Chapter 1: Injectable Smart Hydrogels: Introduction, Preparation, and Applications Free

Hydrogels are crosslinked three-dimensional (3D) networks of hydrophilic polymers that can imbibe copious amounts of water and mimic the extracellular matrix (ECM). These materials continue to find a wide range of therapeutic applications due to their unique physicochemical properties such as softness, biocompatibility, and water permeability, as well as their smart properties such as stimuli-responsiveness (smart hydrogels) and self-healing characteristics.¹  Initially, these hydrogels have been formed outside the body and implanted using invasive surgical procedures.^1,2  However, influential work in the late 1990s has provided researchers with a library of techniques, whereby hydrogel formation can occur in situ after the delivery of a pre-hydrogel through a needle via local crosslinking through transdermal light-induced photopolymerization.³  This approach provided an avenue to develop injectable hydrogels suitable for minimally invasive therapies, fill complex tissue defects, and induce the regeneration of damaged or lost tissues.⁴  They can readily take the shape of a treated cavity. Moreover, various therapeutic molecules and even cells can be incorporated by simply mixing with the precursor solution prior to injection. Hydrogels with shear-thinning and self-healing properties are also recognized as injectable biomaterials as they can be injected directly in the gel state.^5,6  Amongst the several possible applications are tissue engineering, cell transplantation therapies, regenerative medicine, implants, drug delivery, and cancer treatment.^7–12 

In the past few decades, several hydrogels of natural and synthetic origins have been introduced for use in injectable therapies. A wide variety of polysaccharides, proteins and synthetic polymers including chitosan, hyaluronic acid (HA), alginate, collagen, gelatin, poly(ethylene glycol) (PEG) and poly(lactide-co-glycolide) (PLGA) materials have been utilized to fabricate injectable hydrogels.^13–16  Furthermore, hydrogels can be functionalized with growth factors such as transforming growth factor-beta (TGF-β),¹⁷  insulin-like growth factor-1 (IGF-1),¹⁸  platelet-derived growth factor (PDGF),¹⁹  fibroblast growth factor (FGF),²⁰  vascular endothelial growth factor (VEGF)²¹  and bone morphogenetic protein-2 (BMP-2)²²  to achieve required cell proliferation and differentiation.

An injectable hydrogel for biomedical applications needs to satisfy certain important criteria. For example, (1) the sol–gel transition should be processed under mild conditions; (2) the gelation is preferred to occur in time to void the outflow of materials to the surrounding tissue; (3) the formation and degradation of the hydrogel should not require toxic reagents or produce toxic products. The safety of injectable hydrogels is a major concern that should be monitored, as it helps to determine factors for selecting appropriate materials for injectable therapeutics. Other parameters include sufficient biodegradability, biocompatibility, low cytotoxicity, etc.²³  In particular, a smart hydrogel can undergo phase transitions between solution and gel states in response to specific environmental cues, such as changes in temperature, pH, or ionic strength, or the presence of specific molecules. Under the influence of an external trigger, the interactions between the polymer chains and the surrounding solvent can be altered. The driving force of smart hydrogel phase changes from the traditional passive to an active material can be designed according to the needs of specific application sites, such as changes in the pH value. Thermal smart injectable hydrogels can be delivered as liquids via injection, which gel inside the body due to their shear-thinning properties and their ability to form gel upon relaxation post injection. The viscoelastic characteristics of these smart injectable hydrogels enhance their value in surgical applications by facilitating tissue adhesion, preventing fluid leakage, and reducing the risk of infection.

This chapter will cover the fundamental chemistry, synthesis, formation, and physicochemical and biological properties of injectable smart hydrogels. We also explore various mechanisms and strategies employed in designing and synthesizing injectable hydrogels, and their key biomedical applications.

Several mechanisms have been proposed for the development of injectable hydrogels. The basic difference between the preparation of injectable gels and in situ forming gels with the same crosslinking mechanism is the specificity on the control of the gelling kinetics by which the sol–gel transition of an injectable hydrogel should be achieved within the limited time interval. Faster or slower gelation kinetics can affect the injectability/mass transfer of the bulk gel. Based on the physical characteristics injectability and gelation, hydrogels can be categorized into three main groups: (a) in situ forming hydrogels, (b) injectable hydrogels, and (c) injectable particles.

