Trends in Injectable Biomaterials for Vocal Fold Regeneration and Long-Term Augmentation

Abstract

Human vocal folds (VFs), a pair of small, soft tissues in the larynx, have a layered mucosal structure with unique mechanical strength to support high-level tissue deformation by phonation. Severe pathological changes to VF have causes including surgery, trauma, age-related atrophy, and radiation, and lead to partial or complete communication loss and difficulty in breathing and swallowing. VF glottal insufficiency requires injectable VF biomaterials such as hyaluronan, calcium hydroxyapatite, and autologous fat to augment VF functions. Although these biomaterials provide an effective short-term solution, significant variations in patient response and requirements of repeat reinjection remain notable challenges in clinical practice. Tissue engineering strategies have been actively explored in the search of an injectable biomaterial that possesses the capacity to match native tissue’s material properties while promoting permanent tissue regeneration. This review aims to assess the current status of biomaterial development in VF tissue engineering. The focus will be on examining state-of-the-art techniques including modification with bioactive molecules, cell encapsulation, composite materials, and in situ crosslinking with click chemistry. We will discuss potential opportunities that can further leverage these engineering techniques in the advancement of VF injectable biomaterials.

Impact Statement

Injectable vocal fold (VF) biomaterials augment tissue function through minimally invasive procedures, yet there remains a need for long-term VF reparation. This article reviews cutting-edge research in VF biomaterial development to propose safe and effective tissue engineering strategies for improving regenerative outcomes. Special focus is paid to methods to enhance bioactivity and achieve tissue-mimicking mechanical properties, longer in situ stability, and inherent biomaterial bioactivity.

Introduction

Human vocal fold (VFs) have distinct layers of submucosal connective tissue sandwiched between the laryngeal epithelium and the thyroarytenoid muscle and anchored by the arytenoid cartilage (Fig. 1A).¹ During phonation, the VF stretch longitudinally between 5 and 15 mm, while the superficial lamina propria (LP) oscillates with up to 0.5 mm amplitude.^1,2 These three layers are the superficial, intermediate, and deep LP (Fig. 1B). Advanced imaging studies using multiphoton microscopy and optical coherence tomography found a continuous transition between LP layers, further elucidating the intricate organization of VF extracellular matrix (ECM) and microvasculature, which requires tissue-specific regenerative solutions.^3–7

FIG. 1.

VF Anatomy and Injection Laryngoplasty. (A) VFs are housed in the larynx. (B) VF is composed of EP, SLP, ILP, DLP, and M with rich extracellular matrix and vasculature. (C) Laryngoscopic images of injection laryngoplasty with Atelo-Collagen to VF with bilateral atrophy during respiration (top row) and phonation (bottom row). A and B created with BioRender.com. DLP, deep lamina propria; EP, epithelium; ILP, intermediate lamina propria; M, muscle; SLP, superficial lamina propria; VF, Vocal fold.

The ECM structure plays an important role in supporting phonation.^2,8,9 Collagens are the primary constituents of VF ECM, contributing to 40–50% of total protein.¹⁰ Collagen I fibers form the overarching structural network, regulate cell distribution, and recruit cells for wound healing.^10,11 Collagen III and elastin bundles contribute to VF resilience, tensile strength, and viscoelasticity. Hyaluronan is an abundant glycosaminoglycan that helps regulate wound healing and dampens mechanical stress from high-frequency phonatory oscillation.^11,12

VF atrophy, sulcus, scarring, and unilateral paralysis/paresis are candidates for injection laryngoplasty, where a biomaterial is injected to augment VF function and closure.^13–16 Age-associated VF atrophy is characterized by tissue bowing and reduced viscoelasticity that may impact LP or muscle.^17–21 Sulcus vocalis is related to fibrotic alterations characterized by groove formation and epithelium invading the superficial LP.²² VF scarring is generally caused by surgical interventions, radiation, or trauma, where healthy LP is replaced by excessive fibrous collagen deposition.^12,17 VF paralysis/paresis is unilateral or bilateral VF immobilization caused by injury to the recurrent laryngeal nerve from cervical or thoracic surgical intervention, tumor formation, trauma, radiation, and chronic inflammation.^23–26 Bilateral VF paralysis may require reinnervation or airway bypass surgeries. However, unilateral VF paralysis and paresis account for up to 75% of injection laryngoplasty treatments.^27–29

Injection laryngoplasty is performed in operating rooms or doctor’s offices (Fig. 1C).^30–32 In-office injection laryngoplasty offers relatively low risk and cost as well as immediate testing of glottal function. The materials are injected into the lateral aspect of the thyroarytenoid muscle, or the paraglottic space.^32–35 Material placement is critical to treatment outcomes. Superficial injection causes improper biomaterial localization, tissue stiffening, and impeded vibration.^36–38 While there is interest in developing regenerative biomaterials and delivery techniques for the superficial LP, intramucosal injections are limited to free-drug formulations such as steroids and growth factors.^36,39–41

In this article, the primary goal is to assess the successes and limitations of current VF injectable biomaterials and review the state-of-the-art of VF biomaterials under development. We also evaluate advanced tissue engineering techniques that can help stimulate regeneration, mimic mechanical properties, and control material degradation toward permanent VF augmentation.

