Influence of material format and surface chemistry for the sustained delivery and efficacy of silk drug delivery systems in vivo

Abstract

Silk fibroin materials are promising for use in controlled drug delivery in the field of tissue engineering and biomedical applications thanks to silk's generally established biocompatibility and tunable properties for implants and drug storage. Several factors must be considered in the materials design, including material format, drug properties and release kinetics, and the activity and stability of the drug after release. While numerous reviews described silk-based DDS that demonstrated controllable in vitro release, success in vivo has been limited, especially in some material formats. This review therefore aims to provide insight into the current material format and functionalization strategies to maximize in vivo performance by describing the in vivo activity of recently developed silk drug delivery systems. The review also aims to provide a fresh perspective on the suitable format and functionalization strategies for a target biomedical application. Based on the release behavior of drugs in various material formats, silk films, foams, and microneedles were better suited to serve as scaffolds for cell regeneration and improved recovery rate for biomedical applications involving wound healing and tissue engineering. Gels and particles could be incorporated within the films and foams but the purpose would be to serve as additional physical barriers towards drug diffusion in these types of application. For drugs or therapeutics that target internal organs (i.e. brain, liver, intestines, etc.), gels and particles were mainly used due to their size. In the event that the material format selection based on the target application does not contribute a lot to the prolonged release of drugs or therapeutic agents, hybrid functionalization strategies were adapted to make the surface chemistry of the material more responsive to the environmental stimuli for a more tunable silk DDS.

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Christine Jurene O. Bacal

Dr Christine Jurene Bacal is a versatile researcher with a strong background in chemistry and materials science. During her PhD, she fabricated a biocompatible and highly selective filtration membrane for paracetamol capture with potential application in overdosed patients. Currently, as a postdoctoral researcher at Ear Science Institute Australia, she is functionalizing drug candidates and material scaffolds for sustained drug delivery in the middle ear of children with otitis media. Dr Bacal aims to contribute to the faster translation of interdisciplinary research projects by using her knowledge and skills in integrated research fields.

Benjamin J. Allardyce

Dr Ben Allardyce works across the fields of biochemistry and materials science to develop silk-based materials. His research aims to understand silk protein properties and their interactions to design new methods to disassemble silk and re-assemble it into high performance materials, including fibers, aerogels and membranes.

Filippo Valente

Dr Filippo Valente has 12 years of experience in biomedical and materials science, focusing on hearing applications. He leads the Biomaterials Research team at the Ear Science Institute Australia and holds an adjunct position at the University of Western Australia and Curtin University. In the last 5 years, Dr Valente has successfully raised $1.9M through grants and industry investment, managed product development, and built industry partnerships locally and internationally. Since 2020, Dr Valente has led the product development of ClearDrum®, a prosthetic eardrum repair device. He also plays a key role in the regulatory strategy and clinical trial preparation.

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1. Introduction

Drug delivery systems (DDS) are devices or formulations that carry drugs or any therapeutic agents towards a certain target, which could be a pathogen, a cell, or a gene, for the improvement of the therapeutic efficiency, to improve specificity of the drug towards its target.^1,2 DDS are designed based on both the pharmacokinetics (PK) and biodistribution (BD) of the drug and the environment where the therapeutic agent will be distributed.^1,2 Depending on the desired application, DDS can consist of nanoparticles³ and liposomes,⁴ films,⁵ foams,⁶ hydrogels,⁷ injectables,⁸ inhalants,⁹ and even microneedle patches.¹⁰ Single dose DDS are useful to improve patient compliance in instances where delivery is required over prolonged periods.¹¹ In addition, DDS have the potential to reduce the side effects from drug administration by providing local and targeted delivery at therapeutic doses, reducing the amount of systemic drugs that need to be administered over time.¹² Despite their versatility and the tunability, development of DDSs that are highly resorbable and capable of mitigating undesired immunological responses such as hypersensitivity reactions^13–15 to prevent invasive removal of non-resorbable DDS^16,17 is yet to be achieved.

Some biodegradable and resorbable materials approved by the United States Food and Drug Administration for clinical use include synthetic polymers such as poly(glycolic acid) (PGA), poly(lactic acid) (PLA), poly(lactic-co-glycolic acid) (PLGA), poly(caprolactone) (PCL), and poly(dioxanone) (PDS)^18–20 and a natural polymer like collagen.²¹ However, these polymers still trigger inflammation,^22–26 degradation of bone tissue (osteolysis) in bone implants,^27,28 and reduced mechanical strength.^11,29–31 Therefore, natural biopolymers such as chitosan, alginate, and silk are being increasingly explored.³²

Among the natural biopolymers, silk has been extensively explored as a DDS because it was known to preserve the activity and stability of drugs and other therapeutic agents like temperature-sensitive vaccines³³ and labile antibiotics.³⁴ The inherent structure of silk resembles amphiphilic block copolymers which enables controllability of drug release via external stimuli (i.e. pH) owing to the changes in bond conformations and surface charges on the silk backbone.³⁵ In addition, by altering the degree of physical crosslinking through a range of established methods like alcohol treatment³⁶ or water vapor annealing,³⁷ silk could be processed in various formats making it a versatile DDS.

Despite the advantages, silk's inherent amino acid sequence gives it a negative charge at neutral pH. As a result, compounds with a higher isoelectric point are typically released rapidly,^38,39 which could be unfavorable when sustained release of these compounds is preferred. It is therefore necessary to explore a material configuration and functionalization method that will enable prolonged and controllable release of wider range of drugs and therapeutic agents in silk DDS.

In this review, the challenges of developing silk-based DDS that can deliver drugs with controlled release while ensuring the drug functionality were evaluated. Various functionalization strategies were presented to correlate how modifications in chemical functional groups affected the release profiles and other clinically-relevant silk properties. Finally, emerging smart (“stimuli-responsive”) and functional materials were evaluated to provide an insight on the novel materials that can be incorporated with various silk formats. The use of smart materials to improve surface chemistry were assessed on how they influenced the controllability of drug release.

2. Silk functionalization strategies for drug/therapeutic agent delivery

In designing a silk-based DDS, chemical/physical interactions between the drug and silk must be established based on the specific therapeutic agent to be delivered. In this section, techniques to incorporate active agents into silk were discussed as well as the available functionalization methods utilized to control binding and release.

2.1. Conversion of silk fibers to silk solution

Silk fibers must be converted to silk solution prior to fabricating various silk drug delivery systems.⁴⁰ The raw silk from the cocoons must be treated before removing the sericin. Sericin is the protein “glue” that binds silk filaments together into the cocoon.⁴¹ Silk sericin removal is called degumming and various reagents/solvents can be used such as sodium carbonate,⁴² calcium chloride/ethanol solution,⁴³ sodium oleate,⁴⁴ enzymes,⁴⁵ and ionic liquid⁴⁶ to boil the cocoons during pre-treatment. The selection of one or multiple solvents depends on the desired application as some could influence the resulting tensile properties.⁴⁷ The degummed silk fibers are further treated with a chaotropic agent (e.g. lithium bromide,⁴⁸ formic acid,⁴⁹ lanthanide chlorides,⁵⁰ or hexafluoroisopropanol⁵¹) to disrupt hydrogen bonding within the silk structure and dissolve silk resulting in processability of the solution into various forms (i.e. films, hydrogels, foams, etc.). Finally, the dissolved silk fibers are dialyzed against pure water to remove excess salts and centrifuged, resulting in an aqueous solution of pure silk fibroin that can be used to fabricate various silk material formats for drug delivery.

2.2. Silk functionalization strategies in various silk drug delivery systems

Silk drug delivery systems, in any type of format, could be functionalized to control the release of the encapsulated therapeutic agent. This could be performed by tapping into the noncovalent interactions between silk and therapeutic agents or via chemical crosslinking.

2.2.1. Functionalization using noncovalent interactions between silk and therapeutic agents. One method of functionalizing a silk material is via physical interaction. Physical interactions between the silk and the therapeutic agent involve inter- and intramolecular forces of attraction (i.e. electrostatic interactions, hydrogen bonding, van der Waals forces) with no formation of new chemical bonds.

There are several different strategies to incorporate silk with therapeutic agent physically but this could be classified into two modes: by directly incorporating drugs within the silk solution and by coating physically. Strategies to directly incorporate drugs within the silk solution include embedding a drug/therapeutic agent within pure silk (bulk blending), using silk composite materials or by using interfacing or encapsulating agents. Bulk blending is by far, the simplest physical approach to sustain release in silk drug delivery systems (Fig. 1(a)). In this method, the drug is homogeneously blended with the silk solution prior to immobilization into a solid scaffold (i.e. casting to a film or lyophilizing to a foam) or direct administration to target site. Mixing drugs or therapeutic agents to silk blended with another polymer to form silk composite materials is another strategy to physically incorporate drugs or therapeutic agents.⁵² In this process, silk is mixed with another polymer to form a composite material with a different property than the original components (silk composite drug delivery systems) (Fig. 1(b)). Therapeutics are often internalized in this technique by adding them into a mixture of silk with another polymer. The drug is then physically adsorbed via electrostatic interaction or hydrogen bonding. Another strategy is to use interfacing or encapsulating agents such as an amphiphilic polymer, a surfactant, or a phospholipid before blending with silk solution (Fig. 1(c)). This technique is commonly applied in silk particles or hydrogels and has enabled slow release of drugs to facilitate sustained delivery over extended periods.^52–54 In this same technique, the interaction between silk and the encapsulating agent prevents drug leaching due to the indirect interaction between the drug and the release bath.

Fig. 1 Functionalization methods via noncovalent interactions. (A) Bulk blending of silk and therapeutic agents. SNF (silk nanofibers), PTX (paclitaxel drug), DOX (doxorubicin drug) (reprinted with permission from Xiao, et al. ©2024 Wiley-VCH GmbH)⁵⁵ (B) use of silk composite materials. PBS/EtOH (phosphate buffered saline/ethanol solvent system), SF (silk fibroin), PDLLA (poly(D,L-lactic acid)) (reprinted with permission from Deng, et al. ©2024 American Chemical Society)⁵⁶ (C) use of encapsulating agents. (Reprinted with permission from Mao, et al. ©2017, Published by Elsevier B.V.).⁵³

Meanwhile, coating (Fig. 2) is another functionalization technique which involved covering the silk biomaterial with a layer containing the encapsulated drug or another polymer containing the drug to slow down the release of the therapeutic agent.⁵⁷ Strategies to coat the silk material physically are via dip coating to form a layer of surface coating or via the layer-by-layer (LbL) approach to form a multilayer assembly in silk drug delivery systems. Dip coating involves (Fig. 2(a)) silk substrate immersion into a solution that could form a surface barrier via electrostatic interactions which surrounds the silk drug delivery system and prevent the burst release of therapeutic agents. This solution could be pure silk itself, another polymer (forming silk composite coating), or a surfactant. Dip coating can also be applied repeatedly onto the same substrate to form a multilayered assembly in silk drug delivery systems. This technique is known as layer-by-layer (LbL) approach and manifested a more prolonged release behavior in most silk drug delivery systems developed compared to a single-layered coating (Fig. 2(b)).