In situ forming hydrogels are developed based on the concept of sol–gel; in this process, the materials can be prepared as flowable solution and can be converted into hydrogels upon injection into the body (Figure 1.1a).^24–26  These gels can be categorized according to the process that facilitates the process of liquid to hydrogel transition. Chemical crosslinking (e.g., covalent bond, light enzyme, or click crosslinking) and physical crosslinking (e.g., hydrogen, electrostatic or hydrophobic interactions) are the most commonly employed methods for the preparation of in situ forming hydrogels. Studies indicate that gelation through reversible covalent bonding allows the incorporation of cells and therapeutic drugs during gel formation and their controlled release is achieved by applying desired stimuli that lead to hydrogel degradation.^27,28  Moreover, hydrogels prepared by reversible covalent bonds have the ability to absorb aqueous solutions, stimuli-controlled mechanical properties similar to ECMs, and quick gel formation making them suitable for use as tissue engineering scaffolds.^29,30  Among various stimuli, changes in temperature can be considered as the most important stimulus that prompts hydrogel formation after injection into the human body. In situ forming hydrogels can be formed irreversibly via chemical reactions between the functional groups present in the hydrogel or reversibly by physical stimuli such as temperature, light, electricity, pressure, acoustic waves, etc.^31,32 

The formation of in situ forming hydrogels through electrostatic interactions is currently a major area of research for many biomedical applications.^33,34  Based on the concept of electrostatic interaction, the cationic and anionic charges present in the hydrogel network are the main driving force behind in situ formation of hydrogels. To induce electrostatic interactions, polymers must be ionized or undergo electrical changes that can lead to electrostatic interaction between opposite charges to create ionic crosslinking (Figure 1.1b). Basically, electrolytes are monomeric constituents with electric charge, while polyelectrolytes are ionic polymers with high molecular weight prepared from electrically charged monomeric molecules. The properties of electrostatic interaction in the solution depend on many parameters such as the ratios between opposite charges, mobility or flexibility of polyelectrolyte materials, pH, temperature, etc. Recent studies have shown that electrostatically interactive in situ forming hydrogels can be produced from natural polymers such as chitosan, cellulose and alginate.^33,35  Recently, Jafari et al.³⁶  produced double crosslinked chitosan/alginate injectable hydrogels via a green and sustainable approach for regenerating irregular-shaped defects, particularly wound healing applications. The self-healing capability of the hydrogels was observed due to the inter-polymer complex formed by the electrostatic interaction between the negatively charged alginate’s carboxyl group and the positively charged chitosan’s amino group.

The main challenge in the practice of in situ forming hydrogels is the short phase transition time, from solution to gel (sol-to-gel phase transition) in just minutes or even seconds.³⁹  Several factors affect the gel formation, such as gelation time, gelation temperature, pH, mechanical strength and the fabrication method. In addition, biocompatibility, biodegradability and minimal immune response are other deciding factors that govern in vivo applications of in situ forming hydrogels.³⁷  Importantly, hydrogel solidification must occur slowly to allow homogeneous transition of the solution to become a hydrogel (Figure 1.1c).³⁸  Other benefits include: (a) no surgical procedure is needed, (b) easy incorporation of biological components through mixing and (c) shear thinning properties of hydrogels that enable easy adaptation to the geometry of the target site.¹ 

