3D bioprinted microneedles: merging drug delivery and scaffold science for tissue-specific applications

ABSTRACT

Introduction

Three-Dimensional (3D) microneedles have recently gained significant attention due to their versatility, biocompatibility, enhanced permeation, and predictable behavior. The incorporation of biological agents into these 3D constructs has advanced the traditional microneedle into an effective platform for wide-ranging applications.

Areas covered

This review discusses the current state of microneedle fabrication as well as the developed 3D printed microneedles incorporating labile pharmaceutical agents and biological materials for potential biomedical applications. The mechanical and processing considerations for the preparation of microneedles and the barriers to effective 3D printing of microneedle constructs have additionally been reviewed along with their therapeutic applications and potential for tissue engineering and regenerative applications. Additionally, the regulatory considerations for microneedle approval have been discussed as well as the current clinical trial and patent landscapes.

Expert opinion

The fields of tissue engineering and regenerative medicine are evolving at a significant pace with researchers constantly focused on incorporating advanced manufacturing techniques for the development of versatile, complex, and biologically specific platforms. 3D bioprinted microneedles, fabricated using conventional 3D printing techniques, have resultantly provided an alternative to 2D bioscaffolds through the incorporation of biological materials within 3D constructs while providing further mechanical stability, increased bioactive permeation and improved innervation into surrounding tissues. This advancement therefore potentially allows for a more effective biomimetic construct with improved tissue-specific cellular growth for the enhanced treatment of physiological conditions requiring tissue regeneration and replacement.

KEYWORDS:

• Microneedles
• additive manufacturing
• 3D printing
• bioactive delivery
• biomedical applications
• regulatory considerations

1. Introduction

The past decade has seen remarkable progress in the design and fabrication of microneedles (MNs). Due to the associated pain and needle phobias related to conventional injections, in addition to the existing limitations associated with oral drug delivery, MNs have historically been developed for percutaneous administration due to the attractive advantages of convenience, improved patient compliance, self-administration, low injury risk due to bleeding and infection, steady infusions, stable blood levels, and treatment termination flexibility. Depending on the intended application, these micron-scale needles can be arranged onto a patch substrate individually, or as rows, and can vary in length from 25 to 2500 µm, in width from 20 to 250 µm, and from 1 to 25 µm in tip diameter [1–4]. When many MNs are arranged on a normal transdermal patch or polymeric base, this is called an MN patch and MN array (MNA), respectively [5]. Usually, these MNs are designed as arrays in order to improve skin surface contact [6–8], where the MN type, geometry, shape, needle thickness, and needle density dictate MN performance and therapeutic efficacy [9], with drug-release being influenced by drug-binding affinity, the inherent properties of the matrix materials, the interior body environment, swelling, diffusion, and degradation controls [10,11]. Over the years, considerable interest has also been afforded to MNs due to their ability to percutaneously deliver not only synthetic pharmaceutical agents but also macromolecules, such as bovine serum albumin and insulin [12–14], while also having other applications in electro-responsive delivery [15], immunobiological administration [16], disease diagnosis [17] and monitoring [18–20], cosmetics [21] and biosensing [22].

1.1. Conventional MNs

MNs are broadly classified into solid MNs, hollow MNs, coated MNs, dissolving/polymer MNs, and hydrogel MNs (Figure 1), and depending on their application, confer various types of delivery strategies [4,6,12,23,24]. The original solid MNs enhanced permeation by poking holes onto the surface of the substrate, facilitating a ‘coat and poke’ approach, whereas hollow MNs facilitate a ‘poke and flow’ approach, whereby the drug is released through bore-like conduits. Coated MNs additionally have surface drug coatings which dissolves after tissue insertion [23,24], while dissolving MNs are composed of biocompatible and biodegradable polymers and employ the ‘poke and release’ strategy, which facilitates needle dissolution, influencing the release of the encapsulated drugs [4,6,24]. Hydrogel MNs are conversely composed of non-dissolving crosslinked hydrogels, where following the application of a drug-loaded adhesive patch onto the skin surface, the hydrogel MNs swell due to the skin’s aqueous media, thereby creating hydrogel channels which facilitate drug delivery deeper into the skin [23]. Notably, hydrogel MNs can be used for both fluid extraction and drug delivery [25,26].

Figure 1. Types of MNs, step-by-step process of their application, and corresponding mechanisms of drug delivery across a tissue barrier (reproduced from [23]).

MNs can also be classified based on their material composition, i.e. metal MNs (titanium and stainless steel), inorganic MNs (glass and silicon), and polymeric MNs (polyacrylic acid, chitosan, hyaluronic acid, and polymethyl methacrylate). The use of inorganic and metal MNs are limited in biomedical applications due to inorganic MNs being prone to breakage in the skin because of their inherent brittleness, while metal MNs, although offering greater strength and permeability, have a limited drug-loading capacity as well as posing a risk for skin rupture. Polymeric MNs are thus the favorable option due to their biocompatibility and high drug loading potential, and can additionally be engineered to have stimuli-responsive properties [27]. Table 1 highlights the unique characteristics, advantages, and disadvantages of these conventionally designed MNs, based on their type and common material composition.

Table 1. An overview of the different types of conventional MNs based on their characteristics, benefits, limitations, and materials (adapted from [28,29]).