Fig. 2 Physical coating strategies for silk drug delivery systems. (A) Dip coating to form a single surface layer. (B) Layer-by-layer coating approach.

Based on early in vitro studies, it was hypothesized that merely applying bulk blending is insufficient to prolong the release of therapeutic agents for treatment of various diseases.^58–60 However, recent in vivo studies have shown the efficacy of bulk blending in prolonging drug release in various animal models (Table 1, entries 1–8). Bulk blending was applied to develop silk drug delivery systems for combination chemotherapy,⁵⁵ as vaccines for COVID-19,⁶¹ for tissue regeneration,⁶² for circadian rhythm regulation,⁶³ for wound healing,^64,65 for acute liver failure treatment,⁶⁶ and for treatment of sensorineural hearing loss⁶⁷ primarily since it is a facile approach and does not involve a lot of other materials or polymers aside from the therapeutic agent and silk solution minimizing the components that could counteract each other. For instance, when both doxorubicin and paclitaxel were incorporated in silk nanofiber hydrogels and administered intratumorally, tumor growth in female nude mice was inhibited for over 28 days (Table 1, entry 1).⁵⁵ When receptor-binding domain (RBD) linked with fragment crystallizable (Fc) region of human immunoglobulin G1 (IgG1) (RBD-Fc) was directly incorporated with silk fibroin solution to form a silk fibroin hydrogel-based vaccine for COVID-19, it was found that the formulations had 1-1.3-fold increased immunogenicity in female ICR mice after 90 days since intramuscular injection compared with the silk hydrogel control without the therapeutic agent (Table 1, entry 2).⁶¹ However, burst release can still be observed upon transdermal administration of melatonin-containing silk fibroin microneedles for circadian rhythm regulation in Sprague–Dawley rats (Table 1, entry 4)⁶³ as this has immediately induced sleep in the rats. Hence, other functionalization techniques can be explored to apply for this specific silk drug delivery system to prolong drug release.

Table 1 In vivo activity of silk drug delivery systems

Entry no.	Silk drug delivery system	Drug/therapeutic agent carried	Application	Functionalization technique used	Material format	Delivery mode	Model organism	In vivo efficacy	Ref.
*Silk solutions were taken as fibroin solution unless explicitly stated to be sericin.
1	Silk nanofibers	Doxorubicin and paclitaxel	Combination chemotherapy	Bulk blending	Hydrogel	In situ	Female nude mice	Inhibition of tumor growth over 28 days	55
2	Silk fibroin-based hydrogel	Receptor-binding domain (RBD) linked with fragment crystallizable (Fc) region of human immunoglobulin G1 (IgG1) (RBD-Fc vaccine)	Vaccine for COVID-19	Bulk blending	Hydrogel	Intramuscular injection	Female ICR mice	SF hydrogel-based vaccine formulations had 1–1.3-fold increase immunogenicity after 90 days	61
3	Silk fibroin microspheres	Glucose	Tissue regeneration	Bulk blending	Particle (microspheres)	Subcutaneous implantation	Sprague–Dawley rats	Large amount of extracellular matrix distributed peripherally throughout the silk fibroin microspheres after 2-weeks	62
4	Silk fibroin microneedles	Melatonin	Circadian rhythm regulation	Bulk blending	Microneedles	Transdermal	Sprague–Dawley rats	Burst release upon administration as it induced sleep immediately, then melatonin concentration peaked after ∼1 h but sustaining release after 8 h	63
5	Omentum-extracellular matrix/silk fibroin hydroscaffold	Decellularized extracellular matrix	Treatment of nonhealing diabetic wounds	Bulk blending	Hydrogel/hydrogel scaffold hybrid	In situ	Wistar rats	Wounds treated with hydrogels and hydroscaffolds were completely closed after 28 days, untreated wounds showed delayed closure and took another 3–5 days	64
6	Silk fibroin nanofibers	Quercetin containing morpholine-pyridine motifs (QFM)	Wound healing	Bulk blending	Mat (membrane)	Transdermal	Wistar rats	100% wound closure after 9 days	65
7	Silk fibroin electrospun scaffolds	Hepatocyte growth factor (HGF)/fibroblast growth factor-4 (FGF-4)	Acute liver failure treatment	Bulk blending	Scaffolds (membrane)	Implanted	BALB/C mice	Growth factor incorporated scaffolds showed greatest reduction in the biomarkers for hepatocyte damage after 14 days	66
8	Silk fibroin-based hydrogels	Dexamethasone and lipoic acid	Treatment of sensorineural hearing loss	Bulk blending	Hydrogel	Intratympanic injection	Guinea pigs	Dexamethasone undetected after 5 days in the perilymph after injecting dexamethasone loaded silk fibroin hydrogels; had no superior advantages than microcrystals	67
9	Silk/poly(D,L-lactic acid) microspheres	Doxorubicin hydrochloride	Anticancer	Using silk composite materials	Particle (microsphere)	In situ	BALB/C nude mice	Tumor volume is minimal after 15 days	56
10	Silk/poly(vinyl alcohol) microparticles	Metformin	Corneal neovascularization	Using silk composite materials	Particle (microparticle)	In situ	Rats	Reduced neovascularization and inflammation after 7 days	68
11	Silk fibroin/collagen nanofiber	Combination of silver (Ag) and gold (Au) nanoparticles	Wound healing	Using silk composite materials	Scaffolds (membrane)	Implanted	Sprague-Dawley rats	Healing percentage of ∼99% over the course of 21 days	69
12	Silk fibroin nanoparticle hydrogel	NK1R antagonist (CP-99,994)	Dry eye disease treatment	Using silk composite materials	Hydrogel	Topical administration	C57BL/6 mice	Significant recovery of Treg cell function and nanoparticle size retained after 14 days of treatment	70
13	Silk fibroin/chitosan-based scaffolds	Curcumin and 5-aminosalicylic acid	Anti-inflammatory and healing function for anal fistula	Using silk composite materials	Scaffold (membrane)	Subcutaneous implantation	Rats	Sustainable drug release above 400 h, better biocompatibility and anti-inflammatory effect than control groups	71
14	Silk nanofiber hydrogels	Asiaticoside and magnesium ions	Wound healing	Use of silk composite materials	Hydrogel	In situ	Sprague–Dawley rats	Above 60% of wounds exhibited better healing within 7 days compared to other hydrogel groups	72
15	Silk fibroin composite nanofiber membrane	Flavonoids	Wound healing	Use of silk composite materials	Membrane	Transdermal	Sprague–Dawley rats	After day 15, ∼93% and ∼97% wound shrinkage can be observed with modified groups compared to the ∼72% of the control group	73
16	Silk fibroin/alginate 3D-printed scaffold	Placental extracellular matrix	Wound healing	Use of silk composite materials	3D printed scaffolds (hydrogel scaffolds)	Transdermal	C57 mice	Remarkable increase in the expression of pro-angiogenesis genes in the wound bed treated with the 3D-printed scaffold with extracellular matrix at both day 7 and day 21 post-surgery	74
17	Electrospun silk fibroin/poly(hydroxybutyrate-co-valerate) (PHBV) dressing	Berberine	Diabetic wound healing	Use of silk composite materials	Nanofibrous dressing (membrane)	Transdermal	Mice	Wound healed completely at a significantly higher rate (∼99.7%) than the control group (∼83–85%) after 18 days of treatment	75
18	Silk fibroin/poly(caprolactone) nanofibrous mat	Berberine	Diabetic alveolar bone regeneration	Use of silk composite materials	Nanofibrous mat (membrane)	Topical	Sprague–Dawley rats	Most regenerated bone formation in type 2 diabetes milletus rats compared to the control groups after 2–6 weeks	76
19	Cell-free chitosan/silk fibroin scaffolds	Bioactive glass	Regeneration of critical-size bone defects	Use of silk composite materials	Scaffold (foams)	In situ	Sprague–Dawley rats	Bone volume fraction was higher (∼46%) compared to the control (∼42%) at 12 weeks of treatment	77
20	Silk sericin hydrogel	Resveratrol	Wound healing	Use of silk composite materials	Hydrogel	Topical	Wistar albino rats	Maximum wound healing achieved at the end of day 14 compared to the toxic group	78
21	Silk fibroin mesh	Honey extract from bees	Wound healing	Dip coating in a mixture of poly(hydroxybutyrate-co-hydroxyvalerate) (PHBV) and poly(lactic acid) (PLA) or poly(caprolactone) (PCL)	Mesh (membrane)	Implanted	Sprague Dawley rats	∼45% and ∼62% area deposition of extracellular matrix and collagen, respectively from ∼24% and ∼20% in the control PHBV mesh up to 28 days	79
22	Silk fibroin shell	siRNA	Antitumor	Layer-by-layer assembly	Particle	Subcutaneous injection	BALB/C nude mice	The treated group yielded the smallest tumor sizes (0.3 g) and highest tumor inhibition rate (56%) than the control group after 15 days	80
23	Dissolving microneedle system	Mupirocin and Ag nanoparticles	Treatment of delayed wound healing	Bulk blending/using silk composite materials	Microneedles	Transdermal	BALB/C mice	Wound contraction rate increases to ∼78%, significantly higher than control on day 11	81
24	Silk fibroin and hydroxypropyl cellulose composite injectable hydrogel	Folate-modified AC16 cell-derived extracellular vesicles	Myocardial infarction repair	Use of silk composite materials/use of encapsulating agents	Hydrogel	In situ	Sprague–Dawley rats	Best cardiac repair result with left ventricular ejection fraction of ∼71% compared to the myocardial infarction control group with ∼43% after 28 days of treatment	82
25	Silk fibroin coating	Tumor targeting peptide	Cancer therapy	Layer-by-layer assembly/use of encapsulating agents	Particle (nanorods)	Intravenous injection	BALB/C nude mice	Decrease in tumor size after 14 days of irradiation in the treated group compared to the control	83
26	Porous chitosan–alginate hydrogel/electrospun PCL-silk sericin hydrogel scaffold	10-hydroxydecanoic acid (queen bee acid)	Wound healing	Using silk composite materials/chemical crosslinking with glutaraldehyde for the porous chitosan–alginate layer	Hydrogel scaffold hybrid	Transdermal	Wistar rats	Fibrosis is less prominent with 10-HDA-containing wound dressing than controls after 14 days	84
27	Poly(vinyl alcohol)/silk fibroin nanofibrous membrane	Astaxanthin and silver nanoparticles	Wound dressing	Use of silk composite materials/crosslinking using glutaraldehyde	Membrane	Transdermal	Sprague–Dawley rats	Wound area reduced to ∼12% and ∼4% compared with the ∼22% of the negative control after 14 days	85
28	Self-healable and 3D-printable silk fibroin nanocomposite hydrogels	Oxidized salep (OS) and κ-carrageenan nanoparticles	Wound care	Using silk composite materials/Schiff base crosslinking between OS and κ-carrageenan nanoparticles	Hydrogel	Subcutaneous implantation	Wistar albino rats	All animals survived without any malignancy, infection, or abscess formation. No necrosis or muscle degeneration after 14 days	86
29	Conductive hydrogels based on tragacanth and silk fibroin	Vancomycin	Burn wound healing	Blending with another stimuli-responsive polymers/crosslinking using dopamine and to form a Schiff base/crosslinking using EDC	Hydrogels	In situ	Wistar rats	Wound treated with hydrogel dressings had nearly entirely healed with >60% of the wound area showing signs of recovery after 14 days	87
30	Silk fibroin/poly(vinyl alcohol)/ROS-scavenging dendrimer hydrogels	Gallic acid	Wound healing	Using silk composite materials/use encapsulating agents (dendrimers)/amide coupling reaction between gallic acid and the dendrimer using EDC/NHS	Hydrogel	In situ	Male BALB/C mice	Wound size of ∼30% in the dendrimer-modified silk compared to ∼56% in the control group after 10 days over 16-day treatment course	88
31	Silk fibroin/gelatin films	Gelatin	Regeneration of tympanic membrane perforation	Chemical crosslinking with genipin	Film	Surgical implantation	Sprague–Dawley rats	Tympanic membrane regeneration was 60% compared to 40% in control after 5 days	89
32	Silk protein hybrid hydrogel	Teicoplanin-decorated reduced graphene oxide	Diabetic wound healing	Chemical crosslinking with genipin	Hydrogel	Topical	Sprague–Dawley rats/BALB/C mice	Accelerated healing of noninfected wounds in diabetic rat and infected wounds in a diabetic mouse after 14 days of treatment	90
33	Silk fibroin modified decellularized porcine liver scaffolds	Gelatin	Tissue regeneration	Bulk blending/crosslinking using citric acid	Scaffolds (membrane)	Subcutaneous implantation	Wistar rats	Collagen remodeling within the 90-day time point	91
34	Silk fibroin/soy protein isolate hydrogels	Quercetin	Burn wound healing	Use of silk composite materials/chemical crosslinking between silk and soy protein isolate using epichlorohydrin	Hydrogel	In situ	Kunming mice	Wound closure rates were ∼79% for the hydrogel-treated group whereas it was only ∼48% for the control group after day 17	92
35	Dual-network DNA-silk fibroin hydrogels	Bone marrow mesenchymal stem cells	Cartilage repair	Physical crosslinking via DNA base pairing/crosslinking via horseradish peroxidase (HRP)-mediated enzymatic reactions	Hydrogel	Implanted	Sprague–Dawley rats	Significantly accelerated cartilage regeneration after 14 days	93
36	Silk-based photocurable hydrogel	Platelet-rich plasma lysate	Meniscus treatment	Methacrylation via crosslinking (grafting) using lithium phenyl-2,4,6-trimethylbenzoylphosphinate (LAP) photoinitiator and poly(ethylene glycol diacrylate) (PEGDA) crosslinker	Hydrogel	In situ	New Zealand white rabbits	Treatment of full-thickness tears in rabbit menisci compared to the untreated counterparts after 8 weeks	94
37	Silk nanofibril/hyaluronic acid scaffold	Basic fibroblast growth factor	Spinal cord repair	Surface grafting of silk nanofibril/hyaluronic acid blends	Scaffold (membrane)	Implanted	Sprague–Dawley rats	Recovery of motor function 10 weeks after surgery compared to the control	95
38	Silk hydrogels	Adipose stem cells	Bone/wound healing	Photocrosslinking using riboflavin and hydrogen peroxide	Hydrogels (can be extended to microneedles, microcarriers, and bone screws)	Subcutaneous implantation	Sprague–Dawley rats	Bone screws did not cause serious inflammation and maintained bone screw shape after 14 days	96
39	Silk fibroin/hyaluronic acid hydrogel microneedles	Insulin	Diabetes management	Blending with another stimuli-responsive polymer (methacrylated hyaluronic acid)/photocrosslinking using riboflavin/layer-by-layer cumulative solution process	Hydrogel microneedles	Transdermal	C57BL/6 mice	Gentle regulation and prolonged maintenance of healthy blood glucose levels for close to 6 h compared to the blank group	97
40	Silk fibroin microgel scaffolds	Fluorescently tagged bovine serum albumin (BSA)	Tissue engineering and regenerative medicine	Photocrosslinking using tris(2,2′-bipyridyl)dichlororuthenium(II) hexahydrate (Ru) and sodium persulfate (SPS)	Hydrogels (microgel suspensions and scaffolds)	Subcutaneous implantation	C57/BL6 mice	Promoted cell and tissue ingrowth after day 10	98
41	Dual-phase silk fibroin methacryloyl (SilMA) (hydrogel scaffold) (DPSH)	Neurotrophin-3 (NT-3) and angiotensin (1–7) (Ang-(1–7))	Spinal cord repair	Use of encapsulating agents/UV crosslinking for silk fibroin methacryloyl	Hydrogel scaffolds (membrane)	Subcutaneous implantation	Mice	DPSH group promotes recovery of motor function following spinal cord injury after 8 weeks post implantation compared to the control groups	99
42	Silk fibroin hydrogels	Chondroitin sulfate methacrylate	Bone screw	Use of silk composite materials/UV crosslinking using methacrylic anhydride and poly(ethylene glycol diacrylate) (PEGDA)	Hydrogel	Implanted	White rabbit	Bone screws successfully implanted into the femur with appropriate match to the implantation site after day 7	100
43	Silk fibroin/chitosan interpenetrating network hydrogel with microspheres	Methylprednisolone acetate (MPA)	Cartilage regeneration	Using silk composite materials/crosslinking using β-glycerophosphate and sodium dodecyl sulfate	Microsphere encapsulated in a hydrogel	In situ (intra-articular injection of precursor suspensions)	White rabbits	Improvement in locomotor function in treated groups compared to control groups after 12 weeks	101