Injectable gels belong to the second category, which is slightly different from the category of in situ forming gels. In this case, the gels are formed ex vivo, but they can be injected into the target site due to their shear-thinning properties and their ability to regain their hydrogel form upon post-injection relaxation. The concept is illustrated in Figure 1.1b. Injectable gels are able to flow due to shear stress, commonly induced by using needles. The process of return of the gel to its original intact gel form is known as self-healing. The mechanism of shear-thinning and self-healing varies in different systems. For example, a fast-gelling injectable gel based on HA and methylcellulose (MC) was developed as a drug delivery system for localized release of growth factors to aid spinal cord repair.⁴⁰  The HAMC gel was formed prior to injection and the gel strength increased after injection due to increased temperature in the body. In another study, Sun et al. proposed a different strategy to transform the “solid into gel” by the dispersion of large solid materials into alginate-Ca²⁺ gel to fabricate a functional injectable gel.^41,42  Recently, they have incorporated magnetic metals/metal oxide powders into alginate-Ca²⁺ to formulate injectable magnetic gels for in vivo magnetic–thermal tumor therapy.⁴³ 

The third category of injectable hydrogels is injectable particles, where the particles are immersed in a liquid and the mixture is injected. Particle size can be on the nano-, micro or macro-scale depending on the desired outcome. For instance, the particle size of dermal fillers is positively correlated with the depth of injection into the skin and their efficacy.⁴⁴  Typically, these particles can aggregate or self-assemble in situ to form an intact gel after injection (Figure 1.1c). In one study, injectable particles based on fluorescent mesoporous silica nanoconjugates end-capped with HA were fabricated to target cancer.⁴⁵  In this design, mesoporous silica served as the drug carrier and HA played multiple roles as a drug-release cap, a tumor-targeting point and a responsive gel matrix. The developed nanoconjugates were effective in transformation of particles to a hydrogel at the tumor site, and sustained drug release in response to specific enzymes in tumor cells, thus having potential for use in targeting tumors. PLGA colloidal gels have also been studied as injectable controlled release systems to deliver drugs such as dexamethasone (DEX), which is used as an anti-inflammatory and immunosuppressive agent.⁴⁶ 

Based on the gelation mechanism, injectable hydrogels are classified into two categories: physically crosslinked and chemically crosslinked hydrogels.^47,48  Physically crosslinked hydrogels are formed due to the changes in physicochemical parameters such as temperature, pH, ionic strength, glucose concentration or mechanical stress. These parameters function as stimuli for polymer conformation changes, which prompt phase separation and enable physical crosslinking by the aggregation of polymer chains.⁴⁹  Physical crosslinking mechanisms comprise ionic crosslinking, hydrophobic interactions, host–guest interactions and π–π stacking interactions,^50–52  while chemical crosslinking comprises different reactions including Michael addition, Schiff base formation, Diels–Alder reaction, click chemistry reaction and enzymatic crosslinking reaction.^53–56  Various mechanisms are applied to synthesize shear-thinning and self-healing injectable hydrogels by either dynamic covalent bonding (Diels–Alder reactions, Schiff base formation, etc.) or non-covalent interactions (hydrogen bonding, electrostatic interaction and hydrophobic interaction) or multi-mechanism interactions.^5,57,58 

The preparation of an ionic crosslinked hydrogel involves the mixing of ionizable polymers with counter ions. The properties of the resulting hydrogel depend on different factors such as polymer structure and composition, reaction temperature, time, and pH. By changing the concentration of ions, temperature or pH, the dynamics of sol–gel transition is altered, which in turn governs the injectability of the hydrogel.^59,60  The mechanical properties of the hydrogel can further be controlled by the molecular weight of the constituting polymer and the crosslinking density, e.g., by tuning the concentration of the polymer or counter ions. For example, alginate, which has plenty of carboxyl groups, can be crosslinked with a cationic donor such as calcium chloride (CaCl₂) to form an egg-box like conformation. The mechanical properties of the alginate hydrogel were enhanced by increasing the concentration of CaCl₂ in the matrix.⁶¹ 

Hydrogen bonding has dynamic nature and breaks down at higher temperature. Hence, H-bonding can be considered as an appropriate crosslinking strategy for the preparation of injectable hydrogels. Basically, individual hydrogen bonds in hydrogel systems are usually weak but can achieve mechanically strong hydrogels by synergy resulting from combining other types of chemical interactions. Different materials can be incorporated into hydrogen bond-based networks to improve the mechanical properties of the hydrogel. Moreover, the hydrogel can be endowed with many properties including self-healing, conductivity and thermoplasticity. In a recent study, this technique was adopted to produce an injectable conducting hydrogel based on poly(3,4-ethylenedioxythiophene):poly(styrene sulfonate) (PEDOT:PSS)/guar slime that can be utilized for wound healing applications.⁶²  A disadvantage of H-bonding crosslinked hydrogels is their poor stability in water, which causes dissociation of H-bonds in the polymer segments.

