Stimuli-responsive nanogels as intelligent nanocarriers for tumor microenvironment-triggered anticancer drug delivery

Abstract

Stimuli-responsive nanogels have been utilized as perfect nanocarriers for anticancer drug delivery because of their on-demand, controlled, and site-specific drug-releasing chemistry. These HNP are cross-linked hydrophilic polymer nanoparticles with a high-water content, biocompatibility, and adjustable reactivity to chemical or physical stimuli (such as pH, temperature, and redox potential, which are among the most extensively studied triggers in cancer-targeted nanogel systems). Because of their structural flexibility, these nanocarriers can react intelligently and passively to the tumor microenvironment's high glutathione content, acid pH, and overexpressed enzymes, ensuring increased intracellular release and reduced systemic toxicity. Cross-linking strategies, top-down and bottom-up production processes, and core characterization methods concerning size, charge, morphology, and release kinetics are the main topics of this article. Anticancer medications like doxorubicin, paclitaxel, camptothecin, and docetaxel have been shown to be well accommodated in a variety of nanogels, including pH-responsive, thermo-responsive, redox-responsive, magnetic-based, and multi-responsive ones for increased bioavailability and anti-tumor activity. In addition, receptor-mediated endocytosis mediated by targeting ligands such as folic acid, hyaluronic acid, aptamers and monoclonal antibodies improves the cellular uptake and uptake in tumor of drug-loaded nanogels. Collectively, intelligence-triggered nanogels stated above possess outstanding benefits in combination therapy, controlled drug release, and theranostic application and so illustrate these as state-of-the-art intelligent delivery systems for tumor treatment. Future goals include optimizing biocompatibility, removing tumor penetration obstacles using techniques including surface charge modification, PEGylation, and enzyme-sensitive cross-linkers, and guaranteeing scalability and therapeutically transferable formulations.

Graphical Abstract

Introduction

Nanogels

The median dimension of nanogels, also called hydrogel nanoparticles (HNPs), is approximately 100 nm. They are made of water and interconnected hydrophilic polymers. They are stable, have a high moisture content, and have a substantial surface area [1]. In medical fields, nanogel-based technologies have been developed to provide both the extended circulation half-life of cargo in the body and the ability to convey that cargo to the intended site [2]. Furthermore, another crucial concern in a variety of medicinal use is continuous and regulated on-demand drug release. Consequently, there has been a lot of recent interest in ecologically responsive nanogels. In response to an exogenous stimulus, these nanogels—also referred to as stimulus-sensitive or ecologically sensitive nanogels—can alter their physicochemical characteristics, including volume, water content, refractive index, network internal, permeation, and hydrophobicity. An exogenous stimulus is an exterior provided trigger that can cause physical or chemical changes in the nanogel composition, such as swelling, deswelling, or bond breakage, hence modifying drug release behaviour. Examples of such triggers include temperature changes, magnetic fields, and light irradiation [3].

Such stimuli are often divided into two categories: chemical stimuli, that involve fluctuations in pH, ionic concentration, and chemical or biological compounds, and physical stimuli, that comprise variations in light, temperature, and magnetic forces. Investigation on the nanoscale modifications of hydrogels was prompted by all these appealing qualities [4]. Despite their widespread use, these hydrogels had a major disadvantage in stimuli-responsive systems wherein the rate of dispersion hampered signal transduction. Subsequent research showed that effective signal propagation within bulk hydrogels was hampered by the sluggish dispersion rate. In order to get over this restriction, scientists created interconnected capillary-like pore networks inside the polymeric framework and shrunk the hydrogel's size to the nanoscale, which improved diffusion and responsiveness and resulted in the creation of nanogels. There are several benefits that nanogels offer Fig. 1, which enumerates the main physicochemical and medicinal benefits of nanogels; however, these features can also be contextually included in the discussion that follows [1].

Fig. 1.

List of various advantages of nanogel

Drug components can be protected and their dispersion from nanogels can be controlled by adding high-affinity functional groups, stimulus-sensitive configurations, or biodegradable bonds to the polymer network [5]. Using hydrophobic, van der Waals, and/or electrostatic forces, the loaded medication combines with the matrix of polymers. Nanogels' hydrophilicity makes it simple for the particles to dissolve in aqueous conditions and, consequently, to be given in liquid dosage form [6]. By adding hydrophilic polymers, like PEG, nanogels are protected by being absorbed by the mononuclear phagocytic pathway and have a longer circulatory circulation duration, which increases longevity. These nanogels can be targeted with specific moieties like transferrin’s, folic acid (FA), and monoclonal antibodies (mAb) thanks to the functional modification of the various functional groups accessible on their exteriors [7]. Nanogels exhibit remarkable stability by distributing their contents throughout cells and beyond biological barriers while protecting them from cellular metabolic degradation [8, 9]. Therefore, nanogels offer great potential for improving the systemic drug administration of low molecular weight medications as well as their oral and cerebral permeability [10].