MN Type	Characteristics	Advantages	Disadvantages	Material
Solid	Creates channels in the skin to allow drugs reach the lower skin layer. Adequate mechanical strength. Sharper tip.	Allows more drugs to pass into the skin. Easy to manufacture.	Damage to the skin and microincisions need to be closed to avoid infections.	Silicon Polymer Metal
Hollow	Empty shape to be filled with the drug. Ability to control drug release over time.	Handles a large dose/amount of drug solution.	Weak needles. Requires intensive care in terms of needle design and insertion method. Might cause leakage and clogging.	Silicon
Coated	Carries less amount of the drug due to the design. Ability to deliver proteins and DNA in a minimally invasive manner.	Delivers the drug quickly to the skin.	Prone to infection.	Silicon
Dissolving	Facilitates rapid release of macromolecules.	Ease of administration for patients with one-step application.	Requires technical expertise to manufacture. Takes time to dissolve.	Silicon
Hydrogel-forming	Absorbs fluids (due to its hydrophilic nature) and swells, creating channels for the delivery of drug molecules.	Can be applied to the skin leaving no residues, and has no clogging, unlike hollow microneedles.	Potential for localized tissue damage and has a slow swelling rate.	Polymer

1.2. MN fabrication techniques

Due to the wide-ranging and complex architectural subtypes that have been developed, numerous methods have been employed for the fabrication of conventional MNs, namely molding-based techniques [30]; injection molding [31]; cleanroom-free molding [32]; wet and dry etching [33]; micromilling (wet and/or dry cutting) [34]; photolithography (combined with thermal- and photo-polymerization) [35,36]; drawing lithography [37]; laser patterning [38]; and photolithography with an elastocapillarity-driven self-assembly mechanism [39]. These fabrication techniques are able to produce simple MN geometries to more complex bioinspired platforms, such as Honeybee stingers [40], the clawed toes of eagles [41], snake fangs [42], the teeth of limpets [43] and the mosquito/endoparasite Pomphorhynchus laevis’s proboscis [44,45]. It is important to note that the ultimate design of the MNs is highly dependent on the chosen MN type, the fabrication methods employed and the physical, mechanical, and chemical properties of the material used [46]. These variations, although broad, allow for the development of a large range of MNs for varying biomedical applications. Table 2 provides a summary of the various factors and considerations in the manufacture of MNs.

Table 2. Factors affecting the fabrication of MNs (reproduced from [47]).

Factor	Practical Parameters	Practical Considerations
Material properties	Mechanical properties (stiffness, strength, and toughness).	High stiffness and strength may cause materials to be more difficult to be molded or etched, while materials that are too brittle may be prone to breakage.
	Chemical properties (reactivity, solubility, and stability).	Highly reactive or unstable materials may require specialized handling or storage conditions.
	Biocompatibility (nontoxic, non-immunogenic, and non-inflammatory).	Material should be biocompatible for clinical or medical use.
	Compatibility with drug formulations (physical, chemical, and biological interactions).	The material should not adsorb or absorb the drug to avoid any loss of drug efficacy.
Device design	Needle length and diameter.	The length and diameter of the MNs should be optimized for the intended application, taking into account factors such as skin thickness and drug delivery requirements.
	Needle shape and geometry.	The shape and geometry of the MNs can affect their mechanical properties and ability to penetrate the skin.
	Delivery target.	Design considerations shall take into account the target site (depth of penetration topical or deeper penetration), and whether it is a local or systemic action/sensing.
Quality control	Inspection and testing.	Inspected and tested to ensure they meet quality standards and specifications (device dimensions, insertion force, insertion depth, failure force, skin irritation, skin permeation, payload release, lubrication, flexibility, shelf life, etc.)
	Precision and accuracy of alignment.	Alignment and registration of MNs are important for ensuring consistent performance and drug delivery
Cost	Materials and manufacturing costs, and ease of fabrication.	Materials should be economic. The material should be easy to fabricate using the chosen manufacturing technique.
Regulatory requirements	FDA and other regulatory requirements.	MNs may be subject to regulatory requirements, such as safety and efficacy testing, before they can be approved for use.

Abbreviations: FDA, Food and Drug Administration.

Three-Dimensional (3D) printed MNs, either through direct printing or 3D molding, have been incorporated into many of these conventional techniques to increase the accuracy and reproducibility of MN platforms, incorporating the properties and advantages of MNs as described in Table 1, while addressing many of the factors outlined in Table 2. These 3D printed MNs have been developed in numerous studies to deliver drug molecules and more recently for more complex applications utilizing cells and peptides, whereby the known potential for protection and delivery of the bioactive molecules through advanced pharmaceutical design can be combined with the structural properties of MNs for enhanced biomedical applications.

This review focuses on the 3D MNs that have been developed to deliver biological cargo for complex biomedical applications including its potential for use in tissue engineering and regenerative medicine. A brief introduction into the applications, considerations, and limitations of MN 3D printing technologies has been provided in addition to the various 3D printed MNs loaded with labile bioactives, such as proteins, peptides, and cells. Further emphasis has been provided on this new, yet potentially revolutionary field of research, where advanced drug delivery systems and the intricate nature of additive manufacturing can be integrated for the fabrication of complex biological platforms for the treatment of conditions requiring tissue regeneration or replacement.