Amphiphilic polymers can be engineered to fabricate injectable hydrogels by the association of hydrophobic moieties as physical crosslinking points.^63,64  Amphiphilic polymers are synthesized by block or random copolymerization or the modification of hydrophilic macromolecules with hydrophobic pendent groups. Polymeric amphiphiles may have either a low critical solution temperature (LCST) or an upper critical solution temperature (UCST), which ensures the possibility of using them to form sol–gel structures by the change of temperature.⁶⁵  Polymeric amphiphiles, such as Pluronics, can exhibit either a LCST or an UCST, allowing for the formation of sol–gel structures through temperature changes.⁶⁵ 

Host–guest interactions are a type of noncovalent interaction with broad applications in the fabrication of injectable hydrogels. A host molecule interacts with a guest molecule to bind internally to form a complex structure with physical crosslinking that helps to design an injectable hydrogel. Typical host molecules include cyclodextrin (CD), cucurbituril and crown ether,^66,67  which can be linked or grafted onto polymer backbones as pendent groups. CDs are also the most common host molecules capable of coupling with linear polymers such as PEG, modified conducting monomers such as PEDOT and small guest moieties such as adamantine and azobenzene to form complex structures for injectable hydrogels.

π–π stacking interactions are a spatial arrangement of aromatic compounds that occurs between π-electron-rich groups and π-electron deficient groups, resulting in the flow of electron clouds from the electron-rich side to the electron-deficient side in order to increase the electron density on the deficient side.⁶⁸  There are mainly two stacking modes: (a) plane-to-plane and (b) edge-to-plane. Amino acids such as phenylalanine, tyrosine and fluorenylmethyloxycarbonyl (Fmoc) groups are most commonly used to formulate π–π stacking in hydrogels.^69,70 

Michael addition is a type of addition reaction where a nucleophile (a molecule or an ion with a pair of electrons to donate, such as amines, thiols, enolates, etc.) adds to an α,β-unsaturated carbonyl compound (a compound with a carbon–carbon double bond adjacent to a carbonyl group), creating a new carbon–carbon bond.^71,72 

The click crosslinking procedure involves two types of functional groups that can react with each other to form covalent bonds: (A) thiol–ene reaction: this reaction occurs between a thiol group (–SH) and an alkene group (–C═C–), forming a thioether bond (–S–C–C–). Typically, this reaction is initiated by either light or heat to generate free radicals. For example, HA–thiol and HA–acrylate can undergo a thiol–ene click reaction to produce crosslinking.^74,75  (B) Diels–Alder reaction: this reaction occurs between a diene group (–C═C–C═C–) and a dienophile group (–C═C–Z–Z–), forming a cyclohexene ring. The dienophile group usually contains electron-withdrawing groups (Z), such as furan or maleimide, to increase the reactivity.^76,77  (C) Copper-catalyzed azide–alkyne cycloaddition (CuAAC) reaction: this reaction occurs between an azide group (–N3) and an alkyne group (–C≡C–), forming a triazole ring (–N–N–C–N–).⁷⁸  To facilitate ring formation, this reaction necessitates the presence of a copper catalyst, such as Cu(i). However, copper is cytotoxic to cells, rendering this procedure unsuitable for preparing hydrogels in biomedical applications.

