Polymeric microneedles for closed-loop drug delivery: from biosensing to therapy

Abstract

Polymeric microneedles (PMNs) have emerged as a transformative platform for transdermal drug delivery and minimally invasive biosensing. By combining tunable polymer chemistry, microfabrication strategies, and bioelectronic integration, PMNs enable controlled delivery of small molecules, biologics, and vaccines while simultaneously facilitating real-time monitoring of physiological biomarkers. Recent advances have accelerated the development of closed-loop microneedle systems capable of autonomous sensing, feedback-actuation, particularly for chronic disease management. This review provides a comprehensive and critical analysis of polymeric microneedle technologies, covering polymer classification, material properties, fabrication techniques, mechanical and biological performance, biosensing integration, and closed-loop therapeutic applications. Emphasis is placed on stimuli-responsive polymers, hydrogel-forming microneedles, and glucose-responsive systems as paradigms for next-generation precision medicine. Manufacturing scalability, regulatory considerations, and translational challenges are discussed, alongside future perspectives for smart, wearable, and self-regulated microneedle platforms.

Polymeric microneedles enable simultaneous transdermal drug delivery and real-time biosensing, advancing toward closed-loop, stimuli-responsive systems for precision medicine and chronic disease management. The GA image was created using Google Gemini.

1. Introduction

Transdermal drug delivery systems (TDDS) provide minimally invasive alternatives to oral and injectable administration, offering advantages such as avoiding first-pass metabolism, maintaining sustained plasma levels, improving adherence, and reducing pain and infection risk.¹ However, the skin's outermost barrier, the stratum corneum presents a highly efficient barrier that severely limits the passive diffusion of many therapeutic agents, especially macromolecules (>500 Da molecular weight), vaccines, and hydrophilic drugs, which severely restricts the clinical efficacy of traditional transdermal patches and creams.^2,3

Only drugs possessing specific physicochemical properties, such as low molecular weight (typically <500 Da), optimal lipophilicity (log P in the range of 1.5–3.5), a low melting point (<200 °C), and high potency requiring low daily doses (<10 mg) are capable of effectively permeating the skin barrier.⁴ So, chemical enhancements, such as penetration enhancers, and physical enhancements, such as ultrasound-guided therapy and iontophoresis, are currently used in transdermal drug delivery systems to improve efficacy. Although enhancement methods such as chemical permeation enhancers, iontophoresis, and ultrasound have been explored, inconsistent delivery and skin irritation remain concerns.⁵

To address the aforementioned limitations, microneedle (MN) technology has emerged as an advanced transdermal drug delivery approach that offers distinct clinical advantages over conventional delivery systems.⁶ Microneedles (MNs) are micron-scale needle arrays (typically 100–1500 µm in length) designed to penetrate the stratum corneum without reaching deeper pain receptors. Unlike conventional hypodermic needles, MNs enable minimally invasive, painless, and targeted transdermal drug delivery by creating transient microchannels across the stratum corneum barrier (10–20 µm), thereby enhancing drug permeation, therapeutic efficacy, and overall bioavailability.⁷ MNs are particularly advantageous for self-administration, reducing needle phobia and improving patient adherence.⁸

Microneedles have evolved significantly since their conceptualization, transitioning from solid silicon or metal needles to sophisticated polymeric platforms capable of controlled release and integrated sensing^9,10 shown in Fig. 1.

Fig. 1. Development and applications of polymeric microneedles (PMNs).

Early microneedles were fabricated from silicon and metals; however, concerns regarding brittleness, cost, and biohazardous waste limited their clinical translation. Polymeric microneedles overcome these challenges and enable drug encapsulation within the needle matrix, controlled dissolution, and biodegradation.^11,12 This review critically examines polymeric microneedles for closed-loop drug delivery, focusing on material selection, fabrication strategies, biosensing integration, and therapeutic applications. By synthesizing recent advances and translational challenges, this work aims to guide the development of next-generation smart microneedle systems.

2. Transdermal drug delivery and microneedle technology

2.1. Conventional transdermal drug delivery system

The human skin exhibits a stratified architecture with site-dependent thickness. The stratum corneum (SC), the principal permeability barrier, is ∼10–20 µm thick (up to ∼40 µm in palms per soles). Beneath it, the epidermis measures ∼50–150 µm, followed by the dermis (∼1–2 mm), which contains vascular and neural networks, and the hypodermis, extending from ∼1 mm to several millimetres depending on adipose content. The SC consists of keratinised corneocytes within a lipid matrix (“brick-and-mortar” structure), conferring high impermeability; consequently, only small (<500 Da), moderately lipophilic molecules can effectively permeate intact skin, limiting conventional transdermal drug delivery systems (TDDS) to a narrow drug class.¹³

Although traditional transdermal patches are clinically effective for small lipophilic drugs, they fail for hydrophilic molecules, macromolecules, and biologics. Enhancement strategies such as chemical permeation enhancers, iontophoresis, sonophoresis, and thermal ablation partially address this limitation but introduce drawbacks including variable drug delivery, limited bioavailability, skin irritation, safety concerns, and reduced patient compliance.¹⁴ Microneedles overcome these constraints by creating transient, self-healing microchannels (∼100–800 µm depth) that enable drug transport into the viable epidermis and superficial dermis without reaching deeper regions (∼1–3 mm depth from skin surface) rich in nociceptors and large blood vessels. This controlled penetration ensures minimally invasive, pain-free delivery while maintaining therapeutic efficacy.

Thus, unlike conventional TDDS constrained by SC barrier properties or enhancement techniques requiring external energy and complex control, microneedle systems offer a unique combination of minimal invasiveness, consistent delivery, and broad drug compatibility, illustrated in Table 1. Their capacity to bypass the SC enables efficient administration of macromolecules, peptides, and vaccines, establishing them as a transformative platform in transdermal drug delivery.¹⁵

Table 1. Comparison of microneedles and conventional transdermal drug delivery systems.

Parameter	Conventional transdermal patches	Chemical enhancers	Iontophoresis	Sonophoresis	Thermal ablation	Microneedles (MNs)
Mechanism of delivery	Passive diffusion across stratum corneum	Disrupt lipid structure to enhance permeability	Electrically driven transport of charged molecules	Ultrasound-induced cavitation enhances permeability	Heat-induced microchannel formation	Physical microchannels bypass stratum corneum
Drug type suitability	Small, lipophilic drugs (<500 Da)	Small to moderate molecules	Charged and small molecules	Small to moderately sized molecules	Broad range (including macromolecules)	Broad range including biologics, peptides, vaccines
Control over drug delivery	Limited	Poor	Current-controlled	Moderate	Moderate	High (design-dependent; programmable release possible)
Onset of action	Slow	Moderate	Rapid	Moderate	Rapid	Rapid to controlled
Invasiveness	Non-invasive	Non-invasive	Minimally invasive (electrical stimulation)	Non-invasive	Minimally invasive (thermal damage)	Minimally invasive (painless micro-penetration)
Pain/discomfort	None	Possible irritation	Mild tingling/burning	Mild heating sensation	Possible	Minimal or painless
Risk of skin irritation/damage	Low	High (chemical irritation)	Moderate	Moderate	High	Low
Dose delivery capability	Limited	Limited	Moderate	Moderate	High	Moderate to high
Reproducibility	High	Variable	High	Variable	Moderate	High
Suitability for biologics	Poor	Poor	Limited	Limited	Good	Excellent
Need for external device	No	No	Yes (power source)	Yes (ultrasound device)	Yes (thermal device)	Sometimes (for advanced/closed-loop systems)
Patient compliance	High	Moderate (irritation issues)	Moderate	Moderate	Low to moderate	High
Manufacturing complexity	Low	Low	Moderate	High	High	Moderate to high
Cost	Low	Low	Moderate	High	High	Moderate
Potential for closed-loop systems	Not suitable	Not suitable	Limited	Limited	Limited	Highly suitable (integration with sensors and feedback systems)

2.2. Microneedle system

A microneedle system consists of an array of microscopic projections (100–1500 µm in length) mounted on a supporting backing layer (baseplate/substrate). The microneedles enable controlled skin penetration of 100–800 µm, while the backing layer provides mechanical stability, handling support, and, in some designs, serves as a drug reservoir or interface for loading. Together, they form a microneedle patch for minimally invasive transdermal delivery.

Microneedles are engineered in distinct formats based on delivery mechanism. Dissolving microneedles encapsulate drugs within biodegradable polymers that dissolve after insertion. Coated microneedles carry a surface-applied drug layer for rapid release. Hollow microneedles contain an internal lumen for pressure-driven infusion of liquid formulations. Solid microneedles, in contrast, create microchannels followed by drug diffusion from an external reservoir. Upon application, these systems penetrate the stratum corneum to form transient conduits enabling transport via diffusion, dissolution, coating release, convective flow, or hydrogel swelling. The resulting microchannels reseal quickly after removal, preserving skin integrity and minimizing infection risk.^16,17

The versatility of microneedle architectures from dissolving to hydrogel-forming systems, allows precise control over release kinetics and supports integration with sensing modules for closed-loop, feedback-regulated therapy (Table 2). Despite challenges in large-scale manufacturing, mechanical strength, and cost, microneedles provide a favourable balance of efficacy, safety, and patient compliance, reinforcing their role as next-generation transdermal delivery platforms.

Table 2. Various approaches of microneedle system.

Type	Delivery mechanism	Advantages	Limitations	Closed-loop capability	Ref.
Solid MNs (poke & patch)	Creation of transient microchannels in stratum corneum, followed by topical application or passive diffusion of drug through formed pores	Minimally invasive; simple fabrication; reduced infection risk; reusable	Low delivery efficiency; relies on passive diffusion; no controlled release	Indirect compatibility; can be integrated as a pre-treatment enhancer in feedback-controlled transdermal systems	14, 16 and 19–21
Dissolving MNs (poke & release)	Polymer dissolution after insertion release drug into skin	Higher drug loading; no sharp waste; excellent biocompatibility; patient-friendly	Slower release kinetics, complex fabrication, and limited mechanical strength	Can be integrated with biosensors for feedback-controlled release	14, 16, 22, 25 and 26
Coated MNs (coat & poke)	Rapid dissolution of coating upon insertion	Precise dosing; rapid drug release; minimal systemic exposure	Limited drug loading; coating instability; potential dose loss during insertion	Limited real-time control; potential for triggered release systems	4, 8, 14, 16 and 23
Hydrogel MNs (swell & diffuse)	Swelling forms diffusion channels for sustained delivery	Sustained and controlled release; flexible, can function as biosensors; minimal tissue damage	Limited mechanical robustness; careful polymer design required; slower response	Strong potential for closed-loop diffusion-based control	14, 16, 25 and 26
Hollow MNs (poke & flow)	Pressure-driven infusion of drug into dermis	Enables delivery of large doses and macromolecules; controlled infusion; sampling capability	Complex fabrication; clogging risk; leakage; higher cost	Highly compatible with feedback-controlled infusion systems	14, 16, 24 and 27–29

2.2.1. Closed loop drug delivery

Beyond simple drug delivery, recent advances have demonstrated that microneedle technology can be integrated with biosensing elements to enable real-time monitoring of physiological biomarkers, which in turn facilitate feedback-controlled drug release, shown in Fig. 2. This allows the system to dynamically adjust the therapeutic dosing based on the detected analyte levels, thereby achieving automated, responsive, and patient-specific therapy that closely mimics physiological regulation and minimizes over- or under-dosing. In such systems, microneedles can be engineered to simultaneously sample interstitial fluid for biomarker detection and deliver therapeutic agents in response to real-time physiological signals. This feedback-driven approach enables dynamic modulation of drug release, thereby improving therapeutic efficacy and reducing systemic side effects.¹⁷

Fig. 2. Closed loop drug delivery system using microneedles.