2. 3D printed MN technologies

MNs fabricated through additive manufacturing, commonly known as 3D printing, were designed to overcome the limitations of conventional methods and to enhance reproducibility, accuracy, and precision of the manufacturing process [9,48]. Offering versatility and customizability, 3D printing involves the use of a 3D computer-aided design (CAD) model input which is exported to the 3D printer, ultimately translating into the deposition of materials in a layer-by-layer fashion, producing a physical output [9,49–52]. Used in tissue and organ regeneration [36,53–59], wound healing [60,61], personalized medicines, such as 3D printed tablets [62–64], medical devices including implants and prostheses [56,65], organ-on-a-chip devices [66–69], cancer therapy [70,71], stem cell research [72,73] and fertility and embryology [74–76], 3D printing allows for freedom of design, waste minimization, and the ability to manufacture nanostructured complex structures [49,77].

Material deposition and vat polymerization are two of the main methods for 3D printing of MNs. Material deposition examples include fused deposition modeling (FDM) and material jetting (MJ), while stereolithography (SLA), digital light processing (DLP), continuous liquid interface production (CLIP), and two-photon polymerization (TPP) are vat polymerization techniques [4]. TPP is advantageous in that complex structures in the micro- and nano-scale can be fabricated, while SLA is characterized by a slow printing speed compared to DLP, which may influence the viability of biologicals [27]. These methods are further dictated by the characteristics and properties of the materials employed [78] and can vary from polymeric materials, such as polyvinyl alcohol (PVA), polylactic acid (PLA), polymethyl methacrylate (PMMA) and polyetheretherketone (PEEK) [4,79], biocompatible polymers [40,80,81], and medical-grade biocompatible resins, such as Castable Resin [82], IP-S photoresist [83,84] and Dental LT Clear [85]. The wide-ranging availability of materials, each with their own properties and functionality, when combined with the large variability in 3D printing technologies, therefore allows for a broad scope of 3D printed MNs to be fabricated based on its loaded cargo and the intended therapeutic application. Table 3 provides a list of 3D printing technologies used in the fabrication of MNs, the typical materials employed, and the advantages and drawbacks of the respective processes.

Table 3. A summary of typical technologies used in the fabrication of 3D printed MNs, the materials employed for MN fabrication and the advantages and limitations of the respective additive manufacturing processes (adapted from [86]).

3D Printing Technology	Material	Power Source	Layer Thickness	Advantages	Limitations
Fused deposition modeling (FDM)	Thermoplastic polymers, glass, metal	Heat	50–300 µm	Simplicity. High speed. Low cost (for thermoplastic materials). Ability to create complex, innovative, and customized dosage forms.	The high cost (for glass and metal). Weak mechanical properties (delamination due to temperature fluctuations). Limited material. Layer-by-layer finish. High temperature during the extrusion process (possible API degradation). Lack of biocompatible/Biodegradable printable polymers.
Stereolithography (SLA)	Liquid photopolymer	UV light	10–400 µm	Simplicity. Low printing costs. Fine spatial resolution. High quality (minimum mechanical anisotropy). Complex and customized drug delivery systems (<100 µm). Minimum drug decomposition.	Single material. Limited mechanical properties. Limited availability of biocompatible photopolymerizable polymers. Use of UV light to initiate polymerization (possible API degradation). Potential toxicity. Rinsing and the post-curing process is necessary.
Digital light processing (DLP)	Acrylates, Epoxides	UV light	25–100 µm	High resolution. High speed. Low cost. Less affected by oxygen inhibition than SLA. Low initial vat volume.	Limited mechanical properties. Toxicity. Need supports.
Liquid crystal display (LCD)	UV-curable Resins	UV light	50–100 µm	A smaller volume of resin is required. High resolution. Short curing time. Low cost.	Low precision.
Continuous liquid interface printing (CLIP)	UV-curable resins, acrylates	UV light	50–100 µm	High speed. High precision.	High cost. Low viscosity resin is needed. Probable toxicity.
Two-photon polymerization (2PP)	UV-curable resins, acrylates, ceramics	UV light	100 nm–5 µm	High spatial resolution. Scaling up. Low-cost materials.	Low yield of production. Low build speed. Limited material.
Selective laser sintering (SLS)	Thermoplastics, polymer, metal, and ceramics	Laser beam	20–150 µm	No need for support material. High resolution (30 µm) and precision. High quality. High speed. No post-curing required.	Limited mechanical properties. High cost. Slow printing. High printing temperature. Rough surface. Wastage of unsintered powder.
Direct metal laser sintering (DMLS)	Compact fine powder metals and alloys	Laser beam	20–100 µm	Fine accuracy and resolution. Good mechanical properties.	Support structures required. Protective atmosphere required.
Selective laser melting (SLM)	Metals, alloys	Laser beam	20–100 µm	No need for support material. Good mechanical properties.	High cost. Poor dimensional accuracy and quality.

Abbreviations: API, active pharmaceutical ingredient; UV, ultraviolet.

3. 3D MN-Facilitated biological delivery

The rapid advancement of drug delivery using MNs has resulted in the development of novel platforms loaded with biologicals for a variety of therapeutic applications such as for cancer treatment and wound healing. These devices incorporated the positive aspects of advanced drug delivery platforms with the mechanical support and innervation derived from MN arrays, combined with the variable release properties of polymeric materials. This trend was noted in studies by Wang and coworkers, who developed dissolving MNs for the delivery of anti-programmed death-1 (aPD1) protein-loaded nanoparticles for treatment of melanoma [87], Chang and coworkers [88], where cryomicroneedles via the micro molding of cryogenic medium with pre-suspended cells were synthesized, Su and coworkers [89], where polyvinylpyrrolidone MN patches loaded with W379 (sequence: RRRWWWWV) and the monoclonal antibody anti-PBP2a were designed and evaluated for the eradication of wound biofilms, and by Vander Straeten and team [90] who used a vaccine printer to print a thermostable MN containing COVID-19 mRNA vaccine composed of mRNA-loaded lipid nanoparticles into prefabricated molds.