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Silk composite materials have been increasingly used for delivering therapeutics for cancer treatment,⁵⁶ corneal vascularization⁶⁸ and dry eye disease treatment,⁷⁰ wound healing,^{69,72–75,78} healing of anal fistula,⁷¹ diabetic alveolar bone regeneration⁷⁶ and regeneration of critical-size bone defects (Table 1, entries 9–20).⁷⁷ An in vivo study showed that by incorporating bioactive glass into a composite scaffold made from cell-free chitosan and silk fibroin, bone volume fraction of Sprague–Dawley rats was higher (∼46%) compared to the control scaffold (without bioactive glass) (∼42%) after 12 weeks of treatment (Table 1, entry 19). Thus, cell-free chitosan/silk fibroin composite has the potential to be an effective material to regenerate critical-size bone defects. In other studies though, the use of silk composite DDSs such as silk/poly(D,L-lactic acid) microspheres⁶⁸ and silk nanofiber hydrogels⁷² were able to sustain the release of the effective concentration of therapeutic agents in vivo for a week in rats (Table 1, entries 10, 14–15). Hence, to utilize these materials for applications that require longer release of therapeutics, there must be careful consideration on the chemical properties of the incorporated components such as charge or hydrophilicity.

Dip coating of silk fibroin mesh incorporated with honey extract from bees in a mixture of poly(hydroxybutyrate-co-hydroxyvalerate) (PHBV) and poly(lactic acid) (PLA) or poly(caprolactone) (PCL) was utilized for wound healing and has shown more area deposition (∼45–62%) of extracellular matrix and collagen – indicators of healing process, as compared to the silk fibroin mesh dipped in PHBV alone (no honey extract) with ∼20–24% area deposition of extracellular matrix and collagen within 28 days of treatment administration in Sprague–Dawley rats (Table 1, entry 21).⁷⁹ Since PHBV is a relatively large molecule, blending it with PLA and PCL increases the viscosity of the coating, decreasing the porosity of the mesh and therefore, adding a physical barrier to prevent fast diffusion of the honey extract towards the wound site. Layer-by-layer (LbL) coating strategy was adapted for inhibiting tumor growth in BALB/C mice with breast cancer (Table 1, entry 22).⁸⁰ Here, small interfering RNA (siRNA) was loaded in a hollow mesoporous copper sulfide nanoparticle, to silence the PD-L1 gene expression in tumor cells thus, inhibiting tumor growth. The treated group yielded the smallest tumor sizes (0.3 g) and the highest tumor inhibition rate (56%) compared to the control group after 15 days of treatment. LbL assembly was applied by electrostatically adsorbing the copper sulfide nanoparticles onto a positively-charge polymer (poly(ethyleneimine)) (PEI) forming the first coating layer. Since siRNA were easily degraded, an outermost coating of silk was applied to the CuS-siRNA-PEI nanoparticle to stabilize the therapeutic agent. Both dip coating and LbL assembly, though facile, can cover pores especially in porous silk drug delivery systems. Therefore, it is essential to know the physical properties of the material used as a coating to know which physical coating strategy can be applied to achieve the target release profile.