Enzyme crosslinking is a procedure that uses enzymes to induce the formation of covalent bonds between polymer chains (Figure 1.2a). This process depends on the type of enzyme and polymer used. A specific enzyme is chosen based on its ability to catalyze the desired chemical reaction within the polymer solution. Horseradish peroxidase (HRP),^79,80  transglutaminase (TG), tyrosinase,⁸¹  and laccase^82,83  are the enzymes commonly used for the crosslinking of hydrogels. The gelation process can be controlled by adjusting the enzyme concentration, reaction time, and reaction conditions. For example, horseradish peroxidase can oxidize various substrates in the presence of hydrogen peroxide, therefore, the enzyme can crosslink phenolic compounds by forming covalent bonds between their hydroxyl groups. Gelatin modified with tyramine groups can react with hydrogen peroxide in the presence of HRP to form covalent crosslinks, which is an application example.

Photo-crosslinking is a procedure that involves the use of light to trigger chemical reactions between the functional groups present in the polymer chains or molecules (Figure 1.2b). This process employs photoactive molecules, which, upon exposure to light of a specific wavelength, generate free radicals or other reactive species that can form covalent bonds between the functional groups of the polymer, such as disulfide bonds,^84,85  serving as both light-absorbing and cross-linking components. When exposed to light, these polymers can undergo changes in their properties or functions, such as transitioning from a liquid solution to a solid gel or becoming more rigid.⁸⁶  Additionally, the photocrosslinking process occurs rapidly upon light exposure, and the location and timing of the gelation process can be precisely controlled by adjusting the light intensity and exposure time, allowing for detailed adjustments to the properties and shape of the hydrogel. Gelatin methacryloyl (GelMA)^87,88  is a commonly used material in the preparation of injectable hydrogels through photo-crosslinking, and it has found various applications that include 3D bioprinting^89–92  and sensors.^93–95 

To fabricate injectable hydrogels on an industrial scale, the process should be executed in a stepwise manner, considering various parameters as outlined in the Ishikawa model diagram (Figure 1.3). The fabrication of injectable hydrogels depends on many parameters such as the starting material (type, solvents, active ingredients), environment (humidity, temperature, particle size), fabrication process (reaction time, temperature) and sterilizability. Also, other factors such as the reproducibility of the process, production environment and the transportation of raw material need to be considered. The complete process must be optimized from an economic point of view for commercial feasibility. Detailed analyses, from the preparation to the scale-up, are presented in different chapters of the book.

The conventional approach for administering hydrogel-based delivery systems involves implanting a pre-formed hydrogel scaffold that contains cells and/or bioactive agents. Nevertheless, achieving uniform distribution of substances within bulk hydrogels and effectively targeting specific sites in the body remains challenging. Injectable hydrogels are highly favoured biomaterials due to their capacity for local application via minimally invasive techniques.^98,99  As discussed in Section 1.2, this can be accomplished either by injecting a liquid precursor that subsequently transforms into a gel in situ or by utilizing pre-formed hydrogel particles designed to pass through a needle. Injectable hydrogels offer basic structural support, control the spatial and temporal delivery of cells or therapeutic substances, and locally attract and modify host cells to enhance tissue regrowth.¹⁰⁰  The high water content in the hydrogel and the porous structure with interconnected pores allow for the diffusion of nutrients and metabolites to maintain cell viability. Further, the viscoelastic characteristics of hydrogels allow them to serve as cell protectors, shielding cells from the shear stress during injection and preventing cell removal via the bloodstream.¹⁰¹  Hydrogels have been extensively studied for their biocompatibility, adjustable mechanical characteristics, permeability to oxygen and nutrients, and responsiveness to stimuli, to meet the specific requirements of different applications.¹⁰²  Figure 1.4 presents the use of injectable hydrogels in various biomedical applications.