The ability of microneedles to combine minimally invasive delivery, compatibility with a wide range of drug molecules, and potential integration with wearable and smart systems positions them as a critical bridge between conventional transdermal patches and invasive injection-based therapies. Recent works have focused on MN formulations that enable long-acting release, reducing dosing frequency critical for chronic diseases such as diabetes and contraception. Design strategies include biodegradable cores, back-layer reservoirs, and polymers with tunable degradation behaviors that prolong drug residence in the dermis and enable sustained drug release and therapeutic applications,¹⁸ as illustrated in Table 2.

An example of a closed-loop microneedle therapeutic system is a glucose-responsive insulin delivery patch designed for diabetes management. In this system, a dissolving or hydrogel-forming microneedle array is integrated with glucose-sensing elements (e.g., enzymatic sensors based on glucose oxidase). When interstitial glucose levels rise, the enzymatic reaction generates local biochemical triggers (such as pH change or hydrogen peroxide production), which in turn stimulate the controlled release of insulin from the microneedle matrix. This creates an autonomous feedback loop where sensing and drug delivery are coupled in real time, enabling on-demand insulin administration without external intervention. Upon application, the microneedles painlessly penetrate the skin (∼100–800 µm), access interstitial fluid for glucose monitoring, and deliver insulin into the viable epidermis. Such systems have demonstrated rapid response, improved glycaemic control, and reduced risk of hypoglycaemia compared to conventional therapies, highlighting their potential as minimally invasive, self-regulated platforms for chronic disease management.¹⁸

2.2.2. Mechanical basis of microneedle insertion in human skin

Microneedle (MN) performance is fundamentally governed by its ability to overcome the mechanical resistance of human skin, which is dominated by the stratum corneum and varies with anatomical site and physiological condition. The initial barrier requires localized puncture forces typically in the range of ∼0.05–0.5 N per insertion site, with lower values for soft regions (e.g., forearm/abdomen) and higher values (∼0.2–0.5 N) for thicker, keratinized regions such as palm or sole skin. Once this barrier is breached, the viable epidermis and dermis contribute primarily viscoelastic resistance with reduced mechanical opposition. To ensure reliable penetration, microneedles are designed with fracture forces typically >0.5–1 N per needle, providing a 3–10× safety margin over the required insertion force. This ensures structural integrity during insertion, particularly in high-resistance skin sites where puncture thresholds approach the upper limit (∼0.3–0.5 N). Material selection plays a key role in this balance: silicon (Young's modulus ∼130–180 GPa) and metals (∼70–200 GPa) provide high stiffness but are brittle, whereas polymeric microneedles (∼0.5–3 GPa) offer an optimal combination of stiffness and toughness for safe deformation without fracture.¹⁹

Insertion efficiency is also strongly influenced by geometry. Sharp microneedles with tip radii <10 µm significantly reduce insertion force to approximately ∼0.05–0.2 N per needle, while blunt or densely packed arrays increase the cumulative patch insertion force to ∼5–20 N due to collective skin resistance (“bed-of-nails” effect). Successful penetration therefore occurs when fracture and buckling resistance exceed skin puncture forces, while optimized geometry minimizes required insertion load.

Overall, microneedle–skin interaction defines a narrow but critical mechanical window for drug delivery. MNs must be sufficiently rigid to surpass the site-dependent skin puncture threshold (∼0.05–0.5 N) while remaining below their fracture limit (>0.5–1 N per needle). This balance ensures reliable stratum corneum penetration, minimal pain response, and mechanically stable transdermal drug delivery.

3. Polymers in microneedle technology

Among various microneedle materials, polymers have gained significant attention due to their biocompatibility, tunable mechanical properties, degradability, formulation versatility and compatibility with scalable mass-production techniques.^20–22 As summarized in Table 3, the core macromolecular characteristics of polymers used in microneedle fabrication significantly influence their structural and mechanical performance.

Table 3. Core macromolecular characteristics of polymers used in microneedle (MN).

Macromolecular property	General requirement in MN polymers	Representative examples
Biocompatibility & low immunogenicity	Universally required for all MN polymers to ensure safe skin interaction	PVA, PVP, hyaluronic acid (HA), CMC, gelatin, chitosan
Backbone rigidity vs. flexibility	Predominantly flexible or semi-flexible backbones are preferred to prevent brittle fracture during insertion	Flexible: PEG, PVA, semi-flexible: PLGA, PCL
Polymer architecture	Primarily linear or lightly branched structures dominate due to processing advantages	Linear: PVA, PVP, PEG; lightly branched: CMC, modified polysaccharides
Molecular weight range	Typically 10–300 kDa to balance viscosity (processability) and mechanical strength	PVP grades, PVA grades, PLGA
Chain mobility (segmental dynamics)	Moderate-to-high mobility required for deformation without fracture and proper mold filling	PEG, PVA, amorphous PLGA
Crystallinity	Preferably amorphous or semi-crystalline (low crystallinity)	Amorphous: PVP; semi-crystalline: PCL (controlled use)
Crosslinking capability	Required in advanced MN systems (hydrogels) for mechanical tuning and controlled release	GelMA, MeHA, PEGDA
Hydrophilicity/hydrophobicity balance	Tuned depending on drug type and release mechanism	Hydrophilic: HA, PEG; hydrophobic: PLGA
Mechanical integrity versus brittleness	Must achieve high strength without fracture during insertion	PVA, PVP, HA (optimal systems)

The mechanical behavior of MNs varies significantly with the type of materials used for their construction. Silicon MNs exhibit very high stiffness (Young's modulus ∼130–185 GPa; orientation-dependent: ∼169 GPa for 〈100〉, ∼130 GPa for 〈111〉) but are inherently brittle exhibiting low fracture toughness (0.7–1.0 MPa) which raises concerns regarding needle fracture and potential retention in skin, particularly under lateral or bending stresses. Metallic MNs (e.g., stainless steel, Young's modulus ∼190–210 GPa) provide high strength with ductility, reducing abrupt failure; however, their non-biodegradability and generation of sharp biomedical waste limit their applicability in dissolving systems.

Unlike rigid silicon or metallic MNs, polymeric MNs reduce the risk of fracture, enable painless penetration of the stratum corneum, and often dissolve or biodegrade into non-toxic by-products after application. Polymeric MNs such as polyvinyl alcohol (PVA), poly(lactic-co-glycolic acid) (PLGA), and polyvinylpyrrolidone (PVP) exhibit lower stiffness (Young's modulus ∼0.01–4 GPa) but significantly higher toughness (1–5 MPa) and flexibility, allowing them to deform without catastrophic failure.^23,25–30

Beyond their mechanical advantages, polymeric microneedles have evolved into multifunctional platforms for drug delivery, biosensing, and closed-loop therapeutic systems. Interstitial fluid (ISF), which closely reflects blood composition, can be accessed using MNs in a minimally invasive manner, enabling real-time biomarker monitoring with reduced pain and tissue damage compared to conventional methods.³⁰ Recent advances have extended MN-based sensing from glucose to a wide range of analytes, including metabolites, ions, and proteins, while integration with biosensors enables feedback-controlled drug delivery, supporting the development of autonomous and personalized healthcare systems.^31–34

Polymers utilized in MNs must minimize irritation and inflammatory responses. Natural polymers generally offer superior biocompatibility, while synthetic counterparts require careful selection to avoid adverse tissue reactions. U.S.FDA-approved polymers such as PLGA and PVA-PVP have established safety profiles for clinical use in drug delivery systems, biodegradable implants, and transdermal applications including microneedle-based delivery platforms.³⁵ A detailed comparison of polymer characteristics (Table 8) highlights the trade-off between mechanical strength and dissolution rate in MN applications.

Table 8. Comparative mechanical, physicochemical, and application-oriented properties of polymers used in microneedle (mn) fabrication.

Sl no.	Polymer name	Nature of polymers	Modulus of elasticity (GPa)	Tensile strength (MPa)	Elongation at break (%)	Mechanical properties	Suitable for	Fabrication process	Applications	Ref.
1	Hyaluronic acid (HA)	Natural, biodegradable	0.1–0.5	10–30	800–1200	High mechanical strength, high compliance, hydrophilic, excellent biocompatibility, faster dissolving	Dissolving MN	Solvent casting	Delivery of anti-cancer drugs, adenosine, insulin, methotrexate, sustained drug release, treating skin diseases, deeper treatment sites	30, 31, 37, 42, 43, 46, 51 and 52
2	Silk, 2-ethoxyethanol (ECS) modified silk fibroin	Natural, biodegradable	2–17	50–740	4–26	Ability to maintain the activity of biomolecule, biocompatible, non-toxic, swelling capacity	Dissolving/hydrogel forming MN	Micromolding	Substantially enhanced transdermal drug release and vaccination	34, 42, 47 and 53
3	Carboxy methyl cellulose (CMC)	Natural, biodegradable	0.3–1.5	20–50	10–40	Biodegradable, biocompatible,	Dissolving MN	Two step casting	Transdermal insulin delivery in sustained manner	34, 44 and 51
4	Chitosan	Natural, biodegradable	2–7	30–100	10–50	Strong mechanical strength, excellent biocompatibility, antimicrobial properties, high swelling and water retention	Dissolving MN/hydrogel forming MN/coated microneedle	Micromolding/solvent casting	Wound healing, transdermal delivery of macromolecules in sustained manner, transmucosal vaccine delivery	30, 42, 43 and 47
5	PVA	Synthetic, biodegradable	1–2	40–80	50–150	Sufficient mechanical strength and moderate swelling, hydrophilic, dissolvability and nontoxicity, thermal stability	Hydrogel forming MN/dissolving MN	Molding fused deposition method (FDM)	Deliver hydrophilic and hydrophobic polymer, DNA vaccines, nanoparticles, reduced drug loss	43, 46, 50, 53 and 54
6	PLA	Synthetic, biodegradable	3–4	40–60	4–10	Biodegradability, biocompatible, hydrophobic, good mechanical strength, higher rates of tensile strength	Coated MN	Molding	Insulin-coated microneedles exhibited effective glycemic control	34, 38, 47 and 48
7	PLA/PLGA, PLA/PEG (or TWEEN)	Synthetic, biodegradable	0.5–1.5	30–60	100–300	Biocompatible, gradual degradation, ability to safely break down into nontoxic products, increased drug loading, easy fabrication and operation, low cost	Porous, solid MN/dissolving MN	Molding and solvent casting	Dermal delivery of HbSAg to improve immunogenicity, controlled release of encapsulated drug (like estradiol valerate), long-acting sustained release, avoiding frequent dosing	31, 34, 47, 66, 74, 75 and 84
8	PVP/PVA/hydrogels	Synthetic, biodegradable	0.01–0.5	5–50	100–500	Satisfactory mechanical strength and insertion capabilities, high dissolving, good biocompatible	Coated rapidly dissolving MN	Molding and photo polymerization	Controlled release, high drug loading, effective penetration, diabetes treatment, cancer treatment, psoriasis management	30, 35, 37, 43, 45, 47, 51, 53, 54, 56 and 60
9	PEGDA	Synthetic, biodegradable	0.01–0.1	1–10	10–200	Hydrophilic, swellable, tunable mechanical strength, easy penetration	Hydrogel forming MN	Photolithography	Deliver hydrophilic drugs	58 and 61
10	PGMA	Synthetic, biodegradable	1–3	30–70	5–50	Hydrophilic properties, low toxicity, good penetration efficiency	Porous MN, hydrogel-forming MN	Photopolymerization	Useful for solubilizing and delivering hydrophobic drugs. Improved therapeutic efficacy and reduced side effects	30 and 54
11	PCL, Fe3O4/PCL MNs	Synthetic, non-biodegradable	0.2–0.4	10–30	300–500	Hydrophobic, low melting point, and exceptional blend-compatibility, solubility in organic solvents, good thermal stability, high rate of permeability	Solid MN	3D printing, micro-molding	Thermally modulated release of the encapsulated metformin, skin cancer treatment with transdermal controlled drug delivery	30, 54, 57, 84 and 93
12	PMVE/MA	Synthetic, non biodegradable	1–2	30–60	2–10	Biocompatible, water-soluble or swellable. The microneedles remained intact after drug delivery	Back-layer reservoir polymeric microneedles	3D printing	Used to deliver metformin hydrochloride in a sustained manner	44, 49, 53 and 54

3.1. Types of polymers used in microneedles

3.1.1. Natural polymers

Natural polymers such as chitosan and hyaluronic acid (HA) are widely utilized in microneedle (MN) fabrication due to their biocompatibility, biodegradability, and inherent biofunctionality. Chitosan, a cationic polysaccharide derived from chitin, typically exhibits a degree of deacetylation of 70–95% and a molecular weight ranging from ∼50 to 1000 kDa. It offers intrinsic antimicrobial activity, mucoadhesiveness, and hemostatic properties, making it particularly suitable for minimally invasive transdermal applications. However, its relatively high crystallinity and strong intermolecular hydrogen bonding can limit solubility at physiological pH, often necessitating acidic solvents such as acetic acid, which may constrain formulation flexibility and mechanical tuning in MN systems.^36,42

In contrast, HA is a linear anionic glycosaminoglycan with molecular weights spanning from ∼50 kDa to several MDa, characterised by excellent water solubility, high viscoelasticity, and strong skin affinity through CD44 receptor interactions. HA-based MNs rapidly swell (1–3× volume increase) upon insertion, enhancing analyte diffusion and drug release, although their high hydrophilicity can lead to premature dissolution within 5 minutes under humid conditions and comparatively lower mechanical strength, often requiring crosslinking or blending with polymers such as PVA or PVP to improve mechanical strength (>0.5 N per needle), reduce dissolution rate, and stabilize performance during storage and applications.