In some cases, two or more physical functionalization strategies can be applied to silk drug delivery systems to prolong the release and enable maximum in vivo efficacy (Table 1, entries 22–24).^81,82,97 Silk fibroin and hydroxypropyl cellulose (HPC) composite hydrogel was used to deliver folate-modified AC16 cells encapsulated in extracellular vesicles in situ in Sprague–Dawley rats for myocardial infarction repair (Table 1, entry 23).⁸² Best cardiac repair result was recorded for the rats treated with silk/HPC composite hydrogel loaded with folate-modified extracellular vesicles with left ventricular ejection fraction of ∼71% compared to the untreated group with ∼45% after 28 days of treatment. Extracellular vesicles (EVs) are cell membrane particles around 30 nm to 10 μm in diameter that contain lipids, proteins, and nucleic acids and are recently discovered to regulate myocardial fibrosis after infarction.^102–104 In this study, EVs were used to encapsulate folate-modified AC16 cells – stem cells that could differentiate into mature cardiac cells before incorporating in the silk/HPMC composite material and delivering to the target site in the form of a hydrogel. EVs are good encapsulating agents as they are naturally formed by human cells hence they are stable inside the body, naturally biocompatible, and possess low immunogenicity.^104,105 However, some of the limitations of using EVs include low drug loading, lack of standardized procedures for isolation and purification, and insufficient methods to upscale production.¹⁰⁶ Hence, therapeutic agents that can be loaded in EVs should have therapeutic dose that only require nano- to micromolar amounts and other encapsulating agents can be considered if a higher drug dosage is required.

2.2.2. Functionalization using chemical crosslinking between silk and therapeutic agents. In this silk functionalization approach, new chemical or covalent bonds are formed between silk and therapeutic agents or other polymers to form silk composite materials with the aid of crosslinking agents. Some of the established crosslinking agents used to form new covalent bonds between the silk backbone and the drug/other polymers includes glutaraldehyde (Table 2, entry 1),^107–109 EDC or 1-ethyl-3-((3-dimethyl aminopropyl)carbodiimide) (Table 2, entry 2),¹¹⁰ diazonium salts,¹¹¹ and genipin (Table 2, entry 3).¹¹² Newer studies have utilized the use of other chemical crosslinkers such as dopamine,⁸⁷ citric acid,⁹¹ epichlorohydrin,⁹² horseradish peroxidase,⁹³ poly(ethylene glycol diacrylate) (PEGDA),⁹⁴ riboflavin and hydrogen peroxide,⁹⁶ and sodium persulfate (SPS)⁹⁶ in forming covalent bonds between silk and the therapeutic agent. The main driving force for selecting these chemical crosslinkers is the resulting biocompatibility and mechanical stability of the functionalized silk drug delivery system based on the mechanism of covalent bond formation.

Table 2 Chemical crosslinkers for silk-drug covalent interaction

Entry no.	Type of reaction	Chemical crosslinker	Type of bond/products formed	Ref.
a Similar mechanism as with enzyme-catalyzed crosslinking.
1	Glutaraldehyde crosslinking			113
2	Carbodiimide coupling			113
3	Genipin crosslinking			114 and 115
4	Epoxide ring opening			116
5	Enzyme-catalyzed crosslinking			117 and 118
6	Photocrosslinking			49
7	Photocrosslinking			49 and 119
8	Thiol–ene “click” reaction			120

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For instance, dopamine is known to form a covalent bond via Michael addition or Schiff base formation of the catechol, imine, and amine groups of dopamine with free oxygen radicals and therefore, reactive oxygen species (ROS) is decreased when used as crosslinker in silk drug delivery systems which helps in diabetic wound healing.^121,122 Epichlorohydrin (ECH) can react with the hydroxyl and amino groups of various polymers including silk via epoxide ring opening or nucleophilic attack on hydroxyl groups, resulting in a silk drug delivery system with better mechanical stability which could be particularly useful as implants or scaffolds for tissue engineering (Table 2, entry 4).^123,124 Horseradish peroxidase catalyzes crosslinking reactions under mild conditions during chemical functionalization of silk drug delivery systems. Silk, having aromatic phenol group, amine, and phenolic acids, can act as a reducing substrate whereas hydrogen peroxide can be added to the solution to act as the oxidizing substrate to drive the formation of covalent bonds via a cationic radical polymerization using the heme group of horseradish peroxidase (Table 2, entry 5).¹¹⁷ Depending on the amount of the crosslinker and hydrogen peroxide, the mechanical strength and biodegradability of the silk drug delivery system can be tuned for the desired application. Since the polymerization reaction was performed under mild conditions to preserve enzymatic activity, the activity of the therapeutic agent can be preserved as well. Riboflavin, when exposed to UV, is excited to form reactive singlet oxygen (¹O₂) which attacks the tyrosine residue of silk to form a tyrosine radical intermediate. Hydrogen peroxide acts as an electron acceptor forming dityrosine linkages in silk drug delivery systems (Table 2, entry 6).^125,126 Sodium persulfate is another crosslinker that can be used to form dityrosine bonds in silk. It acts as an electron acceptor upon activation of a photoinitiator ruthenium(II) hexahydrate (Ru(II)(bpy)²⁺) (Ru) which could propagate by attacking the silk tyrosine residue to form a tyrosine radical intermediate (Table 2, entry 7).¹¹⁹ Since silk is mainly composed of tyrosine residues (∼5.1 mol% or ∼277 residues per protein),³⁵ there will be more bonding sites for riboflavin and ruthenium, increasing the polymerization duration which might not be suitable for in situ applications.⁹⁶ However, the dityrosine linkages formed make the silk or silk composite material more robust and mechanically stable.¹²⁷ Poly(ethylene glycol diacrylate) (PEGDA) has also been used as crosslinkers in silk drug delivery systems because of its ability to be photocrosslinked and to form three-dimensional (3D) polymer structures that can swell and absorb water. These 3D structures are formed by thiol–ene click reaction of PEGDA with silk. In this reaction, the acrylate groups in PEGDA will form radicals upon UV exposure and will react with the thiol groups from the cysteine residues of silk (Table 2, entry 8).¹²⁰ Unlike when riboflavin is used, thiol–ene click reactions with PEGDA are rapid since cysteine residues in silk are not as abundant as tyrosine residues and therefore, much applicable to in situ applications. The specificity of the click reactions between PEGDA and silk can also enable the functionalization of smart polymers (“stimuli-responsive”) materials which allows for a much controllable release of encapsulated therapeutics. Glutaraldehyde is regarded as the most commonly used chemical crosslinker in silk drug delivery systems because it is readily accessible and does not involve several precursor steps to bind the target molecule of interest.^107–109 The crosslinking mechanism involves formation of Schiff bases between the two carbonyl ends of glutaraldehyde and positively-charged amino groups on the lysine residues of the silk protein (Table 2, entry 1). Meanwhile, 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC) is another crosslinker that couples N-hydroxysuccinamide (NHS) with silk to form amide bonds (Table 2, entry 2).¹²⁸ In this crosslinking reaction, EDC activates the carboxyl group of silk to form an unstable O-acylisourea intermediate. The primary amines in silk attacks the O-acylisourea intermediate via a nucleophilic substitution reaction to form amide bonds. NHS is often added to form a more stable ester with the O-acylisourea intermediate which can also react with the primary amines of silk. As a result, a higher crosslinking efficiency is achieved and with less toxicity than glutaraldehyde.¹²⁹ Genipin, a crosslinker derived from gardenia fruits, has a covalent interaction with silk via the olefinic carbon at C3 position in genipin which is attacked by the amino group of silk under acidic and neutral conditions. Subsequently, a second condensation reaction occurs wherein the dihydropyran ring in genipin opens and the secondary amino group attacks the newly formed aldehyde group (Table 2, entry 3).^112,130,131 Genipin is the least toxic crosslinker among glutaraldehyde and EDC and has been increasingly used for silk drug delivery systems applied in vivo.

Recent strategies used to form covalent bonds in silk drug delivery systems for prolonged drug release include grafting another layer of polymer containing the therapeutic agent on the silk surface to form a covalently-bonded coating (Fig. 3) or direct attachment of the therapeutic agent towards the silk backbone using chemical crosslinkers. Grafting involves rearrangement of chemical bonds (“grafting to” and “grafting through”) (Fig. 3(a) and (c)) or formation of radicals (radical polymerization, “grafting from”) (Fig. 3(b)) on the polymer containing the therapeutic agent which could be triggered by various stimuli such as temperature, pH, UV light, plasma, or enzyme catalyst¹³² before incorporating with silk. On the contrary, direct attachment of the therapeutic agent towards the silk backbone involves activation of the drug to enable binding sites for the chemical crosslinkers. The crosslinkers serve as a bridge that connects the drug towards the silk substrate via covalent bonds such as amide bonds using EDC or Schiff bases using glutaraldehyde to reversibly or irreversibly bind silk to the drug adding a degree of controllability to the release of therapeutics incorporated within the silk drug delivery system while enabling slower release of drugs.

Fig. 3 Coating strategies via covalent interactions (grafting) for silk drug delivery systems. (a) “Grafting to” silk by addition of functional groups on film surface that could bind the drug noncovalently or covalently. (b). “Grafting from” silk by addition of “activated” functional groups that could bind drugs covalently (c) “Grafting through” silk by incorporating the functional groups within and on the film surface.