Water-soluble polymers that are physically or chemically crosslinked form nanogels, nanoscale 3D hydrogels with high permeability, ECM-like mechanical strength, and excellent biocompatibility. Their high-water content allows efficient loading of proteins, nucleic acids, and other bioactives, making them suitable for drug delivery and cancer therapy. Recently, stimuli-responsive nanogels have gained attention for controlled and targeted drug release. These systems undergo swelling or degradation in response to biological cues such as pH, enzymes, ROS, reducing agents, or ATP, which are abundant in the tumor microenvironment. Such bio responsive nanogels are advantageous due to their site-specific drug release, cost-effectiveness, and improved patient compliance compared to externally triggered systems [11]. They are usually synthesized by crosslinking natural or synthetic polymers or by using active crosslinkers. Techniques such as inverse mini-emulsion, nanoprecipitation, microfluidics, and in-situ self-assembly enable precise control of nanogel size and shape [12]. Chemically crosslinked nanogels exhibit high stability in circulation but rapid response in tumor environments, enabling selective drug release [13]. Advances in polymer chemistry and ligand conjugation (e.g., with folic acid, antibodies, peptides, polysaccharides) enhance target specificity and cellular uptake. Moreover, dual-ligand functionalization further improves tumor penetration and drug accumulation, making bio responsive nanogels a promising platform for targeted cancer therapy [14].

Types of stimuli-responsive nanogels

Endogenously responsive stimuli responsive nanogels

Redox-responsive, enzyme-responsive, pH-dependent, and ionic microenvironment-responsive drug delivery systems make up the class of endogenous stimulus of biochemical and chemical roots. Through control of the microenvironment tissues, overexpression of target-oriented enzymes, antibody-antigen contact, and identification between host–guest moieties in a particular state, these DDS activate drug delivery. In the presence of biological stimuli like pH, enzymes, ROS, reducing agents, or ATP all of which are present in greater concentrations within the tumor microenvironment endogenously triggered nanogels are made to expand, deteriorate, or break down [15, 16]. Target-controlled drug release, increased intracellular delivery efficiency, and less systemic toxicity are all made possible by these bio responsive nanogels [17].

For redox species with varying concentrations, redox-responsive materials may be involved in both extracellular and intracellular regions of sick tissues or cells, Table 1. Compared to intercellular (endosomes) and external fluids, the concentration of a reducing agent like glutathione is twice as high in the cellular cytosol and nucleus. Extracellular and intracellular glutathione concentrations are correlated with the redox potential that exists between extracellular (oxidative) and intracellular (reductive) space. Securing and releasing the payload within the cell is crucial when delivering DNA or siRNA [18]. Redox-sensitive nanocarriers are consequently more advantageous for intracellular distribution, especially for gene transfer, in light of these generalizations. The majority of glutathione-responsive entities resemble pH-responsive materials with acid-labile groups in terms of design. In glutathione-responsive block copolymers, it is typically a disulfide bond among hydrophilic and hydrophobic blocks that results in the production of micelles, commonly referred to as "shell-sheddable" [19]. These micelles become unstable and release the medicinal substance when they come into contact with glutathione [20]. The disintegration of the nanostructure is caused by the oxidation of sulfide entities to create hydrophilic sulfoxide and the disruption of the hydrophilic/hydrophobic equilibrium in the solution assemblies from sulfone moieties [21].

Table 1.

Comprehensive Classification and Characteristics of Nanogels

Type of Nanogel	Category	Stimulus / Trigger	Composition or Functional Modification	Mechanism of Drug Release	Applications	References
pH-Responsive Nanogels	Endogenous	pH variation in local microenvironment (acidic tumor tissues, inflamed regions)	Ionizable polymers such as chitosan, poly(acrylic acid), poly(methacrylic acid), or poly(β-amino esters)	Protonation or deprotonation causes network swelling or collapse, triggering controlled drug release	Tumor-targeted therapy, oral drug delivery, infection site-specific release	[34, 35]
Thermo-Responsive Nanogels	Exogenous	Temperature fluctuation near physiological range (37–42 °C)	Polymers showing LCST or UCST such as PNIPAM, poly(N-vinyl caprolactam), PEG-based copolymers	Volume phase transition causes structural change and triggers drug diffusion	Local hyperthermia-mediated cancer therapy, inflammation-triggered release	[36, 37]
Redox-Responsive Nanogels	Endogenous	Elevated intracellular GSH or ROS levels	Incorporation of disulfide, diselenide, or thioketal linkages in polymer or crosslinker	Reductive cleavage of disulfide bonds leads to matrix degradation and drug liberation	Intracellular drug delivery, cancer chemotherapy, antioxidant delivery	[38]
Enzyme-Responsive Nanogels	Endogenous	Enzymes overexpressed in diseased tissues (e.g., matrix metalloproteinases, proteases, lipases)	Incorporation of enzyme-cleavable peptides or polymeric linkers	Enzymatic hydrolysis or degradation of crosslinks releases loaded drug	Targeted cancer or inflammatory disease therapy, wound healing	[39]
Magnetic Nanogels	Exogenous	External magnetic field	Integration of Fe₃O₄ or γ-Fe₂O₃ nanoparticles into polymeric network	Magnetically induced local heating or guidance to target site enables controlled release	Magnetically targeted drug delivery, MRI contrast agents, hyperthermia therapy	[40, 41]
Photo-Responsive Nanogels	Exogenous	Light irradiation (UV, visible, NIR)	Incorporation of photo-labile or photochromic groups (e.g., azobenzene, o-nitrobenzyl)	Light exposure induces structural rearrangement or bond cleavage, causing drug release	Photodynamic and photo-triggered cancer therapy	[42]
Ultrasound-Responsive Nanogels	Exogenous	Application of ultrasound waves	Polymer matrices sensitive to acoustic cavitation or microbubble inclusion	Mechanical disruption or heating due to ultrasound exposure releases drug	On-demand release in deep tissues, sonodynamic therapy	[43, 44]
Ionic Strength-Responsive Nanogels	Endogenous	Variation in ionic concentration of biological fluids	Polymers with ionic groups such as poly(acrylic acid) or polyelectrolyte copolymers	Ion-induced contraction or expansion of gel network modifies diffusion rate	Controlled release in GI tract, ocular and mucosal delivery	[45]
Multi-Responsive Nanogels	Endogenous	Two or more combined stimuli (e.g., pH + redox, pH + temperature, or pH + enzyme)	Hybrid or copolymer systems (e.g., PNIPAM–chitosan–disulfide composites)	Synergistic structural response to multiple triggers enables smart drug delivery	Complex pathophysiological targeting (e.g., tumor + inflammation)	[46, 47]
Stimuli-Sensitive Nanogels (General)	Endogenous	pH, redox, temperature, enzyme, or magnetic stimuli	Smart polymeric matrices with functional monomers sensitive to multiple stimuli	Swelling, shrinking, or bond cleavage causes on-demand release	Controlled release systems, biosensors, diagnostics	[4]

