polymeric nanoparticles have become key players in the field owing to the high design and synthetic flexibility. This review describes the challenges faced for the treatment of neurodegenerative and neurodevelopmental conditions, the molecular and cellular features of the nasal mucosa and the contribution of intranasal nano-drug delivery to overcome them. Then, a comprehensive overview of polymeric nanocarriers investigated to increase drug bioavailability in the brain is introduced.

Graphical abstract

This work reviews the use of polymeric nanocarriers for intranasal (nose-to-brain) drug delivery in neurodegenerative diseases and neurodevelopmental disorders.

1. Introduction

Neurodegenerative disease is an overarching term that describes a variety of progressive conditions that affect the neurons of the central nervous system (CNS), resulting in their damage and death. The complexity of neurodegenerative diseases is manifested not only in their diversity (e.g., Huntington’s disease, amyotrophic lateral sclerosis, Parkinson’s disease), displaying unique combinations of symptoms and progressing at varying rates, but the biochemical pathways involved are not fully understood and thus, their treatment is challenging. Over 46 million people currently live with dementia worldwide, and this number is expected to increase to 131 million cases by the year 2050, and it is urgent to better understand their origin, progression and new treatment strategies¹.

The leading neurodegenerative condition is Alzheimer’s disease (AD), which is considered as the main cause of global memory disorder. AD is characterized by accumulation of amyloid-beta peptide (Aβ; senile plaques) in the brain, memory loss, brain atrophy, presence of hyperphosphorylated tau filaments (neurofibrillary tangles), and cerebrovascular changes^2,3. Cholinesterase inhibitors and glutamate regulators are the main drug classes used in the therapy of AD. The former is considered as the first-line pharmacotherapy for mild to moderate AD, impeding the breakdown of acetylcholine, an important neurotransmitter associated with memory, by blocking the enzyme acetylcholinesterase. However, the side effects of per-os (oral) treatment include vomiting, loss of appetite, nausea, and increased frequency of bowel movements^2,3, which further deteriorate the patient’s lifestyle.

The brain is separated from its blood supply by two physiological barriers, the blood–brain barrier (BBB) and the blood–cerebrospinal fluid barrier (BCSFB)^4,5. The BBB is a semipermeable diffusion barrier formed by astrocytes, endothelial cells bound by tight junctions, neurons, pericytes and basal membrane and controls the influx and efflux of biological mediators involved in metabolic activity and neuronal function as well as drugs^4,5. The BCSFB is established by choroid plexus epithelial cells and controls the transfer of solutes, including drugs, from the bloodstream to the cerebrospinal fluid (CSF) that is secreted by this epithelium⁶.

Drugs for the treatment of neurodegenerative diseases cross the BBB by different pathways that depend on their physicochemical properties, though a limited number has reached the clinical practice and their bioavailability is often jeopardized by the presence of intact BBB and BCSFB^7,8. These crucial obstacles for drug delivery to the CNS also affect the treatment trajectory of many other psychiatric illnesses.

Neurodevelopmental disorders include intellectual disability, autism spectrum disorder (ASD), and attention-deficit/hyperactivity disorder (ADHD) and rely on the disruption of essential neurodevelopmental processes that perturb brain development^9,10. They share similar treatment limitations as those described above for neurodegenerative diseases. Despite recent advances, drug discovery research in neurodevelopmental disorders faces challenges associated with their heterogeneous condition and set major difficulties to bridge the gap for bench-to-bedside translation¹¹. The prevalence of neurodevelopmental disorders exceeds 15% of the population worldwide, with frequent comorbidity with other neurological disorders (e.g., anxiety). Neurodevelopmental disorders exhibit different patterns in the acquisition of motor, cognitive, linguistic and socio-emotional skills, which affect the functioning in different contexts relevant mainly to children¹². These clinical manifestations are varying in severity with differing extents of cognitive and adaptive functioning disability¹³.

Neurodegenerative diseases and neurodevelopmental disorders that together affect approximately 120 million people share a significant reduction in lifespan and high rate of disability and they are detrimental of the quality of life¹⁴. Hence, there is an overwhelming need to develop new treatment strategies that address unmet needs and reduce their associated disability. For example, there is lack of effective pharmacological treatments for ASD core symptoms¹⁵. Moreover, both neurodegenerative and neurodevelopmental conditions encountered a dry pipeline in the last two decades. The attrition rate in these fields is very high and strongly linked to many aspects of drug discovery, including drug efficacy, formulation and delivery, side effects, bioavailability, and the absence of so-called disease-modifying drugs that overall reduce the therapeutic index¹⁶.

2. The nose-to-brain (intranasal) pathway to target drugs to the central nervous system

The use of most therapeutic agents (e.g., small-molecule drugs, biologicals), for the treatment of CNS disorders in general and neurodegenerative diseases and neurodevelopmental disorders in particular, by oral and parenteral routes is precluded by their very low permeability across the BBB and bioavailability in the cerebral parenchyma^8,17. Transient increase of the BBB permeability with osmotic solutions (e.g., hyperosmolar mannitol), radiation therapy, focused ultrasound, and pharmacological manipulation have been explored though, they are non-specific and often encompass risks and side effects^18,19. Other delivery strategies include the synthesis of prodrugs by the conjugation of the active compound to shuttle peptides or other transport vectors (e.g., transferrin) that target pathways overexpressed at the apical side of the BBB endothelium (e.g., transferrin receptor, glucose transporters) through a cleavable linker, and enable transcytosis and delivery into the CNS upon cleavage of sensitive chemical bonds in the extracellular or the intracellular space^20,21. However, since prodrugs are often administered systemically, they could also undergo biodistribution and release in off-target (peripheral) body sites, leading to systemic adverse effects. In addition, they do not necessarily ensure high drug bioavailability in the CNS. The capitalization of anatomical and physiological pathways represents an appealing strategy to target drugs to organs such as the brain, while minimizing systemic exposure and reducing the effective dose and the side affects.

Nasal administration was originally used for local drug delivery in the treatment of local allergic rhinitis^22,23. Features such as minimal invasiveness, painlessness, self-administration and high patient compliance contributed to its success. The nasal route also enables drug delivery to the systemic blood circulation in cases where fast absorption is required or in unconscious patients owing to a large surface area, a very permeable endothelial network and high total blood flow^24,25. This route is commonly named intranasal (i.n.) or transnasal though to discern between delivery into the systemic circulation to treat peripheral conditions and directly from the nose to the brain, which is the main scope of the current review, it would be more convenient to call it transnasal in full analogy to the transpulmonary pathway that capitalizes on the airways to deliver drugs systemically^26,27. Transnasal delivery also bypasses hepatic first-pass metabolism and may result in systemic bioavailability comparable to the intravenous (i.v.) and the intramuscular (i.m.) routes, such as in the case of opioid drugs²⁸. This route has been also proposed to deliver proteins, hormones, anti-inflammatory and anti-migraine drugs, and pain killers^29,30,31. In addition, the presence of the nasal-associated lymphoid tissue (NALT) enables mucosal immunization by the nasal administration of vaccines^32,33.

In 1989, Prof. W.H. Frey II described for the first time and filed a patent on the i.n. delivery of peptides and other therapeutic molecules directly to the brain through an olfactory neural pathway between the nasal mucosa and the CNS that bypasses the BBB³⁴. His pioneering research and the understanding of the nose-to-brain anatomy and the key cellular mechanisms involved in this transport^35,36 led to the comprehensive investigation of this alternative administration route for the passive targeting of small-molecule and macromolecular (e.g., proteins, peptides, genes) drugs in the therapy of brain cancers and infections, stroke and a broad spectrum of CNS conditions including epilepsy, eating disorders, and drug addiction^37,38. Advantages of the i.n. route comprise quick onset of action, avoidance of hepatic first-pass metabolism, and good patient compliance. Since then, numerous clinical trials predominantly using drug solutions have been conducted especially in psychiatric and neurodegenerative diseases38, 39 though with relatively low preclinical success rates and limited bench-to-bedside translation³⁹. A main constraint of i.n. administration is the small volume that can be administered per nostril each time and that limits the maximum dose³⁸. At the same time, owing to self-administration, repeated dosing regimens are possible to reach the therapeutic drug concentration in the CNS.