In vivo studies have presented the use of chemical crosslinking to fabricate more stable silk drug delivery systems in tandem with other physical functionalization techniques to further prolong drug release. Glutaraldehyde crosslinking has been employed to fabricate silk scaffolds as wound dressings (Table 1, entries 26, 27).^84,85 Glutaraldehyde crosslinked the porous chitosan–alginate layer to form the bottom hydrogel layer of the scaffold which incorporates the active therapeutic agent, 10-hydroxydecanoic acid (10-HDA), also known as the “queen bee acid”. It is found exclusively in royal jelly and has anti-inflammatory,¹³³ antimicrobial,¹³⁴ and immune system-modulating properties.¹³⁵ Electrospinning was utilized to form the poly(caprolactone)/silk sericin surface layer for the glutaraldehyde-crosslinked chitosan–alginate hydrogel scaffold. It was observed that fibrosis was less prominent with the 10-HDA-containing wound dressing than the control (without 10-HDA) after 14 days from transdermal application in Wistar rats (Table 1, entry 26).⁸⁴ When astaxanthin, an antioxidant, and silver nanoparticles, an antimicrobial were incorporated in electrospun poly(vinyl alcohol)/silk fibroin nanocomposite membrane, wound area induced in Sprague–Dawley rats was reduced to ∼4–12% as compared to the membranes without the therapeutic agents with wound area reduced to ∼22% after 14 days. In these membranes, glutaraldehyde was used to crosslink poly(vinyl alcohol), silk fibroin, and the therapeutic agent. Meanwhile, crosslinking using EDC/NHS in silk drug delivery systems has also allowed for the wound treatment in Wistar rats and BALB/C mice for approximately two weeks. For the treatment of burn wounds in Wistar rats (Table 1, entry 29),⁸⁷ vancomycin – a potent antibiotic that inhibits growth of methicillin-resistant Staphylococcus aureus (MRSA), was loaded in conductive hydrogels based on tragacanth – an anionic polysaccharide that accelerates wound regeneration,¹³⁶ and silk fibroin. EDC was used to conjugate vancomycin with silk fibroin and to functionalize the silk fibroin/tragacanth composite with a conductive carboxyl aniline pentamer. Meanwhile, EDC was used to conjugate gallic acid to a reactive oxygen species scavenging dendrimer prior to blending with poly(vinyl alcohol) and silk fibroin composite to form a hydrogel that can promote wound healing in BALB/C mice (Table 1, entry 30).⁸⁸ Genipin was used to crosslink silk fibroin with gelatin to form silk fibroin/gelatin films (Table 1, entry 31)⁸⁹ and teicoplanin-decorated reduced graphene oxide to form silk protein hybrid hydrogels (Table 1, entry 32)⁹⁰ to regenerate the tympanic membrane and accelerate diabetic wound healing in Sprague–Dawley rats and BALB/C mice, respectively. There was a 60% tympanic membrane regeneration compared to the control after 5 days⁸⁹ and an accelerated healing of infected wounds in diabetic BALB/C mice was observed after 14 days of treatment.⁹⁰ Quercetin loaded in epichlorohydrin-crosslinked silk/soy protein isolate hydrogel facilitated wound closure rates of ∼79% compared to ∼48% in the control group after 17 days of treatment (Table 1, entry 34).⁹² When bone marrow mesenchymal stem cells were loaded in dual network DNA-silk fibroin hydrogels, it was observed that there was significantly accelerated cartilage regeneration after 14 days in Sprague–Dawley rats (Table 1, entry 35).⁹³ Since enzyme crosslinking using horseradish peroxidase did not involve harsh chemicals that could influence the chemical properties of both the silk/silk composite substrate and the therapeutic agent (stem cells), significant improvement was observed in vivo. PEGDA has been used in a couple of studies as a crosslinker in a photo-initiated functionalization of hydrogels (Table 1, entries 36 and 42).^94,100 Platelet-rich plasma lysate was loaded in a silk-based photocurable hydrogel for the treatment of knee cartilage (meniscus) tears in New Zealand rabbits.⁹⁴ Methacrylate functional groups have been added to the silk and gelatin composite substrate using lithium phenyl-2,4,6-trimethylbenzoylphosphinate as initiator and PEGDA crosslinker which aided in better distribution and defect coverage of the hydrogel in rabbit menisci (Table 1, entry 36).⁹⁴ Chondroitin sulfate, a building block of human cartilage,¹³⁷ was also methacrylated using methacrylic anhydride and PEGDA to form silk fibroin/chondroitin sulfate methacrylate hydrogel bone screws for the treatment of osteoarthritis (Table 1, entry 42)¹⁰⁰ By UV crosslinking of chondroitin sulfate with silk using PEGDA, it was observed that bone screws were successfully implanted into the femur of white rabbits and were integrated well into the implantation site after a week. This could be attributed to the formation of macromolecular network framework which incorporated beta-sheet crystalline building blocks and preserved the mechanical integrity of the bone screws. Riboflavin has been an effective crosslinker as well as it enabled bone/wound healing and diabetes management by carrying adipose stem cells and insulin, respectively (Table 1, entries 38, 39).^96,97 Silk hydrogel bone screws formed from photocrosslinking silk substrate using riboflavin in the presence of hydrogen peroxide did not induce serious inflammation and maintained bone screw shape after 14 days after subcutaneous implantation of the adipose stem cell-carrying hydrogel in Sprague–Dawley rats (Table 1, entry 38).⁹⁶ When silk fibroin was blended with methacrylated hyaluronic acid (HAMA) – a stimuli-responsive polymer, and were photocrosslinked with using riboflavin, insulin was regulated and healthy blood glucose levels were maintained close to 6 h in C57BL/6 mice compared to the control group without insulin due to the regulation of the macromolecular network density, which swells upon absorbing tissue fluid after insertion into the subcutaneous tissue (Table 1, entry 39).⁹⁷ Layer-by-layer assembly was the adapted technique to the blend to ensure that there was controlled regulation of insulin after transdermal application of the hydrogel microneedles.

Chemical crosslinkers stoichiometrically reacted with a certain compound¹³⁸ and excess unreacted drugs could influence release rates through initial burst release. A striking observation when chemical crosslinkers were used to bind drugs to the silk material was some drugs were only partially released despite the prolonged period of drug release monitoring which could be quite counterproductive as this could hinder reaching the drug therapeutic window when applied in vivo. Despite the recorded sustained release of drugs in various silk formats using these crosslinkers, careful evaluation of the release profiles upon consideration of various factors such as the bonds and the by-products formed, the reversibility of the release, and the drug activity, biocompatibility, and biodegradability of the material after crosslinking must be performed to ensure that there are no cytotoxic by-products formed, the release was prolonged, and the drug activity was preserved even after covalent interaction with the chemical crosslinker.

3. Silk formats used in drug delivery systems

Several different material formats could be molded from the silk solution by controlling its physico-chemical properties and conformation. Some of the most common forms utilized in current silk drug delivery systems are gels, films/membranes, particles/spheres, and microneedle patches (Fig. 4).

Fig. 4 Various silk formats adapted as drug delivery systems. Silk can be prepared as (a) hydrogels (reprinted with permission from Wang, et al. ©2014, The Author(s)¹⁰⁹), (b) aerogels (reprinted with permission from Maleki, et al. ©2019, American Chemical Society¹³⁹), (c) films (reprinted with permission from Wang, et al. ©2022, The Author(s)¹⁴⁰), (d) electrospun membranes (reprinted with permission from Xie, et al. ©2019, The Author(s)¹⁴¹), (e) particles or spheres (reprinted with permission from Wang, et al. ©2009, Elsevier Ltd. All rights reserved.¹⁴²), and (f) microneedle patches (reprinted with permission from Gao, et al. ©2019, American Chemical Society.¹⁴³).

3.1. Gels

Gels are semisolid three-dimensional polymer networks, with a solid phase dispersed in a liquid (hydrogels) or gas phase (aerogels/foams).¹⁴⁴ Although the liquid/gas phase comprise the majority of the gel composition, these materials exhibited more solid-like properties.¹⁴⁴

3.1.1. Hydrogels. Silk hydrogels (Fig. 4(a)) have been mostly adapted in the delivery of various drugs and therapeutics in vivo since they are highly tunable and can adapt to the intricate geometries of anatomical structures inside a living organism (Table 1, entries 1, 2, 5, 8, 12, 14, 16, 20, 24, 26, 28–30, 32, 34–36, 38–43). It exists in either injectable or non-injectable forms and can also be combined with another silk format (i.e. hydrogel microneedles⁹⁷ or hydroscaffolds⁶⁴) to enable a more effective therapeutic delivery in vivo. Injectable silk hydrogels take advantage of the metastable nature of regenerated silk solution for it to be delivered on the target site in situ. Silk's tendency to self-assemble into β-sheets can be harnessed by introduction of heat (temperature-induced),^145–147 another polymer (e.g., ethylene glycol/tri(ethylene glycol)¹⁴⁸ or poly(ethylene glycol)¹⁴⁹), and energy through sonication.^150–152 Thus, injectable silk hydrogels can be fabricated by triggering self-assembly via temperature, polymer additive, or sonication before localized delivery. Increase in temperature induces self-assembly via the movement of molecules and interactions of atoms towards one another within the silk solution¹⁵³ and therefore promotes corresponding changes in secondary structure from random coils to β-sheets (Fig. 5(a)). The β-sheet structure of silk can enable drug molecules to be trapped in between “sheets” via hydrogen bonding and electrostatic interactions, depending on the pH of the environment, and therefore mitigated the burst release of drugs which was favorable for applications such as chemotherapy^145,146 and neovascularization¹⁴⁷ with treatment courses that could last for at least three months.^154,155 Polymer additives such as ethylene glycol (EG), triethylene glycol (TEG), or poly(ethylene glycol) (PEG) can also drive β-sheet formation in silk hydrogels.^148,149 These polymer additives are biocompatible¹⁴⁸ and trigger β-sheet formation in such a way that they increase the viscosity of the silk solution, breaking silk hydration layers and restricting the rotation of the β-sheet emerging units thus, facilitating the effective collision which leads to polymer aggregation and rapid gelation (Fig. 5(b)).¹⁵⁶ Ultrasonication or sonication triggers β-sheet formation in silk hydrogels by disrupting hydrophobic forces in silk protein chains resulting in accelerated formation of physical crosslinks responsible for gel stabilization.^150,151,157 During sonication, there is an increase in the local temperature, mechanical/shear forces as a result of forces during cavitation, and air–liquid interfaces at a short period of time depending on the amount of ultrasonic energy applied on the system.¹⁵⁸ Since the gelation rate is dependent on the ultrasonic wave energy applied and the duration of application, the interaction between the drug/therapeutic agent and the silk protein could be tuned. Meanwhile, most non-injectable silk hydrogels were delivered to the target site trans-dermally and could be formed either via physical or chemical means. Physical methods of non-injectable silk hydrogel formation included sonication, vortex mixing, and application of heat similar to the formation of injectable silk hydrogels whereas the chemical methods included photo-crosslinking of silk hydrogel or silk hydrogel composite components.