Overview of Existing VF Biomaterials Used in Clinics

The first injectable VF biomaterial, Teflon, caused severe granuloma for 5–10 years following injection and removal caused by tissue morbidity.⁴² Also, collagen (Zyplast^®) and gelatin-based (Gelfoam^®) injectables were used for VF augmentation for nearly 20 years but have been discontinued due to short-term efficacy and chance of allergic reaction to bovine collagen.^43,44 Atelocollagen, a similar bovine collagen material, continues to be used in Japan.⁴⁵ Cymetra™ is a discontinued micronized decellularized dermal matrix sourced from human cadavers that lasted 6 weeks to 6 months and reduced the risk of allergic reaction, decreased foreign body response, and improved voice outcomes compared to bovine collagen.^46,47

These examples emphasize the need for long-term safety evaluations and follow-up.⁴² At present, common biomaterials for injection laryngoplasty include hyaluronan, calcium hydroxyapatite with carboxymethylcellulose, autologous fat, and silk hyaluronan (Table 1).⁴⁸

Table 1.

Common VF Injectable Biomaterials

Biomaterial type	Commercial name(s)	Status	Advantages	Disadvantages
Autologous Fat	N/A	Off-Label Use	- Longer-lasting results - Tissue remodeling capacity - VF-like mechanical properties	- Unpredictable deleterious outcomes - Frequent resorption - Risk of overinjection - Variation and difficulty of preparation
Calcium hydroxylapatite	Radiesse^®, Prolaryn Plus^®, Renú Voice^®	FDA Approved	- Up to 24-month residence time - FDA approved for VF applications	- Tissue stiffening - Potential for impaired VF function - Potential for migration out of target tissue - Potential or deleterious foreign body response
Collagen	Zyplast^®	Discontinued	- Low swelling	- Risk of allergic reaction - Gluteraldehyde cytotoxicity - Frequent reinjection
Decellularized human dermis	Cymetra™	Discontinued	- Reduced foreign body response - Improved voice quality	- Discontinued for poor ease of use
Gelatin	Gelfoam^®		- Tissue tolerance- Early alternative to permanent medialization	- Pressurized injection through large-gauge needle- Short term use (2–8 weeks)
Hyaluronan	Juvederm^®, Restylane^®	Off-Label Use	- Reduces excess collagen I deposition- Improves voice quality	- Variable degradation time-Frequent reinjection
Hyaluronan-Silk	Silk Voice^®	FDA Approved	- Evidence of tissue remodeling- VF-like mechanical properties- FDA approved for VF applications	- Frequent reinjection
Teflon	Teflon/Polytef Paste	Discontinued	- Long-term injectable treatment	- Delayed inflammatory complications, particularly granuloma- Tissue morbidity upon removal

VF, vocal fold; FDA, United States Food and drug administration.

Restylane^® and Juvederm^® are produced from hyaluronan microparticles and reduce excess collagen I deposition as well as augment the glottal wave and voice quality.^49,50 These hyaluronan-based products are approved as dermal fillers but are used off-label for voice applications.⁴⁸ Treatment with HA-based products may improve phonation quality for at least 6 months.⁵¹

Calcium hydroxylapatite microparticles are a bioceramic injectable filler suspended in a carboxymethylcellulose gel.^47,51 Calcium hydroxyapatite, marketed as Radiesse^® and Prolaryn Plus^®, is FDA approved for injection laryngoplasty. The residence time of calcium hydroxylapatite is relatively long, up to 24 months.^47,51 However, calcium hydroxylapatite injection outcomes may be variable, and cause deposits in the superficial LP and deleterious foreign body response.^52,53 VF stiffening, impaired mucosal wave, edema, and hypervascularity were also reported with calcium hydroxylapatite.^48,52–54

A significant limitation of the aforementioned injectables is a lack of capacity to remodel tissue.^13,55,56 In contrast, autologous fat and silk hyaluronan show potential to stimulate tissue deposition for long-term efficacy.

Autologous fat grafts contain stem and endothelial cells that contribute to tissue remodeling and vascularization.^55–57 About 30–40% of excess volume is commonly used to offset resorption, but optimal volume is difficult to predict as resorption varies from 20% to 60%.^58–60 Overinjection of autologous fat can alter VF shape and mechanics, cause dysphonia, and require surgery to remove the excess.^57–59 Autologous fat may also induce acute and chronic inflammation, necrosis, and fibrosis.^55–57

Silk microparticles in a crosslinked hyaluronan carrier are new FDA-approved injection laryngoplasty biomaterials.^33,61,62 The silk microparticles incorporate bioactive properties to promote cellular infiltration and tissue remodeling.^33,62–65 Silk hyaluronan improves control over degradation, wherein the hyaluronan degrades more rapidly while silk microparticles are retained for up to 18 months.⁶³ While the remodeling effects are promising, about 32% of patients in a large clinical study required additional injection laryngoplasty procedures after a median of 106 days.^33,62 Serious complications including edema, dyspnea, and difficulty swallowing were reported with low incidence.^62,64

There remains an unfilled need for a minimally invasive, injectable VF biomaterial that effectively restores tissue structure and functions to reduce the requirement for repeated treatments. Tissue engineering strategies are required for injectable VF biomaterials with more permanent efficacy. In the following section, we review options for biomaterial composition, enhancing of bioactivity with biomaterial functionalization and encapsulation of growth factors and cells, and control of mechanical and degradation properties through composite materials and crosslinking.

Choosing Polymers for VF Injectables

Injectable VF biomaterials are fabricated from polymers derived from natural sources or synthesized in laboratories, each of which carry advantages and limitations (Table 2). Selecting polymers for VF injectables relies on desired properties, including bioactivity, cell response, structure, mechanics, and biodegradability.

Table 2.