From a mechanistic perspective, the biodegradability of these polymers eliminates the need for post-use needle retrieval, reducing biohazardous waste and improving patient compliance. Both chitosan and HA degrade via enzymatic or hydrolytic pathways into non-toxic byproducts (e.g., glucosamine and HA disaccharides), which are readily metabolized in physiological systems. This property is particularly advantageous for single-use MN patches in point-of-care diagnostics and controlled drug delivery, enabling safer application and sustained or programmable release profiles. While HA provides superior biocompatibility and skin-mimetic behaviour, chitosan contributes structural integrity and antimicrobial functionality. Consequently, polymer selection is highly application-specific, and hybrid systems are often employed to optimize mechanical strength, degradation kinetics, and drug release performance.³⁶

3.1.2. Synthetic polymers

Synthetic polymers are widely preferred in microneedle (MN) fabrication due to their superior batch-to-batch reproducibility, tunable mechanical properties, and precisely controllable degradation kinetics, which are critical for ensuring consistent insertion depth, mechanical failure resistance, and predictable drug release profiles.^36,45 Unlike natural polymers, which often exhibit variability in molecular weight distribution and impurity content, synthetic polymers offer standardized chemical architecture and scalable manufacturing, making them more suitable for translational and industrial MN production.

A key example is poly(lactic-co-glycolic acid) (PLGA). PLGA whose degradation time can be precisely tuned from days to several months by adjusting the lactic acid: glycolic acid ratio (e.g., 50 : 50 for faster degradation; 75 : 25 for slower degradation). This level of tunability is difficult to achieve with natural polymers such as chitosan or hyaluronic acid, whose degradation rates depend largely on enzymatic activity and physiological variability.³⁷

In MN systems, PLGA-based structures maintain mechanical integrity during insertion while enabling sustained and programmable release, making them particularly suitable for long-acting drug delivery and implantable MN platforms. Similarly, polyvinyl alcohol (PVA) forms mechanically robust, water-processable films with high tensile strength (40–80 MPa) and swelling capability 1–5× weight/volume increase (W/V). PVA-based dissolving MNs have demonstrated efficient delivery of vaccines and macromolecules with high reproducibility due to their uniform crystallinity and controlled freeze–thaw crosslinking.^38,45

Polyvinylpyrrolidone (PVP) is another widely used synthetic polymer for dissolving MNs due to its excellent film-forming ability and rapid aqueous dissolution, enabling complete drug release within minutes to hours after skin insertion. However, unlike fully metabolizable natural polymers such as HA, PVP degradation pathways in vivo are less biologically defined, and residual polymer fragments may persist temporarily, necessitating further biocompatibility and clearance studies.^46,47

From a structural perspective, synthetic polymers provide design versatility across linear (PVA, PVP), biodegradable polyester (PLGA), and crosslinked hydrogel systems, enabling precise engineering of MN behaviour. Linear polymers facilitate rapid dissolution, branched architectures enhance drug loading capacity, and crosslinked networks enable hydrogel-forming MNs for sustained diffusion-based delivery. A detailed comparison of natural and synthetic polymers is presented in Table 4.

Table 4. Comparison of synthetic and natural polymers in microneedle fabrication.

Feature	Synthetic polymers	Natural polymers
Common examples	PVA, PVP, PLGA, PLA, PGA, PEG-based hydrogels	Chitosan, hyaluronic acid (HA), gelatin, alginate
Mechanical strength	High (e.g., PVA, PLGA provide strong needle integrity and reliable skin insertion)	Moderate to low (HA and chitosan often require reinforcement or blending)
Reproducibility/batch consistency	Excellent due to controlled synthesis and defined molecular architecture	Variable due to biological origin, molecular weight fluctuations, and impurity content
Degradation behavior	Precisely tunable (e.g., PLGA degradation controlled by lactic: glycolic ratio from days to months)	Enzyme- or hydrolysis-dependent, less predictable in vivo
Drug release control	Highly programmable (sustained, delayed, or rapid release depending on polymer design)	Typically, diffusion- or enzyme-controlled, less tunable
Biocompatibility	Generally high, but depends on degradation products (e.g., PLGA → lactic & glycolic acid)	Excellent intrinsic biocompatibility and bioactivity
Bioactivity	Limited intrinsic bioactivity unless functionalized	High (e.g., HA interacts with CD44 receptors; chitosan has antimicrobial activity)
Dissolution behavior	Fast and controllable (PVP dissolves within minutes to hours; PVA tunable via crosslinking)	HA/chitosan dissolve or degrade biologically, often slower and enzyme-dependent
Manufacturing scalability	Highly scalable, industrially standardized production	Limited by extraction/purification variability and cost of biological sourcing
Example advantage	PLGA enables precise long-acting drug release (weeks–months) not achievable with natural polymers	HA provides excellent skin affinity and wound healing properties
Key limitation	Possible lack of intrinsic bioactivity; some residues (e.g., PVP) require clearance studies	Mechanical weakness and variability under physiological conditions
Overall role in MNs	Structural backbone, controlled release systems, long-acting delivery platforms	Bioactive, biocompatible matrices for skin-interactive and regenerative applications

3.1.3. Hydrogel-forming and composite synthetic polymers

Hydrogel-forming synthetic systems, such as acrylate-modified PVA or PEG-based networks, swell upon insertion to establish controlled aqueous diffusion pathways between drug reservoirs and dermal tissue, enabling sustained delivery or analyte extraction. For example, a PEG-based hydrogel MN patch coupled with an external insulin reservoir can enable sustained insulin delivery over ∼24 hours, while simultaneously allowing interstitial fluid extraction for glucose monitoring. Similarly, composite MN systems enhance performance by integrating functional nanomaterials.

For instance, PVA-based MNs embedded with gold nanoparticles improve electrical conductivity for biosensing, silica nanoparticles enhance mechanical strength, and zinc oxide nanoparticles provide antimicrobial activity. Lipid-based carriers (e.g., liposomes) loaded within polymeric MNs can further improve drug stability and enable controlled or targeted release, demonstrating multifunctional therapeutic and diagnostic capabilities.^39,48,110

3.1.4. Biodegradable and non-biodegradable polymer

Biodegradable and non-biodegradable polymers play distinct roles in microneedle (MN) systems. Biodegradable polymers such as PLGA, PLA, PCL, hyaluronic acid (HA), chitosan, PVP, and PVA are widely used to fabricate dissolving or sustained-release MNs, as they degrade in vivo into non-toxic byproducts, eliminating the need for removal. For example, PLGA-based MNs have been used for long-acting vaccine delivery (release over days–weeks), while HA-based MNs are commonly employed in cosmetic and dermatological applications for rapid drug release.^47,48 Chitosan MNs have demonstrated antimicrobial wound healing, and PVP/PVA-based dissolving MNs have been successfully used for insulin and vaccine delivery, dissolving within minutes to hours.⁴⁹

In contrast, non-biodegradable polymers such as polymethyl methacrylate (PMMA), polycarbonate (PC), cyclic olefin copolymers (COC), and polydimethylsiloxane (PDMS) are primarily used for fabrication molds, solid MN arrays, or reusable platforms due to their high mechanical strength and stability. For instance, PDMS is widely used as a mold material in micromolding processes, while PMMA and COC have been utilized in solid MNs for “poke-and-patch” applications. These materials do not degrade in vivo and are therefore less suitable for dissolving systems but are valuable for structural support and repeated-use devices.^50,51

3.2. Polymeric microneedles (MNS)

Polymeric microneedles (MNs) are minimally invasive transdermal systems that combine the bio-functionality of natural polymers with the mechanical tunability of synthetic polymers. Natural polymers such as chitosan and hyaluronic acid offer superior biocompatibility, biodegradability, and skin affinity, whereas synthetic polymers like PVA, PVP, and PLGA provide enhanced mechanical strength, stability, and controlled fabrication. This contrast enables tailored MN designs that balance insertion efficiency with drug release performance. Consequently, hybrid polymer systems are often employed to integrate bioactivity with structural robustness for advanced therapeutic and diagnostic applications.⁵¹

3.2.1. Factors influencing the performance of polymeric microneedles

The performance of polymeric microneedles is governed by several interrelated factors that determine insertion efficiency, drug delivery, and mechanical stability, illustrated in Table 5. Mechanical properties of polymer, such as Young's modulus, fracture force, and elasticity and physico-chemical characteristics of polymers, such as brittleness, flexibility and degradability, ensure effective MN insertion, integrity and penetration of the skin without bending or breaking and toxicity. Synthetic polymers like PLGA and PVA–PVP blends provide adequate stiffness and toughness for reliable skin penetration, whereas natural polymers often require blending to avoid buckling or fracture during insertion.^45–47 Hydrogel networks balance flexibility with sufficient rigidity to maintain structural integrity while swelling and show the greatest ability for long-term and multiplexed biosensing of biological health parameters. The polymer's hydrophilicity/hydrophobicity, crosslink density, and mesh size govern drug encapsulation efficiency and release profiles.⁴⁸

Table 5. Influence of structural and mechanical parameters of microneedle on its functional performance.