Transnasal and i.n. drug delivery is challenged by a series of permeability barriers in the nasal mucosa that include the mucus layer and the underlying epithelium. Small-molecule lipophilic drugs are absorbed rapidly and efficiently from the nasal cavity, and in the case of transnasal administration result in plasma concentration profiles comparable to those of i.v. injection (∼100% bioavailability)⁴⁰. The same is valid for i.n. administration. Conversely, the systemic bioavailability of low molecular weight hydrophilic drugs is only 10% and of high molecular weight compounds (e.g., peptides and proteins) does not exceed 1% through the nasal cavity. The major obstacle for the absorption of hydrophilic molecules and macromolecules in the nasal mucosa is the low membrane permeability where the epithelial cells are closely connected by the tight junctions, which are the major controllers of the paracellular transport⁴¹. The diameter of a closed tight junction is 4–8 Å, restricting the permeability of small lipophilic molecules with narrow hydrodynamic diameter. However, the dynamic structure of these junctions enables them to open and close to a certain degree, which allows small hydrophilic molecules with molecular weights less than 1000 Da to cross the membrane by paracellular mechanisms. Another factor that limits nasal drug permeability is mucociliary clearance towards the esophagus which leads to undesired oral administration⁴². The half-life of drugs administered in the human nasal cavity is only 15–20 min. In the case of molecules that are not absorbed fast (including hydrophilic molecules and peptides), this short contact time with the mucosa can drastically decrease the drug bioavailability⁴¹. Regardless of the drug features, the difference in the pathway (transnasal or i.n.) upon nasal drug administration mainly relies on the region of the nostril where the formulation is administered. These barriers are less relevant in the case of local nasal treatments (e.g., allergic rhinitis) because in these interventions, drug absorption is not required.

Since i.n. administration localizes the delivery at the CNS, lower doses would be required to reach therapeutic efficacy when compared to systemic routes. At the same time, the complex structure of the nasal cavity and the different absorption rates and transport pathways that drugs, especially high molecular weight ones, could undergo (into the systemic circulation or the CNS) have contributed to low clinical success³⁹. To overcome the barriers in the nasal mucosa, various types of carriers, some of which contain absorption promoters that modulate the permeability of the nasal mucosa, were designed. Among these are surfactants, bile salt derivatives, fatty acids and their derivatives and polymers⁴¹. These molecules improve drug solubility and increase permeability across tight junctions in the nasal epithelium. Solubility enhancement can be achieved by encapsulating the drug within the carrier, such as a cyclodextrin, a microemulsion, or nanoparticles. These carriers have been also evaluated for i.n. drug delivery to the CNS⁴². Another strategy for enhancing drug absorption upon i.n. administration is modifying the permeability of the nasal epithelium using penetration. Slowing down the mucociliary clearance is also used to increase the bioavailability of nasally administrated drugs via prolonged contact time between the formulation and the nasal mucosa. Incorporating mucoadhesive agents such as polymers (e.g., chitosan) is one of the most useful methods to achieve this goal and extend the residence time in the nasal cavity⁴¹.

Regardless of the efforts to elucidate the molecular features that govern the absorption, transport and pharmacokinetics (PK) in the CNS for drugs in solution and in particulate form such as microparticles and nanoparticles (NPs), and the complex cellular pathways involved in it, they remain a matter of investigation and scientific debate³⁸. Most probably, the transport relies on the interplay of diverse intracellular and extracellular mechanisms that contribute differently according to the properties of the drug form (soluble or particulate) and formulation. There exists general consensus about the key role of the olfactory region though how olfactory neurons perform in this transport is not fully understood for active agents of different physicochemical properties (e.g., hydrophilic, hydrophobic) and forms (e.g., solution, suspension). The intracellular mechanism comprises drug internalization by an olfactory neuron, trafficking of the endocytic vesicle within the cell to the neuron's projection site, and the final release via exocytosis. The extracellular pathway starts when the drug crosses the nasal epithelium to the lamina propria, before it is transported externally along the length of the neuronal axon by bulk flow processes. The axon leads to the CNS, and the drug is distributed further via fluid movement⁴³. The extracellular transport is faster than the intracellular one and, therefore, it has been proposed as the predominant mechanism for fast drug transport to the brain⁴⁴. Drugs can also directly reach the brain via the trigeminal nerve pathway which connects to the tail part of the brain such as the spinal cord, the pons, and the medulla. The movement of the drug in the nasal cavity via the trigeminal nerve pathway is by intracellular (axonal) transport or endocytosis. The trigeminal nerve is composed of three branches, ophthalmic, maxillary and mandibular divisions. These branches play an important role in nasal administration to the brain because neurons from these branches pass directly through the nasal mucosa. Some of the segments of the trigeminal nerve end in the olfactory bulbs. Branches from the ophthalmic part of the trigeminal nerve innervate the dorsal part of the nasal mucosa and the anterior nose and the maxillary branch innervates the nasal turbinates that are covered by mucosa. Once a compound diffuses through the nasal mucosa, it reaches the branches of trigeminal nerves in olfactory and respiratory regions, and via the brain stem it is transported to the axonal route. Many mechanisms are involved in this pathway such as paracellular, transcellular, carrier-mediated transport, transcytosis, and receptor-mediated transport^37,45. In addition to the olfactory and trigeminal pathways, the compound can enter the brain indirectly, via the blood vasculature and the lymphatic system. Neuronal processes, which are projections from the neuron body, are protected by multiple glial cells such as oligodendrocytes, astrocytes and microglia⁴⁶. The participation of these supportive cells in the nose-to-brain transport is not clear yet. Recently, our group showed for the first time the possible role of microglia in the nose-to-brain transport of different types of NPs^47,48. Since this work is focused on nanoparticulate drug delivery systems for nose-to-brain administration, the possible transport mechanisms of NPs are discussed more extensively in the next section. The clinical potential of this administration route has also led to the design of administration-metered devices that maximize the drug dose deposited in the olfactory epithelium and improve patient comfortability⁴⁹. Fig. 1 schematically describes the nose-to-brain pathway and the main permeability barriers in the nasal mucosa to deliver drugs to the CNS by the nose-to-brain pathway.

Figure 1.

Schematic illustration of the intranasal (nose-to-brain) administration pathway and the permeability barriers in the nasal mucosa. Drawn with BioRender (https://biorender.com/, 2022). Retrieved from https://app.biorender.com/biorender-templates.