For a variety of biological uses the pH-sensitive nanogel should be stable and responsive to a slightly reduced and physiological pH values (5.0–6.5, 7.4). For the goal of medication delivery, this layout facilitates the release of medicines within the cell and surrounding tissues [1]. The acid-functionalized groups are incorporated into a polymeric backbone in the first method. The pH response, which might coincide with conformational shifts in every segment of the polymeric backbone, may be triggered by these active pharmacological groups. This conformational shift modifies the nanostructures, potentially leading to the polymer's self-assembly [22]. As a result of the cleavage of the pH-triggered link, the copolymer's amphiphilic equilibrium is upset, which leads to the nanocarrier's degradation or self-assembly breaking and the release of encapsulated medicines [23].

Designing responsive elements for DDS heavily relies on enzyme-based modulation and/or disruption in the intracellular milieu and their participation in all biological and metabolic functions. Features including biorecognition, sensitivity, selectivity, and catalytic efficacy are provided by enzyme-responsive DDS [24]. When the polymeric moiety in enzyme-responsive DDS is broken down by enzymes, the medications are liberated from the drug-loaded carriers. By adding attached acidic or basic functional entities to the polymer backbone, ionic microenvironment-responsive nanogels are created [25]. These entities have a significant impact on the degree of ionization. The quantity of pendant acidic or basic groups determines the drug release mechanism. At high pH, a large concentration of acidic groups increases electrostatic repulsion and swelling ratios, while basic groups ionize at low pH, producing comparable effects [26].

Exogenously responsive stimuli responsive nanogels

In contrast to endogenous stimuli-responsive DDS, exogenous ones have the advantage of overcoming inter-patient variability potential (due to the fact that in these systems drug release is dedicated by externally applied field which can be accurately controlled. Thermo-responsive polymers are thus pivotal in the design of these DDS as most of the stimuli work via heat generation [17]. If so, it can trigger the release of selected drugs by the help of the temperature-sensitive materials. Exogenously responsive nanogels offer spatial and temporal control over drug release, enable non-invasive activation, and can be engineered to respond to externally applied physical stimuli for on-demand therapeutic efficacy, Table 1 [27].

Exogenous systems with temperature or thermo-responsive DDS have been extensively studied. When a drug is exposed to a greater temperature, thermoresponsive drug carriers release the drug. Response to either externally applied heat or intrinsically elevated tumor temperature are two ways [28]. Temperature-dependent phase transition behaviour is shown in PNIPAM-based systems; above LCST, the polymer turns hydrophobic and regulates drug release [29]. Additionally, photothermal or magnetic heating can be used to externally stimulate drug release. In order to accomplish targeted drug release with improved spatiotemporal control, photo/light responsive DDS devices are being investigated extensively. Various wavelengths are employed, including visible, UV, and near-infrared (NIR). Because of its superior tissue penetration, NIR is favoured [30]. Upconverting nanoparticles, two-photon activation, and the photo-thermal effect are among the mechanisms. By magnetic targeting and magnetic hyperthermia, magnetic field responsive technologies allow for regulated medication release. Drug release is triggered by the production of heat by super-paramagnetic iron oxide nanoparticles in an alternating magnetic field [31]. Additionally, imaging and theranostic uses are made possible by magnetic systems. Under magnetic stimulation, implantable magnetic hydrogels have demonstrated pulsatile drug release. Deep tissue penetration and spatiotemporal control for medication release are made possible by ultrasound waves [32]. Drug release is stimulated by cavitation, mechanical, and thermal processes. While mechanical vibrations destabilize nanocarriers and increase drug release, ultrasound-induced heat can cleave thermoresponsive linkers [33].