Due to the positive results seen in these studies using conventional fabrication methods for the delivery of biologics, 3D printed MNs have exponentially expanded over the last few years for the delivery of labile pharmaceutical actives including proteins and peptides, with these platforms able to not only protect the active agents allowing for exertion of their required therapeutic effects but also to provide further diagnostic and drug delivery capabilities [4,27,91,92]. One such study that developed a 3D printed biological-loaded MN was by Economidou and coworkers [93], where 3D hollow MNs were combined with microelectromechanical systems for the transdermal delivery of insulin. This device was developed for patient-individualized treatment and was analyzed in vivo using Swiss albino female mice. The developed device consisted of a 49 (7 × 7 symmetrical) configuration array of cone-shaped Dental SG MNs (1000 μm base diameter, 100 μm tip diameter, 1000 μm height, and 800 μm internal bore diameter) and was manufactured using laser SLA, with both the ‘Bevel’ and ‘Ellipsis’ MNs fabricated demonstrating superior printability of the elliptical design (Figure 2). For the in vivo study, the number of the MNs was reduced to 4 in a 2 × 2 symmetrical configuration and exhibited an enhanced and sustained insulin reduction in the blood glucose levels when compared with the subcutaneous injection control (16.1 μIU/mL vs. 9 μIU/mL at 6 h after treatment). Notably, a high fracture strength (98.2 ± 4.6 N) was determined for the fabricated MNs in addition to effective skin penetration being demonstrated. Results from this study have shown that SLA can be an effective technique for the development of biocompatible hollow MNs.

Figure 2. CAD images of the ‘bevel’ and ‘ellipsis’ MN designs and respective cross-sections (a and c), with SEM images of the 3D printed ‘bevel’ and ‘ellipsis’ MNs (b and d) and of 4 MNs of the ‘ellipsis’ design featuring dimensions (e) (reproduced with permission from [94]).

The use of hollow 3D MNs for the delivery of insulin was further investigated by Li and coworkers [94], who developed fast customizable hollow MN patches, composed of a photosensitive resin, using static optical projection lithography (SOPL). The fabricated MN patches demonstrated in vitro biocompatibility in HaCaT and human dermal fibroblast cells. Furthermore, the MN patches exhibited a compression strength of about 4 N and an approximate compression distance of 250 μm, with further in vivo analysis in C57BL/6 mice noting that insulin delivered through the fabricated device maintained similar blood glucose levels when compared to the control group. Results indicated that by using SOPL-based 3D printing, quickly customized and high-quality hollow MN patches can be achieved. Xenikakis and coworkers [95] similarly endeavored to deliver insulin using hollow MNs fabricated via vat polymerization, which offers a relatively high resolution. In this study, however, curved pyramid (average minimum ferret diameter was ~0.6 mm) and syringe-like MNs (average minimum ferret diameter was 0.9 mm) were prepared from NextDent Ortho Rigid resin. Mechanical testing determined that the syringe-like hollow MNs were able to withstand forces up to 650 N, while the curved pyramid MN array fractured at 416.5 N. Permeation studies, using Franz diffusion cells, further determined a cumulative insulin permeation of 4.3 ± 1.1% and 6.0 ± 2.3% from the curved pyramid MNs and the syringe-like MNs, respectively after 60 minutes. It should be noted that in these studies, the hollow MNs were used as a conduit for the delivery of insulin, with the MNs themselves not loaded with the bioactive agent.

3D hollow MNs have additionally been investigated for other conditions, including for potential wound healing applications. In a study undertaken by Barnum and team [96], hollow MNs were prepared using a material jetting 3D printer for the spatial and temporal release of bioactives including Bovine Serum Albumin (BSA) and vascular endothelial growth factor (VEGF). The MNs were prepared from VeroClear resin, with flexible ribs between the MN islands composed of simulated rubber. The hollow MNs were subsequently filled with alginate and polyethylene glycol diacrylate (PEGDA) hydrogels containing the respective bioactives (Figure 3). The results of this study noted the successful incorporation of the bioactive agents, with variations in the hydrogel composition influencing the release rates of BSA. Additionally, the study determined that the MNs platform preserved the bioactivity of VEGF with scratch assays noting positive results for potential wound healing applications. The study further noted that the developed MNs were mechanically stable, with testing determining that the MNs could be easily removed if required and did not break despite exposure to extensive mechanical strain (up to 200 N). As with the research undertaken by Economidou et al. [93], Li et al. [94] and Xenikakis et al. [95], the hollow MNs developed in this study were used an effective conduit for the delivery of BSA and VEGF, however through the use of hydrogel matrices, it was determined that the loaded bioactives were not only protected during fabrication and administration, but further allowed for the controlled delivery of the bioactives through the modification of the hydrogel compositions.

Figure 3. Images of the miniaturized needle arrays for the enhanced bioavailability of drugs at the desired tissue depths with (a) depicting a schematic representation of the MNA-based dressing for the delivery of two different therapeutics targeting different depths or regions of the tissue to target separate cell populations, (b) a typical 3D-printed multilength needle array, in which the needles of different lengths were filled with hydrogels carrying different agents (shown with different colors), (c) the qualitative and (d) the quantitative distribution of different dyes in an agarose skin model released from the multilength MNA bandage with (e) the compliance of the semiflexible MNA on a sample of pig skin (reproduced with permission from [97]).