3. Molecular and cellular features of the nasal mucosa

The external nose is a pyramidal structure with a bony upper portion bound to the frontal bone of the skull and cartilage and soft tissues that form the nostrils that control the airflow into the airways^50,51. The internal nose is the functional part formed by two cavities that extend from the base of the skull to the roof of the mouth and that open up to the face through the nostrils⁵². The nostrils are highly vascularized and present an innervated epithelium with a total surface area of approximately 150–200 cm² and a volume of 15 mL^53,54. Despite the species, four cell types can be identified in the nasal epithelium: (i) an outer squamous epithelium layer that is restricted to the vestibules; (ii) ciliated pseudostratified cuboidal/columnar epithelium (respiratory epithelium) present in the main chamber and the nasopharynx and that extends from the base of the skull to the upper surface of the palate; (iii) non-ciliated/columnar epithelium or transitional epithelium that lies between the squamous and the respiratory epithelium in the anterior portion of the main chamber and (iv) olfactory epithelium located in the dorsal part of the cavity. The surface of the olfactory epithelium is greater in rodents and canines than in primates and humans which represents a challenge to design clinically relevant preclinical trials as mouse and rat are the primary animal models⁵⁵. In addition, the olfactory epithelium is comprised of the olfactory sensory neurons, and the supporting and the basal cells. The olfactory sensory neurons are bipolar cells emerging at the olfactory bulb, crossing the cribiform plate, and ending in the nasal submucosa with dendritic portions at bulbous olfactory knobs that are covered by immotile cilia⁵⁴. For decades, these neurons have been intimately associated with the nose-to-brain transport of soluble and particulate matter. The nasal cavity is also innervated by the trigeminal nerve and displays a rich vascular and lymphatic network involved in the fast and efficient absorption of low and high molecular weight drugs into the systemic circulation⁵⁴. The lymphatic system is formed by a lymphoepithelium that covers the NALT that is located in the underlying lamina propria. The nose also hosts serous glands distributed in the anterior part that secrete a viscous water-rich mixture of proteins and electrolytes and seromucous glands that produce a protein secretion that incorporates mucus. Two-thirds of the vestibules and the nostrils are covered by haired skin, and one-third by squamous epithelium followed by a transitional one⁵⁶. The rest of the cavity presents a typical respiratory epithelium that combines ciliated pseudostratified columnar epithelial cells with goblet cells that produce mucin and react to irritating stimuli⁵⁴.

4. Nanotechnology for nose-to-brain drug delivery

Nanotechnology has been comprehensively investigated to target drugs to the CNS, mainly by i.v. injection^57,58. For example, our group demonstrated the ability of different polymeric NPs surface-modified with shuttle peptides and folic acid receptor-alpha (FRα) to target pathways overexpressed in the BBB endothelium and the BCSFB chorid plexus, respectively, and increase the brain bioavailability^59,60. Regardless of the increase of the CNS bioavailability by severalfold with respect to untargeted counterparts59, 60, these NPs (as most of those described in the scientific literature) accumulate in major clearence organs such as the liver and the spleen according to their size, shape and chemical composition and surface properties^61,62. Conversely, very small NPs (5–10 nm) usually undergo renal filtration and are secreted in the urine. In this context, i.n. administration remains an appealing administration option to better constrain the delivery to the CNS and reduce systemic exposure, also owing to the lower doses required to achieve pharmacologically active drug concentrations.

Even though that there is still limited understanding of the critical formulation properties that control the nose-to-brain transport, several works showed that this pathway is more efficient for NPs than for solutions⁶³. For example, Kumar et al.^64,65 reported on the nanoencapsulation of the antipsychotic drug risperidone within a mucoadhesive nanoemulsion and its targeting to the brain of Swiss albino rats by the i.n. route. The brain-to-blood uptake ratio with the nanoemulsion was substantially increased with respect to a drug solution⁶⁵. Similar findings were reported for ziprasidone hydrochloride loaded into a buffered nanoemulsion⁶⁶. Following this trend, different pure drug nanoparticles⁶⁷, emulsions^64,65, and lipid and polymeric nanocarriers^68,69,^70,71 have been investigated for i.n. drug delivery^72,73. In addition, the use of mucoadhesive excipients^74,75 that prolong the residence time of the nanoformulation in the nasal mucosa have been shown to increase absorption, nose-to-brain transport and CNS bioavailability76, 77.

More than 98% of the small-molecule and macromolecular drugs and new neurotherapeutics under preclinical and clinical investigation fail in crossing the BBB from the systemic circulation and show poor biopharmaceutical and PK properties8, 78. The presence of efflux mechanisms such as those of adenosine triphosphate (ATP)-binding cassette (ABC) transporters family contribute to reduce the drug bioavailability at the CNS⁷⁹. This, together with peripheral accumulation and potential side effects following systemic administration, and the need to ensure high CNS bioavailability, call for novel therapeutic strategies. Nano-drug delivery systems administered by various routes have been proposed to bypass the activity of ABCs at the BBB level⁸⁰. In this conceptual framework, their i.n. administration for local drug delivery to the brain in neurodegenerative diseases and neurodevelopmental disorders, is very promising because small size and high surface area make nanocarriers radically more efficient than other classical technologies to surpasss this dynamic barrier and to release the cargo in the target organ, the brain.

As described above, the complex transport mechanisms from the nasal mucosa to the brain are not fully understood yet and most probably, they differ for drug solutions and particulate formulations of different particle size, shape, surface chemistry and charge, and flexibility. The maximum NP size to enable nose-to-brain transport was initially set at ∼100 nm, though particles as large as 300 nm were also succesfully delivered to the brain⁸¹. These divergent results highlight that a more comprehensive investigation of the molecular and cellular mechanisms and the NP features that govern the transport is critical to reach the clinical stage and increase translation into pharmaceutical products⁸². Moreover, the lessons learned from i.n. nanotherapeutics and the in vitro, ex vivo and in vivo models developed to assess their efficacy would have implications in other fields such as nanosafey and nanotoxicology. For instance, the unintended exposure to air nanoparticulate pollutants such as carbon NPs of different kinds (e.g., carbon dots) and other inorganic NPs with profuse industrial production and use (e.g., silica, zinc oxide) and their accumulation in the brain may result in CNS toxicity (e.g., neuroinflammation) and it has become a growing health concern^83,84. For example, ultrafine NPs (>100 nm) that contact the nasal mucosa, undergo translocation and reach the olfactory bulb⁸⁵. In 1970, De Lorenzo showed by transmission electron microscopy that gold NPs (diameter of 50 nm) instilled into the nose of squirrel monkeys translocated anterogradely in the axons of the olfactory nerves to the olfactory bulb⁸⁶. Moreover, an olfactory bulb–brain translocation pathway has been also shown for airborne zinc oxide NPs in rat⁸⁴. Others suggested that NPs are mainly internalized by neuronal terminals of the olfactory nerve system that emerge in the brain and end in the nasal olfactory region^87,88. Microglia are macrophage-like cells considered the brain’s garbage men⁸⁹. Quiescent microglia are extremely plastic cells and reside near neurogenic centers such as the olfactory bulb where they can change their phenotype upon stimulation and reach the nasal mucosa^90,91. Microglia cells undergo permanent maintenance and in situ proliferation and upon entry of nanoparticulate matter^92,93, they undergo activation and produce cytokines that recruit immune cells from the perivascular region^94,95. Recently, we investigated the interaction of different polymeric NPs (hydrodynamic diameter in the 17–483 nm range) with neutral, negatively- and positively-charged surfaces with different cells of the nose-to-brain axis, namely primary olfactory sensory neurons, cortical neurons, and microglia isolated from olfactory bulb, olfactory epithelium, and cortex of newborn rat and showed that olfactory and forebrain neurons do not internalize these NPs, while, as we hypothesized, they are endocytosed by olfactory and cortical microglia⁴⁷. These results represent the first clear-cut proof of the possible role of microglia in the nose-to-brain transport pathway.

Multicellular and heterocellular spheroids have become solid in vitro tools in drug screening, tissue engineering and biofabrication, and nanomedicine^96,97. However, they should recapitulate the complex cellular composition of the organs in vivo. In this framework, we developed a cellular construct that included, for the first time, primary microglia cells to assess their possible role in NP uptake in the presence of other cells of the BBB⁴⁸. A relevant feature of this model is that hCMEC/D3 endothelial cells together with pericytes form a monolayer on the spheroid surface tightly encasing the rest of the cellular construct and they express vascular endothelium (VE)–cadherin and claudin-5 (CLDN5), characteristic of adherens and tight junctions, respectively⁴⁸. Once the spheroids were characterized, they were exposed to different fluorescently labeled polymeric NPs and their uptake visualized by confocal laser scanning microscopy (CLSM) and ligh-sheet microscopy (LSFM). Results suggested that the enhanced fluorescence of NP-exposed spheroids stems from their internalization by primary microglia or astrocytes (Fig. 2) and shed more light into other possible cellular mechanisms involved in the nose-to-brain transport⁴⁸.