Fig. 5 Formation of injectable silk hydrogels. (A) Temperature-induced formation of Bombyx mori and Antheraea assamensis-blended silk fibroin hydrogels (reprinted with permission from Jaiswal, et al., ©2022, Elsevier B.V. All rights reserved.¹⁴⁵). (B) Polymer additive-induced formation of silk hydrogels using ethylene glycol/tri(ethylene) glycol (formulation of silk fibroin : ethylene glycol : tri(ethylene glycol) is 60 : 20 : 20) (reprinted with permission from Maity, et al., ©2020, American Chemical Society¹⁴⁸). (C) Sonication-induced formation of silk fibroin hydrogels with miR-675-loaded stem cell-derived exosomes (reprinted with permission from Han, et al., ©2019, The Author(s). Published by Elsevier B.V.¹⁵¹).

Doxorubicin (chemotherapeutic drug) and desferrioxamine (drug for neovascularization) release were sustained for up to two months after administering the temperature-induced silk hydrogel intratumorally or in situ in female BALB/c nude mice¹⁴⁶ and Sprague–Dawley adult male rats.¹⁴⁷ The results could be attributed to the more ordered structure of the β-sheets which trapped the drugs upon heating. Although promising, the use of temperature in silk hydrogel formation was not applicable to incorporate drugs or therapeutic agents that were unstable or lose activity at very high temperatures. Therefore, other methods of silk hydrogel formation could be adapted such as the use of biocompatible polymer additives and sonication to deliver temperature-sensitive drugs or therapeutic agents. When polymer additive-induced silk hydrogel delivered insulin in Wistar rats,¹⁴⁸ and vancomycin subcutaneously and in situ¹⁴⁹ as treatment for diabetes and chronic osteomyelitis (bone infection), respectively, release of therapeutics was recorded to be after five days to a week. Typically, diabetes treatment for type 2 patients lasts for up to two days and the silk-EG/TEG hydrogel drug delivery system being able to sustain the release of insulin for up to five days promotes glucose homeostasis in diabetic rats.¹⁴⁸ A sustained release of ≥3 months for bevacizumab, a therapeutic agent (monoclonal antibody) that has an affinity against vascular endothelial growth factor (VEGF), was recorded in Dutch-belted rabbits when sonication-induced silk hydrogels were used to encapsulate this therapeutic agent for the potential treatment of age-related macular degeneration (AMD).¹⁵⁰ Low ultrasonic energy was applied to ensure that there was minimal increase in temperature during sonication that induced rapid gelation, allowing for more interaction between bevacizumab and silk. The therapeutic threshold for typical hydrogel formulations for bevacizumab was one injection per month which meant that sonication was able to prolong the release of the anti-VEGF, bevacizumab. Significantly enhanced blood perfusion was also observed in C57BL6 mice on day 28 when sonication-induced silk hydrogel was used to encapsulate miR-675, a modulator of aging process via a stem cell-derived exosome for the treatment of vascular diseases although it should be taken note of that a temperature of 37 °C was applied too (Fig. 5(c)).¹⁵¹

A non-injectable, sonication-induced heparinized-silk fibroin hydrogel was used to encapsulate fibroblast growth factor (FGF-1) for wound healing applications.¹⁵⁷ A burst release (43.7%) was recorded in the first 12 h after application on the wound site on Sprague–Dawley rats and a slow release (70%) was observed over the course of 5 days. A study also showed that a 3D-bioprinted non-injectable silk fibroin/gelatin hydrogel delivered trans-dermally was able to reduce the wound area of wounded Kunming mice compared to the control group after two weeks which could be attributed to the reduced viability of bacteria S. aureus and E. coli only after 20 min in vitro.¹⁵⁹

Despite the preference on silk hydrogels over other formats, silk hydrogels remain to have limitations on its mechanical integrity and therefore, optimization is still necessary for the formulation of silk hydrogels that enables better mechanical strength in vivo while ensuring degradation rates are rapid enough to minimize adverse physiological reactions.

3.1.2. Aerogels/foams. Silk aerogels (Fig. 4(b)) are advantageous over hydrogels due to their more improved mechanical stability, with higher stress resistance and elastic recovery owing to their macroporosity.^160,161 Silk aerogels can be fabricated via drying at room temperature (xerogels),¹⁶² by using supercritical carbon dioxide (scCO₂),^163,164via freeze-drying (also known as lyophilization) (lyogels)¹⁶⁵ and freezing (cryogels),¹⁶⁶ and via a combination of both lyophilization and 3D printing.¹⁶⁷ Although foams have better mechanical integrity than hydrogels, application as drug delivery systems in vivo has been limited⁷⁷ which could be attributed to burst release of drugs due to the highly porous structure¹⁶³ and the slower degradation rates of foams compared to hydrogels due to higher number of beta-sheets. Therefore, foams are currently used as scaffolds for bone/tissue regeneration (Table 1, entry 19).⁷⁷

In vitro release studies showed that some silk foams were capable of sustaining release for longer periods depending on the fabrication method which influences foam morphology, the functionalization technique applied on the material which influences material-therapeutic agent interaction, and the chemical properties of the therapeutic agent which influences diffusion of the drug from the material. For instance, silk fibroin xerogels sustained the delivery of estradiol (Fig. 6(a)), a hormone used for contraception and hormone replacement therapy (HRT) by enabling slow release (∼90%) after 129 days (∼four months).¹⁶² Silk xerogel morphology for the delivery of estradiol (Fig. 6(a)) showed a compact surface with pores sparsely distributed on the xerogel surface which could have prevented initial burst release. Meanwhile, the therapeutic level for estradiol had a wide range with hardly any published literature for its specific therapeutic index.^168,169 Estradiol was observed to form crystals when the lyophilized foams were treated with ethanol for β-sheet formation, with estradiol lipophilicity also contributing to its slow diffusion in the aqueous release bath. Aerogels fabricated using scCO₂ drying were also utilized to deliver ibuprofen (Fig. 6(b)) and a release of ∼90% after 6 h was observed.¹⁶³ A highly porous surface morphology with more evenly distributed pores could have contributed to this result (Fig. 6(b)). Ibuprofen also had a wide therapeutic range (800–1200 mg day⁻¹ at low, over-the-counter doses; 1800–2400 mg day⁻¹ at prescription doses),^170,171 and the drug release behavior must be controlled depending on whether it was used acutely, as an antipyretic or temporary pain reliever or chronically, for the long-term treatment of arthritis, ankylosing spondylitis, and other long-term conditions.¹⁷¹ For ibuprofen-loaded aerogels, the majority of the drug was released within a short time period (6 h) since physical adsorption between the silk and the drug was the prevailing interaction for the whole system which was ideal since ibuprofen should not exist in increased levels in the body for a prolonged period to prevent hepatotoxicity.¹⁷² Lyogels were also previously used to deliver murine anti-TGFβ monoclonal antibody and it was observed that using this silk format, sustained release of the antibodies (∼15–70%) can be achieved until 38 days.¹⁶⁵ Lyogels, most particularly those with higher silk content, were more resistant to swelling due to its solid structure. In addition, the silk-antibody interaction was strengthened upon water removal resulting in prolonged antibody release. It was also found that using this silk format, murine anti-TGFβ monoclonal antibody activity lie within the expected range which meant that the released antibody was stable even after 38 days. Ciprofloxacin and a protease enzyme were also loaded into a cryogel (Fig. 6(c)) and showed a controllable drug release (∼20%) until about 50 days when the drug in powder form was loaded.¹⁶⁶ Cross-sectional micrographs (Fig. 6(c)) showed a stacked sheet-like structure which could have served as a barrier for the antibodies that were bound to the innermost region of the lyogel enabling the release to be sustained for around a month. The antibiotic also has retained its activity as manifested in the clearing zones observed from the antibacterial assay. 3D-printed aerogels were also used to deliver sorafenib (SFN) (Fig. 6(d)) during photothermal bone cancer therapy,¹⁶⁷ and it was observed that ∼70–90% was released for up to around seven hours. 3D printing allowed for the fabrication of a dual network structure leading to the attainment of a hierarchically organized porosity for a more controlled release (Fig. 6(d)). Still, careful consideration must be given to the method of drug incorporation (i.e. physical blending or chemical crosslinking),¹⁷³ the composite materials to be added for better dissolution or dispersion of the drug (i.e. silica, MXenes),^{167,174–176} and the solubility of the drug towards the chosen solvent (i.e. use of binary solvents)¹⁷⁷ for the silk bio-ink to ensure that more therapeutics were internalized within the 3D-printed structure.

Fig. 6 Morphological characteristics of silk (a) xerogels loaded with estradiol (reprinted with permission from Krizman, et al., ©2021, Published by Elsevier B.V.),¹⁶² (b) aerogels made from drying using supercritical carbon dioxide loaded with ibuprofen (reprinted with permission from Marin, et al., ©2014, Elsevier B.V. Published by Elsevier B.V. All rights reserved.),¹⁶³ (c) cryogels loaded with ciprofloxacin hydrochloride (reprinted with permission from Chambre, et al., ©2020, American Chemical Society.),¹⁶⁶ (d) lyophilized 3D printed aerogels loaded with bismuth sulfide nanorods (reprinted with permission from Al-Jawuschi, et al., ©2023, American Chemical Society).¹⁶⁷

A current study on the in vivo bone regeneration in Sprague–Dawley rats showed that bone volume fraction was higher (∼46%) in cell-free chitosan/silk fibroin composite scaffolds loaded with bioactive glass compared to the control (without bioactive glass) (∼42%) at 12 weeks of treatment. The sustained release could be attributed to the very complex pore architecture brought by the fabrication method (radial freezing technique) which prevented the ready diffusion of bioglass from the composite scaffold (Table 1, entry 19).⁷⁷

The silk foam format is mainly for the purposes of increased drug loading compared with other silk formats (i.e. aerogel porosity) and mechanical stability (i.e. flexibility). The three-dimensional (3D), porous structure of silk aerogels results in a high surface area (∼200–400 m² g⁻¹)¹⁶⁴ which can be utilized to load higher drug concentrations for localized delivery of therapeutics.¹⁶⁷ Opting for silk aerogel format would be beneficial if there was an involvement of organic solvent or other stimuli such as extreme temperature¹⁷⁸ and pH¹⁷⁹ during the fabrication process which could alter the mechanical stability of the silk aerogel upon delivery of a drug or therapeutic agent to the target site.