Nanogels' inherent characteristics

Biocompatibility

Because they function without any negative effects, nanogels are regarded as biocompatible. As it aids in determining suggested biomedical uses, cytotoxicity analysis is crucial in the study of biomaterials. Tests for cytotoxicity and biological compatibility are frequently conducted in vitro. While this is occasionally dose- and time-related, nanogels are generally considered to be non-cytotoxic. According to this procedure, cytotoxicity experiments can be validated using morphological assessments of cell growth, impairment, and specific features of cell metabolism [48].

The superior biocompatibility and cellular compatibility of nanogels as possible drug delivery methods have been shown in numerous research. Human mesenchymal stem cells (hMSCs) were not harmed by a doxorubicin-loaded nanogel, which demonstrated an encapsulation rate of about 86% and successfully killed MG-63 carcinoma cells in vitro [49]. According to Pereira et al., glycol chitosan nanogels (GC-nanogels) exhibited great biocompatibility and were appropriate for systemic distribution because they failed to activate the complement mechanism, react with RBC, or cause thrombosis [50]. In a comparable way, Wei et al. created photoinitiator-functionalized nanogels composed of PEGDMA and OEOMA that demonstrated exceptional cellular compatibility on HeLa cells in MTT tests [51]. Additionally, Guo et al. developed positively charged chitosan nanogels (CS/HCPT) that contained 10-hydroxycamptothecin, and this demonstrated significant cytotoxic activity toward 4T1 breast cancer cells and high biocompatibility, demonstrating their efficacy to be secure and effective means of drug delivery [52].

Aqueous media's swelling property

One of hydrogels' inherent characteristics is the swelling phenomena. Nanogels can swell more than traditional hydrogels because of their huge surface area. A high fluid interchange capability with the environment is produced by nanoformulation [53]. The spacing among cross-link sites, the amount of reactive molecules per cross-link chain, the hydrophilic nature of the cross-linker, along with the nanogel shape all influence the swelling process. Pikabea et al. used a multifunctional reagent and two dual-purpose cross-linkers to create poly methacrylate-based nanogels. Because of their greater hydrophilicity, the PEGDA cross-linked nanogels proved to expand the most, whereas the multipurpose Dex40MA86 interconnected nanogels had a lower swelling capacity than the other forms of nanogels [54]. Furthermore, endogenous (such as enzymes, pH, reduction, etc.) or exogenous (such as light, temperature, and magnetic forces) stimuli affect the pace at which nanogels inflate. [54]. Tamura et al. showed that PEGylated nanogels swell in response to pH, with cross-linking density influencing size variations. Protonation caused the nanogels to inflate below their pKa, demonstrating their capacity for pH-sensitive release of medications [55]. Daniel et al. created κ-carrageenan-based nanogels that demonstrated thermo-responsive swelling between 37 and 45 °C, where the hydrodynamic diameter increased depending on both temperature and polymer composition, suggesting their applicability for medication delivery purposes that are sensitive to temperature [56].

Increased drug delivery and loading efficiency

Each medication delivery mechanism must incorporate drug molecules within nanogels; the greater the drug's loading ability, the smaller the drug's nanocarrier. Nanogels possess a significant medication loading capability and encapsulating efficacy due to their tendency to swell. Physical retention, including hydrophilic–hydrophilic, hydrophobic–hydrophobic, bonds of hydrogen, Van der Waals/electrostatic interactions, covalent attachment of physiologically active compounds, or controlled self-assembling (that can be regulated by a variety of variables like as pH, light, temperature, interaction of ions, etc.), is how drug loading happens, just like with all other nanoparticles. The chemical makeup, molecular mass, nanogel particle dimension, hydrogel, and various functional groups of the hydrogel unit can all have an impact on the load-carrying potential [57].

Furthermore, a number of variables, such as the kind of drug, how it interacts with the nanogel, the composition of the nanogel, and the surrounding environment, influence the release of medicines from nanogels [58]. Diffusion regulation, swelling regulation, and chemical regulation are the three categories of drug release methods for nanogels. The drug release technique is determined by the physicochemical properties of the nanogel and how the medication gets integrated into it. [59].

Stability of colloids

The fundamental physicochemical characteristic of nanoparticles is their colloidal resilience, which is dictated by their physical and chemical composition. Environmental elements including pH, salt, temperature, proteins, organisms, and so forth have a significant impact on colloidal stability [60]. The coating material controls the stability of colloidal characteristic since it is the very initial component of the particle that gets into proximity to the surroundings. As a result, the coating material offers colloidal strength and instant biocompatibility all over the duration of its life. Alginate, chitosan, hyaluronic acid, and charged dextran analogues are examples of charged polysaccharides that have been shown to enhance and prolong stability of colloids in biological mediums. Additional biopolymers that are becoming more significant in the stability of colloids and nanomedicine are proteins [61].