Figure 2.

Characterization of the interaction of different polymeric nanoparticles with biofabricated 5-cell spheroids. CLSM micrographs of 5-cell heterocellular spheroids exposed to (A and B) mixed nanoparticles of chitosan-g-poly(methyl methacrylate) containing 30% w/w poly(methyl methacrylate) and poly(vinyl alcohol)-g-poly(methyl methacrylate) containing 17% w/w poly(methyl methacrylate) (CS-PMMA30:PVA-PMMA17) ionotropically crosslinked with sodium tripolyphosphate, (E and F) nanoparticles of chitosan-g-poly(methyl methacrylate) containing 33% w/w poly(methyl methacrylate) (CS-PMMA33), (I and J) nanoparticles of poly(vinyl alcohol)-g-poly(methyl methacrylate) containing 17% w/w poly(methyl methacrylate) (PVA-PMMA17) crosslinked with boric acid, (M and N) nanoparticles of hydrolyzed galactomannan-g-poly(methyl methacrylate) containing 28% w/w poly(methyl methacrylate) (hGM-PMMA28). LSFM micrographs of spheroids exposed to (C and D) CS-PMMA30:PVA-PMMA17, (G and H) CS-PMMA33, (K and L) PVA-PMMA17, and (O and P) hGM-PMMA28 nanoparticles. Cell spots in D, H, and L, and P are included to ease the visualization of the cells and exactly to locate the fluorescent nanoparticles underneath the spheroid surface. Up-taken nanoparticles are highlighted by dotted yellow squares. Reproduced from Ref. 48 with permission of Cell Press.

5. Polymeric nanocarriers in the nose-to-brain drug delivery

As described above, different types of nanocarriers have been investigated for nose-to-brain drug delivery. Polymeric NPs emerge as one of the most versatile ones owing to their safety, good drug encapsulation capacity and the ability to engineer their compositional and structural features to confer mucoadhesiveness and reduce mucociliary movement and hence, prolong their residence time in the nasal mucosa⁹⁸ as well as display permeation-enhancing properties for increased paracellular permeability across the nasal epithelium by transiently opening tight junctions^99,100. Different types of polymeric NPs have been investigated for i.n. delivery with variable preclinical success. At the same time, the research at the interface of polymeric NPs and nose-to-brain delivery remains scarce. The most relevant works in this field are described below.

5.1. Polymeric micelles

Polymeric micelles are spherical nanostructures consisting of a hydrophobic core and a hydrophilic shell formed by the self-aggregation of polymeric amphiphiles in aqueous medium above the critical micelle concentration (CMC) and accordingly can encapsulate hydrophobic drugs within their micellar core¹⁰¹. Polymeric micelles have been investigated to increase the bioavaiability of hydrophobic drugs by different mucosal administration routes^102,103 including intranasally¹⁰⁴. For example, aiming to target the reservoir of the human immunodeficiency virus in the CNS, we administered the antiretroviral efavirenz encapsulated within polymeric micelles made of linear and branched poly(ethylene oxide)-b-poly(propylene oxide) (PEO-PPO) block copolymers containing 20% and 30% w/w cargo by the i.v. and i.n. routes and tracked the drug concentration in plasma and in the brain¹⁰⁵. Efavirenz is highly lipophilic and its aqueous solubility is extremely low (<10 μg/mL)¹⁰⁶. For example, the i.v. administration of polymeric micelles of Pluronic® F127 and mixed polymeric micelles of Pluronic® F127 and Tetronic® 904 with a 20% w/w drug loading resulted in maximum concentration (C_max) values in plasma of 1.792 and 1.632 μg/mL, regardless of the difference in micellar size and composition¹⁰⁵. The area-under-the-curve in plasma for the first 2 h (AUC_plasma0–2h) followed the same trend, values being 1.465 and 1.487 μg/mL/h. When polymeric micelles containing 30% w/w payload, the drug dose increased because the administration volume remained constant and thus, a sharp increase of C_max from 1.632 to 2.403 μg/mL was recorded. The biological half-life (t_1/2) in plasma did not change much. After i.v. injection, efavirenz was also detected in the brain though at very low concentrations when compared to plasma¹⁰⁵. The i.n. administration of the different efavirenz-loaded polymeric micelles resulted in a substantial 3- to 4-fold increase of the CNS bioavailability (as expressed by the increase of the AUC_{CNS0–2 h}) with respect to the i.v. route and the concomitant decrease of the systemic exposure which is an advantageous feature of local drug delivery. As discussed above, the NP size would play a key role in the nose-to-brain transport^36,72,81, though the optimal size range is still under debate and other properties such as surface chemistry, shape and flexibility should be also considered¹⁰⁷. Nanocarriers smaller than 100 nm often display facilitated mucosal and transcellular transport via endocytic pathways. In our work, regardless of the size gap between mixed Pluronic® F127:Tetronic® 904 polymeric micelles with 20% and 30% w/w efavirenz loading (hydrodynamic diameter of 14 and 247 nm, respectively), upon i.n. administration, a 50% increase of the drug dose increased both C_max and AUC_CNS0–2h to a similar extent¹⁰⁵. These findings indicated that changes in the micellar size do not necessarily affect the efavirenz transport and suggest the role of other mechanisms. The inhibition of the breast cancer resistance protein (BCRP) for which efavirenz is a substrate could also contribute to increase the CNS bioavailability after i.n. administration¹⁰⁸ because this efflux pump is also overexpressed in the olfactory ephithelium and some PEO-PPO block copolymers with relatively low hydrophilic-lipophilic balance such as Tetronic® 904 have demonstrated their ability to block them and reduce the efflux of different drugs^109,110. Results suggested that these polymeric micelles would fulfill a dual role of nanocarrier and BCRP inhibitor.

Polymeric micelles loaded with neurotherapeutics have been also attempted in the treatment of neurodegenerative and neurodevelopmental conditions. Nour et al.¹¹¹ encapsulated the benzodiazipine drug clonazepam within mixed polymeric micelles of Pluronic® P123:Pluronic® L121 for direct delivery in epilepsy. Biodistribution studies showed a significant increase of the drug concentration in the brain and a prolonged protection from seizures in comparison with the i.n. administered solution. Wang et al.¹¹² investigated polymeric micelles (88 nm) that undergo gelation upon heating to a T > 32 °C loaded with rotigotine, a non-ergoline agonist of dopamine receptor used in Parkison’s disease. Intranasal administration increased the CNS bioavailaibiltiy with respect to i.v. injection in rat (Fig. 3)¹¹².

Figure 3.

(A) Rotigotine (ROT) concentration in plasma over time upon administration of a solution (ROT solution), polymeric micelles (PM) and polymeric micelles thermosensitive gel (ROT-PM-TSG) to rats (n = 6, mean ± SD). (B) ROT biodistribution in rat brain tissues upon i.n. administration of ROT-PM (n = 4, mean ± SD). (C) ROT biodistribution in rat brain tissues upon i.v. administration of a solution (n = 4, mean ± SD). (D) ROT biodistribution in rat brain tissues upon i.n. administration of ROT-PM-TSG (n = 4, mean ± SD). Reproduced from Ref. 112 with permission of Elsevier.