Injectable hydrogels are ideal for repairing non-weight-bearing bone defects that do not require a high degree of mechanical resilience. Additionally, they serve as effective systems for the delivery of therapeutic biomolecules and cells to the target area. The rapid sol–gel transition of hydrogels promotes a regenerative response by improving the survival, adhesion, and integration of injected cells.¹⁰³  Injectable hydrogels have found clinical applications in various orthopaedic procedures, including cartilage repair,¹⁰⁴  bone void filling¹⁰⁵  and nucleus pulposus replacement.¹⁰⁶  Their responsiveness to stimuli like temperature, pH, and ionic strength, as well as their ability to mimic the ECM, makes them versatile for creating conducive environments for cell proliferation and tissue regeneration.^107,108  For example, injectable aldehyde-cellulose nanocrystalline and silk fibroin (ADCNCs/SF) hydrogels containing polyetheretherketone (PEEK) have accelerated internal bone fixation and bone cage assimilation.¹⁰⁹  In the craniomaxillofacial area, injectable hydrogels effectively regenerate bone by filling irregularly shaped and contoured defects.¹⁰⁹  Customizable injectable hydrogels, created by blending type I and type II collagens to control compressive strength, offer versatility for various applications.¹¹⁰  Another study introduced an injectable, click-crosslinked cytomodulin-enhanced HA acid hydrogel for cartilage tissue engineering.¹¹¹  This modified hydrogel, known as Cx-HA, exhibited increased stiffness and persisted in vitro and in vivo for an extended period of time. The findings of this study highlight the influence of hydrogel stiffness on chondrogenic potential, which can be modified through composition or chemical cross-linking. Furthermore, injectable hydrogels have been investigated as potential nucleus pulposus replacements, holding promise for restoring disc biomechanics and preventing height loss.^112,113 

An injectable conductive hydrogel composed of polypyrrole (PPy), gelatin and oxidized xanthan gum was used to restore electrical transmission at the myocardial infarct site to preserve cardiac function and enhance repair.¹¹⁴  Other studies have explored the use of injectable hydrogels composed of alginate and fibrin to locally deliver cardiomyocytes and vascular endothelial growth factor (VEGF) to the infarct area.¹¹⁵  Injectable hydrogels made of chitosan-graft-aniline tetramer (CS-AT) and dibenzaldehyde-terminated poly(ethylene glycol) (PEG-DA) were also investigated for cardiac regeneration, benefiting from their conductivity and self-healing properties.¹¹⁶ 

Injectable hydrogels have attracted significant attention for use in wound dressing due to their exceptional fluidity enabling them to fill deep and irregular wounds that conventional hydrogels cannot do. Moreover, the hydrated microenvironment created by these hydrogels promotes cell mitosis and migration, accelerating the wound healing process. Alginate and HA-based materials,^117,118  dopamine-modified gelatin@Ag nanoparticles,¹¹⁹  and injectable silk nanofiber hydrogels¹²⁰  have been successfully employed to enhance wound healing. Their viscoelastic properties make injectable hydrogels valuable for surgical applications, promoting tissue adhesion, preventing fluid leakage, and reducing infection risks.^121–123  The remarkable potential of injectable hydrogels for intraocular applications arises from their high water content, which enables transparency and a structure resembling the natural vitreous humor. These smart hydrogels have clinical potential and have been utilized as drug delivery systems or substitutes for tamponade in the treatment of eye diseases (Figure 1.5).^124,125 

Injectable hydrogels serve as controlled drug carriers due to their ability to release drugs gradually. The release rate can be tuned by adjusting hydrogel properties such as composition, crosslinking density, and the degradation rate. For instance, increasing crosslinking density leads to slower drug release due to enhanced diffusion resistance within the tighter network. Similarly, slow degrading hydrogels result in lower drug release rates, extending drug retention within the matrix. This precision allows for tailored drug delivery, improving treatment efficacy while minimizing systemic exposure and potential side effects. For example, in post-orthopaedic surgery pain management, conventional methods face limitations like short-duration analgesia and potential systemic toxicity. Injectable hydrogels offer an alternative for long-acting analgesia. In one study, a single peripheral nerve blockade using PLGA–PEG–PLGA gel-immobilized bupivacaine-loaded microspheres maintained sensory and motor blockade for over seven days without neurotoxicity.¹²⁷  Similarly, an injectable chitosan/β-glycerophosphate hydrogel delivering dexamethasone extended drug release over 28 days and led to reduced inflammation and joint damage in a rheumatoid joint rat model.¹²⁸  These hydrogels can respond to multiple stimuli, such as a chitosan-graft-polyaniline copolymer responding to electric fields and acidic pH in infected tissue environments.¹²⁹  Moreover, these gels can be used for the delivery of different drugs. For example, as shown in Figure 1.6, an injectable hydrogel (PPNPs/EPB@HA-Gel) combining paclitaxel nanoparticles (PPNPs) and epirubicin (EPB) into a HA hydrogel sustained drug release over 30 days, inhibiting breast cancer cell growth and improving survival in a mouse model, preventing postoperative recurrence and metastasis of breast cancer.¹³⁰ 