3.2. Films/membranes

Silk films/membranes (Fig. 4(c) and (d)) are two-dimensional drug delivery systems. Like foams, silk films are highly applicable as scaffolds for bone and tissue regeneration or as implants or dressings for wound healing because of its sheet-like geometry and higher mechanical integrity than hydrogels (Table 1, entries 6, 7, 11, 13, 15, 17, 18, 21, 26, 27, 31, 33, 37, and 41). This silk format was primarily fabricated via the conventional dry casting wherein the silk solution is drop-casted in a poly(styrene) Petri dish,^108,180,181 a glass plate,⁵² or relatively inert mold¹⁸² and is allowed to dry in ambient conditions but in some applications (i.e. corneal tissue engineering¹⁸³ and scaffolds for neuronal growth¹⁸⁴), centrifugal casting technique (centrifugal force is applied on an aluminum tube with silk fibroin solution), and electrospinning¹⁸⁵ are also utilized. Electrospinning¹⁴¹ is another common method used to fabricate silk membranes, which can be also loaded with drugs. Homogeneity/solubility of the therapeutic agent on the spinning solution was mainly important for slower release of the drug.^141,184,185 Film/membrane structure is more controllable in electrospinning than solution casting since formation of solid matrix occurs inside an electrospinning machine wherein the main driving force of film/membrane formation is the tunable electric current whereas in solution casting, it happens mainly via solvent evaporation. However, a more mechanically stable film/membrane can be formed via solution casting than electrospinning since silk proteins can self-assemble during casting therefore reinforcing film/membrane mechanical properties compared to irregularly oriented fibers than can be obtained on the electrospun mats.¹⁸⁶ In both cases, drug chemistry, particularly its solubility and charge, has a significant effect in its incorporation inside the film/membrane which can ultimately influence its release. In terms of release behavior, the key to prevent burst release and sustaining the amount of drugs released over time for this material format was to increase the number of β-sheets to improve silk film crystallinity and increase the interpenetrating networks (physical crosslinks) which could trap the small drug molecules and restrict film swelling thus hindering fast diffusion of drugs towards the release bath.¹⁸⁷ The most common solvents used to form silk films/membranes with high β-sheet content were formic acid,¹⁸⁸ and 1,1,1,3,3,3-hexafluoroisopropanol (HFIP).¹⁸⁹ Conformational change that promoted tighter interactions between silk molecules causing increased number of β-sheets in silk films was brought by the hydrogen bonds that could be formed between the hydroxyl groups of formic acid,¹⁸⁸ HFIP,¹⁹⁰ and silk upon the evaporation of the solvents.

Annealing is another strategy used to increase the β-sheets in the films by providing sufficient thermal energy for large-scale molecular motion of protein chains; therefore, allowing for faster crystallization due to the kinetic motion of the molecules.¹⁹¹ The annealing process could be water-based^37,181 or alcohol-based (commonly methanol/ethanol).^181,192 Glycerol or poly(vinyl alcohol) could be added as a plasticizer to improve film flexibility for annealed films.^193,194

A number of current studies has incorporated flavonoids in silk nanofibrous mats for wound healing in rats (Table 1, entries 6 and 15).^65,73 Flavonoids are a class of plant-derived compounds characterized by polyphenol groups in their chemical structure which are found to have wound–healing properties attributed to their anti-inflammatory, angiogenesis (formation of new blood vessels), re-epithelialization (restoration of skin barrier), and antioxidant effects.¹⁹⁵ Full wound closure can be observed from 9–15 days after transdermal application of the silk nanofibrous mats. In other studies, silver nanoparticles has been incorporated with silk composite materials for better wound healing in rats (Table 1, entries 11 and 27).^69,85 Silver nanoparticles has been known to aid in wound healing by controlling the release of cytokine (proteins that act as immune system messengers) which are responsible for inflammation.¹⁹⁶ A healing percentage of ∼88–99% was recorded after 14–21 days of treatment when silver was combined with gold nanoparticles⁶⁹ or astaxanthin, a carotenoid found to have high antioxidant properties and promotes collagen production⁸⁵ prior to incorporation with silk fibroin/collagen blend or silk fibroin/poly(vinyl alcohol) composite, respectively. Berberine, also a plant-derived compound that belong to alkaloid class with similar functions as flavonoids, has been loaded in electrospun silk fibroin blended with biodegradable polymer such as poly(hydroxybutyrate-co-valerate) (PHBV) or poly(caprolactone) (PCL) for wound healing in diabetic mice (Table 1, entries 17 and 18). It was observed when berberine-loaded electrospun silk fibroin/poly(hydroxybutyrate-co-valerate) (PHBV) wound dressing was applied on diabetic mice, wound healed completely at a significantly higher rate (∼99.7%) than the control group (∼83–85%) after 18 days treatment (Table 1, entry 17).⁷⁵ A more sustained release was also recorded for berberine-loaded silk fibroin/poly(caprolactone) (PCL) nanofibrous mat where bone regeneration can be observed after 2–6 weeks (Table 1, entry 18).⁷⁶ The differences in release rates could be due to the faster degradation rate exhibited by PHBV than PCL which enabled more berberine to be released at a higher dose at a shorter period of time. Another therapeutic agent such as honey extract from bees loaded in silk fibroin mesh fabricated by dip coating silk membrane in a mixture of poly(hydroxybutyrate) (PHBV) or poly(caprolactone) (PCL) also resulted in faster wound healing 28 days after implantation (Table 1, entry 21).⁷⁹

Gelatin was blended with silk fibroin to form composite films that were used for tissue regeneration in rats (Table 1, entries 31 and 33).^89,91 It was observed that tympanic membrane has regenerated and 60% of the perforation has been covered compared to 40% in control five days after surgical implantation (Table 1, entry 31).

Growth factors have been also incorporated within silk and silk composite scaffolds for acute liver failure treatment and spinal cord repair in rats and mice (Table 1, entries 7, 37, and 41).^66,95,99 A combination of growth factors such as hepatocyte growth factor (HGF)/fibroblast growth factor-4 (FGF-4) loaded in silk electrospun scaffolds aided in reducing the biomarker concentration for hepatocyte damage in BALB/C mice after 14 days of transplantation on injured livers (Table 1, entry 7).⁶⁶ Basic fibroblast growth factor was also loaded in silk nanofibril/hyaluronic acid scaffold for spinal cord repair in Sprague–Dawley rats (Table 1, entry 37). In the case of spinal cord injury repair, motor function was recovered 8–10 weeks after subcutaneous implantation of silk nanofibril/hyaluronic acid scaffold loaded with basic fibroblast growth factor (Table 1, entry 37)⁹⁵ and dual-phase silk fibroin methacryloyl (SilMA) hydrogel scaffold (DPSH) loaded with neurotrophin-3 (NT-3) and angiotensin (1–7) (Ang-(1–7)) (Table 1, entry 41),⁹⁹ respectively.

Ultimately, the severity of the wound or tissue damage will determine how the silk film/membrane can be fabricated to ensure the highest therapeutic agent loading and delivery towards the target site for faster recovery. Sustained release of therapeutic agents is necessary for diabetic wound healing and spinal cord repair hence the use of materials that could form composites with silk must not degrade at a relatively fast rate. Otherwise, if the silk film/membrane is to be applied for normal wound healing, faster degradation rate of the material is preferable.

3.3. Particles

Silk particles (Fig. 4(e)) are one-dimensional formats used to encapsulate the drug/therapeutic agent of interest. They can exist from nano- to micro- range and display fast release rates when directly applied to the target site due to high surface area.^197–199 Silk particles can also be embedded or incorporated within scaffolds/implants to sustain the release of the therapeutic agent for a longer duration.^200–202

Silk particles can be synthesized via salting out,²⁰³ nanoprecipitation,^198,204,205 emulsification,^200,206 coaxial electrospraying,²⁰¹ spray drying/spray freeze drying,²⁰² and semi-batch silk desolvation. Salting out is a process wherein silk particles are formed by adding high concentrations of salt in the silk solution.²⁰⁷ The salt ion can be a kosmotropic agent which competitively interacts with the water molecules that forms hydrogen bonds with the amide moieties in silk, therefore facilitating dehydration and favoring precipitation of the protein resulting in the formation of silk particles.²⁰⁷ By selecting the proper type of salt to use during the “salting out” process, the dynamic transition of the aggregate and the controllability of particle formation can be achieved.²⁰⁸ Nanoprecipitation is another method of silk particle formation wherein the aqueous silk solution is dissolved in a partially miscible organic solvent from which the heavy-chain hydrophilic blocks have low solubility.²⁰⁹ Organic solvents can also serve as kosmotropic agents as they can induce protein aggregation by lowering dielectric constant and increasing protein–protein interactions.²¹⁰ Emulsification is another technique of silk particle formation wherein an aqueous phase is dispersed in an oil phase to form water-in-oil (W/O), oil-in-water (O/W), or water-in-oil-in-water (W/O/W) emulsions upon application of an external force (i.e. ultrasonication). For the fabrication of silk particles using this technique, silk aqueous solutions are often dispersed in a continuous oil phase, or another immiscible liquid, to form the particles or spheres.²¹¹ Silk, being an amphiphilic multi-block copolymer by itself, serves as an emulsion stabilizer as it could interact and self-assemble between two phases, decreasing interfacial tension.²¹² Coaxial electrospraying involves ejecting the silk polymer solution via electrostatic forces forming droplets which eventually turn into nanoparticles upon solvent evaporation.²⁰¹ This technique is promising as it does not require the emulsification step or further purification which could alter silk properties and reduce the activity of the encapsulated therapeutic agent.²⁰¹ Meanwhile, for spray drying, a spray dryer is used to aerosolize the silk solution and spray it into a heated drying chamber to evaporate the solvent and leave a solid particle. In spray freeze drying, silk solution is atomized using a nozzle and sprayed into liquid nitrogen prior to lyophilization.

Silk particles has been mostly used to encapsulate drugs to treat tumors that cause cancer (Table 1, entries 9, 22, and 25).^56,80,83 Doxorubicin hydrochloride and small interfering RNAs (siRNAs) were loaded in silk/poly(D,L-lactic acid) microspheres and silk fibroin shell, respectively whereas a tumor-targeting peptide was coated with silk to form nanorods as cancer treatment in BALB/C nude mice. When these were administered intratumorally, tumor volume is minimal after about two weeks.