5.2. Solid polymeric nanoparticles

Solid polymeric NPs represent one of the most popular nanotechnology platforms for drug mucosal delivery¹¹³. In a recent work, we investigated a novel mechanism of folate uptake mediated by FRα and receptors expressed at the choroid plexus epithelium (e.g., low density lipoprotein receptor-2) that bind the FRα-folic acid complex for the delivery of nanomedicines to the CNS⁶⁰. For this, we synthesized NPs (hydrodynamic diameter of 58–98 nm) of a highly hydrophobic poly(ethylene glycol)-b-poly (epsilon-caprolactone) (PEG-b-PCL) block copolymer (molecular weight of ∼20,000 g/mol; PCL content of ∼80% w/w) that were surface-functionalized with the FRα-folic acid complex. After demonstration that the modification of the NP surface significantly increases the uptake and permeability across standard and inverted primary human choroid plexus epithelial cell monolayers in vitro, the biodistribution of unmodified and modified NPs following i.v. and i.n. administration to ICR male mice was comparatively assessed. In these experiments, 200 μL of NP suspension (0.1% w/v) was injected i.v. through the tail vein or 20 μL of the same suspension was administered i.n. Then, at predetermined time points, live animal screening was performed using IVIS Spectrum In Vivo Imaging System and the fluorescence radiance efficiency of each organ was measured and expressed as average radiance with the basal signal of the respective organ in untreated mice (control) used as blank. The highest NP accumulation following i.v. injection was observed in the liver (Fig. 4A and B), while all the other organs showed relatively lower fluorescence signal⁶⁰.

Figure 4.

Biodistribution of PEG-b-PCL nanoparticles (A) without modification (i.v. injection), (B) with FRα-folic acid modification (i.v. injection), (C) without modification (i.n. administration) and (D) with FRα-folic acid modification (i.n. administration). Nanoparticles were fluorescently labeled with NIR-797 and administered to Hsd:ICR mice (n = 3). The average radiance was measured using Living Imaging analysis software. Bars represent the average of mice at each time point. The error bars are S.D. of the mean. Reproduced from Ref. 60 with permission of the American Chemical Society.

As for the i.n. administration, their biodistribution in the brain and peripheral organs was lower than for the i.v. injection (Fig. 4C and D)¹¹⁴, despite sizes <100 nm. To further analyze the differences in brain accumulation among the different NPs and administration routes, at different time points, animals were dissected and brains were imaged (Fig. 5A)^60,114. The highest brain accumulation was observed for modified NPs 1 h after i.v. injection with an AUC of (11 ± 1) × 10⁶ p × h/s/cm²/sr that was significantly higher than that of unmodified counterparts that showed an AUC of (4 ± 1) × 10⁶ p × h/s/cm²/sr (Fig. 5B and C)^60,114.

Figure 5.

Ex vivo analysis of the distribution of PEG-b-PCL nanoparticles without and wth FRα-folic acid modification in the brain following i.v. and i.n. administration to Hsd:ICR mice (n = 3). Nanoparticles were fluorescently labeled with NIR-797. (A) IVIS images, (B) average radiance obtained after the subtraction of the control (untreated mice brain) radiance at different time points, and (C) calculated AUC values. Reproduced from Ref. 60 with permission of the American Chemical Society.

In the case of i.n. adminstration, no significant difference in terms of NP accumulation in the brain was observed with AUC values of (0.5 ± 0.1) × 10⁶ and (0.5 ± 0.2) × 10⁶ p × h/s/cm²/sr for unmodified and FRα-folic acid-modified NPs, respectively¹¹⁴. On one hand, the AUC of both NPs after i.n. administration was approximately 10 times lower than the counterparts administered intravenously. This could be explained by a 10-times smaller NP dose. On the other, the accumulation in the liver decreased by approximately 40 times, confirming that the i.n. route reduces the relative systemic exposure to the NPs. Regardless of size <100 nm, their nose-to-brain transport is not very efficient which suggests that size is not the only key NP feature to ensure good i.n. delivery to the brain and other properties such as flexibility could be involved in it.

5.3. Cationic polymers

Solid polymeric NPs have excellent mucoadhesive properties owing to their ability to interact via electrostatic bonds with negatively-charged mucins in mucus. Chitosan is one of the most extensively used cationic polymers in drug delivery because it also transiently opens epithelial tight junctions and increases paracellular transport^74,75. Chitosan NPs produced by different methods such as ionotropic crosslinking have been also intended for i.n. delivery^104,115. For example, we demonstrated the ability of amphiphilic mixed nanoparticles of chitosan-g-poly(methyl methacrylate) containing 30% w/w poly(methyl methacrylate) and poly(vinyl alcohol)-g-poly(methyl methacrylate) containing 17% w/w poly(methyl methacrylate) (CS-PMMA30:PVA-PMMA17) ionotropically crosslinked with sodium tripolyphosphate to cross a model of the nasal epithelium in vitro¹¹⁶. Upon i.v. administration to Hsd:ICR mice, these NPs mainly accumulated in the liver, while i.n. administration of a 10-times lower dose resulted in a substantial decrease of the biodistribution into off-target organs (20-fold decrease in hepatic accumulation) with respect to i.v. injection¹⁰⁴. Intriguingly, i.v.-injected NPs reached mainly the “top” brain and the i.n. counterparts were detected in the “bottom” brain and the head because of their retention in the nasal mucosa. In addition, the brain bioavailability was similar for both administration routes but with a 10-fold smaller dose in the case of i. n. and a more limited accumulation in off-target organs¹⁰⁴. These NPs have shown excellent encapsulation capacity for cannabidiol¹¹⁷ and are currently under investigation for i.n. delivery of different active compounds to treat CNS conditions. Other mucoadhesive polymers such as acrylated Eudragit® E PO have been also proposed as a platform for this administration route¹¹⁸.

As described above, neurodegenerative diseases and neurodevelopmental conditions represent a major health burden and treatments are ineffective due to low BBB penetration and CNS bioavailability¹¹⁹. In advance, the use of different types of polymeric NPs for the i.n. administration to treat them will be overviewed.

Depression is a psychiatric condition associated with the abnormalities in neuronal transport in the brain, this disease is the second most common cause of disability worldwide and it affects approximately 4% of the world’s population¹²⁰. In recent years, some works investigated the encapsulation of antidepressants within polymeric NPs to improve the treatment efficacy. Tong and coworkers encapsulated desvenlafaxine within poly(d,l-lactic-c-glycolic acid) (PLGA) NPs (size of ∼173 nm) surface-modified with chitosan for i.n. delivery¹²¹. Behavioral tests performed in rodent depression models to determine efficacy showed the reduction of depression symptoms and the increase of the monoamine (e.g., serotonine, noradrenaline and dopamine) levels in the brain of the group treated with i.n. desvenlafaxine-loaded PLGA-chitosan NPs when compared to the orally administered counterpart. In another work, the ability of thiolated chitosan nanoparticles to improve the nasal delivery of selegiline hydrochloride was investigated¹²². The antidepressant activity was evaluated by the forced swim and the tail suspension tests and results confirmed the promise of this polymeric nanocarrier for i.n. delivery.