Further research has been undertaken by Derakhshandeh and coworkers [97], who endeavored to fulfill the need of multidrug delivery and customized therapy by developing a 3D printed wearable and programmable MNA bandage consisting of VEGF-loaded hollow MNs for the treatment of chronic diabetic wounds. The developed device consisted of Vero Clear hard resin MNAs (diameter of 9 mm), printed using an FDM 3D printer, a RPQ1 peristaltic micropump for drug administration and a wireless HC-05 Bluetooth communication module, which could be controlled through an in-house smartphone application. The results of this study determined that in vitro, the cumulative concentration of protein delivered was 70% after 180 min, while further studies using human umbilical vein cells to assess the angiogenic effectiveness of VEGF noted that the group receiving VEGF via the MNAs had a migration rate comparable to the positive control group, with a 100% scratch closure achieved in 4 hours. Additional in vivo studies using this device for the treatment of 5-day-old full thickness cutaneous wounds in diabetic mice determined that the average wound size in the MNA group decreased to 0.04 cm² with an average of 95% closure after 19 days, while ~55% and 40% closure was observed in the topical delivery and negative control groups, respectively. Furthermore, the device exhibited improved re-epithelialization, wound closure, and angiogenesis compared to topical VEGF delivery, with the results of this study proving that the fabrication process utilized maintained the viability of the loaded VEGF and allowed for topical delivery of the bioactive. The final resolution of the FDM approach is lower than that of vat polymerization method, and when comparing the results of the study by Xenakis and coworkers [95], skin penetration from the vat polymerization-fabricated hollow MNs occurred at 5 N for the curved pyramid MNs and at 10 N for the syringe-like MNs, whereas with Derakhshandeh and coworkers [97], full penetration through the pig skin using the FDM-prepared hollow MNs was only achieved with about 7 N.

The delivery of VEGF has also been investigated using self-healing 3D-printed hydrogel MNs based on eutectic Gallium-Indium (EGaIn) and spidroin (ESMNs) [98], with a photothermal response of the MNs conferred by EGaIn nanoparticles. In this study, the ESMNs were fabricated by printing the scaffold structures directly on polydimethylsiloxane microporous array templates using an extrusion 3D printer. Evaluation of the developed MNs VEGF on arm wounds in mice noted that the VEGF-loaded ESMNs exhibited a slightly better healing rate when compared to its control groups. Furthermore, accelerated growth of granulation tissue and collagen synthesis were observed in addition to reduced TNF-α and IL at the wound site. With the positive results seen in this study, it is apparent that the integration of nanoparticles into MN arrays allows for the controlled delivery of the bioactives, with protection afforded to the loaded cargo during the fabrication process. Also working with spidroin, 3D printed self-healing and photothermal-responsive MXene and spidroin MN scaffolds were developed by Shao and team using an Ecoflex microneedle mold [99]. Inspired by cacti, these MNs comprised of a hydrogel precursor of MXene, spidroin, polyurethane, and aloe vera gel. Results from the study demonstrated that the MN scaffold was capable of stable motion sensing and excellent biocompatibility using 3T3 cells. Furthermore, outstanding wound healing capabilities were demonstrated using human epidermal growth factor (hEGF), and mupirocin-encapsulated MNs, as per the presence of dense collagen deposition and significantly lower IL-6 and TNF-α levels. In both these studies hydrogel-based MNs printed into molds were effectively utilized for the delivery of the respective bioactives, noting that with varied hydrogel compositions combined with the appropriate additive manufacturing processes, MNs can be fabricated for varying biomedical applications.

A further focus has been made for the transdermal delivery of peptides using 3D printed MNs for cosmetic effects. One such molecule is the anti-wrinkle small peptide, Acetyl-hexapeptide 3 (AHP-3) which was delivered using 3D printed personalized MN patches (PMNP) [100]. In this study, Lim and coworkers focused on optimizing the PMNP MN geometry with varying curvatures. The results of this follow-up study noted that PMNP skin pre-treatment demonstrated a penetration depth of ∼750 μm, compared with the flat MN patches (∼480 μm) and non-treated skin (∼220 μm). Furthermore, enhanced cumulative permeation of AHP-3 was exhibited by the PMN across human cadaver skin compared with flat MN patches (∼45X higher). Expanding on this, this research team developed PMNPs of vinyl pyrrolidone (VP) and polyethylene glycol diacrylate (PEGDA) in a photopolymer resin of 7:3 (VP: PEGDA) ratio, using phenylbis (2,4,6-trimethylbenzoyl) phosphineoxide as a photoinitiator [101]. The MNs which had a 400 μm base diameter, 100 μm tip diameter, 800 μm height, and a 800 μm needle interspacing spacing totalling 927 MNs, were evaluated using human cadaver skin, with the PMNP (average weight of 2.668 g (±0.034 g) loaded with an average of ~13 mg AHP-3 peptide) shown to have over 98% intact MNs remaining post compression. Appreciable biocompatibility in human dermal fibroblast cell lines was also observed with >80% of the cells surviving after 24 hours of exposure. Notably, increasing the AHP-3 peptide concentration from 0.1% to 0.5% resulted in a general increase in overall drug release.