Parameter	Typical range	Effect on performance	Optimization strategy	Key ref.
Array density (number of microneedles per unit area)	100–10 000 needles per cm2	Higher density increases drug flux and sensing signal, but excessive density causes the “bed-of-nails” effect, reducing penetration efficiency	Use moderate density (500–3000 needles cm−2) to ensure effective penetration and delivery	1, 2, 6, 29 and 41
Base geometry (shape of MN base)	Circular and polygonal	Influences stress distribution and fracture resistance; circular base provides uniform stress distribution	Circular or hexagonal bases provide optimal mechanical stability	6, 8, 40 and 41
Microneedle shape (structural profile)	Conical, pyramidal, cylindrical	Conical needles require lower insertion force; pyramidal needles provide higher mechanical strength	Conical or pyramidal shapes offer optimal performance	2, 8, 40 and 41
Height (length) (distance from tip to base)	150–1000 µm	Determines penetration depth and access to interstitial fluid; longer needles improve delivery but increase pain risk	Optimal height: 300–800 µm	8, 14 and 39
Base width (diameter at base)	50–300 µm	Larger base improves fracture resistance but increases insertion force	Balance width for strength and penetration	4, 6, 8, 39 and 40
Tip diameter (diameter of needle tip)	1–25 µm	Smaller tip diameter reduces insertion force and improves penetration efficiency	Optimal tip diameter: 5–15 µm	2, 3, 6, 24 and 39
Aspect ratio (height : base width)	2 : 1 to 10 : 1	High aspect ratio improves penetration but increases risk of buckling	Optimal aspect ratio: 2 : 1 to 5 : 1	8, 9, 24 and 40
Insertion force	0.1–1 N per needle	Determines successful penetration and mechanical reliability	Minimize insertion force through sharp tips and optimal geometry	1, 3, 8, 9 and 24

The performance of polymeric microneedle (MN) systems with respect to skin penetration, drug delivery efficiency, release kinetics, and biocompatibility is predominantly dictated by the structural design of the MN array, shown in Table 5. Tip sharpness, needle length, and array density determine insertion efficiency and consistent penetration of the stratum corneum into viable skin layers, while polymer composition, dissolution behavior, and coating thickness (1–20 µm) regulate drug loading and release kinetics, thereby controlling dose and duration.³⁹ Sharper and slender microneedles generally require lower insertion forces and enable more reproducible skin penetration, although such geometries may compromise mechanical strength. In contrast, microneedles with wider bases or conical and pyramidal configurations exhibit improved fracture resistance but typically demand higher insertion forces.^40,41

Array density and inter-needle spacing influence the total drug loading capacity, spatial uniformity of drug delivery, mechanical stress distribution within the skin, and user comfort. While increasing needle density can enhance overall drug dose, excessively dense arrays may elevate insertion force requirements and promote skin deformation rather than effective penetration.^29,41

Tip architecture and sharpness, such as bevelled, pyramidal, hollow, or solid designs, further modulate insertion efficiency, penetration depth, and tissue interaction. Hollow microneedles permit liquid drug infusion, while solid and dissolvable microneedles rely on diffusion or polymer dissolution mechanisms, which may constrain drug loading capacity and release kinetics.

Drug incorporation strategies, including matrix encapsulation, surface coating, core–shell architectures, and multilayer configurations, significantly affect drug loading efficiency, release profiles, and stability, particularly for protein- and peptide-based therapeutics. Drug-loaded dissolvable matrices generally provide sustained release characteristics, whereas surface-coated designs enable rapid release but are limited in drug payload.⁴² Optimizing the design of polymeric MN therefore involves balancing skin penetration and structural integrity with efficient drug release, safety, and patient comfort.

The key factors influencing the fabrication of polymeric microneedles are summarized in Table 6. Material choice critically affects mechanical integrity, drug compatibility, and functional performance of polymeric micro needle.⁴³

Table 6. Factors influencing the fabrication of polymeric microneedle.

Fabrication factor	Key practical parameters	Microneedle characteristics influenced	Fabrication and translational considerations	Key ref.
Polymer selection	Biocompatibility, hydrophilicity/hydrophobicity, polymer class	Drug compatibility, biodegradability, release profile	Selection depends on therapeutic need (rapid vs. sustained release)	8, 9, 11, 13 and 20
Polymer concentration	Polymer wt%, viscosity, rheological behavior	Tip sharpness, structural integrity, fracture resistance	High concentration enhances strength but hinders cavity filling; low concentration risks tip collapse	8, 9, 11, 17, 20, 21, 25, 31, 38, 41, 43 and 46
Polymer molecular weight (MW)	PVP: ∼10–360 kDa; PVA: ∼30–150 kDa; PLGA: ∼10–100 kDa	Mechanical strength, viscosity, mold filling, structural integrity	Higher MW improves strength but increases viscosity, limiting mold filling and tip sharpness	25, 29–31 and 35
Glass transition temperature (Tg)	PVP: ∼150–180 °C; PVA: ∼85 °C; PLGA: ∼40–60 °C	Rigidity, thermal stability, deformation behavior	Higher Tg increases stiffness and dimensional stability but reduces flexibility and processability	42, 44 and 45
Crystallinity	PVA: semi-crystalline (∼30–60%); PLGA: amorphous; PVP: amorphous	Mechanical strength, dissolution rate, drug diffusion	Higher crystallinity enhances stiffness but slows dissolution and may increase brittleness	31 and 47
Mechanical properties (material level)	Young's modulus: ∼0.01–4 GPa; fracture force: ∼0.1–1 N per needle	Skin penetration, fracture resistance, insertion efficiency	Lower modulus polymers require optimized geometry; higher toughness reduces catastrophic failure risk	1, 2, 11 and 24
Polymer architecture	Linear, lightly branched (dominant); crosslinked (hydrogels)	Moldability, tip formation, swelling behavior	Linear polymers enable uniform mold filling; crosslinked systems enable controlled release	13, 20 and 25
Polymer concentration/rheology	wt%, viscosity, shear-thinning behavior	Tip sharpness, structural uniformity, mechanical stability	High viscosity limits cavity filling; low viscosity leads to weak or collapsed needles	8, 11, 17 and 20
Hydrophilicity/hydrophobicity	Hydrophilic (PVA, PVP); hydrophobic (PLGA)	Drug loading efficiency, release kinetics	Hydrophilic polymers dissolve rapidly; hydrophobic systems enable sustained release	9, 11 and 20
Crosslinking capability	Crosslink density, network formation (e.g., GelMA, PEGDA)	Swelling behavior, mechanical strength, release control	Increased crosslinking improves strength but reduces swelling and drug diffusion	13, 25 and 31
Polymer degradation behavior	Degradation rate (days–weeks), hydrolytic stability	Sustained release profile, biocompatibility	Faster degradation → rapid release; slower degradation → long-term delivery	20, 30 and 35
Solvent system	Solvent type, evaporation rate, polymer solubility	Surface morphology, internal porosity, drug stability	Aqueous solvents preferred; rapid evaporation induces voids and non-uniform tips	1, 11, 24, 26, 27, 31 and 35
Mold material and geometry	Mold type, cavity depth, tip radius, aspect ratio	Needle height, tip radius, penetration capability	Sharp cavities are essential for skin insertion; PDMS molds ease demolding but may swell	7, 9, 19–21, 24, 25, 30, 36 and 41
Filling technique	Casting method, centrifugal force, vacuum pressure, filling time	Tip fidelity, drug distribution uniformity	Centrifugation improves tip filling; vacuum minimizes trapped air	10, 11 and 22
Drying and curing conditions	Temperature, time, humidity	Mechanical stability, dimensional accuracy, residual stress	Rapid drying causes cracking; controlled drying improves reproducibility	1, 6 and 11
Drug incorporation	Drug loading, drug–polymer compatibility, distribution	Drug release profile, MN mechanical strength	High drug loading weakens MNs; phase separation causes non-uniform release	14, 15, 18, 23, 26, 27, 35, 38 and 41
Mechanical properties	Fracture force, Young's modulus, geometry	Skin insertion efficiency, failure resistance	MNs must exceed minimum insertion force without fracture	1, 2, 11, 16, 24, 28, 29, 31, 36, 40 and 41
Reproducibility and scalability	Batch variation, fabrication yield, automation	Dimensional consistency, dose uniformity	Manual fabrication increases variability; scalable methods favor translation	11, 13, 34 and 57
Sterilization compatibility	Sterilization method, polymer stability	Mechanical integrity, drug activity, and geometry retention	Gamma irradiation may degrade polymers; EO requires degassing	11, 32, 41 and 43

3.3. Classification of polymeric microneedles

Compared with metallic or silicon microneedles, polymeric systems offer superior safety, tunable degradation, and enhanced compatibility with biological payloads such as proteins, vaccines, and enzymes. Polymeric microneedles (PMNs) are minimally invasive microstructures—typically 300–900 µm in length with tip diameters <10–50 µm fabricated from biocompatible or biodegradable polymers, enabling effective penetration of the ∼10–20 µm thick stratum corneum without reaching deeper pain receptors. These systems can be engineered with controlled molecular weights (∼10–1000 kDa) and degradation times ranging from minutes (dissolving MNs) to hours or days (hydrogel-forming systems), allowing precise tuning of drug release kinetics. Additionally, polymeric MNs exhibit sufficient mechanical strength (fracture forces typically ∼0.1–1 N per needle) for reliable skin insertion while maintaining flexibility (0.5–3 GPa) and safety compared to brittle silicon or rigid metal counterparts.

Based on structural design and polymer behavior, polymeric MNs are commonly classified into solid, coated, dissolving, hydrogel-forming, and hollow microneedles, summarized in Table 7 and illustrated in Fig. 3. The selection of a polymeric microneedle (MN) type depends primarily on key factors such as intended application, drug properties, release profile, and mechanical and safety requirements.

Table 7. Types and applications of polymeric microneedles.

Types of microneedles	Polymers used	Applications	Release duration	Ref.
Dissolving	PVP, CMC, PVA, PLGA, HA, chitosan, dextran	Vaccines, insulin, local anaesthesia	Minutes to several hours	12, 22, 26, 35, 37, 38, 45, 50 and 51
Coated	PLGA, carboxymethyl chitosan, PLA/PVA (coating)	DNA vaccination, ocular drug delivery	Minutes to hours	8, 44, 47, 53 and 54
Biodegradable	PLA, PLGA, PCL, GelMA	Long-acting drug delivery, implants, and cancer therapy	Days to weeks	18, 30, 31, 34, 42, 46 and 54
Hollow	Polycarbonate, PMMA, SU-8, PDMS	Interstitial fluid (ISF) extraction, large volume drug infusion, and electrochemical monitoring	Minutes to hours	21, 24, 28, 29, 32, 33, 49 and 55
Stimuli-responsive	Hypoxia-sensitive polymers, glucose-responsive gels	Closed-loop diabetes therapy (smart patches)	Triggered swelling/degradation-on demand	43, 48, 52 and 56
Hydrogel-forming	PVP, PMVE/MA, PEGDA, PEG, PVA, poloxamers, silk fibroin	Continuous monitoring, controlled/sustained release	Hours to several days(weeks)	48, 57, 58 and 59

Fig. 3. Representative types of MNs for transdermal drug delivery. (A) Solid microneedles (MNs) create temporary microchannels in the skin to enhance permeability and facilitate subsequent application of a drug formulation. (B) Coated microneedles deliver a drug through a thin surface coating that rapidly dissolves upon insertion into the skin. (C) Dissolving microneedles encapsulate the drug within a biodegradable matrix that dissolves after insertion, enabling either immediate or controlled release. (D) Hollow microneedles Penetrate the skin to allow administration of liquid formulations through their internal lumen via infusion or passive diffusion. (E) Hydrogel-forming microneedles absorb interstitial fluid upon insertion, swell within the tissue, and permit diffusion of drug. Reproduced from ref. 41 under terms of the CC-BY 4.0 license.

3.3.1. Dissolving polymeric microneedles

Dissolving microneedles are fabricated from water-soluble polymers such as PVP (M_w ∼10–360 kDa), PVA (M_w ∼30–150 kDa), hyaluronic acid (M_w ∼50 kDa–2 MDa), and carboxymethyl cellulose (90–700 kDa), which encapsulate drugs within the needle matrix and fully dissolve in interstitial fluid, eliminating biohazardous sharps waste.⁵⁰ Typically measuring 300–800 µm in height with tip diameters <10–30 µm, these MNs penetrate the ∼10–20 µm thick stratum corneum and dissolve within minutes (≈5–30 min), releasing therapeutics into the viable epidermis or upper dermis.³⁹

Drug release kinetics can be tuned from rapid “burst” to sustained profiles (up to several hours) by varying polymer concentration (5–40%), molecular weight (100–700 kDa), or incorporating crosslinkers. For example, insulin-loaded PVA–sucrose MNs demonstrate adequate mechanical strength (∼0.2–0.5 N per needle) and efficient transdermal delivery, while gelatin/CMC MN patches provide sustained drug release (4–12 hours) thereby achieve prolonged glucose reduction in diabetic models. Due to their high loading efficiency (often >80–90%) and preservation of biomolecule activity, dissolving MNs are particularly suitable for vaccines and macromolecular therapeutics, supporting controlled and minimally invasive delivery.^12,50,51

3.3.2. Nano-/microparticle-loaded dissolving microneedles

In this advanced subclass, therapeutic agents are first encapsulated within polymeric nano/microparticles (typically ∼50–500 nm for nanoparticles and ∼1–50 µm for microparticles) and then incorporated into dissolving microneedles. Upon insertion, the microneedle matrix dissolves within ∼5–30 minutes, releasing these particulate carriers into the viable epidermis or dermis, where they provide sustained drug release over extended durations (hours to days). This system improves drug stability (encapsulation efficiencies often >70–90%), enables controlled or long-acting delivery, and supports stimuli-responsive or targeted release mechanisms (e.g., pH- or enzyme-triggered degradation).^47,48,52 Such designs are particularly advantageous for fragile biomolecules and chronic therapies, where prolonged therapeutic levels and reduced dosing frequency are required.