PD is a progressive neurodegenerative disorder characterized by the presence of α-synuclein-containing Lewy bodies in the substantia nigra of the brain and widespread accumulation of α-synuclein which leads to decreased levels of dopamine in the striatum and disrupted motor control^123,124. Loss of dopaminergic neurons leads to reduced facilitation of voluntary movements. The most important characteristic of PD is motor impairment¹²⁵. In the last decade, several works investigated the encapsulation of dopamine, levodopa, and monoamine oxidase Type B (MAO-B) inhibitors within gels and NPs for nasal administration to reduce the motor impairment in PD patients¹²⁶. Rao et al.¹²⁶ formulated ropinirole in a thermo-reversible in situ nasal gel containing Pluronic® F127 and hydroxy methyl propyl cellulose as gelling agents. The optimized formulation was administered to mice by the i.v. and i.n. routes. Results indicated an AUC_brain/AUC_plasma ratio of 4.29 for the nasal gel compared to 0.16 for the i.v. counterpart¹²⁶ though the efficacy was not assessed. The neurochemical pathways of several CNS conditions are not fully understood and thus, the choice of the proper drug to treat them is difficult. The lack of clinically relevant animal models is another hurdle of the research in this field and, in most works, the correlation between increased drug bioavailability in the brain and improvement of the symptoms is not demonstrated. This is a serious limitation, since the drug could reach regions of the CNS that are not relevant to the treatment of the disease or the disorder. In another work, the MAO-B inhibitor rasagiline was encapsulated within chitosan-coated PLGA NPs (size of ∼122 nm) and the PK compared to a drug solution administered intravenously¹²⁷. The brain concentrations were higher upon i.n. nasal administration, the AUC in the brain being 430.32 ± 16.33% in the nasal administration group compared to only 27.90 ± 2.06% for the i.v. one. Also here, no efficacy study was carried out. In another work, the dopamine agonist pramipexole dihydrochloride was loaded into chitosan NPs and administered to rats with induced motor impairment¹²⁸. The group treated with drug-loaded NPs intranasally showed improved score of locomotor activity and reduced motor deficit in the form of catalepsy than other groups.

Psychotic disorders are severe mental disorders characterized by abnormal thinking and perceptions. Polymeric NPs were also loaded with antipsychotic drugs and tested in animal models for the treatment of schizophrenia. Piazza et al.¹²⁹ developed poly(ethylene glycol)–b-poly(d,l)-lactic-co-glycolic acid (PEG–PLGA) NPs surface-modified with Solanum tuberosum lectin and loaded with the antipsychotic drug haloperidol. In vivo experiments assessed the efficacy of the NPs to deliver haloperidol to the brain for which rats received drug-loaded NPs injected intraperitoneally (i.p.) and i.n., lectin-modified NPs i.n. (in all cases a single dose of 2 mg/kg) and drug-free NPs i.n. and catalepsy tests were performed. Results showed that the loaded NPs significantly reduce the cataleptic response in comparison to the empty counterparts. To estimate the concentration of the drug in the brain, tissue samples from the striatal region of the basal ganglia, and the olfactory bulb were removed after 1 h of the catalepsy test. A statistically significant difference of haloperidol concentration in the striatum between the empty and the loaded NPs, regardless of administration route, were observed (Fig. 6)¹²⁹. In addition, after i.p. injection, there was no difference between the solution and the nanoparticles. Conversely, haloperidol concentration in the striatum after i.n. administration of the NPs was significantly higher than via i.p. These findings confirmed the benefit of the i.n. route.

Figure 6.

Haloperidol concentrations in (A) striatum tissue and (B) olfactory bulb tissue in the different treatment groups: empty nanoparticles administered intranasally [Empty NPs (IN)], unencapsulated haloperidol-only intraperitoneal injection (HP only), haloperidol nanoparticles injected intraperitoneally [HP-NPs (IP)], haloperidol nanoparticles administered intranasally [HP-NPs (IN)], and haloperidol STL-nanoparticles administered intranasally [HP-NPs (IN)]. Reproduced from Ref. 129 with permission of Elsevier.

The in vitro experiments and the behavioral test (catalepsy) showed positive results and suggest the promise of this nanoformulation in the therapy of schizophrenia. However, a control group (untreated rats) should have been included in the catalepsy test to unequivocally demonstrate the beneficial effects of the treatment¹²⁹. Moreover, no PK studies were conducted and the improvement in the tests upon administration of the formulations cannot be correlated with an increase of the drug bioavailability in the brain. These are critical aspects that demand careful investigation to conclude about the therapeutic potential a nano-drug delivery platform.

Anxiety is a type of CNS disorder and its frequency and intensity is overwhelming and interfere with daily functioning. Anxiety disorders are associated with impaired workplace performance and heavy economic costs, as well as an increased risk of cardiovascular morbidity and mortality. Buspirone hydrochloride was first of the azapirone chemical class and a centrally acting anxiolytic agent primarily used to treat general anxiety disorder. Bari et al.¹³⁰ encapsulated this drug within thiolated chitosan NPs with a size of ∼227 nm produced by ionotropic gelation for i.n. delivery to the brain to treat anxiety disorders. The maximum brain concentration after the i.n. of a solution and the NPs to rats was 417.77 ± 19.24 and 797.46 ± 35.76 ng/mL, respectively, and significantly higher than after i.v. administration of a solution that resulted in 384.15 ± 13.42 ng/mL (Fig. 7)¹³⁰. In addition, concentrations for the i.n. formulations were always higher than the i.v. one at all time points. Again, no efficacy studies were conducted which limits the clinical relevance of the study.

Figure 7.

Buspirone hydrochloride (BUH) concentration after intranasal adminstration of loaded nanoparticles in brain and blood at different time points. Reproduced from Ref. 130 with permission of Elsevier.

Polymeric nanocarriers were also investigated for the treatment of post-traumatic stress disorder (PTSD). For example, CBD was loaded in temperature-sensitive hydrogels made of poloxamers 188 and 407 and administered intranasally¹³¹. Even though in this work the carrier was not defined as a nanoparticulate one, these gels are formed by the interaction of highly concentrated PEO-PPO polymeric micelles. These polymeric micelles are considered drug nanocarriers. PK and pharmacodynamics (PD) studies showed a significant improvement in the delivery of CBD via the nasal route and a potentially more efficacious treatment of PTSD than the oral route though efficacy studies are missing.

Other few research work investigated polymeric NPs to improve the treatment of other CNS conditions such as epilepsy. In this context, antiepileptic thyrotropin-releasing hormone (TRH) analogues were encapsulated within PLGA NPs coated with chitosan (size between 85 and 92 nm) for i.n. administration¹³². The chitosan coating prolonged the release of the cargo in vitro, as quanitified by high performance liquid chromatography (HPLC)¹³². To visualize the uptake of the NPs from the nose to the brain, they were loaded with fluorescent quantum dots (QDs). Results of fluorescence microscopy showed that the NPs reached the brain after 30 min as opposed to the free QDs that were visualized only after 4 h (Fig. 8)¹³².

Figure 8.

Fluorescence microscopic images of different parts of rat's brain after nasal administered for two different treatment groups (1) unencapsulated quantum dots (QDs) and (2) QD-loaded PLGA-chitosan nanoparticles. Reproduced from Ref. 132 with permission of Elsevier.

Behavioral tests were performed in a pentylenetetrazole model of epilepsy in mice and the NPs found to significantly inhibit the seizures and protect from death.

5.4. Dendrimers

Dendrimers are regularly hyperbranched polymers with a three-dimensional architecture, a low dispersity, a high and tunable functionality and sizes at the nanometer scale¹³³. Polyamidoamines (PAMAMs) were the first class of dendrimers introduced by D. Tomalia and coworkers. Owing to the high functionalization and depending on the branching level and the steric hindrance, these NPs enable the encapsulation of drugs within the dendrimeric core or their complexation and conjugation on the surface^134,135. Cell compatibility studies suggested the concentration-dependent toxicity of amine-terminated derivatives due to the cationization of the molecule and the damage of the cell membrane¹³⁶. A similar effect was observed in contact with red blood cells. Conversely, carboxylic acid- or hydroxyl-terminated dendrimers are often more biocompatible. This rationale has been capitalized on for the design of a broad variety of hyperbranched molecules with improved biocompatibility though their use in i.n. delivery has not been extensive¹³⁷.

The use of different types of polymeric nanocarriers in the i.n. delivery of drugs aimed at treating CNS disorders is summarized in Table 1.

Table 1.