With the positive results noted for the delivery of VEGF and smaller peptides, research has also focused on the use of 3D printed MNs for vaccine delivery. In a study by Tran and coworkers [102] transdermal core – shell MNs (600 mm height, 400 mm core height, 300 mm base diameter, 200 mm core diameter) were manufactured for vaccine delivery. Three separate parts of the MNs – the shell, the cap, and the dry medication or vaccine core – were assembled together using a 3D printing method. The experimental setup was used to deliver the Prevnar-13 vaccine via a Poly(d,l-lactide-co-glycolide) shell, which modulated the observed delayed burst release. In another study using the CLIP technique, Caudill and coworkers [103] investigated OVA and CpG oligonucleotide-coated 3D MNs for vaccine self-application. The fabricated square pyramidal MNs in this study were noted to display enhanced retention and improved immune cell activations with the MN vaccine exhibiting a 20X higher OVA-specific IgG response after the prime immunization on day 21 and a 50X higher response on day 30 after the boost immunization when compared to the control groups. Additionally, dose sparing was also achieved by the MN vaccine formulation, with comparable levels of humoral responses seen with lower doses delivered through the fabricated MNs. CLIP is a faster and more advanced fabrication process, allowing for continuous output rather than layer-by-layer deposition, which may be a valuable 3D printing technique for labile and sensitive materials.

The use of 3D printed MNs have further been investigated for complex conditions, such as for chronic wounds, which are therapeutically difficult to treat due to surface scar formation, clots, and exudates that are present [104]. For these reasons, silk has been extensively explored over the years as a potential bioink for use in the fabrication of MNs [105]. In a study which explored the production of protein-based MNs, silk fibroin (SF) MNs were synthesized by Shin and Hyun [80] through one-step fabrication DLP 3D printing. The DLP printing was performed via curing of SF-riboflavin solutions, whereby the riboflavin served as an enzymatic photoinitiator, forming a dityrosine bond between radicals in the SF structure. The fabricated MNs were noted to be capable of skin penetration to a depth of 200 µm and was able to withstand ~300 mN of compressive force. Compared with other 3D printing techniques, DLP is capable of producing smooth surfaces with the rapid production of high-resolution MNs on a micrometer scale, which may allow for the preparation of consistent, painless MNs for the delivery of bioactives.

3D printed MNs have additionally been used for other biomedical applications including for sample extraction and due to its minimally invasive nature, can facilitate the development of point-of-care blood collection. This was highlighted in a study by Razzaghi and coworkers [106], where DLP was used to fabricate MNAs using PEGDA, tartrazine (photo-absorber), and lithium phenyl-2,4,6-trimethylbenzoylphosphinate as a photo-initiator. The MNAs had different cross-section shapes but were 800 µm in length and had a circular diameter of 300 µm. The MNAs were shown to effectively extract skin interstitial fluid and colorimetrically detect pH and glucose biomarkers.

The use of 3D printing technologies for the fabrication of biological-loading MNs has therefore been determined, through various studies, to be effective in protecting and delivering the loaded materials for the treatment of their respective ailments. With many of these studies utilizing polymeric materials traditionally utilized for drug delivery applications, as well as including bioactive-loaded nanoparticles within the printed bioinks, it has become apparent that the integration of additive manufacturing principles with drug delivery applications has the potential to develop accurate and effective treatment platforms for a variety of physiological conditions. Many of these platforms have been subsequently analyzed in clinical trials, as reported in Table 4.

Table 4. Clinical trials using MN systems for the delivery of biologics (adapted from [107]).

Study number	Study Title	Condition	Drug/Device	Phase	Actual Completion
NCT00837512	Insulin Delivery Using Microneedles in Type 1 Diabetes	Type 1 Diabetes	Device: Microneedle Device: Subcutaneous insulin catheter	Phase 2 Phase 3	07–2013
NCT00602914	A Pilot Study to Assess the Safety, PK and PD of Insulin Injected Via MicronJet or Conventional Needle (MicronJet)	Diabetes Mellitus Type II	Device: MicronJet Device: MicronJet	Early Phase 1	07–2009
NCT03274674	Use of Injectable-platelet-rich-fibrin (I-PRF) to Thicken Gingival Phenotype	Periodontoclasia Gingiva; Injury Blood clot Gingiva Disorder	I-PRF	N/A	01-06-2018
NCT02837094	Enhanced Epidermal Antigen Specific Immunotherapy Trial −1 (EE-ASI-1)	Type 1 Diabetes	Drug: C19-A3 GNP	Phase 1	31-12-2019
NCT02459938	Safety and Efficacy of ZP-Glucagon to Injectable Glucagon for Hypoglycemia	Hypoglycemia	Drug: Glucagon (ZP-Glucagon) Drug: Glucagon (GlucaGen)	Phase 1	10–2015
NCT00489918	Dose Ranging Study – Macroflux Parathyroid Hormone (PTH) in Postmenopausal Women With Osteoporosis	Osteoporosis	Drug: teriparatide Drug: teriparatide	Phase 2	08–2008
NCT02478879	A Study to Determine the Patient Preference Between Zosano Pharma Parathyroid Hormone (ZP-PTH) Patch and Forteo Pen	Postmenopausal Osteoporosis	Drug: ZP-PTH Drug: FORTEO	Phase 1	12–2015
NCT01674621	Phase 2 Study of BA058 (Abaloparatide) Transdermal Delivery in Postmenopausal Women With Osteoporosis	Postmenopausal Osteoporosis	Drug: Abaloparatide Transdermal (50 mcg) Drug: Abaloparatide Transdermal (100 mcg) Drug: Abaloparatide Transdermal (150 mcg) Drug: Abaloparatide Injection (80 mcg) Drug: Abaloparatide Placebo	Phase 2	02-08-2012
NCT03607903	Adalimumab Microneedles in Healthy Volunteers	Pain Injection site	Biological: Adalimumab ID Biological: Adalimumab SC Other: Saline ID Other: Saline SC	Phase 1 Phase 2	30-10-2018