Advantages and Disadvantages of VF Biomaterials Under Investigation

Biomaterial	Advantages	Disadvantages
Alginate^79,80,143	- VF-like mechanical properties - Readily modifiable with bioactive molecules - Antimicrobial	- No innate cell adhesion sites - Ionic crosslinking is heterogenous and uncontrollable - Ion leaching causes loss of stability
Collagen/Gelatin^10,46,72,160	- Predominates structure of ECM - Stimulates cell recruitment, migration, and proliferation	- Some types carry high risk of allergic reaction - High contractility - Rapid degradation
Chitosan⁷⁹	- Stimulates ECM deposition - Low immunogenicity - Antimicrobial - Reduces fibroblast proliferation and excess collagen deposition	- High swelling - Unstable at neutral pH
dECM^41,85,92,162	- Consists of innately regenerative ECM components - Stimulates native wound healing - Antifibrotic - Can be VF-specific	- Rapid degradation - Processing damages mechanical properties - Methods of producing can cause different immune and cell responses
Hyaluronan^72,121	- Sustained growth factor release - Improves vibration - VF-like viscoelasticity - Reduces deposition of fibrotic ECM molecules	- Rapid degradation - Lack of remodeling effects
Poly ε-Caprolactone^98,102,103	- Viscoelastic - Hydrolytic degradation - Biocompatibility	- Hydrophobicity - High mechanical properties compared to VF
Polyethylene glycol^99–101,154	- Biocompatibility - Low immunogenicity - Readily modifiable and crosslinkable - Tunable degradation	- Lacks innate bioactivity - Potential for pseudoallergy - Derivatives may be cytotoxic
Polylactic coglycolic acid^163,164	- Hydrolytic degradation - Biocompatibility	- Hydrophobicity - Acidic degradation byproducts
Polyurethane^105–107	- Design versatility - Durability	- Toxic impurities and degradation byproducts
Silk Fibroin^{33,62–64,69,78,165}	- Biocompatible - Mechanical tunability - Stimulates cellular infiltration and ECM remodeling	- Nonstandardized fabrication - Expensive - Stored in unstable aqueous solution

ECM, extracellular matrix; dECM, decellularized extracellular matrix; VF, vocal fold.

Natural materials

Natural polymers are selected for biocompatibility, bioactivity, porous ECM-like structure, or immunomodulatory capacity.^66,67 On the other hand, natural polymers have high variability, and may lack mechanical tunability or degrade before achieving permanent effect unless crosslinked or in composite.^66,68 In addition to those in clinical VF biomaterials, other natural polymers including alginate, chitosan, gelatin, and decellularized extracellular matrix (dECM) are under active investigation for favorable regenerative, mechanical, and degradation properties.

Collagen and gelatin are used in VF biomaterials for their capacity to promote cell adhesion, migration, and proliferation.^66,69–71 However, collagen-based hydrogels exhibit rapid degradation and poor mechanical and structural stability.^66,69 Despite the discontinuation of existing collagen and gelatin VF injectables, crosslinked and composite collagen and gelatin formulations are under investigation.^68,72,73

Hyaluronan and silk-based biomaterials are under development to improve treatment duration and remodeling capacity.^74–76 Hyaluronan provides a damping function within the VF, regulates cell function, and modulates inflammation with reduced immunogenicity compared to collagens.^44,77 Silk fibroin has well-established biocompatibility due to its use in sutures with highly tunable mechanical and structural properties.^69,70 However, biomedical grade silk fibroin is expensive, depends on nonstandardized hand processing and environmental conditions, and is sold in unstable aqueous solution.⁷⁸ Production variability leads to differences in bioactivity, degradation, and mechanical properties.⁷⁸

Sodium alginate is a biocompatible, antibacterial polysaccharide derived from brown seaweed that degrades into glucose-like monosaccharides that can be safely metabolized.^66,69,79,80 Alginate lacks innate cell adhesion sites, but its bioactivity and mechanical properties are highly tunable through variable molecular weight, chemical modification, and crosslinking.⁸⁰ In VF injectables, alginate is under investigation for cell and growth factor delivery, and as a composite to provide improved mechanics and durability to regenerative biomaterials.^74,76,79,81

Chitosan is a biocompatible, antimicrobial polysaccharide derived from the exoskeletons of arthopods with low immunogenicity.⁶⁶ In VF applications, chitosan has been shown to reduce fibroblast proliferation and excess collagen deposition.⁸² However, chitosan has a high swelling ratio, lacks elasticity, and is unstable at neutral pH, limitations that may be circumvented by crosslinking and use in composites.^71,79,83,84

dECM is produced by removing cells from tissue. Its composition of ECM proteins, glycosaminoglycans, and other regenerative molecules helps stimulate innate wound healing processes.^41,85,86 VF-specific dECM has high batch to batch variability and poor scalability, but commercial sources from larger organs such as small intestinal submucosa (SIS) show regenerative potential in the VF.^87,88 Both VF and SIS dECMs have been shown to downregulate fibrotic genes, reduce excess collagen I deposition, and stimulate LP restoration, though these hydrogels possessed weak mechanical properties.^86–90 dECM also has potential to stimulate other VF-relevant processes including recruitment of endogenous stem cells, angiogenesis, and formation of tissue-specific structures.^41,91–93

Synthetic materials

Synthetic polymers have standardized fabrication processes and high tunability. However, synthetic polymers do not possess innate bioactivity and may integrate poorly with surrounding tissue or produce toxic byproducts.⁶⁶ Also, polymer synthesis may involve toxic materials, and require extensive purification.^94,95 Synthetic polymers under investigation for VF applications include polyethylene glycol (PEG), poly ε-caprolactone (PCL), and polyurethane (PU).^79,96–98