Polymeric nanoparticles used in the i.n. delivery of drugs for the therapy of CNS disorders.

Disease	Drug	Type of nanoparticle	Main result	Ref.
Alzheimer’s disease	Donepezil hydrochloride	Thiolated chitosan hydrogel. The size of the particles was 438.7 ± 28.3 nm.	The pharmacokinetics study in rabbit indicated that the hydrogel increased the mean peak drug concentration and area-under-the-curve (AUC) by 46% and 39%, respectively, through nasal route compared to the oral tablets.	138
Alzheimer’s disease	Tacrine	PAMAM dendrimers (generations 4.0 and 4.5)	In vivo experiments in zebrafish larvae that administration of both DG4.0-TAC and DG4.5-TAC reduced the morphological and hepatotoxic effects of tacrine.	139
Anxiety	Buspirone hydrochloride	Thiolated chitosan nanoparticles. The particle size was 226.7 ± 2.52 nm.	The brain concentration achieved after intranasal administration of the nanoparticles was 797.46 ± 35.76 ng/mL and significantly higher than after intravenous administration of a drug solution (384.15 ± 13.42 ng/mL) and intranasal administration of a drug solution (417.77 ± 19.24 ng/mL).	130
Anxiety/depression	Venlafaxine hydrochloride	In situ-forming mucoadhesive thermoreversible gel formulated using Lutrol® F127 (18%) as a thermo-gelling copolymer. The gel is formed by interaction of concentrated polymeric micelles.	Pharmacodynamic experiments, showed that venlafaxine hydrochloride was more effective as an antidepressant by nasal administration as in situ gel nasal drops in comparison to oral administration of equivalent dose.	140
Depression	Desvenlafaxine	PLGA-chitosan nanoparticles. The mean size was 172 nm.	Nasal administration of desvenlafaxine-loaded nanoparticles improved the pharmacokinetic profile of drug in brain together with their brain/blood ratio at different time point.	121
Depression	Selegiline hydrochloride	Thiolated chitosan nanoparticles. The size of the nanoparticles was 215 ± 34.71 nm.	The highest concentration in the brain of rats was achieved with intranasal nanoparticles.	122
Post-traumatic stress disorder	Cannabidiol	Thermo-sensitive gels made of poloxamer 188 and poloxamer 407. The gel is formed by interaction of concentrated polymeric micelles.	Pharmacokinetics studies and brain and liver tissue distributions showed that nasal administration had more advantages in comparison with the oral administration. The time to the maximum concentration (Cmax), namely the Tmax, of nasal administration was shorter, and the drug could accumulate in the brain to achieve a brain targeting effect.	131
Parkinson’s disease	Geraniol-ursodeoxycholic acid complex (prodrug)	PLGA nanoparticles and solid lipid nanoparticles. The particle size was 200 nm.	The nasal treatment was only conducted with solid lipid nanoparticles due to better encapsulation capacity. The prodrug was detected in the cerebrospinal fluid of rats 30–150 min after the administration but not in the bloodstream.	141
Parkinson’s disease	Pramipexole dihydrochloride	Chitosan nanoparticles. The size was 292.5 ± 8.80 nm.	The pharmacodynamic studies of behavioral testing revealed improved score of photoactometer and reduced motor deficit in the form of catalepsy in pramipexole-loaded chitosan nanoparticles treatment when compared to the nasal solution or oral marketed tablets treatments.	128
Parkinson’s disease	Ropinirole	Thermoreversible nasal gels prepared using Pluronic® F127 and hydroxy methyl propyl cellulose as gelling agents.	The nasal administration achieved five-fold higher bioavailability in brain than i.v. administration.	126
Parkinson’s disease	Rasagiline	Chitosan-coated PLGA nanoparticles. The size was 122.38 ± 3.64 nm.	The nasal administration of the nanoparticles showed significant enhancement of bioavailability in the brain of rat.	127
Parkinson’s disease	Rasagiline mesylate	Thermosensitive gel prepared by combination of poloxamer 407 and poloxamer 188 with mucoadhesive polymers (carbopol 934 P and chitosan). Poloxamers form polymeric micelles.	Pharmacokinetic experiments in rabbits showed significant improvement in bioavailability (four- to six-fold) of the drug with intranasal gels whe compared to an oral and a nasal solution.	142
Parkinson’s disease	Selegiline hydrochloride	A gel of poloxamer 407 and chitosan. Poloxamers form polymeric micelles.	Behavioral studies showed an improved score of photoactometer and reduced motor impairment (the catalepsy score) in the animals treated with intranasal formulation. Significant increase in brain dopamine, and reduction in monoamine oxidase B levels.	143
Parkinson's disease	Rotigotine	Polymer micelles in a thermosensitive gel. The size was 88.62 ± 1.47 nm.	In vivo studies show that the mean residence time of the polymeric micelles and the gel after nasal administration was extended 1.43 and 1.79 times with respect to the intravenous group.	144
Epilepsy	Antiepileptic TRH analogues (NP335 and NP647)	PLGA-chitosan nanoparticles. The nanoparticles sizes were between 110.7 ± 9.7 and 163.6 ± 8.0 nm.	Behavioral tests were performed in epilepsy animal model to estimate the efficacy of the treatment. Results indicated that intranasally administered loaded nanoparticles significantly inhibit the seizures and protect them from death.	132
Epilepsy	Catechin hydrate	Chitosan-coated–PLGA nanoparticles. The particle size was 93.46 ± 3.94 nm.	The results of pharmacokinetics study indicated that the intranasal administration of the nanoparticles significantly enhance the brain bioavailability of the drug.	145
Epilepsy	Carbamazepine	Carboxymethyl chitosan nanoparticles. The size was 218.76 ± 2.41 nm.	Pharmacokinetics results show that the brain AUC0–∞ after intranasal administration of loaded nanoparticles was approximately 8.1 times higher than that using a solution.	146
Epilepsy	Carbamazepine	PAMAM dendrimers (generations 4.0 and 4.5)	Drug solubility was increased by using the dendrimers DG4.5 and controlled release profile was achieved (40% of the carbamazepine was retained after 28 h dialysis). High in vivo biocompatibility as no neurotoxicity, cardiotoxicity or malformations in a zebrafish model.	147
Status epilepticus	Clonazepam	Polymeric micelles, mixture of Pluronic® copolymers (L121 and P123). The size was between 83.77 nm and 132.7 nm.	Results indicated that the optimized formula had significantly higher drug-targeting efficiency (DTE = 242.3%), drug-targeting index (DTI = 144.25), and nose-to-brain direct transport percentage (DTP = 99.30%)	148
Schizophrenia	Haloperidol	Poly(ethylene glycol)–block-poly (d,l)-lactic-co-glycolic acid nanoparticles	Brain concentrations of the drug were higher by 1.5–3-fold via the intranasal administration of lectin-functionalized nanoparticles compared to unmodified ones and other administration routes.	129
Schizophrenia	Quetiapine fumarate	Chitosan nanoparticles. The size was 131.08 ± 7.45 nm.	Two-fold higher nasal bioavailability in brain with the nanoparticles in comparison to a drug solution following i.n. administration.	149
Schizophrenia	Haloperidol	PAMAM dendrimer. The size was 15.1 ± 5.4 nm.	Dendrimer-based formulation showed a significantly higher distribution of haloperidol in the brain and plasma compared to the intraperitoneal injection of haloperidol. Additionally, 6.7 times lower doses of nasal administration of dendrimer−haloperidolproduced behavioral responses that were comparable to those induced by intraperitoneal injection.	150

The types of drugs attempted for nanoencapsulation within polymeric NPs and nose-to-brain delivery is relatively small with respect to the large number of small-molecule drugs under clinical use to treat neudegenerative diseases and neurodevelopmental disorders^151,152.