3.3.3. Coated microneedles

Solid MNs coated with drug molecules applied for instant delivery, also called the coat and poke method. Coated MNs (typically 300–700 µm in length) carry thin layer of drugs (∼1–20 µm coating thickness) on their surface, allowing immediate dissolution in interstitial fluid within seconds to a few minutes. Drug coatings composed of hydrophilic polymers (e.g., PVP, CMC) dissolve rapidly upon insertion, enabling rapid bolus delivery. However, the drug loading capacity is inherently limited (typically ∼1–100 µg per patch) due to surface area constraints, which restricts their use to potent drugs or vaccines. Despite this limitation, coated MNs are highly effective for bolus delivery applications requiring precise dosing and rapid pharmacological response.^47,53,54

3.3.4. Biodegradable (solid) polymeric microneedles

Solid microneedles (MNs) create transient microchannels (∼10–50 µm diameter) via the “poke-and-patch” approach, enhancing skin permeability., polymeric solid MNs with length typically 300–800 µm, fabricated from PMMA or cyclic olefin polymers exhibit adequate mechanical strength (∼0.5–2 N per needle) with reduced brittleness compared to silicon or metals, improving safety and handling.19,36,54 These microneedles are composed of polymers such as PLA, PLGA, and PCL (M_w ∼10–200 kDa), retain structural integrity after insertion and enable sustained drug release through matrix degradation over extended periods (days to weeks). Unlike dissolving MNs, these systems function as implant-like depots, supporting long-acting transdermal delivery with controlled release kinetics.^25,26,36,54

3.3.5. Hollow polymeric microneedles

Hollow polymeric microneedles (MNs) contain internal microchannels (∼10–100 µm lumen diameter) that enable active or passive transport of liquids for drug infusion or interstitial fluid extraction. They function as miniaturized hypodermic needles of typically 500–1000 µm in length, supporting controlled delivery rates (∼1–100 µL min⁻¹) under passive diffusion or applied pressure. Compared with silicon, polymeric hollow MNs offer improved mechanical compliance (fracture forces ∼0.2–1 N per needle) and reduced brittleness, enhancing safety and skin conformity. Advances in microfabrication and additive manufacturing have further enabled precise control over geometry and flow characteristics, making them highly suitable for continuous drug delivery and biosensing applications.⁵⁵

3.3.6. Stimuli-responsive polymeric microneedles

Stimuli-responsive microneedles are engineered to respond to physiological or external triggers such as pH, glucose concentration, temperature, light, or enzymatic activity. These “smart” systems enable self-regulated or closed-loop drug delivery, exemplified by glucose-responsive insulin microneedle patches for sustained drug release and diabetes management.^{16,33,43,45,48,52,56}

3.3.7. Hydrogel-forming microneedle

Biodegradable microneedles (MNs) enable sustained drug release via polymer erosion (typically over days to weeks), whereas hydrogel-forming MNs (length ∼500–800 µm) swell upon insertion (swelling ratio ∼2–10×), creating hydrated microchannels for controlled diffusion between a drug reservoir and the skin. Unlike dissolving MNs, hydrogel-forming systems remain structurally intact, supporting sustained delivery from hours to days depending on crosslink density and reservoir design, while also permitting continuous interstitial fluid extraction (∼0.1–5 µL) for biosensing. Their tunable swelling behavior, adequate mechanical strength (∼0.2–1 N per needle), and excellent biocompatibility make them highly promising for long-term therapy and real-time physiological monitoring in personalized healthcare.^57–59

4. Fabrication techniques for polymeric microneedles

The fabrication of polymeric microneedles (MNs) is a critical step that determines their mechanical performance, drug-loading capacity, and clinical applicability. A variety of techniques have been developed to produce MNs with precise geometries, high reproducibility, and scalability, taking into account polymer properties such as solubility, viscosity, and crosslinking behaviour^{49,50,53,60,61} shown in Fig. 4 and 5.

Fig. 4. Fabrication technique of polymeric microneedle: micromolding. Reproduced from ref. 10 under terms of the CC-BY 4.0 license.

Fig. 5. Fabrication technique of polymeric microneedle: 3D printing of microneedle mold using SLA printer. Reproduced from ref. 58 under terms of the CC-BY 4.0 license.

4.1. Micromolding

Micromolding, illustrated in Fig. 4 remains the most widely adopted fabrication technique for polymeric microneedles (MNs) due to its scalability, reproducibility, and compatibility with a broad range of polymer systems. The process typically involves casting polymer solutions or melts into microfabricated molds (e.g., silicone, PDMS, or metal), followed by vacuum or centrifugal filling to ensure complete cavity occupation, and subsequent drying or crosslinking to form solid MN arrays. This method is particularly suitable for the fabrication of dissolving, biodegradable, and hydrogel-forming MNs, enabling high-throughput production and precise control over needle geometry.^{48–50,58–60}

A representative example from the literature is the fabrication of hydrogel-forming MNs using crosslinked polymer networks (e.g., poly(methyl vinyl ether-co-maleic acid) with polyethylene glycol), where micromolding combined with vacuum-assisted filling produced mechanically robust MN arrays with high insertion efficiency and consistent swelling behavior. These MNs demonstrated efficient interstitial fluid uptake and sustained drug diffusion when coupled with an external reservoir, highlighting the suitability of micromolding for complex, multifunctional systems. The results illustrate key advantages of micromolding, including excellent shape fidelity, uniform drug distribution, and scalability. However, limitations such as air entrapment, incomplete tip filling in high aspect ratio molds, and sensitivity to polymer viscosity can affect needle sharpness and mechanical performance.

Centrifugal filling is widely employed to enhance mold filling by driving polymer–drug solutions into microcavities, resulting in improved uniformity and higher drug loading efficiency.^49,50 Nevertheless, this approach may still suffer from air bubble entrapment, particularly in viscous formulations. Vacuum-assisted filling addresses this limitation by removing trapped air prior to or during casting, thereby improving tip sharpness, structural integrity, and reproducibility of MN arrays.^58–60 Importantly, the physicochemical properties of polymers such as molecular weight, viscosity, crosslinking behavior, and mechanical strength directly influence micromolding efficiency. For instance, highly viscous or rapidly crosslinking polymers may hinder complete mold filling, whereas polymers with optimal flow and curing characteristics enable better replication fidelity and mechanical robustness.

4.2. Additive manufacturing

Stereolithography (SLA) and two-photon polymerization (2PP) have emerged as advanced additive manufacturing techniques for polymeric microneedle (MN) fabrication, offering exceptional precision, reproducibility, and architectural complexity. These layer-by-layer photopolymerization methods enable fabrication with resolutions in the range of ∼25–100 µm for SLA and submicron (<1 µm) for 2PP, allowing precise control over critical geometrical parameters such as needle height (300–1000 µm), tip radius (<10–20 µm), and aspect ratio. Such control is essential for optimizing skin penetration efficiency and mechanical stability.^55,58,59

SLA, shown in Fig. 5 is particularly effective for fabricating MN arrays using photocurable polymers such as polyethylene glycol diacrylate (PEGDA) and gelatin methacryloyl (GelMA). These materials offer tunable crosslinking density, enabling modulation of mechanical strength (typically >0.1–0.5 N per needle, sufficient for skin insertion) and swelling behavior for controlled drug release. SLA-fabricated PEGDA MNs exhibit smooth surface morphology, uniform geometry, and the ability to incorporate internal microchannels for drug delivery or fluid extraction. SLA enables high-resolution structures with feature sizes of ∼10–50 µm and geometric fidelity >90–95%, producing microneedles with tunable fracture strength (∼0.3–1.5 N per needle) depending on crosslink density. GelMA-based MNs additionally provide enhanced biocompatibility and biofunctionality, making them suitable for biosensing and tissue-interfacing applications.^28,58,61

Compared to micromolding of PEGDA-based MNs, additive manufacturing offers superior design flexibility and geometric precision, particularly for complex structures such as hollow or porous MNs. Micromolding, while advantageous for scalability and high-throughput production, may suffer from limitations such as incomplete mold filling, variability in tip sharpness, and restricted design adaptability. In contrast, SLA enables superior precision, rapid prototyping, better control over structural design and customization without the need for molds, although it is generally associated with slower fabrication speeds and higher costs.

Two-photon polymerization (2PP) further extends these capabilities by enabling submicron-resolution fabrication of hollow, porous, and biomimetic MNs with interconnected lumens, supporting applications in interstitial fluid extraction, biosensing, and theranostics.^55,62 Continuous Liquid Interface Production (CLIP) addresses some limitations of conventional SLA by enabling rapid, layer less fabrication through oxygen-inhibited photopolymerization. This approach produces MNs with improved surface smoothness, structural fidelity, and significantly faster production rates compared to traditional SLA, while maintaining precise control over geometry and mechanical properties.^38,47,55,62

4.3. Drawing lithography

Drawing lithography illustrated in Fig. 6 utilizes the controlled elongation of polymer droplets under thermal or mechanical forces to fabricate microneedles (MNs) with high aspect ratios and sharp tip geometries. This technique is particularly suitable for thermoplastic and hydrogel-forming polymers, enabling precise control over MN dimensions, including needle height (∼300–1500 µm), base diameter (∼50–300 µm), and tip radius (<5–20 µm), which are critical for efficient skin penetration. The resulting MNs typically exhibit high aspect ratios (>5–10), enhancing insertion efficiency while maintaining structural integrity. Mechanical strength can be tuned through polymer concentration and crosslinking density with insertion forces commonly exceeding ∼0.1 N per needle, sufficient to penetrate the stratum corneum without fracture.

Fig. 6. Fabrication technique of polymeric microneedle: drawing lithography. Reproduced from ref. 10 and 68 under terms of the CC-BY 4.0 license.

Drawing lithography is especially advantageous for hydrogel-based MNs used in the extraction of interstitial fluid (ISF) as a key sampling medium in microneedle-based biosensing, as it contains physiologically relevant analytes (e.g., glucose, lactate, electrolytes, and drugs) that closely correlate with blood concentrations and biosensing applications. These MNs can achieve swelling ratios of ∼100–500%, enabling efficient ISF uptake and continuous analyte transport to integrated sensing platforms. Additionally, the process allows fabrication of uniform, tapered structures with smooth surfaces, improving reproducibility and reducing insertion force variability. Compared to micromolding, drawing lithography offers superior control over aspect ratio and tip sharpness without requiring complex molds; however, limitations include challenges in large-scale uniformity, batch-to-batch reproducibility, and restricted compatibility with highly viscous or rapidly crosslinking polymer systems.^47,51,60,64

4.4. Injection molding

Injection molding is widely employed for the large-scale fabrication of robust thermoplastic microneedles (MNs), particularly using polymers such as poly(lactic acid) (PLA) and poly(lactic-co-glycolic acid) (PLGA).⁶⁵ In this process, molten polymer is injected into precision-engineered molds under high pressure (typically ∼50–150 MPa), followed by controlled cooling to solidify the MN structures. The technique enables fabrication of MNs with well-defined geometries, including needle heights of ∼300–1000 µm, base diameters of ∼100–300 µm, and tip radii typically below ∼20–50 µm, ensuring reliable skin penetration.