4. 3D printed MNs for tissue engineering applications

From a cell delivery perspective, MNs have been effectively utilized for a few regenerative efforts, although these have not been produced through 3D printing fabrication methods. Such studies include the development of cardiac stromal cell-integrated [108] and cardiomyocyte-integrated MN patches for the treatment of myocardial infarction [109], and MN depots for the delivery of mesenchymal stem cells [110], while further studies include the development of double-layered adhesive MN bandages (DL-AMNBs) by Lim and coworkers [111] and microspheroids on MNAs for cartilage tissue engineering by Grogan and team [112]. This research has shown the undoubted potential of these platforms to not only protect but also control the delivery of these agents. With this in mind, further research, although limited, has also focused on utilizing 3D printed MNs for potential tissue engineering and regenerative medicine applications through the delivery of viable cell cargo.

This trend was exemplified in a study by Farias and coworkers [113] who used SLA for the fabrication of 3D printed hollow MNs (tip diameter = 400 µm, MN diameter = 1000 µm, MN height = 600 µm), loaded with shear-sensitive human hepatocellular carcinoma (HepG2) cells contained in atomized alginate capsules. While this study did not specifically deliver cellular cargo for specific tissue engineering or regenerative medicine applications, it did provide for a platform capable of protecting its cellular cargo during the 3D printed fabrication process with results showing no statistical difference in cell viability between the extruded material and its non-extruded controls. Additionally, by innovating the field of bioprinting, Reid and coworkers [114] printed human induced pluripotent stem cells (hiPSCs) into Geltex using a modified consumer-based 3D printer incorporating MNs as part of the bioprinting process. The results of this study also noted the successful bioprinting of the pluripotent hiPSCs, with positive TRA-1-81 staining observed, indicative of printing in differentiation conducive environments.

5. Microneedle regulatory policies and patents

The regulatory mandate of any agency/administration is to ensure timeous production of safe and effective medical devices, thereby promoting and protecting public health. A search generated on the European Patent and the United States Patent and Trademark Office databases yielded 20 and 13,071 records, respectively while MNs for vaccine, drug delivery, and bioactive compounds were searched on the World Intellectual Property Organization Espacenet Patent Search Database using ‘microneedles’ and ‘drug delivery’ as descriptors yielded 618 results (as of 2 April 2024). MNs are therefore still a relatively new pharmaceutical technology and thus requires standardized guidelines on MN manufacture, testing procedures, and evaluation both for regulatory approval and quality control. The Food and Drug Administration’s (FDA’s) Regulatory Considerations for Microneedling Devices, addresses criteria that MN platforms have to fulfill in order to be classified as a medical device [115]. Also illustrated in this document is the regulatory pathway to market MN devices for aesthetic use. The MN platform is considered a medical device if used ‘in the diagnosis of disease or other conditions, or in the cure, mitigation, treatment, or prevention of disease’ or ‘to affect the structure or any function of the body of man.’ Mention of ‘penetration into living layers of skin’ is also made in this FDA document. As a result, the length, sharpness, arrangement, and puncture rate are important MN parameters. In 2017, the FDA issued a draft guidance on the Technical Considerations for Additive Manufactured Medical Devices for the manufacture of 3D printing-based devices [116]. Although, whether the FDA will consider just the 3D printed product or the 3D-printer or both is still questionable.

6. Conclusion

Compared to the traditional manufacturing methods of MNs, the key advantages of 3D printing are improved efficacy, safety, and therapeutic drug delivery, cost-effectiveness, and simplified production. Therefore over the last decade, 3D MNs have been extensively evaluated for the delivery of pharmaceutical actives and have shown their versatility to not only protect the loaded bioactives, but spatially and temporally control their delivery. With the rapid advancement of technology and the introduction of additive manufacturing for the development of precise drug delivery systems, MN research has followed this trend allowing for the development of microscale platforms that can be modeled virtually and then fabricated accurately and consistently. This exponential increase in capabilities, however, in isolation often does not allow for the loading of labile biological materials, with the printing processes used being detrimental to the viability of the loaded cargo. The incorporation of biological material-loaded advanced drug delivery systems into the additive manufacturing process allows for the opportunity to protect and deliver these compounds with potential uses in complex treatments, such as for tissue engineering and regenerative medicine applications. While limited research has been undertaken to date in this field of study, the results that have been published point to potential effective, versatile fabrication processes that have the potential to overcome many of the shortcomings of current MN manufacturing as well as those experienced in 3D bioprinting. Through these advancements in pharmaceutical design, complex physiological conditions that were previously difficult to effectively treat can be engaged and overcome.

7. Expert opinion

The invention of MN platforms was a pioneering moment in the fabrication of drug delivery devices to overcome the common concerns associated with the transdermal administration of pharmaceutical agents, including needle phobias, injury at injection sites, and inconsistent and erratic dosing. The fabrication of these MNs further allowed for the site-specific delivery of the bioactives, minimizing side effects and decreasing therapeutic doses. These MNs varied in structural design for example, hollow MNs which may be regarded as miniaturized conventional hypodermic needles, allowing for easier manufacture and formulation, with solid MNs similarly not associated with drugs directly, thus simplifying device development and regulatory approval. Dissolvable and coated MNs significantly differ from these MN counterparts but do offer more potential benefits due to their ease-of-use patch -based format with dissolvable MNs particularly offering the unique advantage of not generating any medical waste.