Low immunogenicity

Nanogels have exceptional electromagnetic, biological, chemical, and physical attributes. Nanogels as drug carriers exhibit low toxicology, non-immune reactions or insignificant immunogenic reactions, and appropriate biocompatibility and ability to degrade with non-toxic products of degradation to improve their targeted efficiency in cancer treatment, along with to their highest curative effectiveness with few adverse reactions and preferred site-specific discharge responses. The majority of nanogels are now only used in pre-clinical laboratory testing; they need to be investigated more effectively for in vivo trials in order to move toward next-generation clinical applications for precise and customized healthcare [62].

Development of stimuli-responsive nanogels

The polymerization of monomers, the physical or chemical cross- linkage of polymer precursors or natural polymers, and template-aided nanofabrication are the three primary ways of nanogel formation, Fig. 2 [63].

Fig. 2.

A graphical representation of various nanogel preparation techniques [11]

There are two main methods for engineering nanogels at the nanoscale: "top-down" and "bottom-up." The top-down approach uses mechanical, chemical, or physical techniques to create nanogels from bigger hydrogel particles [64]. The photolithographic technology known as Particle Replication in Non-Wetting Templates (PRINT) was developed by Rolland et al. It involves shaping liquid precursors within non-wetted elastomer casts made of perfluoropolyether frameworks on patterned silicon frames. This technique produces monodisperse particles that are appropriate for encasing sensitive biomolecules and medications by precisely controlling particle dimensions, form, and content while reducing interfacial film development [65].

Peptide-based hydrogels (HGs) and nanogels (NGs) were created by Gallo et al. employing hydrogelator Fmoc-FF either by itself or in conjunction with (FY)₃ or PEG₃-(FY)₃ peptides. Using a top-down method, they were sustained and submicronized using biocompatible surfactants. Doxorubicin (Dox) was well enclosed by both HGs and NGs, which also demonstrated outstanding biocompatibility, retaining > 95% of cell activity for emptied gels and exhibiting 49–57% vitality when Dox-loaded in MDA-MB-231 cells. While NGs showed component-dependent drug loading (DLC 0.093–0.137) and rates of release (20–40% after 72 h), the hydrogel makeup affected gelation kinetics (24–40 min) and drug release (16–28% over 72 h) [66]. Similarly, Rosa et al. used water/oil emulsion, top-down, and water nanogelling to create peptide nanogels from Fmoc-FF. The top-down approach was beneficial since it created biocompatible nanogels (~ 200 nm) that are perfect for use in medicine, eliminated mineral oils and organic solvents, and required fewer processes [67].

The bottom-up method, on the other hand, creates nanogels from molecular precursors that undergo chemical or physical cross-linking. By combining acryloyl-modified cholesterol-bearing pullulan (CHPOA) with thiol-modified PEG (PEGSH), Shimoda et al. created nanogels. Particle sizes of 50–150 nm were achieved by varying the percentage of nanogel, acryloyl replacement, and acryloyl:thiol ratios. Injectable carriers for prolonged cytokine release were provided by the resultant CHPOA–PEGSH nanogels [68]. Similarly, Sekine et al. produced biodegradable hydrogels by connecting four-armed PEGSH with CHPOA, resulting in nanogels and nanogel-coated liposomes with molecular chaperone-like function that can encapsulate hydrophobic medicines, proteins, and DNA [69]. To create stimuli-responsive nanogels, monomer synthesis and prepared polymer cross-linked in heterogeneous mediums are frequently used. As water-soluble monomers nucleate to produce colloidal suspended particles, their starting charge keeps nanogels stable. This technique is useful for trapping tiny molecules and biomacromolecules, but it frequently requires complicated purification and may be contaminated by solvents and surfactants, which limits its application in biomedicine [70].

On the other hand, biological macromolecules can be mildly encapsulated without the use of toxic solvents through physical self-assembly in aqueous circumstances. Especially interesting are self-constructing nanogels made of peptides and polysaccharides. Tai et al. used the self-assemblage peptide Fmoc-FRGDF in conjunction with fucoidan to create extracellular matrix-mimicking nanofibrils. Their usefulness as biodegradable tissue-specific transports was highlighted by the resulting hydrogel's enhanced mechanical strength, increased hesperidin encapsulation productivity, and regulated release patterns [71]. However, because they depend on weak hydrophobic, electrostatic, or hydrogen-bond bonds, physiologically cross-linked nanogels may be unstable. These may split into polymers when diluted in biological fluids, resulting in early release of medication. Chemical cross-linking after physical self-assembly is a potential enhancement that improves nanogel stability and eliminates the need for surfactants [72].

To adjust medication distribution, bioavailability, and cytotoxicity, cross-linking is essential. By utilizing solely polymer and water as substrates while minimizing hazardous catalysts or additives, photo- and radiation-triggered cross-linking offer safer substitutes for chemical activation. In radiation-induced production, radicals are produced in aqueous polymeric solutions by ionizing radiation (gamma or electron beam), which starts intramolecular cross-linking. Managing circulation duration, viscosity, and loading of drugs in clinical settings requires oversight of nanogel size. Sütekin et al. developed a consistent method for producing poly(N-vinyl pyrrolidone) (PVP) nanogels using electron beam (NGEB) or gamma irradiation (NGG). By modifying factors including absorbed dose, dose speed, polymer content, and molecular weight, nanogels between 30 and 250 nm, ideal for biomedical use, were created. Radiation-based manufacturing is consequently a versatile, sanitary, and efficient method for a range of water-soluble polymers [73].