Injection-molded MNs exhibit excellent mechanical strength due to the high stiffness of thermoplastics (Young's modulus ∼1–3 GPa), with insertion forces often exceeding ∼0.5 N per needle, making them suitable for repeated or long-duration applications. The method offers high reproducibility, low batch-to-batch variability, and compatibility with industrial-scale production, including cycle times of seconds to minutes per batch depending on mold complexity. These features make injection molding particularly attractive for manufacturing long-acting biodegradable MN systems for sustained drug release.⁶⁶

However, the process requires elevated temperatures (typically ∼150–250 °C), which may limit the incorporation of thermolabile drugs such as proteins or vaccines. Additionally, mold fabrication costs and limited flexibility in design modification can restrict rapid prototyping compared to additive manufacturing approaches. Despite these limitations, injection molding remains a highly efficient and scalable method for producing mechanically robust MN arrays with consistent performance.

4.5. Laser cutting and etching

Laser-based techniques are primarily employed for prototyping non-dissolving polymer microneedles (MNs) or for fabricating high-precision master molds used in subsequent micromolding processes. Techniques such as laser cutting and laser ablation enable micron-scale patterning with typical feature resolutions of ∼10–50 µm, allowing fabrication of customized MN geometries with controlled needle heights (∼300–1000 µm) and tip diameters often below ∼20–40 µm. These methods are particularly useful for rapid design iteration and precise structuring of hard polymer substrates or mold materials.^34,50,67

Laser-fabricated MNs exhibit good geometric accuracy and reproducibility, and when used for mold fabrication, they significantly enhance the fidelity of replicated polymeric MN arrays. However, direct laser processing of polymers may introduce thermal effects such as localized melting, surface roughness, or microcracks, which can compromise mechanical integrity and tip sharpness. In addition, the achievable mechanical performance depends on the base polymer, with non-dissolving systems typically exhibiting higher stiffness but limited biodegradability.

Despite their precision, laser-based techniques are generally limited by low throughput and higher processing time compared to micromolding or injection molding, making them less suitable for large-scale production. As a result, they are predominantly used in research and prototyping stages or as a complementary tool for mold fabrication rather than as a standalone manufacturing method for commercial MN production.

4.6. Hot embossing

Hot embossing is a widely used replication technique for polymeric microneedle (MN) fabrication, in which a thermoplastic polymer is heated above its glass transition or melting temperature (typically ∼80–200 °C, depending on the polymer) and pressed against a micro structured mold under controlled pressure (∼1–10 MPa) to replicate high-aspect-ratio MN arrays.^53,60,67 This method enables fabrication of MNs with well-defined geometries, including needle heights of ∼300–1000 µm, base diameters of ∼100–300 µm, and tip radii typically below ∼10–30 µm, ensuring effective skin penetration and structural consistency.

Hot embossing offers several advantages, including simple tooling, low material waste, and good dimensional fidelity, making it suitable for thermoplastics such as PMMA, PLGA, and cycloolefin polymers. The resulting MNs typically exhibit sufficient mechanical strength (Young's modulus ∼1–3 GPa for thermoplastics) and insertion forces exceeding ∼0.1–0.5 N per needle, enabling reliable penetration of the stratum corneum. Additionally, the technique supports fabrication of both solid and porous MN structures, which can be tailored for drug diffusion or controlled release applications.⁶⁷

Process parameters such as embossing temperature, applied pressure, holding time (∼1–10 min), and cooling rate critically influence mold filling efficiency, tip sharpness, and final mechanical performance.⁶⁸ Improper optimization may lead to defects such as incomplete cavity filling, tip blunting, or residual stresses. While hot embossing is more scalable than laboratory-scale prototyping techniques and offers better reproducibility than some additive methods, it is still relatively slower than injection molding and less suitable for thermolabile drug incorporation due to elevated processing temperatures.

4.7. Hybrid and advanced methods

Electrospinning assisted molding, along with laser micromachining and injection molding, further expands the design flexibility and scalability of polymeric microneedle (MN) fabrication.⁶⁹ In electrospinning-assisted approaches, nanofibrous polymer mats (fiber diameters typically ∼100–1000 nm) are integrated into micromolds or layered within MN structures, resulting in composite architectures with enhanced mechanical strength and increased surface area for drug loading. This hybrid technique enables precise control over microstructure, where parameters such as fibre density, alignment, and composition can be tuned to modulate drug loading capacity and release kinetics, supporting sustained or biphasic drug delivery profiles.

Electrospun MNs have demonstrated improved mechanical robustness, with reinforced structures capable of achieving insertion forces >0.1–0.3 N per needle, while maintaining high porosity that facilitates diffusion-controlled release. This makes them particularly suitable for delivering macromolecules such as peptides, proteins, and vaccines, where preservation of bioactivity and controlled release are critical. Additionally, the interconnected fibrous network enhances permeability and enables efficient incorporation of sensitive therapeutics without exposure to high temperatures or harsh processing conditions.^58,63,70

4.8. Droplet air blowing

Droplet air blowing (DAB) is a mold-free fabrication technique for polymeric microneedles (MNs), particularly suited for dissolving and drug-loaded systems. In this method, polymer droplets are deposited onto a substrate and elongated into conical MN structures using a controlled air stream, followed by rapid solvent evaporation and solidification. This process enables fabrication of MNs with typical heights of ∼300–800 µm and tip diameters in the range of ∼10–50 µm, although precise geometric control remains limited compared to mold-based techniques. DAB allows direct incorporation of therapeutic agents into water-soluble polymers such as polyvinyl alcohol (PVA), polyvinylpyrrolidone (PVP), carboxymethyl cellulose (CMC), and hyaluronic acid, achieving high drug-loading efficiencies due to minimal processing steps and reduced material loss.

Compared with micromolding and hot embossing, DAB offers advantages including lower fabrication cost, simplified processing, and elimination of mold-related constraints, making it attractive for rapid prototyping and drug-loaded MN fabrication. However, challenges such as variability in needle geometry, limited reproducibility, and relatively lower mechanical strength (typically insertion forces ∼0.05–0.2 N per needle) may affect penetration efficiency and reliability. Consequently, DAB is primarily applied to solid or dissolving MN systems for transdermal drug delivery rather than structurally complex or load-bearing applications.⁷¹

Overall, each fabrication technique presents distinct advantages and limitations, and the selection depends on the intended MN application, polymer properties, and desired drug release profile, as summarized in Table 9. While micromolding and injection molding remain preferred for large-scale manufacturing due to their reproducibility and throughput, advanced techniques such as additive manufacturing and electrospinning offer greater flexibility for complex architectures and personalized therapeutic systems. Continued advances in polymer design and fabrication technologies are expected to further expand the capabilities of polymeric MNs in drug delivery and biosensing applications.^67,69–71

Table 9. Key fabrication methods for polymeric microneedles.

Fabrication technique	Principle/process description	Compatible polymers	MN characteristics	Key advantages	Limitations	Scalability & translational potential	Ref.
Mold casting (micromolding)	Polymer solution or melt cast into microcavities, followed by drying/curing	PVP, PVA, PVP, HA, PL, PLGA, CMC	High geometric fidelity, sharp tips, uniform dimensions	Simple, low-cost, widely adopted, compatible with drugs	Limited throughput, drying-induced defects, batch variability	Moderate; suitable for lab-to-pilot scale	10, 35, 42, 47, 49, 50, 57, 58 and 60
Solvent casting + centrifugation	Centrifugal force drives polymer solution into mold cavities	PVP, PVA, HA, chitosan	Improved tip filling, enhanced mechanical strength	High tip replication accuracy, reduced air entrapment	Requires centrifugation equipment, limited scale	Moderate; semi-scalable	16, 25, 35, 37, 58 and 60
Vacuum-assisted micromolding	Vacuum applied to remove trapped air during mold filling	PVA, PVP, gelatin, HA	Uniform needles with reduced porosity	Better cavity filling, improved reproducibility	Additional process complexity	Moderate	16, 21, 35, 37, 58 and 60
Hot embossing/thermal molding	Heated polymer pressed into rigid microstructured mold	PLA, PLGA, PMMA	High mechanical strength, precise geometry	Solvent-free, good dimensional control	High temperature may degrade drugs	High; suitable for mass production	42, 53, 60, 67 and 68
Injection molding	Molten polymer injected into metal micro-molds under pressure	PLA, PLGA, CMC	Excellent uniformity, high strength	High throughput, industrial scalability	Expensive tooling, limited drug incorporation	Very high; commercial manufacturing	49, 57, 65 and 66
3D printing (additive manufacturing)	Layer-by-layer fabrication via micro-stereolithography or two-photon polymerization	PEGDA, resins, acrylates	Customizable geometries, complex architectures	Design flexibility, rapid prototyping	Limited resolution, low throughput	Low to moderate; mainly prototyping	10, 28, 34, 38, 41, 49 and 60
Electro-drawing/drawing lithography	Polymer drawn into needle shapes using mechanical pulling	PVA, PEG, silk fibroin	Smooth surfaces, tapered needles	No mold required, simple setup	Poor dimensional uniformity	Low	10, 28, 34, 37, 38, 47, 49, 51, 53, 60 and 64
28 laser cutting and etching	Focused laser beams (e.g., excimer, CO2, femtosecond) are used to ablate or pattern polymer substrates, forming microneedle geometries	PMMA, polycarbonate (PC), PLA, PLGA, PET	Solid microneedles with sharp tips, moderate aspect ratios, good dimensional accuracy	Maskless, high precision, rapid prototyping, suitable for hard thermoplastics	Thermal damage, limited tip sharpness, material wastage, not ideal for dissolving MNs	Moderate scalability; mainly suitable for prototyping and niche production rather than mass manufacturing	30, 34 and 67
Droplet air blowing (DAB)	Polymer solution droplets are stretched into microneedle shapes by a controlled air stream, followed by solvent evaporation and solidification	PVA, PVP, CMC, hyaluronic acid, chitosan	Solid/dissolving microneedles, drug-loaded, tapered geometry	Mold-free, low cost, high drug-loading efficiency, simple processing	Limited geometric control, variable mechanical strength, unsuitable for hollow MNs	High potential for low-cost scalable production of dissolving MNs	38, 47, 60, 70 and 71

5. Applications of polymeric microneedles: the paradigm shift-from passive drug delivery to smart, closed-loop therapeutics

Polymeric microneedles (PMNs) are a versatile, minimally invasive platform for drug delivery, biosensing, and therapeutic interventions. They are applied in transdermal drug delivery, vaccines, chronic disease management, biosensing, cosmetic therapy, and localized treatment of autoimmune disorders, leveraging polymers' tunable properties such as biodegradability, mechanical strength, and controlled or stimuli-responsive drug release,^13,60,68 as illustrated in Fig. 7. PMNs are central to closed-loop therapeutic systems, where drug administration responds dynamically to physiological signals, e.g., glucose in diabetic patients.⁶¹

Fig. 7. Illustration of applications of polymeric microneedle.

5.1. Long-acting, controlled transdermal drug delivery

Polymeric microneedles have gained significant attention as advanced transdermal platforms capable of enabling sustained and long-acting drug delivery while overcoming the limitations of conventional dosage forms. By painlessly breaching the stratum corneum, microneedles create micro-scale pathways (∼50–150 µm depth) that allow therapeutic agents to diffuse into viable skin layers, enabling controlled systemic absorption over extended periods. Biodegradable and biocompatible polymers such as PVA, PLGA and hyaluronic acid have been extensively employed, where drug release can be tuned via polymer dissolution, swelling, or degradation kinetics.^16,72 Quantitatively, dissolving MNs typically release >70–90% of loaded drug within minutes to hours, whereas hydrogel-forming and PLGA-based systems can sustain release over several days to weeks, depending on crosslinking density and polymer composition.