PEG is a nonimmunogenic, amphiphilic polymer comprised of ethylene oxide monomers that degrades slowly through oxidative degradation.^99–101 PEG has extensive chemical modification capacity for modulating bioactivity, degradation, and mechanical properties.^96,97,99 However, some patients exhibit pseudoallergy to PEG, and PEG derivatives used at high concentrations or functionalized with molecules like acrylates may cause cytotoxicity.^100,101

PCL possesses favorable viscoelasticity at physiological temperature and is biodegradable over 2–3 years, though its hydrophobicity limits cell adhesion and proliferation.^102,103 The mechanical properties of PCL are high compared to soft tissues, but may be tuned by using composites.¹⁰³ In rabbits with VF paralysis, PCL microspheres in a Pluronic F127 carrier augmented VF function for 12 months.^98,102

PU has been used as a composite material for injectable VF biomaterials.¹⁰⁴ PU are flexible and durable block copolymers that can be synthesized with hydrolytic or enzymatic degradability depending on linkages between its constituent monomers.¹⁰⁵ PU degradation products and impurities retained after synthesis may be cytotoxic, though safer methodologies are under development.^105–107

Enhancing Bioactivity of Injectable VF Biomaterials

To better match material degradation with the wound healing rate, the bioactivity of injectable VF biomaterials may be enhanced by functionalization with bioactive motifs and encapsulating growth factors or cells within a bioactive material.

Conjugation with bioactive motifs

Conjugating polymers with biomolecular motifs can increase bioactivity.^96,98,108 Conjugation is performed via crosslinking the functional groups of the bioactive molecule and the polymer.^109,110 For example, arginylglycylaspartic acid (RGD), a peptide found in fibronectin and laminin, is a commonly used to enable cellular adhesion to polymers without innate adhesion sites.^108–111 Heparin, a glycosaminoglycan, is another common conjugation choice that is used to attract and sequester growth factors released by surrounding cells or control the release rate of encapsulated growth factors.^{96,98,112,113}

Injectable microparticles for VF augmentation were produced from PEG functionalized with maleimide and thiol groups that crosslinked by thiol-Michael addition.¹⁰⁸ The PEG-maleimide backbone was modified with thiol-functionalized RGD. After injection to rabbit thyroarytenoid muscle in annealing polymer solution, exposure to visible light produced a microporous annealed particle (MAP) hydrogel.¹¹⁴ RGD-conjugated PEG and the MAP structure provided a favorable environment for cellular adhesion, infiltration, and vascularization.^108,114

The Heparin and RGD-conjugated MAP hydrogel is the first to demonstrate tissue remodeling with long-term VF augmentation in an animal model.⁹⁶ The RGD-conjugated MAP hydrogel was further conjugated with thiolated heparin to sequester growth factors and improve tissue integration (Fig. 2A).⁹⁶ MAP hydrogels and a Restylane^® control were injected intramuscularly to a rabbit VF paralysis model. After 14 months, all six rabbits treated with MAP maintained augmentation. The biomaterial was replaced by collagen-rich neo-tissue with distinctive morphology compared to surrounding muscle.

FIG. 2.

Regenerative Effects of Enhancing the Bioactivity of VF Injectable Biomaterials. (A) RGD and Heparin-conjugated PEG-based MAP hydrogel. 1. MAP hydrogel structure showing thiol-ether crosslinks, MethMal annealing, and conjugation with RGD and heparin within microparticles. 2. Change in airway volume over 12 months. The MAP hydrogel maintained consistent augmentation while the airway volume increased with the hyaluronan control, indicating volume loss. 3. Picosirius red staining of neo-tissue produced by MAP hydrogel after 12 months under brightfield (left) and polarized light (right). Black box outlines neo-tissue. (B) SIS with encapsulated urinary stem cells in a rabbit model of VF injury promotes a regenerative immune response. 1. Proposed regenerative timeline following biomaterial injection. 2. Regenerative processes correlating to the proposed timeline. 3. Immunofluorescence and quantified integrated density (IntDen) of inflammatory (Cluster of Differentiation 68, Tumor Necrosis Factor alpha) and anti-inflammatory (CD206) markers in excised rabbit VF4. Immunofluorescence and quantified ratio of collagen I (red) to collagen III (green) in excised rabbit VF. *p < 0.05, **p < 0.01, ***p < 0.001, ****p < 0.0001, ns, not significant. Panel A is reproduced from Ref.⁹⁶ (Open Access: https://creativecommons.org/licenses/by-nc-nd/4.0/). Panel B is reproduced from Ref.⁸⁶ (Open Access: https://creativecommons.org/licenses/by-nc-nd/4.0/). MAP, microporous annealed particle; PEG, polyethylene glycol; RGD, arginylglycylaspartic acid; SIS, small intestinal submucosa; VF, vocal fold.

Encapsulation of growth factors

Growth factors are signaling cytokines secreted by cells that stimulate cellular activity, inflammation, and tissue remodeling.¹¹⁵ Basic fibroblast growth factor (bFGF) and hepatocyte growth factor (HGF) are used clinically to treat VF scarring due to fibroblast modulation and antifibrotic activity.^39,115 Free growth factors have several limitations including instability, rapid inactivation, and poor targeting resulting in safety concerns including tumorigenesis.^116,117 Encapsulating growth factors in biomaterials for intramuscular injection may slow release and improve target site retention.^74,98

Encapsulation in a VF injectable has been performed by calcium chloride crosslinking, trapping the HGF within the “egg-box” network of a hyaluronan-alginate hydrogel.⁷⁴ In vitro, HGF release was sustained over 3 weeks as the porous “egg-box” facilitated slow diffusion.^74,118 When the hydrogel was injected into injured VF LP, vibrational and viscoelastic properties improved compared to controls treated with HGF alone.