Injectable hydrogels are not only used for drug delivery but also designed to release growth factors that promote tissue healing and regeneration. For instance, a thermosensitive system was found to achieve a sustained release of bone morphogenetic protein-2 (BMP-2) through injectable polymeric nanoparticles (poly-phosphazene).¹³²  In another example, an in situ crosslinked hydrogel, composed of glycol chitosan (GC) and multialdehyde functionalized 4-arm PEG (4-arm PEG-CHO), incorporated transforming growth factor beta-1 (TGF beta-1) and bone mesenchymal stromal cells (BMSCs) to facilitate tissue regeneration.¹³³  Recent interest has emerged in exploring the biomedical applications of extracellular vesicles (EVs), which are released by cells of various types and contain diverse biological macromolecules critical for cellular functions.^134–136  However, a current limitation lies in the need for multiple administrations to achieve adequate tissue regeneration. Injectable hydrogels offer a solution to this drawback, allowing for more efficient delivery.^137,138  Furthermore, hydrogels can be engineered for complete degradation within the body, release of their payload and enhanced therapeutic outcomes. While injectable hydrogels hold immense potential in biomedical applications, addressing challenges such as safety, efficacy, mechanism of action, structural integrity, stability under physiological conditions, and regulatory compliance is crucial for successful clinical translation.

The viscoelastic properties of injectable hydrogels make them valuable for surgical applications, promoting tissue adhesion, preventing fluid leakage, and reducing infection risks.^123,139  Specifically, injectable smart hydrogels have shown promising properties for biomedical applications including in situ gelation in response to environment sensitivities, tunable mechanical properties, filling of irregular defects, the capability to be incorporated with bioactive compounds and biodegradability. Despite notable attempts in the fabrication of injectable hydrogels, several critical challenges still remain to be addressed. The safety of the regime is the primary priority. High risk based “new materials” should be considered for further testing before clinical studies, and the use of natural polymers such as polysaccharides for the fabrication of injectable hydrogels needs to be further explored. The exploration of simple methods or facile design for the production of injectable hydrogels needs to be carefully considered. The mechanical strength of hydrogels must be assessed appropriately, according to recognized standards. In drug delivery applications, sustained and acceptable release profiles represent an important factor that needs to be thoroughly investigated. The influence of the production process on the resultant hydrogel requires more investigation. The gelation parameter is another crucial parameter for in situ forming or injectable gels that must have a sol–gel transition near the targeted site and thus the gelation should not be too slow or too fast, as this may lead to clogging of the needle. Moreover, more in vivo studies need to be carried out to investigate the effect of degraded products on biological systems, or the metabolism pathways and the elimination processes of the byproducts. Advances in various fields of technology should guide us to design smarter injectable hydrogel systems with unique and multiple functionalities that can have suitable viscosity, desired gelation kinetics, robust mechanical properties, biomolecule release and stimuli-responsiveness, making them applicable for the treatment of more challenging clinical problems in future.

This publication was supported by the project MEBioSys with reg. no. CZ.02.01.01/00/22_008/0004634, co-funded by the ERDF as part of the MŠMT and funding received from the V4-Korea 2023 Joint Call. NA acknowledges funding from the National Institutes of Health (1UG3TR003148-01), the American Heart Association (23IPA1053441), the Estonian Research Council (PRG1903), and Henry Ford Health, Department of Surgery (00679169).