Sustained-release polymeric microneedles, including dissolving, hydrogel-forming, and implantable designs, have demonstrated the ability to maintain therapeutic drug levels for days to weeks following a single application. These systems reduce dosing frequency and improve patient adherence, which is particularly critical for chronic conditions such as diabetes, hormonal disorders, and infectious diseases. For instance, hydrogel-forming MNs can swell up to ∼100–500% of their dry weight, forming diffusion pathways that enable controlled drug flux from an attached reservoir at rates of ∼0.1–1 mg cm⁻² h⁻¹, thereby maintaining stable plasma concentrations. In insulin delivery, MN-based systems have shown prolonged pharmacodynamic effects lasting up to 24–48 h with reduced peak–trough fluctuations compared to subcutaneous injections, leading to improved glycemic control and reduced risk of hypoglycemia. Similarly, PLGA-based implantable MNs have demonstrated sustained release of small molecules and proteins over periods extending to weeks, highlighting their suitability for chronic disease management.⁷²

MN systems such as PLGA-based implantable MNs, hydrogel-forming MNs, and dissolving polymeric MNs enable sustained release ranging from several hours to weeks/month, depending on polymer degradation and diffusion kinetics. Reported quantifiable outcomes include initial burst release within minutes–hours followed by controlled release extending up to ∼7–30 days (hydrogel/implantable systems) or longer for PLGA-based depots, with drug loading efficiencies typically >80–90% and bioavailability improvements of ∼1.5–3× compared to conventional transdermal patches in preclinical models.

Kirkby et al., emphasized that polymer-based microneedle systems significantly expand the transdermal delivery landscape for biologics, addressing challenges related to stability, dose precision, and patient compliance. These attributes are particularly advantageous for long-term therapies requiring consistent plasma concentrations.⁷³ Advances in fabrication technologies, such as micromolding, photopolymerization, and additive manufacturing, have further enhanced the design flexibility of polymeric microneedles for sustained delivery applications. Tailoring microneedle geometry, polymer composition, and drug distribution within the needle matrix enables precise control over release profiles.⁷⁴ These developments support the feasibility of scaling polymeric microneedles for clinical use. Among long-acting applications, insulin delivery has emerged as a prominent example demonstrating the clinical relevance of polymeric microneedles. Sustained insulin release from polymeric microneedle patches has shown improved glycemic control while minimizing hypoglycemia risks associated with bolus injections. Zhao et al., reported that microneedle-based insulin systems can achieve prolonged pharmacodynamic effects, highlighting their potential to transform diabetes management.⁷⁵

5.2. Vaccine and immunization delivery

Polymeric microneedle platforms have rapidly evolved as next-generation immunization technologies, providing a minimally invasive, pain-free alternative to conventional intramuscular and subcutaneous injections. By penetrating the stratum corneum, dissolving polymeric microneedles deliver antigens directly to skin-resident dendritic and Langerhans cells, enabling efficient immune activation at lower antigen doses (10–50% of conventional injected doses)^31,76 This targeted interaction with dermal antigen-presenting cells enhances both humoral and cellular immunity, with studies reporting elevated IgG titre and balanced Th1/Th2 responses in animal models—often comparable or superior to traditional needle-based vaccination.^50,70 In clinical and translational studies, influenza vaccine-loaded MN patches have demonstrated equivalent or enhanced immunogenicity compared to intramuscular injection, while also improving patient compliance and eliminating sharps waste.^14,76

Recent advancements have focused on integrating nanoparticle-based vaccine formulations within polymeric MNs to further enhance antigen stability and immune response. For example, PLGA nanoparticle-loaded dissolving MNs for Clostridium perfringens toxin delivery have demonstrated sustained antigen release over several days, resulting in prolonged immune stimulation and improved protective efficacy.⁷⁷ Controlled-release MN systems can extend antigen presentation duration from hours to multiple days, which is critical for enhancing germinal centre responses and long-term immune memory. These systems typically achieve release efficiencies of >70–90% while preserving antigen integrity, addressing key challenges associated with protein degradation and rapid clearance. Evidence from sustained transdermal drug delivery system including 3D-printed PLA microneedle arrays for prolonged estradiol valerate delivery⁷⁸ and hydrogel-forming PVA/poly(vinyl pyrrolidone) microneedles for methotrexate administration highlights how polymer composition, architecture, and swelling behavior can be precisely engineered to regulate release kinetics. These design principles are directly applicable to vaccine delivery, where prolonged antigen exposure can enhance immune memory and potentially reduce booster requirements.^79–81

5.3. Biosensing and diagnostic applications

Polymeric microneedles provide minimally invasive access to interstitial fluid (ISF), the extracellular fluid located within the intercellular spaces of tissues, primarily in the viable epidermis and dermis, with a composition similar to plasma but lacking red blood cells and large proteins which closely reflects blood biomarker composition, making them highly suitable for biosensing applications. Hydrogel-forming MNs can exhibit swelling ratios of ∼100–500%, facilitating efficient ISF uptake (in the µL range) for continuous or periodic monitoring physiologically relevant analytes (e.g., glucose, lactate, electrolytes, and drugs) that closely correlate with blood concentrations. Their mechanical compliance, biocompatibility, and tunable swelling or dissolution properties enable safe, repeated, or long-term use without significant tissue damage.^13,33,55,81 Compared with conventional blood sampling, microneedle-based sensing reduces pain, infection risk, and procedural complexity, supporting continuous and point-of-care diagnostics. Their flexible, skin-conformal polymeric structures are particularly advantageous for wearable health monitoring systems.

Recent advances have integrated polymeric microneedles with electrochemical, optical, and hydrogel-based sensing modalities for real-time detection of analytes such as glucose, lactate, electrolytes, and bilirubin.⁸² Electrochemical MN sensors have demonstrated continuous glucose monitoring with response times typically within ∼1–5 minutes and strong correlation with blood glucose levels (correlation coefficients often >0.9). Hydrogel-forming MNs enable passive ISF extraction through swelling-driven diffusion without leaving residues in the skin, supporting sustained monitoring over several hours to days.⁸³ Optical MN-based systems further provide high sensitivity and non-electrical readouts, improving safety and compatibility with wearable devices.^84,85 Collectively, these developments establish polymeric microneedles as sustained transdermal bio interfaces capable of reliable, long-term diagnostic monitoring.⁸⁶ Despite these promising results, several challenges remain. Signal stability can be affected by biofouling and enzyme degradation during long-term use, while limited ISF extraction volume and variability in skin insertion efficiency may impact measurement accuracy.

5.4. Cosmetic and dermatological applications

In dermatology and cosmetology, polymeric microneedles (PMNs) enhance transdermal delivery by bypassing the stratum corneum, significantly improving the penetration of peptides, antioxidants, and other active agents compared to conventional topical formulations. Dissolving microneedles additionally stimulate collagen production, providing a minimally invasive approach for skin rejuvenation and the treatment of acne, hyperpigmentation, and melasma. Innovations such as collagen-based microneedles with flexible pedestals improve mechanical adaptability and conformal skin contact, enabling efficient drug deposition into epidermal and dermal layers for anti-aging and chronic skin disorder therapies.⁸⁷

PMNs also enable targeted treatment of deep cutaneous infections and inflammatory conditions by delivering antifungal and antimicrobial agents directly to affected sites.⁸⁸ Biodegradable microneedle patches loaded with clindamycin hydrochloride, for instance, provide sustained drug release within pilosebaceous units, reducing bacterial load, inflammation, and dosing frequency.⁸⁹ Through careful selection of polymer composition and microneedle geometry, these systems optimize localized bioavailability and patient compliance, establishing PMNs as promising platforms for long-acting dermatological therapy and infection management.^88,89

5.5. Wound healing and tissue regeneration

Polymeric microneedles (MNs) play a dual functional role in wound healing by serving as localized drug delivery systems and temporary micro-scaffolds that actively regulate the wound microenvironment.⁹⁰ By enabling direct transdermal administration of growth factors, anti-inflammatory agents, and antimicrobial compounds, MNs enhance cellular migration, angiogenesis, and extracellular matrix remodeling while minimizing systemic exposure. The use of biodegradable polymers allows gradual matrix degradation and sustained therapeutic release, which are critical for chronic and non-healing wounds such as diabetic ulcers and pressure sores.

Recent studies have demonstrated that MN-based delivery systems can significantly accelerate wound closure rates, with reported improvements of ∼30–60% faster healing compared to conventional topical treatments. Hydrogel-based MNs, which can exhibit swelling ratios of ∼100–400%, maintain a moist wound environment and enable continuous release of therapeutic agents, while also supporting tissue regeneration. Additionally, photothermally active polymeric MNs have shown the ability to achieve localized temperature increases (∼40–45 °C) under external stimulation, promoting antibacterial activity and enhancing tissue repair processes.⁹¹ These multifunctional systems enable controlled and stimuli-responsive drug delivery tailored to the dynamic wound environment. In parallel, nanofiber-based regenerative dressings have demonstrated that biomaterial-mediated modulation of wound architecture and biochemical signalling significantly accelerates diabetic wound healing, reinforcing the importance of localized and sustained delivery strategies.⁹² Collectively, these findings highlight the strong potential of polymeric MN platforms as advanced tools in regenerative medicine, offering precise, sustained, and minimally invasive therapeutic delivery for improved wound care.

5.6. Pain management and local anesthesia

Polymeric microneedles (PMNs) enable controlled, localized delivery of analgesics such as lidocaine, ibuprofen, and opioid-based agents, maintaining therapeutic concentrations at the target site while reducing systemic exposure and dosing frequency.⁹³ Sustained-release systems, including hydrogel-forming and biodegradable polymeric matrices, allow prolonged drug diffusion over several hours to days, maintaining therapeutic concentrations at the target site with reduced dosing frequency. This localized delivery approach significantly lowers systemic side effects compared to oral or injectable administration and improves patient compliance, particularly in pediatric, geriatric, and needle-phobic populations.^57,93,94

Recent advances include rapidly dissolving polymeric microneedles for on-demand local anesthesia. For example, PVP/PVA-based dissolving systems incorporating lidocaine hydrochloride demonstrate sufficient mechanical strength for skin penetration (>0.1–0.3 N per needle) and can release >80–90% of encapsulated lidocaine within minutes, achieving faster onset of anesthesia (∼5–10 minutes) compared to conventional topical formulations (∼20–60 minutes). Emerging “smart” microneedles responsive to inflammation-related stimuli, such as pH or enzymatic changes, are also being explored to dynamically modulate analgesic release according to physiological conditions, paving the way for personalized pain therapy.⁹⁴

5.7. Cancer therapy

Polymeric microneedle (MN) platforms enable localized and controlled delivery of chemotherapeutic and immunotherapeutic agents, making them particularly effective for superficial malignancies such as melanoma and recurrent breast cancer. By delivering drugs directly into tumor-associated skin or tissue (∼200–800 µm depth), MNs significantly enhance local drug concentration while reducing systemic exposure and off-target toxicity. Studies have reported local drug accumulation increases of ∼2–5 fold compared to conventional topical or systemic administration, with reduced systemic drug levels and associated side effects. Their minimally invasive nature, combined with compatibility for combination therapies and adaptability through advanced fabrication techniques such as additive manufacturing and 3D printing, further supports their role as precision cancer delivery systems.