In general, both top-down and bottom-up production methods have definite advantages. Top-down approaches provide precise geometric control and simple production, but bottom-up and radiation-based approaches enable configurable chemical activity and biocompatibility. The intended drug delivery application anticipated release kinetics, and mechanical strength all influence the approach choice [11].

The application of stimuli-responsive nanogels in the delivery of anti-cancer drugs

Nanogels have opened up new treatment options with their advantageous ability to administer anticancer drugs with non-specific toxicity and a short therapy duration. Because smart nanogels are dependent on pH and temperature and enable the adherence of further functional units targeted to cancer cells, this nanoscopic carrier provides an enhanced framework for delivering anticancer medications [74, 75].

Stimuli-sensitive nanogels

Responsive nanogels are increasingly regarded as intelligent medication delivery systems for targeted and regulated delivery of medications, particularly in cancer therapy, since they enable precise and extended administration of medicines [76]. When such nanogels experience structural or chemical modifications in reaction to a range of external stimuli, such as temperature, pH, ion concentration, light, and magnetism, the encapsulated molecules of medicine are released. Because they exhibit reversible swelling and deswelling in accordance with variations in temperature, temperature-responsive nanogels are particularly important for biomedical applications [77]. For example, Juan Wu et al. created doxorubicin (Dox)-loaded PEG–poly(ethyl ethylene phosphate) triblock copolymer-based nanogels that self-assembled when heated [78]. These nanogels, which ranged in size from 100 to 450 nm, displayed high cytotoxicity in A549 lung cancer cells and decreased hydrodynamic width with rising temperature, demonstrating their effectiveness in drug administration [17, 79]. In a similar vein, Weitai Wu et al. created core–shell nanogels for integrated chemotherapy and photothermal treatment that had a bimetallic (Ag–Au) core, a polystyrene shell, and an outside PEG layer. This resulted in controlled dissolution of curcumin and notable cytotoxicity in B16F10 melanoma cells [80]. In a different work, You-Yong Yuan et al. described biodegradable and biocompatible branching nanogels made from TEGDP and TREN that showed effective absorption by MDA-MB-231 cancerous breast cells and enzyme-dependent doxorubicin discharge [81]. Super-expandable nanogels with an 800-fold reversible volume shift (from 150 nm at 37 °C to 1.4 µm at 15 °C) were created by YLee et al., causing necrosis in HeLa cells and potentially acting as carriers for hydrophobic anticancer medications [82]. All of these results demonstrate that thermo-responsive nanogels combine biologic compatibility, variable physicochemical traits, and multifunctional medicinal properties to provide stimuli-triggered, site-specific, and controlled delivery of medication, rendering them highly desirable for innovative cancer therapy utilization [83].

pH-responsive nanogels

Because pH-sensitive nanogels can change the kinetics of drug release showing delayed absorption throughout circulation and rapid dissolution at acidic tumor or endosomal pH they have drawn a lot of interest as clever ways to deliver medicines for cancer treatment. Chitosan–ovalbumin nanogels (100 nm) with pH-related size and hydrophobicity were created by Shaoyong Yu et al. [84] and show promise for uses in medication delivery. In order to improve anticancer efficacy in HuH-7 cells, Motoi Oishi et al. [85] created PEGylated Dox-loaded nanogels that crossed a volume phase shift (145–165 nm) and demonstrated improved release of drugs at endosomal pH. Doxorubicin (Dox) and camptothecin (CPT) were successfully encapsulated by other chitosan-based nanogels made by inverse microemulsion or crosslinking techniques, demonstrating enhanced removal of tumors and decreased cardiotoxicity [86, 87].

In the same way, NIPAAm/chitosan nanogels demonstrated pH-sensitive CPT and paclitaxel diffusion, resulting in increased cytotoxicity at acidic pH (6.8) and total tumor regression in over half of the treated mice. CS-g-PNIPAm nanogels for Oridonin delivery were described by Duan et al. [88], allowing for acid-induced discharge and increased cytotoxicity in HepG2 cells. Additionally, Dox-loaded chitin nanogels (130–160 nm) demonstrated robust fluorescence-based imaging of several cancer cell lines and pH-sensitive release. Under acidic endosomal conditions, PEG-b-P nanogels containing Dox prodrugs discharged medications preferentially, demonstrating effective intracellular administration. Lastly, a virus-mimetic nanogel made of PEG-BSA shells and poly(L-histidine-co-phenylalanine) core showed pH-responsive release of drugs and improved curcumin skin penetration, exhibiting specific cytotoxicity against melanoma cells (A375) [89].