Growth factors may also be bound to injectable microspheres with growth factor binding sites.^98,119 Gelatin microspheres with bFGF have reduced scarring when injected to injured rabbit and canine VF LP.^119,120 Heparin-immobilized PCL microspheres with bound bFGF and HGF in a Pluronic F127 carrier showed ECM deposition and muscular regeneration 1 month after injection to a VF paralysis rabbit model, a regenerative effect that could not be achieved with PCL alone.⁹⁸

Cell delivery

Codelivery of cells with a biomaterial may provide a synergistic effect on VF regeneration. In VF animal models, stem cells have been shown to remodel fibrotic tissue and restore mechanical function.^121–123 In a nonrandomized clinical study, patients received injections of autologous mesenchymal stem cells (MSCs) to the LP and muscle following surgical resection of scarring.¹²² Up to 75% of patients (n = 16) showed improved VF function. Encapsulating cells in a biomaterial improves cellular viability and retention in the target tissue and provides the cells with bioactive and mechanical cues to enhance VF augmentation.^122,123

Encapsulation of stem cells in hyaluronan has improved retention in VF LP or muscle and reduced fibrosis in rat and rabbit models.^121,124 However, a hyaluronan-MSC group performed comparably to MSCs alone in a preliminary clinical study.¹²² Encapsulation in a biomaterial with remodeling capacity or longer-term stability improve outcomes.

Delivery of urine-derived stem cells (USCs) in a SIS dECM hydrogel was recently explored in a heat-injury rabbit VF LP model (Fig. 2B).⁸⁶ SIS dECM-only, USC-only, and SIS-USC groups all demonstrated lower inflammatory marker expression (CD86, TNF-α) and reduced collagen I/collagen III ratio compared to LP without treatment after 2 and 8 weeks, respectively. The SIS dECM-only and USC-only groups induced comparable responses, while the SIS-USC group induced an enhance synergistic effect on VF regeneration.

Modulating the Mechanical, Physical, and Degradation Properties of VF Biomaterials

Mimicking in vivo VF mechanical properties and degradation at a rate that matches regeneration are necessary for biodegradable materials to permanently augment VF function. Alongside biomolecular cues, the mechanical (stiffness, viscoelasticity) and physical (porosity, swelling) properties of a biomaterial help direct cells to generate neo-ECM characteristic of VF.¹²⁵ VF mimicry is complicated by its layered structure and varied thickness.^3,126 VF exhibit compressive moduli of 3.9–5.4 kPa on the medial LP surface, 2.7–3.2 kPa on the superior LP surface, and 1.3–2.7 kPa on the thyroarytenoid muscle.¹²⁶ A range representative of the muscle or LP is typically targeted in biomaterial design.

Degradation properties are also critical, as bioactive materials cannot achieve permanent VF augmentation if they degrade before tissue remodeling is complete.⁶⁶ Despite the bioactivity and VF-mimicking storage modulus of SilkVoice^®, its stability varies, and reinjection are still frequently required.^62,63 The use of composite materials and chemical crosslinking can tune biomaterial properties to achieve VF mimicry and controlled degradation.

Composite VF biomaterials

Natural, ECM-derived polymers with innate bioactivity may be combined with more stable polymers and the composition ratio and crosslinking method tuned to produce an injectable composite that better mimics human VF than either alone (Table 3).^72,81

Table 3.

Advantages and Disadvantages of Conventional Crosslinkers

Crosslinker	Advantages	Disadvantages
1,4-butanediol diclycidyl ether¹³¹	- Reliable for clinically approved in situ gelling materials - Excess removed prior to gel formation - Slows degradation rate compared to uncrosslinked hyaluronan	- Toxic in concentrations greater than 2 ppm - Reinjection required due to degradation
Dialdehydes (glutaraldehyde, glyoxal)^{43,83,119,132–136,166}	- High crosslinking efficiency - Soluble in water and organic solution - Mechanical tunability - Versatility with varied materials - Resistance to thermal and enzymatic degradation	- Toxic at low concentrations - Degrades by hydrolysis - Methods for reducing gluteraldehyde exposure may lower crosslinking efficiency - Unsuitable for in situ gelation due to protein crosslinking
Genipin^{136,138–141}	- Low cytotoxicity compared to other conventional crosslinkers - Mechanical tunability - Crosslinking specificity to primary amines - Low swelling ratio - Resistance to enzymatic degradation	- Low cellular infiltration - Cytotoxicity increases at physiological temperature - Unsuitable for in situ gelation due to protein crosslinking
1-ethyl-3-(3-dimethylaminpropyl) carbodiimide hydrochloride (EDC)^135,136,167	- Rapid gelation - Neutral reaction conditions - Forms favorable network structure for cellular attachment	- Rapid degradation - Toxic side products inhibit use in injectables - Unsuitable for in situ gelation due to protein crosslinking

For instance, thiolated hyaluronan was crosslinked with PEG-diacrylate (PEGDA) to produce a hydrogel with tunable viscoelasticity for VF injection, and supplemented with collagen I collagen III to achieve a native-like structure.⁷² Inclusion of collagens was necessary to retain encapsulated fibroblasts and increased the storage modulus from 0.16–0.21 kPa to 0.65–1 kPa within the 0.3–1.2 k range of adult VF LP.^72,127 HA-PEGDA prevented the fibroblast-mediated contraction associated with collagen-only hydrogels. The inclusion of natural materials and a synthetic polymer improved tissue-mimicking and bioactivity.