Metformin-loaded silk/poly(vinyl alcohol) microparticles were used to treat corneal neovascularization (abnormal growth of blood vessels in a typically avascular cornea) in rats and showed reduced neovascularization and inflammation after seven days (Table 1, entry 10).⁶⁸ Methylprednisolone acetate (MPA), a type of corticosteroid, was also loaded in silk fibroin/chitosan interpenetrating network hydrogel for cartilage regeneration in white rabbits (Table 1, entry 43).¹⁰¹ It was found that locomotor function in treated groups has improved after 12 weeks. Silk particles can typically sustain the release of drugs for at least two weeks and the introduction of a hydrogel network during fabrication can sustain the release of MPA from silk fibroin particles for up to 12 weeks.

Functionalized silk particles allowed for easier achievement of therapeutic dose since more drugs could be encapsulated or loaded and due to their size, they could also penetrate cell membranes or even the blood–brain barrier which was highly favorable for targeted or localized drug delivery.^197,213 However, their size could also pose some disadvantages when it came to introducing the drug.²¹⁴ Since cells internalized these materials easily at this form,²¹³ careful optimization of the maximum noncytotoxic concentration of the drug loaded must be performed prior to and upon encapsulating with silk particles to ensure that the drugs would not agglomerate resulting in increased size.

3.4. Microneedles

Microneedles (Fig. 4(f)) are an array of micron-sized needles typically ranging from 150–1500 μm in length, 50–250 μm in width, and 1–25 μm in thickness attached on top of a supporting component that allow these needles to penetrate through the skin and deliver therapeutic agents transdermally without incurring pain to the patient.²¹⁵ This format can be solid, coated, hollow, or dissolvable depending on the required mode of delivery. Solid silk microneedles (Fig. 7(a)) are often administered following a “poke-and-patch approach”²¹⁶ wherein the microneedles directly puncture the skin surface.²¹⁷ One limitation of solid microneedles was it was difficult to maintain therapeutic dose for this microneedle type due to the inherent ability of skin to heal over time.²¹⁸ Silk-coated microneedles (Fig. 7(b)) is another type of microneedle applied using the “coat-and-poke approach”²¹⁶ wherein therapeutic agents are coated on the tip of the microneedle before administration. This technique might not be the best for some drugs that were attached only via electrostatic interaction or physical adsorption since the premature release could occur unless initial burst release was sufficient to deliver the therapeutic level of the drug.^219,220 For instance, immobilization of a peptide within a coated silk microneedle for insulin delivery only required 15 min for a therapeutic level to be delivered. Dissolvable microneedles (Fig. 7(c)) have tips made up of soluble polymers with encapsulated therapeutic agents.^216,221 Although less expensive and enabled higher dose of drugs via encapsulation than coated microneedles, the “poke-and-dissolve approach” has its limitations in terms of other “non-drug” components which can also be degraded during the delivery process.²²² Finally, silk swelling or hydrogel microneedles also existed (Fig. 7(d)) wherein a swellable material was used to control drug release.^216,223 This microneedle type was particularly interesting to use in smart release systems triggered by various stimuli due to the maneuverability of its degree of swelling.^223,224 However, one downside to this approach was that highly tunable silk hydrogel microneedles had trade-offs in mechanical integrity. To counter this, a new class of porous microneedle had emerged using a mechanically stable photocurable resin as substrate resulting in the preservation of the hardness of the microneedle during delivery.²²²

Fig. 7 Silk microneedle types. (a) Solid microneedles for levonorgestrel delivery (reprinted with permission from Yavuz, et al., ©2020, American Chemical Society²¹⁷), (b) coated microneedles for Rhodamine B (model drug) delivery (reprinted with permission from Lin, et al., ©2020, The Author(s)²⁵), (c) dissolvable microneedles loaded with antigen for vaccine delivery (reprinted with permission from Boopathy, et al., ©2019, The Author(s)²²¹), (d) swellable microneedles for FITC-Dextran (model drug) release (reprinted with permission from Yin, et al., ©2018, Elsevier B.V. All rights reserved.²²³).

Silk fibroin microneedles loaded with melatonin for circadian rhythm regulation in Sprague–Dawley rats have exhibited burst release upon administration as it induced sleep immediately but was able to sustain the release after 8 h after melatonin concentration peaked after about an hour (Table 1, entry 4).⁶³ A dissolving silk microneedle system loaded with a topical antibiotic mupirocin and silver nanoparticles was used to treat delayed wound healing in BALB/C mice and it was observed that wound contraction rate increases to ∼78% 11 days post-treatment (Table 1, entry 23).⁸¹ When insulin was loaded in silk fibroin/hyaluronic acid hydrogel microneedles for diabetes management in C57BL/6 mice, it was reported that there was a gentle regulation and prolonged maintenance of healthy glucose blood levels for close to six hours compared to the blank group (Table 1, entry 39).⁹⁷ Silk microneedles used in current in vivo studies did not require high doses of therapeutic agents for effective activity which made it suitable for these applications. In addition, a rapid effect is required for some instances (i.e. sleep induction or diabetes management) hence, burst release is not an issue. However due to the high surface area brought by the morphology of silk microneedles (pointed geometry), it is recommended to maximize the utility of this format in terms of drug/therapeutic agent loading for a more effective and rapid drug action.

4. Conclusions and perspectives

Based on recent studies, most silk functionalization methods are oriented more towards blending or coating another biocompatible/biodegradable material with silk rather than conjugating silk with the drug or therapeutic agent for sustained release in vivo. This technique is highly attributed to the effect of the interpenetrating networks between silk and the other material in “caging” the drug of interest and thereby slowing down its release. Although this strategy is effective in prolonging the release in some studies, maximum efficiency could not be achieved at a shorter period of time which could ultimately result in decreasing the overall treatment duration resulting in better patient comfort. Some factors that hinder the maximum efficiency of these silk drug delivery systems include limitations in drug loading due to the morphological properties of material format selected for a certain type of application resulting in the inability to reach the therapeutic threshold, the impact of the reagents used in functionalization of the drug delivery system in the mechanical stability of the material, and the biodegradability of the drug delivery system considering the standardized treatment protocol for a certain disease. There has also been an inadequate evaluation of the biocompatibility of other components aside from silk that were utilized to coat the therapeutic agent or the substrate. For instance, in the case of using chemical crosslinkers, increased concentration may probably be required to bind the therapeutic agent and the silk or silk composite substrate together to form a homogeneous surface coating or to sustain the release in vivo. Even though there are already individual studies suggesting the limited biocompatibility of some crosslinkers (e.g. EDC²²⁵ and genipin²²⁶), in vivo assessment of the concentration that could have leached from the substrate is lacking.

To overcome these limitations in the practicality of current silk drug delivery systems towards in vivo testing, therapeutic agents or silk drug delivery systems were modified using hybrid techniques such as combining a coating strategy with other surface modification techniques (i.e. coating the silk substrate after embedding the therapeutic agent via physical means) to allow for higher drug loading concentrations consequently ensuring that the target release profile and the drug therapeutic level could be achieved. The concentration of toxic components that could leach from the sample could be minimized as well.

Interestingly, the preference on the use of silk hydrogel drug delivery systems in current in vivo studies than films or foams could be attributed to silk hydrogel biocompatibility. Aqueous silk solutions, which is the precursor of most silk hydrogels, could straightforwardly be treated with heat or sonication to drive the silk hydrogel formation and do not involve potentially harsh solvents or other polymers which could have a cytotoxic by-product. When it comes to aerogels (foams), 3D printing is gaining traction as a precise technique to achieve the desired material format and to also mitigate the use of excess chemical reagents (i.e. organic solvents used for the polymer dope solution) for enhanced biocompatibility since the amount of reagents can be automatically input in a 3D printing software.¹⁷⁶ Despite the increasing popularity of 3D printing in aerogel fabrication, success in vivo is yet to be demonstrated since the form of the drug (i.e. pure powder or as conjugates) that could be encapsulated with foams are very limited.¹⁶⁶ It might be necessary to synthesize several forms of drug to ensure maximum solubility or dispersibility with the silk solution hence leading to better release behaviors.

It must also be emphasized that consideration of drug properties and target application (i.e. skin patch, implant, inhalant, etc.) is highly significant. Despite silk's well-established reputation as versatile biomaterial, designing the most appropriate silk drug delivery system for a particular application is quite challenging since the silk format (i.e. gels, films, particles, etc.) and silk physico-chemical properties (i.e. molecular weight, isoelectric point, temperature stability, etc.) must be tailored to the drug's pharmacokinetics and intended application via various functionalization strategies to attain the desired release profile. Not to mention, it must be ensured that the activity of the drug/therapeutic agent after release was preserved.

Drug properties such as the size of the drug, lipophilicity, solubility, surface charge, half-life, isoelectric point, and acid-dissociation constant were some of the factors that must be considered along with the selection of the silk material format. For instance, the charge at a certain pH must be determined to know whether it would electrostatically bind to silk with a pI of 4.5 at that certain pH.²²⁷ During internalization of the drug with the silk, it would also be beneficial to know whether the drug would aggregate at the silk fabrication pH as this could possibly affect the drug release profile later on. By carefully studying the chemistry of the components and their compatibility with silk, it is possible to combine two or more therapeutic agents in one format. Selecting a simple approach to functionalize the silk drug delivery system (i.e. bulk blending) does not necessarily lead to burst release of drugs or therapeutic agents from the silk DDS but by comparing bulk-blended silk DDS over silk composite materials showed that shorter release times can be recorded from silk composite DDS, mainly due to insufficient understanding of the chemical properties of the two polymers being combined prior to encapsulating the therapeutic agent. Layer-by-layer technique can be very much applicable in nonporous silk drug delivery systems, wherein the release will be activated upon the degradation of the material. Dip coating can be used if there is a need to preserve the porosity of the substrate for the faster diffusion of drugs to the target site. Crosslinking between silk and another polymer is common to facilitate formation of drug delivery systems with better mechanical strength but studies showing crosslinking between silk and the therapeutic agent is less prominent because as this can alter the drug structure and hence, its activity/effectivity. Therefore, these factors must be considered in designing the preclinical models for the continuously emerging silk drug delivery systems.

Data availability

No primary research results, software or code have been included and no new data were generated or analyzed as part of this review.

Conflicts of interest

There are no conflicts to declare.

Acknowledgements

Dr Filippo Valente and Dr Christine Bacal would like to acknowledge Ear Science Institute Australia and the Channel 7 Telethon Kids Grant of Dr Filippo Valente for the support in the completion of this article.

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