Redox-responsive nanogels

Redox-responsive nanogels degrade apart and deliver drugs within the cytosol by utilizing the intracellular reducing conditions. The method's primary intracellular redox agent, glutathione (GSH), is found in cells at concentrations 100–1000 times higher (2–10 mM) compared to external fluids (2–20 µM). This strong redox gradient enables accurate and quick distribution of drugs into cells [90]. In order to achieve size control, high stability, GSH-responsive release of encapsulated medicines such as DOX, and effective cellular absorption in MCF-7 cells, Ryu et al. [91] produced polymeric nanogels with oligo(ethylene glycol) (OEG) and pyridyldisulfide (PDS) side chains. All things considered, redox-responsive nanogels offer an effective and targeted intracellular delivery method for nucleic acids and anticancer drugs [92].

Magnetic nanogels

Han Wen et al. experimented with using magnetic nanogels in MRI. They coated Fe3O4 nanoparticles beneath UV light to create superparamagnetic poly (ethylene glycol) methyl ether methacrylate (PEGMA) nanogels. They created 68 nm-sized spherical core–shell nanogels on average. They investigated the drug discharge from the nanogels using DOX as a test drug and discovered that it was sluggish, which makes it a viable option for targeted drug administration [93]. N-2-(aminoethyl) methacrylamide hydrochloride was reported to be used in the photochemical preparation of amino functionalized magnetic nanogels, and its diagnostic capability as MRI contrast compounds was assessed. According to Gong et al., the magnetic nanogels exhibit super paramagnetic reactions, high magnetic density, substantial saturation magnetizations, and excellent durability for storage [94].

Targeted delivery strategies of nanogels

Intravenous injection and intratumoral injection of nanogels for anticancer drug delivery

To address issues and enhance treatment efficacy, the intratumoral route of nanoDDS was suggested. To maximize the continuous retention of nanoDDSs in tumor’s, RES capture can be omitted due to their direct injections into tumors tissues [95]. Additionally, a potential aspect of nanoDDSs for intratumoral delivery is their ability to provide diffusion control, which will prevent medications from rapidly dispersing and washing out from tumor tissues into adjacent normal tissues and distant circulatory systems. Compared to systemic treatment, intratumoral delivery of nanoDDSs is known to ensure better therapeutic efficacy and less off-target toxicity [96].

Intratumorally given NanoDDSs showed much more local tumor retention and reduced systemic spread. For instance, compared to intravascular delivery, intratumorally administered nanoparticles showed 13.1 times greater levels inside the tumor and decreased hepatic accumulation [95]. Similarly, when delivered locally as opposed to intravenously, nano shells localized to tumor tissue retained 8.2 times more [97]. It was confirmed that better tumor accumulation of nanoDDSs translated into a considerable reduction (> 30%) in tumor’s when administered intratumorally as opposed to intravenously [98]. Because of their longer intratumoral retention period and slower discharge, nanoDDSs provide better therapeutic results than tiny molecular medicines. Small molecular medicines are rapidly cleared from tumor tissues due to their extreme diffusivity; however, nanoDDSs control diffusivity and lessen the tumoral escape effect, resulting in a targeted therapeutic activity. However, the optimal nanoDDSs for intratumoral delivery must be uniformly distributed throughout the tumor parenchyma with extended retention periods and high cancer cell active uptake. Particle size, surface properties, the tumor microenvironment, and injection parameters are some of the factors that affect intratumoral dispersion and retention [99].

The transport of nanomedicines in a solid tumor may be severely hampered by the features of the tumr microenvironment, such as high interstitial fluid pressure, aberrant vasculature, and extensive extracellular matrix. These obstacles may cause washout from tumor areas and prevent uniform adhesion [100]. The most popular method for practical usage of nanoDDSs, such Doxil, which has been approved to treat cancer, is intravenously. However, quick liver and spleen uptake, inadequate tumor accumulation, and variable increased permeability and retention impact hinder nearly all systemic administration [95, 101].

On the other hand, intratumoral injection can result in reduced tumor retention, less systemic toxicity, and a high local drug concentration. Through intratumoral administration (e.g., direct or image-guided injection), they can be administered directly into disease areas, where precision dosage delivery can optimize on-tumor effects while limiting off-target exposure [102]. Benefits of intra-tumoral administration using nanogels and other nanoDDS include improved tumor penetration, extended drug release kinetics, and increased encapsulation efficiency. Intratumoral delivery of nanoparticle-based formulations may have better anticancer activity, better spatial dispersion, and fewer systemic side effects than systemic treatment [102].

Limitations

While it cover a broader range of stimuli-responsive nanogels, several drawbacks have not been sufficiently addressed. Importantly, the heterogeneity of the tumor microenvironment, which results in highly varied and historically unpredictable clinical drug release, is a big concern because pH, enzymes/lox production, and GSH levels fluctuate greatly among tumor types even in different parts of an individual tumor [103]. Higher intracellular GSH concentrations are essential for redox-responsive systems in particular, but normal tissues also contain substantial levels of GSH, raising concerns regarding off-target drug release and increased toxicity in relation to biodistribution [104]. Furthermore, the penetration and uniform distribution of intratumorally injected nanogel within solid tumors are significantly hampered by physical barriers unique to solid tumors, such as dense extracellular matrix, abnormal vascularization, and high interstitial fluid pressure. The adsorption of serum proteins can cover targeting ligands (such as folic acid, hyaluronic acid, and antibodies), resulting in reduced in vivo targeting efficacy. This is another significant disadvantage of actively targeted nanogels: they do not take into account the development of protein[105]. The majority of the systems discussed are virtually limited to in vitro or preclinical study, and surprisingly few are even close to clinical-grade application. Concerns about large-scale manufacturing, batch-to-batch reproducibility, and regulatory prospects have been mostly overlooked. The intrinsic trade-off between stability and responsiveness, where chemically crosslinked systems compromise their responsiveness but physically crosslinked nanogels may prematurely disintegrate under physiological conditions, remains unresolved [106].