A key advancement in this field is the integration of tumor microenvironment (TME)-responsive polymers into MN platforms to achieve spatiotemporally controlled drug release. These systems respond to tumor-specific triggers such as acidic pH (∼pH 6.5–6.8), elevated glutathione concentrations (∼2–10 mM), or overexpressed enzymes, enabling selective intratumoral drug release and minimizing systemic toxicity.⁹⁵ For example, dissolving MN arrays incorporating pH-responsive micelles loaded with aggregation-induced emission luminogens (AIEgens) have demonstrated efficient intradermal deposition, followed by acidity-triggered destabilization and enhanced photothermal conversion, achieving significant tumor ablation efficiencies (>70–90% tumor reduction in preclinical models).⁹⁶ Hydrogel-forming and dissolving MNs further enable sustained intratumoral drug release over several days, overcoming limitations such as poor tumor perfusion and multidrug resistance by maintaining high local drug concentrations.⁹⁷

5.8. Hormonal and reproductive health

Polymeric microneedles (MNs) are highly suited for hormonal and reproductive health applications requiring sustained, controlled, and patient-friendly drug delivery. Early studies demonstrated effective transdermal hormone transport with improved pharmacokinetic stability compared to oral and injectable routes. More recently, long-acting biodegradable MN patches have been developed for contraception and hormone replacement therapy, maintaining stable plasma hormone levels for weeks to months after a single application.⁹⁸

A notable example is the development of self-administered MN patches for contraceptive delivery, where levonorgestrel-loaded biodegradable MNs have demonstrated sustained release for >30 days, maintaining plasma hormone levels within the therapeutic window while significantly reducing dosing frequency. Quantitatively, such systems achieve controlled release rates in the range of µg per day, with consistent drug diffusion governed by polymer degradation kinetics and MN geometry. Advances in polymer design, including poly(lactic-co-glycolic acid) (PLGA)-based systems, enable tunable release profiles by adjusting molecular weight, crystallinity, and degradation rates, ensuring prolonged therapeutic efficacy with minimal user intervention.⁹⁹

5.9. Smart microneedle system – theragnostic applications

Smart polymeric microneedles (PMNs) are emerging as advanced theragnostic platforms that integrate real-time diagnostic sensing and controlled therapeutic delivery within a single minimally invasive device. By combining stimuli-responsive polymers, embedded biosensors, and drug reservoirs, these systems enable closed-loop, on-demand treatment guided by physiological feedback. Typically penetrating ∼100–500 µm into the skin, PMNs access interstitial fluid (ISF) for continuous biomarker monitoring while delivering therapeutics with precise spatiotemporal control. Their tunable mechanical strength (>0.1–0.5 N per needle), biodegradability, and compatibility with wearable formats support prolonged, stable skin interfacing and improved patient compliance.^100,101

The core of smart PMN functionality lies in stimuli-responsive polymer chemistry, where materials undergo predictable physicochemical changes in response to cues such as pH, glucose levels, redox conditions, enzymes, or hypoxia. These responsive transitions convert biological signals into programmable drug release behaviors, forming the foundation for intelligent, personalized transdermal therapy systems.¹⁰²

Among the most mature examples are glucose-responsive insulin delivery platforms. For instance, hypoxia-sensitive vesicle-integrated MN patches have demonstrated glucose-triggered insulin release within minutes, maintaining normoglycemic levels (∼80–120 mg dL⁻¹) for extended periods in diabetic animal models while reducing hypoglycemia risk.^103–105 These polymeric microneedle arrays demonstrated rapid glucose-responsive insulin release, effective glycemic regulation, and reduced hypoglycemia risk in diabetic animal models. More recently, Huang et al. (2024) advanced this concept by integrating electronic glucose with fluidic polymeric microneedles, enabling simultaneous glucose monitoring sensing (response time ∼1–5 min, correlation >0.9 with blood glucose) and controlled insulin delivery within a single wearable theragnostic system as a critical step toward fully autonomous, closed-loop diabetes management.¹⁰⁴

Smart PMNs are also widely explored for continuous biosensing applications. Electrochemical MN sensors enable real-time monitoring of analytes such as glucose, lactate, electrolytes, and inflammatory mediators with high sensitivity and rapid response times (∼minutes), while hydrogel-based systems facilitate ISF extraction through swelling (∼100–500% expansion). To improve long-term stability, enzyme-free sensing strategies such as phenylboronic acid (PBA), aptamers, and molecularly imprinted polymers (MIPs) have been developed, offering reversible binding, enhanced selectivity, and reduced biofouling. Integration with conductive polymers and nanostructured electrodes further enhances signal stability and durability during prolonged wear.^32,55,83,106

In oncology, smart PMNs exploit tumor microenvironment (TME) characteristics such as acidic pH (∼6.5–6.8), elevated glutathione levels (∼2–10 mM), and hypoxia to enable site-specific drug release and imaging. For example, pH-responsive PMNs incorporating aggregation-induced emission luminogens (AIEgens) have demonstrated enhanced photothermal conversion and tumor ablation efficiencies exceeding ∼70–90% in preclinical models.⁹⁶ Similarly, redox- and hypoxia-responsive systems enable selective drug activation within tumor tissues, improving therapeutic efficacy while minimizing systemic toxicity.^96,107

Despite these advances, challenges remain, including limited drug loading capacity, potential signal drift in long-term sensing, integration complexity of multi-component systems, and scalability for clinical translation. Nonetheless, the convergence of polymer chemistry, nanotechnology, and flexible electronics is rapidly advancing PMNs toward fully autonomous, closed-loop theragnostic systems. These multifunctional platforms capable of simultaneous diagnosis, monitoring, and responsive therapy hold significant promise for personalized medicine, particularly in chronic disease management and precision oncology.¹⁰⁷

6. Manufacturing, regulatory, and translational considerations

The translation of polymeric microneedles (PMNs) from lab research to clinical and commercial use requires careful attention to design, manufacturing, scalability, and regulatory compliance. Manufacturing must ensure reproducibility, precise needle geometry, mechanical strength, uniform drug loading, and sterility. Techniques such as micromolding, injection molding, 3D printing, and hybrid electrospinning should be optimized for high-throughput production while maintaining consistent pharmacokinetics and performance. Quality control through mechanical testing, dissolution profiling, and release kinetics evaluation ensures batch-to-batch consistency and patient safety.^79,102

From a translational perspective, challenges include scaling laboratory fabrication methods to industrial production, achieving cost-effectiveness, ensuring patient adherence, and integrating PMNs with electronic or closed-loop systems. Material selection must balance biocompatibility, mechanical robustness, degradability while aligning with regulatory acceptance criteria. Additional considerations such as shelf stability (often targeting >12–24 months), cold-chain independence, and ease of self-administration are critical for widespread adoption, particularly in resource-limited settings.¹⁰⁸

Regulatory approval pathways for PMNs often classified as combination products (drug–device) require compliance with established frameworks such as the United States Food and Drug Administration (US FDA) combination product regulations and the European Union Medical Device Regulation (EU MDR 2017/745). These frameworks mandate comprehensive evaluation of safety and efficacy, including sterility assurance, biocompatibility (ISO 10993 standards), skin penetration performance, local tolerability, immunogenicity, systemic exposure, packaging integrity, and storage stability. Human factors such as application technique, wear comfort, and training requirements must also be addressed. Ultimately, successful clinical translation of PMNs will depend on coordinated advances in polymer science, scalable manufacturing, device engineering, biosensor integration, and regulatory strategy to ensure both safety and real-world usability.^109,110

7. Future directions and conclusions

Polymeric microneedles (PMNs) continue to represent a cornerstone technology for next-generation transdermal therapeutics and biosensing, offering minimally invasive, patient-friendly drug delivery and real-time monitoring capabilities. The convergence of smart, stimuli-responsive polymers with advances in additive manufacturing and wearable bioelectronics is expected to accelerate the development of personalized, closed-loop MN systems for precision medicine. However, widely used polymers such as PVP, PVA, PEGDA, and GelMA—while effective—still present limitations including protein fouling in biosensing applications, limited long-term mechanical stability, restricted drug loading capacity, and suboptimal control over prolonged release kinetics.

Emerging polymer classes offer targeted solutions to these challenges. Zwitterionic hydrogels, for example, exhibit strong anti-fouling properties due to their charge-balanced structures, which resist nonspecific protein adsorption and biofilm formation. This significantly enhances sensor stability and accuracy during long-term interstitial fluid monitoring, addressing a key limitation in continuous biosensing platforms. Polysaccharide derivatives (e.g., modified hyaluronic acid, chitosan, and cellulose-based systems) provide excellent biocompatibility, intrinsic bioactivity, and enzymatically tunable degradation, enabling safer, more controlled drug release and improved tissue integration. Biodegradable polyesters such as advanced PLGA variants or polycaprolactone-based systems offer tunable degradation rates (ranging from days to months), making them particularly suitable for long-acting therapeutics and sustained-release MN implants. Meanwhile, hybrid composite polymers incorporating nanoparticles, conductive materials, or nanofibrous networks enhance mechanical strength, electrical conductivity, and drug-loading capacity, thereby supporting multifunctional applications such as theranostics and electrochemical sensing.

These material innovations directly address current limitations in MN systems, including biofouling, insufficient mechanical robustness, limited payload capacity, and lack of multi-modal functionality. As a result, they enable expanded applications in long-acting drug delivery, multi-drug combination therapy, continuous biosensing, and integrated theragnostic platforms.

Looking forward, polymeric microneedles are poised to play a transformative role in closed-loop drug delivery by integrating minimally invasive biosensing, real-time physiological monitoring, and responsive therapeutic release within a single platform. Advances in polymer chemistry, microfabrication, and system integration have already enabled MNs capable of accessing dermal interstitial fluid with high fidelity, supporting continuous monitoring of clinically relevant biomarkers with minimal discomfort. The incorporation of stimuli-responsive polymers—sensitive to glucose, pH, hypoxia, redox potential, or enzymatic activity—further enables autonomous, self-regulated drug delivery, marking a shift from conventional open-loop dosing toward adaptive, patient-specific therapy. These systems have demonstrated strong potential across diverse applications, including metabolic disorders, cancer theranostics, pain management, and immunotherapy, where precise spatiotemporal control of drug release is essential. Continued interdisciplinary advances will be critical to translating these next-generation PMN platforms into clinically viable and commercially scalable healthcare solutions.

Despite these advances, the clinical translation of closed-loop polymeric microneedle systems remains challenged by system integration, long-term stability, manufacturing scalability, and regulatory complexity. Achieving reliable coupling between biosensing accuracy and therapeutic response, ensuring polymer and sensor stability under prolonged skin contact, and validating performance across diverse patient populations are key hurdles that must be addressed. Future progress will likely depend on multidisciplinary convergence—combining smart polymers, microelectronics, artificial intelligence–assisted control algorithms, and standardized regulatory frameworks—to realize fully autonomous, wearable therapeutic systems. As these challenges are overcome, polymeric microneedle-based closed-loop platforms are poised to redefine precision medicine by enabling continuous monitoring, on-demand therapy, and improved patient adherence, ultimately advancing toward next-generation personalised healthcare solutions.

Conflicts of interest

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.

Biographies

Biography

Mythily V.

Mrs Mythily V. is a PhD candidate in the Department of Biomedical Engineering at Anna University, Chennai, Tamil Nadu, India. She obtained her BTech. degree in Biotechnology from Bharathidasan University, Tiruchirappalli, Tamil Nadu, and her ME in Biomedical Engineering from Anna University, Chennai. Her research interests focus on microneedle-based transdermal drug delivery systems, biosensors, and wearable biomedical devices. She has published several research articles in peer-reviewed journals and presented her work at international conferences. Her current research is centered on the development of polymer-based microneedles for minimally invasive disease diagnostics.

Biography

Sreeja B. S.

Dr Sreeja B. S. is a distinguished academic with 20 years of teaching experience across various universities. She obtained her BE degree from Bharathidasan University in 2002, followed by ME and PhD degrees in 2004 and 2012, respectively. Currently, she is a faculty member at Anna University in Chennai, India. Dr Sreeja's research focuses on smart devices, MEMS and NEMS devices, sensors and actuators, and biosensors and their applications. Her extensive expertise and contributions to these fields underscore her commitment to advancing technological innovation and education. She is a life member of ISSS, a member of IEEE, and a life member of IETE. Additionally, she is a recipient of the DST International Travel Support and the Young Scientist Award. To her credit, Dr Sreeja has published 86 peer-reviewed articles, executed several research projects, and holds several patents.

Biography

S. Fouziya Sulthana.

Dr S. Fouziya Sulthana received the BE degree in Electronics and Communication from Bharadhidasan University, Tamil Nadu, in 2002, and the MTech and PhD degrees in Electronics and Communication from Pondicherry University, Puducherry, in 2008 and 2017, respectively. She has been in the teaching profession, since 2008. She is currently an Associate Professor with the Department of Mechatronics Engineering, SRM Institute of Science and Technology, Kattankulathur Campus. She has published many papers in refereed journals and conferences. Her current research interests include micro and nano sensors, and flexible and wearable electronics. She is a member of IETE, ISTE, SAE and IE(I).

Data availability

All data supporting the findings of this study are available within the article and its referenced literature.