Table 4.

Advantages and Disadvantages of Click Crosslinking Methods

Crosslinking method	Advantages	Disadvantages	Vocal fold applications
Thiol-Michael^{72,75,149,151}	- Rapid reaction kinetics- High conversion rate	- Requires external catalyst-Potential for multiple products- Potential for heterogeneity and reduced stability	- RGD-conjugated PEG “MAP” hydrogel with tissue integration¹⁰⁸-HA-PEGDA hydrogel with collagens I and III⁷²
Thiol-ene^152,153,168	- Rapid reaction kinetics - Stable reactants - Potential for visible light catalyzation	- UV light may induce toxicity - Ester bonds subject to hydrolysis	- Trilayered Hyaluronan in vitro model of the Lamina Propria¹⁵³
SPAAC^97,154	- Selectivity prevents reaction with biomolecules - No external catalyst	- Cyclooctene reactants are difficult to synthesize - Low reaction yield	- four arm PEG functionalized with azide and dibenzocyclooctyne⁹⁷
IEDDAC^156–158	- Crosslinks at room and physiological temperature - Selectivity prevents reaction with biomolecules - No external catalyst - Nitrogen gas only side product	- Typical alkenes may react slowly (norbornene) or lack aqueous stability (TCO)	- Core-Shell PCL-Hyaluronan in vitro model for modulating HVFF inflammatory response¹⁵⁷ - MSC differentiation into fibroblasts with delayed stiffening of Hyaluronan hydrogel by secondary click reaction¹⁵⁸

HVFF, Human Vocal Fold Fibroblast; IEDDAC, inverse electron demand Diels–Alder cycloaddition; MSC, mesenchymal stem cell; MAP, microporous annealed particle; PCL, poly ε-caprolactone; PEG, polyethylene glycol; RGD, arginylglycylaspartic acid; SPAAC, strain promoted azide-alkyne cycloaddition; TCO, transcyclooctene.

The impact of material composition on mechanical properties is illustrated by the varied properties of calcium-crosslinked alginate microspheres in composite with poly-L-lysine or chitosan (Fig. 3A).⁷⁹ The VF Young’s modulus by nanoindentation ranges between 0.5 and 6 kPa, matching the Alginate-only (∼3 kPa) and Alginate-poly-L-lysine (∼2 kPa) microspheres but not alginate-chitosan (∼12 kPa).^79,128,129 Low swelling ratios are favorable for stability and biological response predictability. Alginate-poly-L-lysine microspheres showed the lowest ratio at 9% after 48 h, while alginate-only microspheres swelled by 34%. Alginate-chitosan microspheres swelled by over 80%, contraindicating use due to risk of airway obstruction.

FIG. 3.

(A) Alginate-based microspheres as injectable VF biomaterials. 1. Calcium Ion Crosslinking with “Egg-Box” Structure. Created with BioRender.com. 2. Young’s modulus and 3. Swelling ratio of ● alginate, ● alginate-poly-L-lysine, and ● alginate-chitosan microspheres. (B) Alginate-chitosan VF injectable hydrogel with tunable mechanical and physical properties. 1. Swelling Ratio 2. Hydrolytic degradation and 3. Cyclic stress-relaxation between 300% and 1% strain at varying ratios (1:1, 2:1, 1:2) of alginate to chitosan. 4. Remodeling of Collagens I and III in response to the 1:1 alginate-chitosan hydrogel compared to hyaluronan and untreated injury controls. Panel A is reproduced from Ref.⁷² Panel B is reproduced from Ref.⁷⁹ Panel C is reproduced from Ref¹³⁰ (Open Access: https://creativecommons.org/licenses/by-nc-nd/4.0/).

Alginate-chitosan hydrogels exhibited more favorable swelling characteristics when crosslinked with Schiff base and borate ester bonding and visible light methacrylate polymerization (Fig. 3B).¹³⁰ Tuning the alginate-chitosan ratio revealed that a 1:1 ratio was most stable with a swelling ratio of ∼26% after 48 h compared to 1:2 (∼35%) and 2:1 (∼45%) ratios. The 1:1 alginate-chitosan hydrogel retained the greatest mass by at least 1.5 times after a 28-day hydrolytic degradation test. Intramuscular injection of the 1:1 hydrogel in a VF paralysis rabbit model better maintained volume, exhibited a smaller glottal gap, and stimulated neo-ECM production after 12 weeks compared to a hyaluronan control.¹³⁰ In sum, composite biomaterials provide vast opportunities for tuning biomaterial properties to produce a synergistic effect on regeneration.

Conventional crosslinking of VF biomaterials

Crosslinking method fundamentally impacts network structure, mechanical properties, and degradation rate of a biomaterial. Desirable properties for crosslinkers in VF injectables include reaction specificity, nontoxic byproducts, and capacity for safe in situ crosslinking.

Chemical and ionic crosslinking are used to fabricate VF injectables, such as 1,4-butanediol diclycidyl ether (BDDE) in Juvederm^®, Restylane^®, and SilkVoice^®.¹³¹ BDDE slows enzymatic degradation by obstructing hyaluronidase cleavage sites. Although BDDE is toxic, it is removed before patients are exposed.¹³¹ However, BDDE-crosslinked hyaluronan degrades without repairing tissue, necessitating reinjection.