Discussion

In this review, we discuss numerous unresolved problems that restrict their current translational potential, stimuli-responsive nanogels exhibit considerable promise as intelligent drug delivery systems. The accuracy and dependability of controlled drug release are hampered by relying on endogenous triggers like pH and redox gradients, which are highly dependent on biological circumstances that can change greatly (and frequently overlap with normal physiological ranges). Overall, a lower therapeutic index and early drug leakage from the drug system can result from the slight pH changes between tumor and normal tissues as well as the abundance of GSH in healthy cells. Accordingly, while active targeting strategies using ligands such as folic acid, hyaluronic acid, and monoclonal antibodies have been thoroughly investigated to enhance the uptake of nanoparticles by cells in vitro, their effectiveness is frequently reduced in vivo due to protein crown effects, variable receptor expression, and endocytic pathways; there is a gap between experimental and clinical success. Furthermore, while intratumoral administration improves local retention with a loss of post-injection distribution, intravenous treatment is limited by rapid clearance and RES uptake along with inconsistent tumor accumulation.

However, the manuscript speculates on the benefits of intratumoral and intravenous delivery routes without comparing their drawbacks. Lastly, the aforementioned literature primarily reports favorable results, such as enhanced cytotoxicity or drug release efficacy, with a few exceptions showing negative effects and ostensibly no comparison to alternative standards, which introduces bias in evaluating improvement by nanogel systems. Crucially, while problems like limited tumor penetration and scalability are mentioned, there is very little characterization of the mechanisms underlying these behaviors or strategic outlines for solutions. This highlights a glaring imbalance in the review, leaving many questions about nanogels unanswered with only vague observations on what currently works, what doesn't work but might work well, and where future efforts should be concentrated.

Summary and conclusion

Because of their high drug-loading value, tunable physicochemical characteristics, and capacity to react selectively to internal or external stimuli, stimuli-responsive nanogels represent a broad and novel class of nanocarriers with substantial potential for targeted anticancer drug delivery. They have demonstrated promising preclinical outcomes for several anticancer drugs and enable regulated and site-specific drug release. However, a number of significant obstacles, including the variety of tumor microenvironments, variations in biological triggers (pH, glutathione), and similarities to normal tissues that limit the specificity of stimulus-driven release, limit their clinical application. Furthermore, biological barriers such low interstitial pressure (high), high interstitial pressure (glycoprotein), biofilm softening, and inadequate vascular perfusion severely limit targeting accuracy and treatment consistency. Despite having theoretical advantages over chemotherapeutics, ligand-based targeting and multiresponsive systems frequently work poorly in vivo due to biological obstacles such cellular competition and patient-specific variability. Furthermore, issues with regulatory approval, repeatability, and large-scale synthesis are still not adequately addressed. Therefore, future research should concentrate on ways to get around these restrictions through improved biological knowledge of tumors and the creation of practical methods to lessen the effects of protein corona; enhance tumor penetration and distribution; and establish scalable and standardized manufacturing procedures. The ability of nanogel systems to transition from experimental platforms to clinically useful cancer treatments will depend on an honest assessment of both their advantages and disadvantages.

Abbreviations

DLS

Dynamic light scattering

PEG

Polyethylene glycol

PNIPAM

Poly(N-isopropylacrylamide)

GSH

Glutathione

Dox

Doxorubicin

CPT

Camptothecin

PTX

Paclitaxel

DTX

Docetaxel

HA

Hyaluronic acid

FA

Folic acid

mAb

Monoclonal antibody

CLSM

Confocal laser scanning microscopy

SEM

Scanning electron microscopy

Cryo-EM

Cryogenic electron microscopy

Author contributions

Conceptualisation (Lead), Methodology (Lead), Investigation (Lead), and Writing—review and editing (Supporting): S.P.N.B. Formal analysis (Lead), Interpreted the data (Lead) and writing—original draft preparation (Lead): A.W., S.J. and K.K.S. Validation (Lead), Visualisation (Lead), Project administration (Lead) and Writing—review and editing (Supporting): A.K. and A.D. Data curation (Lead), Supervision (Lead) and Writing—review and editing (Supporting): N.V., P.S. and A.G. The authors read and approved the final manuscript.

Funding

The author(s) received no specific funding for this work.

Data availability

No datasets were generated or analysed during the current study.

Declarations

Competing interests

The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper.

Ethics approval and consent to participate

Not applicable.

Clinical trial number

Not applicable.

Consent for publication

Not applicable.