With these wide MN architectural options, research into MNs for pharmaceutical applications subsequently increased exponentially with studies using a large variety of biocompatible MN materials (both novel and existing), incorporating various classes of drug molecules, which could be further controlled to allow for a required release profile. While predominantly focused on transdermal delivery, numerous studies began to utilize MNs for other therapeutic applications including for cardiovascular, central nervous, hepatobiliary, and musculoskeletal conditions. Applications in these physiological systems benefited from the precision and accuracy and ‘pain-free’ properties of MNs while also taking advantage of its structural integrity and biocompatibility in areas of the body that could be exposed to physiological strain. Clinical research into this area also noted successful results with numerous MN platforms undergoing clinical trials.

Additive manufacturing in recent years has also garnered significant attention in the drug delivery field, with researchers noting the benefits of its rapid, precise, accurate, and consistent fabrication processes. These processes additionally allowed for the virtual modeling of the designs prior to fabrication, allowing for the development of complex, personalized structures which were not possible using traditional molding processes. The use of 3D printing in pharmaceutical design subsequently became more widespread with dosage forms including tablets, capsules, implants, and patches being developed. It was through this advancement that bioprinting, the incorporation of biologics such as cellular material within the printing materials, became possible. Numerous studies have since been undertaken utilizing 3D printing for the biofabrication of complex architectures, including platforms for tissue replacement and regeneration (3D bioprinting). It was further through the advancement of additive manufacturing that the incorporation of blood vessel and nerve innervation structures became possible with 3D printing resolutions reaching micro-scale, while also being able to print multiple substrates concurrently on the same platform to reproduce some of the complex architecture of physiological tissues. It should, however, be noted that some of the serious challenges with currently experienced with 3D (bio)printing are its slow print speeds, limited material selection, and poor biocompatibility, with resolution still remaining a challenge for more complex physiological environments despite the large amount of research being undertaken in this area. Additionally, of the materials available for 3D printing, many exhibit limited bioavailability and/or poor mechanical properties, ultimately warranting future studies on the development of safer materials and optimal printing techniques without compromising MN resolution. Particularly with SLA, the materials used must fulfill the requirement of being both light-sensitive and biocompatible with the initiator, absorber, and rinsing materials required in this method posing a toxicity risk to the loaded biological material and the application site.

With reference to the scientific journal and patent literature available, a trend exists attesting to the global interest surrounding 3D MN technology. Because MNs are capable of delivering vaccines, biologics and biomolecules, diagnosing and monitoring diseases, and sample extraction, commercialization of MN systems will offer substantial benefits for patients. MN pentype and syringe-type systems have been approved for commercial use, but their development appears to be at a standstill. There are still no products on the market for MNs other than vaccine delivery systems, and some companies seem to have stopped developing them. For effective product evaluation and regulatory approval, regulatory agencies are working to understand material properties, 3D geometric design attributes, and 3D printing process parameters on the performance of 3D printed devices. Although there are currently no regulatory requirements defined for MN-based products, this issue will have to be addressed in the near future when companies intending to commercialize MN patches apply for marketing authorization, with the lack of any structured regulatory guidelines potentially delaying their introduction into the market.

Regarding the use of 3D printed MNs for tissue engineering applications, to date, very few studies have focused on this area of research, with many current 3D printing processes limiting their use in bioprinting due to the use of high temperatures or UV light, which can affect the viability of any loaded cell-cargos. Furthermore, extrusion-based technologies may physically damage biological cargo during printing. It however, should be noted that the study undertaken by Farias and coworkers [113] has shown that the potential to develop 3D bioprinted MNs still exists with the results of this study highlighting the important benefits of these MN devices. It should be further acknowledged that numerous pharmaceutical platforms, such as coated nanoparticles, exist for the protection of loaded actives against the potentially damaging 3D printing processes, with these platforms still to be fully investigated for their application in 3D bioprinted MNs. There is additionally the potential of utilizing tissue-specific cells within these advanced pharmaceutical platforms to allow for more accurate tissue engineering and regenerative medicine applications. Nevertheless, with the constant advancement in additive manufacturing technologies, as well as in advanced drug delivery design, the use of 3D printed MN architectures merging these two research fields can be successfully incorporated into tissue engineering and regenerative medicine interventions, with the ultimate goal of fabricating intricate platforms for complex physiological applications.

Article highlights

• The various 3D printing techniques along with their advantages and limitations in the fabrication of MNs and MN arrays (MNAs) have been provided.

• The factors affecting MN manufacture have been identified.

• Conventional and 3D printed MNs have been discussed along with their applications in biological delivery.

• 3D biofabricated microneedles (MNs) which combine drug delivery strategies with biological cargo have been discussed for their potential biomedical applications including tissue engineering and regenerative medicine.

• MN clinical trials, regulatory policies, and available patents have been highlighted and discussed.

Declaration of interest

The authors have no relevant affiliations or financial involvement with any organization or entity with a financial interest in or financial conflict with the subject matter or materials discussed in the manuscript. This includes employment, consultancies, honoraria, stock ownership or options, expert testimony, grants or patents received or pending, or royalties.

Reviewer disclosures

Peer reviewers on this manuscript have no relevant financial or other relationships to disclose.

Additional information

Funding

This research was supported by the National Research Foundation (NRF) of South Africa SARChI grant [UID: 64814].