6. The clinical impact of the intranasal route

Due to numerous challenges and obstacles at different stages of development, relatively few nanoparticle-based medicines have been approved for clinical use¹⁵³. Doxil® (liposomal doxorubicin) was the first nanomedicine approved by the US Food and Drug Administration (FDA) in 1995¹⁵⁴. Since then, the FDA approved more than 50 clinical trials of nanopharmaceuticals. As of May 2022, 112 clinical trials including the term “nano” were listed as “recruiting or “active” on ClinicalTrials.gov¹⁵⁵. Among them, only eight nanoformulations for nasal delivery are currently in clinical trial and just two of them for nose-to-brain delivery with no clinical trials for i.n. polymeric nanoparticles¹⁵⁵, indicating that regardless of their great therapeutic potential, the field is in its infancy. The nanoformulations for nasal delivery under clinical trial are summarized in Table 2¹⁵⁵. It is worth stressing that regardless of the fact that some clinical trials include the word “nano” in their name or abstract, this does not necessarily mean that the trialed formulation is a nanoformulation based on the nanonization of a pure drug for the production of pure drug nanoparticles or its encapsulation within a nanocarrier. Since in these cases the use of the word “nano” is misleading, these clinical trials were not included.

Table 2.

Nanoformulations for nasal administration under clinical trials as listed on ClinicalTrials.gov¹⁵⁵.

Name	Drug	Disease or medical condition	Country	Phase	Collaborator	Sponsor
Silver Nanoparticle Investigation for Treating Chronic Sinusitis (SNITCH)	Colloidal silver nanoparticles	Chronic rhinosinusitis (diagnosis)	USA	1	Washington University School of Medicine	Washington University School of Medicine
Impact of Chronic Rhinosinusitis on the Index of Ciliary Beat Efficiency Using Fluorescent Nanosticks: (R-IMPAC) (R-IMPAC)	Nasal brushing and bacteriological sample	Chronic rhinosinusitis (diagnosis)	France	–	Foundation for Medical Research	Centre Hospitalier Intercommunal Creteil
Inhaled Nanosilver Study	Silver nanoparticles	Inflammatory disease	USA	–	National Institutes of Health Clinical Center (CC) (National Institute of Environmental Health Sciences (NIEHS)	National Institute of Environmental Health Sciences (NIEHS)
Role of Ivermectin Nanosuspension as Nasal Spray in Treatment of Persistant Post covid19 Anosmia	Intranasal ivermectin spray	Anosmia	Egypt	2 and 3	Zaky Aref, South Valley University	South Valley University
A Safety and Immunogenicity of Intranasal Nanoemulsion Adjuvanted Recombinant Anthrax Vaccine in Healthy Adults (IN NE-rPA)	BW1010: Intranasal nanoemulsion-based mucosal vaccine in spray	Anthrax	USA	1	Porton Biopharma Ltd.	BlueWillow Biologics
Evaluation of Silver Nanoparticles for the Prevention of COVID-19	Mouthwash and nose rinse with the silver nanoparticles suspension	Coronavirus disease 2019 (COVID-19)	Mexico	–	Bionag SAPI de CV General Hospital Tijuana	Cluster de Bioeconomia de Baja California, A.C.
Study of APH-1105 in Patients with Mild to Moderate Alzheimer's Disease	APH-1105 (alpha-secretase modulator)	Dementia Alzheimer’s disease 1 Alzheimer’s disease 2 Alzheimer’s disease 3	UK	2
A First in Human Study of the Safety, Tolerability and Pharmacokinetics of PRV-002 in Healthy Volunteers	PRV-002 (Synthetic non-naturally occurring neurosteroid) nasal spray	Traumatic Brain Injury (TBI)	Australia	1	Avance Clinical Pty Ltd.

– not applicable.

7. Nasal formulations on the market

Despite the delivery route, transnasal or i.n., pharmaceutical products of nasal administration are attractive in many therapeutic fields such as pain, frigidity, migraine, erectile dysfunction, insomnia, panic attacks, Parkinson’s disease rigidity, seizures, hot flushes, Alzheimer’s disease, cardiovascular events, and emesis. Tyzine (tetrahydrozoline hydrochloride) was the first nasal spray approved by the FDA in 1954 to treat nasal congestion due to colds, allergies, sinusitis, or hay fever¹⁵⁶. Since then, more than 32 nasal products have been approved for marketing by the FDA. Several active pharmaceutical ingredients for systemic or CNS treatments are being marketed as nasal products. Among them, sumatriptan (Imitrex®), zolmitriptan (Zomig®), and the opioid analgesic butorphanol tartrate (Stadol NS) for migraine treatment, fentanyl (PecFent®, Instanyl® for pain, and estradiol hemihydrate (Aerodiol®) for menopausal syndrome, and the number of products is increasing over the years¹⁵⁷. There are a few intranasal products on the market for CNS conditions, e.g., esketamine (Spravato®, Ketanest®) for treatment-resistant depression and depressive symptoms¹⁵⁸. However, none of the products takes advantage of nanotechnology to improve the drug efficacy. As described above, nanocarriers are more efficiently transported to the CNS than their soluble counterparts and, to succesfully translate them into pharmaceutical products, the challenge in the near future will be to demonstrate their advantage over conventional formulations (e.g., solutions).

8. Conclusions and future perspectives

Mucosal drug delivery has emerged as a strategy to target drug delivery and increase their bioavailability while reducing systemic exposure and thus, increase the therapeutic index. This strategy has become even more sound in the field of CNS therapeutics because of the limited brain bioavailability of systemically administered drugs. Owing to the more efficient transport from the nose-to-brain, nano-drug delivery systems have gained the attention of researchers and clinicians for the delivery of small-molecule and macromolecular drugs. Polymeric nanocarriers are among the most versatile platforms, enabling the tailoring of key features such as size, shape, surface chemistry and charge that control the interaction with the mucosal tissue and the molecular and cells cues governing this transport. For example, the use of cell-penetrating peptide-modified polymeric micelles to intranasally deliver small interfering ribonucleic acids has been reported¹⁵⁹. Following the same concept, shuttle peptides currently used in the targeting of the CNS from the bloodstream upon i.v. administration20, 21, 59, 160 could be also investigated for nose-to-brain delivery. Even though there exists consensus about the potential and the benefits of i.n. administration to target the CNS and relatively solid preclinical proofs of concept with a broad spectrum of nanoformulations, this therapeutic approach still faces critical issues which finds an expression in the very small number of clinical trials going on. Main issues include relatively small administrable volumes, differences in the nose anatomy between the common animal models (e.g., rodents) and humans that jeopardize the correlation of preclinical and clinical results, the need for the development of formulations and dosing devices that fit this administration route, and the still limited understanding of the pathways involved. Moreover, there is a lack of scientific evidence about the regions in the brain that can be targeted by this route and how the nanocarrier features and the administration in different areas of the nostril affect the biodistribution in the CNS. This limits our ability to target regions associated with different diseases and disorders of the CNS and they need further investigation. Finally, the improved efficacy has to be confirmed in clinically relevant animal models to then conduct clinical studies. These hurdles could be only overcome by multidisciplinary research that will address the design of advanced drug delivery systems in a more comprehensive and rational way.

Acknowledgements

This work was funded by the NEVET Nanotechnology Grant of the Russell Berrie Nanotechnology Institute (RBNI) at Technion – Israel Institute of Technology (Israel). Alejandro Sosnik thanks the Tamara and Harry Handelsman Academic Chair (Israel) for financial support.

Author contributions

Rania Awad, Avi Avital and Alejandro Sosnik wrote the original and revised versions of the manuscript. All of the authors have read and approved the final manuscript.

Conflicts of interest

The authors have no conflicts of interest to declare.