Dialdehydes such as glutaraldehyde and glyoxal react with amine groups and are used to strengthen mechanical properties and provide resistance to thermal and enzymatic degradation in collagen, gelatin, and glycol-chitosan VF injectables.^83,119,132 However, dialdehydes are cytotoxic at low concentrations and the reaction is difficult to control with multiple possible reaction products.^43,133–136

Carbodiimide crosslinking with 1-Ethyl-3-(3-dimethylaminopropyl) carbodiimide (EDC) and N-hydroxysuccinimide (NHS) couples amine and carboxylic acid groups and is used to improve mechanical stability. EDC-NHS forms a favorable network structure for cellular integration but crosslinks DNA, proteins, and cell surface molecules and produces a toxic isourea byproduct.^135–137

Genipin crosslinking is less cytotoxic and reacts specifically with primary amines, making it useful for controlling crosslinking density.^138,139 In VF applications, genipin-crosslinked gelatin hydrogels exhibited low swelling ratios, slow enzymatic degradation, and improved mechanical strength compared to carbodiimide-crosslinked hydrogels.⁶⁸ However, cytotoxicity and hinderance of cell infiltration in genipin-crosslinked biomaterials may increase at physiological temperature, and there is a shortage of clinical safety evaluation.^138–141

Dialdehyde, carbodiimide, and genipin crosslinkers may be safe for microparticles that can be sufficiently purified prior to injection but are not recommended for in situ gelation to prevent cytotoxicity and crosslinking of ECM proteins in the native tissue.¹³⁶

Ionic crosslinking with calcium chloride, used to fabricate alginate and chitosan biomaterials, is rapid and biocompatible.^74,79,142 However, limitations include rapid, uncontrollable crosslinking causing heterogenous bond distribution and difficulty of injection and ion leaching resulting in loss of structural stability.¹⁴³

The side reactions, cytotoxicity, and heterogeneity inherent to conventional crosslinkers led to the development of bioorthogonal click chemistry, which may be more suitable for in situ gelling VF injectable applications.¹⁴⁴

Bioorthogonal click chemistry in VF biomaterials

Bioorthogonal click chemistry describes selective reactions that link molecules at high yield, do not require purification, and proceed safely in vivo.^144–148 Bioorthogonal click chemistries for in situ gelling VF injectables include Thiol-X, strain promoted azide-alkyne cycloaddition (SPAAC), and inverse electron demand Diels–Alder cycloaddition (IEDDAC) (Table 4).^144,148

The MAP and HA-PEGDA hydrogels discussed previously were fabricated by the thiol-Michael thiol-X reaction.^{96,144,148,149} The thiol-Michael reaction occurs through a thiol attacking an electron-deficient alkene, catalyzed by a base, primary amine, or phosphine.¹⁴⁹ A limitation is that catalysts may cause side reactions including disulfide bond formation, acrylate dimerization, and Aza-Michael addition.^149,151

The light-catalyzed, radical-mediated thiol-ene reaction is a thiol-X without side reactions that proceeds between a thiol and an electron-rich alkene.¹⁵² It has not yet been used to fabricate a VF injectable, but has been used to model the VF LP by reacting hyaluronan-norbornene of increasing concentrations (1–2.5%) with a UV-activated dithiol to produce three layers with differing stiffness. (Fig. 4A).¹⁵³ A bench test showed that the scaffold could oscillate within the normal human speaking range (about 150 Hz).¹⁵³ To prevent UV-induced tissue damage, VF injectables may be produced using visible light-catalyzed thiol-ene chemistry.¹⁵²

FIG. 4.

Click Chemistry in VF Tissue Engineering (A) Trilayer hyaluronan hydrogel produced by thiol-ene click chemistry vibrates to reproduce lamina propria structure. 1. Compressive modulus of the trilayer hydrogel is reduced by collagenase exposure. Red, yellow, and blue layers represent measurement of the composite, but do not have quantitative value within Figure. 2. Sound pressure level and frequency for air flow at 95 and 115 mL/s. * fundamental frequency * harmonics. + subharmonics 3. Vibration of trilayer hydrogels under airflow. *p < 0.05. (B) Injectable PEG hydrogel fabricated by SPAAC for permanent VF augmentation. 1. Storage and loss moduli of a Click 4% PEG hydrogel. 2. Hydrogel volume 1 month after injection in mice. 3. Location of VF at maximum and minimum amplitude of vibration. Lower values equate to more favorable results. 4. Hemotoxylin and eosin staining of VF 16 weeks after injection. Blue arrow: Calcium-hydroxylapatite migrated to muscle tissue. Black arrows: vasculitis in adipose tissue. *p < 0.05. (C) Hyaluronan hydrogels with delayed stiffening fabricated by IEDDAC for differentiation of MSCs into fibroblasts. 1. Schematic representing hydrogel fabrication process for each crosslinking condition. nLP or “newborn lamina propria” is homogeneously crosslinked only with norbornene. sLP0 or “stiff lamina propria 0” is homogenously crosslinked with added TCO on Day 0. sLP7 or “stiff lamina propria 7” is homogeneously crosslinked with added TCO on Day 7. mLP or “mature lamina propria” has a stiff top layer and soft bottom layer produced by stopping the TCO reaction after 4 h. 2. Expression of fibroblast gene markers by MSCs on Day 14. 3. Expression of extracellular matrix (ECM) remodeling enzyme gene markers by MSCs on Day 14. 4. Expression of ECM gene markers by MSCs on day 14. 5. Secretion of regenerative growth factors VEGF and HGF by MSCs on day 14. Panel A is reproduced from Ref.¹⁵³ Panel B is reproduced from Ref.⁹⁷ Panel C is reproduced from Ref.¹⁵⁸ MSCs, mesenchymal stem cells; PEG, polyethylene glycol; SPAAC, strain promoted azide-alkyne cycloaddition; IEDDAC, inverse electron demand Diels–Alder cycloaddition; VF, vocal fold.

SPAACs involve the reaction between strained ring cyclooctynes and aliphatic azides to form a triazole bond.¹⁵⁴ SPAACs do not require an external catalyst, but disadvantages include difficulty of cyclooctyne synthesis and low reaction yield.¹⁵⁴ A VF hydrogel was produced from by SPAAC from four-arm PEG functionalized with azide and dibenzocyclooctyne (Fig. 4B).⁹⁷ The storage modulus of the PEG hydrogel were low compared to native VF, less than 100 Pa. Nevertheless, the click PEG hydrogel showed increased stability compared to Radiesse^® in a VF paralysis rabbit model after 16 weeks, albeit minimal cellular infiltration. Additionally, the PEG hydrogel remained within the thyroarytenoid muscle, while Radiesse^® migrated to surrounding tissue and caused foreign body giant cell accumulation and vasculitis.⁹⁷

IEDDACs are rapid reactions between dienes and alkenes or alkynes that forms a six-membered ring, do not require catalysts or react with moieties in tissue, and can occur at room and physiological temperatures.^154–156 IEDDACs have not yet been used for VF injectables, but have been used to model fibroblast differentiation during LP maturation.^157,158

MSCs were first encapsulated in soft hydrogels formed by reaction between hyaluronan-tetrazine and a norbornene-functionalized MMP-degradable peptide, then stiffened by secondary reaction with hyaluronan-transcyclooctene after7 days (Fig. 4C). Compared to hydrogels without secondary crosslinking, MSCs within the delayed-stiffening hydrogels upregulated expression of fibroblast (Fibroblast Activation Protein, Ferroptosis Suppressor Protein 1) and neo-ECM (collagen I, collagen III, fibronectin) markers matrix metalloproteinases, hyaluronan synthase 3, and regenerative growth factors (VEGF, HGF). To achieve delayed stiffening in an injectable biomaterial, a slow secondary crosslinker could be included in a rapidly gelling hydrogel. Overall, click chemistry shows great potential in VF tissue engineering for producing bioactive, injectable composites with tunable mechanical, physical, and degradation properties.

Discussion and Future Prospects

Injection laryngoplasty provides a minimally invasive, relatively low-cost procedure. However, there remains a need for a reconstructive injectable biomaterial capable of permanently augmenting VF function. To achieve this goal, a balance is required between biodegradability and sufficient biomaterial longevity to stimulate VF repair. Neither biomaterials requiring reinjection nor nondegradable materials are ideal for long-term VF augmentation.

Standard injection laryngoplasty may be capable of permanently restoring VF function if a regenerative biomaterial is replaced with neo-ECM as it degrades, but cannot treat fibrotic conditions in the superficial LP.⁹⁶ Surgical techniques to preventing vibratory disruption in intramucosal biomaterial injection are under investigation alongside biomaterials with scar-remodeling potential.^39,82 Ultrafast laser surgery has been proposed to aid with localization by subsurface ablation, i.e., the precise removal of tissue to form a void within the LP.^36,159 When subsurface ablation was performed in a preliminary canine study (n = 1), biomaterial localization was observed within the void indicating the surgical techniques may aid in retention at the target region by creating a defect for the injectable biomaterial to fill.¹⁵⁹

Intramucosal injection with autologous fat has been tested clinically with variable outcomes.⁶⁰ Researchers have begun to explore the possibility of using dECM hydrogels to remodel scars in the LP, though these studies have not yet reached the clinical phase.^41,86 In future development of intramucosal biomaterials, it is critical to mimic the native mechanical and physical properties of the LP, as mechanical mismatch is one of the reasons superficial injection of existing VF biomaterials disrupts vibratory function.

The need for minimally invasive permanent VF restoration may be fulfilled by injectable VF biomaterials with improved stability that stimulates reconstruction. Promising results have been achieved in vitro and in animal models for VF-mimicking composites of mechanically tunable polymers and regenerative ECM-based proteins or polysaccharides. Bioorthogonal click chemistry can enable improved safety for in situ gelling crosslinked biomaterials and incorporate features such as delayed stiffening to stimulate cellular behavior toward wound healing. These advanced tissue engineering techniques will greatly improve the efficacy of minimally invasive intramucosal and injection laryngoplasty treatments.

Footnotes

Authors’ Contributions

M.B.: Conceptualization (lead), writing—original draft (lead). H.O.: Writing—review and editing (supporting). M.Y.: Writing—Review and Editing (supporting). M.T.: supervision (supporting), writing—review and editing (supporting). N.Y.K. L.-J.: Supervision (lead), conceptualization (supporting), writing—review and editing (lead).

Author Disclosure Statement

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

Funding Information

This review was supported by the National Sciences and Engineering Research Council of Canada (RGPIN-2018–03843, RGPIN-2024-04235, and ALLRP 548623-19), Canada Research Chair research stipend (M.T. and N.Y.K.L.-J.), and the National Institutes of Health (R01 DC-018577-01A1). The presented content is solely the responsibility of the authors and does not necessarily represent the official views of the above funding agencies.

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